Nanostructures for the Engineering of Cells, Tissues

Nanostructures for the Engineering of Cells, Tissues and Organs

Related titles Nanobiomaterials in Soft Tissue Engineering (ISBN 978-0-323-42865-1) Electrospun Materials for Tissue Engineering and Biomedical Applications (ISBN 978-0-081-01022-8) Nanotechnology Applications for Tissue Engineering (ISBN 978-0-323-32889-0)

Nanostructures for the Engineering of Cells, Tissues and Organs From Design to Applications

Edited by

Alexandru Mihai Grumezescu Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Romania

William Andrew is an imprint of Elsevier The Boulevard, Langford Lane, Kidlington, Oxford, OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2018 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-813665-2 For Information on all William Andrew publications visit our website at https://www.elsevier.com/books-and-journals

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Contents List of Contributors .............................................................................................. xvii Series Preface: Pharmaceutical Nanotechnology...................................................xxi Preface ................................................................................................................. xxiii

CHAPTER 1 Cell and organ drug targeting: Types of drug delivery systems and advanced targeting strategies....................................................... 1 Imane Himri and Abdelkarim Guaadaoui 1.1 Introduction ....................................................................................1 1.2 Drug Targeting: What, Why and How? ........................................2 1.2.1 Definition and Reasons for Drug Targeting....................... 2 1.2.2 Drug Targeting Strategies................................................... 4 1.2.3 Properties Influencing Drug Targeting............................. 20 1.3 Cellular Targeting: Normal Cells and Abnormal Cells ..............30 1.4 Organ Targeting ...........................................................................31 1.4.1 Drug Delivery to Brain ..................................................... 31 1.4.2 Drug Delivery to Lung ..................................................... 32 1.4.3 Drug Delivery to Eye........................................................ 35 1.4.4 Drug Delivery to Neoplastic Disease ............................... 37 1.5 Conclusion ....................................................................................39 Glossary ....................................................................................... 39 Abbreviations............................................................................... 40 References.................................................................................... 41

CHAPTER 2 Cell-penetrating peptides in nanodelivery of nucleic acids and drugs ............................................. 67 Canan Ozyurt, Ozge Ugurlu and Serap Evran 2.1 Introduction ..................................................................................67 2.1.1 Cationic Cell-Penetrating Peptides ................................... 68 2.1.2 Amphipathic and Hydrophobic Cell-Penetrating Peptides ............................................................................. 69 2.2 Applications of Cell-Penetrating Peptides in Delivery of Therapeutic Molecules .................................................................70 2.2.1 Delivery of Drugs and Proteins ........................................ 70 2.2.2 Delivery of Nucleic Acids ................................................ 76

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2.3 Discovery and Design of Novel Cell-Penetrating Peptides ........83 2.4 Cell-Penetrating Bacterial Effector Proteins as Novel Tools......89 2.5 Recombinant Production of Cell-Penetrating Peptides for the Delivery Purpose ..............................................................90 2.6 Conclusion and Future Respects..................................................91 Acknowledgments ....................................................................... 92 References.................................................................................... 92

CHAPTER 3 The current perspectives of nanoparticles in cellular and organ-specific drug targeting in biological system...................................................... 105 3.1 3.2 3.3

3.4 3.5 3.6

3.7

Arunachalam Muthuraman Introduction ................................................................................106 Role of Nanoparticle Action in the Biological System.............107 3.2.1 Endocytosis Mechanism of Nanomedicine .................... 107 Mechanism of Nanoparticle Action in Cellular and Subcellular System.....................................................................111 3.3.1 Antigen-Specific Action of Nanoparticles ..................... 112 3.3.2 Receptor-Mediated Action of Nanoparticles.................. 113 3.3.3 Folate Receptor-Mediated Action of Nanoparticles....... 113 3.3.4 Transferrin Receptor-Mediated Action of Nanoparticles................................................................... 114 3.3.5 Epidermal Growth Factor Receptor-Mediated Action of Nanoparticles .................................................. 114 3.3.6 Integrins Receptor-Mediated Action of Nanoparticles................................................................... 115 3.3.7 Neonatal Fc-Receptor-Mediated Action of Nanoparticles................................................................... 115 Role of Nanoparticle Action in Pathophysiological Condition ....................................................................................116 Limitation of Nanoparticles Action in Biological Systems ......117 Intracellular and Subcellular Targeted Action ..........................117 3.6.1 Endosome/Lysosome-Targeted Action........................... 117 3.6.2 Cytoplasm-Targeted Action............................................ 119 3.6.3 Endoplasmic Reticulum and Golgi ApparatusTargeted Action ........................................... 119 3.6.4 Mitochondria-Targeted Action ....................................... 119 3.6.5 Nucleus-Targeted Action ................................................ 121 Interaction of Nanoparticle in Biological System .....................122

Contents

3.7.1 3.7.2 3.7.3 3.7.4

Interaction of Nanoparticle With Lipids ........................ 123 Interaction of Nanoparticle With Proteins ..................... 124 Interaction of Nanoparticles With DNA ........................ 125 Interaction of Nanoparticles With Other Smaller Biomolecules................................................................... 126 3.8 Pharmacological Action of Nanoparticle...................................126 3.8.1 Genomic Action of Nanoparticles .................................. 126 3.8.2 Proteomic Action of Nanoparticle.................................. 127 3.8.3 Metabonomic Action of Nanoparticle ............................ 127 3.9 Therapeutic Application of Nanoparticle ..................................128 3.9.1 Effect of Nanoparticles in Cancer .................................. 128 3.9.2 Effect of Nanoparticles in Vascular Disorders............... 129 3.9.3 Effect of Nanoparticles in Neurological Disorders........ 129 3.9.4 Effect of Nanoparticles in Infectious Disorders............. 130 3.9.5 Miscellaneous Action of Nanoparticles.......................... 131 3.10 Future Scopes .............................................................................132 Abbreviations............................................................................. 132 Acknowledgments ..................................................................... 134 References.................................................................................. 134 Further Reading ......................................................................... 154

CHAPTER 4 Precision medicine and drug targeting: The promise versus reality of target-specific drug delivery ............................................................. 155 4.1 4.2 4.3 4.4

Karel Petrak Precision Medicine.....................................................................155 Precision Drugs ..........................................................................157 Progress Towards Precision Drugs ............................................158 Conclusion ..................................................................................162 Abbreviations............................................................................. 163 References.................................................................................. 163

CHAPTER 5 Brain targeting of payload using mild magnetic field: Site specific delivery...................... 167 Murali M. Bommana and Sangram Raut 5.1 Magnetic Nanoparticles..............................................................167 5.1.1 Polymers Used in Magnetic Targeting........................... 167 5.1.2 Physical Characterization of Magnetic Nanoparticles................................................................... 171

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5.1.3 In Vivo Distribution of Nanoparticles: Mechanism....... 173 5.1.4 Passive and Active Targeting ......................................... 174 5.1.5 Targeting Brain Delivery Using Magnetic Nanoparticles................................................................... 174 5.1.6 Targeting Brain Tumors With Magnetic Nanoparticles................................................................... 175 5.1.7 In Vitro Characterization of Nanoparticles: In Vitro BloodBrain Barrier Model............................. 176 5.1.8 Feasibility of Superparamagnetic Iron Oxide Nanoparticles for Gene Delivery Systems: Magnetofection ............................................................... 177 5.2 Diagnostic Applications Using Magnetic Nanoparticles ..........178 5.2.1 Magnetic Resonance Imaging and Other Applications .................................................................... 178 5.2.2 Cancer Theranostics........................................................ 179 5.2.3 Magneto Acoustic Tomography ..................................... 180 5.2.4 Hyperthermia................................................................... 180 5.2.5 Challenges and Future Directions................................... 181 References.................................................................................. 182

CHAPTER 6 Nanoparticles influence in skin penetration of drugs: In vitro and in vivo characterization........ 187

6.1 6.2 6.3 6.4

6.5

Camila N. Lemos, Francieli Pereira, Luciana F. Dalmolin, Camila Cubayachi, Danielle N. Ramos and Renata F.V. Lopez Introduction ................................................................................187 Skin Structure .............................................................................189 6.2.1 Skin Appendages............................................................. 192 Mechanisms and Routes of Nanoparticles Skin Penetration.....192 Characteristics of Nanoparticles for Drug Skin Penetration .....196 6.4.1 Type of Nanoparticle ...................................................... 196 6.4.2 Size and Surface Area..................................................... 199 6.4.3 Charge ............................................................................. 200 6.4.4 Shape ............................................................................... 201 Physical Methods to Enhance Nanoparticle Skin Penetration..................................................................................201 6.5.1 Iontophoresis ................................................................... 201 6.5.2 Electroporation ................................................................ 205 6.5.3 Microneedles ................................................................... 207

Contents

6.6 Experimental Techniques for Studying Nanoparticle Skin Penetration..................................................................................208 6.6.1 In Vitro Studies ............................................................... 209 6.6.2 Ex Vivo Skin Penetration Experiments.......................... 223 6.6.3 In Vivo Skin Penetration Experiments........................... 224 6.7 Conclusion ..................................................................................229 References.................................................................................. 229

CHAPTER 7 DNA aptamer-based molecular nanoconstructions and nanodevices for diagnostics and therapy ........ 249 Elena Zavyalova and Alexey Kopylov 7.1 DNA Aptamers in Diagnostics and Therapy.............................250 7.2 Basic Principles of DNA Nanoconstruction Creation ...............252 7.2.1 Double Helices................................................................ 252 7.2.2 Triple Helices.................................................................. 254 7.2.3 G-Quadruplexes .............................................................. 255 7.2.4 i-Motifs............................................................................ 256 7.3 DNA Nanoconstructions for Aptamer Oligomerization............257 7.3.1 Functional Activity of Aptamer Homodimers ............... 258 7.3.2 Oligomeric Aptamers in Sensors.................................... 260 7.3.3 Lifetime of Oligomeric Aptamers In Vivo..................... 261 7.4 DNA Nanoconstructions With Different Aptamers ..................262 7.4.1 Thrombin Aptamer Hetero-Oligomers ........................... 263 7.4.2 Aptamer Heterodimers for Cell Targeting ..................... 266 7.5 Designing Extensive DNA Nanoconstructions..........................269 7.6 Examples of Extensive DNA Nanoconstruction Geometry ......271 7.6.1 DNA Tiles ....................................................................... 271 7.6.2 Two-Dimensional DNA Origami ................................... 271 7.6.3 Three-Dimensional DNA Origami ................................. 272 7.7 Promising Applications of DNA Tiles and DNA Origami .......273 7.7.1 Drug Delivery Systems ................................................... 273 7.7.2 Membrane-Associated DNA Nanoconstructions............ 275 7.7.3 Spatial Arrangement of the Molecules........................... 276 7.7.4 Biosensors ....................................................................... 279 7.7.5 Molecular Machines........................................................ 279 7.8 Conclusions ................................................................................282 Acknowledgments ..................................................................... 282 References.................................................................................. 282

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CHAPTER 8 Nanobiodevices for electrochemical biosensing of pharmaceuticals ................................................... 291 Sevinc Kurbanoglu, Bengi Uslu and Sibel A. Ozkan 8.1 Nanobiodevices ..........................................................................292 8.1.1 Calibration....................................................................... 294 8.1.2 High Sensitivity............................................................... 294 8.1.3 Wide Measurement Range.............................................. 295 8.1.4 Stability ........................................................................... 295 8.1.5 Lifetime ........................................................................... 295 8.1.6 Response Time................................................................ 296 8.1.7 Selectivity........................................................................ 296 8.1.8 Rapid Response Time ..................................................... 296 8.2 Nanobiodevices Based on Bio-Constituents..............................297 8.2.1 Biomaterial Immobilization in Nanobiodevices ............................................................... 298 8.2.2 Enzyme-Based Nanobiodevices...................................... 302 8.2.3 Microbial Nanobiodevices .............................................. 306 8.2.4 Immunosensors................................................................ 310 8.2.5 DNA-Based Nanobiodevices .......................................... 312 8.2.6 Tissue-Based Nanobiodevices ........................................ 318 8.3 Electrochemical Methods in Biosensing ...................................318 8.4 Conclusion ..................................................................................322 References.................................................................................. 323

CHAPTER 9 Imprinted polymeric nanoparticles as nanodevices, biosensors and biolabels .................. 331 9.1 9.2 9.3 9.4 9.5 9.6 9.7 9.8

´ Monika Sobiech and Piotr Lulinski Introduction ................................................................................331 Principles of Imprinting Process................................................332 Overview of Formats of Imprinted Nanomaterials ...................334 Imprinted Drug Delivery Nanodevices......................................335 Imprinted Nanosorbents for Sample Preparation ......................344 Imprinted Nanocomposites for Biosensors................................352 Biolabeling With Imprinted Polymeric Nanostructures............................................................................363 Conclusions ................................................................................363 References.................................................................................. 364 Further Reading ......................................................................... 374

Contents

CHAPTER 10 Poly(lactic-co-glycolic acid) (PLGA) matrix implants..................................................................... 375 10.1 10.2 10.3 10.4 10.5 10.6 10.7

10.8

10.9 10.10

Joana A.D. Sequeira, Ana C. Santos, Joaõ Serra, Francisco Veiga and Antońio J. Ribeiro Introduction ................................................................................376 Implantable Drug Delivery Systems..........................................376 Ability to Sustain and to Control Drug Delivery ......................377 The Issue of Biocompatibility....................................................379 Poly(Lactide-co-Glycolic Acid) (PLGA)...................................380 Biodegradability .........................................................................381 PLGA Matrix Implants ..............................................................381 10.7.1 PLGA Matrices as Sustained Drug Delivery Systems.......................................................................... 382 10.7.2 Manufacturing Techniques ........................................... 384 10.7.3 Drug Release ................................................................. 387 10.7.4 Factors Affecting Degradation and Also Drug Release From Degradable PLGA Matrices .................. 389 10.7.5 Therapeutic Peptides and Proteins Incorporated in PLGA Matrix Implants............................................. 391 Successful Case Studies .............................................................393 10.8.1 Zoladex.......................................................................... 393 10.8.2 Suprefact Depot............................................................. 394 10.8.3 Ozurdex ......................................................................... 395 Problems to Overcome and Opportunities.................................395 Conclusions ................................................................................397 References.................................................................................. 398

CHAPTER 11 Hydrogels for biomedical applications.................... 403 Luciane R. Feksa, Eduardo A. Troian, Cristina D. Muller, Fabian Viegas, Aline B. Machado and Virgıńia C. Rech 11.1 Hydrogels: Concepts and Definitions ........................................404 11.2 Classification of Hydrogels........................................................407 11.2.1 Classification Based on Response ................................ 407 11.2.2 Classification Based on Type of Cross-Linking .......... 409 11.2.3 Classification According to the Method of Preparation .................................................................... 409 11.2.4 Classification Based on Source .................................... 409 11.2.5 Intelligent Hydrogels .................................................... 412 11.3 Applications of Hydrogels .........................................................414

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11.4 11.5 11.6 11.7 11.8 11.9

11.3.1 Tissue Regeneration...................................................... 414 11.3.2 Tissue Dressing ............................................................. 417 11.3.3 Contact Lenses .............................................................. 419 11.3.4 Drug Delivery System .................................................. 420 11.3.5 Hygiene Products .......................................................... 421 11.3.6 Others Applications....................................................... 423 Hydrogel Technical Features .....................................................424 Metabolism and Hydrogels ........................................................424 Regulation of Hydrogels ............................................................425 Potential Risks of Hydrogels .....................................................426 Nanoparticle Biosafety...............................................................427 Final Remarks ............................................................................428 References.................................................................................. 429 Further Reading ......................................................................... 438

CHAPTER 12 Silk-based matrices for bone tissue engineering applications.......................................... 439 12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10 12.11 12.12

Promita Bhattacharjee, Prerak Gupta, M. Joseph Christakiran, Samit K. Nandi and Biman B. Mandal Introduction ................................................................................440 Global Perspective on Orthopaedic Trauma Management .......441 Rising Needs for Bone Grafts....................................................441 Ideal Scaffold for Bone Tissue Engineering .............................442 Lacunae of Current Materials and Practices for Bone Regeneration...............................................................................443 Bioderived and Synthetic Materials for Bone Tissue Engineering.................................................................................444 Context and Structure of the Chapter ........................................445 Various Sources of Silk .............................................................445 Silk From Silkworms .................................................................447 Spider Silk ..................................................................................448 Benign Aspects of Silk for Bone Tissue Engineering...............448 Processing Silk Into Various Formats .......................................450 12.12.1 Particulate Leaching.................................................... 450 12.12.2 Nanofibrous Scaffolds Using Electrospinning ........... 451 12.12.3 Biopatterning............................................................... 452 12.12.4 Knitted Scaffolds......................................................... 452 12.12.5 Freeze-Drying.............................................................. 452

Contents

12.12.6 Silk Microparticles and Microfibers as Reinforcements ........................................................... 453 12.12.7 Hydrogels .................................................................... 453 12.12.8 Computer-Aided Fabrication of SF Scaffolds (3D Printing) ............................................................... 453 12.13 Silk Composites for Bone Tissue Engineering..........................454 12.13.1 Hydroxyapatite ............................................................ 454 12.13.2 Clay- or Silica-Based Additives ................................. 456 12.13.3 Bioactive Glasses ........................................................ 456 12.13.4 Silk Inclusion in Other Substrates .............................. 457 12.13.5 Other Miscellaneous Additives for Silk Scaffolds..... 458 12.14 Beyond the Classic Mulberry Silk Fibroin................................458 12.14.1 Nonmulberry Silk........................................................ 458 12.14.2 Spider Silk................................................................... 459 12.14.3 Silk Sericin.................................................................. 460 12.15 Recent Trends in Bone Tissue Engineering ..............................460 12.15.1 Use of Bioreactors ...................................................... 460 12.15.2 Cocultures of Multiple Cells on Silk Scaffolds ......... 461 12.15.3 Growth Factors Delivery Through Silk Scaffolds...... 462 12.15.4 Gene Therapies ........................................................... 462 12.15.5 Conclusion and Future Perspectives........................... 463 References.................................................................................. 464

CHAPTER 13 Implantable drug delivery systems: An overview............................................................... 473 13.1 13.2

13.3

13.4

Anoop Kumar and Jonathan Pillai Introduction ................................................................................473 Classification of Implantable Drug Delivery Systems ..............479 13.2.1 Passive Implants............................................................ 479 13.2.2 Dynamic Implants ......................................................... 484 13.2.3 Electromechanical Systems .......................................... 487 Design Approaches.....................................................................489 13.3.1 Implant Material Selection ........................................... 489 13.3.2 Mechanisms of Drug Release From Implantable Drug Delivery Systems............................. 490 Current Therapeutic Applications..............................................492 13.4.1 Women’s Health ........................................................... 492 13.4.2 Chronic Diseases........................................................... 492

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13.4.3 Infectious Diseases (Tuberculosis) ............................... 495 13.4.4 Neurology and Central Nervous System Health .......... 496 13.5 Current Challenges and Future Perspectives .............................496 13.5.1 Biocompatibility-Related Issues ................................... 496 13.5.2 Patient Compliance ....................................................... 498 13.5.3 Regulatory Aspects ....................................................... 498 13.5.4 Cost-Effectiveness......................................................... 499 13.5.5 Future Perspectives ....................................................... 499 13.6 Conclusion ..................................................................................500 Acknowledgments ..................................................................... 501 References.................................................................................. 501

CHAPTER 14 Nanobionics and nanoengineered prosthetics ........ 513 14.1 14.2 14.3 14.4

14.5

14.6

Hemant K.S. Yadav, Ghufran A. Alsalloum and Noor A. Al Halabi Introduction ................................................................................514 History ........................................................................................517 Definition....................................................................................520 Types and Classifications...........................................................523 14.4.1 Orthopedic Prostheses................................................... 524 14.4.2 Plastic and Reconstructive Prostheses.......................... 526 14.4.3 Neuroprostheses ............................................................ 526 14.4.4 Cerebrospinal Fluid Drainage Systems ........................ 527 14.4.5 Ophthalmic Prostheses.................................................. 527 14.4.6 Cardiovascular Prostheses............................................. 527 14.4.7 Myoelectric Prostheses ................................................. 528 14.4.8 Dental Prostheses .......................................................... 528 Manufacture................................................................................529 14.5.1 Lithography ................................................................... 529 14.5.2 Photolithography ........................................................... 530 14.5.3 Beam Lithography ........................................................ 530 14.5.4 Micro and Nano Contact Printing ................................ 530 14.5.5 Jet Printing .................................................................... 531 14.5.6 Scan Probe Lithography................................................ 531 14.5.7 Dip-Pen Nanolithography ............................................. 532 Nanobiomaterials........................................................................532 14.6.1 Polymeric Materials ...................................................... 533 14.6.2 Nanotitanium (NanoTi)................................................. 535 14.6.3 Carbon Nanotubes......................................................... 536

Contents

14.7

14.8

14.9 14.10

14.6.4 Nanodiamonds............................................................... 539 14.6.5 Nanobioceramic ............................................................ 540 14.6.6 Nanocomposite.............................................................. 543 14.6.7 Peekpolymer.................................................................. 549 14.6.8 Hydrogel........................................................................ 549 Applications................................................................................550 14.7.1 Orthopedic Prostheses................................................... 550 14.7.2 Neuroprostheses ............................................................ 566 14.7.3 Cardiovascular Prostheses............................................. 571 14.7.4 Cerebrospinal Fluid Drainage Systems ........................ 572 14.7.5 Plastic and Reconstructive Prostheses.......................... 573 Ethical Issues..............................................................................574 14.8.1 Safe Use: Benefits Versus Risks .................................. 575 14.8.2 Justice ............................................................................ 575 14.8.3 Identity, Privacy, and Accountability ........................... 575 14.8.4 Autonomy...................................................................... 576 14.8.5 Validity of Informed Consent....................................... 576 14.8.6 Problems of Ambition: Treatment Versus Enhancement ................................................................. 577 Safety Issues Pertinent to Nanobionics and Prosthetics............577 Conclusion ..................................................................................579 References.................................................................................. 580 Further Reading ......................................................................... 587

Index ......................................................................................................................589

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List of Contributors Noor A. Al Halabi RAK Medical and Health Sciences University, Ras al Khaimah, United Arab Emirates Ghufran A. Alsalloum RAK Medical and Health Sciences University, Ras al Khaimah, United Arab Emirates Promita Bhattacharjee Indian Institute of Technology, Kharagpur, West Bengal, India Murali M. Bommana Impax Labs, Middlesex, NJ, United States Camila Cubayachi School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil Luciana F. Dalmolin School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil Serap Evran Ege University, Izmir, Turkey Luciane R. Feksa Feevale University, Novo Hamburgo, RS, Brazil; Federal University of Rio Grande do Sul, Porto Alegre, RS, Brazil Abdelkarim Guaadaoui University Mohammed Premier, Oujda, Morocco Prerak Gupta Indian Institute of Technology Guwahati, Guwahati, Assam, India Imane Himri University Mohammed Premier, Oujda, Morocco M. Joseph Christakiran Indian Institute of Technology Guwahati, Guwahati, Assam, India Alexey Kopylov Chemistry Department of Lomonosov Moscow State University, Moscow, Russian Federation Anoop Kumar Translational Health Science & Technology Institute (THSTI), Faridabad, India; Indo-Soviet Friendship College of Pharmacy (ISFCP), Moga, Punjab, India

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List of Contributors

Sevinc Kurbanoglu Ankara University, Ankara, Turkey Camila N. Lemos School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil Renata F.V. Lopez School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil ´ Piotr Lulinski Medical University of Warsaw, Warsaw, Poland Aline B. Machado Feevale University, Novo Hamburgo, RS, Brazil Biman B. Mandal Indian Institute of Technology Guwahati, Guwahati, Assam, India Cristina D. Muller Feevale University, Novo Hamburgo, RS, Brazil Arunachalam Muthuraman JSS University, Mysuru, India Samit K. Nandi West Bengal University of Animal and Fishery Sciences, Kolkata, West Bengal, India Sibel A. Ozkan Ankara University, Ankara, Turkey Canan Ozyurt Ege University, Izmir, Turkey Francieli Pereira School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil Karel Petrak NangioTx Inc., New York, NY, United States Jonathan Pillai Translational Health Science & Technology Institute (THSTI), Faridabad, India Danielle N. Ramos School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil Sangram Raut Texas Christian University, Fort Worth, TX, United States

List of Contributors

Virgıńia C. Rech Federal University of Rio Grande do Sul, Porto Alegre, RS, Brazil; Franciscan University, Santa Maria, RS, Brazil Antońio J. Ribeiro University of Coimbra, Coimbra, Portugal; Institute for Molecular and Cell Biology, Porto, Portugal Ana C. Santos University of Coimbra, Coimbra, Portugal; Institute for Molecular and Cell Biology, Porto, Portugal Joana A.D. Sequeira University of Coimbra, Coimbra, Portugal Joaõ Serra Tecnimede Group SA, Dois Portos, Portugal Monika Sobiech Medical University of Warsaw, Warsaw, Poland Eduardo A. Troian Feevale University, Novo Hamburgo, RS, Brazil Ozge Ugurlu Ege University, Izmir, Turkey Bengi Uslu Ankara University, Ankara, Turkey Francisco Veiga University of Coimbra, Coimbra, Portugal Fabian Viegas Feevale University, Novo Hamburgo, RS, Brazil Hemant K.S. Yadav RAK Medical and Health Sciences University, Ras al Khaimah, United Arab Emirates Elena Zavyalova Chemistry Department of Lomonosov Moscow State University, Moscow, Russian Federation

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Series Preface: Pharmaceutical Nanotechnology Due to its immense applicative potential, nanotechnology is considered the leading technology of the 21st century. The science and engineering of nanometer-sized materials is currently employed for the development of numerous scientific, industrial, ecological, and technological fields. Biology, medicine, chemistry, pharmacy, agriculture, food industry, and material science are the main fields which have benefited from the great technological progress developed in nanoscience. In the pharmaceutical field, nanotechnology has revolutionized traditional drug-design concept and the art of drug delivery. The idea of a highly specific nanoscale drug for the targeted therapy of diseases is now considered a feasible treatment for severe health conditions. Some scientists believe that the pharmaceutical domain has been reborn by the important contribution of nanotechnology. The field of pharmaceutical nanotechnology has the potential to offer innovative solutions for all diagnosis, therapy, and prophylaxis domains. Application of nanotechnology tools in pharmaceutical research and design is likely to result in moving the industry from a “blockbuster drug” model to “personalized medicine.” The current main focus of clinicians is to treat patients individually, not their general diagnosed diseases, which are usually difficult to diagnose or incorrectly diagnosed. There are compelling applications in the pharmaceutical industry where suitable nanotechnology tools can be successfully utilized. By designing and modifying drugs at nanoscale, pharmaceutical nanotechnology could be useful not only for the development of completely new therapeutic solutions, but also to add value to existing products. This possibility opens perspectives of success for pharmaceutical companies in existing markets, but also for new markets. Scientists have manifested an impressive interest on the field of pharmaceutical nanotechnology research in recent years. However, we face today a true dilemma of data unavailability, due to the multitude of existing information which can be highly inaccurate and contradictory. This is because of the lack of an efficient model for sorting the plethora of nanotechnology tools and information that exists, and strategically correlate those with potential opportunities into different segments of pharmaceutical research and design. This series is trying to cover the most relevant aspects regarding the great progress of nanotechnology in the pharmaceutical field and to highlight the currently emerging trend of pharmaceutical nanotechnology towards the personalized medicine concept. The 10 volumes of this series are structured to wisely offer relevant information regarding basic concepts and also to reveal the newest approaches and perspectives in pharmaceutical nanotechnology.

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Series Preface: Pharmaceutical Nanotechnology

Nanoscale Fabrication, Optimization, Scale-Up and Biological Aspects of Pharmaceutical Nanotechnology, introduces the readers into the amazing field of nanoscale design. Also, this volume facilitate understanding of the biological requirements of nanostructured pharmaceutical formulations for advanced drugs. In Design and Development of New Nanocarriers, the most recent progress made on the field of nano-delivery is discussed. Modern nanostructured drug carriers employ innovative solutions for the detection and treatment of various diseases in a personalized and efficient manner. Design of Nanostructures for Theranostics Applications, highlights the impressive impact of nanotechnology in the development of combined diagnosis and therapy concept: theranostics. Design of Nanostructures for Versatile Therapeutic Applications, offers a dynamic solution for immune modulation, treatment of diseases by natural-based products and infection control, while employing nanostructured solutions to achieve top results. Nanostructures for the Engineering of Cells, Tissues and Organs: From Design to Applications, is a highly investigated and debated field; tissue engineering, is dissected through this volume. Here is shown how nanotechnology has advanced research and applications in the manipulation and engineering of cells and tissues in vitro. Organic Materials as Smart Nanocarriers for Drug Delivery, deals with the specific world of organic nanomaterials, revealing their wide applications, types, and advantages in drug delivery. In the volume entitled: Inorganic Frameworks as Smart Nanomedicines, the main focus is to discuss the variety and properties of inorganic nanostructures for therapy and drug delivery in the context of improved personalized medicine. Lipid Nanocarriers for Drug Targeting, deals with recently developed lipid nanostructures and the advances made in drug targeting. Drug Targeting and Stimuli-Sensitive Drug Delivery Systems, dissects smart stimuli-responsive nanosystems employed to specifically detect various biochemical conditions and control the release of drugs. Fullerens, Graphenes and Nanotubes: A Pharmaceutical Approach, reveals major findings made on widely applied drug-design nanosystems, namely fullerens, graphenes and nanotubes. The impact of these nanostructures in pharmaceutical research is highlighted. All 10 volumes are nicely illustrated and chapters are organized into a logical manner to be accessible to a wide audience. The series is a valuable resource of new and comprehensive scientific proof on the intriguing and emerging field of pharmaceutical nanotechnology, which could be of a great use for scientists, engineers, pharmaceutical representatives, clinicians, and any non-specialist interested user. Alina M. Holban University of Bucharest, Bucharest, Romania

Preface The aim of this reference book is to present the novel progress from recent years in the field of pharmaceutical nanotechnology, with special attention to engineering of cell, tissues, and organs. The book provides an up-to-date overview about organ targeting and cell targeting using nanotechnology. Different biomedical applications are presented with pros and cons, as follows: nanobionics and nanoengineered prosthetics, molecular nanoconstructions and nanodevices for diagnostics and therapy, precision medicine and drug targeting, brain targeting and many others. The book entitled Nanostructures for the Engineering of Cells, Tissues and Organs: From Design to Applications, contains 14 chapters, prepared by outstanding researchers from United Arab Emirates, Portugal, Poland, Russian Federation, Brazil, United States, India, Turkey, and Morocco. Chapter 1, Cell and organ drug targeting: Types of drug delivery systems and advanced targeting strategies, prepared by Himri Imane et al., gives an up-to-date overview about the targeted delivery of drugs to diseased tissue without affecting the original characteristics of the tissue. It highlights a smart approach to increase the therapeutic index of a drug to specifically deliver the therapeutic molecule in its active form, not only into target tissue, nor even to target organs, but more importantly, into the targeted cells. Chapter 2, Cell-penetrating peptides in nanodelivery of nucleic acids and drugs, prepared by Canan Ozyurt et al., summarizes the use of cell-penetrating peptides in enhancing the gene transfer efficiency of nonviral vectors. Besides the clinical potential of currently known cell-penetrating peptides, they also discuss the limitations and the need for designing novel cell-penetrating peptides. Chapter 3, The current perspectives of nanoparticles in cellular and organspecific drug targeting in biological system, prepared by Arunachalam Muthuraman, focuses on the cellular and organ-specific action of nanoparticles in a biological system. In addition, it also emphasizes the possible mechanism to overcome the limitation of nanoparticles. Furthermore, the interaction between nanoparticle and cells are explained by the cellular uptake and the action on subcellular compartments process. It also covers the factors affecting nanoparticle action i.e., activity on cell membranes, ion channels, cytoskeletal proteins, mitochondria, and nucleus; interaction with proteins, lipids, DNA and small molecules; alteration of cellular signaling, genomic, proteomic and metabonomic processes in the biological system. The detailed overview of this chapter can deliver successful nanomedicine in various disorders including cancer, vascular, and neurodegenerative disorders. Chapter 4, Precision medicine and drug targeting: The promise versus reality of target-specific drug delivery, prepared by Karel Petrak, give an up-to-date overview about precision medicine, which is an initiative is to improve disease

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treatment and prevention by taking into account individual variability in genes, environment, and lifestyle for each person. Successful application of the new knowledge generated by precision medicine will ultimately require the availability of precision drugs; drugs that act exclusively on the disease molecular targets without impacting adversely on the rest of the body. This review examines the recent progress in developing such site-specific targeting drug systems. Chapter 5, Brain targeting of payload using mild magnetic field: Site specific delivery, prepared by Murali Mohan Bommana et al., reviews the applications of magnetic nanoparticles in brain-targeting and diagnostic applications. Chapter 6, Nanoparticles influence in skin penetration of drugs: In vitro and in vivo characterization, prepared by Camila Nunes Lemos et al., presents the recent progress of the nanoparticle characteristics affecting skin drug penetration, their possible penetration pathways, in addition to in vitro and in vivo techniques commonly used in assessing penetration and distribution of drugs into the skin where nanoparticles are used as topical delivery systems. Chapter 7, DNA aptamer-based molecular nanoconstructions and nanodevices for diagnostics and therapy, prepared by Elena Zavyalova et al., provides a thorough discussion of DNA aptamer-based nanoconstructions and nanodevices. The challenges and the opportunities in the field are also discussed. Chapter 8, Nanobiodevices for electrochemical biosensing of pharmaceuticals, prepared by Sevinc Kurbanoglu et al., presents the recent progress of nanobiodevices for electrochemical biosensing of pharmaceuticals related to different biorecognition parts, such as enzymes, DNA, tissues, bacteria, yeast, antibodies, antigens, liposomes, and organelles. Nanobiodevices such as Lab-on-a-chip platforms in biosensors and their applications in pharmaceutical analysis are discussed Chapter 9, Imprinted polymeric nanoparticles as nanodevices, biosensors and biolabels, prepared by Monika Sobiech et al., presents the recent advances in synthetic approaches for fabrication of imprinted nanomaterials, together with diversity of formats for possible applications in pharmaceutical science. The physicochemical behavior of imprinted nanostructures are discussed in context of their practical utility. Finally, the current limits and future prospects for the imprinted nanomaterials were pointed out. Chapter 10, Poly(lactic-co-glycolic acid) (PLGA) matrix implants, prepared by Joana A. D. Sequeira et al., begins offering a brief review about poly(esters) in general, with special emphasis on PLGA. It discusses current methods used to produce PLGA matrix implants, focusing on PLGA matrix implants for the delivery of therapeutic peptides. Commercially successful examples are revealed, and a reflection is made on future directions and potentialities. Chapter 11, Hydrogels for biomedical applications, prepared by Luciane Rosa Feksa et al., reports recent progress in the field of hydrogels used for biomedical purposes, demonstrating that most of the definitions are still under construction. In this context, new drug delivery systems, wound dressings, and contact lenses

Preface

are being developed, although not limited to these applications. Because they are considered recent technology, there is still little information or studies regarding their behavior in biological systems. Finally, comments are made about the marketing of products containing nanomaterials. Chapter 12, Silk-based matrices for bone tissue engineering applications, prepared by Promita Bhattacharjee et al., gives an overview of the field from a perspective of materials and fabrication. Silk is a biopolymer with several characteristics, including excellent biocompatibility and mechanical strength that make it a potential candidate for various tissue engineering applications. A growing trend is observed towards designing mineralized nanofibrous and composite scaffolds. Chapter 13, Implantable drug delivery systems: An overview, prepared by Anoop Kumar et al., starts with a review of various types of implantable drug delivery systems from biomaterial-based to electro-mechanical systems. Furthermore, design approaches to optimal drug delivery, including methods to tailor drug release profiles and the mechanism of release kinetics, are presented. Potential therapeutic applications and biocompatibility related issues are briefly discussed. Finally, this chapter concludes with a summary of future perspectives of implantable drug delivery systems, particularly in their applicability to precision and personalized medicine. Chapter 14, Nanobionics and nanoengineered prosthetics, prepared by Hemant K.S. Yadav et al., focuses on the impact of nanotechnology on medical bionic devices. The contributors discuss the engineering and manufacture of prosthetics, their uses and applications, and recent advances in this field. Alexandru M. Grumezescu University Politehnica of Bucharest, Bucharest, Romania

www.grumezescu.com

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Cell and organ drug targeting: Types of drug delivery systems and advanced targeting strategies

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Imane Himri and Abdelkarim Guaadaoui University Mohammed Premier, Oujda, Morocco

CHAPTER OUTLINE 1.1 Introduction ......................................................................................................... 1 1.2 Drug Targeting: What, Why and How?.................................................................... 2 1.2.1 Definition and Reasons for Drug Targeting ............................................2 1.2.2 Drug Targeting Strategies ....................................................................4 1.2.3 Properties Influencing Drug Targeting.................................................20 1.3 Cellular Targeting: Normal Cells and Abnormal Cells............................................ 30 1.4 Organ Targeting ................................................................................................. 31 1.4.1 Drug Delivery to Brain .......................................................................31 1.4.2 Drug Delivery to Lung .......................................................................32 1.4.3 Drug Delivery to Eye .........................................................................35 1.4.4 Drug Delivery to Neoplastic Disease ...................................................37 1.5 Conclusion ........................................................................................................ 39 Glossary ................................................................................................................... 39 Abbreviations............................................................................................................ 40 References ............................................................................................................... 41

1.1 INTRODUCTION Nanotechnology is a multidisciplinary branch which consists of manufacturing nanometer-sized structures and materials. It combines the elements of molecular biology, engineering and chemistry. Nanotechnology is one of the most dynamically developing branches of science and technology (Niemirowicz and Car,

Nanostructures for the Engineering of Cells, Tissues and Organs. DOI: http://dx.doi.org/10.1016/B978-0-12-813665-2.00001-6 © 2018 Elsevier Inc. All rights reserved.

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CHAPTER 1 Cell and organ drug targeting

2012). In the past two decades, nanotechnology has been developing quickly and is widely used in medical sciences, specifically for disease diagnosis and treatment (Desai, 2012; Gao, 2016). Delivering drugs or biocompounactives to the target site is a major problem in the treatment of many diseases (Wilczewska et al., 2012). Reducing the size of a selected material to nanometric scale makes it possible to utilize them in numerous potential applications, including drug targeting (Niemirowicz and Car, 2012). Recent developments in nanotechnology have shown that nanostructures have great potential as drug carriers. Due to their small sizes, nanocarriers exhibit unique physicochemical and biological properties that make them a favorable material for biomedical applications (Wilczewska et al., 2012). The pharmacokinetic behavior of the drug-loaded nanocarriers depends on the nanosystems and the modes of targeting (passive or active) (Martin Scha¨ffler et al., 2014).

1.2 DRUG TARGETING: WHAT, WHY AND HOW? 1.2.1 DEFINITION AND REASONS FOR DRUG TARGETING Drug targeting is a nano-biotechnological method of delivering a biocompounactive (bioactive compound) or an active pharmaceutical ingredient to a patient, in order to increase its concentration in the intended site of action, in a specific part of diseased tissue, and avoiding interaction with healthy tissues. This nanopharmaceutical method requires various disciplines (biologists, chemists, engineers, etc.) and is believed to improve efficacy, while reducing side effects (maximum efficacy with minimal toxicity). The drug discovery process has been accelerated with the development of modern technology in pharmaceutical chemistry and molecular biology (i.e., drug design, combinatorial chemistry, high throughput screening, etc.). Nevertheless, this increase in complexity does not necessarily offer more efficient drugs, even if these molecules often possess physicochemical and/or biological characteristics that make their use suboptimal in humans (Bertrand and Leroux, 2012; Chang et al., 2015). In fact, new drug candidates often exhibit many problems and challenges such as: (1) poor solubility, (2) insufficient in vitro stability (shelf life), (3) too low bioavailability, (4) too short in vivo stability (half-life), (5) strong side effect, (6) need for targeted delivery, (7) regulatory issues/hurdles, and/or (8) lack of largescale production (Muller and Keck, 2004; Chang et al., 2015). Consequently, the introduction of biotechnological methods for the production of drugs delivery brought a revolution to this biopharmaceutical field for advanced drug development (Muller and Keck, 2004).

1.2 Drug Targeting: What, Why and How?

In chemotherapy, for example, which remains the main form of treatment for cancer, only a small portion of drugs administered typically reach the organ to be affected (the tumor) since there is no clinically available antineoplastic drug that acts selectively on the tumor mass. For this reason, the scientific research is focused towards the development of novel cancer therapies and drug delivery strategies, such as drug targeting, that would enhance the therapeutic efficacy of drugs while reducing their side effects (Basile et al., 2012). Moreover, to reach the site of action, the biocompounactive has to transit many biological barriers, such as organs, cells, and intracellular compartments (blood, kidney, liver, spleen, etc.), where it can be inactivated or express undesirable effects on organs and tissues that are not involved in the pathological process. As a result, to achieve a required therapeutic concentration of an active pharmaceutical ingredient in a certain body compartment or certain tissue, one has to administer the drug in large quantities (thus increasing the cost of the therapy), the great part of which, even in the best case scenario, is wasted in normal tissues; cytotoxic and/or antigenic/immunogenic agents can become the cause of many negative side effects. Drug targeting can bring a solution to all these problems (Torchilin, 2010; Bertrand and Leroux, 2012). The main goal of drug targeting is to localize, target, prolong, and have a protected drug interaction with the diseased tissue (Fig. 1.1). Drug targeting may resolve many problems currently associated with systemic drug administration (orally or as injectables), such as: (a) pharmaceuticals biodistribution, (b) the necessary dose of a drug, (c) lacking affinity between drug-pathological site, (d) the adverse side effects, etc. (Torchilin, 2000; Mishra et al., 2013).

FIGURE 1.1 Reasons for drug targeting as referred by Agnihotri et al. (2011)-modified.

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1.2.2 DRUG TARGETING STRATEGIES 1.2.2.1 Common approaches of targeted drug delivery Targeted drug delivery is the ability of the drug to accumulate in the target tissue or organ selectively and quantitatively, independent of the site and methods of administrations. The aim of targeted drug delivery is to obtain high local concentrations of drug in the target area without any side effects in normal tissues, together with low systemic exposure (Fallis, 2002; Hirsja¨rvi et al., 2011; Mishra et al., 2013). Targeting may have spatial and temporal properties which deliver the right amount of drug to the right place (double-targeting) (Goodman et al., 2008; Shaji and Lal, 2013). Consequently, targeted drug delivery presents many advantages, the most important of which are (Fallis, 2002): (1) simplification of administration protocols; (2) drastic reduction in the cost of therapy and drug quantity required to achieve a therapeutic effect; and (3) sharp increase in drug concentration in the required sites without negative effects on nontarget areas. Currently, the principal strategies of drug targeting include many schemes (Torchilin, 2000). Among these various approaches, passive and active targeting seem to be most advanced, and could be relied onto achieve organ-based targeting (first order), specific cell-based targeting in an organ (second order) and cell organelle-based targeting (third order) (Torchilin, 2000, 2010; Danhier et al., 2010). Both passive and active drug targeting reduce toxic side effects, increase efficacy, and enhance delivery of poorly soluble or sensitive therapeutic molecules (Hirsja¨rvi et al., 2011).

1.2.2.1.1 Passive drug targeting In the human body, some molecules (i.e., hormones, growth-factors) have a natural tendency to target their receptors (sites of action) by the action of physicochemical and pathophysiological factors. This process is called “passive targeting,” and can also be applied to drugs (Garnett, 2001). In fact, passive drug targeting can benefit from the presence of (1) physicochemical modifications under diseased conditions, like internal stimuli (pH, temperature, etc.) (X. Zhang et al., 2010), and (2) modified physiologies, such as structural changes (i.e., leaky vasculature) in the microenvironment of inflammatory tissues (Crielaard et al., 2012). Passive drug targeting (or enhanced permeation and retention (EPR) effectmediated targeting) is based on the longevity of the pharmaceutical carrier in the blood and its accumulation in pathological sites with compromised vasculature (Torchilin, 2010). For example, drugs can penetrate the tumor vasculature through its leaky endothelium and, in this way, accumulate in several solid


tumors. This is called the enhanced permeation and retention (EPR) effect (Hirsja¨rvi et al., 2011; Nakamura et al., 2016). The EPR effect is specifically responsible for passive drug targeting in cancer tissues (Greish, 2010; Torchilin, 2011; Maeda, 2012). Drug targeting systems will be stimulated by such modifications to release the drug only at the diseased site and spare the untargeted tissues. However, the targeting potential of such a strategy is relatively low and often associated with partial nonspecific localization of therapeutics in the normal tissues, which needs to be considered while employing such therapies. (Thanki et al., 2015; Nakamura et al., 2016). Additionally, passive drug targeting relies on the basic defense mechanism of the reticulo-endothelial system (RES), which is part of the immune system, consisting on the phagocytic cells, such as monocytes and macrophages. Drugs or drug carriers can be taken up by the RES in the liver, spleen, lung, lymph nodes, etc. by opsonization (via C3, C4, and C5 complements) and phagocytosis processes (Nie, 2010). In this case, the passive targeting strategy may be designed outside the RES (i.e., RES blockade), and can be explored for conditions wherein the RES is the target site of action (Allen and Chonn, 1987; Lammers et al., 2012; Liu et al., 2015). Another approach to prolong drugs’ blood circulation time is by employing polymers, such as polyethylene glycol (PEG). PEGylation of therapeutic agents is an established technology used to enhance the bioavailability and prolongs blood circulation of an active pharmaceutical ingredient in the body of patients (Mohs et al., 2014). PEGylation is a process by which one or more PEG-chains are attached to a biocompounactive. PEGs are hydrophilic polymers that are nontoxic, nonimmunogenic, nonantigenic, and FDA approved. The PEG-drug conjugates have several advantages: a prolonged residence in body (stealth characteristics), a decreased degradation by metabolic enzymes, and a decreased uptake by mononuclear phagocyte system (MPS) cells. By these favorable properties, PEGylation improves pharmacokinetics and enhances the potentials of therapeutic agents (Harris and Chess, 2003; Veronese and Pasut, 2005; Huynh et al., 2010; Gokarn et al., 2012; Salmaso and Caliceti, 2013).

1.2.2.1.2 Active drug targeting Unlike passive targeting, which is a nonspecific strategy, active drug targeting is a specific approach that involves interactions between specific biological pairs/ systems, such as ligand-receptor, antigen-antibody, and enzyme-substrate. In active targeting, therapeutics can be also transported specifically to relevant cells through stimuli responsive nanocarriers (temperature, ultrasound, magnetic field) (Fleige et al., 2012; Kong et al., 2012; Mura et al., 2013). Active targeting strategy is based on the anchoring (attachment) of active agents or ligands to the surface of drug delivery system (DDS), which is selectively and specifically recognized by the target in concern. Attaching these active

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agents/ligands can be mediated by various mechanisms and facilitates greater uptake. This approach provides selectivity, recognizability and potential to interact (bind) with pathological cells and specific tissues in the body (Nishioka and Yoshino, 2001; Minko et al., 2004; Torchilin, 2010; Kong et al., 2012). It is imperative that, in physiological conditions, very high-binding affinity moiety-receptors are required. In tumor cells for example, receptors are (highly) overexpressed relative to normal cells, and molecular targets are usually employed toward such overexpression of surface receptors for site-specific delivery of therapeutics. Using DDSs exerts precise effects also on cells with low expression (Allen, 2002; Danhier et al., 2010; Lipson et al., 2012; Muro, 2012). Appropriate modifications and functionalization on the pharmaceuticals/ drugs or drug nanocarriers for specific affinity to receptors/markers on targeted cells, tissues, or organs can involve the use of several targeting components. These include peptides, antibodies and their fragments, sugar residues, lectins, vitamins (i.e., folate), transferrin, mRNAs, etc. The targeting moieties will enable the drug/drug carriers to efficiently reach only the intended sites of action and avoid nonspecific accumulations and related side effects (Torchilin, 2000; Ferris et al., 2011; Mahon et al., 2012; Morachis et al., 2012; Wang et al., 2013; De Oliveira et al., 2016). Active targeting could be divided into organ level, cellular level, and subcellular level, depending upon the extent of penetration. For intracellular/subcellular targeting, the active form of the therapeutic substance must reach the intracellular target for its mechanism of action. The targets in question may be located on the plasma membrane or cell components such as endosome, lysosome, endoplasmic reticulum, nucleus, mitochondria, or even mRNA-binding complexes (Harris et al., 2010; Heller et al., 2012; Lammers et al., 2012; Thanki et al., 2015).

1.2.2.1.3 Other approaches’ classification Combination targeting: An active targeting process is the more efficient way to obtain targeting and it provides the widest opportunities. But to yield better results, passive and active targeting approaches are often combined (i.e., PEGcoupled transferrin) (Arima et al., 2012; Krukemeyer et al., 2012; Schleich et al., 2014). The advantages of each approach need to be weighed prior to designing nanocarriers for active or passive targeting (Torchilin, 2010; Hirsja¨rvi et al., 2011; Allen and Cullis, 2013). As an example, EPR effect permits inherent passive accumulation of nanocarriers in tumors, therefore there is a need for active targeting strategy to evade the RES system and/or reach the specific (intra)cellular target (Bae and Park, 2011). Additionally, a combination of active targeting techniques is used to provide additional benefits of targeting, and this is evident from various examples of multifunctional nanocarriers in the literature. This multifunctionality includes modifications of the surface of nanocarriers with the targeting ligands (Du et al., 2012; Lee and Nan, 2012; Xun et al., 2013). Polymer- or liposome-antibody


fragment conjugates (immunopolymers or immunoliposomes) present an example of this combination between polymer- or liposome-based drug delivery and antibody-mediated targeting, for improving stability, solubility, immunocompatibility, pharmacokinetics, etc. (Sapra and Shor, 2013; Srivastava et al., 2014). For increasing targeting efficiency, a combination between nonphysical/ biochemical-based approaches with physical-based approaches is established. For example, ultrasound exposure improves the targeted therapy effects of some galactosylated nanoparticles by increasing vascular permeability on hepatocellular carcinoma (cellular targeting) (Wei et al., 2013). The same method is used for gene therapy with microbubbles for enhancing the delivery of nucleic acid (NA)-containing particles in ultrasound-targeted region (intracellular targeting) (Cool et al., 2013). Physical targeting: Physical targeting is a physical/biophysical stimuli-based approach with topical (local) characteristic, where the therapeutic agents are administered systemically. For drug targeting, we distinguish endogenous physical stimuli such as temperature, pH, redox potentials, etc., and exogenous physical stimuli, which demands the employment of an external driving force (i.e., magnetic, ultrasound) for preferential localization and destabilization of nanocarriers (Torchilin, 2000; Florence, 2012). A variety of (bio-) physical approaches have been widely explored for their immense potential to preferentially localize drug carriers in the targeted areas; biocompatible physical targeting systems have emerged as interesting material in biological application (Watanabe et al., 2013; Chen et al., 2016). As mentioned previously, the microenvironment of pathological tissues may present abnormal pH value and/or temperature, which could be specifically exploited as endogenous stimulus for biophysical targeting. In chemotherapy, physical targeting systems are designed to degrade at (low) acidic pH and/or elevated temperatures (hyperthermia) (Torchilin, 2000; Giustini et al., 2010; Huang and Hainfeld, 2013; Weerakkody et al., 2013). As example of exogenous stimulus, magnetically modulated drug targeting, is a particular system wherein the drug in concern is immobilized on ferromagnetic nanoparticles and allowed to circulate in the patient body (physiological medium). The external magnetic field is applied at the diseased tissue/organ which accumulates the circulating nanoparticles at the desired site of action. The method using ultrasound energy may be similarly established (external ultrasound field guided targeting) (Issa et al., 2013; Hu et al., 2014). Physical targeting approaches may eliminate the multiple chemical processing steps and reduce the number of components required in conventional preparation of drug targeting systems (core formation, bioconjugation with targeting ligands, etc.), thereby improving the chances of reproducibility (biophysicochemical properties) and decreasing difficulties in predicting behavior (mechanism of action) of physical targeting systems in a site of action (Desai, 2012; Florence, 2012; Hu et al., 2014).

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Finally, it should be noted that there are other classifications of approaches, according to targeting-based strategies. We can find a distinction on the basis of: (1) The purpose (subject) of targeting, such as disease-based approaches (cancer, infections), route-based approaches (oral, parenteral, nasal, transdermal, etc.), disease- or location-based approaches (brain, lung, skin) and order- or level-based approaches (organelle, cell, tissue or organ); (2) The mode of target, such as local-targeting, in which a biocompounactive is applied directly into the affected zone; and systemic targeting that requires active or passive targeting mechanism to carry the therapeutic agents to the intended site; and (3) The carrier- or ligand-based targeting (peptide, vitamin, transferrin, antibody, inorganic components, etc.) with different characteristics (size, density, etc.) and interactions (drug, physiological microenvironment, etc.).

1.2.2.2 Drug targeting systems A conventional application of drugs is characterized by the lack of selectivity, poor biodistribution, and limited effectiveness (Nevozhay et al., 2007). In fact, approximately 95% of all new potential therapeutics have poor pharmacokinetic and biopharmaceutical properties, hence there is a need to develop suitable drug systems that distribute, in a very efficient way, the therapeutically active drug molecule only to the site of action, without affecting healthy tissue or organ (Agnihotri et al., 2011). Drugs can be incorporated into nanosystems or nanocarriers, which can carry therapeutic drugs and deliver them to the target site. The incorporation of drug molecules into nanosystems can protect a drug against degradation, as well as offer possibilities of targeting and controlled release. The drug targeting systems involve (nano-)technology, designed to maximize therapeutic efficacy of drugs by controlling their biodistribution profile (Agnihotri et al., 2011; Wilczewska et al., 2012; Yamashita and Hashida, 2013; Gao, 2016). Nanocarriers, as DDSs, are designed to improve the pharmacological and therapeutic properties of conventional drugs (Wilczewska et al., 2012). The major advantages of using such systems are improving the solubility, protecting cargoes from rapid degradation or clearance, and enhancing drug concentration in target tissues; therefore, the required doses of drugs are lower (Nevozhay et al., 2007; Gao, 2016). In comparison with the traditional form of drugs, nanocarrier-drug conjugates are more effective and selective. They can reduce the toxicity and other adverse side effects in normal tissues. The way of conjugating the drugs or other biocompounactives to the nanocarriers is highly important for a targeted therapy. A drug may be entrapped or encapsulated into the nanocarriers, or it can be adsorbed or covalently attached to its surface (Wilson et al., 2010; Wilczewska et al., 2012; Masood, 2015).


Nanocarriers, especially nanoparticles, are interesting because they are easy to synthesis and have numerous applications. In addition, the surface of nanoparticles is fairly easily conjugated with specific ligands or antibodies for recognition and binding to target cells (Faraji and Wipf, 2009; Paulo et al., 2011). There is a need for further improved particulate carrier systems, with as many as possible of the following ideal characteristics and properties (Muller and Keck, 2004; Agnihotri et al., 2011; Hahn et al., 2011; Gujral and Khatri, 2013): (1) easy or reasonably simple to be produced and qualified, at a large scale and cost effective, and acceptable by regulatory authorities; (2) eontoxic, nonimmunogenic (biocompatible) and physicochemically stable in vivo and in vitro (minimal drug leakage during transit); (3) resistant to aggregation, able to cross biophysiological barriers, and have prolonged circulation times in the body (resistant to RES uptake); (4) release the drug moiety inside the target with controllable and predictable rate; (5) have high sensitivity and selectivity for the target, and maintain the specificity of the surface ligands (stability in biofluids); (6) biodegradable or readily eliminated from the body without any problem after drug delivery; and (7) applicable to as many drugs as possible. Various kinds of nanoparticulate platforms for the targeted delivery of drugs have gained increasing attention in the biomedical field (Gao, 2016) and are currently under development using emerging novel nanomaterials (Park, 2014). Some of those pharmaceutical carriers have already made their way into clinic, whereas others are still under preclinical development (Torchilin, 2012). These include drug carrier systems such as polymer nanoparticles, liposomes, dendrimers, micelles, inorganic nanoparticles, nanogels, carbon nanotubes, etc. (Torchilin, 2007; Agnihotri et al., 2011; Srikanth and Kessler, 2012; Gujral and Khatri, 2013; Rani and Paliwal, 2014). Some chosen examples of the main drug targeting systems are presented here.

1.2.2.2.1 Lipid-based nanosystems In the past few years, lipids have been of major interest and importance in field of DDS. Reports on different lipid-based nanosystems have increased immensely, with a major focus on liposomes and solid lipid nanoparticles (SLNs) (Pathak and Thassu, 2009; Mehnert and Ma¨der, 2012; Ramteke et al., 2012; Kawadkar et al., 2013). Liposomes: Liposomes represent an important class of nanocarriers known as vesicular DDSs. They are the first to be investigated as drug carriers and have since received a lot of attention as pharmaceutical carriers of great potential (Torchilin, 2005; Wilczewska et al., 2012). Liposomes are spherical, self-closed structures, with one or several aqueous compartments surrounded by one or many

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concentric bilayered (phospho)lipid membranes (Torchilin, 2005; Faraji and Wipf, 2009; Gujral and Khatri, 2013). These small, artificially designed vesicles are nano/microparticular or colloidal carriers with a size ranging from 20 nm to 10 μm (Wilczewska et al., 2012; Rani and Paliwal, 2014). A drug is incorporated in liposomes through the encapsulation process. Additionally, the interactions of liposomes with cells can be realized by adsorption, fusion, endocytosis, and lipid transfer (Wilczewska et al., 2012; Yamanaka and Yasuda, 2012; Alekseeva et al., 2015). Liposomes are being investigated on a large scale due to their attractive biological properties and favorable characteristics. The structural similarity to human cell membrane components (phospholipids) makes liposomes biocompatible and provides a unique opportunity to deliver pharmaceuticals into cells, or even inside individual cellular compartments, without undesirable side reactions (Torchilin, 2005; Moro et al., 2010). Due to their amphiphilic nature, liposomes can entrap both hydrophilic and hydrophobic pharmaceutical agents. Hydrophilic biocompounactives remain encapsulated in the aqueous interior, and hydrophobic ones may diffuse in the phospholipid membrane. Finally, the high loading capacities provide a flexibility in designing and an ease of modification of liposomes, which makes them an appealing solution for increasing the drug delivery advantages in various applications (Agarwal et al., 2000; Nobs et al., 2004; Faraji and Wipf, 2009; Chen et al., 2010). Different methods have been suggested to prepare liposomal vesicles of different sizes and properties. Depending on the number of bilayers and size, liposomes are classified into (1) multilamellar vesicles, formed by several concentric bilayers (0.55 μm); and (2) uni-lamellar vesicles which consist of a single bilayer, with two different sizes: small (around 100 nm) and large (from 200 to 800 nm) (Torchilin, 2005; Gujral and Khatri, 2013). Other types of liposomes are based on their charge, so are distinguished neutral, cationic and anionic liposomes (Jain et al., 2010; Balazs and Godbey, 2011). Both positively and negatively charged liposomes are extensively researched for their application in gene therapy (Srinivasan and Burgess, 2009; Kapoor and Burgess, 2012a; Kapoor et al., 2012; Zhi et al., 2013; Yang et al., 2016). Variants of liposomes can be designed to adhere to cellular membranes for delivering a drug payload or simply to transfer drugs following endocytosis (Faraji and Wipf, 2009). For that, new varieties of liposomes have been developed, such as “stealth” liposomes (incorporating PEG) (Nag and Awasthi, 2013; Luo et al., 2016), immunoliposomes (attaching antibodies to the surface) (Rothdiener et al., 2010; Paszko and Senge, 2012; Saeed et al., 2016), proteoliposomes (incorporating a crude mitochondrial membrane fraction) (Ciancaglini et al., 2012; Bolean et al., 2015), transferosomes (elastic/deformable liposomes) (Rajan et al., 2011; Ghanbarzadeh and Arami, 2013; Ali et al., 2015), LeciPlex (containing methyl-ammonium bromide and soybean lecithin) (Date et al., 2011; Shah et al., 2015), magnetoliposomes (encapsulating maghemite nanocrystals)


(Sabate´ et al., 2008; Soenen et al., 2009; Monnier et al., 2014), etc. Such formulations may prevent liposomes, and accordingly, have longer circulation times and increased duration of action in targeted tissues/cell compartments (Wilczewska et al., 2012; Gujral and Khatri, 2013). Other liposome formulations are rapidly taken up by macrophages, and this can be exploited either for macrophagespecific delivery of drugs, or for passive drug targeting (Torchilin, 2012; Rani and Paliwal, 2014; Bozzuto and Molinari, 2015; Deng et al., 2016). Liposomes are still being investigated for various novel applications. They are used for delivering peptides and proteins (Reddy and Couvreur, 2011; Storka et al., 2015), and treating various diseases (Wilczewska et al., 2012; Kim, 2016), but gene delivery and cancer therapy are the principal areas of interest for liposomal drugs (Torchilin, 2005; Thanki et al., 2015). In fact, liposomes have been widely studied to deliver the NAs (Khatri et al., 2008; Balazs and Godbey, 2011; Buyens et al., 2012; Kapoor and Burgess, 2012a; Allen and Cullis, 2013; Chen et al., 2013; Vhora et al., 2015). Generally, lipids are associated with NAs either via surface complexation or encapsulation of hydrophilic NA molecules within the aqueous core. Biophysical properties of liposomes can be modulated to achieve high NA entrapment, efficient cellular uptake, and endosomal escape (Kapoor and Burgess, 2012b; Angart et al., 2013). For example, cationic liposomes interact electrostatically with NAs to form lipid-NA complexes known as lipoplexes (Balazs and Godbey, 2011; Cool et al., 2013), which protect NA from nuclease degradation, enhance cellular transfection, and facilitate NA release from the intracellular vesicles before they reach the destructive lysosomal compartments (Brand and Nicholls, 2011). Lately, the focus has been diverted towards anionic liposomes for efficient gene delivery (Srinivasan and Burgess, 2009; Balazs and Godbey, 2011; Kapoor and Burgess, 2012a,b; Aoki et al., 2015). Moreover, liposomes are used as thermo-responsive carriers (Ponce et al., 2006; McDaniel et al., 2013). The usefulness of pH-sensitive liposomes has been well exhibited in a wide variety of applications, especially in nanomedicine (Culver et al., 2014). Liposomal delivery systems, modified with pH-responsive, were developed to deliver NA therapeutics and peptides (Q. Zhang et al., 2013; Aoki et al., 2015; Q. Zhang et al., 2015) and used to explore multiple possibilities to treat cancer. The acidic microenvironment of tumors, due to increased glycolysis, permits pH-triggered drug release from liposomes (Karanth and Murthy, 2007; Ferreira et al., 2013; Liu et al., 2014). Various immunoliposomes were also reported for efficient management of cancer (Rochlitz and Mamot, 2009; Herrmann et al., 2012; Nishikawa et al., 2012; Limasale et al., 2015). It was determined that conventional liposomes present some negative points, such as rapid clearance (rapidly captured by the RES) (Yokoyama, 2005). However, problems associated with lower circulation half-life and stability in blood have been resolved (Hong et al., 2002; Bertrand et al., 2010) by functionalized liposomes containing one or many specific substances (proteins, antigens, PEG, magnetic compounds, etc.). Functionalization of liposomes is a promising

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approach for targeted delivery of therapeutics, and can be achieved via many techniques (ligand-binding, chemical modification, etc.) (Zhang et al., 2009; Nicolas et al., 2013; Naumovska et al., 2014; Vabbilisetty and Sun, 2014). This approach facilitates the targeting and uptake (Kelly et al., 2011; Jain et al., 2013). It was observed that modified liposomes were more efficient in uptake and localization into the targeted tissues, compared to unmodified/conventional liposomes. The circulation times of these particles can be greatly increased simply by hydrophilic surface modification (Faraji and Wipf, 2009; Yoshizawa et al., 2011). Multifunctional liposomes which can conjugate a variety of ligands were also reported to act selectively on particular tissues, especially for targeting cancer cells (Biswas et al., 2011; Perche and Torchilin, 2013; Xiang et al., 2013; Balducci et al., 2014; He et al., 2014; Rengan et al., 2014). For example, a breast cancer study shows that conjugation of a lipophilic cationic ligand to the liposomes induces more apoptosis compared to unconjugated liposomes, and the amount of drug required for an effective therapeutic response by functionalized liposomes was less compared to nonfunctionalized liposomes (Patel et al., 2010b). The same results were obtained for induction of immune response; the coated-liposomes (with lectin or IgA) showed higher immune response in comparison to nonconjugated formulation (Behera et al., 2011; Gupta and Vyas, 2011; Q. Zhang et al., 2016). Solid lipid nanoparticles: Solid lipid nanoparticles are one of the major types of lipid-based nanoparticles that have attracted special interest during recent decades. They are submicron colloidal carriers ranging from 50 nm to 1 μm (Faraji and Wipf, 2009; Asawale et al., 2014; Mahajan et al., 2015). This system possesses a solid biodegradable lipid core matrix (dispersed in water) which is stabilized by a shell of various surfactants/emulsifiers (i.e., soy lecithin or poloxamer) (Nair et al., 2011; Wilczewska et al., 2012). Generally, SLNs are made of a solid hydrophobic core with a monolayer of phospholipids coating. The solid core contains the drug dissolved or dispersed in the solid high-melting fat matrix. The hydrophobic chains of phospholipids are embedded in the fat matrix (Ramteke et al., 2012; Asawale et al., 2014; Nikam et al., 2014). Nair et al. (2011) and Mahajan et al. (2015) have provided an overview of lipids and surfactants used for preparation of SLNs (Table 1.1). Several production methods of SLNs are detailed in the literature (Faraji and Wipf, 2009; Ekambaram et al., 2012; Ramteke et al., 2012; Yadav et al., 2013; Mahajan et al., 2015). SLNs were developed at the beginning of the 1990s to overcome the limitations of other traditional colloidal carriers, like liposomes and polymeric nanoparticles (PNPs), and have been reported as an alternative DDS (Vishvajit et al., 2010; Nair et al., 2011; Ekambaram et al., 2012; Yadav et al., 2013). SNLs are frequently studied for their effective delivery (L. Zhang et al., 2010; Wong et al., 2012; Mahajan et al., 2015). They have the potential to carry lipophilic or hydrophilic drugs, or diagnostics (Ramteke et al., 2012; Nikam et al., 2014). Several studies investigated the potential of SLNs to improve the oral bioavailability of

Table 1.1 Lipids and Surfactants Used for Preparation of SLNs as Referred by Nair et al. (2011) Lipids Triacylglycerols: Tricaprin Trilaurin Trimyristin Tripalmitin Tristearin

Acylglycerols: Glycerol monostearate Glycerol behenate Glycerol palmitostearate

Fatty acids: Stearic acid Palmitic acid Decanoic acid Behenic acid

Waxes: Cetyl palmitate

Cyclic complexes: Cyclodextrin

Hard fat types: Witepsol W 35 Witepsol H 35

Ethylene oxide/propylene oxide Copolymers: Poloxamer 188 Poloxamer 182 Poloxamer 407 Poloxamine 908

Sorbitan ethylene oxide/propylene Oxide copolymers: Polysorbate 20 Polysorbate 60 Polysorbate 80

Alkylaryl polyether alcohol Polymers: Tyloxapol

Bile salts: Sodium cholate Sodium glycocholate Sodium taurocholate Sodium taurodeoxycholate

Alcohols: Ethanol Butanol

Surfactants Phospholipids: Soy lecithin Egg lecithin Phosphatidylcholine

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poorly water soluble drugs (Ekambaram et al., 2012; Thukral et al., 2014), but they can be also used to deliver drugs orally, topically, or via inhalation (Faraji and Wipf, 2009). SLN formulations for various application routes (parenteral, oral, dermal, ocular, pulmonar, rectal) have been developed and thoroughly characterized in vitro and in vivo (Nikam et al., 2014; Sanap, 2014). SLN combine the advantages of different colloidal carriers and simultaneously avoid some of their disadvantages (Nair et al., 2011; Ekambaram et al., 2012). SLNs are attractive for their potential to improve performance of neutraceuticals, pharmaceuticals, and other materials (Asawale et al., 2014; Aditya and Ko, 2015). They offer possibility to develop new prototypes in drug targeting because of their characteristics including an excellent reproducibility and physical stability (due to their relatively rigid core), enhanced bioavailability and biocompatibility, protection of incorporated labile drugs from degradation inside the gut, controlled release along with good tolerability, less toxicity (ease of biodegradation), etc. However, some disadvantages have been observed, such as poor drug loading capacity (limited by the lipid matrix structure and/or the solubility of drug in the lipid), drug expulsion after crystallization, and generally a relatively high water content of the dispersions (70%99.9%) (Nair et al., 2011; Ramteke et al., 2012; Wilczewska et al., 2012; Nikam et al., 2014; Sanap, 2014). Despite these disadvantages, it has been proven that an SLN is more beneficial than a colloidal DDS because of its more controlled and targeted properties (Mahajan et al., 2015; Kalaycioglu and Aydogan, 2016). So, SLNs hold great promise for reaching the goal of controlled and site-specific drug delivery (Asawale et al., 2014). They are at the forefront of the rapidly developing field of nanotechnology with several potential applications in (targeted) drug delivery system (TDDS), clinical medicine and other science (Vishvajit et al., 2010; Ekambaram et al., 2012). SLNs have shown great potential for insulin delivery (Dolatabadi et al., 2015; Ansari et al., 2016); they are used to improve the diffusion of the drugs across the bloodbrain barrier (BBB) (Ramalingam and Ko, 2015; Singh et al., 2015; Kuo and Cheng, 2016) and enhance permeation through skin, especially for skin cancer treatment (Geetha et al., 2015; Kelidari et al., 2015; Akbari et al., 2016). Some SLN formulations were prepared as novel anticancer chemotherapeutics (Peters and Brown, 2015; Talluri et al., 2015) and for treatment of tuberculosis (Gaspar et al., 2016). The next generation of carrier systems based on the solid lipid matrix, nanostructured lipid carriers (NLCs) and lipid drug conjugates (LDCs) were designed as modified SLNs to improve their colloidal stability and drug loading capacity, and decrease drug leakage during shelf life of the product. NLCs comprise both liquid and solid lipids, which leads to a special nanostructure that makes them comparatively more versatile than conventional SLNs. Three structural models of NLCs have been proposed (imperfect, multiple, and amorphous types) (Muller and Keck, 2004; Chen et al., 2010; Wilczewska et al., 2012; Naseri et al., 2015; Beloqui et al., 2016). LDCs are insoluble drug-lipid conjugates prepared by salt


formation or by covalent linking followed by homogenization. They were developed in order to expand the applicability of lipid-based carriers to lipophobic drug molecules (Muller and Keck, 2004; Yadav et al., 2013; Khatak and Dureja, 2015). These lipid nanoparticles have potential applications in the drug delivery field, clinical medicine, cosmetics (especially dermal applications), etc. (Kelidari et al., 2015; Naseri et al., 2015). Both SLNs and NLCs have also been reported to be amenable to functionalization with attachment of targeting moieties (ligands, antibodies, PEG, magnetics, etc.). Modifications are reported either by functionalizing the lipid or the surfactant (employed in stabilization) (Rostami et al., 2014; Campos et al., 2015; Devi et al., 2015; Lo´pez-Garcia and Ganem-Rondero, 2015; Neves et al., 2016). Several studies reported improved drug bioavailability and therapeutic efficacy of functionalized SLNs/NLCs compared to nonfunctionalized nanoparticles (Yu et al., 2010; Alukda et al., 2011; Soni et al., 2014; Luo et al., 2015; Arranja et al., 2016).

1.2.2.2.2 Polymer-based nanosystems Polymeric nanoparticles: PNPs are solid colloidal particles occurring in multiple nanoforms with a diameter ranging from 10 to 100 nm (Vauthier and Bouchemal, 2009; Niemirowicz and Car, 2012; Masood, 2015). According to their structural organization, PNPs are classified as nanocapsule and nanosphere (Fig. 1.2) (Vauthier and Bouchemal, 2009; Kumari et al., 2010). Based on their chemical composition, PNPs can originate from natural polymers (i.e., gelatin, chitosan, albumin, DNA) (Sharma et al., 2006; Saraogi et al., 2011; Charoenphol and Bermudez, 2014; Ghaz-Jahanian et al., 2015), synthetic ones (i.e., poly-caprolactone: PCL; poly-methyl methacrylate: PMMA) (Martıńez et al., 2011; Niemirowicz and Car, 2012), or semisynthetic polymers (Niemirowicz and Car, 2012; Wilczewska et al., 2012). Depending on their in vivo behavior, PNPs may be classified as biodegradable (i.e., poly-L-lactide: PLA; polyglycolide: PGA) (Kumari et al., 2010; Kumar et al., 2013; X. Zhang et al., 2013), or nonbiodegradable (i.e., polyurethane) (Fritzen-Garcia et al., 2009). A few decades ago, PNPs emerged as a promising and viable technology platform for targeted and controlled drug delivery (Cheng et al., 2013). A targeted delivery system based on PNPs as a drug carrier is becoming a system of choice, and presents many pivotal characteristics, which include stability and prolonged circulation in blood, biodegradability, nontoxicity, and biocompatibility (Des Rieux et al., 2006). PNPs are widely investigated for their distinctive size and shape properties for tissue penetration via active and passive targeting, specific cellular/subcellular trafficking pathways, and easy control of cargo release by sophisticated material engineering (Masood, 2015); thus making these carriers useful for the treatment of chronic diseases (Panyam and Labhasetwar, 2003). In cancer, for example, targeted PNPs can be used to deliver chemotherapies to tumor cells with greater efficacy and reduced cytotoxicity on peripheral healthy tissues (Chan et al., 2010).

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FIGURE 1.2 Forms of nanocapsule and nanosphere with adsorbed or entrapped drug as referred by Kumari et al. (2010).

Many studies have reported that hydrophilic and hydrophobic biocompounactives are delivered by using PNPs as carriers (Anand et al., 2010; Nanjwade et al., 2010). Drugs can be encapsulated on the PNP structure during a polymerization step (nanocapsule) or can be immobilized on PNP surfaces after a polymerization reaction (nanosphere) (Qiu and Bae, 2006; Mora-Huertas et al., 2010; Nicolas et al., 2013; Moulton and Wallace, 2014). Also, PNPs are of great interest for drug delivery, since they can be tailored from a range of polymers (Ungaro et al., 2012). A number of different polymers have been investigated for formulating nanoparticles by using several methods (Pal et al., 2011; Rao and Geckeler, 2011; Moulton and Wallace, 2014). However, biodegradable polymers are the main materials frequently used as drug delivery vehicles in nanopharmacology because of their good bioavailability, better encapsulation, control release, complete degradation inside the human body, and less toxic properties with negligible side effects (Kumari et al., 2010; Niemirowicz and Car, 2012).


These polymers consist of ester groups that hydrolyze and cause polymer degradation in aqueous environments, thereby releasing the entrapped material. Drugs may be released by desorption, diffusion, or nanoparticle erosion in the target tissue. The rate of degradation and release is controlled, based on polymer properties (i.e., molecular weight, crystallinity, end groups, etc.) (Torchilin, 2008, 2012). Several polymers have been demonstrated in nanomedicine to improve the performance of the therapeutic agents. The use of biodegradable PNPs for controlled drug delivery has shown significant therapeutic potential (Chan et al., 2010). Synthetic biodegradable polymers, such as PCL, polyacrylamide, and polyacrylate, are well-known materials for the synthesis of nanoparticles (Kalaria et al., 2009), and poly(D,L-lactic-co-glycolic acid) (PLGA) polymer seems to be the most studied for drug delivery application (Jain, 2000; Betancourt et al., 2007). Nevertheless, detailed nanotoxicity studies are needed to ensure the safety of PNPs (Cenni et al., 2008; Kawaguchi et al., 2009; Soenen et al., 2011). Additionally, recent progress in the field of TDDS is based upon the rational design of polymers tailored for particular drugs (Masood, 2015). PNPs can be chemically conjugated and modified to targeting drugs/ligands. They are usually coated with nonionic surfactants in order to reduce immunological and intermolecular interactions (Torchilin, 2008). Moreover, the engineering of multifunctional PNPs (combining several properties in one particle) can enhance the therapeutic efficacy of nanoparticulate drugs, and provide a promising method for treatment and/or diagnosis of many chronic diseases (Torchilin, 2012). For example, sophisticated PNPs have been aggressively pursued for responding to dual/multiinternal and/or external stimuli, such as pH/temperature or redox, temperature/enzyme, temperature/pH/magnetic, pH/redox/magnetic, etc. (Cheng et al., 2013; Zan et al., 2014; Guragain et al., 2015). Dendrimer nanocarriers: Dendrimers, referred to as the “polymers of the 21st century” are a novel class of synthetic materials, based on well-defined cascade macromolecules that are characterized by their three-dimensional and extensively branched globular nanopolymeric architectures, which provide a high degree of surface functionality and versatility (Garg et al., 2011; Priya and Jeyapragash, 2013; Kesharwani et al., 2014). These structural characteristics allow dendrimers to play an important role in a wide range of bionanotechnological and pharmaceutical applications (Singh et al., 2009; Meena et al., 2010; Challa et al., 2011; Patel and Patel, 2013). The dendritic architecture is one of the most popular structures, widely represented in biological systems (bronchioli structure, tree branch, and root system arrangements, etc.). At the nanometer level, molecules possessing dendritic structure include: glycogen, amylopectins, and proteoglycans (Niemirowicz and Car, 2012; Wilczewska et al., 2012).

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In contrast to linear macromolecules, the structure of dendrimers is characterized by three basic components, including (Sekowski et al., 2008; Garg et al., 2011; Dwivedi and Singh, 2014; Selin et al., 2016): (1) Central/inner core: constituting a single atom or a symmetric molecule containing two identical function groups, from which the polymeric branches (dendrons) emanate; (2) Interior layers/shells (arms): composed of repeating units, radically attached to the interior core. The number of monomers corresponds to subsequent generations, and their nature determines the microenvironment of the interior and thus the solubilization ability of the dendrimer; (3) Multivalent surface (nanoscaffolding): formed by surface functional groups (terminal functionality) attached to the outermost interior generations. The terminal functional groups provide the dendrimer features. Several reviews have described the synthesis strategies and types of dendrimers with their different functionalities, properties, and potential applications (Nanjwade et al., 2009). Generally, dendrimers have a symmetrical structure, and are built by gradually adding monomeric or oligomeric units, such that each layer of branching units doubles or triples the number of peripheral groups (Fig. 1.3) (Faraji and Wipf, 2009). Dendrimers are prepared using either a divergent or a convergent method (Dwivedi and Singh, 2014). Therefore, dendrimers can be synthesized starting from the central core and working out toward the periphery (divergent synthesis) or in a top-down approach starting from the outermost residues (convergent synthesis) (Niemirowicz and Car, 2012; Selin et al., 2016). Higher generations can accommodate greater payload and there are more sites available for conjugation (McNelles et al., 2015; Valencia-Gallegos et al., 2015). Recent successes in simplifying and optimizing the synthesis of dendrimers provide a large variety of structures with reduced cost for their production (Garg et al., 2011; Selin et al., 2016). Attractive features of dendrimers distinguish them from the available polymers. They have proved themselves to be very challenging and applicative, as the structure provides unique properties like uniform nanoscopic size, high degree of branching, multivalency, water solubility, monodispersity, stability, well-defined molecular weight, and the availability of multiple functional groups at the periphery and interior cavities (Nanjwade et al., 2009; Ina, 2011; Patel and Patel, 2013). Additionally, terminal functionalities provide a platform for conjugation of the drug and targeting moieties (Kesharwani et al., 2014; Thakur et al., 2015). All these advantages clearly distinguish these structures as unique and optimum carriers in targeted delivery. For example, the size properties allow dendrimers to penetrate through a network of vessels to a target site; a high level of monodispersity and a definite number of surface functional groups makes them possible to connect a specified number of drug molecules to the carrier surface at a stoichiometric ratio (Niemirowicz and Car, 2012; McNelles et al., 2015; Yang, 2016).


FIGURE 1.3 Structure (A) and synthetic methods (B) of dendrimers: convergent (1) and divergent (2) methods (Dwivedi and Singh, 2014).

This unique coreshell design of dendrimers helps in the incorporation of both hydrophilic and hydrophobic moieties (Serrano et al., 2015). The biocompounactives can be easily immobilized on a dendrimer in two basic ways. The first method consists of encapsulating the drug molecule into the interior of the dendrimers (cavities inside the core and interior branches). The second method consists of attaching/conjugating covalently (chemically) or electrostatically adsorbing (physically) the compound onto the dendrimer surface (terminal groups), serving the desired properties of the carrier to the specific needs of the active material and its therapeutic applications (Singh et al., 2009; Challa et al., 2011; Dwivedi and Singh, 2014). Both methods have been used to immobilize many drug substances (Niemirowicz and Car, 2012; Astruc et al., 2015).

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There is no doubt that architectural properties are the main reason for showing significant interest in using dendrimers as carriers in the DDSs (Caminade and Turrin, 2014; Magdalena and Elzbieta, 2016). These properties are dominated by the functional groups on the molecular surface (Garg et al., 2011; Niemirowicz and Car, 2012). The surface of dendrimers provides an excellent platform for their poly-functionalities, which facilitates the immobilization of hydrophobic/hydrophilic drugs and the attachment of specific ligands as pharmacokinetic modulators (PEG, fatty or folic acid, antibodies, etc.). The attached compounds can improve surface activity as well as the biological and physical properties of dendrimers, depending on the intended use (Niemirowicz and Car, 2012; Wilczewska et al., 2012; Caminade and Turrin, 2014; Wu et al., 2015). A number of dendritic polymers designed and developed for use in drug delivery applications have shown a higher efficacy in a large variety of therapeutic fields. For example, poly(amidoamine) (PAMAM) is one of the most commercially available dendritic scaffolds, widely used in biology (Serrano et al., 2015; Oddone et al., 2016; Wan and Alewood, 2016). Some dendrimeric systems, like poly(propylene imine) (PPI), have shown significant cytotoxicity, especially for cationic surfaces (Ziemba et al., 2011, 2012; Kaur et al., 2016). However, PEGylation or derivatization with fatty acid can reduce toxic effects by reducing the overall charge density and minimizing contact between cell surfaces and dendrimers (Meena et al., 2010; Challa et al., 2011; Priya and Jeyapragash, 2013; Thakur et al., 2015; Luong et al., 2016). Several studies have demonstrated the possible applications of dendrimers in various areas of science. The impact of dendrimer applications in the biomedical field shows major potential and holds a promising future for the treatment of several disorders (Ina, 2011; Patel and Patel, 2013; Selin et al., 2016). Scientists are especially enthusiastic about possible use of dendrimers as drug delivery tools (Kesharwani et al., 2014). The use of dendrimers as a DDS certainly promises a reliable, safe, selective, and precise method of drug delivery to enhance biocompatibility and help in achieving increased bioavailability, and sustained, controlled, as well as targeted release of drug (Ina, 2011; Dwivedi and Singh, 2014; Mane et al., 2014). This approach is being used for various treatments (anticancer, HIV prevention, antiinfective diseases, etc.) (Patel et al., 2010c; Ina, 2011; Patel and Patel, 2013; Wu et al., 2015; Selin et al., 2016; Singh et al., 2016). Dendrimers can be also used as promising nanocarriers for other applications, like gene and oligonucleotide delivery, in vivo diagnostics, photodynamic therapy, industrial processes, etc. As research progresses, newer applications of dendrimers will emerge and the future should see an increasing numbers of commercialized dendrimer-based DDSs (Nanjwade et al., 2009; Garg et al., 2011; Magdalena and Elzbieta, 2016; Wang et al., 2016).

1.2.3 PROPERTIES INFLUENCING DRUG TARGETING Despite their potential advantages, only a relatively small number of nanoparticle-based medicines have been approved for clinical use, with numerous


challenges and hurdles at different stages of development. The complexity of DDSs as multicomponent three-dimensional constructs requires careful design and engineering, detailed orthogonal analysis methods, and reproducible scale-up and manufacturing processes to achieve a consistent product with the intended physicochemical characteristics, biological behaviors, and pharmacological profiles. The safety and efficacy of nanomedicines can be influenced by minor variations in multiple parameters and need to be carefully examined in preclinical and clinical studies, particularly in the context of biodistribution, targeting sites, and potential immune toxicities (Desai, 2012). From this, those with new ideas involving multidisciplinary approaches can confer the ability to overcome biological and physicochemical barriers; and must be taken into consideration to control the pharmacokinetics, pharmacodynamics, nonspecific toxicity, immunogenicity, biorecognition, and efficacy of drugs (Agnihotri et al., 2011).

1.2.3.1 Drug properties Depending on their chemical composition and physical properties, two groups of drugs can be differentiated: water soluble drugs (biopharmaceuticals), and poorly water soluble drugs, obtained generally by biotechnological processes (Muller and Keck, 2004). These molecules often possess physicochemical and/or biological characteristics that make their use suboptimal in living tissues (Bertrand and Leroux, 2012; Chang et al., 2015), but they present different problems in delivery such as solubility, instability, absorption, specificity, etc. (Rani and Paliwal, 2014). In this case, a drug needs to be combined with a DDS to make it clinically applicable for the treatment of patients (Muller and Keck, 2004). Moreover, an ideal targeted drug delivery approach would not only increase therapeutic efficacy of drugs, but also decrease the toxicity associated with drugs to allow lower doses of the drug to be used in therapy (Vasir et al., 2005). The efficient and safe delivery of hydrophobic therapeutic compounds remains a serious hurdle for the pharmaceutical industry, because the formulation of many hydrophobic drugs requires toxic solvents and surfactants (e.g., Tween), which often impair drug distribution and are associated with severe side effects (Desai, 2012). Moreover, and as we have seen, a drug may be encapsulated into the nanocarriers or be adsorbed/attached to its surface. The way of conjugating the drugs to the DDS is highly important for targeting (Wilson et al., 2010; Masood, 2015). These data indicate that change in the physicochemical properties of the drugs, due to processing as well as conjugating, changes their release properties at the intended site of action (Bhardwaj and Burgess, 2010). TDDS are designed to improve the pharmacological and therapeutic properties of conventional drugs by protecting them against degradation and controlling their biodistribution profile and release (Agnihotri et al., 2011; Wilczewska et al., 2012; Yamashita and Hashida, 2013; Gao, 2016). For this

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reasons TDDS is preferred over conventional DDSs. However, many TDDS require special formulation technologies to overcome drug-associated problems (Muller and Keck, 2004). Several studies have reported to deliver a variety of hydrophilic/hydrophobic drugs, taking different advantages of DDS. For example, the amphiphilic nature of liposomes allows them to encapsulate hydrophobic drugs in the aqueous interior and hydrophobic ones in the phospholipid membrane (Faraji and Wipf, 2009; Chen et al., 2010). For dendrimers, the surface provides an excellent platform for the immobilization of both hydrophobic and hydrophilic drugs (Caminade and Turrin, 2014; Serrano et al., 2015; Wu et al., 2015). While PNPs structure permits the encapsulation of the drug or its immobilization on the surface (Nicolas et al., 2013; Moulton and Wallace, 2014).

1.2.3.2 Carrier properties The main goal of drug targeting is to localize, target, prolong, and protect a drug interaction with the diseased tissue. A drug-loaded carrier with optimal targeting properties should contain structural features that specifically interact with the designated target (Torchilin, 2000; Petrak, 2005; Mishra et al., 2013). Data accumulated over the past 40 years have revealed a few concepts that are fundamental to DDS. First, DDS efficacy is closely related to the chemical structure of the material. For example, minor chemical modifications to polymer structure can drastically affect material degradation, safety, and targeting. Second, the physical shape and size of DDS matters; this can affect material properties and even interactions with the immune system. Third, DDS actively engage with the body, even when they are not designed to (Tibbitt et al., 2016).

1.2.3.2.1 Physicochemical properties Properties and surface characteristics of the nanoparticle play a key role in compound delivery (Faraji and Wipf, 2009). Systemically administered nanoparticles should remain in circulation for a long time to increase their accumulation in targeted tissues (before being cleared by the RES), and be effectively internalized by the targeted cells. This can be influenced significantly by the physicochemical characteristics of nanoparticles, such as particle size, surface properties, and particle shape (Duan and Li, 2013; Gatoo et al., 2014). The majority of these characteristics are interrelated and are important to understand the biological activity of nanoparticles, including their effects on nanoparticle transport behavior in blood, their uptake and clearance by macrophages, and their consequent biodistribution, as well as their interaction with targeted cells and organs (Duan and Li, 2013; Luyts et al., 2013). Size: Size seems to be the most important property of nanoparticles, not only dictating their pharmacological behaviors, but also determining their hazardous properties (Luyts et al., 2013). It has been established that various biological mechanisms, including endocytosis, cellular uptake, and efficiency of particle


processing in the endocytic pathway, are dependent on the size of the material (Gatoo et al., 2014). The effects of size on targeting ability have been investigated, and it has been shown that this factor affects adsorption, blood circulation half-life, renal clearance, and the ability to deliver bicompounactives to the target site (Gujral and Khatri, 2013; Rahi et al., 2014; Masood, 2015; Yang, 2016). The size of nanostructures should range from 5 to 10 nm; however, the sizes of the most commonly obtained structures range from 100 to 500 nm (Niemirowicz and Car, 2012). A decrease in the size of the materials leads to an exponential increase in surface area relative to volume, thereby making the nanomaterial surface more reactive on itself and to its contiguous milieu, thus the potency of the nanocarrier to cause adverse effects (Gatoo et al., 2014; Singh, 2016). It is generally accepted that nanoparticles ranging from 10 to 100 nm have optimal pharmacokinetic properties for in vivo applications (Wilczewska et al., 2012). New data show that nanoparticles ranging from 5 to 200 nm in diameter can easily escape both renal clearance and RES during blood circulation, which leads to longer blood circulation time and a higher concentration in blood, in comparison with the former (X. Zhang et al., 2016). Because of the opsonic effect, larger particles (.200 nm) are more quickly opsonized and removed from the bloodstream by macrophages of the RES than small particles (Cole et al., 2011; Ho¨rmann and Zimmer, 2016). Smaller particles (,5 nm) are taken up by biological systems more easily than larger molecules, but they are more reactive and more toxic. As a consequence, they are subjects to tissue extravasations and are typically eliminated from the blood circulation by renal clearance (Ai et al., 2011; Shin et al., 2015). A growing number of studies show that smaller nano-sized particles often cause more toxicity than larger particles. However, some conflicting data can be found (Luyts et al., 2013). Finally, even if the carrier targeting is influenced in a size-dependent manner, no one particular nanoparticle size suits all clinical applications, because the specific biological (micro) environment around the disease also influences targeting (Toy et al., 2013; Shin et al., 2015). Shape (morphology/geometry): Shape is another essential property of nanoparticles that plays an important role in various biological processes associated with therapeutic delivery (Duan and Li, 2013). Recently, the design of nanoparticles has gained a lot of attention and resulted in the production of particles with various shapes, such as spheres, rods, tubes, fibers, and disks. More extraordinary geometrical shapes, such as worms, squares, urchins, and ellipsoids are also produced (Luyts et al., 2013; Gatoo et al., 2014; Wu et al., 2016). There are not many conclusive evidences on shape parameter with regard to how it governs cellular uptake. However, with the studies concerning biocompatibility of nanomaterials targeted for medical use, some research describing the active cellular uptake, circulation, and biodistribution of nanoparticles have been performed in vivo (Arnida et al., 2011; Lin et al., 2011; Luyts et al., 2013;

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Rahi et al., 2014). There is a need, however, for advancement in understanding of the interplay between particle shape and living tissue for the development of more efficacious nanomaterial-based targeted delivery systems. Nanomaterial shapes may influence biological activity (Tibbitt et al., 2016). The shape of the nanoformulations has an effect on biodistribution and clearance (X. Zhang et al., 2016). It also influences the speed of internalization and the different pathways used to enter the cells (Doshi and Mitragotri, 2010; Murugan et al., 2015). Additionally, one of the mechanisms related to nanoparticles shape is their ability to cause direct physical damage in cells or tissues (Luyts et al., 2013; Rahi et al., 2014). It has been observed that endocytosis of spherical nanoparticles is easier and faster, compared to rod shaped or fiber like nanoparticles, and, more importantly, spherical nanoparticles are relatively less toxic irrespective of whether they are homogenous or heterogeneous (Gatoo et al., 2014). For example, a study on gold nanoparticles showed that spherical shape has a 375%500% higher cellular uptake in mammals, compared to rod-shaped gold nanoparticles (Rahi et al., 2014). Nevertheless, some studies have shown that, in comparison to their spherical counterparts, nonspherical materials are more disposed to blood flow through capillaries, which causes distinct biological consequences such as longer circulation times (Bertrand and Leroux, 2012; Gatoo et al., 2014). In some cases, the various geometries of the TDDS are associated with differences in sizes and surface properties, which complicate interpretation of the data (Bertrand and Leroux, 2012; Wu et al., 2016; Zhu et al., 2016). Surface characteristics: In addition to size and shape, the particle surface itself and its properties also plays a critical role in biological activity (Shin et al., 2015). These surface properties can significantly influence opsonization, phagocytosis, the circulation in blood, and biodistribution of nanoparticles. They can also affect nanoparticle stability and their interactions with cells or tissues (Duan and Li, 2013). The concept of “nanomaterial surface” includes many aspects (surface area, pore, charge, etc.), but the surface charge is the most important, since it constitutes a major determinant of colloidal behavior, which influences the organism response by changing other nanoparticle characteristics, like shape and size, through aggregate or agglomerate formation (Li et al., 2015). Surface defects and roughness. Some defects on nanoparticle structure may cause surface reactions (i.e., production of reactive oxygen species) which increases the toxicity of nanoparticles in vivo (Rahi et al., 2014). Some nonspecific bindings due to surface roughness also play a role in increasing cellular uptake of nanoparticles, which makes the reaction rate of nanomaterials with cells ineffective, and may create a disorder or transient hole in plasma membrane, leading to cytotoxicity (Mahmoudi and Serpooshan, 2011; Gatoo et al., 2014). Surface area and porosity. In relation to the size and shape, it has been observed that a decrease in size/diameter increases exponentially the surface area of the particle, thereby making the nanomaterial surface more reactive on itself


and to its contiguous milieu, thus its biological activity increases substantially (Duan and Li, 2013; Shin et al., 2015). The total number per unit volume may be important. Smaller particles occupy less volume, such that a larger number of particles can occupy a unit area, resulting in potency to cause adverse effects (Shin et al., 2015). Additionally, it is understandable that the elongated particles are likely to adhere more strongly to the endothelial cell walls via multivalent bonding as a function of particle aspect ratios (Duan and Li, 2013). A larger surface area may cause higher reactivity with nearby particles, resulting in possible harmful effects when used in drug carriers (Shin et al., 2015). Conversely, the surface of nanoparticles is not only related to their size, but depends also on the pores present and the smoothness/roughness of the surface (Luyts et al., 2013). For example, nanoparticles of porous silica are highly biocompatible, and present a lower hemolytic activity, compared to silica nanoparticles without these pores (Luyts et al., 2013; Barisik et al., 2014; Rahi et al., 2014). Even when particles have the same composition, they can have significantly different levels of bioactivity, depending on both particle size and surface reactivity (Shin et al., 2015). Several studies employing diverse classes of nanoparticles indicate that the increased surface area of the nano-scaled particles was the most likely factor contributing to increased pathophysiological toxicity mechanisms, such as oxidative stress, reactive oxygen species (ROS) generation, mitochondrial perturbation, etc. (Holgate, 2010; Duan and Li, 2013; Gatoo et al., 2014). Surface charge. Upon an increase in the medical applications of nanomaterials, considerable attention has been given to the effects of surface charge of nanoparticles on their cellular uptake, translocation to different tissues, and bioactivity (Kim et al., 2010; Wilhelm et al., 2016). Surface charge is a key factor determining behavior after inhalation. It specifically influences the organism response upon exposure to nanoparticles by changing their size and shape through aggregate or agglomerate formation. With high surface energy, the particles will tend more to aggregate (Rahi et al., 2014). Moreover, various aspects of nanomaterials such as selective adsorption and permeability are primarily regulated by surface charge of nanoparticles, as it largely defines their interactions and stability with the biological systems (Verma and Stellacci, 2010; Georgieva et al., 2011; Lesniak et al., 2013; Palmela et al., 2015). Another important perspective is that surface charge also determines the encapsulation efficiency of nanoparticles and their interaction with drug release. Accordingly, optimization of physicochemical properties of nanocarriers is a prime objective while developing a TDDS (Shoyele and Cawthorne, 2006; X. Zhang et al., 2016). Surface charge impacts opsonisation, phagocytosis, and biodistribution of nanoparticles (Schipper et al., 2009). Reports suggest that positively charged

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nanoparticles showed significant cellular uptake and clearance, compared to negative and neutral nanoparticles (Bertrand and Leroux, 2012; Shin et al., 2015; Chen et al., 2016). Neutral surfaces remain more biocompatible (Rahi et al., 2014). In vivo studies on cell cycle show that positively charged nanoparticles may cause DNA damage and activate checkpoints (Liu et al., 2011; Ruenraroengsak and Tetley, 2015). Furthermore, positively and negatively charged nanoparticles induce cell death by apoptosis, whereas the neutral ones induce necrosis (Luyts et al., 2013). Changes in surface charge result in considerable differences in the bioactivity (Shin et al., 2015). Composition and crystalline structure: Targeting is attained by means of chemical design of the different delivery systems (Vasir et al., 2005). One of the main functions of different nanocarriers is to mask all unwanted interactions between the drug and the (micro)environment, until the drug is released from the carrier at the target site (Petrak, 2005). Several studies highlighted that the composition and crystalline structure of nanoparticles influence their bioactivity (Griffitt et al., 2008; Yang et al., 2009; Wang and Wang, 2014). The most frequently employed approach in the design of a DDS is to use water soluble, inert macromolecules as drug carriers, or to attach them chemically or physically to the surface of drug-carrying particles (Petrak, 2005). Furthermore, the application of biodegradable nanosystems for the development of nanomedicines is one of the most successful ideas (Wilczewska et al., 2012). Natural and synthetic biodegradable nanomaterials are the main DDS frequently used in nanopharmacology because of their magnificent bioavailability, improved encapsulation, control release, complete degradation inside the human body, and less toxic properties with negligible side effects (Kalaria et al., 2009; Chan et al., 2010; Kumari et al., 2010; Niemirowicz and Car, 2012), although nanoparticles can change crystal structure after interaction with water or other dispersion medium (Gatoo et al., 2014). For metallic nanoparticles, the composition and inherent chemical properties of materials may lead to different bioactivities. As an example, nanosilver and nanocopper, with their soluble forms, cause more toxicity in vivo compared to titanium dioxide (TiO2) of the same dimensions. Generally, nanoparticulate forms of metals were less toxic than soluble forms (Griffitt et al., 2008; Jiang et al., 2008). Additionally, it was reported that modifying the composition of a nanomaterial may cause variable effect. Fortifying carbon nanoparticles by metallic iron, for example, leads to increased reactivity and oxidative stress (Jiang et al., 2008; Rahi et al., 2014). Nevertheless, detailed nanotoxicity studies are needed for further improved particulate carrier systems (Cenni et al., 2008; Kawaguchi et al., 2009; Soenen et al., 2011).

1.2.3.2.2 Functionalization and surface coating As for physicochemical characteristics, modifications of the surface carrier can also affect the absorption, distribution, metabolism, and excretion (ADME)


process (X. Zhang et al., 2016). Surface coating can affect the carrier characteristics by changing their physicochemical properties and chemical reactivity (Gatoo et al., 2014). Since most nanomaterials are hydrophobic, they are rapidly processed by RES and do not have a stable suspension in aquatic environments. Consequently, they are usually modified by the addition of surfactants or stabilizers. Selecting the appropriate functional groups and specific coating materials should also be considered (Gatoo et al., 2014; Rahi et al., 2014; Ho¨rmann and Zimmer, 2016). Moreover, surface carriers are engineered differently, depending on the chosen targeting strategy. For passive targeting, “stealth” systems are formed by modulating some physicochemical properties (size, shape and surface characteristics) and functionalization employing a variety of natural and synthetic materials (i.e., PEG) for significant prolongation in the circulation half-life (Salmaso and Caliceti, 2013; X. Zhang et al., 2016). Whereas in active targeting, appropriate functionalization on surface nanocarriers can be involved by the use of several targeting components for specific affinity to receptors on targeted tissues (antibodies, lectins, vitamins, etc.). Thus, the functionalized DDS reach efficiently only the desired sites of action and avoid nonspecific accumulations and related side effects (Torchilin, 2000; Nie, 2010; Ferris et al., 2011; Mahon et al., 2012; Morachis et al., 2012; Wang et al., 2013; De Oliveira et al., 2016). In general, surface coating can moderate or eliminate the adverse effects of nanoparticles and enhance their therapeutic properties (Gatoo et al., 2014). But, some studies indicate that chemical properties of stabilizers and functional groups may cause toxicity and immune responses in the body (Coccini et al., 2010; Rahi et al., 2014). For example, PEGylation is often associated with numerous advantages such as stability, high solubility, biocompatibility, simple anchoring process, etc. (Park et al., 2010; Nag and Awasthi, 2013; Ho¨rmann and Zimmer, 2016). Although, adding PEG or its derivatives may also be associated with a variety of disadvantages such as immunogenicity, poor stability and problems of drug leakage due to higher hydrophilicity, and concerns about metabolism safety and renal clearance (Ishida et al., 2007; Wilczewska et al., 2012; Yang et al., 2013).

1.2.3.2.3 Aggregation and concentration One of the ideal characteristics of carrier systems is to be resistant to aggregation (Agnihotri et al., 2011; Gujral and Khatri, 2013). The aggregation states of nanoparticles depend on size, surface charge, concentration, and composition, among others (Gatoo et al., 2014). With high surface energy, for example, the nanoparticles change behavior and have a tendency to aggregate more, which influences their bioactivity in the organism (Rahi et al., 2014; Li et al., 2015). Biological or experimental medium conditions may also affect particle dispersion and agglomeration state of nanoparticles, which, in turn, have an effect on some of their

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physicochemical properties (i.e., size), thereby positively or negatively influencing the bioactivity associated with nanoparticles as already mentioned. Therefore, the same nanoparticles may exhibit different manifestations according to the dissolution mediums (Hou et al., 2013; Shin et al., 2015). Generally, agglomerated nanoparticles have more adverse effects compared to well-dispersed nanoparticles. Moreover, research conducted on the bioactivity of different kinds of nanoparticles in vivo have shown that the increase in the concentration of nanoparticles induces a significant increase in toxicity and other adverse effects (Gong et al., 2014; Li et al., 2015). As examples, the accumulation of carbon nanotubes aggregates for long periods in liver, spleen or lungs induces cytotoxic effects (Yang et al., 2008; Li et al., 2010). Similarly, the evaluated toxicity of carbon fullerene at higher concentration (200 μg/mL) induced a significant increase in malformations in embryonic zebrafish used as a biological model (Usenko et al., 2007; Li et al., 2015).

1.2.3.3 Physiological and anatomical properties (biological barriers) Targeting of drugs offers enormous advantages but is equally challenging. The challenge in drug targeting is the targeting of a drug to a specific site, as well as retaining it for the desired duration to elicit pharmacological action (Vasir et al., 2005). Drug carriers are expected to stay in the blood for long time, accumulate in pathological sites with affected and leaky vasculature (tumors, inflammations, and infarcted areas) via the EPR effect, and facilitate targeted delivery of specific ligand-modified drugs and drug carriers into poorly accessible areas (Torchilin, 2010). To reach the intended site of action with the required therapeutic concentration, however, the drug-loaded carrier must cross many biological barriers (Bertrand and Leroux, 2012). Therefore, the complexity of the biological system presents significant hurdles to the site-specific delivery of therapeutic drugs (Sriraman et al., 2014). Hurdles to DDS design criteria vary with the route of administration. Drugs can be administered in a variety of ways, and their successful delivery requires different design criteria. For example, systemic delivery requires the drug to avoid clearance by the RES, and enter the correct tissue. Oral delivery systems must overcome extreme changes in pH as well as accommodate changes in biomolecule concentrations that vary with food intake. Similarly, skin, nasal, and pulmonary drug delivery requires efficient transport of drugs across the epithelium. DDS for local delivery must avoid damage to the surrounding tissue, and must control release to prevent dose “dumping” (Desai, 2012; Tibbitt et al., 2016). Therefore, a comprehensive understanding of the body characteristics is necessary to predict and prevent all problems related to TDDS, such as inadequate pharmacokinetic/biodistribution profiles and/or unacceptable toxicities, and thus may help nanomaterial conception from the early stages (Vasir et al., 2005; Bertrand and Leroux, 2012).


Better understandings of the biological barriers which a drug or a drug-loaded nanocarrier needs to overcome include the following (Vasir et al., 2005; Bertrand and Leroux, 2012; Desai, 2012; Mu et al., 2014; Ho¨rmann and Zimmer, 2016; Tibbitt et al., 2016; X. Zhang et al., 2016):

Physiological barriers: Several physiological elements/manifestations affect the in vivo biodistribution of the TDDS and may be involved in their elimination. These include the interactions with the systemic circulation components such as lung capillaries, blood cells, plasma proteins and opsonin system; the glomerular filtration by the kidney; the capture by the liver (phagocytosis in Kupffer cells); and the sieving through the spleen. Physiologically, the circulation times of TDDS must be tailored according to the intended therapeutic purposes. Long DDS circulation times can be correlated with positive clinical manifestations such as sustained pharmacological effects. However, prolonged residence of the DDS in the bloodstream can also provoke undesired side effects. The physiological behavior of nanosystems is governed by their physicochemical characteristics. Anatomical barriers: First and foremost implies a single barrier, whereas vasc endo and base member can be seen as two. Another barrier is the high heterogeneity in the body organs and tissues; even in the same diseased tissue. For example, the tumor presents high density in cells and stroma; vasculature is highly heterogeneous in distribution, which means large areas of tumors may be poorly perfused, but some places may be more permeable. Another example that constitutes a major obstacle for treatment of most central nervous system (CNS) and brain disorders is the BBB, which restricts the diffusion of large or hydrophilic molecules into the cerebrospinal. Biochemical challenges: Nanoparticles enter the human body through various pathways, reaching different organs and contacting tissues and cells; interactions of nanoparticles with biological systems may occur at different levels. All of these interactions are based on nanoparticle-biomacromolecule associations, from which challenges are identifying and validating the molecular targets at the desired level/site of action. If the whole tissue constitutes a target, the cross of extracellular matrix and the uniform distribution of drug throughout the tissue are the main barriers. Targeting nanosystems to specific receptors or antigens on the cell surface provides the driving force for diffusion of the system to the specific cells. For subcellular targeting, further barriers need to be crossed to allow internalization of nanosystems into specific cells.

If drug targeting can bring a solution to all these problems, it is because one of the key features of engineered nanoparticles is that they can be tailored to execute specific functions (Torchilin, 2010; Bertrand and Leroux, 2012; Sriraman et al., 2014). For example, long-circulating pharmaceutical carriers is one of the

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key properties of nanoparticulate DDS for both passive and active targeting (Torchilin, 2010). Moreover, passive and active drug targeting can profit the pathophysiological and anatomical opportunities and modified biochemical factors under diseased conditions (internal stimuli, structural changes, overexpressed receptors, etc.) to reach the targeted site of action (X. Zhang et al., 2010; Crielaard et al., 2012; Fleige et al., 2012; Kong et al., 2012; Mura et al., 2013). The combined approach for drug targeting and multifunctional nanocarriers engineering may improve the biophysicochemical properties of nanoparticles to overcome many biological barriers.

1.3 CELLULAR TARGETING: NORMAL CELLS AND ABNORMAL CELLS Different nanoparticulate platforms for targeted delivery of drugs have been studied for possible use in nanomedicine in the past few decades (Ignjatović et al., 2014). Some researchers bring out the benefits of nanosystems, ensuring local and controlled delivery of a drug compared to traditional systems (Gu et al., 2013). There are three different levels of targeting (Chandna et al., 2013): (1) First order targeting or organ targeting: When a DDS liberates the therapeutic molecule only in a particular organ it is described as an organ-targeting process. For example, drugs may be targeted to the liver because its vasculature is normally leaky. This is organ targeting. In this case, the drug is not liberated in other tissues because their vasculature is not leaky; (2) Second order targeting or cellular targeting: When a DDS liberates the therapeutic molecule to a specific cell within an organ or tissue, it is described as second order or cellular targeting process. When we connect an antibody to our delivery system, it specifically distinguishes and connects to a specific antigen on a cell surface. This is second order targeting; (3) Third order or subcellular targeting: When the delivery system can go through particular cells and leave the drug intracellularly, then it is described a third order or subcellular targeting process. When the delivery system carries the gene, it goes through specific cells and leaves the gene intracellularly. This is the third order and most complicated variety of targeting (Kirti Rani, 2014). Cellular targeting is beneficial for therapeutic action under several scenarios. First, the quantity of drug required for the desired effect may be significantly reduced because of its specificity, ultimately leading to a decrease in side effects. Just second, and above all, cellular targeting will overcome the most critical limitation of drug actions, i.e., multidrug resistance, which has been a major hurdle in cancer chemotherapy (Sakhrani and Padh, 2013; Cheng et al., 2006). Compared to the huge numbers of research in the field of nanotechnology, few information is available for cellular or intracellular targeting and generally highlights that the drug targeting to these sites is really challenging. The majority of research work is focused on organ targeting (Shegokar, 2013).

1.4 Organ Targeting

1.4 ORGAN TARGETING 1.4.1 DRUG DELIVERY TO BRAIN The CNS is well protected from exposure to xenobiotics and toxic substances by several barriers. These interfaces between the CNS and the periphery with varied degrees of permeability include the blood-cerebrospinal fluid (CSF) barrier, the blood-spinal cord barrier, the blood-retinal barrier and the BBB (Wong et al., 2012; Loureiro et al., 2014; Gao, 2016; Cipolla, 2009). Drug delivery to the brain plays a significant role in modern drug expansion for the CNS (Pardridge, 1997). The major challenge for new drugs is to cross the BBB; 100% of large-molecule pharmaceutics do not cross the BBB, and more than 98% of small molecules cannot cross this barrier (Hauptman, 2014; Madsen and Hirschberg, 2010; F. Zhang et al., 2015; Pardridge, 2003). Only a small class of drugs—small molecules with high lipid solubility and a low molecular mass (Mr) of ,400500 Daltons (Da)—actually cross the BBB. There are only four groups of CNS disorders that systematically respond to such molecules, and these include affective disorders, chronic pain, epilepsy and migraine headache. Conversely, most CNS disorders, such as Alzheimer’s disease, Parkinson’s disease, Huntington’s disease, Autism, etc., require large-molecule drug therapy (Pardridge, 2005). The development of new drugs for brain disorders is a big challenge, and there is no efficient cure for the majority of brain diseases.

1.4.1.1 Bloodbrain barrier: the most important barrier of drug delivery to the central nervous system The BBB is the most selective barrier in brain targeted delivery; it restricts the influx of most compounds from circulating blood to neural tissue while supplying the brain with the essential nutrients for proper function. It was first discovered by Ehrlichin in 1885, who established that intravenously injected dye could stain most organs, excluding the brain (Alyautdin et al., 2014; Eyal and Unadkatjd, 2009). The BBB is mainly composed of several kinds of cells, including brain capillary endothelial cells (BCECs), tightly connected by tight junctions (Novakova et al., 2014), discontinuous layer of pericytes, astrocytes, and neuronal cells (Bhowmik et al., 2015; Posadas et al., 2016; Nimjee et al., 2011). The BBB carefully confines transport into the brain through both physical (tight junctions) and metabolic barriers (enzymes) (Begley, 2004). Tight junctions play an important role in maintaining the integrity of the BBB and keeping out the entry of molecules from blood to brain (Gao, 2016). There are big differences in structure between the brain capillaries and the nonbrain capillary endothelium. Tight junctions are structures that form a thin and continuous close surrounding of each endothelial and epithelial cell and are intended to rigorously regulate the movement of molecules through the paracellular pathway (Wolburg and Lippoldt, 2002; Neha et al., 2013).

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The brain capillaries are unfenestrated (tight junctions without openings) between the endothelial cells throughout, whereas in nonbrain capillaries, the endothelium epithelial has fenestrations (openings) through which solutes can move easily via passive diffusion.

1.4.1.2 Strategies to overcome the bloodbrain barrier To deliver drugs to the brain and bypass the BBB, three different approaches are currently used: the invasive approach (neurosurgical-based), the pharmacological approach, and the physiological approach. For neurosurgical-based delivery approaches, these include intracerebral implantation, intracerebroventricular (ICV) infusion, and convection enhanced diffusion (CED) (Pardridge, 1997). The pharmacologic-based delivery approaches include the use of lipid carriers or liposomes. They consists of modifying a molecule, which is known to be active against CNS target, by increasing the solubility and stability in plasma to enable it to bypass BBB. The physiological-based delivery approaches take advantage of the normal physiological transport processes. They involve either carrier-mediated transport of nutrients or receptor-mediated transport of peptides (Pardridge, 1997).

1.4.2 DRUG DELIVERY TO LUNG The respiratory system is an organ in direct contact with the environment and it embodies a possible door for the entrance of drugs into the body (Carlotta Marianecci et al., 2011). The respiratory tract is one of the ancient routes used for drugs administration (Chaturvedi and Solanki, 2013). Pulmonary drug delivery has become one of the indispensable aspects for both local and systemic DDSs (Bharti et al., 2013). It can be used as an alternative to oral delivery (Velayutham et al., 2011; Shah et al., 2012; Chaturvedi and Solanki, 2013). Pulmonary drug delivery is emerging as a noninvasive, nonsystemic delivery approach to directly target disorders of lung (such as asthma, chronic obstructive pulmonary diseases (COPD), emphysema, cystic fibrosis, lung cancer, tuberculosis, pulmonary hypertension, and diabetes, etc.) (Carlotta Marianecci et al., 2011; Jawahar and Reddy, 2012; Chaturvedi and Solanki, 2013; Bharti et al., 2013; Paranjpe and Mu¨llerGoymann, 2014), and some nonrespiratory diseases such as type I diabetes (Jawahar and Reddy, 2012). Drugs are inhaled through the lungs and enter the bloodstream through the alveolar epithelium. Inhalation gives the most direct access to drug target. This route can be used to deposit the drug to the target site at high concentration, reducing systemic side effect (Jawahar and Reddy, 2012).

1.4.2.1 Advantages of pulmonary drug delivery A targeted pulmonary delivery system can offer several therapeutic advantages and disadvantages for the treatment of respiratory diseases (Shah et al., 2012; Mansour et al., 2013). The respiratory system has anatomical and physiological advantages: (1) a large absorptive surface area (up to 100 m2) but extremely thin

1.4 Organ Targeting

(0.10.2 mm) absorptive mucosal membrane (for comparison: the columnar intestinal epithelium is B2030 μm) (Da Silva et al., 2013; Chaturvedi and Solanki, 2013); (2) a thin physical barrier of alveolar epithelium (Mansour et al., 2013); (3) hepatic drug degradation can be avoided (Chaturvedi and Solanki, 2013; Mansour et al., 2013); (4) reduction in systemic side effect (Shah et al., 2012); (5) rich blood supply; and (6) express systemic delivery from the alveolar region to the lung (Chaturvedi and Solanki, 2013; Mansour et al., 2013).

1.4.2.2 Disadvantages of pulmonary drug delivery Pulmonary drug delivery can offer some disadvantages: (1) the interval of activity is frequently short-lived, owing to the fast removal of drug from the lungs or because to drug metabolism (Shah et al., 2012); (2) involves frequent dosing (Ozer, 2007; Gonda, 2006); (3) some drugs may produce irritation or toxicity (Velayutham et al., 2011); (4) the lungs have less transporters and channels than liver and intestine (Bur et al., 2009).

1.4.2.3 Anatomy and physiology of the lung The physiological functions of the pulmonary system are: (1) to inhale oxygen for blood which is supplied to all parts of the body; (2) to exhale carbon dioxide endogenous wastes and toxins from the blood passing the pulmonary capillary bed; and (3) to defend against aggressive intruders (Mansour et al., 2013; Velayutham et al., 2011). The lungs are composed of a total of five lobes, the left lung consisting of two lobes and the right lung of three (Paranjpe and Mu¨llerGoymann, 2014). The human respiratory system can be divided in two main parts: (1) the conducting airways (the nasal cavity and associated sinuses, the pharynx, larynx, trachea, bronchi, and bronchioles); and (2) the respiratory region (bronchioles, the alveolar ducts, and the alveolar sacs) Mansour et al., 2013; (Jawahar and Reddy, 2012; Karhaleashish et al., 2012; Chaturvedi and Solanki, 2013).

1.4.2.4 Mechanism involved in deposition of particles in lung The size of particles is vital in the mechanisms of drug deposition in the different regions of the respiratory tract. The drug is deposited in the airways by three main mechanisms, namely impaction (inertial deposition), sedimentation (gravitational deposition), and diffusion (Mortonen and Yang, 1996; Shivanand, 2009; Shah et al., 2012; Chaturvedi and Solanki, 2013; Paranjpe and Mu¨ller-Goymann, 2014). The first two mechanisms are essential for larger drug particles deposition within the airways, whereas the deposition mechanism is used for the smaller particles (Shah et al., 2012; Chaturvedi and Solanki, 2013). Impaction is the main mechanism of deposition that occurs throughout the lung in the upper respiratory passages, and at or near bronchial branching point. Owing to centrifugal force, the particles collide with the respiratory wall

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and are deposited in the oropharynx regions. The air velocity, breathing frequency, and particle size increase the probability of impaction (Velayutham et al., 2011; Shah et al., 2012; Chaturvedi and Solanki, 2013; Paranjpe and Mu¨ller-Goymann, 2014). Sedimentation occurs when the gravitational force acting on a particle is more than the air flow force. Inspired particles, with sufficient mass and sizes 15 μm, will then be deposited slowly at a constant rate due to the low flow of air. Thus, sedimentation is also influenced by residence time in the airway and to particle size, and breathing rate (Velayutham et al., 2011; Shah et al., 2012; Chaturvedi and Solanki, 2013; Paranjpe and Mu¨ller-Goymann, 2014). Diffusion is the main mechanism of deposition for small particles (less than the diameter of 0.5 μm) in the upper respiratory tract, trachea, and larger bronchi (Manca et al., 2012). The diffusion takes place when the collision of small particles with gas molecules applies discrete nonuniform pressures at the particles’ surfaces. Decreasing particle size and flow rate may increase diffusion. More deposition happens in the alveoli region, due to longer residence time and smaller airway (Velayutham et al., 2011; Shah et al., 2012; Chaturvedi and Solanki, 2013; Paranjpe and Mu¨ller-Goymann, 2014).

1.4.2.5 Drug delivery devices Drug delivery devices play an important role in the effectiveness of pulmonary delivery. It is hard to administer a drug through pulmonary route without appropriate drug delivery devices. The three common inhalation devices used for respiratory delivery (Velayutham et al., 2011; Jawahar and Reddy, 2012; Chaturvedi and Solanki, 2013; Bharti et al., 2013) are: (1) metered dose inhalers (MDI) (2) dry powder inhalers (DPI) (3) nebulizers The choice of device depends on the drug, the formulation, the site of action, and the pathophysiology of the lungs (Jawahar and Reddy, 2012).

1.4.2.5.1 Metered dose inhalers An MDI is a complex drug delivery device designed to provide a fine spray of medicament, commonly with a particle size of less than 5 μm. An MDI is generally used for treatment of respiratory diseases such as asthma and COPD; it can be given in the form of suspension or solution. The most important inconvenience of the MDI is that patients must be educated on operating the device. Another problem in MDI is that just 10%20% of the expelled dose reaches the lung (Velayutham et al., 2011; Jawahar and Reddy, 2012; Chaturvedi and Solanki, 2013; Bharti et al., 2013).

1.4 Organ Targeting

1.4.2.5.2 Dry powder inhalers A DPI contains drug in solid form of drug alone or mixed with suitable carriers (e.g., lactose) that is fluidized when the patient inhales. The most important advantages of a DPI include product and formulation stability, ease of usage, and not expensive, compared to MDI (Velayutham et al., 2011; Jawahar and Reddy, 2012; Chaturvedi and Solanki, 2013; Bharti et al., 2013).

1.4.2.5.3 Nebulizers There are two basic types of nebulizer systems, namely jet and ultrasonic. In a jet-type nebulizer, liquid is transformed and sprayed as fine droplets by use of compressed gas. The disadvantages include time consumption, drug wastage, poor delivery efficiency, etc. Only 10% of the dose is actually deposited in the lungs by jet type. In ultrasonic type, ultrasound waves are produced in an ultrasonic nebulizer chamber when electrically excited. The aerosol droplets produced through high frequency vibrations of an air jet nebulizer are generated when compressed air is forced through an orifice; an area of low pressure is created where the air jet exists. Nebulizers are widely used as aerosolized drug solutions or suspensions for drug delivery to the respiratory tract and are mostly useful for the treatment of hospitalized patients (Velayutham et al., 2011; Jawahar and Reddy, 2012; Chaturvedi and Solanki, 2013; Bharti et al., 2013).

1.4.3 DRUG DELIVERY TO EYE Drug delivery to the eye is one of the most challenging tasks for pharmaceutical scientists, due to the eye’s unique anatomy, physiology, and biochemistry. This organ is well protected from exogenous substances and external stress by various barriers. Ocular barriers include: static barriers (e.g., different layers of cornea, sclera, iris/ciliary body via blood-aqueous barrier (BAB) and retina via blood-retinal barriers), dynamic barriers (e.g., tear dilution, choroidal, and conjunctival blood flow, lymphatic clearance), and efflux pumps (Gaudana et al., 2010; Patel et al., 2010a, 2013; Jaleh Barar et al., 2016). Many of these barriers are similar to the BBB that restrict influx of molecules into brain. These ocular barriers are as follows: •

•

•

The blood-retinal barrier (BRB): disconnects the retina and the vitreous from the systemic circulation, prevents free passage of exogenous substances from the choroid into the retina and vitreous (Mitra et al., 2006; Yasukawa et al., 2004); The BAB: controls the transport of molecules from the blood to the inner part of the eye, and maintains the transparency of the eye (Jaleh Barar et al., 2016; Campbell et al., 2010); The tear film: with a pH range of 7.37.7, the first protective layer of the cornea and conjunctiva. The tear film is an important antibacterial/viral solution, due to its composition. It is composed of nutrients, electrolytes,

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•

•

proteins (e.g., lysozyme, secretory immunoglobulin IgA, lactoferrin, lipocalin, and peroxidase), lipids and mucin (Jaleh Barar et al., 2016; Yavuz et al., 2013); The inner limiting membrane (ILM): a barrier that separates and controls the exchange and entry of particles from the vitreous to the retina (Jaleh Barar et al., 2016); and Intact structure of corneal epithelium connected together with desmosomes and forms tight junctions: these tight junctions build a diffusion barrier in the surface of the epithelium and offer resistance to the passage of most drugs due to the presence of layers: hydrophobic epithelium, hydrophilic stroma, and hydrophobic endothelium (Gaudana et al., 2010; Jaleh Barar et al., 2016).

1.4.3.1 Anatomy and physiology of the eye The eye is an isolated, spherical structure, highly complex and specialized organ with a unique anatomy and physiology (Mitra et al., 2006; Patel et al., 2013; Yavuz et al., 2013; Vadlapudi et al., 2015). Anatomically, the structure of the eye can be divided into two segments: the anterior segment, which occupies approximately one-third. It contains the cornea, crystalline lens, iris, ciliary body, and fluid-filled aqueous humor. Whereas the posterior segment, or the back of the eye, contains the sclera (the white of the eye), choroid vessels, retina, macula, optic nerve, and fluid-filled vitreous humor (Gaudana et al., 2009; Vadlapudi et al., 2015; Patel et al., 2013). The eye is well defended with a variety of specialized cellular modifications that produce a variety of barriers that in part isolate the eye from the rest of the body, which can be a challenge for drug delivery (Diebold and Calonge, 2010; Yavuz et al., 2013).

1.4.3.2 Classification of ocular drug delivery systems A large number of ocular dosage forms are accessible for delivery of drugs to the eye. These can be arranged on the basis of their physical forms (Khokhar and Shukla, 2014; Patel et al., 2010a) as follows: • • • •

Liquids: solutions, suspensions, sol to gel systems, sprays Solids: ocular inserts, contact lenses, corneal shield, artificial tear inserts, filter paper strips Semi-solids: ointments, gels Miscellaneous: ocular iontophoresis, vesicular systems, mucoadhesive dosage forms, particulates, ocular penetration.

1.4.3.3 Route of ocular administration The drug can be delivered to the damaged tissues by six modes of administrations: (1) topical administration, (2) systemic administration, (3) intravitreal drug delivery, (4) intravitreal implants, (5) the conjunctival/scleral drug delivery, (6) subconjunctival (Mitra et al., 2006; Gaudana et al., 2010).

1.4 Organ Targeting

(1) Topical drug delivery (nearly 90% as eye drops) is the most commonly preferred noninvasive route of drug administration to treat anterior segment diseases (glaucoma, conjunctivitis, dry eye syndrome and keratitis) due to ease of administration, low cost and patient compliance. After topical administration, a large fraction (about 90%) of the applied dose is lost due to removal mechanisms (Mitra et al., 2006; Vadlapudi et al., 2015). (2) Systemic administration is frequently preferred for the treatment of posterior segment diseases (scleritis, episcleritis, cytomegalovirus retinitis, posterior uveitis) (Vadlapudi et al., 2015). However, a small volume of eye compared to entire body and the exclusion of the organ from systemic circulation makes systemic administration an unusable route of ocular administration (Mitra et al., 2006). (3) Intravitreal drug delivery is principally employed to treat posterior segment diseases (endophthalmitis, uveitis, proliferative virreoretinopathy, and viral retinitis) that involve light sensitive structures like neural retina, retinal pigmented epithelium, and retinal blood vessels. In intravitreal administration, drugs can directly be injected into the vitreous and retina, sustain drug levels, and evade the BRB. However, the intravitreal injection procedure is associated with a high risk of complications (Mitra et al., 2006; Patel et al., 2010a; Gaudana et al., 2010). (4) Intravitreal implants have been used as an alternative to overcome the problem of high intravenous dose-related toxicity. Intravitreal implants employ biodegradable or nonbiodegradable polymer technology (Mitra et al., 2006). They have many advantages compared to traditional delivery systems, including drug delivery being closer to the target tissue in the posterior segment, dose reduction due to localized delivery (and consequential minimization of systemic side effects), and reduced complications, including infections due to low frequency dosing (Kompella et al., 2010). Now, it is possible to inject implants into the vitreous, avoiding the need for their surgical placement. (5) Scleral drug delivery is the most permeable path, where drug is delivered across sclera and fundamental barriers to treat disorders, due to the high permeability of sclera compared to cornea (Unlu and Robinson, 1998). (6) Subconjunctival administration delivers drugs to anterior and posterior segment by avoiding the conjunctival epithelial barrier, which is rate-limiting for the permeation of water soluble drugs. The subconjunctival route is principally employed to treat posterior segment diseases (glaucoma, cytomegalovirus retinitis, age-related macular degeneration and posterior uveitis) (Gaudana et al., 2010).

1.4.4 DRUG DELIVERY TO NEOPLASTIC DISEASE By definition, cancer is a pathophysiological condition in which normal cells transmute into abnormal cells (that have no physiological function) displaying

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properties of uncontrolled and undifferentiated growth, invasion, and sometimes metastasis (Sakhrani and Padh, 2013). Normal cells are incapable to compete with cancerous cells which rapidly consume the nutrient supplied from the bloodstream (Foster et al., 1999). The normal cells will at last be overcrowded by the tumor cells, as a result the tumors expedite the bloodstream leading to irregular and leaky vasculature (Pattni and Torchilin 2015; Chowdhury et al., 2016). Cancer is one of the major leading causes of death worldwide, with no effective treatments. According to the mortality data from World Health Organization (WHO) in 2015, they are expected in the last decade (Travis et al., 2015; Chen et al., 2016). The most common cancer treatments are restricted to surgery, radiation, chemotherapy, and immunotherapy (Zhang et al., 2011; Masood, 2015; Da Silva et al., 2013; Chen et al., 2016; Q. Zhang et al., 2016a; X. Zhang et al., 2016). Even with the considerable advances in these methods, there is still no universal cure for the cancer, due to their side effects. For example, chemotherapy treatment relies on the hypothesis that rapidly proliferating tumor cells are more possible to be destroyed than normal cells by chemotherapeutic agents. But, in reality, these chemotherapeutic agents have nonspecific systemic distribution and inadequate drug concentrations reaching the tumor, leading to cruel and doselimiting side effects such as neutropenia, anemia, hair loss, and damage to liver, kidney, and bone marrow (Iwao Ojima et al., 2012). Compared to regular treatments, drug delivery to neoplastic disease is an efficient approach to improve the delivery of drugs by enhancing drug accumulation, reducing side effects (by limiting drug access to sensitive noncancerous tissue), and improving efficacy and flexibility in treating the diverse cancer types (Nichols, 2013). For anticancer drugs to be effective in cancer treatment, they should primarily be able to bypass various barriers in the body and to enter the target tumor tissues. Secondly, after getting to the tumor tissues, they should selectively kill cancer cells without harming normal cells. Thirdly, they should be delivered to tumor cells and tissues (Warangkana Lohcharoenkal, 2014). Drug delivery to neoplastic disease occurs via two different procedures: passive targeting and active targeting (Baptista, 2009; Peng Mi, 2014; Nie, 2010; Allen, 2004). The passive targeting of tumor cells by small molecules, particularly nanoparticles, is mainly due to the defective structure of angiogenic vessels and improper lymphatic flow, achieving a higher accumulation in tumor sites compared to normal cells (Tiwari, 2012). Small molecules diffuse more easily from normal vessels in the body, which offers small molecules a short plasma half-life. Concerning large particles, they have a long plasma half-life because they are too large to bypass the normal vessel walls, unless they are recognized and cleared by the RES (Rigon et al., 2015) in different organs, so might not have a chance to reach the target site (Matsumura, 2013). The increased accumulation of a drug within the tumor interstitium achieved by nanoparticles can be ten times higher, compared to the drug alone. In spite of the increased accumulation of a drug within the tumor sites, passive targeting raises several concerns

Glossary

about the targeting specificity of this mechanism, founded in the controversial influence of the EPR effect on drug externalization, which encourages a wide distribution all around the tumor (Baptista, 2009). The lack of selectivity and specificity of passive targeting led to the development of active targeting, also named tumor-specific delivery. This strategy of targeting is achieved by the binding of marker components (as antibodies, peptides, etc.) to the nanoparticle’s surface, which facilitates their uptake by a specific diseased or cancerous cell (Baptista, 2009; Nie, 2010). The choice of this procedure depends on its greater abundance in cell surface and its unique expression, and therefore the ability for internalization of the nanoconjugate (Baptista, 2009; Conde, 2013; Gao and Yang, 2013). In spite of the fact that it is considered that active targeting does not have a direct association to the whole nanoparticles accumulated into the tumor, it will influence the uptake of nanoparticles through receptor-mediated internalization pathways and improve the efficiency of antitumor agents that have intracellular targets (Bartlett, 2007; Kirpotin, 2006). Nevertheless, there are two important approaches to active targeting: (1) to target the overexpressed proteins on the surface of the cells, and (2) to target the tumor microenvironment such as low extracellular pH, relative micro-acidosis, mild hyperthermia, etc. In the former case, ligands, such as antibodies, are commonly used to target the cell surface antigens (Allen, 2002). Both approaches are used to achieve significant payload at the target site (Byrne, 2008).

1.5 CONCLUSION In conclusion, targeting increases efficacy, decreases side effects, and reduces systemic drug exposure. The organized structures of nanoparticles allow for incorporation of various targeting moieties to enhance drug delivery to the target sites, reduce off-target organ toxicities, and facilitate cellular uptake of therapeutic agents (Desai, 2012; Lammers et al., 2012). The true challenge of nanomedicines lies in conciliating innovative nanomaterial advancements with a realistic clinical perspective (Bertrand and Leroux, 2012).

GLOSSARY Bioavailability The rate and the degree to which a biocompounactive is absorbed and becomes available at the site of action (target) in a biological system (Guaaˆdaoui et al., 2015). Biocompatibility Ability to integrate with a biological system without eliciting immune response or any negative effects (Wilczewska et al., 2012).

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Biocompounactives Natural or synthetic compounds which have the capability and/or the ability to interact with one or more component(s) of the living tissue by presenting a wide range of probable effects (Guaaˆdaoui et al., 2014a,b). Carrier Engineered vectors (special molecule or system), which retain a biocompounactive via encapsulation and/or via spacer moiety, and transport or deliver it into the preselected site of action (target) (Agnihotri et al., 2011). Drug delivery system (DDS) A system that is capable of releasing carried bioactive agents in a specific location within the body at a specific rate (Guney and Kutlu, 2011). Drug targeting Selective drug delivery to specific physiological sites (targets) where a drug’s pharmacological activities are required. It may be classified into two general methods: active and passive targeting (Yokoyama, 2005). Enhanced permeability and retention (EPR)-effect Phenomenon wherein the drugtargeting systems accumulate into the diseased tissues because of loose fenestrations and/or poorly formed lymphatic drainage (Torchilin, 2011; Stylianopoulos and Jain, 2013). Nanoparticles Structures smaller than 100 nm in at least one dimension (Wilczewska et al., 2012). PEGylation The coating/coverage of a drug delivery system by the covalent grafting, entrapping, or adsorbing of polyethylene glycol (PEG), a water soluble polymer which is nontoxic, nonimmunogenic, nonantigenic, and with no significant interference with drug release (Kouchakzadeh et al., 2010; Ho¨rmann and Zimmer, 2016). Reticulo-endothelial system (RES) The physiological system responsible for the elimination of foreign macromolecules and particles from the body; macrophages of liver, spleen and lymphatic system play a key role in this elimination (Torchilin, 2005). Target Specific organ or a tissue (group of cells) or a cell which, in chronic or acute condition, needs treatment (Agnihotri et al., 2011).

ABBREVIATIONS ADME BAB BBB BCECs BRB CED CNS COPD CSF Da DPI EPR FDA ICV ILM

absorption, distribution, metabolism, and excretion blood-aqueous barrier bloodbrain barrier brain capillary endothelial cells blood-retinal barrier convection enhanced diffusion central nervous system chronic obstructive pulmonary diseases cerebrospinal fluid Daltons dry powder inhalers enhanced permeability and retention U.S. Food and Drug Administration intracerebroventricular inner limiting membrane

References

LDCs LNPs MDI MPS Mr NAs NLCs nm PAMAM PCL PEG PGA PLA PLGA PMMA PNPs PPI RES ROS SLNs (T)DDS WHO

lipid drug conjugates lipid nanoparticles metered dose inhalers mononuclear phagocyte system molecular mass nucleic acids nanostructured lipid carriers nanometer poly(AMidoAMine) poly-caprolactone poly(ethylene glycol) polyglycolic acid poly-l-lactide poly(D,L-lactic-co-glycolic acid) poly-methyl methacrylate polymeric nanoparticles poly(propylene imine) reticulo-endothelial system reactive oxygen species solid lipid nanoparticles (targeted) drug delivery system World Health Organization

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CHAPTER

Cell-penetrating peptides in nanodelivery of nucleic acids and drugs

2

Canan Ozyurt, Ozge Ugurlu and Serap Evran Ege University, Izmir, Turkey

CHAPTER OUTLINE 2.1 Introduction ....................................................................................................... 67 2.1.1 Cationic Cell-Penetrating Peptides .....................................................68 2.1.2 Amphipathic and Hydrophobic Cell-Penetrating Peptides .....................69 2.2 Applications of Cell-Penetrating Peptides in Delivery of Therapeutic Molecules ......................................................................................................... 70 2.2.1 Delivery of Drugs and Proteins ...........................................................70 2.2.2 Delivery of Nucleic Acids ..................................................................76 2.3 Discovery and Design of Novel Cell-Penetrating Peptides ..................................... 83 2.4 Cell-Penetrating Bacterial Effector Proteins as Novel Tools .................................. 89 2.5 Recombinant Production of Cell-Penetrating Peptides for the Delivery Purpose............................................................................................................. 90 2.6 Conclusion and Future Respects ......................................................................... 91 Acknowledgments ..................................................................................................... 92 References ............................................................................................................... 92

2.1 INTRODUCTION Working with cells mostly requires methods to facilitate intracellular transport of nucleic acids and drugs. A number of different approaches have been developed to achieve this goal. Cell-penetrating peptides (CPPs) are short amino acid sequences that can pass through cellular barriers. This property enables CPPs to be employed as components of nanodelivery systems. Since the first identification of natural CPPs, numerous studies have been devoted to understanding the structure and mechanism of CPPs. As reviewed extensively (Zou et al., 2017), the first characterized CPP was HIV-1 transactivating transcriptional activator protein-derived Tat peptide (Green


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CHAPTER 2 Cell-penetrating peptides in nanodelivery

and Loewenstein 1988; Frankel and Pabo, 1988). The discovery of Tat peptide was followed by other CPPs, such as the third helix of the Antennapedia homeodomain (Derossi et al., 1994), transcription factor VP22 of herpes simplex virus type 1 (HSV-1) (Elliott and O’Hare, 1997), and transportan (Pooga et al., 1998). Today, the cell-penetrating property of protein-derived, synthetic or chimeric peptides is a valuable tool in several delivery applications (Tan et al., 2014; Ramsey and Flynn, 2015; Tai and Gao, 2017; Tashima, 2017). There is a continuous interest in therapeutic applications of CPPs, and various types of peptide sequences have been reported. CPPs share some common characteristics and they use similar mechanisms to pass through the cells they interact with. However, there are questions still unanswered regarding the uptake mechanism and interactions of various CPPs. Different methods are employed to understand how CPPs can localize inside the cell. Most of the studies rely on fluorophore labeling of CPPs and then monitoring cellular uptake of CPPs by fluorescence-activated cell sorting (FACS) or fluorescence microscopy (Orosco et al., 2016; Rezgui et al., 2016). Intracellular penetration of CPPs depends on several factors, but uptake mechanisms are essentially classified into two groups: direct uptake and endocytosis (Madani et al., 2011). CPPs can also be classified according to their structural and biological activity properties. Structural classification involves three main families, based on physical and chemical properties. According to this approach, CPPs are categorized in three groups as cationic, amphipathic, and hydrophobic (Milletti, 2012). Different kinds of CPP-mediated nanocarriers can be designed for intracellular delivery of nucleic acids, proteins and drugs (Fig. 2.1). Since the discovery of Tat in 1988, a huge number of CPPs have been reported. In 2012, the CPPsite was established as a comprehensive web server to search and submit peptide sequences (Gautam et al., 2012). The database, with around 1850 peptide entries, was updated in 2016 (Agrawal et al., 2015).

2.1.1 CATIONIC CELL-PENETRATING PEPTIDES One of the first examples of a cationic CPP is 9-amino acid Tat peptide (RKKRRQRRR), which contains eight positively charged amino acids (Tunnemann et al., 2006; Green et al., 1989). Lysine and arginine are the most common residues in both natural and synthetic cationic CPPs. Arginine-rich CPPs are the model sequences, which can pass through eukaryotic cells. Interaction of arginine-rich CPPs with the environment and their uptake mechanisms have been studied in detail (Schmidt et al., 2010). It was described by Milletti, that peptides with a cationic charge of less than eight exhibited a lower cell-penetrating efficiency (Milletti, 2012). It was shown that arginine-rich CPPs could exhibit a protective effect against brain trauma injuries (Chiu et al., 2016). Recently, it was shown that localization of charges were effective on cellular uptake (Nagel et al., 2017).

2.1 Introduction

FIGURE 2.1 Intracellular delivery of nucleic acids, proteins, and drugs can be achieved by CPP-mediated and nanoparticle-based delivery systems.

Recent studies are mostly focused on conjugation of CPPs with various types of nanoparticles. For this reason, some characteristics, such as molecular weight, charge type, and distribution are the parameters that should be optimized to handle interactions between nanoparticles and cationic peptides. For instance, the effect of polycations on polyelectrolyte nanoparticles was discussed in one recent study (Alhakamy et al., 2016).

2.1.2 AMPHIPATHIC AND HYDROPHOBIC CELL-PENETRATING PEPTIDES Amphipathic CPPs generally include two different regions: hydrophobic and charged. Two of the well-characterized amphipathic peptides are MPG (Morris, 1997) and Pep-1 (Morris et al., 2001). The MPG peptide was constructed by combining the fusion sequence of HIV gp41 and the SV40 T antigen nuclear localization sequence. The MPG peptide vector could deliver oligonucleotides into mammalian fibroblast cells (Morris et al., 1997). Amphipathic peptides are generally divided into two main groups as primary and secondary. Chimeric MPG and Pep-1 peptides belong to the primary group (Pooga and Langel, 2015). Secondary amphipathic CPPs generally have an α-helical structure (Milletti, 2012). Based on the classification by Milletti, hydrophobic CPPs are described as the peptides consisting of only apolar residues or the peptides having a low net charge (Milletti, 2012). One exception is the case when a hydrophobic group is

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mainly responsible for cell penetration, regardless of content of the whole sequence. The hydrophobic class is the less abundant group, while the number of studies on cationic and amphipathic CPPs is much higher. In particular, cationic HIV-1 Tat peptide has been extensively studied for its use in therapeutic applications (Rizzuti et al., 2015).

2.2 APPLICATIONS OF CELL-PENETRATING PEPTIDES IN DELIVERY OF THERAPEUTIC MOLECULES 2.2.1 DELIVERY OF DRUGS AND PROTEINS Overcoming biological barriers and specific targeting are the major obstacles of drug delivery (Blanco et al., 2015). CPPs are unique molecules for drug delivery into cells, as they possess cell-penetrating ability. Another requirement is that drug carriers should be structurally compatible with living organisms. CPPs attract much interest, because they meet all of these requirements (Ye et al., 2016); they can be modified in different ways and multi-functional drug delivery systems can be constructed (Zhang et al., 2016a). Different strategies can be used for covalent or noncovalent attachment of CPPs to their cargos (Tashima, 2017). Table 2.1 summarizes some of the recent studies on delivery of proteins/drugs. In contrast to injection-based administration of drugs, oral administration of peptide- or protein-based drugs is often the preferred method. Oral use of drugs enables a convenient way for patients, especially in the case of chronic diseases, which require frequent injections. Moreover, production and storage costs of oral drugs are considerably lower. One limitation of oral drugs is that they must efficiently penetrate mucus layers, such as the intestinal epithelium. CPPs hold great promise for noninjectable drug delivery systems (Kristensen and Nielsen, 2016). In a recent study, gold nanoparticles were modified with Tat peptide, and the effect of modification on doxorubicin (Dox) release profiles in brain metastatic breast cancer cells was investigated (Morshed et al., 2016). In this study, Morshed et al. conjugated the PEGylated gold nanoparticles with doxorubicin and Tat peptide (YGRKKRRQRRR). The resulting Tat-Au-Dox structure was used both in vitro and in vivo experiments. Due to the acidic microenvironment of cancer cells, doxorubicin was attached to gold nanoparticles via an acid-labile hydrazone linker. The pH-sensitive structure was observed to adjust the release and targeting of the drug. Compared to the unconjugated form, Tat-Au-Dox conjugate significantly enhanced the delivery of doxorubicin. Moreover, the IC50 of TatAu-Dox conjugate decreased 80%. In this study, employing Tat peptide played a critical role to penetrating the blood-brain barrier (BBB). Gold nanoparticles were also the subject of a study by Khamehchian et al. The C-terminal truncated form of maurocalcine animal toxin was employed as the CPP agent and it was conjugated to gold nanoparticles via the introduced cysteine

2.2 Applications of Cell-Penetrating Peptides in Delivery

Table 2.1 summarizes some of the recent studies on delivery of proteins/drugs CPP

Nanoconjugate

Target and Aim

Reference

Pep-1

E7/Pep-1 Nanoparticles

Mardani et al. (2016)

Tat

Dextran-coated Iron oxide Nanoparticles (IONP)

Tat

Amphipathic chitosan derivative Gold nanoparticles

Vaccination of tumor cells by delivering a tumor antigen named HPV16 ROS generation on A549 and H358 cell lines by altering magnetic field Improving oral colon absorption of insulin Dox delivery to breast cancer cells Carrying of PTX and 4-HPR drugs through blood-brain barrier and penetrating gliomas Decreasing brain metastasis of breast cancer cells Insulin delivery

Improved cell penetration

Steinbach et al. (2016)

pH-sensitive responses of polymers and doxorubicin delivery Tumor-penetrating peptide delivery via urokinase mediation Ultrasound-stimulated doxorubicin delivery to CD13 1 tumors Improving the use of insulin Improving cell penetration for GNP Targeting autoimmunity

Ng et al. (2016)

Enhancing cellular uptake

Peng et al. (2017)

Tat LMWP

Albumin nanoparticles

iRGD tumorpenetrating peptide

iRGD nanoparticles

R9 peptide

Porous silicon Nanoparticles (PSi NPs) Poly(lactic-coglycolic acid) nanoparticles

Penetratin (AP), end-binding protein 1 (EB1), MPG, and MPGDNLS pArg

SGR, RSGR, or RPARSGR

Triblock Cholic acid based (CA) copolymers Silver nanoparticles

NGR

DSPE-PEG2000based nanobubbles

R8, Tat, Penetratin

PLGA nanoparticles

Tat-C-Cy5 and Tat-C-GNP Myelin peptide (myelin oligodendrocyte glycoprotein, MOG) Tat

Gold nanoparticles Polyplex-based nanoparticles

Poly(ethylene glycol)b-polystyrene stomatocyte nanomotors

Hauser et al. (2016)

Guo et al. (2015) Morshed et al. (2016) Lin et al. (2016a)

Hamilton et al. (2015) Shrestha et al. (2016)

Braun et al. (2016) Lin et al. (2016b) Zhu et al. (2016) Khamehchian et al. (2016) Hess et al. (2017)

(Continued)

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Table 2.1 summarizes some of the recent studies on delivery of proteins/drugs Continued CPP

Nanoconjugate

Target and Aim

Reference

RGD and RGDTat peptides Mgpe9

PAMAM dendrimer core Mgpe9-EGFP-C1 nanocomplexes Dox-CPPs encapsulated magnetic liposomes Ru-CD-RGD nanoparticles Nanoliposomes

Tumor targeting

Ma et al. (2017) Vij et al. (2016) Lin et al. (2016c)

CGRRMKWKK

Adamantaneadded RGDyK GHHNGR homing peptide

Skin penetration Doxorubicin delivery into the MCF-7 cells Tumor targeting and treatment Curcumin delivery into MCF-7 and MDA-MB468 cells for cancer treatment

Xue et al. (2017) Kangarlou et al. (2017)

residue. The same strategy was also applied for Tat peptide. Both nanoconjugates were able to penetrate into different cell types without any severe toxic effects (Khamehchian et al., 2016). Zhang et al. used a gold nanorod-based approach to combine chemotherapy and thermotherapy. In this study, specific targeting was achieved by conjugation with low density lipoprotein receptor (LDLR)-targeted peptide (Zhang et al., 2016b). Quantum dots (QDs) are fluorescent semiconductor nanoparticles, which are excellent tools for imaging, bioconjugation, and delivery purposes (Bilan et al., 2016). CPPs are also used as components of QD-based delivery systems. In one study, Tat peptide was conjugated with a silica QD structure; it was used for controlled release of doxorubicin inside the tumor cells (Li et al., 2014). Breger et al. (2016) studied the effect of branched arginine-rich peptide-conjugated CdSe/ZnS core-shell QD on cellular uptake. They used the method of Medintz et al. (2008) to synthesize a peptide sequence composed of polyarginine attached to a dendritic sequence. They observed that the CCP motif could be used for targeted QD delivery. In a similar study, CdTe QDs were modified with CPPs (Farkhani et al., 2016). It was proven that covalent binding to a cationic CPP enhanced the uptake of QDs. The success of QD and CPP conjugation in these studies is promising and may open the way to design novel QD-CPP-based drug delivery systems. Intrinsic fluorescence of QDs provides another advantage to easily monitor and handle drug molecules. Overcoming the BBB is another challenge in drug development studies. The conjugation of nanomaterials with CPPs has unique advantages to overcome this problem. Low molecular weight protamine (LMWP) fragment can facilitate crossing various biological borders, including skin (Huang et al., 2010), intestinal


mucosa (He et al., 2013; Kang et al., 2014), and tumors (Wang et al., 2014). Lin et al. (2016a) employed LMWP-modified albumin nanoparticles to penetrate the BBB for glioma therapy. In this study, BSA was denatured by a series of chemical reactions. Following the denaturation step, the hydrophobic drugs paclitaxel (PTX) and fenretinide (4-HPR) were loaded to albumin nanoparticles. It was shown that modification of PTX and 4-HPR with LMWP-conjugated albumin nanoparticles increased BBB penetration and cellular intake. Delivery and controlled-release of doxorubicin was achieved by ultrasoundsensitive nanobubbles (NBs) and a penetratin-originated CPP (CKRRMKWKK) (Lin et al., 2016b). First, CPP-Dox conjugate was synthesized, then the conjugate was entrapped in NBs. To overcome nonspecific cell penetration by CPP, NBs were modified with the tumor-selective asparagine-glycine-arginine (NGR) peptide. The NGR motif functioned as a guide to gather a CPP-Dox/NGR-NB cluster in tumor zones. The resulting conjugate was designated as CPP-Dox/NGR-NB. Ultrasound-induced stimulation resulted in release of CPP-Dox from the nanoconjugate, and CPP-enabled penetration of drug molecule through the membrane in a classical manner. In addition to gold nanoparticles and QDs, CPP-modified magnetic nanoparticles are also promising tools for cancer therapy. Hauser et al. (2016) modified iron oxide nanoparticles with Tat peptide and employed CPP-mediated energy delivery into lung cancer cells. The study relied on alternating magnetic field exposure, causing an energy conversion to heat, and rotation with two different processes called Nee´l relaxation and Brownian rotation (Dennis et al., 2009). They studied the effect of magnetic field after penetration of the nanoparticles into cells. The results indicated that Tat peptide facilitated penetration of nanoparticles into A549 and H358 lung cancer cells and provided a lysosomal protective effect. Moreover, CPP-conjugated nanoparticles induced mitochondrial membrane depolarization. The authors discussed that targeted delivery to specific organelles, such as mitochondria, could enhance the therapeutic efficiency of nanoparticles. Amphipathic chitosan and CPP-modified nanoparticles were used to study oral colon absorption of insulin (Guo et al., 2015). The authors modified polyvinyl alcohol (PVA) with Tat peptide and chitosan (Tat-CS-NPs) to enhance cellular uptake and transport of insulin. The conjugates PVA-NPs, CS-NPs and Tat-CS-NPs were tested on Caco-2 cell monolayers. It was shown that Tat-CS-NPs increased the cellular uptake. The colon-targeting property of the system was also confirmed by using single-photon emission computer tomography with 99mTc isotope labeled nanoparticles. Modification of polymer nanoparticles with photo- and pH-sensitive polypeptides was another important contribution to the CPP-based therapeutic methods (Yang et al., 2016a). For instance, a new biodegradable polymeric nanoparticlebased peptide delivery system was developed to overcome telomerase deficiencies (Egusquiaguirre et al., 2015). The authors aimed to enhance the delivery of GSE24.2 peptide, which was proven to increase telomerase activity in earlier

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studies. First, the therapeutic peptide GSE24.2 was produced in Escherichia coli Rosetta-2-gami cells and purified. Then, the purified peptide was modified with biodegradable poly(lactic-co-glycolic acid) (PLGA) nanoparticles. To achieve intracellular delivery, PLGA nanoparticles were modified with either CPP or polycations. Cellular uptake efficiency of CPP and polycations were compared. It was concluded that employing CPP enabled an enhanced delivery. In a similar study, PLGA nanoparticles were modified with CPP, and cellular uptake of surface-modified nanoparticles was investigated (Steinbach et al., 2016). The peptide MPG (SG-GALFLGFLGAAGSTMGAWSQPKKKRKV) enabled the highest uptake efficiency among other peptides (penetratin, end binding protein 1 and MPGΔNLS). In addition, the authors tested two different surface modifications with either avidin-MPG or 1,2-distearoyl-sn-glycero-3phosphoethanolamine (DSPE)-MPG. It was concluded that both surface modifications contributed significantly to internalization. Lee et al. aimed to understand the effect of CPP type on the delivery of green fluorescent protein (GFP). They prepared Tat-conjugated and 9R-conjugated GFP. The two CPPs showed different effects. Tat enabled a high cellular uptake efficiency, whereas 9R-conjugated GFP had better retention ability. Dual modification of GFP with both Tat and 9R resulted in high transfection and long retention properties (Lee et al., 2015). Gao et al. (2014) identified dual-functionalized nanoparticles with angiopep-2 and activatable cell penetrating peptide (ACP). The authors aimed to overcome both BBB and blood-tumor barrier. To achieve this goal, they used two different peptides to increase targeted delivery of the drugs into glioma. ACP was constructed by linking EEEEEEEE(E8) and RRRRRRRR (R8). Due to high expression level of metalloproteinase in the glioma site, ACP was designed to include a linker sequence for matrix metalloproteinase-2 (MMP-2) recognition. MMPsensitive ACP (EEEEEEEE(E8)-6-aminohexanoyl-PLGLAG-RRRRRRRR(R8) was conjugated with nanoparticles. Targeting nanoparticles to glioma was achieved by conjugating angiopep-2 (TFFYGGSRGKRNNFKTEEY). In order to evaluate the targeting ability of angiopep-2 and ACP dual-conjugated nanoparticles, docetaxel (DTX) was chosen as a model chemotherapeutic drug. Dualconjugated and DTX-loaded nanoparticles enabled both specific targeting and enhanced glioma penetration. In another study, triple negative breast cancer (TNBC) cells were targeted by the peptide-conjugated nanoparticle (Hu et al., 2015). Similar to the study by Gao et al., angiopep-2 was used as the targeting agent. Dendrigraft poly-L-lysine (DGL) was used to enhance tumor penetration of the doxorubicin-loaded gelatin nanoparticles (GNP). The final construct Angio-Dox-DGL-GNP remarkably inhibited tumor development. Gautam et al. showed that a novel CPP, IMT-P8, was more efficient than Tat peptide. The most remarkable result of the study was that GFP and the proapoptotic peptide KLA fusion constructs of skin-penetrating IMT-P8 could be


applied topically to mice skin. IMT-P8-KLA and IMT-P8-GFP were shown to be internalized into the hair follicles and dermal tissue (Gautam et al., 2016). Fang et al. designed a Dox-loaded nanoconjugate for use in glioma chemotherapy. For targeting, they conjugated the cyclic RGD peptide to reversibly corecrosslinked biodegradable micelles. The synthesized nanoconjugate was superior in terms of high drug loading capacity, small size, active internalization by U87MG glioma cells, long circulation time, nontoxicity, as well as efficient accumulation and retention (Fang et al., 2017). In the study by Wang et al., liposomes were modified with the pentapeptide, QLPVM, which was derived from Ku70 Bax-binding domain. Then, liposomes were loaded with either the estrogen receptor modulator (TAM), or the drug Dox. The conjugated peptide was found to enhance the cellular uptake and increase cytotoxicity of the drugs. One significant outcome of this study was that the combination of Dox and the endocrine therapy agent TAM was evaluated by using the QLPVM-conjugated liposome delivery system (Wang et al., 2016a). Suzuki et al. described the cellular interactions of liposomes, which had been previously reported by them. Following their previous study, the authors observed that liposomes modified with slightly acidic pH-sensitive peptide could penetrate to deeper regions in tumor tissues. Therefore, the authors proposed that the modified liposomes could be effectively used for drug delivery to tumor core regions (Suzuki et al., 2017). In order to achieve specific targeting and reduce the negative effect of PEG on the uptake, Ding et al. used a different strategy for conjugation of CPP. They designed pH-sensitive PEGylated liposomes and CPP was attached onto the surface of liposomes. The PEG groups were derivatized by the acid-sensitive hydrazone bond. Due to the acidic microenvironment in tumor tissues, PEG could be eliminated and CPP was exposed on the surface of liposome. Intracellular uptake studies showed that the designed nanocarrier showed enhanced blood circulation in vivo, tumor-specific targeting, and, more importantly, high cell-penetrating and endosomal escape properties (Ding et al., 2015). In order to investigate the potential of a novel radioprotective therapeutic agent, Pan et al. produced a fusion protein (GST-Tat-SOD) composed of two antioxidant enzymes and Tat peptide. The authors showed that the fusion protein gained cell-penetrating activity, due to the presence of Tat peptide. Moreover, the bifunctional fusion protein was more efficient than monofunctional protein (SOD-Tat) and the clinically used drug amifostine. Although it was concluded that further studies would be needed, this study provided a useful approach to design CPP-based bifunctional protein fusions (Pan et al., 2016). In another fusion protein-based strategy, Oh et al. employed a hybrid peptide of CPP and collagen (COLL) domains. The authors used this fusion peptide due to its coil-to-helix transition below physiological temperature. The hybrid peptide was shown to be delivered efficiently only when it had the helical form. This

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significant property allowed the authors to design a temperature-controlled delivery system (Oh et al., 2016). As a novel application, Peng et al. recently described construction of poly(ethylene glycol)-b-polystyrene stomatocyte nanomotors and conjugation with Tat peptide. Penetration ability of the Tat-modified nanomotor was confirmed by cellular uptake studies (Peng et al., 2017). Apart from the delivery studies described above, Teramura et al. employed CPPconjugated poly(ethylene glycol)-lipid structure as a cell glue that was able to induce adhesion of floating cells. In addition to CPPs, the authors also tested other cationic short peptides. Although cationic peptides R4, RA7, and GPI showed a similar binding property with CPPs, they were not able to induce cell adhesion. In contrast, in the case of CPPs, a significant cell adhesion was observed. The authors assumed that cationic peptides were so much surrounded by other membrane proteins that they were not available for interaction. One other important factor was the length of PEG chains that were employed. It was observed that CPP-PEG-lipids with 20-kDa or 40kDa PEG were sufficient for firm adhesion without any cytotoxicity. In this study, cells were successfully immobilized onto microfibers. This study contributed to the field, since it provided a novel method for fabricating CPP-based three-dimensional cell immobilization platforms (Teramura et al., 2017).

2.2.2 DELIVERY OF NUCLEIC ACIDS Gene therapy involves the delivery of genetic material into cells or tissues and it has the potential to treat inherited disorders and acquired diseases like Parkinson’s disease and different types of cancers (Dizaj et al., 2014). Gene therapy may be used to replace a defected gene or silence gene expression in order to block production of specific proteins (Alhakamy et al., 2013). For an effective gene therapy, the nucleic acid of interest must be efficiently delivered into target cells and the designed delivery system should not cause any toxic effect. The carrier system must effectively penetrate the cellular membrane and transport nucleic acid cargo to nucleus. Viral vectors (adenoviruses and retroviruses) are commonly used in gene therapy, due to their natural ability to deliver genetic material. There are ongoing clinical trials of viral vectors. However, oncogenesis and immunological problems arisen by viral vectors make it necessary to design novel and safe vectors. For this reason, nonviral vectors are considered as alternative systems for gene delivery (Alhakamy et al., 2013; Dizaj et al., 2014). Functionalized nanoparticles with unique physicochemical properties show promise as nonviral gene delivery systems. However, nanoparticle-based gene delivery vectors suffer from low transfection efficiency, inefficient targeting to the nucleus, and degradation of genetic cargo. Therefore, recent research interest is focused on improving the properties of such nanoparticle-based gene delivery systems. Enhanced transfection efficiency and targeting to nucleus are the most important requirements.


Table 2.2 Recent Studies on CPP-Conjugated Nanostructures for Delivery of Nucleic Acids CPP

Nanoconjugate

Target and Aim

Reference

NF55

NF/pDNA nanoparticles

DNA delivery and potential gene therapy

Tat

Lipid-PEI hybrid nanoparticles PLGA Nanoparticles and pH and photo sensitive peptides DSPE-PEG2000-MAL based nanobubbles

Docetaxel and pDNA delivery siRNA delivery to cancer cells

Freimann et al. (2016) Dong et al. (2016) Yang et al. (2016a,b)

CGRRMKWKK

RKKRRQRRR-Cys

R8 peptide

Nano-sized polyion complexes (PICs)

R16 KMPNWTYRFRMTPRK L5a

LPD- and HLPD-based particles

CKRRMKWKK Oleoyloctaarginine (OA-R8) R8 peptide

CKRRMKWKK

H-rich peptides

CPP/DNA calcium chloride nanoparticles Ultrasound sensitive nanobubbles Lipid nanoparticles Dexamethasone (Dex)loaded PHEA-g-C18Arg8 (PCA) nanoparticles Thermal and Fe3O4 core magnetic dual-responsive liposomes H-rich peptide/DNA nanostructures

Improvement of EGFR siRNA delivery for breast cancer treatment Advanced siRNA delivery and gene silencing in cancer cells Gene delivery

Gene delivery to Human A549 Cells Anti-c-myc siRNA delivery siRNA delivery to HepG2 cancer cells Gene delivery to HEIOC1 cells as a model for inner ear therapy c-Myc siRNA delivery for cancer therapy Gene delivery to NIH3T3 and 293T cells

Jing et al. (2016) Golan et al. (2016) Zhang et al. (2016c) Liu et al. (2016a) Xie et al. (2015) Li et al. (2016) Yoon et al. (2016) Yang et al. (2016a,b) Meng et al. (2017)

CPPs, as well as nuclear-localized peptides and proteins, are essential components for a successful nonviral vector system. Table 2.2 summarizes some of the recent studies on delivery of nucleic acids. CPPs were used in various types of application to increase the delivery efficiency of different cargo molecules, such as peptides, proteins, drugs, nucleic acid, nanomaterials, and viruses into different types of cells (Bolhassani et al., 2017; Farkhani et al., 2014). Nucleic acids cannot pass through cellular membranes because of their negatively charged phosphate backbone. On the contrary, cationic CPPs can interact

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with nucleic acids to mask the negative charge, leading to a more efficient delivery, comparable to commercial transfection agents (Freire et al., 2017). Choosing the right CPP is significant for delivery systems, because the internalization mechanism of each CPP varies, depending on cell type and cargo type. It is also important whether CPP is covalently or electrostatically bound (Bolhassani et al., 2017; Alhakamy et al., 2013). When CPPs are covalently bound to the cargo, it can induce a loss of biological activity in some cases. Nevertheless, covalently attached CPPs were demonstrated to increase the transfection efficiency and biological activity. Conversely, in some cases, noncovalently bound CPPs may result in decomposition of the CPP/cargo complex before transfection (Alhakamy et al., 2013). Chen et al. designed a delivery system to monitor DNase activity in living cells. In this study, it was concluded that CPP preferably bound to singlestranded DNA (ssDNA) rather than double-stranded DNA (dsDNA). This property of CPP was useful to construct ssDNA-assisted CPP-DNA fluorescent probes (Chen et al., 2016).

2.2.2.1 Delivery of plasmid DNA Delivery of DNA is a promising approach to alter a defective or missing gene to cure the progression of a disease (Ibraheem et al., 2014). There are several strategies to enhance the ability and specificity of DNA delivery systems. One common approach is based on employing positively charged arginine-rich peptides, such as Tat, Penatratin and VP-22. These peptides were shown to enhance uptake of nanoparticle-based DNA delivery systems (Liu et al., 2013a,b; Lee et al., 2013; Rudolph et al., 2004; Kato et al., 2016). Liu et al. concluded that charge and size properties of CPP-cargo nanocomplexes had an influence on delivery efficiency (Liu et al., 2013a). The peptide IR9 was employed for the construction of stable IR9/QD and IR9/DNA complexes, which were delivered into human bronchoalveolar carcinoma A549 cells (Liu et al., 2013b). Muroski et al. used AuNPs modified with N-terminal cysteine modified Ku70 peptide. Then, the peptide-modified AuNPs were conjugated with the linearized expression vector (6.6 kbp) encoding brain-derived neurotrophic factor (BDNF)/ mCherry fusion protein. The conjugate was successfully delivered into rat-derived mesenchymal stem cells (MSCs), and 80% of the treated MSCs expressed the protein within 4 days (Muroski et al., 2014). Safe, efficient, and specific delivery of genes is essential for gene therapy. For this reason, active targeting is very important to enhance the ability to deliver DNA (Scha¨tzlein, 2003; Shao et al., 2015; Jiang et al., 2011). For instance, Jiang et al. targeted tumor cells and designed a cationic polymeric nanovector including both folic acid (FA) and CPP octaarginine (R8). In conclusion, they showed that the nanocomplex (FA-PC/R8-PC/pDNA) enhanced the gene transfection efficiency in folate receptor (FR)-positive tumor cells in vitro and in vivo (Jiang et al., 2011). Freimann et al. rationally designed a novel amphipathic α-helical peptide NF55 for in vivo application. Furthermore, the resulting stable and


resistant-protease NF55/DNA nanoparticles were proven to be promising tools for transfection in various mouse tumor models (Freimann et al., 2016). In another strategy, a combination of therapeutic agents via different mechanisms was effectively used for treatment of particular cancer cells (Chu et al., 2015; Gaspar et al., 2015). For example, Dong et al. designed a promising codelivery nanosystem of anticancer drug (docetaxel) and plasmid DNA (pDNA). Cytotoxicity and transfection efficiency parameters were evaluated by using prostate cancer cell lines in vitro and in vivo. The authors covalently conjugated Tat peptide (RKKRRQRRR) with biocompatible lipid-PEI nanocarrier (200 nm) carrying DTX and pDNA. The nanocarrier enhanced the antitumor effect against cancer cells (Dong et al., 2016). Proapoptotic peptide KLA, responsible for regulation of mitochondria-based apoptosis, was conjugated with mitochondrial targeting triphenylphosphonium (TPP). The resulting complex was bound to poly-R8 peptides through a reversible disulfide bond. Due to the presence of poly-R8 peptides, both internalization efficiency and p53 gene (p53 tumor suppressor gene) stability were increased. The delivery system was designed to release proapoptotic peptide and p53 gene so it provided a more effective mechanism. The authors concluded that this strategy was promising to control apoptotic mechanism of cancer cells (Chen et al., 2015). In another study, Zhao et al. aimed to inhibit proliferation of cancer cells. To do this, they designed a KALA (positively charged CPP)-calcium carbonatebased nanoparticle and loaded it with both p53 expression plasmid and doxorubicin hydrochloride (Dox). It was shown that co-delivery of p53 and Dox into HeLa and 293T cells was enhanced as a result of KALA modification (Zhao et al., 2012). The enzyme-sensitive substrate strategy is based on the unique extracellular environment of a tumor. CPP linked to enzyme-sensitive substrate provides advantages for the delivery of the cargo (Shi et al., 2012). By using this approach, Huang et al. designed a construct consisting of a cationic CPP, pH-sensitive masking peptide, and matrix metalloproteinase-2 substrate named dtACPP. Then, dtACPP was conjugated with DGLs, which could interact with plasmid DNA and form a stable nanoconjugate. Low pH and the increased MMP activity of tumor microenvironment triggered the release of masking peptide. In the case of that activation, the exposed CPP was able to mediate the selective internalization of nanoparticles into tumor cells (Huang et al., 2013). Ohta et al. revealed the effect of poly(ethylene glycol) (PEG) on the efficiency of peptide-mediated penetration of single-walled carbon nanotubes (SWCNTs); SWCNT-(KWKG)7 modified with 12 units of PEG chains enhanced the transfection efficiency significantly. This enhancement was contributed to by adjusting the optimum dispersion stability (Ohta et al., 2017). Majidi et al. attempted to mimic viral delivery mechanism and they designed CPP-based chimeric biopolymeric nanostructures with average sizes of 200 nm. In this study, the chimeric structure involved DNA-binding motif, amphipathic CPP (MPG), and nuclear

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localization signal (NLS). The DNA-binding motif of the nanostructure functioned to complex with plasmid DNA, which was successfully transfected into HEK293T cells (Majidi et al., 2016). A CPP with an unnatural amino acid side-chain was designed and synthesized by Oba et al. In this study, they introduced a guanidinylethyl (GEt) amine [Lys (GEt)] side chain to the lysine residue by using chemical synthesis method. Then, they prepared Lys(GEt)-peptide/pDNA complex and compared its transfection efficiency. As control peptides, they also synthesized Lys and Arg oligopeptides, as well as Lys(AEt) with aminoethyl (AEt) group in the side chain amine. The authors showed that Lys(GEt)-peptide was more efficient than arginine-peptide at a low concentration. This result was due to Lys(GEt)-peptide/pDNA complexes that were able to escape from endosomes without any cytotoxicity (Oba et al., 2016). Liu et al. derived a five amino-acid CPP from bovine lactoferricin. This short peptide was named L5a CPP (RRWQW). Plasmid DNA encoding the reporter fluorescent protein was noncovalently conjugated with L5a, and human lung cancer A549 cells were transfected. Moreover, the L5a and L5a/DNA complexes did not show cytotoxic effect, which was a promising result for developing nonimmunogenic gene delivery systems (Liu et al., 2016a). Hu et al. designed a nonviral gene delivery system, which employed the specific tumor-targeting, as well as cellular and nucleus penetration properties of the R18 peptide. They synthesized a degradable polyethylenimine (PEI) derivate by crosslinking low-molecular-weight (LMW) PEI with N-octyl-N-quaternary chitosan (OTMCS). Then, the peptide was covalently attached and the OTMCS-PEIR18 polymer was complexed with plasmid DNA-encoding fluorescent protein. The nanoparticle was found to have a mean size of 100300 nm, which was appropriate for gene delivery studies. Transfection efficiency of OTMCS-PEIR18/DNA was higher compared to OTMCS-PEI-R13, OTMCS-PEI, and PEI 25 kDa. Moreover, the nanoparticle showed controlled degradation, high buffer capabilities and low cytotoxicity properties (Hu et al., 2016). Genome editing by CRISPR-Cas systems attracts much interest and recent studies employed CPPs to deliver the genome-editing machinery. Ramakrishna et al. covalently conjugated recombinant Cas9 protein to CPP. In addition, the positively charged CPP was complexed with the negatively charged guide RNA. Targeted gene disruption was achieved by employing CPP-mediated delivery system (Ramakrishna et al., 2014). For a similar aim, the BBB-crossing RVG peptide was genetically fused to the gene-editing enzyme Cre recombinase. The fusion protein could specifically deliver the enzyme through the BBB and genome editing was achieved in the brain (Zou et al., 2016).

2.2.2.2 Delivery of RNA interference molecules siRNA is known as short interfering RNA or silencing RNA. RNA interference (RNAi) relies on the formation of double-stranded RNA molecules that manipulate the expression of specific genes (Huang et al., 2015; Kanazawa, et al., 2012; Margus et al., 2012).


Jing et al. employed siRNA targeting the epidermal growth factor receptor (siEGFR). siRNA was transfected into TNBC cells by using Tat-conjugated siEGFR NBs in combination with ultrasound-targeted microbubble destruction (UTMD). Transfection efficiency of siEGFR into MDA-MB-231 cells was found to be higher, and silencing of EGFR mRNA was improved. This study presented an example of how NBs carrying genetic material could be loaded with CPPs. The authors also discussed the requirement of more specific targeting to the tissue, as well as long-term stability of NBs and protection of siRNA in blood (Jing et al., 2016). Owing to the positive charge, arginine-rich CPPs can form stable complexes with nucleic acids. This interaction is often used to improve the transfection efficiency of nucleic acids (Kim et al., 2010). Nonaarginine (R9), which consists of nine repeating arginines, is one of these cationic peptides. By using this principle, R9-modified chitosan nanoparticles were prepared. Improved stability of chitosan-siRNA nanoparticles was obtained and transfection efficiency of siRNA in vitro was increased (Park et al., 2013). In another study, Veiseh et al. modified iron oxide nanoparticles with PEG, siRNA, and coated the nanoparticles with each of three different cationic polymers: polyarginine (pArg), polylysine (pLys), and PEI. They evaluated the efficiency of transfection by comparing the gene-silencing levels in brain, breast, and prostate tumor cells. The authors showed that the pArg coated-nanoparticle provided a more efficient way to deliver siRNA into tumor cells. In this study, cell uptake mechanisms of the cationic polymers were different from each other. Transcytosis was the mechanism used by pArg-coated nanovectors, while penetration of pLys and PEI-coated nanovectors was achieved by endocytosis. The results indicated that cell transcytosis-based internalization mechanism could provide a more safe and efficient way for siRNA delivery (Veiseh et al., 2011). Gold nanoparticles were also attempted for use as siRNA delivery vectors. Conde et al. used a functional screening approach to compare the transfection efficiency of several covalent and noncovalent conjugates of siRNA and AuNP. The nanostructures were conjugated with Tat and KKK-Tat peptides. Then, they evaluated the gene silencing effect of multifunctional nanostructures in three biological models (Conde et al., 2012). Chu et al. used the stearylated H16R8 (STRH16R8) peptide and showed that the amphiphilic CPP contributed to stability and significantly assisted release of Bcl-2 siRNA into the target cells. The designed nanocarrier was significant, since it enabled codelivery of siRNA and the anticancer agent Ellipticine (EPT). Therefore, the designed system with dual functions promised as an efficient anticancer nanocarrier (Chu et al., 2015). PepFect (PF) and NickFect (NF) families of CPPs can form nanocomplexes with their cargo through electrostatic and hydrophobic interactions. Therefore, there are different studies on analysis of different types of siRNA-CPP nanoparticles (van Asbeck et al., 2013). For example, Margus et al. investigated the nanocomplexes which were made of pDNA/SCO/siRNA and four transportan-based CPPs belonging to the PepFect and NickFect families. Stearyl-modified peptides, which were prepared by noncovalent interaction strategy provided more

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stable CPP/nucleic acid nanocomplexes. By using transmission electron microscopy (TEM), the authors evaluated the morphology of nucleic acid nanoparticles modified with novel transfection peptides. Furthermore, they concluded that the NeutrAvidingold (NAgold) complex with plasmid DNA was appropriate for monitoring the mechanisms of CPP-mediated nucleic acid (Margus et al., 2016). Wang et al. designed the MS2 bacteriophage virus-like particle (VLP), which displayed Tat peptide on its surface. The protein-based delivery system successfully delivered microRNA-122 (miR-122) to hepatocarcinoma cell lines, albeit with poor targeting. The authors discussed the advantages of the delivery system, such as high biocompatibility and biodegradability, low cost, and simple preparation. They also argued that MS2 VLP could induce an immune response because of its protein-based character (Wang et al., 2016b). As shown in Fig. 2.2, Yang et al. designed thermal and magnetic dualresponsive liposomes, which were conjugated to CPP and siRNA. Due to the presence of Fe3O4 nanoparticles and thermosensitive lipids, the nanoconjugate

FIGURE 2.2 Design of thermal and magnetic dual-responsive liposomes conjugated with siRNA and CPP. This figure was taken from Yang et al. (2016b).

2.3 Discovery and Design of Novel Cell-Penetrating Peptides

could be controlled by an alternating magnetic field to produce heat and trigger the drug release. The authors concluded that CPP (CKRRMKWKK) efficiently delivered siRNA and that CPP-siRNA conjugate could escape from endosomal entrapment. The design strategy was unique, because it successfully combined several functionalities (thermosensitive release, targeted cellular uptake, gene silencing) in the same delivery system (Yang et al., 2016b).

2.3 DISCOVERY AND DESIGN OF NOVEL CELL-PENETRATING PEPTIDES Following the discovery of the first CPP in 1988 (Green et al., 1989), several CPPs with different sensitivities and properties were identified and designed. There are several natural CPPs derived from quite different organisms, from viruses (Elliott, O’Hare 1997) to human (Young Kim et al., 2015). The first synthetic CPPs were made by mimicking motifs commonly found in natural CPPs. The microenvironment of a solid tumor is characterized by low pH. Therefore, Iwasaki et al. aimed to design low-pH selective CPPs. Their design strategy was based on replacing arginine or lysine residues with histidine. Because of its pKa value, histidine was assumed to have a better sensitivity and selectivity to tumor cells at low pH values. To test their hypothesis, the authors used the R8 peptide (RRRRRRRR) as a template and chemically synthesized different lengths of R8-hybrid peptides or polyhistidines. They observed that cellular uptake of these novel CPPs was increased as the chain length was increased. Among them, the cell-penetrating efficiency of the H16 peptide was up to 14.6-fold higher than the R8 peptide. The fluorescent protein cargo was fused to the H16 peptide, which functioned as both a delivery vector and a His-tag for purification process. The authors obtained unexpected results regarding the pH dependency of cellular uptake. The R8-hybrid peptides did not respond to the pH changes and cellular uptake level was same at both pH 6.0 and pH 7.4. However, the H16 peptide showed a significantly large cellular uptake at both pH values (Iwasaki et al., 2015). One of the recent studies showed that the A2 domain of a heat-labeled enterotoxin could play a role in delivery (Liu et al., 2016b). This claim was confirmed by the transport of mCherry, a fluorescent protein, into HCT116 cells. The mCherry-Linker-LTA2 fusion protein was expressed in E. coli BL21 (DE3) cells and homology modeling of the recombinant protein was carried out. Cyclic CPPs also have an importance in medical use. It was discovered that cyclization of arginine-rich peptides enhanced cellular uptake (Qian et al., 2016; Traboulsi et al., 2015). Qian et al. synthesized cFΦR4 analogue peptides and elucidated their effects on cellular uptake. Cellular uptake was monitored by using naphthofluorescein (NF) (Qian et al., 2016). Menegatti et al.

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developed peptides to facilitate the delivery of peptide drugs through skin (Menegatti et al., 2016). For this purpose, they used computational methods to generate and scan libraries of disulfide-cyclic peptides against keratin and cyclosporine A (CsA). In this study, pentameric, hexameric, heptameric, and octameric cyclic peptides were scanned for their keratin-binding capacities. The lowest KD value (9.34 3 1026) was exhibited by an octapeptide with the sequence of ACNAHQARSTCG. Pescina et al. derived novel peptides from the well-known PEP-1 peptide. The six PEP-1-derived peptide sequences were analyzed. The corneal-penetrating ability of each 50 -FAM-labeled peptide sequence was monitored by confocal microscopy (Pescina et al., 2016). A library consisting of overlapping sequences with PEP-1 was designed. Only one of the designed peptides (PEP-3) could not be obtained at high production scale. Glycine linkers were used to avoid possible effects of FAM on peptides. The permeability coefficients of five different peptides (PEP-2; PEP-4; PEP-5; PEP-6; and PEP-7) were similar with PEP-1 peptide. Neundorf et al. derived a novel peptide (sC18) from the antimicrobial peptide CAP18 (Neundorf et al., 2009). This chimeric peptide enhanced cellular uptake, without significant cytotoxicity on different cell lines, including MCF-7, HeLa, and HT-29 cells. The same study also investigated the effect of a different form of previously reported E5L-peptide on uptake and toxicity. The derivative of E5L-peptide was referred to as N-E5L. It was identified that this peptide couldn’t pass through cells by itself. Cellular uptake and other parameters were confirmed by using fluorescence microscopy and flow cytometry. The peptide sC18 generated by Neundorf et al. was further subjected to other CPP design studies. In other studies, the sC18 peptide was subjected to a dimerization process (Hoyer et al., 2012) and imidazolium salt conjugation (Gronewold et al., 2017). Dimeric form of the peptide was named (sC18)2. Effects of the dimeric peptide on HEK-293, HT-29, and MCF-7 cells were investigated by resazurin-based cell viability assay. Dimerization of the sC18 resulted in a cytotoxic effect on MCF-7 and HT-29 cells. Horn et al. performed a study to change some properties of the cyclized sC18 peptide that was designed by rational approach (Agarwal et al., 2015; Gronewold et al., 2017; Horn et al., 2016). The GLRKRLRKFRNK peptide sequence was used as the linear structure. The peptide was converted to a cyclic structure by using copper (I)-mediated alkyne-azide click (CuAAC) reaction. Glycine and lysine residues were substituted with propargylglycine and Ɛ-azidolysine, respectively. Cyclization resulted in three different peptides (named Cyc1, Cyc2, and Cyc3). Cytotoxicity studies on different human cell lines with triazolebridged cyclic peptides indicated that none of them exhibited cytotoxic effect up to a concentration of 200 μM. Furthermore, cyclic peptides were much more preventive against proteolytic attacks than the linear counterpart. Cyc3 was found to be promising as a good DNA carrier. This effect was contributed to burial positioning of Cyc3 in the membrane, as well as more comfortable interactions of the peptide with DNA.


A peptide belonging to the inner mitochondrial membrane-targeting SzetoSchiller peptide class was derived and named SS-31 by Cerrato et al. Mitochondrial uptake and accumulation profiles were monitored by using fluorescence microscopy after a labeling step with 5(6)-carboxyfluorescein (FAM) (Cerrato et al., 2015). The original sequence of SS-31 peptide was identified as D-Arg-Dmt-Lys-Phe-NH2. The authors synthesized different peptides by using solid-phase chemistry. The net charge of the nine synthesized peptides was (13) and it was the same with the original peptide. However, the net charge of one peptide (mtCPP-6) was found to be (14) at pH 7. Labeling with FAM changed the net charge of nine peptides to 12. The net charge of mtCPP-6 was shifted to 13. The positively charged L-arginine, D-arginine, L-ornithine, and D-ornithine were chosen within the framework of the design strategy. In order to protect peptides against degradation by proteases, D-amino acid derivatives were introduced in the first and third positions of CPPs. As mentioned by the authors, a further protection against hydrolysis was obtained by adding an amide group to the end of CPPs. Furthermore, the authors used phenylalanine and tryptophan amino acids to introduce lipophilic property. Both cell-based studies and the experiments on isolated mitochondria showed that uptake capability of the mtCPP-1 was better than the SS-31 peptide. Moreover, none of the peptides showed cytotoxic effect up to a concentration of 100 μM. The protective effect of each peptide on reactive oxygen species (ROS) formation was also investigated. The findings showed that mtCPP-1, -8, and -9 peptides exhibited better protection effect against ROS formation than the original SS-31 peptide. This design strategy, in which natural and unnatural amino acids were used, presented a high potential for mitochondrial targeting and further therapeutic applications of CPPs. Branching of the known peptides was used as another design approach to developing novel CPPs. Tat peptide and some regions of the sequence were exposed to different branching processes to enhance some properties, such as uptake, interaction with certain materials, and stability. One of the striking examples of these designs was the effort to convert Tat peptide to a nickel complex transporter (Szyrwiel et al., 2015). For this purpose, Tat sequence was branched between the 47 and 57 positions, and the peptide was elongated by two additional histidine (His) residues. The resulting branched structure was able to penetrate into human fibrosarcoma cells. In a similar study, Tat peptide was branched and dimerized to enhance the ability to enter into the cells (Monreal et al., 2015). The process was carried out by using microwave-assisted solid phase peptide synthesis. Flow cytometric results showed that the branched dimerized form Tat-D was the most permeable, compared with other peptides in the range of 0.255 μM. Moreover, any cytotoxic effect was not observed up to a concentration of 50 μM in HeLa cells. One recent study described new branched peptides, which were derived from Tat peptide after DMSO oxidation (Jeong et al., 2016). Gene transfection efficiencies of the newly synthesized peptides were confirmed by using flow cytometry and confocal laser scanning microscopy (CLSM).

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For this purpose, the plasmid encoding green fluorescent protein (pGFP) was used as the reporter gene. All peptides described in this study were conjugated with pGFP. Disulfide bonds were constructed between mTat peptides by using cysteine (Cys) amino acid. Disulfide bonds were designed in both horizontal and vertical directions. Transfection efficiency of the branched peptide was found nearly 40 times better, compared with original Tat and mTat peptides. It was concluded that these kind of branched CPPs could have a great potential for gene delivery studies. As a novel CPP class, cell-penetrating poly(disulfide)s (CPDs) can be considered as synthetic mimics of poly-arginine. Instead of the polypeptide backbone, CPDs are based on poly(disulfide)s. Fu et al. designed and synthesized novel CPDs, which were attached to fluorescently labeled recombinant avidin, BSA, and the epigenetic reader BRD-4 protein. Cellular delivery efficiency of CPD was found to be higher than that of a commercial liposome-based protein delivery system. This novel CPD-mediated protein delivery system was also proved to be successful for caspase-3, antibody, and doxorubicin (Fu et al., 2015). The study by Najjar et al. revealed the importance of stereochemistry. This study investigated the effect of L- to D-amino acid conversion on stability, endosomal escape, and uptake properties (Najjar et al., 2017). The authors used the dimeric fluorescent form of Tat (dfTat), composed of L-amino acids, which showed an efficient endosomal escape. dfTat was compared to D-dfTat, which was chemically synthesized. It was proven that L- to D-conversion contributed positively to protease resistance and stability of the D-peptide. However, the D-peptide showed a reduced endocytosis. In contrast to this negative effect, the D-peptide, once in the endosome, showed a significant ability for endosomal escape. Najjar et al. provided very essential data useful for future design of CPPs with improved properties. As reviewed by Wu et al., phage display technology is useful to identify target-specific peptides (Wu et al., 2016). In this technology, a library of billions of different peptide sequences is displayed on the surface of phages. Following incubation of phage-displayed library with the target and washing steps, the bound phages are eluted. Then, E. coli cells are infected by the eluted phages and the peptide sequences are revealed by DNA sequencing (Fig. 2.3). Phage display technology enabled identification of specific peptides for prostate-specific membrane antigen (PSMA). The peptide GTI (GTIQPYPFSWGY) showed high affinity and specificity for the prostate cancer cell lines LNCaP and C4-2. Moreover, fusion of the proapoptotic peptide to GTI enabled specific delivery, which enabled cell death (Jin et al., 2016). A modified form of phage display is mirror image phage display, which enables the identification of D-peptides (Funke and Willbold, 2009). Mirror image display technology could help to identify novel CPPs composed of D-peptide structures. In their study, Jiang et al. employed D-peptides that had been previously obtained by mirror-image phage display. The arginine-rich peptide D3


FIGURE 2.3 Schematic view of phage display used to identify novel CPPs.

showed similar properties to the arginine-rich motif of Tat. It was shown that D3 penetrated the BBB and specifically bound to Aβ plaques (Jiang et al., 2016). As an alternative to phage display, plasmid display can be used for a better differentiation of the peptides uptaken by cells (Gao et al., 2010). Plasmid display technology relies on formation of a noncovalent linkage between randomized DNA sequence and displayed fluorescent protein. First, a gene library of randomized CPP sequences is formed. Then, this plasmid library is transformed into E. coli cells and protein-DNA complexes are purified. In order to screen the library, target cells are transfected by protein-DNA complexes. Expression of fluorescent transgene indicates that the plasmid DNA has been uptaken by the cells. In other words, it means that the peptide sequence coded by this particular plasmid shows CPP activity. By employing plasmid display technology, Gao et al. (2011) obtained a peptide named SG3, which was not more efficient than Tat, but interestingly, using SG3 in combination with the transfection reagent, Lipofectamine2000, resulted in a higher transfection efficiency than Tat under the same conditions. The aim of the study by Schmidt et al. was to design novel and less-cationic CPPs to overcome the side-effects, such as toxicity and accumulation. First, the authors classified the randomly selected peptides according to polarity and

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hydrophobicity properties. Then, they designed peptide sequences with a maximum of 13 charge. The designed peptides were screened for cellular uptake and subcellular distribution. The authors enriched short and hydrophobic peptides after this first screening. As a second phase of their design approach, Schmidt et al. introduced D-amino acids and lactam bridges to the enriched peptide sequences. In conclusion, they compared the peptides based on their ability to deliver the reporter protein GFP (Schmidt et al., 2017). Begum et al. investigated the optimum peptide sequence required to target the gastrin releasing peptide receptor (GRPR). First, they compared the expression level of the receptor in various cancer cell lines, as well as noncancerous ones. They found that the PC-3 (prostate cancer) cell line had the highest level of GRPR expression, while Caco-2 (colon cancer) cell line showed the lowest level. By using this information, they compared the cellular uptake levels of full-length (14 amino acid) Bombesin (BBN) or the truncated BBN (614) sequences, which were reported to bind to GRPR with high affinity and specificity. The authors showed that cellular uptake of the BBN peptides positively correlated with GRPR expression levels. Since both the full-length peptide and the truncated version showed similar cellular internalization, the authors concluded that the truncated BBN (614) peptide would be sufficient and more appropriate to use in further drug delivery studies (Begum et al., 2016). In addition to CPPs, novel properties of known proteins have been discovered. For instance, SERPINA5 is a member of the alpha-1-antitrypsin clade (clade A) of serpins. In addition to its well-known function, Yang and Geiger described a novel cell-internalization property of SERPINA5. It was shown that it could be translocated to the nucleus (Yang and Geiger, 2017). Ryu et al. (2016) engineered Bombyx mori 30Kc19 protein to function as a fusion partner for intracellular delivery of transcription factors. A cysteine-rich CPP (CyLoP-1) was derived from nuclear localization sequence of snake toxin (crotamine). CyLoP-1 could be used a cargo-delivery vector, both in mammalian cells and plant cells (Ponnappan et al., 2017). Similarly, a novel CPP was derived from the heparin-binding domain of epidermal growth factor-like growth factor and it was named HBP. HBP was shown to have CPP-like properties. It functioned as an effective carrier and cell-penetrating agent when it was conjugated to the C-terminus of MAP30 protein. It was concluded that this novel peptide had promise for use in drug delivery to cancer cells (Luo et al., 2016). Sciani et al. attempted to identify proline-rich CPPs from Bothrops jararaca venom. The venom was chosen due to its high peptide content. In this study, the authors also described an interesting approach for screening CPPs, which eliminated the labeling step. Their label-free approach relied on mixing peptides and liposomes first and then identification of the peptides that penetrated into liposomes. Identification of the BPP13a peptide inside the liposomes confirmed that this novel approach was successful as a novel screening method for CPPs (Sciani et al., 2017).

2.4 Cell-Penetrating Bacterial Effector Proteins as Novel Tools

2.4 CELL-PENETRATING BACTERIAL EFFECTOR PROTEINS AS NOVEL TOOLS Bacterial effector proteins, which play an important role in penetrating the host cells, display a great potential as cell-penetrating tools. The transport potential of Type III bacterial effector proteins was emphasized in a previous study by Galań et al. (2014). EHEC, EPEC, Legionella, Shigella, Salmonella and Yersinia species are the members of gram-negative bacteria expressing T3SS-dependent effector proteins (Ru¨ter and Schmidt, 2017). YopM protein from Yersinia and SspH1, SspH2, Slrp proteins from Salmonella, as well as IpaH from Shigella belong to the family of LPX effector proteins. In contrast to YopM, these proteins have a conserved domain encoding a novel class of ubiquitin E3 ligase (NEL), located at the C-terminal (Ho¨fling et al., 2015; Ru¨ter and Schmidt, 2017). In a previous study, it was demonstrated that recombinant SspH1 effector protein of Salmonella typhimurium was able to pass through the membrane of eukaryotic cells without additional factors (Lubos et al., 2014; Miao and Miller 2000). Yersinia is a striking organism secreting several effector proteins such as YopJ (Orth, 2002), YopN and YopM (Boland et al., 1996; Ho¨fling et al., 2015), YopE (Schotte et al., 2004) and YopT (Iriarte and Cornelis, 1998). The mechanism of these proteins to pass through the host cells is inspiring for development of different cell penetration systems. In contrast to others, YopM protein of Yersinia includes leucine-rich repeating (LRR) motifs. Benabdillah et al. (2004) revealed that the first three LRRs (LRR13) and the 32 terminal residues of YopM behaved as NLSs, which were dissimilar to classical NLSs. YopM is an acidic protein and does not have any enzymatic activity. It was shown by Ho¨fling et al. (2015) that this protein was able to penetrate the cell barrier, localize in nucleus, and transport the protein. YopM is composed of two conserved antiparallel α-helices at its N-terminal, a short conserved C-terminal, as well as a leucine-rich repeat domain ranging between 12 to 21 LRRs, depending on the strain (Ru¨ter et al., 2010). Ru¨ter et al. used recombinant YopM for the first time to design a protein delivery system which did not require any additional factors. For this aim, they used the truncated versions of YopM and they concluded that one of the two N-terminal α-helices of YopM was sufficient for translocation of the protein cargo into the cells (Ru¨ter et al., 2010). Therefore, these bacterial effector proteins may have a great potential for delivery of various cargo molecules into target cells (Ru¨ter and Schmidt, 2017). For instance, in order to obtain a therapeutic protein against breast cancer cells, YopJ was genetically fused to glutathione S-transferase (GST). The resulting fusion protein was able to form self-assembled protein nanoparticles with sizes of 100 nm. In this study, the apoptosis inducer protein YopJ functioned as both a cell penetrating protein and a therapeutic protein. The protein nanoparticles showed selective cytotoxicity to breast cancer cell lines, including doxorubicin-resistant

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cells (Herrera Estrada et al., 2016). Similarly, Deng et al. constructed a fusion protein consisting of Pseudomonas exotoxin A (PE), and the hyperthermophilic archaeal histone (HPhA). The cellular-penetration ability of PE and DNA-binding ability of histone protein were the essential for design strategy. The results indicated that the fusion protein enabled a higher transfection efficiency than cationic liposomes (Deng et al., 2015) In our attempts to design a novel gene delivery nanocarrier, we used the cell-penetrating function of YopM for delivery of plasmid DNA-encoding GFP protein. For this aim, we produced recombinant YopM proteins from three different strains of Yersinia. Then, we prepared QD conjugates, including the purified YopM and plasmid DNA. The nanoconjugate did not require using the commercial gene transfection reagent and was able to deliver its plasmid DNA cargo into the cells. Moreover, the designed QD nanoconjugate enabled a higher GFP signal than the transfection reagent. In addition, we observed that three different YopM proteins resulted in different levels of GFP expression (unpublished study).

2.5 RECOMBINANT PRODUCTION OF CELL-PENETRATING PEPTIDES FOR THE DELIVERY PURPOSE Chemical synthesis of peptides is time-consuming, expensive, and difficult to scale-up. Conversely, CPPs should be produced in large yields for use in therapeutic applications. In terms of economical production, E. coli is the most appropriate host system. However, E. coli cells do not possess the post-translational modification machinery. Therefore, in the need for glycosylation or phosphorylation, eukaryotic hosts must be used. The small-size, possible toxic effects to the host cells and cationic charge make it difficult to produce peptides. Therefore, small-sized peptides are often produced as fusion proteins. Peptides can be fused to solubility enhancer or selfcleavable proteins (Li, 2011). For instance, Zhao et al. (2016) proposed a cleavable self-aggregating tag and optimized the purification conditions to produce a variety of therapeutically important peptides. In order to produce the toxic peptide GKY20 in E. coli, Pane et al. genetically fused the peptide to the C-terminal of onconase and introduced an acid-labile Asp-Pro linker sequence between the two partners. The fusion protein was heterologously produced in E. coli cells in the form of inclusion bodies. The recombinant fusion protein was solubilized from inclusion bodies, and then was partially purified by immobilized metal ion affinity chromatography (IMAC). Onconase was attempted to be removed by a mild acid-cleavage reaction, but the authors observed that treatment with dilute acetic acid resulted in the formation of undesired shorter fragments. Therefore, the authors needed to optimize the onconase sequence. Onconase was subjected to

2.6 Conclusion and Future Respects

rational design to introduce mutations replacing the acid-labile cysteine and methionine residues. The engineered onconase enabled a simple and low-cost production of the target peptide (Pane et al., 2016). Genetic fusion of CPPs to cargo proteins enables carrier protein design. Alternatively, direct fusion of two different partners often causes problems, such as poor stability, low production yield, inclusion body formation in E. coli cells, and the lack of rotational freedom of two different domains. To overcome these problems, the studies focused on designing the optimum linker sequences. For this aim, Gong et al. (2016) used a glycine-serine linker, which improved the production yield without any adverse effects on cellular uptake. In a study by Lin et al., the fusion construct of enhanced GFP (EGFP), Tat, and A2 subunit of cholera toxin (CTA2) was produced in soluble form. The His-tagged fusion protein was then purified from E. coli BL21 (DE3) bacterial lysis supernatant. In this study, penetration ability of Tat and localization function of CTA2 enabled a remarkable carrier system for EGFP (Lin et al., 2017). Saffarian et al. (2016) optimized recombinant production conditions for the fusion of Tat peptide and the light chain of botulinum toxin type A (Saffarian et al., 2016).

2.6 CONCLUSION AND FUTURE RESPECTS Overcoming cellular barriers is the major obstacle for the delivery of therapeutic molecules. CPPs, which have the intrinsic ability to penetrate cell membranes, offer a solution to this problem. In particular, BBB-penetrating CPPs hold great promise. CPPs, which have smaller structures, are less prone to inducing immune response. There are several successful attempts to employ CPPs as the components of nanocarrier systems. CPPs are robust structures that can be conjugated to various nanoparticles, such as liposomes, magnetic nanoparticles, gold nanoparticles, gold nanorods, and QDs. As described throughout the text, there are several successful examples of CPP-conjugated therapeutic nanodelivery systems. The number of newly discovered CPPs continues to increase. Some CPPs are derived from natural proteins and some of them are designed by using the existing sequences as templates. As an alternative to rational design approach, combinatorial peptide design involves screening several different peptide sequences. Phage-display is a frequently used method in combinatorial approach, but the cellpenetrating activity of the displayed peptides should be differentiated carefully. Due to the genetically encodable property, CPPs can be fused to the protein of interest. Once genetically fused, the fusion protein can be heterologously produced in an appropriate host. The most preferred expression system is E. coli, due to its fast growth kinetics, simple cultivation conditions, and wellcharacterized genetic manipulation techniques. For overcoming the problems related to producing cationic or toxic peptides, different solubility tags and

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carrier proteins are also available. During optimization studies and tracking the intracellular route of the designed delivery system in particular, CPPs are often fused to fluorescent proteins. Similarly, fluorescent protein can also be used as the reporter gene during CPP-mediated gene transfection studies. Chemical synthesis of CPPs offers an advantage when different functional groups need to be incorporated. For instance, unnatural side chains can be introduced by chemical synthesis approach. From an economical view, CPPs should be produced in large amounts, especially for therapeutic applications. When large-scale chemical synthesis becomes a limiting factor, recombinant DNA technology could be a convenient way to produce peptides. When designing CPP-based nanodelivery systems, the cargo capacity of the nanoconjugate, stoichiometric ratio of nanoparticle:CPP, as well as stability, biodegradation, and retention time properties should be optimized and reconsidered for each different application. As it is a common problem for all intracellular delivery systems, CPP-mediated nanocarriers should also be designed to escape from endosomal entrapping. Future studies are focused on designing novel CPPs with improved cell, tissue, or organelle-specific targeting properties. Tat peptide was the first CPP identified, and it was followed by many studies on novel CPPs with improved uptake properties. There is a great potential for employing CPP-conjugated nanoparticles for therapeutic delivery of proteins, nucleic acids and cancer drugs. In order to use CPPs as components of therapeutic agents efficiently, possible toxic effects of CPPs and the issues related to specific targeting should be considered and investigated. Moreover, there are still some issues that should be clarified regarding the uptake mechanisms and intracellular interactions of CPPs.

ACKNOWLEDGMENTS We thank The Scientific and Technological Research Council of Turkey (TUBITAK) for the financial support (project number: TUBITAK 113Z379).

REFERENCES Agarwal, R., Domowicz, M., Schwartz, N., Henry, J., Medintz, I., Delehanty, J., et al., 2015. Delivery and tracking of quantum dot peptide bioconjugates in an intact developing avian brain. ACS Chem. Neurosci. 6 (3), 494504. Agrawal, P., Bhalla, S., Usmani, S., Singh, S., Chaudhary, K., Raghava, G., et al., 2015. CPPsite 2.0: a repository of experimentally validated cell-penetrating peptides. Nucleic Acids Res. 44 (D1), D1098D1103. Alhakamy, N., Nigatu, A., Berkland, C., Ramsey, J., 2013. Noncovalently associated cell-penetrating peptides for gene delivery applications. Ther. Deliv. 4 (6), 741757.

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CHAPTER

The current perspectives of nanoparticles in cellular and organ-specific drug targeting in biological system

3

Arunachalam Muthuraman JSS University, Mysuru, India

CHAPTER OUTLINE 3.1 Introduction ...................................................................................................106 3.2 Role of Nanoparticle Action in the Biological System .......................................107 3.2.1 Endocytosis Mechanism of Nanomedicine ......................................107 3.3 Mechanism of Nanoparticle Action in Cellular and Subcellular System..............111 3.3.1 Antigen-Specific Action of Nanoparticles........................................112 3.3.2 Receptor-Mediated Action of Nanoparticles ....................................113 3.3.3 Folate Receptor-Mediated Action of Nanoparticles...........................113 3.3.4 Transferrin Receptor-Mediated Action of Nanoparticles ....................114 3.3.5 Epidermal Growth Factor Receptor-Mediated Action of Nanoparticles...............................................................................114 3.3.6 Integrins Receptor-Mediated Action of Nanoparticles.......................115 3.3.7 Neonatal Fc-Receptor-Mediated Action of Nanoparticles ..................115 3.4 Role of Nanoparticle Action in Pathophysiological Condition ............................116 3.5 Limitation of Nanoparticles Action in Biological Systems..................................117 3.6 Intracellular and Subcellular Targeted Action ..................................................117 3.6.1 Endosome/Lysosome-Targeted Action .............................................117 3.6.2 Cytoplasm-Targeted Action ............................................................119 3.6.3 Endoplasmic Reticulum and Golgi ApparatusTargeted Action .........119 3.6.4 Mitochondria-Targeted Action ........................................................119 3.6.5 Nucleus-Targeted Action ...............................................................121 3.7 Interaction of Nanoparticle in Biological System..............................................122 3.7.1 Interaction of Nanoparticle With Lipids ..........................................123 3.7.2 Interaction of Nanoparticle With Proteins .......................................124 3.7.3 Interaction of Nanoparticles With DNA ...........................................125

Nanostructures for the Engineering of Cells, Tissues and Organs. DOI: http://dx.doi.org/10.1016/B978-0-12-813665-2.00003-X © 2018 Elsevier Inc. All rights reserved.

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CHAPTER 3 The current perspectives of nanoparticles

3.7.4 Interaction of Nanoparticles With Other Smaller Biomolecules .........126 3.8 Pharmacological Action of Nanoparticle..........................................................126 3.8.1 Genomic Action of Nanoparticles ...................................................126 3.8.2 Proteomic Action of Nanoparticle...................................................127 3.8.3 Metabonomic Action of Nanoparticle..............................................127 3.9 Therapeutic Application of Nanoparticle..........................................................128 3.9.1 Effect of Nanoparticles in Cancer...................................................128 3.9.2 Effect of Nanoparticles in Vascular Disorders ..................................129 3.9.3 Effect of Nanoparticles in Neurological Disorders ............................129 3.9.4 Effect of Nanoparticles in Infectious Disorders................................130 3.9.5 Miscellaneous Action of Nanoparticles ...........................................131 3.10 Future Scopes................................................................................................132 Abbreviations..........................................................................................................132 Acknowledgments ...................................................................................................134 References .............................................................................................................134 Further Reading ......................................................................................................154

3.1 INTRODUCTION The nanometer range of tiny particles are ready to interact with biological molecules due to its size and mimicking action of biological materials like various proteins and cells. The size of particles between 1 and 100 nm is used as a nanoparticle for the preparation of nanomedicine. Nanomedicine is a current approach in the field of medicine and some of the medicine has been proved in clinical trials (Satalkar et al., 2016). The nanostructure of medicines overcome various medical problems, target-specific drug delivery, and improves the efficiency of drug actions (Satalkar et al., 2016). Furthermore, modification of nanoparticle properties (i.e., physical and chemical properties) plays a key in the regulation of biological responses (Satalkar et al., 2016). However, the current hurdle of the nanoparticle usage is the elimination of nanoparticle after delivery of nanomedicine from the target site, because deposition or accumulation of nanoparticles is able to cause potential toxicity in the biological system (HofmannAmtenbrink et al., 2015; Leung et al., 2015). Therefore, understanding the role of the nanoparticle in cellular, tissue, and organ-specific actions for drug delivery and handling nature of nanoparticle in the biological system, is essential in drug discovery process, becausenanoparticles are interacts with different variety of cellular enzymatic, ion channels and receptor signaling associated proteins (Sniadecki, 2010; Yun et al., 2013). Generally, nanoparticles are used for the delivery of nanomedicine. However, currently, nanomedicine is also used for diagnosis and imaging of the biological system. The major limitation of nanotoxicity of the nanoparticle can be reducing by the study of endocytosis, exocytosis,

3.2 Role of Nanoparticle Action in the Biological System

and clearance mechanisms of nanoparticles form the biological system (Chithrani and Chan, 2007; Hsueh et al., 2015).

3.2 ROLE OF NANOPARTICLE ACTION IN THE BIOLOGICAL SYSTEM The nanoparticle has multidisciplinary action in the field of medicine, especially in targeting specific drug delivery for diagnosis and treatment. Currently, it is a “magic bullet,” for cell-specific diagnosis and target-specific therapies (Mai and Meng, 2013; Nguyen and Zhao, 2015). It overcomes the various problems of conventional medicine therapy, such as poor solubility, nonspecific cytotoxicity, poor bioavailability, and lack of optimal pharmacokinetic and pharmacodynamics (Semisch et al., 2014). Nanoparticles play a key role in the drug carrier action of target-specific and site-specific locations (Cheng et al., 2013). These target- and site-specific actions vary, based on their physical characteristics, such as size, surface charge, shape, and mechanical strength; and chemical characteristics such as biological interaction, zeta potential, affinity, releasing of the drug, and releasing of the particle from various levels like systemic circulation, organ, tissue, and cell (Ma et al., 2013; Salatin et al., 2015). The current developments are focused on the hydrophilic and lipophilic action of nanoparticles because biological entries vary with these properties. The hydrophilic molecules, like polyethylene glycol (PEG) and polycarboxybetaine (PCB) posses the efficient permeation and retention (EPR) capacity via the interaction of immune system and circulatory cells (Charoenputtakun et al., 2015). In addition, this hydrophilic complex of nanomedicine enhances the prolonged systemic circulation, which leads to release the medicine in target- and site-specific locations like tumor cells and damaged tissue with EPR action (Huang et al., 2016; Rammohan et al., 2016). The internalization and removal of nanomedicine from targeted cells are also known as endocytosis and exocytosis action. These actions occur in the cellular (membrane permeability) and subcellular (cytosolic secondary messenger) levels (Halo et al., 2014).

3.2.1 ENDOCYTOSIS MECHANISM OF NANOMEDICINE Endocytosis is one of the key mechanism of nanomedicine entry into the cell and it follows the two kinds of reactions: nonligand (nonreceptor)-mediated entry and ligand (receptor)-mediated entry. The efficient entry of nanoparticles into the cells overcomes the drug delivery problem in target-specific cells via alteration or escaping from the barrier-blocking action (Fuchs et al., 2015; Schubertova et al., 2015). Generally, the endocytosis of nanoparticles across the plasma membrane barrier is known as translocation. The translocation of nanoparticles in the plasma membrane is a major challenge due to variables of the individual cell membrane

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and complex reaction, such as cellular adhesion, interaction, communication, and selection of entry pathway (ion channel, lipid carrier, osmosis, exchanger, and receptor) (Ali and Singh, 2016; Maisel et al., 2015; Rathje et al., 2013). Some of the endocytosis mechanisms are identified for nanomedicine, i.e., receptor interaction, conjugation, enzymatic action, phagocytosis, macropinocytosis, caveolindependent endocytosis, clathrin-dependent and independent endocytosis actions (Meng et al., 2011; Zielinski et al., 2016). In contrast, exocytosis transports the nanoparticle from plasma membrane to extracellular space. The dynamic endocytosis and exocytosis trafficking action are well regulated. There are two biological processes of cells that play a role in the endocytosis: phagocytosis and pinocytosis. The phagocytosis action occurs in dendritic cells, neutrophils, and macrophages (Fytianos et al., 2015; Meng et al., 2011). Whereas, the pinocytosis process occurs in all types of cells: viaclathrin/caveolae-dependent entendocytosis; clathrin/caveolaeindependent endocytosis and macropinocytosis (Chaudhary et al., 2014) In addition, the nutrients and solutes entry pathways are also identified as achieving efficient uptake of nanoparticles in the cellular system. It is also known as “biological pathway” and it facilitates the precise translocation of specialized cells such as cancer cells and enhances the successful therapeutic outcome (Kou et al., 2013; Schmalz et al., 2007).

3.2.1.1 Phagocytosis The primary stage of nanoparticle translocation is cellular recognization. The phagocytosis-associated endocytosis process is observed in epithelial cells; fibroblast; immune cells; specific phagocytic cells (monocyte, macrophages, and neutrophils); inflammatory response cells (basophils, eosinophils and mast cells); and natural killer cells (Baby et al., 2014). In normal conditions, these phagocytic cells engulf the disabled particles, senescent cells and infectious microorganisms such as bacteria and viruses (Petazzi et al., 2015). This characteristic is involved in the translocation of nanomedicine. In addition, phagocytosis processes are triggered via ligands and the foreign agentassociated interaction of cell surface receptors (Yameen et al., 2014). Furthermore, some of the soluble factors, such as complement system, antibodies, acetylcholine, laminin, fibronectin, C-reactive protein and type-I collagen recognize the foreign materials and facilitate the phagocytosis process (Bartneck et al., 2014; Yameen et al., 2014). This is also known as opsonization. Moreover, three major receptors are identified in the phagocytosis process: (1) Fc receptor family for IgG (FcγRI, FcγRIIA and FcγRIIA); (2) complement receptors (CR1, CR3, and CR4); and (3) α5β1 integrin receptor (Dalzon et al., 2016; Kim et al., 2014). The second stage of nanoparticle translocation is cellular internalization. This is varied, based on attractive forces such as van der Waals, electrostatic, ionic, hydrophobic, and hydrophilic actions between the phagocytosis cells like macrophages and nanoparticle surfaces (Lee and Lim, 2016). Nanoparticle translocation is also triggered by receptor-mediated recognition, i.e., opsonins (Mirshafiee et al., 2016).

3.2 Role of Nanoparticle Action in the Biological System

3.2.1.2 Pinocytosis Pinocytosis is one of the endocytosis processes and its role is in the cell trafficking process of eukaryotic cells (Twomey et al., 2016). Clathrin protein is identified in the process of endocytosis, due to its cellular signaling, membrane recycling, and nutrient uptake processes. Clathrin-dependent endocytosis (CDE) makes the proteinligand complex (such as vesicles) formation with multiple varieties of proteins such as epsin; amphiphysin; endophilin; adaptor protein complexes (i.e., AP-2 heterotetrameric complex and AP180); and clathrin assembly lymphoid myeloid leukemia (CALM) protein (Allard-Vannier et al., 2016; Isas et al., 2015; Schreiber et al., 2015). In the cellular process, these vesicle release from the plasma membrane viaguanosinetri phosphatase (GTPase) actions. Once vesicles are entered into the cytoplasm, proteinligand complexes are disassembled with auxilin and heat shock cognate 70 (HSC70) action (Chuang et al., 2015; Diesel et al., 2013). The internalized particles, i.e., uncoated-vesicles, are recycled and come back to the plasma membrane surface. Here, the vesicle can be a target for multiple levels of ligand binding and endocytosis process. It also supports the receptor-associated entry of nanoparticles. The cationic nanoparticles with polylactide-co-polyethylene glycol (PLA-PEG) coating are documented to endocytosis via CDE (Yang et al., 2015). The cationic polymer, such as poly(Llysine) and poly(lactide-co-glycolide) (PLGA), nanoparticles are widely involved in the cellular uptake process via CDE (Pongrac et al., 2016). In addition, mesoporous silica nanoparticles also are known to possess the potential cell entry action via CDE in human mesenchymal stem cells (hMSCs) and adipocytes (3T3-L1) cells (Li et al., 2015). This reveals that the positively charged nanoparticles have high affinity to clathrin and raise the internalization (endocytosis) process. Furthermore, systemic study of nanoparticle endocytosis process are limited due to changes of its structure and shape for efficient cellular uptake and internalization (Chakraborty and Jana, 2015). The nanoparticle entry occurs with clathrin-independent endocytosis (CIE) action. Generally, it occurs via bacterial toxins and cell surface proteins. The target of cell surface proteins mainly involves the normal cellular regulatory functional actions like cell repair, polarization, spreading, and modulatiory proteins of intercellular signaling (Zhou et al., 2014). The major difference between CIE than CDE is that CIE does not make the vesicle formation with a protein coat. Whereas, actin and actin-associated proteins play a key role in the cellular uptake process (Maltas and Ertekin, 2015). The CIE undergoes multiple pathways with the interaction of different varieties of cellular proteins such as Arf-6, RhoA and Cdc42 proteins (Yameen et al., 2014). Among all proteins, Arf-6-dependent endocytosis shows the internalization action with major histocompatibility complex (MHC) class I, β-integrins, and glucose transporter 1 (GLUT-1) via amino acid and extracellular matrix interactions (Parvani et al., 2015; Venturelli et al., 2016; Xu et al., 2013). Furthermore, RhoA and Cdc42 proteins make lipid rafts to carry the nanoparticles via β-chain of the interleukin-2 receptor (IL-2Rβ) in immune

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cells and fibroblasts (Frohlich and Roblegg, 2012). However, Cdc42 interacts with cholera toxin B (CtxB) and the Helicobacter pylori vacuolating toxin (VacA) proteins via cellular uptake process (Walker et al., 2016). Furthermore, caveolae proteins also contribute to the nanoparticle entry process in various cellular systems. Caveolae are flask-shaped proteins, are located in the plasma membrane and play a key role in multiple cellular processes, such as cholesterol homeostasis, endocytosis of large size of the proteins, and signal transduction process. They are widely located in several types of cells like fibroblasts, smooth muscle, adipocytes, and endothelial cells, whereas, they are absent in nerve and leukocyte cells (Rewatkar et al., 2016). In endothelial cells in blood capillaries, concentration is a higher percentage (70%) of caveolae; when adipocytes caveolae are accumulated in the plasma membrane site, it is 50% of caveolae in endothelial cells (Wang, 2014). The caveolae proteins consist of three major types of proteins such as caveolin-1 (CAO1), caveolin-2 (CAO2) and caveolin-3 (CAO3). In addition, the four major coat proteins are also identified and named as cavins proteins 14 (Cav1 to Cav4). The major proteins (CAO13) and coat proteins (Cav14) work together and potentially contribute to the endocytosis process (Bohmer and Jordan, 2015; Nassar and Parat, 2015). The caveolin-mediated endocytosis pathway localizes the caveosomes. At neutral pH, these caveosomes escape from the hydrolytic degradation by lysosomes. The negative surface charge of caveolae triggers the cellular internalization process. The poly(amido amine) (PAMAM) dendrimer combines with rabies virus glycoprotein 29-amino-acid peptide (RVG29) and alters the brain DNA functions (Rewatkar et al., 2016; Shen et al., 2014). The cellular internalization mechanism of PAMAMRVG29 is clathrin and caveolae-mediated endocytosis action with the interaction of brain capillary endothelial cells, crossing of bloodbrain barrier (BBB) and γ-amino-butyricacid-B (GABA-B) receptor and nicotinic acetylcholine receptor (AchR) (Chen et al., 2011; Liu et al., 2009).

3.2.1.3 Macropinocytosis Currently, one more mechanism is identified in the uptake of nanoparticles, known as macropinocytosis. This newer mechanism named as micropinocytosisassociated endocytosis (MAE). It interacts with actin filaments and dynamin proteins. Trisaccharide-substituted chitosan oligomers (TSCO) and linear chitosan (LCO) possess the MAE process with high efficacy in cellular uptake of the nanoparticle. TSCO initially enter into the cell via CIE system and successfully escape from the endocytic vesicles. Compared to LCO, TSCO has greater nanoparticles aggregation and more efficient cellular internalization (Garaiova et al., 2012). Actin-driven endocytic process (macropinocytosis) also known as macropinosomes. It is a typical route of cellular uptake process and it specifically occurs in apoptotic cells, viruses, and bacteria (Aleksandrowicz et al., 2011). It also has greater interaction with antigen-presenting cells of major histocompatibility complex II (MHCII). Macropinocytosis process is not similar to that of

3.3 Mechanism of Nanoparticle Action in Cellular

receptor-mediated endocytosis and phagocytosis. The activation of tyrosine kinase receptor (TRK) pathway supports the macropinocytosis process via epidermal growth factor (EGF) receptor and the platelet-derived growth factor (PDGF) receptor actions. This leads to enhancing the actin polymerization, actin-mediated ruffling, and formation of macropinosome proteins (i.e., Cdc42, Arf-6, and Rab5) (Vilella et al., 2014; Xu et al., 2013). Macropinosomes are more sensitive to acidic pH in the cytoplasmic region, which leads to enhancing the fusion events. The micron size of lapatinib-loaded nanoparticles with the formulation of egg yolk lecithin were shown to produce the efficient energy-dependent endocytosis via clathrin-dependent pinocytosis and macropinocytosis (Zhang et al., 2014).

3.2.1.4 Biological pathway of endocytosis Nutrients like entry process of nanomedicines are also known as “biological pathway of endocytosis.” The various primary nutrients such as lipids, fat-soluble vitamins and carotenoids are cross the plasma membrane by simple diffusion manner. Some nutrients, such as glucose and lipoproteins control the endocytosis process (Georgieva et al., 2014). The normal cellular uptake of nutrients is a programmed process. Whereas, in a pathological condition, the nutrient uptake process is an abnormal condition. Sometimes it is a unprogrammed process, for example high proliferative cells such as cancer cell progression undergo the abundant and higher rate of a nutrient uptake process. In addition, human melanoma cells express a significant number of amino acid transporters for leucine and glutamine (Yameen et al., 2014). Pancreatic tumor cells also raise the nutrient uptake via macropinocytosis process. Cancer cells utilize a higher rate of albumin to meet the glutamine demand via an active form of macropinocytosis process. This occurs due to an over expression of oncogenic Ras proteins (Prabhu et al., 2015; Srikar et al., 2016). Consequently, this specialized property and environment of abnormal cells may be a target for endocytosis and it makes a great impact on target-specific action of nanomedicine. Therefore, the biological pathway of endocytosis for nanomedicine may be a primary and potential gateway for the successful use of nanomedicine therapy (Fig. 3.1).

3.3 MECHANISM OF NANOPARTICLE ACTION IN CELLULAR AND SUBCELLULAR SYSTEM The efficient therapeutic action of nanomedicine is achieved by the selection of suitable nanoparticles for fabrication of nanomedicine and the interaction of nanoparticle in cellular and subcellular systems. The method and mechanism of endocytosis of nanoparticle and interaction of relevant plasma and cytosolic proteins are discussed in previous sections. This section focused on the mechanism of nanoparticle action in the antigens, receptors, ion channels, cytoskeletal proteins, mitochondria, and nucleus.

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FIGURE 3.1 This illustration shows the endocytosis mechanism of nanomedicine. Nanomedicine enters into targeted cells by multiple mechanisms: phagocytosis, pinocytosis (macropinocytosis) and biological entry. The phagocytosis process occurs in specialized phagocytic cells (monocytes, macrophages and neurophilics); inflammatory cells (basophols, eosinophils and mast cells); and natural killer cells. Whereas, the pinocytosismediated nanomedicine entry occurs via eukaryotic cells. In addition, the biological pathway of nanomedicine occurs via simple diffusion like carbohydrate, vitamins, and carotinoids action.

3.3.1 ANTIGEN-SPECIFIC ACTION OF NANOPARTICLES Prostate carcinomas cells express various types of antigen. The integral membrane of prostate glycoprotein is integrated with the neovasculature of various solid tumors. Neovasculature is also known as “angiogenesis.” This angiogenesis process is essential for the growth of cancer cells, due to excess requirements of essential nutrients. The docetaxel loaded polymeric (PLA-PEG and PLGA-PEG) nanoparticle (BIND-014) shows the anticancer potential via interaction with S, S-2-[3-[5-amino-1-carboxypentyl]-ureido]-pentanedioic acid (ACUPA) against prostate cancer (Jin et al., 2014; Yameen et al., 2014). The efficient action of nanomedicine is achieved by the modification of various physiochemical properties, such as particle size, target-specific ligand density, surface hydrophilicity, drug-loading efficiency, life span in the systemic circulation, and drug releasing ability in targeted tissues (Bharali et al., 2016; Saraiva et al., 2016; Siafaka et al., 2016). The similar physiochemical modification of docetaxel formulated nanomedicine is reported to ameliorate prostate cancer.


3.3.2 RECEPTOR-MEDIATED ACTION OF NANOPARTICLES In the pathological condition, cells are overexpressing the specialized proteins. Some proteins act as a receptor for specialized biomolecules and exogenous ligands. Some proteins are documented to carry the nanoparticle for cellular uptake process. Disease targeting receptor proteins are identified; such as folate receptor (FR), transferrin receptor (TfR), epidermal growth factor receptor (EGFR), G-protein coupled receptor, low-density lipoprotein receptor and lectins. The maximal intracellular delivery of nanoparticle is undertaken via receptormediated actions (Xu et al., 2013; Yameen et al., 2014). The receptor-mediated cellular entry of nanoparticle carries various small molecular proteins, such as peptides, aptamers, antibodies, DNA, siRNA, and miRNA because it has a higher affinity towards the receptor proteins (Leucuta, 2016).

3.3.3 FOLATE RECEPTOR-MEDIATED ACTION OF NANOPARTICLES Folic acid has 441 Da of proteins and it shows a high affinity (KD 5 1029 M) towards the FR. Furthermore, it allows selective folate-conjugated nanocarriermediated endocytosis in diseased cells. The over expression of FR is identified in different kinds of tumor cells, such as ovarian, lung, brain, and colorectal cancer. This receptor has 3844 kDa of glycoprotein and exists in two isoforms, i.e., FRα and FR-β. The alpha type of FR is expressed in epithelial cancer cells (Bandara et al., 2014). The expression rate of an alpha type of FR in ovarian cancer cells is 90%; in endometrial cells 90%; in brain tumor cells 90%; and renal carcinoma cells 75% expression (Muller and Schibli, 2013; Zhou and Xu, 2015). Whereas, beta type of FR is found to express in activated macrophages and hematopoietic malignant cells, like chronic myelogenous leukemia. In addition, it is also expressed in various pathological conditions with multiple cellular systems, i.e., type 2 diabetes, atherosclerosis, and rheumatoid arthritis. The major advantages of FR-delivery nanomedicine are: easy scale-up process; an efficient and wide variety of clinical applications; facilitation of the chemical modification; no toxicity or immune reactions because FR act as a vitamin; high stability with extreme temperature, acidic, and/or basic environment (Zwicke et al., 2012). The PEGylated liposomal formulation of doxorubicin (DOX) has higher therapeutic activity via FR conjugation (i.e., folate-PEG-distearoyl-phophatidyl-ethanolamine) in human nasopharyngeal epidermoid carcinoma cell lines and xenograft model (Andriyanov et al., 2014; Scomparin et al., 2015). In addition, human serum albumin and docetaxel formulation of nanomedicine possess antitumour activity in human hepatoma cell line via interaction of FR (Yata et al., 2010).

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3.3.4 TRANSFERRIN RECEPTOR-MEDIATED ACTION OF NANOPARTICLES Transferrin is one of the iron-containing plasma glycoproteins. The transferrinbinding proteins is also known as transferring receptor (TfR). It is a homodimer by nature and both dimers is identical (mirror image type) trans membrane subunits. It enhances the cellular uptake of iron. Each subunit of TfR has 84,910 Da of molecular weight with 760 amino acid residues. It has two varieties of active site, i.e., N-linked glycosylation site and O-linked glycosylation site; these sites play a major role in the action of TfR (He et al., 2013; Hulsmeier et al., 2016). The main function of the TfR is regulation of cellular uptake process of iron and cell growth. TfR is over expressed in various pathological conditions, such as immature erythroid, placental tissue, rapidly dividing cells, and cancer cells (100-fold higher) (Hulsmeier et al., 2016). Therefore, its plays an essential target receptor for the nanoparticle endocytosis. The TfR targeted nanomedicines, such as adamantane functionalized PEG, siRNA, and a cyclodextrin (CALAA-01); liposome-based oxaliplatin (MBP-426); p53 plasmid DNA (SGT-53); and RB94 plasmid DNA (SGT-94) are used for the treatment of various ailments, especially cancer cell proliferation (Camp et al., 2013; Yameen et al., 2014). In addition, TfR over expression is also identified in Neuro2A cells; therefore it also plays a key target for neuro carcinogen disease (Daniels et al., 2012; Dixit et al., 2015). The TfR-based formulation of anticancer drugs, i.e., doxorubicin and paclitaxel, treat the brain glioma cells in U-87 MG-luc 2 xenograft mice. The major advantage of this TfR-based nanomedicine delivery is more effective in the biological medium than in buffer saline; it also shows the rapid levels to reach optimum ligand density in systemic circulation and enhance efficient targeted cellular uptake (Arosio and Casagrande, 2016; Marelli et al., 2013).

3.3.5 EPIDERMAL GROWTH FACTOR RECEPTOR-MEDIATED ACTION OF NANOPARTICLES The growth of cellular systems is based on various growth factors. The dermal region undergoes the growth with the help of EGF. This factor-binding protein is also known as EGFR. Currently, it is also known as HER1. It is a member of the ErbB tyrosine kinase family. Sometimes, this factor also acts on HER2 (ErbB2), HER3 (ErbB3) and HER4 (ErbB4). The expression levels of EGFRs are higher in solid tumors such as colorectal, brain, breast, ovarian, pancreatic, and prostate cancers (Chung et al., 2017; Seshacharyulu et al., 2012). Furthermore, it stimulates the tumor cell growth, invasion, and metastasis. Small molecules and monoclonal antibodies are used to regulate the EGFR actions. EGFR-targeted ligands are EGF, transforming growth factor-α (TGF-α), heparin-binding EGF-like growth factor, betacelluin and epiregulin. The ligands and EGFR complexes enhance the homodimerization or heterodimerization with monomeric proteins of


EGFR family and ErbB family, which leads to alters the cellular signaling process (Seshacharyulu et al., 2012; Singh et al., 2016). The EGFR-targeted nanomedicine, i.e., Fab0 fragments of anti-EGFR antibody ceuximabconjugated DOX (anti-EGFR ILs-DOX), regulates the cellular signaling process in malignant brain tumors (Huang et al., 2015; Mamot et al., 2012; Yang, 2010). Furthermore, cisplatin-encapsulated gelatin nanoparticles show an anticancer effect in A549 cells and tumor-bearing mice via interaction of EGFRs (Huang et al., 2015). In addition, DOX-loaded PEG-PLGA-Au nanoparticles interacting with Herceptin leads to reduction of breast cancer cell growth (Luk and Zhang, 2014). HER2-targeted scFv antibody conjugation with DOX (MM302) possess anticancer activity. The humanized HER3 antibody, U3-1287 (AMG 888), nanomedicine is documented to reduce the growth of cancer cells and improves the tumor suppression levels (Nishio et al., 2015).

3.3.6 INTEGRINS RECEPTOR-MEDIATED ACTION OF NANOPARTICLES Integrins is a membrane-binding protein and it enhances the integration of cells and ligands via membrane receptors. These receptors are known as integrins receptors (InRs); they are a family of heterodimeric trans membrane receptors. They contribute to various vital functions, such as adhesion, migration, invasion, stress responses, proliferation, differentiation, survival, and apoptosis. Furthermore, they also modulate the endothelial cells-extracellular matrix proteins, which leads to regulating the intracellular signaling. The InR is one of the surface receptors; it is a group of dimerized proteins consisting of 18α and 8β subunits in αvβ3 receptors (Desgrosellier and Cheresh, 2010; Seguin et al., 2015). These receptors are over expressed in the endothelial cells of the tumor. The cyclic-arginyl-glycyl-aspartic acid (RGD) peptidesassociated nanomedicine complexes efficiently bind with InR due to its high affinity for αvβ3 receptors (Laitinen et al., 2013). The platinum drug and PLGA-PEG complex of nanoparticles are also known to bind with cyclic-RGD (Graf et al., 2012; Prabhu et al., 2015).

3.3.7 NEONATAL FC-RECEPTOR-MEDIATED ACTION OF NANOPARTICLES The expression of neonatal Fc-receptor (FcRn) in the intestine is responsible for safe transport of breast milk immunoglobulin (IgG) to the intestinal epithelium barrier. In addition, it is also expressed in the adult apical region of epithelial cells. The small intestinal expression of FcRn is equivalent to the neonatal expression (Baker et al., 2013). The oral route of nanomedicine administration shows the inability to cross the intestinal epithelium barrier. Whereas, FcRn-targeted

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nanomedicine overcomes associated pharmacokinetic problem in the oral route of administration. The parenteral administration avoids this kind of problem; however, the conventional oral route of administration is clinically important, especially where frequent administration of medicine is required in chronic disorders (Fokkink et al., 2016). Furthermore, the affinity of the Fc region of IgG and FcRn are pH-dependent; at acidic pH, i.e., ,6.5, it was shown as high binding property, compared to normal physiological pH (B7.4) (Baker et al., 2013; Schoch et al., 2015). The PLA-PEG polymeric nanoparticles have a greater ability of trans-epithelial transport with polyclonal IgG Fc fragment (NP-Fc). The targeted delivery of nanoparticles is 11.5 times higher than nontargeted nanoparticles. Moreover, the oral administration of insulin-loaded NP-Fc is efficient towards hypoglycemic actin in mice (Pridgen et al., 2013; Qin et al., 2016). Therefore, neonatal Fc receptor-targeted nanomedicines are shown as the most effective management in various chronic disorders supporting the pharmacokinetic and pharmacodynamic changes. Overall nanoparticle densities are raised in the bloodstream and interact with opsonin proteins (Magdelaine-Beuzelin et al., 2009; Yameen et al., 2014). The opsonin proteins attachments on the surface of nanoparticles enhance the macrophages recognization and eventually accumulate in liver and spleen. Unlike, these phenomena cause low efficiency in targeted tissue and raise severe systemic toxicity (Borges da Silva et al., 2015; Moghimi and Hunter, 2001).

3.4 ROLE OF NANOPARTICLE ACTION IN PATHOPHYSIOLOGICAL CONDITION The nanomedicine is the newer scope of medicine in the field of clinical as well as preclinical setup. Because it interacts with the different level of the biological system, such as the cellular membrane of all kind of human cell, tumor cell, and cancer cells Chang et al. (2016) and its subcellular location, it acts on multiple proteins and organelle, and interacts with nuclear materials (Liu and Franzen, 2008; Zheng et al., 2015). In pathological conditions, every cell undergoes various abnormal changes; this makes a newer environment of inter and intraellular space via alteration of ion channel protein, receptor proteins, pore opening, the reverse operation of exchanger, pH alteration, and expression of newer toxic proteins (Barry et al., 2014; Falaschetti et al., 2013). The fabrication of nanomedicine is able to bind effectively in pathological environments due to the close mimicking action of newer proteins; stability in the newer environment; maintaining of physiochemical properties of nanoparticle and biomolecular system; and identification of specific target and modulatory action of biomolecules as well as cellular signaling processes (Miao et al., 2014). These properties of nanomedicine are able to render the homeostasis condition from the pathological environment.

3.6 Intracellular and Subcellular Targeted Action

3.5 LIMITATION OF NANOPARTICLES ACTION IN BIOLOGICAL SYSTEMS The nanoparticle is known to regulate the biological system and makes the homeostasis condition from the pathological condition of the cellular environment. However, the complete utilization of nanoparticle has some limitation, such as: (1) the releasing capacity of nanoparticle from biological system is limited; (2) various nanoparticles are metal in nature, which have additional binding properties of some specialized human cells, which can cause specific organ toxicity; (3) the biological elimination remains to be clarified; (4) the predominant activation of cellular signaling may cause toxicity; and (5) activated signals follow the multiple signaling processes, but net biological effects may undergo the unwanted toxic effects (Kim, 2016; Renukaradhya et al., 2015). The overcome these nanoparticle limitation, is to make a better treatment approach of nanomedicine for the multiple genetic and chronic disorders. However, nanomedicine treatment also has some limitation, i.e., circadian, age, gender, racial, and pathological variability; hence these variabilities need to be considered to the development of successful therapy of nanomedicine (Manzoor et al., 2012).

3.6 INTRACELLULAR AND SUBCELLULAR TARGETED ACTION Cellular systems have a different molecular mechanism in nanoparticle entry and therapeutic action via the interacting the membranes channels and receptors; and cytosolic alteration of cytoskeletal proteins, mitochondria, nucleus, etc. Other than this target, some of the endosomes, lysosomes, phagosomes, ribosomes and their proteins also contribute to the successful therapy of nanomedicine (Sakhrani and Padh, 2013).

3.6.1 ENDOSOME/LYSOSOME-TARGETED ACTION Endosomes and lysosomes are a primary site in the cytoplasmic action of the nanoparticle. The ligand-conjugated nanoparticle entry and tissue binding action with specific protein and drugs are discussed in the previous section. Various cytotoxic drugs act on cytoplasmic proteins and some cytoskeletal proteins, such as microtube polymerization (Banerjee et al., 2016; Shibata et al., 2013), whereas, antiviral drugs bind to nuclear protein-specific actions. Therefore, nanomedicine must display the ability to cross specific cells and interact with cytosolic and nuclear proteins. The endosome-specific enzyme activities are reduced with lower luminal pH, i.e., 6.06.5. At this pH, the activity of endosome-specific enzymes level is raised and triggers nanoparticle response in the cytosol region (Kou et al., 2013; Lucas et al., 2015). Protonated polyaminesbased nanomedicine is reported to enhance the luminal transport and raise bioavailability. However, the

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polyamine-based nanomaterials show poor biocompatibility and a high level of cytotoxicity with high pKa, i.e., (B9) and raise the membrane lysis at acidic as well as physiological pH (Koren and Torchilin, 2011; Nehoff et al., 2014). Intracellular trafficking is regulated by Ras-related protein (RAB) proteins (i.e., one of the G-proteins superfamily) and it contributes to the multiple cellular signaling processes via alteration of effectors or regulatory proteins. In cancer cell progression, it is shown to raise the levels of RAB proteins and it is therapeutically implicated with various growth factor receptors’ recycling action. The controlled endocytosis and endosomal trafficking is an essential cellular function (Wandinger-Ness and Zerial, 2014; Zhang et al., 2016). In neurodegenerative diseases conditions such as Alzheimer’s disease, Huntington’s disease, and autism are shown with the abnormal functioning of endosomal trafficking action (Ghanizadeh, 2012; Steketee et al., 2011; Yin et al., 2015). Sodium hydrogen exchangers (NHEs) are a family of nine Na1/H1 exchanger isoforms found in mammals. NHE-6 is identified in the endosomal trafficking process in different types of cell and has the regulatory function of lumen pH, recycling endosomes and lysosomes when pH value decreased at 5 (Fan et al., 2016). In addition, some of the NHEs, i.e., NHE15, are localized in cancer cell plasma membranes and initiate the progress of cancer. All malignant tumors reported to produce the reverse mode of hydrogen gradient, which leads to raising the acidic environment in the cytosolic region (Amith and Fliegel, 2013). In a tumor cell, intracellular pH is alkaline, i.e., 7.127.7, whereas the pH of the extracellular environment is acidic (6.26.9) in nature. The abnormal homeostasis of ion balance is due to the functional alteration (high activity) of NHE1. The NHE1 inhibitors are being employed for effective reduction of a cancer cell with selective antitumor drugs (Amith and Fliegel, 2013; Fan et al., 2016). Lysosomes contribute to the function of cellular digesting or recycling process with cathepsins proteases enzymes. The tumor cell shows the higher activity of lysosomal cathepsin levels, leading to promote tumor cell growth. The destabilization of lysosomal enzymes and lysosomal membrane permeabilization (LMP) raise the hydrolytic contents in the cytoplasmic region, which causes cell death (Domenech et al., 2013). However, drug resistance developed by cancer cells is due to changes of LMP proteins. Acid sphingomyelinase (ASM) is identified to alter the lysosomal membrane integrity, whereas, the inhibition of ASM with cationic amphiphilic drugs (CADs) supports the lysosomal membrane integrity (Baltazar et al., 2012; Funk and Krise, 2012). CADs are documented to produce the cancer cell toxicity, leading to control the cancer cell proliferation via efficient lysosomal functional action (Funk and Krise, 2012). Polymer nanoparticle-targeted liposomes are reported to produce biocompatible, bioresponsive, and biodegradable nanocarrier actions (Liu et al., 2016). pH-sensitive nanoparticles, i.e., PEGylated polyphosphazene and N,N-diisopropylethylenediamine (DPA), play a role in the alteration of cancer cell growth via activation of the lysosomal enzymatic system (Niazi et al., 2016).


3.6.2 CYTOPLASM-TARGETED ACTION Cytoplasm is a pool of biomolecular systems and it shows a wide range of biological reactions with the contribution of signaling, metabolic, and pathogenic actions. The availability of cytoplasmic nanomedicine and some cell-penetrating peptides (CPPs) are based on the ability to transfer the plasma membrane (Jafari et al., 2015). The direct translocation of CPP loses this property with conjugation of ligand. Some CPPs act as an anticancer agent, such as Azurin and XG-102, which are shown to reduce solid tumors via c-Jun-N-terminal kinases action (Liu et al., 2010). Coadministration of CPP and DOX or nab-paclitaxel are documented to produce the therapeutic efficacy against cancel cell growth via binding of αv integrins (Zhang et al., 2015).

3.6.3 ENDOPLASMIC RETICULUM AND GOLGI APPARATUSTARGETED ACTION In the cytoplasm, transport of nanomedicine is also follows an alternate way: bypassing the acid pH and hydrolytic lysosomal environment through the action of “retrograde trafficking mechanism” with the help of Golgi apparatus (GA) and endoplasmic reticulum (ER) (Sakhrani and Padh, 2013). Golgi apparatus and endoplasmic reticulum are regulated by the cytosolic calcium ion homeostasis, folding of membrane and secretory protein, and lipid biosynthesis (Parodi et al., 2015). The retrograde trafficking mechanism is also known as recycling of certain receptors, such as mannose-6 phosphate receptor and certain toxins, such as ricin, shiga, anthrax and cholera toxins, with the interaction of ER action (Fang et al., 2015; Sandvig et al., 2004; Spooner et al., 2006). Irregular ER function raises the ER stress, which leads to inhibition of protein synthesis, refolding of proteins, and clearance of misfolded proteins. ER stress is known to cause various diseases, such as cardiac hypertrophy, degeneration of cardiomyocytes, liver diseases, neurodegenerative diseases, and diabetes (Christen and Fent, 2012; Iurlaro and Munoz-Pinedo, 2016; Zhuang and Forbes, 2014). The nanoparticle has the ability to bind with ER and alters the ER stress environment. In addition PLGA-based nanocarriers have an affinity to the Golgi apparatus of human bronchial epithelial (HBE), and renal tubule cells (Barati et al., 2016; Panariti et al., 2012). Moreover, the albumin coated nanoparticles also bind GA via the retrograde route (2016). The PCL and PLGA nanoparticles that targeted Golgi-associated endosome process ameliorate human HeLa cell proliferation (Miles et al., 2001; Yameen et al., 2014). However, nanoparticle uptake and localization in subcellular compartments varies, based on cell type and type of nanomaterial used.

3.6.4 MITOCHONDRIA-TARGETED ACTION Mitochondria is an energy-generating organelle and has a two-membrane structure, i.e., inner and outer mitochondrial membranes, along with mitochondrial

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DNA. In addition, it comprises the specific genome code for expression of mitochondrial proteins. In pathological condition of mitochondria, the abnormal expression of mitochondrial and nuclear genomes are stimulated leads to release the cytotoxic proteins (Kwon et al., 2016). The mitochondrial dysfunction is observed in multiple diseases, such as cancer, neurodegeneration, neuromuscular diseases, obesity, and diabetes. In this condition, the alteration of inner membrane materials, i.e., increase of the saturated phospholipids, and increased mitochondrial membrane potential, i.e., negative charge via an influx of anions (Kwon et al., 2016; Mallick et al., 2016). Cations are a key target to regulate the membrane potential. The cationic substance, such as triphenylphosphonium (TPP), is shown to produce the balance of delocalized positive charge and lipophilicity, which leads to cross the barrier. Moreover, TPP-conjugated small molecules and drugs are allowed entry into mitochondria. Generally, the liposome-based nanocarrier molecule is used for the fabrication of mitochondrial-targeted action (Pathak et al., 2015). TPP is conjugated with stearyl moieties, which mimic the lipid bilayer membrane of liposomes and it actively carries anticancer drugs, such as silanol and ceramide to mitochondria (Morad et al., 2016). Furthermore, a TPP-associated octaarginine liposome is also carried and delivers the nanomedicine into the mitochondria via macropinocytosis, due to the increased mitochondrial surface density with octaarginine (Kawamura et al., 2013a). In contrast, at low octaarginine surface density, liposomes supports the endocytosis via clatherin-mediated actions, which leads to raising the lysosome transfer for degradation (Kawamura et al., 2013b). TPP with PLGA-PEG nanoparticle-targeted mitochondria-acting drugs has optimal surface charge and mitochondrial uptake. Then, different ratios of TPP with PLGA-PEG showed efficient mitochondrialmediated action. Such functional groups are COOH group of PLGA (PLGACOOH) and the hydroxyl group of PLGA-PEG (PLGA-PEG-OH). The positively charged nanoparticles exhibit high accumulation in the mitochondria of human cervical cancer (HeLa) cell, when compared to negative surface charge molecules (Abe et al., 2016). Consequently, the buffering effect of positively charged nanoparticles prevents the acidification of endosomal vesicles, which leads to raising the ATPase-mediated influx of protons (Marrache and Dhar, 2012). Furthermore, it produces the osmotic swelling, and ruptures of endosomal membrane leads to release the materials from the cytosol. The nanoformulation of mitochondria-acting drugs, such as nidamine and α-tocopheryl succinate, reduced the HeLa cells growth (Mallick et al., 2016). Mitochondrial antioxidant formulation with curcumin treats the neurodegeneration of human neuroblastoma IMR-32 cells (Wongrakpanich et al., 2014), and mitochondrial uncoupled regulator, i.e., 2,4-dinitrophenol, treats the obese 3T3-L1 cells (Fu et al., 2013). Therefore, TPP nanoparticles exhibit the greater therapeutic index via an efficient mitochondrial function. Currently, TPP nanoparticles poly-L-lysine (PLL) have shown efficient nanocarrier actions and are needed for the potential management of various mitochondrial disorders (Wang et al., 2013; Yue et al., 2016).


3.6.5 NUCLEUS-TARGETED ACTION The nucleus plays a core controlling mechanism in every cell. It is wrapped with the double lipid bilayer and it has essential therapeutic targets, such as proteins, nuclear receptors, mRNA, and DNA (Mallick et al., 2015). The alteration of these targets is observed in various cardiovascular and neurovascular diseases (Marcus et al., 2016; Santoso and Yang, 2016). Nucleus-targeted nanomedicines produced therapeutic action via regulation of DNA and gene delivery. The major hurdle of this nucleustargeted nanomedicine is entry or diffusion of nanomedicine to the nuclear target through the cytosol. Polymeric material has a promising role in the entry and reach of the nucleus. Polyplex derivative of N-(2-hydroxypropyl) methacrylamide (HPMA) and methacrylamide monomers bear the L-lysine groups and enhance the nanomolecule entry, improving the nuclear accumulation of plasmid DNA (Abd Ellah et al., 2016; Shi et al., 2013). However, the alteration of internalization and localization of endosomes and lysosomes delays the nuclear delivery. Whereas, caveolin-mediated endocytosis is a support to overcome nuclear delivery problems. Caveolin-mediated endocytosis involves the cell internalization pathway and it delivers the endocytosed material at neutral pH (Dalal et al., 2016). However, it can be bypass the acidic and enzymatic degradation in lysosomes and sort the nanomaterial cargo function to Golgi apparatus and ER to produce the safer entry of nanomedicines into the nucleus. The primary amino acid and other biomolecules such as L-arginin and saccharide moieties are identified as carrier polymers for sorting the GA and ER (Chinen et al., 2015; Sapsford et al., 2013). The major function of saccharide mimics the natural process of glycosylated protein transport from ER to the nucleus (Luk and Zhang, 2015; Molinaro et al., 2016). Moreover, the new polymers with glycosylated proteins or arginine, such as β-cholanic acid and glycol chitosan polymer nanoparticles, also provide the caveolae-mediated nucleus entry (Bryde and de Kroon, 2009; Suarato et al., 2016). 20% of these nanoparticles are localized in lysosomes within 60 min and most of the nanoparticles accumulate in the perinuclear region. The success rate of this approach is limited, due to the activation of drug resistance machinery in the cellular and subcellular systems. The precise reaching of the nuclear target and escaping from cytosolic machinery is a challenging task (Dalal et al., 2016). Recently, the understanding of nuclear pore complexes (NPCs) supports the transportation of nanoparticles across the nuclear envelope. NPC consists of nuclear proteins and nucleoporins (Nup). They plays a role in nuclear assembly and maintenance of structure and function of the nuclear system. The active and passive transport are involved in the entry of nanomedicine into the nucleus via NPCs (Higby et al., 2016; Kodiha et al., 2015). The oligopeptide sequences support the active transport of nanomedicine via NPCs through the specific binding nature to the receptors (Yang et al., 2014). The phenylalanine-glycine (FG) domains of NPCs are reported to alter the inner channels of NPC, which leads to entry of the nanoparticles in targeted site of the nucleus (Doello, 2013; Zwerger et al., 2016).

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In addition, the conjugation of nuclear localization signal (NLS)-derived nanoparticle such as CGGGPKKKRKVGG, functionalized PLGA nanoparticles. Quantum dotconjugated PLGA nanoparticles penetrated to the nucleus of HeLa cells (Sims et al., 2016; Yameen et al., 2014). Furthermore, NLS-functionalized doxorubicin-loaded PLGA nanoparticles were shown to produce efficient delivery potential of the drug into the nucleus of MCF7 cells (Jain et al., 2015). The dualtargeted (cellular and nucleus) nanoparticles now have potential cellular uptake and localization in the nuclear compartment. The loading of Indium-111conjugated trastuzumab, Fab (HER2 antibody) and CGYGPKKKRKVGGconjugated methotrexate possessed the greater delivery and targeted action on the nuclear system (Hoang et al., 2013). The subcellular fraction expresses the higher levels of HER2 in MDA-MB-231, MDA-MB-361, and SK-BR-3 cells lines; and NLS copolymer micelles have an antiproliferative effect via dual-targeting mediated nuclear action (Altenburg et al., 2011; Ampuja et al., 2013). The newer formulation of DOX-loaded folic acid and Ac-CGYGPKKKRKVGG functionalized chitosan micelles have higher cellular uptake in KB cells due to folic acid, and increase nucleus localization, due to the NLS action. The dual-targeted medicines possess a promising role in the management of genetic as well as an abnormal nuclear protein-associated chronic disorders (Yu et al., 2014).

3.7 INTERACTION OF NANOPARTICLE IN BIOLOGICAL SYSTEM The oral and intravenous route of entry of nanoparticles in the biological system involves multiple mechanisms encountering the various biomolecules. A detailed description is explained in the previous section. The biological interaction of nanoparticles exhibits strong interactions with various cellular proteins, phospholipids, nuclear materials, and multiple small molecules in the cytosolic regions (Zhang et al., 2012). The common route of nanomedicine treatment is systemic administration, i.e., intravenous injection. The primary interaction of injected nanoparticles is blood proteins followed by distribution in various target-specific organs (Blanco et al., 2015). Thereafter, these nanoparticles also enter the cell nucleus by the interaction of membrane proteins, ion channels, and carrier proteins, followed by cytosolic proteins and subcellular organelles. Crucially, a nanoparticle enters into the nucleus and interacts with nucleic acids like DNA and mRNA (Albanese et al., 2012; Jin and Kim, 2014). This process may be sequential or direct action on cellular or subcellular levels. Sometimes, these actions of nanoparticles are terminated by various subcellular proteins and enzymatic systems. The adsorption of nanoparticles in the blood or cells occurs with the support of endogenous biomolecules such as amino acids, biotin, and folic acid (Shang et al., 2014). Whereas, the depletion of small metabolites in abnormal location

3.7 Interaction of Nanoparticle in Biological System

(Saptarshi et al., 2013), inability of dissociation of nanoparticle and drugs (Alexander et al., 2012), higher irreversible binding of nanoparticles with biomolecule (Saptarshi et al., 2013), or difficulty of nanoparticle exit may cause the toxicity of nanoparticles (Rahi et al., 2014). In this section, explore the interactions of nanoparticles with proteins, lipids, DNA and small molecules.

3.7.1 INTERACTION OF NANOPARTICLE WITH LIPIDS The various mammalian cell membranes are made up of phospholipids. The understanding of nanoparticle interactions with lipids supports the efficient delivery of nanoparticle; and safe nanotechnology and better treatment of nanomedicine. The nature of nanoparticle and phospholipid interactions is vary due to the alteration of physicochemical properties of the nanoparticle (Frost and Svedhem, 2013). The physiochemical properties can be asses by different techniques like nanoparticle sizes by using differential light scattering (DLS), transmission electron microscopy (TEM) and scanning electron microscopy (SEM) (Tuoriniemi et al., 2014). The surface charge of nanoparticles is assessed by ζ-potential analysis. This property is essential for the interaction lipid and nanoparticle interaction in the biological system. In addition, the binding potential of surfactant molecules with lipid surface can be measured by surface plasma resonance (SPR) absorption and UVVis absorption spectroscopy. Dipalmitoylphosphatidylcholine (DPPC) is a common surfactant molecule and it is strongly adsorbed on the surface of nanoparticles, i.e., carbon-based nanoparticles (Mu et al., 2014). The long aliphatic hydrocarbon chains of the carbon nanoparticle surfaces are shown the raising of ζ-potential, size and stability. It is ready to bind to biological membrane lipids and raise the reactive oxygen species (ROS) in bronchial epithelial cells and A549 alveolar epithelial carcinoma cells (Jing et al., 2015). In addition, DPPC-coated single-walled carbon nanotubes did not show any cytotoxic or fibrogenic effects. Whereas, the colloidal gold nanoparticles are stabilized the membrane phospholipids structure and reduce the surface tension (You et al., 2016). Even though, the size of the nanoparticles also affects the property of lipid interactions. Titanium-di-oxide (TiO2) nanoparticles are raising the surface tension of the surfactant by the interaction of phospholipid (Pera et al., 2014). The binding events of nanoparticle with cholesterol, triglycerides, and phospholipids in human plasma are depended upon the surface properties of the nanoparticle (Pownall et al., 2016). The hydrophobic nanoparticles have a higher content of cholesterol compared to hydrophilic nanoparticles (Harisa and Alanazi, 2014). Interactions between nanoparticles and membranes are essential to generate the integrity of the cell membrane. Hydrophobic and electrostatic potential plays a key role in the interaction of the nanoparticle-lipid membrane. Guanidinylated dendrimers carry a positive surface charge and rapidly adhere with the liposomal membrane due to the availability of negatively charged phosphate groups in membrane phospholipids (Marie

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et al., 2014). After binding, the dendrimer surfaces become hydrophobic in nature, due to charge neutralization. Moreover, the polycationic polymers, i.e., polyamidoamine dendrimers, pentanol-core polyamidoaminedendrons, polyethyleneimine and diethylaminoethyl dextran, are also shown to induction of lipid bilayers disruption leads to the formation of holes, membrane thinning and membrane erosion. Whereas, surface-charged polystyrene nanoparticles are ready to bind with liposomal membrane and induce the surface reconstruction of phospholipids (Mu et al., 2014). Quantum dots interact with lipid membranes and induce sudden changes of the electric field in the neuronal and muscular membrane. This biolipid interaction with nanoparticles is essential for the delivery of nanomedicine and efficient treatment of various disorders (Cavallaro et al., 2016). Polystyrene nanoparticles have 20 nm diameter and show greater interactions with the endothelial lipid membrane. Guanidinylated dendrimeric nanoparticle derivatives also adhered with liposomal membrane and increased the liposome fusion (Yamada et al., 2016). Consequently, the development of nanoparticle-lipid membrane interaction-based nanomedicine is useful for drug delivery and the reduction of nanotoxicity.

3.7.2 INTERACTION OF NANOPARTICLE WITH PROTEINS The cell membrane and subcellular system have abundant quantities of different kinds of proteins. The interactions of nanoparticles with proteins have ample scope for the efficient treatment with nanomedicine of multiple disorders (Saptarshi et al., 2013). Systemic administration of nanoparticles or membranecrossing nanoparticles are covered by plasma proteins. The successful distribution and cell entry of nanoparticles is due to interaction with cellular proteins. Protein binding of nanoparticles changes the surface properties of nanoparticles and alters their biodistribution and clearance. In addition, it also induces the conformational changes of bound proteins and exposes unexpected epitopes, which leads to activating multiple cell signaling. These nanoparticle-protein interactions are dynamic processes that depend upon the physiochemical properties of both components, such as binding affinity, stoichiometry, and kinetic properties (Zhu et al., 2013). The gold nanoparticles interact with the surface of ubiquitin protein and its regulates various cancer cell proliferations (Saptarshi et al., 2013). In addition, the polyacrylic acid-coated Fe3O4 nanoparticles and 1-(3-dimethylaminopropyl)3-ethylcarbodiimide hydrochloride are efficiently bound with human serum albumin (HSA), which also induces the protein cleavage (Madrakian et al., 2013). The surface of quantum dots molecules readily interact with hexahistidine-tagged cytochrome P450BSβ molecules (Khan et al., 2015). The study of nanoparticleprotein interaction helps to treat the multiple diseases as well as metabolic regulation interacting with the major enzymes of metabolism (Madrakian et al., 2013). Nanoparticle-protein binding events depend on the nature of proteins and micro environmental factors such as pH and ionic strength. Nanotubes interact

3.7 Interaction of Nanoparticle in Biological System

with bovine fibrinogen, IgG, transferrin, bovine serum albumin (BSA), soybean peroxidase and α-chymotrypsin, which alter pathophysiological functions (Mahmoudi et al., 2011). The C60 nanoparticle interacts with the human and BSA and HIV-protease and enhances penetration to the cellular lipid membrane. Graphene types of nanoparticle have planar thin layers and interact with serum protein, which leads to mitigating the cytotoxicity and promoting bone formation (Saptarshi et al., 2013). Gold nanoparticle surfaces have a strong binding property with thiol groups in cysteine proteins, BSA, and heparin-binding growth factors, which inhibit tumor angiogenesis (Arvizo et al., 2011). Anionic nanoparticles interact with cytochrome C and it causes the conformational changes of apoptotic proteins and proteolysis. Chymotrypsin also interacts with gold nanoparticle, which leads to regulating the various coagulative proteins (Choimet et al., 2016). In addition, it interacts with fibrinogen which produces Macrophage-1 antigen complement (MAC) receptor activation and antiinflammatory action (Deng et al., 2011). Iron oxide nanoparticles also increase conformational changes of proteins such as human transferrin, which leads to the release of iron from the storage site (Xu et al., 2014). In addition, various nanoparticles such as silica nanoparticles and polymeric nanoparticles also interact with circulatory, membrane, and cellular proteins (Saptarshi et al., 2013).

3.7.3 INTERACTION OF NANOPARTICLES WITH DNA Nanoparticles bind to nucleic acids like DNA in biological systems are used for various biomedical applications like gene therapy, siRNA delivery, DNA repair, cell and growth regulation. The interactive properties of nanoparticles and DNA depend upon the physiochemical properties of both molecules (Li et al., 2013). Carbon nanoparticles have a large surface area and bind to DNA spontaneously. Furthermore, this binding ability is increased by aqueous solubility of CNTs’ unique shape and electronic properties (Mu et al., 2014). DNA molecules of GT-rich sequences support the efficient nanotube wrap. Other types of carbon nanoparticles, such as graphene oxide nanosheets and fullerenes also induced strong interactions with DNA molecules (Sharma et al., 2016). The gold nanoparticle interaction with DNA molecules is due to the electrostatic force between the two molecules. Tiopronin-modified gold nanoparticle interaction has three steps: the first step is diffusion; the second step is the formation of DNA-GNP complex I; and the third step is the formation of a more compact DNA-GNP complex II, leading to induce the DNA conformational changes (Gianvincenzo et al., 2015). Miscellaneous nanoparticles, such as quantum dots, Ag and Pt nanoparticles also interact with DNA molecules with electrostatic and hydrophobic actions. This leads to DNA degradation via generation of hydroxyl radicals and DNA damage via oxidative stress (Friehs et al., 2016). Similarly, titanium oxide nanoparticles and copper nanoparticles also cause DNA degradation via induction of single oxygen radical (Li et al., 2014). Cationic PLL-modified

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silica nanoparticles have been reported to produce the inhibition of DNA transcription in a size-dependent manner (Yeo et al., 2015). Consequently, nanoparticleDNA interactions are used to develop the efficient treatment of various cancer disorders and control multiple cellular processes.

3.7.4 INTERACTION OF NANOPARTICLES WITH OTHER SMALLER BIOMOLECULES The nanoparticle interacts with various small molecules in the biological system; such molecules are carbohydrates, vitamins, hormones, amino acids, and nucleic acid bases (Sapsford et al., 2013). Normally, these molecules play a key role in cell signaling and cell physiological process. The rapid depletion of these molecules leads to a change in the cellular functions and it cause cytotoxicity (Adjei et al., 2014). Nanotubes are adsorbed by the various amino acids, such as arginine, histamine, methionine, phenylalanine and tyrosine; and vitamins like folate, riboflavin, and thiamine (Guo et al., 2008). The interaction of nanoparticles with folic acid and vitamin B1 causes cellular pathophysiological changes (Kumar et al., 2014).

3.8 PHARMACOLOGICAL ACTION OF NANOPARTICLE Nanomedicine interacts with and alters various biochemicals, which produces multiple pharmacological actions by modulating the cellular signaling, genomic, proteomic, and metabonomic functions (Volker et al., 2013). The exact mechanism and usage of nanomedicine are limited, due to a lack of information about the functional action of nanoparticles within the biomolecular system (Kim, 2016; Ragelle et al., 2016). Nanoparticle interactions with biochemical pathways are useful for the various pathophysiological conditions. In addition, nanomedicine is act on replication, transcription and translation process of genes which enhances the biosynthesis of small proteins, which alters the biochemical signaling process, response to intracellular demands, and regulates cellular stress (Shi et al., 2017; Yilmazer et al., 2014).

3.8.1 GENOMIC ACTION OF NANOPARTICLES The various endogenous and exogenous small molecules regulate greater than 5000 genes in the biological system. Similarly, nanoparticles also produce the regulatory function of multiple genes, due to the mimicking action of nanoparticles in biological materials (Thai et al., 2015). The genomic action of nanoparticles has been identified in multiple cellular systems, such as carboxylated nanotubes inhibiting Id genes of HEK293 cells (Rastogi et al., 2014). Nanotubes are identified as genomic interactive action, which regulates the cell cycle, signal

3.8 Pharmacological Action of Nanoparticle

transduction, apoptosis, oxidative stress, metabolism, transport, and immune responses in human cells (Brun et al., 2014; Frohlich et al., 2014). In contrast, PEG-modified nano-tubes affect fewer genes which reduces biological activity. PEG-modified silica quantum dots produce a modulating action on 38 genes at low doses (8 nM), whereas, at high doses (80 nM) they make changes in 12 genes. These genes are linked to carbohydrate, endocytosis, and vesicular protein action in specific stress conditions (Lam et al., 2015; Vladuta et al., 2010). Fe3O4 magnetic nanoparticles alter three varieties of mouse macrophage genes. Furthermore, this nanoparticle also alters the MAPK, TLR and JAK-STAT signaling pathways via interaction of respective genes (Liu et al., 2011). This action activates the inflammatory and immune reactions and inhibits biosynthesis and metabolism. Another nanoparticle, TiO2 nanoparticles, are also upregulated in the inflammatory response gene and cell adhesion genes; however, it does not alter the oxidative stress genes in human (Tsaryk et al., 2013). Gold nanoparticles also interact with heat shock and stress-related genes which increase cellular necrosis (Pan et al., 2009). The administration of copper oxidederived nanoparticles upregulates the MAPK genes and downregulates the cell cycle genes (Luo et al., 2014). Hence, the genomic approach of the nanoparticle is useful in genetic disorders.

3.8.2 PROTEOMIC ACTION OF NANOPARTICLE The alteration of a gene subsequently alters the protein synthesis process in the biological system. However, some of drugs and nanomedicine directly interact with proteins which manage the abnormal protein-dependent pathogenesis. This approach is also known as proteomic action (Gioria et al., 2016). The global protein expression profile with nanomedicine treatment is examined by using two-dimensional (2D) gel electrophoresis. A protein databank is prepared and analyzed by microarray techniques (Triboulet et al., 2015). In addition, nanomedicines are focused on the action of antibodies-associated protein targets and the detection of multiple proteins. More than 51 proteins are identified in human hepatic carcinoma (HepG2) cells (Nguyen et al., 2015) and the treatment of nanotubes significantly reduced the expression of cancer cell proteins that control the metabolic pathways, signaling pathways, cytoskeleton formation, and cell growth of HepG2 cells (Ju et al., 2014). However, the proteomic approach of nanoparticles is regulated by multiple biological actions such as stress response, metabolism, adhesion, cytoskeletal dynamics, cell growth, cell death, and signaling pathways (Ge et al., 2011; Pisani et al., 2015; Tarasova et al., 2017).

3.8.3 METABONOMIC ACTION OF NANOPARTICLE Metabolic abnormalities cause various cellular and subcellular changes in different pathological conditions. Nanoparticles are also known to cause various physiological and metabolic functional changes in the biological system systems. The

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abnormalities also observed in nanoparticles causes nanotoxicity. Therefore, the basic understanding of the metabonomic action of the nanoparticle is useful for the correct choice and treatment of various kinds of metabolic disorders (Ratnasekhar et al., 2015; Tang et al., 2016). Silver nanoparticles reduce the levels of reduced glutathione (GSH), lactate, taurine, and glycine; whereas it raises the levels of most amino acids, choline analogs, and pyruvate (Ashraf et al., 2016). This indicates that nanoparticles depleted endogenous antioxidant substances and induces the conversion of lactate and taurine to pyruvate which produces an oxidative stress environment in the biological system (Kummara et al., 2016). In addition, the treatment of magnetic iron oxide nanoparticles cause the reduction of triglycerides, essential amino acids, and choline metabolites levels, and increase glycerophospholipids, tyrosine, phenylalanine, lysine, glycine, and glutamate (Feng et al., 2010; Oliveira et al., 2016). It indicates that nanoparticles are ready to change the various biological metabolic process, such as phagocytosis and cell membrane perturbation. This nanoparticle interactive action undergoes two varieties of action: direct effect and indirect effects. In direct effects, nanoparticles are bound to membrane phospholipids, receptors, cytosol proteins, and DNA molecules, which modulate the physicochemical properties of specific biomacromolecules and nanoparticles (Mendes, 2013; Radu Balas et al., 2015). With indirect effects, nanoparticles activated the cellular signaling process via sequence reaction with cellular small molecules, leading to the generation of free radicals and it causes cytotoxicity (Bonnemay et al., 2015; Fu et al., 2014). Therefore, the consideration of metabonomic action of the nanoparticle is essential for the management of multiple disorders with nanomedicine.

3.9 THERAPEUTIC APPLICATION OF NANOPARTICLE The medical application of nanoparticles is the delivery of medicines in cellular and organ-specific location in the biological system. The major problem is nanomedicine-associated toxicity (Wang et al., 2015). Nanomaterial interferes with biological molecules and/or structures, due to their structural and molecular similarities. The application of nanoparticles is more because it has a wide range of actions, such as diagnostic devices, contrast agents, analytical tools, physical therapy, and nanocarriers of drug delivery process (Youns et al., 2011).

3.9.1 EFFECT OF NANOPARTICLES IN CANCER Drug-loaded nanoparticles are used for the treatment of various cancers, such as albumin-bound paclitaxel for breast cancer, lung cancer, and pancreatic cancer; doxorubicin-loaded liposomes treats HIV-related Kaposi’s sarcoma, ovarian cancer, and multiple myelomas; irinotecan-loaded liposome treats metastatic pancreatic

3.9 Therapeutic Application of Nanoparticle

cancer; paclitaxel, loaded with cetuximab-coated micelles, bind the EGFR and treats the multiple cancers; and loteprednol etabonate-loaded nanoparticles treated cataracts and inflammation (Yhee et al., 2015; Young et al., 2016). In addition, the nanostructured particle is also reported to produce the antimicrobial action (Babu et al., 2014; Weingart et al., 2013). Some of them nanoparticles is also known to produce potential therapeutic action of regenerative medicine, stem cell therapy, gene therapy, brain tumor implants, bone repair, drug discovery, and cosmetic applications (Auffinger et al., 2013; Baetke et al., 2015; Cheng et al., 2014; Jain, 2005; Nitta and Numata, 2013; Tautzenberger et al., 2012). Furthermore, nanomedicine is documented to treat life-threatening disorders like cancer, acquired immune deficient syndrome (AIDS) and neuroinfectious disease (Masserini, 2013). The administration of drug-loaded quantum dot nanoparticles treated lymph node tumors (Helle et al., 2012). Furthermore, superparamagnetic iron oxide (SPIO) possesses potential action against the prostate cancer (Sterenczak et al., 2012). However, clinical reports of these nanomedicines is limited.

3.9.2 EFFECT OF NANOPARTICLES IN VASCULAR DISORDERS The various cellular and molecular mechanisms play a role in the vascular effect of nanoparticles; undergoing the endocytosis and protein interactive actions. The administration of nanoparticles ameliorated the ischemic condition of peripheral vascular disease. In the ischemic stage, the blood vessel gets narrow via obstruction of peripheral arteries which reduces or blocks the nutrients and oxygen supply (Agyare and Kandimalla, 2014). Nanoparticles carried the drug towards the ischemic tissues and induced angiogenesis action via expression of pro-angiogenic proteins (Mukherjee and Patra, 2016). In addition, nonviral vectorloaded nanoparticles carried the pro-angiogenic genes and induced angiogenesis process via enhancement of cell retention, cell survival, and secretion of angiogenic factors (Tu et al., 2015). The vascular grafts are able to transform blood into the targeted organs; however, the biological system produces a rejection reaction against vascular graft via activation of inflammatory and immunological reactions. Regardless, the nanomedicine-associated tissue-engineered vascular grafts (TEVG) system is documented to produce beneficial action on transplant therapy (Rathore et al., 2012). Similarly, nanomedicines reduce multiple vascular disorders, such as stroke, myocardial infarction, and diabetic complication.

3.9.3 EFFECT OF NANOPARTICLES IN NEUROLOGICAL DISORDERS The primary goal of nanoparticle therapy in brain neurological diseases, such as stroke, Alzheimer’s disease, Parkinson’s disease, amyotrophic lateral sclerosis, multiple sclerosis, and vascular dementia is of nanoparticle crossing of the BBB. In physiological condition, the BBB restricts the nanoparticle entry into the brain.

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The specialized size and target-specific nanoparticle cross the BBB and treat the neurological disorders (Ashton et al., 2015). The brain microvasculature structure of different cells, i.e., endothelial cells, pericytes, astrocyte, and oligodendrites, recognize the nanoparticle (Gomes et al., 2014; Shilo et al., 2015). The lipidderived nanoparticle, i.e., liposomes, posses the potential therapeutic action in neurological disorders (Fiandaca et al., 2011; Fonseca-Santos et al., 2015). The primary mechanism of drug-loaded nanoparticle crossing the BBB and targeting the low-density lipoproteins (LDL) delivers nanomedicine into neurological tumors such as glioblastoma (Feng and Mumper, 2013). Similarly, various nanomedicines are reduce the symptoms of neurological disorders. Therefore, nanomedicine is a future therapeutic agents for CNS disorders. In addition, some nanomedicine also acts neurovascular disorders such as stroke, vascular dementia, migraine, cerebral aneurysm, and intracranial hemorrhage (Chen et al., 2014; Liu et al., 2014; Sehba et al., 2011). The paclitaxel-loaded nanoparticles are shown to reduce vascular smooth muscle cell proliferation in the brain (Chan et al., 2011). In addition, antiangiogenic agents, i.e., fumagillin-loaded nanoparticles, reduce the neurocancer progress via aVb3-integrin receptor action (Lanza et al., 2010). The administration of taxol-loaded albumin nanoparticles reduced the restenosis of cerebral blood vessels in experimental animals (Chan et al., 2011).

3.9.4 EFFECT OF NANOPARTICLES IN INFECTIOUS DISORDERS Infection is a process of microbial growth in the biological system causing various toxic effects by alteration of multiple physiological processes. Nanoparticles also target microbes, reduce microbial growth (static) and kills the microbes (cidal) (Li et al., 2015; Monopoli et al., 2012). Polymeric nanoparticles are reported to reduce the sepsis condition of the biological system. The antitubercular drug-loaded nanoparticle binds to Microbactium tuberculosis and its fight against tuberculosis organism (Garg et al., 2015; Pandey et al., 2003). The oral administration of rifampin, isoniazid, and pyrazinamide-encapsulated poly (-DL-lactide-co-glycolide) nanoparticle produces complete bacterial clearance in mice and guinea pigs (Pandey et al., 2003; Sharma et al., 2004). Moreover, ethionamide encapsulated poly(-DL-lactide-co-glycolide) nanoparticles are also shown to produce a beneficial effect against multi-drug resistant tuberculosis (Kumar et al., 2011). Accordingly, nanoparticles also played a role in the management of infectious disease. The infection of peripheral nervous system (PNS) is causes neuronal inflammation and neurodegeneration (Ramesh et al., 2013; Wada et al., 2013). This causes peripheral neuropathy. In addition, peripheral neuropathy is developed by various infectious organisms, such as herpes varicellazoster (known as shingles), herpes simplex and cytomegalo viruses (Guedon et al., 2015; Muneshige et al., 2003; Sansone and Sansone, 2014). Furthermore, the infection of West Nile virus also produces severe motor neuropathy with multiple inflammatory reactions (Sansone and Sansone, 2014). The infection of human

3.9 Therapeutic Application of Nanoparticle

FIGURE 3.2 Role of nanoparticle in cellular and subcellular action in the biological system. The nanoparticles act on target-specific cellular system via antigen and receptors like FR, TfR, EGFR, InR and FcRn. In addition, it targeted various subcellular organelle, such as endosomes, lysosomes, endoplasmic reticulum, golgi apparatus, mitochondria and nucleus. It interacts with various biomolecules like lipids, proteins, DNA and other small molecules (i.e., carbohydrates, vitamins, hormones, amino acids and nucleic acid). The net biological effect undergoes genomic, proteomic and metabonomic actions. This figure explores the mechanism of nanoparticle action in various pathological conditions like cancer, tumor, vascular, neurological and infectious disorders. Fc, folate receptor; TfR, transferrin receptor; EGFR, epidermal growth factor receptor; InR, integrins receptor; FcRn, neonatal Fc-receptor; GA, golgi apparatus; ER, endoplasmic reticulum.

immunodeficiency virus (HIV) was shown to produce a neuropathic pain syndrome with reduction of immunological reaction and viral protein-associated neuronal damage (Kokotis et al., 2013). Furthermore, some bacterial attacks like diphtheria, leprosy, and Lyme disease damage the nervous tissue, causing the peripheral neuropathy (Benoliel et al., 1999; Reis et al., 2014; Zimering et al., 2014). Therefore, nanoparticle treatment has potential restorative effect against the infectious disorders (Fig. 3.2).

3.9.5 MISCELLANEOUS ACTION OF NANOPARTICLES The nanoparticle is a current concept in the field of medicine due to its close mimicking action of the biomolecule. Drug-loaded nanoparticles (nanomedicine) predominantly contribute to the amelioration of multiple pathological conditions like cancer, and vascular, neurological, and infectious disorders (Kanwar et al., 2012), as discussed in the previous sections. However, the nanoparticle also plays a major role in other disorders; metabolic disorders like diabetes mellitus; autism

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disorders, malnutrition, endocrine disease, bone disorders, reproductive disorders, respiratory disorders, neuromuscular disorders, genetic disorders, and hepatic and renal disorders (Ackerson et al., 2014; Arefian et al., 2015). Nanoparticles’ role in these disorders is under the investigation. Therefore, the nanoparticle has a wide variety of therapeutic potential in the multiple disorders.

3.10 FUTURE SCOPES The current perspectives of nanoparticles are discussed in book chapter, along with possible cellular and organ-specific drug actions in the biological system. In addition, this chapter also discussed the pharmacological action and therapeutic beneficial of drug-loaded nanoparticles in pathophysiological conditions. However, the established toxicity of nanoparticles needs to be considered before the fabrication of nanomedicine. Furthermore, the selection of nanoparticle for potential therapeutic action should consider the following points: (1) type of targeted cell (specific); (2) available biological environment (pH, proteins, enzymes and membrane potential); (3) interactive proteins and other targeted biomolecules; and (4) metabolic or exocytosis action of nanoparticle. Consequently, cellular and pharmacological target-specific nanomedicine can play a promising role in the field of medicine.

ABBREVIATIONS AchR ACUPA AIDS ASM BBB BSA CADs CALM CAO1 CAO2 CAO3 CDE CIE CPP CtxB DOX DLS DPA DPPC

nicotinic acetylcholine receptor S,S-2-[3-[5-amino-1-carboxypentyl]-ureido]-pentanedioic acid acquired immune deficient syndrome acid sphingomyelinase bloodbrain barrier bovine serum albumin cationic amphiphilic drugs clathrin assembly lymphoid myeloid leukemia caveolin-1 caveolin-2 caveolin-3 clathrin-dependent endocytosis clathrin-independent endocytosis cell-penetrating peptides cholera toxin B doxorubicin differential light scattering N,N-diisopropylethylenediamine dipalmitoylphosphatidylcholine

Abbreviations

EGF EGFR ER FG FR GA GABA-B GLUT-1 GSH GTPase hMSCs HBE HepG2 HPMA HIV HSA HSC70 IgG IL-2Rβ InR LCO LDL LMP MAE MHC MHCII NLS NPCs Nup PAMAM PDGF PCB PEG PLA-PEG PLA-PEG PLGA PLL PNS RVG29 ROS SEM SPIO SPR trk TEM TEVG TfR TGF-α

epidermal growth factor epidermal growth factor receptor endoplasmic reticulum phenylalanine-glycine folate receptor Golgi apparatus γ-amino-butyricacid-B receptor glucose transporter 1 reduced glutathione tri-phosphatase human mesenchymal stem cells human bronchial epithelial human hepatic carcinoma N-(2-hydroxypropyl)methacrylamide human immunodeficiency virus human serum albumin heat shock cognate 70 immunoglobulin interleukin-2 receptor integrin receptor linear chitosan low-density lipoproteins lysosomal membrane permeabilization macropinocytosis-associated endocytosis major histocompatibility complex major histocompatibility complex II nuclear localization signal nuclear pore complexes nucleoporins polyamidoamine platelet-derived growth factor polycarboxybetaine polyethylene glycol polylactide-co-polyethylene glycol docetaxel loaded polymeric PEG nanoparticle poly(L-lysine) and poly(lactide-co-glycolide) poly-L-lysine peripheral nervous system rabies virus glycoprotein 29-amino-acid peptide reactive oxygen species scanning electron microscopy super paramagnetic iron oxide surface plasma resonance tyrosine kinase receptor transmission electron microscopy tissue-engineered vascular grafts transferrin receptor transforming growth factor-α

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TiO2 TPP TSCO VacA

titanium-di-oxide triphenylphosphonium trisaccharide-substituted chitosan oligomers Helicobacter pylori vacuolating toxin

ACKNOWLEDGMENTS The authors are thankful to Department of Pharmacology, JSS College of Pharmacy, Jagadguru Sri Shivarathreeswara University, Mysuru-570 015, Karnataka, India for their unconditional support and providing the technical facilities for the preparation of this book chapter.

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FURTHER READING Mocan, L., Matea, C., Tabaran, F., 2016. Photothermal treatment of liver cancer with albumin-conjugated gold nanoparticles initiates Golgi apparatusER dysfunction and caspase-3 apoptotic pathway activation by selective targeting of Gp60 receptor. Int. J. Nanomedicine 11, 1025.

CHAPTER

Precision medicine and drug targeting: The promise versus reality of target-specific drug delivery

4 Karel Petrak

NangioTx Inc., New York, NY, United States

CHAPTER OUTLINE 4.1 Precision Medicine ..........................................................................................155 4.2 Precision Drugs ...............................................................................................157 4.3 Progress Towards Precision Drugs ....................................................................158 4.4 Conclusion ......................................................................................................162 Abbreviations..........................................................................................................163 References .............................................................................................................163

4.1 PRECISION MEDICINE Let us start with how “precision medicine (PM)” has been defined. According to the National Institutes of Health (NIH), PM is “an emerging approach for disease treatment and prevention that takes into account individual variability in genes, environment, and lifestyle for each person” (Collins and Varmus, 2015). This medical model proposes that healthcare should be patient-customized, with medical decisions, practices, and ultimately therapeutic products being tailored to the individual patient. The intent of the initiative is to enable more accurately prediction which prevention and treatment of a given disease should work in a given group of people. It is different from the current generally applied approach, by which drugs are developed for the “average” person. So how this will be tested and approved? The initial “testing ground” of the precision medicine approach will be in oncology in which “targeted therapies” have been developed and some found to


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offer benefits (de Bono and Ashworth, 2010). Novel approaches, based on advances in immunology, have recently produced some significant outcomes (Snyder et al., 2014). The precision medicine initiative plans to achieve its goals as follows: • • • •

analyze many more cancer genomes perform more clinical trials with novel designs (Abrams et al., 2014) develop and validate more reliable models for preclinical testing store the resulting molecular and medical data in digital form and make these easily available to all. Specifically, for oncology,

•

•

•

generate new knowledge on current barriers for effective treatment, such as drug resistance, physical and genomic heterogeneity of tumors, limited drug efficacy and severe side effects, etc.; facilitate a new generation of scientists, through education and training, capable of creating new paradigm for handling a wide range of biomedical information, and adopting a holistic approach in considering molecular, genomic, cellular, clinical, behavioral, physiological, and environmental parameters; promote the adoption of new ways of doing science that embraces engaged participation and open, responsible data sharing, to further form strong partnerships between academia, industry, and patient.

Such an approach to science will likely generate observations and information needed for the development of new drugs, provided the efforts are supported strongly by the regulatory frameworks. Addressing the development of new drugs, 21st Century Cures Act: Key Provisions Related to Drugs (August 7, 2015) has recently been passed by the US House of Representatives. The Act parallels the aims of the PM Initiative, and adds the following: • • • • •

removing barriers to increased research collaboration incorporating the patient perspective into the drug development and regulatory review process measuring success and identifying diseases earlier through personalized medicine providing new incentives for the development of drugs for rare diseases investing in 21st century science and next generation investigators

Taken together, it could be argued that precision medicine does not offer any fundamentally new science-based approach. However, it does offer a new and perhaps a better way to organize scientific efforts. Before discussing drug targeting in the context of PM, let us recognize the difference between the terms “precision medicine” and “personalized medicine.”

4.2 Precision Drugs

The word “personalized” often implies that treatments and preventions are being developed uniquely for each individual; in PM, the focus is on identifying which approaches will be effective for which patients based on genetic, environmental, and lifestyle factors. The concept of PM is not entirely new; it has been widely applied in many countries, for example in blood transfusion and in the use of therapeutic antibodies (Wang et al., 2016). A “low likelihood of success of truly personalized medicine deployed across the healthcare spectrum” was discussed by Mi et al. (2010) who observed that “fully individualizing therapy is likely to be intractable due to complexity and cost” and argued that “data-driven patient segregation” into “quantitatively-related patient sub-groups” is much more likely to be achievable and effective. An and Vodovotz (An and Vodovotz, 2015) claimed that they indeed described earlier a concept that is now called “precision medicine”.

4.2 PRECISION DRUGS Several issues become very apparent when attempts are made to put PM into clinical practice. New precision drugs (PDs) will need to become available; an obvious question comes up: “How are such precision drugs to be developed?”; and “What approval process will be used?” Much emphasis has been placed on genome analysis and the use of new data for the development of gene therapy. Are the currently used tools good enough to give information for generating the therapeutic products such as those PDs use? As pointed out by An and Vodovotz (2015), the huge amount of data gathered, often under varying healthcare settings, coupled with a sketchy understanding of what most of the genes in the Human Genome actually do, it is hard to identify what differences may actually be significant in relation to a given disease. However, let us assume that relevant data will become available in time; how should we define a PD? • • • • •

the drug should be acting only on the targets of the disease and no other it should be fully effective at all times when present at sites other than those of the disease it should be fully inert; it should have no undesirable side effects it would be preferable if the drug is also easy to administer and not too frequently, ideally just once it should not be difficult to make and thus not too expensive (Joy et al., 1999).

This ideal, first postulated by Paul Ehrlich may be difficult to achieve. It is unfortunate that approaches that are being used and have been repeatedly utilized over decades, generating “promising” results, are repeatedly found to be flawed (Petrak, 2005, 2013, 2015a,b, 2006, 2012).

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New paradigms for generating “ideal” PDs are needed. Such paradigms need to involve “self-targeting carriers,” such as antibodies combined with highpotency drugs specifically selected or developed de novo that fully meet the specific pharmacokinetic requirements of targeted-drug delivery. Hopefully the PDs initiative will be successful in this direction (Petrak, 2015a,b). Before examining what has been achieved recently in the area of drug delivery let us recognize that the term “targeting/targeted drug” is being used to describe two different situations: (1) Drug action is directed to a molecular structure associated with the mechanism of disease but also present on a nondiseased/normal tissue; (2) Drug action is directed to a molecular structure uniquely associated with the mechanism of disease and only at a specific disease sites. Obviously employing the approach (1) above is not suitable for generating PDs.

4.3 PROGRESS TOWARDS PRECISION DRUGS Many drugs have been approved by the FDA as “targeted cancer therapies” (https:// www.cancer.gov/about-cancer/treatment/types/targeted-therapies/targeted-therapiesfact-sheet). Such targeted drugs are used to block the growth and spread of cancer by acting on specific molecules that are instrumental in the growth, progression, and spread of cancer. One type of target-specific drugs uses monoclonal antibodies that deliver toxic molecules specifically to cancer cells. Antibody is supposed to bind to its target cell, the toxic molecule that is chemically bound to the antibody is supposed to be taken up by the cell, and ultimately killing that cell. Overall, the therapeutic efficacy of such targeting systems appears to be low. The “targeting” has been frequently evaluated by the overall efficacy of the drug without demonstrating that the intended target-specificity was actually achieved. Unfortunately, none of these approved “targeted” drugs are without often-numerous and serious side effects (http://www.fda.gov/Drugs/InformationOnDrugs/ApprovedDrugs/ ucm279174.htm). Consequently, the wished-for, still fictitious “magic bullet” drug targeting has not yet been “reduced” to data-supported reality. Unique biological mechanisms responsible for the initiation and promotion of growth of cancer that are not also associated with the functioning of normal cells have not been identified as yet. However, structural features that have been found only on cancer cells, or at least found on cancer cells at higher concentration than on normal cells, have been found and are being exploited for drug targeting. However, it needs to be kept in mind that a structure found only on cancer cells may not be unique when more sensitive and more specific analytical techniques become available.

4.3 Progress Towards Precision Drugs

A comprehensive and up-to-date review of molecular drug targets has recently been published (Santos et al., 2017). The conclusions of the review are instructive. It recognizes that successful drug discovery and therapeutic utility development needs to be based on: (1) identification of pathological mechanisms of disease at the molecular and cellular level, with (2) clinical diagnoses involving identification of the disease at the body organ or cellular or both levels, and (3) therapeutics that can act on and modify clinical manifestations at the molecular level. Effective treatment of disease, including new drug development and drug targeting, remains a complex, costly and at times unpredictable “art.” Achieving the goals of PM and thus providing new insights into disease would certainly make the connection between disease and an efficient discovery of a new generation of targeted medicines, i.e., PDs. Overall, the therapeutic efficacy of existing targeting systems appears to be low. The “targeting” has been frequently evaluated by the overall efficacy of the drug without demonstrating that the intended target-specificity was actually achieved. This review examines recent literature published on this topic to gauge what progress has been made towards making site-selective drug delivery possible. While progress may be slow, there has at least been a gradual move to the recognition of needs for developing drugs directed towards molecular targets of disease. He et al. (2015) argue that a “variety of nanomedicine formulae provide a plethora of opportunities for developing new strategies and means for tackling metastasis.” A range of nanoparticles are presented that have been used for controlled release of drugs. They address the issue of targeting by stating that “. . .the preparation of desired functional nano-carriers as well as the integration of nanocarriers with drugs, targeting molecules and other functional elements. . .” is involved. However, they offer not a single suggestion as to how such targeting will be achieved. Furthermore, any targeting system needs to consider a given drug-carrier combination and its body distribution and drug pharmacokinetics. Their conclusion that “. . .more advanced and effective antimetastatic nanomedicines are expectable to be engineered in the future” is not at all convincing or helpful. Bazak et al. (2015) examined available literature to evaluate whether active targeting (such as antibody and antibody fragment-based targeting, antigen-based targeting, aptamer-based targeting, as well as ligand-based targeting) of nanoparticles facilitates uptake of drugs by the tumor cells themselves. They concluded that an optimum targeting strategy has not yet been found; improved precision of drug delivery is needed. In this context, it is relevant to mention that little if no attention is being paid by researchers developing nanoparticles to the potential toxicity of the new materials. Nanoparticles made, even from well-known materials, will in vivo be exposed to a new biological environment, according to their biodistribution and

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should not be assumed to exhibit the same toxicology as the micro and macro items made from the same material. An up-to-date comprehensive review of the topic has been published by Bahadar et al. (2016). Efforts over the last 50 years or so to use nonbiological materials such as solid colloids, polymers, liposomes, colloids, and particles of various sizes (with currently a popular size in the nano-range) have not generated an effective sitetargeting delivery system. Apart from the lack of active targeting with such materials, another misconception has been that existing drugs, developed using conventional drug-selection process, would benefit from targeting. Such drugs, when delivered to a specific site, would diffuse away before their pharmacologically effective concentration at the site could be reached (Petrak, 2005). Drug pharmacokinetics and the overall body biology must be considered in the overall process of developing drug-delivery systems (DDSs). Progress in cell and cell membranebased drug-delivery systems have been reviewed by Tan et al. (2015). They believe that cells and extracellular vesicles (EVs), being endogenous and hence “much safer and friendlier,” may become the next generation of DDSs and exhibit biological effects and targeting specificity that are needed to meet the demands of personalized medicine. Let us hope so. Combining understanding of mutations that activate or inactivate cancer signaling pathways that drive cancer, with advances in molecular immunology, enabled Sharma and Allison (2015) to target immune checkpoints to open up antitumor T cell responses that produced long-lasting responses in some patients. However, such approaches may not generate site-specific targeting. Guiding targeted drug delivery with imaging may provide an opportunity for an improved delivery of drugs to disease sites. Chakravarty et al. (2015) state that using conventional drug carriers (polymers, liposomes, micelles, dendrimers, microparticles, nanoparticles) together with single-photon emission computed tomography (SPECT) have the potential to “lessen the invasiveness of conventional treatment and rapidly monitoring the therapeutic efficacy.” Without providing any evidence, the authors state that “SPECT-IGDD is not only effective for the treatment of cancer but might also find utility in the management of several other diseases,” while recognizing that “progress toward the clinical adaptation of this technology may be slow as of now.” The authors do not acknowledge that delivering a drug to a site of disease does not automatically guarantee its effective residence there. The same authors expressed a similar optimism earlier (Chakravarty et al., 2014) for PET, saying that “PET image-guided drug delivery has great potential to revolutionize patient care.” This misconception likely stems from the assumption that success with delivering an imaging agent to a site “guarantees” that the same can be done with drugs; this could only be the case if the physical properties (for example solubility) and the drug’s pharmacokinetics are the same as those of the imaging agent. Another approach that is being explored, which does not depend critically on the biology of the disease site, is the use of magneto-electric carriers loaded with

4.3 Progress Towards Precision Drugs

drug. Such carriers would have the ability to drive particles by electro-mechanical means precisely to the required location and allow the release of payloads ondemand, without depending on the physiological conditions of the target sites (Kaushik et al., 2014). Here again, however, the authors do not consider that inherent pharmacokinetics of the drug to be delivered will determine its efficacy after it has been released. The definition of how the “site” is to be determined is missing but one can anticipate that combining this with imaging might make it work, up to a point, at least. An increasing number of investigators (Kim et al., 2013) are recognizing that, while identifying biomarkers for the purpose of diagnosis and in turn of treatment, is a key for treating disease such as cancer, the current DDS including those based on nanotechnology are simply not yet sufficient for PM applications and, as such, PDs. Lanza et al. reported on the deliberations of a team of experts on the challenges associated with the applicability of DDSs in clinical medicine (Lanza et al., 2014). Examining, amongst other issues, the efficacy and safety of imageguided drug delivery, the team considered (1) optimization of drug concentrations delivered to the target cells mediating the disease, (2) how to avoid premature clearance of therapeutic particles/drugs before effective drug delivery is achieved, and (3) how to design particles to avoid undesirable biological responses, such as complement activation and adaptive immune response. The team concluded that, while vascular-targeted delivery of drug is feasible, an effective access of particles into sites of disease outside the vasculature has not yet been achieved. Furthermore, while novel approaches on utilizing natural transport mechanisms have been suggested and some demonstrated “in principle,” not enough is yet known about biological transcytosis mechanisms (for example their initiation, transport capacities, and biological control mechanisms) to reduce these ideas to practice. The importance of drug selection was addressed by Minko et al. (2013). Given the inherent significant fluctuation in response to anticancer drugs in different patients, the authors argue for an individualized selection of drugs, based on the molecular characteristics of tumors in individual patients, in order to improve treatment outcome. Viewing this from the point of view of drug delivery, such individual selection of drugs should also take into consideration the pharmacokinetics of drugs to be selected. Tumor microenvironment changes in time. Paulmurugan et al. (2013) suggest that therapeutic assessment should be done using dynamic and nonstatic imaging, and to bring it to the bedside. In this way, targeted therapies could be assessed in “real time.” This would indeed be a very welcome development for the development of site-targeting drug systems. Biomarkers remain popular for differentiating between genotypic or phenotypic cases (Cimino et al., 2013); however, these are frequently not suitable as targets for the delivery of chemotherapy and other drugs. In line with the aims of the PM Initiative, Liu (2012) argues that new cancer therapies should be based on

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comprehensive disease biology, right therapeutic reagents and drug-delivery vehicle, and should also take into consideration personal attributes such as age, sex, nutrition, psychological conditions, gut microbiome, circulating system, etc. According to Liu, nanocarriers can (1) improve the therapeutic efficacy of anticancer drugs, (2) selectively deliver drugs to cancer cells, (3) can be conjugated to antibodies or tumor-specific ligands, and (4) overcome drug resistance through escaping from P-glycoprotein. As an example, he offers the case of monoclonalantibody nanoparticles that can enhance cancer therapy by “targeting drug to [the] tumor specifically and sparing the surrounding normal tissues,” without, however, offering any data or reference. Zhang et al. (2012) examine interactions of nanomaterials and biological systems, and in the process provide a thorough review in the area of nanomaterials. Their focus claims to emphasize systems that have reached clinical trials, however, no data on effective nanomaterial-based therapeutic systems have been presented. There is at least one class of nanoparticles (typically in the size range from 3 to 5 nm) that could be called “magic bullets”: therapeutic antibodies (Petrak, 2015a,b; Larrick et al., 2016; Szamosi et al., 2016; Schumacher et al., 2016; Glukhova et al., 2016; Lai and Dong, 2016; Ferl et al., 2016; Suzuki et al., 2015; Wohlrab, 2015; Redman et al., 2015); we do have a proof that it can be done. A number of such antibodies have been approved by the FDA for cancer treatment (http://www.cancer.org/treatment/treatments-and-side-effects/treatment-types/ immunotherapy/monoclonal-antibodies.html; Scott et al., 2012). Antibodies have been found very effective to treat hematological cancers but not so for the treatment of extravascular malignancies, since access to such disease sites is limited by the size of antibodies. However, using antibodies that target molecular features of tumors for carrying anticancer drugs is fraught with the same obstacles shared by other nanoparticles. Even with intra-tumor administration of DDS, a major limitation is the inability of delivery systems to penetrate throughout the entire mass of the tumor (Aznar et al., 2017). Combination therapies using external stimuli (hyperthermia, radiofrequency ablation, magnetic field, radiation, and ultrasound) have been tried, but as yet, have made little impact on the efficacy of DDSs. Direct intra-tumoral administration of drugs deserves to be mentioned for the sake of completeness. For example, antibodies, pathogen-associated molecular patterns, recombinant viruses, and cells (such as autologous dendritic cells and tumor-reactive T lymphocytes) can all be delivered by this route. Delivering a viral vector expressing GM-CSF received recently an FDA approval (Goins et al., 2016).

4.4 CONCLUSION There remains a large chasm between the discipline of material science that creates new nanoparticles and tests them largely in vitro, and biology that works towards understanding of mechanisms that govern healthy and disease biological

References

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ABBREVIATIONS EVs DDS FDA NIH PET PD PM SPECT SPECT-IGDD

extravascular vesicles drug-delivery system Federal Drug Administration National Institute of Health positron emission tomography precision drug precision medicine single-photon emission computed tomography single-photon emission computed tomography image guided drug delivery

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CHAPTER

Brain targeting of payload using mild magnetic field: Site specific delivery

5

Murali M. Bommana1 and Sangram Raut2 1

Impax Labs, Middlesex, NJ, United States 2Texas Christian University, Fort Worth, TX, United States

CHAPTER OUTLINE 5.1 Magnetic Nanoparticles ...................................................................................167 5.1.1 Polymers Used in Magnetic Targeting ...............................................167 5.1.2 Physical Characterization of Magnetic Nanoparticles .........................171 5.1.3 In Vivo Distribution of Nanoparticles: Mechanism..............................173 5.1.4 Passive and Active Targeting............................................................174 5.1.5 Targeting Brain Delivery Using Magnetic Nanoparticles ......................174 5.1.6 Targeting Brain Tumors With Magnetic Nanoparticles.........................175 5.1.7 In Vitro Characterization of Nanoparticles: In Vitro BloodBrain Barrier Model .................................................................................176 5.1.8 Feasibility of Superparamagnetic Iron Oxide Nanoparticles for Gene Delivery Systems: Magnetofection.......................................177 5.2 Diagnostic Applications Using Magnetic Nanoparticles ......................................178 5.2.1 Magnetic Resonance Imaging and Other Applications ........................178 5.2.2 Cancer Theranostics .......................................................................179 5.2.3 Magneto Acoustic Tomography.........................................................180 5.2.4 Hyperthermia .................................................................................180 5.2.5 Challenges and Future Directions.....................................................181 References .............................................................................................................182

5.1 MAGNETIC NANOPARTICLES 5.1.1 POLYMERS USED IN MAGNETIC TARGETING In order to synthesize and stabilize magnetic nanoparticles, the primary requirement is the presence of a ligand or polymer with the appropriate functional group. Many of the compounds used in the literature are already FDA-approved for use


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CHAPTER 5 Brain targeting of payload using mild magnetic field

in pharmaceutical formulations, which makes them safer/biocompatible in developing magnetic formulations. The stability against aggregation and maintaining the monodispersed colloidal nature is highly important to the success of these nanoparticles as drug delivery vehicles. For ligand- or polymer-coated particles, steric repulsive forces play an important role in stabilizing the colloidal suspension (Vincent et al., 1986). As mentioned earlier, having an appropriate functional group on the ligand or polymer is equally important. Carboxylates and phosphates are widely used/found functional groups when it comes to synthesizing the magnetic nanoparticles. Carboxylates: Use of ligands such as citric acid, gluconic acid, lauric acid, and oleic acid is widespread in the published literature. The advantage of using citric acid is that prepared nanoparticles are easily soluble in aqueous environments, owing to the presence of at least one free carboxyl and a hydroxyl group. However, oleic acid, having a long carbon chain and single COOH group, leads to particles that are the only soluble in organic nonpolar solvents. It requires an additional step to replace the oleic acid with water soluble ligands such as dimethyl succinic acid. Phosphates: Alkanesulphoic acid and its surfactant derivatives have been tried in order to synthesize magnetic nanoparticles (Portet et al., 2001; Kreller et al., 2003). The proposed mechanism of binding is the bidentate attachment of oxygen atoms on the adjacent iron oxide surface. Sahoo et al. have reported the magnetic nanoparticle synthesis with excellent colloidal stability with alkyl phosphonates and phosphates (Sahoo et al., 2001). Moreover, researchers have demonstrated that phosphate derivatives of polyvinyl alcohol (PVA) can also be used to synthesize magnetic nanoparticles using a coprecipitation method (Mohapatra et al., 2006). Cysteine (-SH) has also been used successfully in the stable production of magnetic nanoparticles. Thiol groups also possess excellent affinity towards the iron oxide surface and these nanoparticles are relatively more stable in the biological environment. Gang et al. have used cysteine-terminated polyethylene glycol (PEG) polymers to produce stable particles (Gang et al., 2009). Amino-terminated ligands and dopamine have also been used successfully, however, dopamine stabilized particles are reported to be less stable in water and biological fluids due to hydrolysis issues (Shultz et al., 2007). Fig. 5.1 shows the different functional groups that are required to attach to iron oxide surface. Because the scope of this chapter is limited to the polymeric materials used in magnetic drug delivery systems, we will provide brief information about several widely used polymers as shown in figure below (Fig. 5.2).

FIGURE 5.1 Different functional groups found on ligands used in magnetic nanoparticle synthesis. A represents any aliphatic or aromatic side groups.

FIGURE 5.2 Several widely used ligands/polymers for synthesis of magnetic nanoparticles.

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Citric acid/polycitric acid: Citric acid is an inexpensive and biocompatible/ biodegradable compound that is used on a large scale in the food and drug industries. Citric acidstabilized magnetic nanoparticles are found to be common in the literature (Van Ewijk et al., 1999; R˘acuciu et al., 2006; Goetze et al., 2002). Although it is produced on a large scale and despite the intrinsic importance of this material, there are only a few reports on using citric acid polymer to produce magnetic nanoparticles, especially in drug delivery systems (Mashhadi Malekzadeh et al., 2017). Citric acid is a small molecule with three carboxyl groups. However, after coating nanoparticles with a monolayer of citric acid, the polycondensation/polymerization reaction under vacuum is followed to generate the polycitric acid layer. Dextran: This is a first-generation carbohydrate-derived polymer, offering excellent stability and biocompatibility considering the in vivo use of this formulation. It is composed of α-D-glucopyranosyl unit, with varying degrees of chain length and branching. It is a biopolymer of choice, due to its ability to adsorb on the iron oxide surface and confer stability in blood plasma (Berry et al., 2003; Laurent et al., 2004; Gamarraa et al., 2005). One of the stability is a limitation, due to lack of chemical bond with iron oxide surface unless it is modified to do so. Despite this limitation, many of the commercially available magnetic resonance imaging (MRI) agents, such as Ferimoxtran-10 and other Ferumoxides, use dextran as a stabilizing agent. An important factor is the choice of right size/molecular weight of the dextran in controlling the particle size. It appears that 40,000 g/mol offers particles with excellent colloidal stability and smaller size (Pardoe et al., 2001). PEG: PEG is a hydrophilic, water soluble, biocompatible and a classic pharmacokinetic stabilizer polymer. Many investigations have reported the use of PEG to increase the biocompatibility of the iron oxide dispersions and blood circulation times owing to its ability to make “stealth” particles (Kohler et al., 2006; Ghosh et al., 2011). Due to the PEG coating, these particles can escape opsonization by immune cells thereby increasing the mean residence time in blood circulation. PEG not only increases the circulation time but also provides excellent colloidal stability in water, cell media and blood (Cole et al., 2011; Zhang et al., 2008). Feruglose is the commercially available PEG-coated iron oxide formulation for MRI imaging. Several copolymers of PEG have been reported to be better choice due to the fewer step synthetic protocol; often called as “one-pot” synthesis. PEG and aspartic acid, PEG and glutamic acid, PEG and polyethyleneimine, are several of the copolymers that have been reported (Kumagai et al., 2007, 2009; Yoon and Jang, 2010; Kievit et al., 2009). PVA: PVA is a synthetic, biocompatible, water soluble polymer owing to its high number of hydroxyl groups. PVA also plays an important role in material science research, due to its inert nature and ability to provide its matrix for many thin film studies. Increasing/higher concentrations of PVA during

5.1 Magnetic Nanoparticles

synthesis of iron oxide nanoparticles have been known to affect the crystallinity of iron oxide nanoparticles (Lee et al., 1996). Such imperfections in the crystal strucutre are looked upon as a negative factor for MRI signal, so it is important to optimize the PVA concentrations and crystallinity of the obtained preparations. Ferrofluids and magnetic films using PVA have also been reported, due to its gel-forming abilities (Albornoz and Jacobo, 2006). Alginates: The linear copolymer of 14-linked β-D-mannuronic acid and its c-5-epimer α-L-guluronic acid. Presence of multiple carboxyl groups in alginates makes them suitable for synthesis of metal nanoparticles. Additions of divalent cations to alginates leads to gel formation, and thus can be used for making ferrofluids and magnetic films. Food, drinks, and pharmaceutical industries routinely use this copolymer. Several investigators have used alginates in making magnetic nanoparticles (Finotellia et al., 2004; Rocher et al., 2008; Zhang et al., 2010). Chitosan: Chitosan is also a widely used biopolymer derived from crustacean exoskeleton. This is an alkaline, nontoxic, biocompatible, and biodegradable polymer (Zhi et al., 2006). In many cases, chitosan serves as matrix or polymer coat for pre-synthesized iron oxide nanoparticles, rendering them water soluble and biocompatible (Chang and Chen, 2005; Chen et al., 2011; Zhu et al., 2009). Many times, amino groups in polysaccharide chains can coordinate with the iron surface to stabilize the particles. Amino groups can form stable amide bonds with the free carboxyl group (carbodimide reaction) of the ligand, which is already present on the iron oxide surface. Similar to other biopolymers, this can also be used to produce ferrofluids and thin magnetic films. Layers of chitosan were found to reduce the toxicity effects of magnetic formulations in in vitro studies (Shukla et al., 2015). Several other synthetic polymers have been used along with aforementioned ones. These include polymethacrylic acid (Ma et al., 2009), polyethyleneoxide (Aqil et al., 2008), polyvinylpyrrolidone (Liu et al., 2007), polylactic acid (Zhao et al., 2009) to name a few. Moreover, copolymers of these polymers have also been investigated to attain the desired surface properties (Fig. 5.3).

5.1.2 PHYSICAL CHARACTERIZATION OF MAGNETIC NANOPARTICLES A detailed approach involves systematic characterization of the variety of properties of prepared formulations. One needs to follow a tiered approach towards achieving the goal of characterizing the nanoparticles. Hall et al. have reported a series of steps to do so (Hall et al., 2007). This involves three tiered approaches: physico-chemical characterization, sterility and pyrogenicity test, and in vivo distribution and therapeutic testing. Physical and chemical characterization involves size and morphology, composition, surface chemistry, purity, stability, and

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Type

Surface modifications

Application under the external magnetic field

1. Targeting Surface embellishments Ligands Magnetic nanoparticles

2. Hyperthermia 3. Magnetic resonance imaging

Polymer coating

4. Optical imaging

Radionuclides

5. Nuclear medicine

Antibodies Conjugation :biotherapeutics

6. Cancer theranostics 7. Magnetofection

FIGURE 5.3 Schematics of the different applications of magnetic nanoparticles.

solubility. Pyrogenicity testing involves identification of contaminants and pyrogen that can possibly cause degradation of sample and unwanted immune reaction in small animals. Lastly, in vivo approach involves biodistribution of the injected dose and validating the intended action and use of the formulation using appropriate small animal models. Size: The size of nanoparticles is an elusive term, as no single method accurately determines the size. However, use of multiple techniques will help validate and establish the size and morphology of the prepared nanoparticles. In most cases, TEM is the method of choice for determining the size of the core iron oxide. High resolution TEM will further help to get an idea about the crystalline (crystal structure and lattice orientation) and amorphous nature of the material. Complementary dynamic light scattering (DLS) measurements will help establish the hydrodynamic size of the magnetic nanoparticles. This will reveal the size of the iron oxide core, along with ligand/polymer shell size. It is normal to obtain slightly higher mean particle size in DLS than TEM. Moreover, atomic force microscopy measurements will further establish the findings. All of these methods also help achieve particle size distribution of the sample, leading to calculating the polydispersity index. Such measurements will also reveal the aggregation and disaggregation behavior of sample under investigation.


X-ray diffraction: X-ray diffraction measurements can be done to obtain the crystal structure of iron oxide. Crystal size also is calculated from the line broadening using Scherrer equation. The proportion of different iron oxides can also be deduced from the intensity of the XRD peaks, by comparing the reference molecule peaks. Surface coating: As we saw earlier, the size from TEM and DLS will give different numbers, due to the presence of the protective shell of polymer. If done properly, this information can be used to calculate the thickness of the shell. Moreover, the binding and bonding interactions of different functional groups with iron oxide can be evaluated using NMR and FTIR. NMR is considered the so-called gold standard for structural characterization of organic molecules. Nanoparticles need to be soluble in order to display good NMR data. 1H NMR has been routinely utilized to characterize the binding of acid, phosphonic, amine groups to iron oxide surface. FTIR can be used to identify the functional groups present on the bare iron oxide nanoparticles, as well as to study the interaction of the polymer to iron surface. One can easily compare the reference vibrational peaks of different functional groups to the expected peaks in the sample. X-ray photoelectron spectroscopy can provide information on elemental composition. Relative contribution of Fe21 and Fe31 can be recovered from such measurements. Moreover, mass spectrometry will also help characterize the small molecule ligands and macromolecules attached to iron oxide nanoparticles. Techniques such as thermogravimetric analysis and differential scanning calorimetry will help determine the amount of the polymer coating weight by weight of iron oxide particles. The organic molecule coat on magnetic nanoparticles will not survive the temperature above 400 C but iron oxide will. This loss in weight can be used to quantify the amount of macromolecules used.

5.1.3 IN VIVO DISTRIBUTION OF NANOPARTICLES: MECHANISM The circulation half-life of nanoparticles dictate the distribution, which further depends on nanoparticle physical attributes like surface charge and particle size. As previously described, the particle size of 10100 nm plays a crucial role in the uptake of nanoparticles across biological membranes. Particle size .100 nm is opsonized and rid from the in vivo circulation by macrophages. This challenge was circumvented by the surface coating of PEG, which enhances the steric hindrance and increases the circulation half-life of the nanoparticles in vivo. Particles which are smaller than 10 nm are cleared by glomerular filtration and do not show their efficacy, except toxicity. Even with all the enhancements, nanoparticles, once injected, are cleared by the reticulo endothelial system (RES) like liver, spleen, and kidney, etc., which accounts for a larger portion of the dose of nanoparticles administered. The two ways to target the magnetic nanoparticles was by passive and active targeting which are discussed below.

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5.1.4 PASSIVE AND ACTIVE TARGETING Magnetic nanoparticles utilize the concept of enhanced permeability and retention (EPR) effect and passively accumulate a large concentration through the leaky vasculature of the tumor site. This process is very selective, based on the particle size distribution of the nanoparticles. There are two different types of targeting, based on the nanoparticle size and charge or surface embellishments. The nanoparticle size and surface charge dictate the passive targeting, and surface attachments like ligands or specific antibodies for the receptor play a role in the nanoparticle uptake by active targeting. Polymers or nanoparticles are used as drug carriers, prevent the renal excretion and increase the circulation half-life and controlled release of the drug. The tumor or the tumor cells have a leaky vasculature and impaired lymphatic drainage, which allows the high molecular weight polymers and the nanoparticles of size ,500 nm to accumulate in the tumor cells. The nodules of the tumor show enhanced permeation and retention effect, which retains drug or nanoparticle which was passively targeted by the carrier nanoparticles. Regarding active targeting, it involves the addition of the ligands/antibodies on the surface of drug delivery carrier system, specific to the membrane bound receptors, in order to have the payload reach the target site. Furthermore, active targeting could be synergistically enhanced by the application of external magnetic field. Under the influence of external magnetic field, the magnetic nanoparticles remain in the vicinity of the target tissue, when the magnetic force is substantial enough to compensate the hydrodynamic forces of the blood circulation to resist the movement of the carriers. Magnetic force is directly proportional to magnetic core and influences the targetability of the nanoparticles, but the hydrodynamic radius of the nanoparticles influences their clearance. The faster the clearance of the nanoparticles, the circulation half-life is impacted and, accordingly, the targeting ability of the nanoparticles. There should be a fine balance between the two parameters to have an optimal targeting of the magnetic nanoparticles. The surface coating thickness should be compromised, and an increase of the magnetic core size could be a good attribute to increase efficiency of the magnetic nanoparticles (Cole et al., 2011).

5.1.5 TARGETING BRAIN DELIVERY USING MAGNETIC NANOPARTICLES The bloodbrain barrier (BBB) is composed of endothelial cells supported by astrocyte and pericyte function, which act as a bulwark for the entry of xenobiotic compounds. This fortified wall with tight junctions does not demarcate between xenobiotic and benign components for the entry of the substances through it. Furthermore, the brain endothelial cell receptor shows multidrug resistance and presents an efflux strategy for lipophilic drugs. Only hydrophilic drugs and small


molecules diffuse via the paracellular pathway through the BBB, presenting a tough challenge for the entry of large hydrophilic (proteins and peptides) and lipophilic molecules. Targeting the drug or nanocomposites through BBB into the brain parenchyma is a difficult task, and a lot of research has been performed in recent decades to target drugs, using novel targeting carriers. With this in mind, superparamagnetic nanoparticles have an edge over paramagnetic nanoparticles. They can be magnetized over long distance, compared to paramagnetic materials. Based on the above information, nanoparticles could be good candidates for surface coatings, or conjugation with bio-therapeutics for targeted delivery. Magnetic particles, when injected, are cleared by reticuclo-RES and get lost in spleen, liver, and bone marrow, etc. These systemically cause toxicity and defeat the purpose of the use of nanoparticles. In order to overcome this, the nanoparticles are surface-coated with hydrophilic polymers to increase the systemic half-life and prevent or decrease the clearance from the in vivo system. Different coating materials, like chitosan, dextran and PEG coating increase the half-life of the circulation time, and surface embellishments like ligands or antibodies enhance the targeting efficiency of the nanoparticles. Magnetolipososmes are also embedded magnetic nanoparticles in the liposomes. Phospholiposomes are also cleared the same way as the naked Superparamagnetic iron oxide nanoparticles (SPIONs), hence the magnetoliposomes are coated with PEG to have an increased circulation half-time of the carriers. In the next level, magnetic nanoparticles are surface-coated with polymers and further embellished with antibodies towards the transferrin receptors, which are abundant on the surface of the endothelial cells. The targeting effect to brain can be divided into two stages; in the first place, the encapsulated magnetic nanoparticles embedded in the polymer coating system are coupled with the antibodies gearing towards the RES. The second stage is the influence of magnetic field to retain and transport the drugs across the endothelial cells (Thomsen et al., 2015).

5.1.6 TARGETING BRAIN TUMORS WITH MAGNETIC NANOPARTICLES Brain tumors at the malignant stage are the most lethal form of cancer and have a survival rate of 48 weeks from the day of diagnosis. At present, the scenario for the treatment, which includes surgery, radiotherapy, and chemotherapy. Despite available potent therapeutic agents with strong therapeutic potential, the most important challenge was to carry the payload to the tumor site after the systemic administration of the carrier. Recently, colloidal carriers are being evaluated for the brain tumor delivery and Victor Yang’s group had seen some success with magnetic nanoparticles. The accumulation of the nanoparticles in the vicinity of the tumor site releases the drug in a controlled manner to improve the concentration of the antitumor agent in the tumor region. The magnetic nanoparticle

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retention at the tumor vasculature or lesions under the influence of external magnetic field had a great potential for the scientists to work on iron oxide nanoparticles for brain tumor delivery. Studies have been performed in animal models; higher concentration of magnetic nanoparticles has been localized at the tumor site, when injected into systemic circulation. Targeting subcutaneous tumor is not a problem, compared to brain tumors, because of the position of the tumor (deep cranial tissue). Magnetic nanoparticles injected intravenously reached the site of action, but under the influence of magnetic field, their localized retention increased five times, compared to without the influence of the magnetic field. More research is progressing towards the optimization of the nanoparticles and the permissive retention of them for prolonged times at tumor lesions. Magnetic nanoparticles have a very short plasma half-life and present an advantage by being administered through the carotid artery. Because of advanced angiographic methods, carotid administration would be a feasible alternative for better targeting of magnetic nanoparticles to brain tumors. The carotid administration of magnetic nanoparticles poses a serious challenge in occlusion of the artery because of the external magnetic field and blood flow rate dynamics. The occlusion of the artery leads to serious complication of neurological issues. Chertok et al. experimented with magnetic force to minimize nanoparticle aggregation and thereby decrease occlusion of the carotid artery; further optimizing the retention of the magnetic nanoparticles in the tumor lesions. The group researched the effect of intravenous to intra-carotidal injection of the nanoparticles and their influence on retention at the tumor site under the magnetic field.

5.1.7 IN VITRO CHARACTERIZATION OF NANOPARTICLES: IN VITRO BLOODBRAIN BARRIER MODEL In vitro characterization of nanoparticles could be tested using cell culture studies. Recently, a lot of emphasis has been put on the development of a BBB model. Thomsen et al. tested superparamagnetic nanoparticles in an in vitro BBB model and tested the integrity of the model by studying the uptake of the nanoparticles by brain cells. The BBB is composed of endothelial cells, supported by astrocytes, pericytes and neurons, which act as a barricade to everything except small or lipophilic molecules. Human brain microvascular endothelial cells and rat astrocytes were cultured and were used in the establishment of in vitro BBB model in a transwell plates. The team evaluated the SPIONs under the influence of external magnetic field and has seen the particles reaching the cultured astrocyte region of the brain side of the in vitro BBB culture. Transepithelial electrical resistance (TEER) was used to check the monolayer integrity and the formation of BBB layer. SPIONs were stained with dye and further coated with hydrophilic starch. Because of the hydrophilic coating, the surface charge is slightly anionic and is better for increased circulation time in the plasma. The magnetic particles


without the influence of magnetic field were taken up by the endothelial cells. The high TEER values before and after the treatment confirms the integrity of the monolayer, and there is no leaky uptake of the nanoparticles. Under the influence of the magnetic field under the transwell insert, the magnetic nanoparticles showed a significant difference in concentration dependency, compared to controls. In another experiment with cultured rat astrocytes, the magnetic particles under the influence of external magnetic field have driven the particles to enter the astrocytes. Consequently, it showed that the magnetic particles under the influence of magnetic force could drive nanocarriers to the target site of interest. In conclusion, this study warrants that nanostructures are available in high concentrations in vivo at the site, given the attachment of targeting moieties to the nanoparticles specific to the receptors of the endothelial cells (Thomsen et al., 2013). In another study, Hong Ding et al. used transferrin-coated magnetoliposomes under the influence of magnetic field in an in vitro BBB model. In this approach, transmigration across the endothelial cells was done synergistically by the transferrin linkage and influence of magnetic field in magnetized liposome. The transmigration was higher than the individual variants of magnetic force or transferrin receptor targeting. In independent mode, when the nanoparticles are surface embellished with ligands, they can still undergo RES clearance, and the magnetic nanoparticles alone under the influence of magnetic field would reach the site, but will have no targetability. Accordingly, marrying both technologies would benefit the rate of uptake and avoid clearance because of the movement of magnetic nanoparticles being faster to the site of action under the influence of magnetic field (Ding et al., 2014).

5.1.8 FEASIBILITY OF SUPERPARAMAGNETIC IRON OXIDE NANOPARTICLES FOR GENE DELIVERY SYSTEMS: MAGNETOFECTION For efficient gene delivery systems, in addition to the superparamagnetic properties, nucleic acids (cargo load) must be transferred to the target site. Passive targeting can help in this process, but targeting the active process would be more beneficial in terms of targeting abilities of the nanoparticles. The concentrationdependent magnetic nanoparticles under the influence of magnetic field have shown potential for their use as gene carriers. Magnetic nanoparticles protect the payload from degradation, as well as help in the targeting of the nucleic acids to the target site. Initially, magnetic nanoparticles are used as nonviral transfection agents. The transfection can be guided by magnetic field, termed magnetofection, was studied in vitro and in vivo and shown the merits. It is a nonviral transfection method to target the cells using magnetic nanoparticles surface-coated with nucleic acids using an external magnetic field (Vladimir et al., 2013). McBain et al. (2008) reviewed an in vitro experiment where heparin sulfate linker was

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used to attach coated adeno-associated virus encoded with green fluoroscent protein, which gave increased transduction to the cells in vitro and in vivo through intramuscular injection to mice.

5.2 DIAGNOSTIC APPLICATIONS USING MAGNETIC NANOPARTICLES Nanoparticles have shown great advancements and prominent advantages in the field of drug delivery. The advantages have extended into the products approved by FDA in the last two decades. Recently, a lot of research has emphasized the utility of these nanoparticles in the field of diagnostic imaging and basic/translational research. Magnetic nanoparticles have unique properties such as superparamagnetism, sensitivity to temperature, and high saturation index properties with enhanced colloidal stability (physical/chemical), making them particularly useful for diagnostic applications. Furthermore, these particles are used in organic synthetic schemes and have widespread utilization in the in the field of biotechnology and medicine. Iron oxide nanoparticles are the most common magnetic nanoparticles used in the field of diagnosis and drug delivery. Magnetic nanoparticles coated with polymers or targeting moieties could direct the influence of the payload to the target regions; they are specifically used for MRI as contrast agents in the deciphering of the specific location and help with the isolation of the component for further diagnosis. They are also used in some of the widespread diagnostic biomedical applications (Niemirowicz et al., 2012). Magnetic nanoparticles are noninvasive imaging modalities, used traditionally for disease imaging, but recently improved versions of these nanoparticles are used for multimodal imaging. For efficient diagnostic applications, the surface chemistry of these modalities must be designed and optimized for optimal performance of the nanoparticles.

5.2.1 MAGNETIC RESONANCE IMAGING AND OTHER APPLICATIONS Iron oxide nanoparticles are an excellent choice for cost effective imaging in a clinical setting. With the conventional modalities available, they show potential effects and visualize small tumors and metastases with high resolution offered by the imaging technique. The high spatial resolution and sensitivity of this technique could diagnose cancers, and also be of great use for surgical procedures or chemotherapeutic intervention by imaging the decrease in volume of the tumor size. Because of the enormous vasculature and EPR effect, the iron oxide nanoparticles are accumulated in enhanced quantities and therefore increase the contrast and better imaging for the clinicians in making judgment for the infinitesimal tumors (Ronak and Tamara, 2016). Highly aggressive cancers like

5.2 Diagnostic Applications Using Magnetic Nanoparticles

pancreatic cancer need to be detected in the early stages for an effective treatment. For enhancing the efficiency for effective diagnosis for cancer, Rosenberg et al. prepared diagnostic magnetic nanoparticles for imaging of the cancer. Biodegradable nanoparticles using iron oxide incorporated in the recombinant human serum albumin were manufactured. The prepared nanoparticles were surface-modified with glycosylated peptides targeting the galectin receptors on the pancreatic tumors. Since the tumor environment is conducive to the targeted nanoparticles, there are high chances of accumulation of the nanoparticles and thereby enhancement of contrast of MRI imaging techniques (using a hand held gamma camera). Multimodal imaging has recently gained importance in visualizing the in vivo system by combining with surface coating the SPIONs with fluorophores (optical imaging) and radionuclides (nuclear imaging). The excellent combination of multimodal imaging produces a better fluorescence and MRI contrast for a better judgement in terms of the tumor surgical procedures or the treatment options, furthered by combining positron emission tomography with MRI for better spatial resolution and anatomical imaging of the in vivo biological system (Stephen et al., 2011). Luo et al. experimented with the targeting moiety lactoferrin to target the superparamagnetic nanoparticles as MRI contrast agents for rat glioma in the brain. The group manufactured magnetic nanoparticles by thermal decomposition method, further encapsulated in the copolymer of novel amphiphilic poly(aminoethyl ethylene phosphate)/poly(L-lactide). The polymeric nanoparticles are surface-enhanced by conjugation with lactoferrin to efficiently target the tumor lesions. The targeted nanoparticles have long-lasting tumor targeting and enhanced contrast for MR reading, and showed great promise for the better enhancement of MRI imaging (Luo et al., 2016). Shevtsov and their team identified epidermal growth factor (EGF) as a targeting moiety for enhanced contrasting of the magnetic nanoparticle MR agents. The SPIONs were synthesized by precipitation technique and further conjugated with EGF. The prepared nanoparticles were tested in vitro on C6 glioma cells and showed no toxicity, further tested in vivo and showed high contrast MR in the EGF receptoroverexpressed C6 glioma cells in the brain. High contrast MR imaging has great potential using SPION-EGF conjugates for targeting (Shevtsov et al., 2014).

5.2.2 CANCER THERANOSTICS Colloidal carriers have been applied for medical evaluation and treatments like imaging, diagnostic, and therapeutic applications. Diagnosis and therapy in combination defines the new term “theranostics”; the use of magnetic nanoparticles for therapy and diagnosis defines the term cancer theranostics. The basic advantage of cancer theranostics is to detect the cancer and enhance the targetability of the payload to the cancer, and also decrease the systemic toxicity associated with the use of cancer nanomedicine.

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5.2.3 MAGNETO ACOUSTIC TOMOGRAPHY To date, magnetic nanoparticles have been used for MRI. New imaging modalities are being investigated to see the in vivo distribution more accurately. The magneto acoustic tomography (MAT) method uses the magnetomotive force to induce ultrasound in the magnetic nanoparticle, to induce ultrasound imaging of the in vivo distribution of them with ultrasound imaging resolution. To detail the mechanistic point of view, the secondary effect induced by the short-pulsed field is used to characterize the ultrasound images. Photoacoustic imaging uses the principle of optical absorption properties of detection in the presence of nanoparticles. The nanoparticles absorb the applied laser pulses, and the energy converts to heat; further acoustic waves generated are received by ultrasound imaging receivers. Leo et al. used short-pulsed magnetic field for in vivo imaging of the magnetic iron oxide nanoparticles. The short-pulsed field applied to magnetic nanoparticles produced acoustic vibrations moving in all directions throughout the medium, captured by ultrasound transducer. They used superparamagnetic nanoparticles in the size range of 1020 nm for better magnetic properties, and also have a better colloidal stability. The animal model was mice, and the ultrasound images had good resolution because of short-pulsed magnetic field. Consequently, MAT would be a good technique for imaging soft tumor tissues with a greater imaging depth (Leo et al., 2016).

5.2.4 HYPERTHERMIA Magnetic nanoparticles have unique properties in drug delivery and hyperthermia. Under the influence of alternate current field, magnetic nanoparticles are excited and the electromagnetic energy converts to heat energy, thereby generating high localized temperature, which could be utilized for killing cancer cells or microbes. Magnetic hyperthermia was an approved therapy for brain tumor treatment in Europe, but due to the gap in the optimization, this therapy is still in the budding stages of utilization in other parts of the world. Furthermore, hyperthermia is recently connected to diagnostic imaging to improve the utility of these nanoparticles. The extreme side effect of the hyperthermia is overheating of the tissues, with a lot of discomfort. With diagnostic imaging in place, the localized environment would be visualized with concentration of delivered nanoparticles to the region. Optimization could be an easier task by combining the two technologies and harness the advantages. Emphasis in research is being placed on the optimization of nanoparticles, and the excitation of the AC field to optimize the treatment options (Eric et al., 2016). Furthermore, with the introduction of so-called smart polymers, which are responsive to external stimuli like changes in temperature or pH gradient, technological advance in the field of hyperthermia is geared towards the combinatorial advantages of the smart polymers with the encapsulation of magnetic

5.2 Diagnostic Applications Using Magnetic Nanoparticles

nanoparticles. Medeiros et al., in their review on stimuli responsive magnetic systems, specified the use of smart polymers in combination with magnetic systems and their use in the biomedical applications like MRI, drug delivery, chemotherapy, and hyperthermia, etc. Nanobiomagnetism defines the purpose of the magnetic nanoparticles as intervention with medicine in their process or biomedical applications. The term “nano” does not interfere with the normal functions and the particles under the influence of nanobiomagnetism can traverse through the intricate systems of the body. Recently, the polymer shell used to encapsulate nanomaterials is surface-functionalized to target specific regions of the body, serving the dual purpose of targeting and treatment options. Magnetic biomaterials should be biocompatible, nontoxic, and should not make any aggregates in the living system. The use of magnetic nanoparticles as hypothermic agents is as a nanoscale heater to kill the malignant cancer cells by cooking the tumor cells with heating. It is recognized as a therapeutic intervention for the treatment of malignant cancers. The concept of hyperthermia depends on the magnetic particles and their properties. The magnetic nanoparticles either injected IV or at the tumor site, under the influence of the magnetic field, produce heat in the cells. Some magnetic nanoparticles produce heat by magnetic hysteresis loss, but could be negligible. Cancer cells undergo cell necrosis, carbonization, and cell coagulation at a temperature greater than 46 C, compared to the normal cells. Accordingly, magnetic nanoparticles are becoming an integral part of the customized cancer therapy (Medeiros et al., 2011; Wei et al., 2012).

5.2.5 CHALLENGES AND FUTURE DIRECTIONS The first challenge in commercialization is the reproducibility of nanoparticles in the same size and shape with established biocompatible properties. The other challenge is the toxicity of the nanoparticles. Magnetic nanoparticles, when injected intravenously, are aggressively cleared by the reticulo endothelial system and deposit the nanoparticles in the liver, kidney, spleen, etc. This has elevated the toxicity concern and presents a big hurdle in the progress of this technology from lab to clinic. After administration, many intricate factors play a role which dictate the fate of the nanoparticles for their efficient use in the system. Furthermore, in vivo brain imaging was always elusive because of the BBB, which acts a bulwark and prevents the entry of the nanoparticles. Recently, emphasis was put onto targeting moieties have received promising results. Future directions are to increase in optimization of the nanoparticles, their biocompatibility, and targeting approach to obtain good diagnostic tools for all available diseases (Cheng et al., 2016). Cost-efficient generic medicines will need a lot of expertise to get the product on the shelf, with a lot of variables to meet the bioequivalence option to achieve approval from the FDA. With better understanding

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of the molecular mechanism and the transport options of receptors, magnetic nanoparticles could be harnessed to see a better future in the coming years.

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Nanoparticles influence in skin penetration of drugs: In vitro and in vivo characterization

6

Camila N. Lemos, Francieli Pereira, Luciana F. Dalmolin, Camila Cubayachi, Danielle N. Ramos and Renata F.V. Lopez School of Pharmaceutical Sciences of Ribeirao Preto, University of Saõ Paulo, Ribeirao Preto, SP, Brazil

CHAPTER OUTLINE 6.1 Introduction .....................................................................................................187 6.2 Skin Structure..................................................................................................189 6.2.1 Skin Appendages............................................................................192 6.3 Mechanisms and Routes of Nanoparticles Skin Penetration ................................192 6.4 Characteristics of Nanoparticles for Drug Skin Penetration.................................196 6.4.1 Type of Nanoparticle .......................................................................196 6.4.2 Size and Surface Area.....................................................................199 6.4.3 Charge...........................................................................................200 6.4.4 Shape............................................................................................201 6.5 Physical Methods to Enhance Nanoparticle Skin Penetration ..............................201 6.5.1 Iontophoresis .................................................................................201 6.5.2 Electroporation...............................................................................205 6.5.3 Microneedles .................................................................................207 6.6 Experimental Techniques for Studying Nanoparticle Skin Penetration..................208 6.6.1 In Vitro Studies ..............................................................................209 6.6.2 Ex Vivo Skin Penetration Experiments ..............................................223 6.6.3 In Vivo Skin Penetration Experiments ...............................................224 6.7 Conclusion ......................................................................................................229 References .............................................................................................................229

6.1 INTRODUCTION The skin is the organ with the largest surface area of the human body, and the largest interface between it and the external environment. The skin’s main Nanostructures for the Engineering of Cells, Tissues and Organs. DOI: http://dx.doi.org/10.1016/B978-0-12-813665-2.00006-5 © 2018 Elsevier Inc. All rights reserved.

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function is to establish a protective complex barrier against external agents, including biological, physical, and chemical threats. It also protects the body against the loss of water and other substances. Due to skin’s easy accessibility and large surface area, it is a potential site for noninvasive and painless drug administration. Drugs administered over the skin can have a topical or transdermal effect. For transdermal treatment, the skin is only a route of administration. In this way, no interactions within skin layers should occur, guaranteeing a maximum permeability so the drug can reach the systemic circulation at an effective concentration. The skin is, however, the site of action for topical treatments. Therefore, topically administered drug should remain in the skin with a minimal amount reaching systemic circulation. Either way, the drug needs to cross the outermost layer of the skin, the stratum corneum (SC), in order to be effective. The SC maintains the water content in the body in variable climatic conditions, limits the absorption of toxic substances from the environment, and, consequently, also hinders the entry of drugs into the skin. Therefore, topically administered drug needs to have adequate physicochemical characteristics to diffuse through the SC and reach, in therapeutic concentrations, the viable layers of the skin, where the majority of cutaneous diseases are located. To increase skin penetration of drugs, several strategies are employed. Among these is the use of chemical penetration enhancers, being the simplest strategy, which causes a temporary and reversible disorganization of the lipid bilayers of the SC. The use of physical methods, such as iontophoresis and microneedles, are other strategies that are gaining prominence, mainly due to advances in biotechnology and bioinformatics. Specifically, advances in biotechnology enable peptides and proteins to be produced with high biological activity, high specificity and low toxicity. These substances are normally unstable in biological fluids and need to be administered by invasive approaches, usually poorly tolerated by patients. The administration of these molecules through noninvasive strategies to the skin would certainly be a very welcome and advantageous approach. However, most peptides and proteins have high molecular mass, hydrophilic character, and are often ionized at physiological pH. These characteristics make them extremely poor candidates for passive diffusion through the lipid-rich intercellular space of the SC. Application of low electric current (iontophoresis) or microneedles in the formulation can “force” the penetration of these drugs through this barrier, without significantly compromising skin physiology. The ability to overcome the SC barrier conferred by these physical methods makes them strategies of potential interest for increasing cutaneous penetration of hydrophilic and high molecular weight drugs. In addition to the chemical and physical methods that promote drug penetration, the development of nanoparticulate delivery systems can contribute significantly to the controlled skin penetration of substances. Polymeric nanoparticles, liposomes, solid lipid nanoparticles (SLNs), dendrimers, among others, are

6.2 Skin Structure

modern drug delivery systems that can protect drugs against degradation, sustain their release, and also, with the aid of physical methods, penetrate the skin and enable the release of high concentrations of drugs at specific sites within the skin. Targeting the release of the drug to specific sites of the skin, minimizing systemic effects, is one of the great pursuits of topical administration, both for dermatological and cosmetic purposes. The ability to design nanoparticles with specific characteristics has increased the chances of achieving these feats. Understanding how the characteristics of nanoparticles influence skin interaction and release of the drug, in order to enable the development of nanoparticles optimized to treat specific skin problems, is the greatest challenge today. The aim of this chapter is to provide a base for understanding the role of nanoparticulate delivery systems in skin penetration of drugs. In this way, the skin’s main barriers to drugs penetration will be briefly described, as well as the basic principles of physical methods associated with nanoparticles used in increasing drug penetration. The influence of physicochemical characteristics of nanoparticles in drug skin penetration and the experimental methods used to evaluate this influence will be presented and discussed.

6.2 SKIN STRUCTURE The skin can be divided into three layers: epidermis, dermis, and hypodermis; the presence of appendages along these layers should be considered (Junqueira and Carneiro, 1999; Kanitakis, 2001; Menon, 2002) (Fig. 6.1). The first skin layer is the epidermis, crucial for defensive purposes. Human epidermis thickness can vary, depending on the body part, but usually it is thicker in palm hands and sole of the feet, while eyelid epidermis is thinner. Some authors describe its thickness around 1 mm, but it can vary according to the literature (Kanitakis, 2001; Sandby-Moller et al., 2003). The epidermis can be divided into two sublayers: the SC and the viable epidermis (Elias, 1996; Junqueira and Carneiro, 1999; Kanitakis, 2001; Menon, 2002). The epidermal cells are in continued renewal and are hierarchically organized, differentiating from the deeper layers to the SC, the outermost skin layer. The SC has lipophilic characteristics and is the main skin barrier against drug penetration and other foreign agents, due to its particular features. Interfacing the environment, SC shows a 613 μm thickness (Johnson et al., 1997; Cevc, 2004), depending on the hydration state and the region of human body. It is composed of 15 layers of dead, dehydrated protein-enriched cells and highly specialized keratinocytes, called corneocytes (Elias, 1988; Cevc, 2004). Corneocytes are polygonal, anucleated cells, surrounded by a lipid matrix and forming a very organized structure, arranged as a “brick and mortar” (Plewig et al., 1983; Elias, 1988, 1996) (Fig. 6.1B). The main lipids are ceramides, free

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FIGURE 6.1 Schematic cross-sectional representation of skin anatomy. (A) Structural organization of corneocytes as the “brick and mortar” model, showing its hexagonal shape, organized in clusters (3D perspective). (B) Simplified representation of lipid matrix configuration, composed of phospholipids self-assembled in bilayers.

fatty acids, cholesterol, and sterol or wax esters, mostly with nonpolar characteristics (Baroli, 2010; Zhang et al., 2015a). Cells are connected by corneodesmosomes and packed with keratin filaments, providing a highly water impermeable layer (Kubo et al., 2013). Proteins, such as small proline-rich involucrin and transglutaminases, are also present in the SC, playing a role in the structural envelope formation and cellcell interactions. However, there are some hydrophilic regions in the SC near to the polar head groups of the amphiphilic lipids that

6.2 Skin Structure

compose the lipid matrix (Fig. 6.1C). These regions are known as aqueous porous, which have a diameter varying between 0.4 and 3.6 nm (Cevc, 2004). This structural set ensures that SC fulfills its barrier function with selectivity, avoiding water loss, and high molecular or charged compound penetration (Johnson et al., 1997; Mathes et al., 2014; Ba¨sler et al., 2016; Planz et al., 2016). The viable epidermis is an avascular but dynamic and self-renewing tissue in which SC cell surface loss, i.e., desquamation, is balanced by the cell growth of the epidermis (Bouwstra and Ponec, 2006). Epithelial cells, or keratinocytes, are divided, from the inner to the outer, into stratum basale, stratum spinosum, stratum granulosum and SC (Menon, 2002). Each layer is defined by position, shape, morphology, and state of differentiation of keratinocytes. In contrast to the SC, the viable epidermis contains a high amount of water, which determines its hydrophilicity (Baroli, 2010). Keratinocytes communicate and adhere to each other and/or to the underlying basement membrane through various bonds. Specialized connections, such as hemidesmosomes, desmosomes, gap junctions, adherens junctions, and tight junctions, are responsible for the merger and mechanical resistance of epithelial cells. They also mediate signals from the neighboring cells or from the extracellular matrix to the cytoskeleton and the cytoplasm of the cell (Proksch et al., 2008). Therefore, keratinocytes are very active and secrete various cytokines and growth factors that regulate the proliferation and differentiation of other cell types, thus responding to a wide variety of environmental stimuli to maintain homeostasis and control tissue damage. Basal cells of the viable epidermis are highly proliferating cells that form the stratum basale and are connected to the dermis by proteins such as collagen type IV, laminin, and proteoglycan. Protection against UV-radiation is given by the melanocyte cells. Immune cells, such as Langerhans cells, are also presented (Ba¨sler et al., 2016). The second skin layer, dermis, is highly vascularized by blood and lymph vessels which permeate collagen and elastin fibers that compose this skin layer. Collagen, produced by fibroblasts, is the most abundant extracellular component and represents about 80% of the dry weight of the skin, giving tensile strength to the dermis. The elastin network is responsible for the elasticity and represents about 2%4% of the extracellular matrix in skin protected from sunlight (Junqueira and Carneiro, 1999). Several adjoining structures, such as hairs, nails, and sebaceous and sweat glands are embedded in the dermis and are supported by collagen (Moser et al., 2001). Moreover, immune cell such as mast cells, lymphocytes, T and B cells, and macrophages are also included in this layer. Depending on the location of the body, dermis thickness varies from 0.3 to 5 mm. The transport of substances in this area is rapid, and can reach the systemic circulation through the blood vessels. The dermis’ main function, besides nutrition and oxygen supply, is skin mechanical strength and flexibility (Baroli, 2010; Wong, 2014; Planz et al., 2016).

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The hypodermis or subcutaneous tissue is a layer beneath the dermis, composed mainly of adipocytes and fibroblasts, storing fat and contributing to thermal isolation (Ba¨sler et al., 2016; Planz et al., 2016). Fibrous connective tissue is part of the constitution of hypodermis.

6.2.1 SKIN APPENDAGES Skin appendages, such as sweat glands, sebaceous glands, and pilosebaceous units, are structures adjacent to skin that have particular functions, such as contractility, heat loss, and sensation. They have a special role in nanoparticle penetration of the skin, which will be discussed further in this chapter. Considering skin extension, skin appendages occupy a small area, about 0.1% of skin surface (Scheuplein, 1967; Otberg et al., 2004; Baroli, 2010; Ma and Hadzija, 2012). Sweat glands are responsible for the thermoregulation and excretion of body wastes: a hypotonic aqueous solution of organic acids, carbohydrates, aminoacids, nitrogenous substances, electrolytes, and vitamins. They are tubular glands and are located in the dermis or hypodermis, extending to the SC (Baroli, 2010). Sebaceous glands excrete sebum, a mixture of lipophilic substances that can vary in composition, but mostly composed of fatty acids, cholesterol derivatives, triglycerides, waxes and squalene (Ba¨sler et al., 2016; Vogt et al., 2016). The pilosebaceous unit is a hair follicle associated with the sebaceous gland. They extended from the dermis until the body exterior, providing a small orifice in the SC that may serve as a penetration route. Besides a penetration pathway, hair follicles can be considered reservoir structures for nanoparticles, but they are heterogeneously distributed in the skin surface (Baroli, 2010).

6.3 MECHANISMS AND ROUTES OF NANOPARTICLES SKIN PENETRATION SC organization, structural integrity, and an acid environment with pH gradient are all crucial for the maintenance of skin functions. Its nanoporosity contributes to the limited penetration of molecules bigger than 500 Da and its composition and organization impairs the penetration of ionized molecules (Baroli, 2010; Vogt et al., 2016). Metabolic enzymes in viable epidermis and dermis, in addition to presence of immunological cells, also contribute to external substance neutralization or degradation (Baroli, 2010). Together, these skin characteristics greatly hinder drug penetration into the skin and can be overcome by a delivery system. For greater permeation into or through the skin, the drug needs to surpass the SC. Drug permeation through this layer is generally the slowest phase of the permeation process and therefore is the limiting step for substance entrance in the skin (Darlenski and Fluhr, 2012).

6.3 Mechanisms and Routes of Nanoparticles Skin Penetration

The main mechanism of substances penetration in the skin is passive diffusion, which involves drug partition and diffusion by different skin layers, following a concentration gradient until the blood and lymphatic vessels. Therefore, passive drug penetration through the skin can be described by the Fick’s equation (Eq. (6.1)), which expresses the molecule steady-state flux (J) as a function of the area and the drug concentration gradient (Higuchi and Higuchi, 1960): J5

Q Dcs;m cv 5 3 At h cs;v

(6.1)

where J is the flux of the drug through an area A of the skin, Q is the amount of drug entering the skin, t is the time, D is drugs diffusion coefficient through the SC, Cs,m is its solubility in the skin, h is the thickness, Cv is the concentration of drug and Cs,v is the solubility of the drug in the formulation. According to Eq. (6.1), increasing the exposure time and the area of exposition can increase the drug permeation rate. Nanoparticulate delivery systems, which sustain drug delivery and have a high surface area, may take advantage of Fick’s law to increase drug skin penetration (Mu¨hlen et al., 1998). Nanoparticles can also carry high drug concentrations and increases the saturation degree, which may increase the thermodynamic activity (the Cv/Cs,v) and improve drug flux through the skin (Mitragotri et al., 2011; Schwarz et al., 2013; Scalia et al., 2015; Kassem et al., 2017). Finally, the partition coefficient SC/vehicle (Cs,m/Cs,v) is another variable that can be improved by using nanoparticles. For instance, encapsulation of the hydrophilic doxorubicin (DOX) hydrochloride in SLNs improved the lipophilicity of the drug, facilitating DOX penetration into the skin (Huber et al., 2015). Therefore, the physicochemical properties of the drug, as well as the drug delivery system can influence drug skin diffusion through skin as well as the penetration route. Drug passive diffusion through the SC can occur via three different potential routes (Fig. 6.2): (1) intercellular route, with partitioning into the lipid matrix (2) transcellular route (3) appendageal route The intercellular route includes the partition and diffusion of a substance around the corneocytes, in the lipid matrix. It prevents multiple partitions from the more hydrophilic corneocytes into the lipid intercellular layers in the SC, as happen with the transcellular route. The transcellular route, therefore, involves drugs crossing through the corneocytes and the lipid matrix (Scha¨tzlein and Cevc, 1998). In general, the partition through these routes happens when the molecule exhibits a log octanol/water partition coefficient (log P) of 13 (Benson, 2005; Alexander et al., 2012). A parabolic relationship between the log P and the penetration rate is often observed: a drug with a low log P, i.e., highly hydrophilic,

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FIGURE 6.2 Schematic representation of main routes for nanoparticles diffusion through the skin: intercellular route, transcellular route and appendageal route.

has difficulty partitioning in the lipid matrix of the SC and exhibit a low permeability; highly lipophilic drugs, with a high log P, have difficulty partitioning out of the SC to the viable epidermis (Guy and Hadgraft, 1988). The intercellular route is considered the main pathway for several drugs’ diffusion through the skin (Hadgraft, 2001). However, high molecular weight drugs and rigid micro- or nanoparticles do not use neither intercellular nor transcellular route to penetrate the skin. Specifically, the lack of skin penetration in these drugs or delivery systems is related to the organization and composition of the SC. Besides keratin filaments, desmosomes, and corneodesmosomes that prevent exogenous substances penetration, the intercorneocyte space is filled with a lipid matrix that is composed of phospholipids self-assembled in bilayers (Fig. 6.1C), which are covalently bound with the corneocytes. These bilayers have repeated hydrophilic and lipophilic areas within the headhead and tailtail regions, respectively, assembling an orthorhombic lateral packing, which confers impermeability to SC. Other three-dimensional (3D) configurations are possible, but they are not related to a healthy SC (Bouwstra and Ponec, 2006; Baroli, 2010). In the orthorhombic configuration of the lipid matrix, the distance between hydrophilic regions delimited by lipid heads in a single bilayer is 2.8 6 1.3 nm, and the distance of the more fluidic areas in the tailtail regions can be estimated between 5.42 and 6.94 nm (Bouwstra and Ponec, 2006). To penetrate through these layers, particles need to be smaller than this distance; as a consequence microorganisms and most particles do not cross the SC through the intercellular route, since they present larger sizes. Corneocytes present a hexagonal shape, diameter of 40 μm and thickness of 0.30.8 μm. They are packed in 1015 horizontal layers (Plewig et al., 1983)

6.3 Mechanisms and Routes of Nanoparticles Skin Penetration

and also arranged in longitudinal structures, called clusters. These clusters are formed by the agglomeration of 12 corneocytes and are related to the skin surface until the basal lamina. The cluster boundary area is defined as furrows, which can be considered as areas of low penetration resistance delimiting the intercluster penetration pathway (Cevc, 2004; Baroli, 2010). Therefore, nanoparticles can distribute and accumulate within these areas (Alvarez-Romań et al., 2004; Ku¨chler et al., 2009; Huber et al., 2015). Besides the furrows, nanoparticles can also benefit from the skin appendages to penetrate the skin (Patzelt and Lademann, 2013). The hair follicle can extend deep into the skin tissues, up to 2000 μm (Toll et al., 2004). The funnel shape of the follicles increases the surface area and disrupts the epidermal barrier toward the lower parts, acting as an efficient reservoir for drugs and nanoparticles, since substances stored therein can continuously diffuse to the surrounding spaces, cross the capillary walls, and reach the blood system (Rancan and Vogt, 2014). However, the follicular barrier is similar to the SC at the upper part and presents tight junctions in the lower part (Langbein et al., 2002), hindering nanoparticles penetration in the viable epidermis. Furthermore, follicular density varies in different areas of the body (Jacobi et al., 2007; Lademann et al., 2010), which also affects the magnitude of the drug penetration. The penetration of nanoparticles through the transfollicular pathway is related to particle size (Patzelt et al., 2011; Toll et al., 2004) and to the type of the nanoparticle (Patzelt et al., 2011). The way these characteristics influence nanoparticle skin penetration will be addressed throughout this chapter. It is also important to consider the role of sebaceous glands in the appendage route of penetration. The sebaceous glands open into the hair follicles, releasing sebum, which provides a lipid-enriched environment. Good dispersion into the sebaceous unit is interesting to store more particles therein and sustain drug release (Fang et al., 2014). Studies that assess the follicular deposition of drugs from different vehicles showed that the use of microparticles (Lademann et al., 1999; Gelfuso et al., 2015) and nanoparticles (Shim et al., 2004; Lademann et al., 2007, 2015) increase the amount of the administered drug in the follicles, a suggestion that nanoparticulate delivery systems enter the skin especially through this pathway and accumulate in it. Therefore, to target drug delivery to skin appendages, formulations should be developed concerned with the body region of the treatment, maximum particle size, and the interaction between nanocarrier and the lipid-enriched environment. Due to the importance of hair follicles as a route of nanoparticle penetration, and to the small skin surface area occupied by them (typically 0.09% 1.3% of the total skin area) (Scheuplein, 1967; Otberg et al., 2004; Baroli, 2010; Ma and Hadzija, 2012), physical methods have been studied to improve their penetration and will be further discussed in details in this chapter (Section 6.5).

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6.4 CHARACTERISTICS OF NANOPARTICLES FOR DRUG SKIN PENETRATION One of the major challenge in nanoparticle skin delivery is finding the right balance between enhancing the penetration of drugs while ensuring sufficient retention to maintain a sustained drug delivery. Therefore, a variety of nanoparticles parameters has been studied to improve drug delivery. Considering the mechanism of drug penetration in the skin and the routes of penetration, there are several nanoparticles characteristics that may influence drugs penetration. Physicochemical interactions between nanoparticles and the skin, kinetics, and thermodynamic exchanges can all influence drug skin penetration from nanoparticles. Size, shape, surface properties and charge, and type of nanoparticles (lipid or polymeric, for instance) appear to be the properties of nanoparticles that most influence the penetration of drugs into the skin and will be discussed forthwith (Table 6.1).

6.4.1 TYPE OF NANOPARTICLE Nanoparticles of different types present a variety of composition and flexibility, characteristics that can influence their interactions with the skin and the mechanism of penetration (Zhang et al., 2016). Lipid nanoparticles, for instance, may fuse when in contact with the SC, whereas polymeric nanoparticles should likely maintain their initial characteristics and accumulate in the skin appendages. Liposomes, which are phospholipid vesicles in which a well-defined aqueous nucleus is present within one or multiple lipid bilayers, have a composition similar to the lipid matrix of the SC and cell membranes. In this sense, they interact with SC and destabilize the lipid matrix by mixing or fusing with it and thereby increasing the drug flux in the skin (Kirjavainen et al., 1996, 1999; El Maghraby et al., 2008). They can also serve as a drug reservoir in the skin, sustaining the release of drugs such as antibiotics, corticosteroids, and retinoic acid, among others (Pierre and Costa, 2011; Clares et al., 2014). Liposomes can also increase the hydration of SC by hindering the loss of transepidermal water, increasing drug penetration (Cevc and Blume, 1992; Hofland et al., 1995; El Maghraby et al., 2008). Deformable liposomes, also called elastic or ultraflexible liposomes, respond to external stress by shape deformation, which allows them to squeeze and to penetrate deep through the intercellular regions of the SC (Qiu et al., 2008; Pandit et al., 2014; Zhai and Zhai, 2014; Ashtikar et al., 2016). The main penetration mechanism of these liposomes seems to be the transepidermal osmotic gradient (Cevc and Blume, 1992; Benson, 2009). Due to their flexibility and differences in skin interaction, deformable liposomes are shown to increase skin penetration of some drugs, when compared to the conventional liposomes (Honeywell-Nguyen

6.4 Characteristics of Nanoparticles for Drug Skin Penetration

Table 6.1 Specific Nanoparticle Characteristics That Influence Cutaneous Penetration of Drugs Characteristics Size/surface area

Charge

Type

Shape

Functional groups

Influence on Drug Skin Penetration Smaller particles have a larger surface area-to-volume, increasing the interaction with skin surface and serving as a drug depot in hair follicles, further sustaining drug release Negative nanoparticles are generally repelled by the negatively charged skin domains (carboxylate groups). Positively charged nanoparticles presents higher affinity to the stratum corneum Lipid nanoparticles may fuse when in contact with the stratum corneum, whereas polymeric nanoparticles would likely maintain their initial characteristics and accumulate in the skin appendages Rod-like shaped nanoparticles, compared to spherical and triangle-shaped nanoparticles, seem to facilitate drug penetration through the skin Nanoparticle surface modification, such as cell-penetrating peptides, lipid spacers, oleic acids, and pyrrolidinium lipids can increase drug permeability and retention time in addition to reducing systemic side effects

References Zhai and Maibach (2001) Toll et al. (2004) Vogt et al. (2006) Adib et al. (2016) Blank (1939) Bernfield et al. (1999) Wu et al. (2009) Honary and Zahir (2013) Uchechi et al. (2014) El Maghraby et al. (2008) Baroli (2010) Desai et al. (2010) Lademann et al. (2011) Prow et al. (2011) Clares et al. (2014) Zhang et al. (2012) Dasgupta et al. (2014) Fernandes et al. (2015) Tak et al. (2015) Shah et al. (2012) Wang et al. (2013) Desai et al. (2013) Mohammed et al. (2016) Boakye et al. (2017)

et al., 2003, 2004; Abd et al., 2015). Some recent studies combined the use of elastic liposomes with transdermal vaccination for enhancement in drug penetration and immune response (Tyagi et al., 2015, 2016). Other mechanisms proposed for the increase in drug penetration caused by deformable liposomes are: disorganization of the intercellular lipid arrangement and corneodesmosomes, and intracellular keratin degradation (Subongkot et al., 2014). SLNs are other types of lipid nanostructures greatly studied for topical delivery of drugs in the skin. They are nanosystems composed of solid lipids at room

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temperature and present a structure similar to that of nanoemulsions, except the inner liquid lipid of the nanoemulsion is replaced by a solid lipid in the SLN (Wissing et al., 2000). One variation of SLN is nanostructure lipid carriers (NLCs), which have a core less crystalline as the one presented by SLN, since NLC have a mix of solid and liquid lipids. SLNs provide a reduced mobility of the drug into the lipid matrix, which enable sustained drug delivery (Westesen et al., 1997). When applied in the skin, SLNs and NLCs show adhesiveness, occlusion, and skin hydration properties. These properties, when combined, are responsible for the formation of a hydrophobic monolayer film on the skin, hampering transepidermal water loss (TEWL) (Desai et al., 2010; Gokce et al., 2012; Al-Amin et al., 2016; Roberts et al., 2017). These events consequently reduce the corneocyte packing and open some intercorneocyte gaps, facilitating the skin permeation of the encapsulated drug (Wissing and Muller 2002; Desai et al., 2010). For instance, ropivacain-loaded NLC applied in vivo, on mice skin, showed to be in close contact with the surface of the skin and formed a thin film adhered to it, which promotes occlusion and increased ropivacain penetration (Zhai et al., 2014). The influence of SLN and NLC over the skin penetration of psoralen was evaluated and compared with the drug permeation from a simple emulsion and from an aqueous solution (Fang et al., 2008). The lipid nanoparticles enhanced drug permeation and controlled its delivery, with NLCs showing higher improvement. The best performance presented by the NLC was explained based on this nanoparticle’s higher occlusive power and lower viscosity, when compared with SLNs; the lower the viscosity of the vehicle, the faster is the drug release. However, the smaller size of NLC developed (B200 nm) compared to SLN (B300 nm) may also explain the higher occlusion caused by NLC application (Fang et al., 2008). Nanoemulsion, liposome, and SLN characteristics were compared in a study involving skin penetration of a lipophilic drug, retinyl palmitate (Clares et al., 2014). Deep skin layers were reached when the drug was administered from the nanoemulsion, followed by liposomes, with SLNs holding the drug in the most superficial layers. Different interactions of these drug delivery systems with the SC certainly influenced the extension of retinyl palmitate skin penetration. Different from lipid nanoparticles, rigid polymeric nanoparticles are not generally composed of components present in the SC and do not disperse easily in the lipid matrix. Studies have indicated that polymeric nanoparticles primarily accumulate in the skin surface, furrows, and hair follicles (Baroli, 2010; Prow et al., 2011). Nanoparticle accumulation in skin furrows and appendages offers a shortcut to releasing drugs directly into deep skin layers (Lademann et al., 2011). Furthermore, they can form a concentrated drug reservoir, that gradually releases the drug deep in the skin (Desai et al., 2010; Du et al., 2016).

6.4 Characteristics of Nanoparticles for Drug Skin Penetration

Several studies have described the increase of drug skin penetration after encapsulation in polymeric nanoparticles. Examples of drugs that had skin permeability enhanced by these nanoparticles are: antiretrovirals, hormones, nonsteroidal antiinflammatory, and antitumoral drugs (Luengo et al., 2006; Sheihet et al., 2008; Venuganti and Perumal, 2008; Hasanovic et al., 2009; Tomoda et al., 2012). Besides nanoparticle composition, surface modification can be a strategy to enhance nanoparticle interactions with skin and, as a consequence, drug penetration. The nanoparticle surface modifications that showed to increase drugs skin permeability were performed mainly with cell-penetrating peptides (Hsu and Mitragotri, 2011; Wang et al., 2013), lipid spacers (Shah et al., 2012), oleic acids (Desai et al., 2013), and pyrrolidinium lipid (Boakye et al., 2017).

6.4.2 SIZE AND SURFACE AREA There are plenty of studies demonstrating the influence of the size and surface charge of nanocarriers in biological effects. They can change, for instance, drug pharmacokinetics, cellular uptake, and toxicity (Gatoo et al., 2014; Bhattacharjee, 2016). Particle size is directly involved with skin penetration through the SC barrier. Particles sized below 10 nm may have a positive influence in drug transdermal absorption via intercellular route, while nanomaterials smaller than 1 nm may act like a molecule (Baroli, 2010). However, conventional rigid nanoparticles have sizes ranging from about 50 to 200 nm and, as a general rule, do not penetrate a healthy SC, but rather enhance the drug diffusion through the barrier, increasing the drug concentration gradient (Baroli et al., 2007). Furthermore, size is a property of nanoparticles strictly related to the surface area. Smaller particles have a larger surface area-to-volume ratio, making the nanomaterial surface more reactive to interact with skin surface in different mechanisms, depending on the nanoparticle nature. Polymeric particles up to several micrometers (0.756.0 μm) are reported to enter the hair follicles, with 1.5 μm microparticles showing a maximum penetration depth of .2300 μm (Toll et al., 2004). Polymeric nanoparticles made with poly(lactic-co-glycolic acid) (PLGA), with different sizes, revealed that increasing size from 122 to 643 nm resulted in significantly deeper penetration in hair follicles (122 nm , 230 nm , 300 nm , 470 nm , 643 nm), while particles larger than 860 nm remained closer to skin surface (Patzelt et al., 2011). Silica nanoparticles, on the other hand, showed low penetration depths when presenting small (300 nm) and large (920 and 1000 nm) size, with the deepest penetration showed as 646 nm nanoparticles (Patzelt et al., 2011). Different particle sizes of the same type of nanoparticle, therefore, seem to target different regions of the hair follicle, such as the infundibulum or the bulge. Moreover, small nanoparticles of 40 nm

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were observed in the hair follicles and also through the perifollicular dermis (Vogt et al., 2006). Gold nanoparticles, with 15, 102 and 198 nm size, showed size-dependent permeation through rat skin. Maximum permeation was observed for 15 nm size gold nanoparticles and the increase in particle size revealed a decrease in skin permeation. Gold nanoparticles of 102 and 198 nm present a lag time of 3 and 6 h, respectively. When the localization of these nanoparticles in the skin was analyzed, it was observed that those of smaller size accumulated in the deeper region of the skin, while the larger ones were mostly in the epidermis region (Sonavane et al., 2008). Particularly for lipid nanoparticles, e.g., SLNs, size is related to the skin penetration because smaller particles (larger surface area) can form a more dense and cohesive film on the surface of the skin, increasing the occlusion and skin hydration. Due to the reduced water loss caused by occlusion, skin hydration is increased, influencing percutaneous absorption (Pardeike et al., 2009). In a study comparing three different mean volume diameters of SLN (77, 329, 932 nm), rhodamine B loaded in 77 nm particles was able to cross the SC through the hair follicles, while in the larger nanoparticles, the dye was not able to penetrate into deep layers, even after 24 h treatment (Adib et al., 2016). The small size ensures the close contact of the nanoparticles with the SC and can increase the amount of drug penetrated into the skin (Zhai and Maibach, 2001).

6.4.3 CHARGE Nanoparticle surface charge is also an important property to be considered. When in contact with a formulation at physiological pH, skin surface is negatively charged, due to carboxylate groups present in the SC that have a pKa of B4 (Blank, 1939). Therefore, the affinity of nanoparticles to SC is affected by electrostatic interactions (Bernfield et al., 1999; Uchechi et al., 2014). In this way, negative nanoparticles are generally repelled by the negatively charged skin domains (Honary and Zahir, 2013). It was observed by Wu et al. (2009), for instance, that negatively charged polymeric nanoparticles were not retained on the skin surface, nor released a significant amount of a fluorophore into the SC; in contrast to the positively charged nanoparticle, which presented a higher affinity to the skin, and improved the penetration of the fluorophore into SC. Moreover, positively coated nano-TiO2 significantly increased the skin flux of amphotericin, an antifungal agent, whereas no effect or structural alterations of SC lipids were noticed for negatively coated nano-TiO2 particles (Peira et al., 2014). The influence of nanocarrier surface charge in drug skin penetration can be noticed for solid nanoparticles as well as for nanoemulsions. Positively charged nanoemulsions showed to improve skin permeation of fludrocortisones acetate

6.5 Physical Methods to Enhance Nanoparticle Skin Penetration

and flumethasone pivalate more than the negatively charged ones (Hoeller et al., 2009).

6.4.4 SHAPE The nanoparticles shape has been repeatedly mentioned as a possibly relevant parameter for skin penetration and cellular uptake, but only rarely studied. In the few studies described in the literature, radius curvature of nanoparticles has been emphasized as an important factor to drug penetration. Some studies with human and mouse skin revealed that small rod-like nanoparticles have enhanced skin penetration, when compared to spheres, with a deposition only in the top skin layers (Fernandes et al., 2015). Additionally, small silver nanorods, compared to spheres and triangle-shaped particles, showed the highest silver-blood concentration after topical application in hairless mice (Tak et al., 2015). Experimental (Zhang et al., 2012) and theoretical (Dasgupta et al., 2014) works show enhanced binding with cell surfaces, but lower uptake of ellipsoidal compared with spherical particles.

6.5 PHYSICAL METHODS TO ENHANCE NANOPARTICLE SKIN PENETRATION Physical enhancement techniques, such as iontophoresis, electroporation, microneedles, laser and radiofrequency ablation, and sonophoresis, have been developed in an attempt to increase drug penetration through the skin. The advance in the development of nanoparticle delivery systems and the limited routes of their penetration led to the study of the influence of these methods in nanoparticle skin penetration (Cevc and Vierl, 2010; Mitragotri, 2013; Azagury et al., 2014). The association of a physical method in nanoparticle formulation aims to overcome the SC barrier, targeting the nanoparticle to specific site of skin. Microneedles, laser and radiofrequency ablation, for example, make possible the delivery of nanoparticles into viable epidermis, whereas iontophoresis predominantly promotes a follicle-targeted delivery of the nanostructures (Gratieri et al., 2012; Borgheti-Cardoso et al., 2016) (Table 6.2).

6.5.1 IONTOPHORESIS Iontophoresis is a noninvasive technique based on the application of a low density current to facilitate the penetration of substances into biological membranes, with a predominant net charge and permselectivity, especially in the skin (Hu et al., 2011; Huber et al., 2015; Bernardi et al., 2016), but also in the mucosa

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Table 6.2 Main Physical Methods and Parameters Used in Studies Associated With Nanoparticles Applied in Skin Physical Method Iontophoresis

Electroporation

Microneedles

Most Applied Parameters Range Electric current: direct or alternate Current density: 0.250.5 mA/cm2 (constant) Electrode: Ag/AgCl Polarity: cathodic or anodic Time: 15 min8 h Type: cathodic or anodic Voltage: 50200 V Number of pulses: 540 Pulse length: 5100 ms Needle size: 501400 μm Needle geometry: cylindrical, rectangular, pyramidal, conic, octagonal or quadrangular Needle type and mechanism • Hollow: “Poke and Flow” • Solid: “Poke and Patch,” “Poke and Release,” “Coat and Poke” Needle material: steel, silica, glass, hyaluronic acid, polymeric (poly (vinylpyrrolidone)), silicone

(Hu et al., 2011; Cubayachi et al., 2015) and cornea (Souza et al., 2014; Gratieri et al., 2016). Iontophoresis was first performed by Leduc, in 1900, who noticed that the application of an electric current in a solution of strychnine and cyanide applied over rabbits’ skin resulted in tetanic seizures and poisoning by the cyanide (Leduc, 1900). However, since the middle of the 1980s has the interest to study iontophoresis grown, which correlated with the growing importance of the transdermal route in drug administration. Since then, iontophoresis has been used in a wide range of fields, such as anesthesia (Glaviano et al., 2011; Hu et al., 2011), cancer treatment (Taveira et al., 2014; Byrne et al., 2015; Huber et al., 2015; Lemos et al., 2016), gene therapy (Venuganti et al., 2015; Aljuffali et al., 2016), and transcutaneous immunization (Bernardi et al., 2016; Ita, 2016). Initially, iontophoresis was performed using simple drug electrolyte solutions, however, with the development of nanotechnology, this technique has been studied in combination with different nanocarriers (Taveira et al., 2014; Huber et al., 2015; Bernardi et al., 2016; Malinovskaja-Gomez et al., 2016). In iontophoresis, the electric current is provided by a battery and distributed by two electrodes, the positive (anode) and negative (cathode), through an electrolyte solution, which can be the nanoparticulate formulation. When a current is applied, the cations in solution that are in contact with the anode move through


the skin toward the cathode, while the anions present in the cathode move in the opposite direction. To maximize the efficiency of iontophoretic delivery, the electrodes must be mechanically and chemically stable, and could be inert or reversible. The most commonly used electrodes are reversible Ag/AgCl, because their electrochemical reactions occur at a lower voltage than the one required for water electrolysis (Kalia et al., 2004). The density of electric current applied in the skin varies from 0.1 to 0.5 mA/cm2 (Barry, 2001). Iontophoresis facilitates the transport of substances into the skin by two mechanisms: electromigration and electroosmosis. Electromigration refers to the orderly movement of ions in the presence of an electric current. Therefore, the electromigration flux of an ionized drug is directly related to the flux of electric current (Eq. (6.2)). The transport number of drug depends on the mobility, charge, and concentration of the substance. It also depends on the current density, duration of application and the area of skin surface in contact with the electrode (Kalia et al., 2004). Electroosmosis refers to a flow of solvent and movement of charges when an electric potential is applied to the skin. The electroosmotic term derives from the fact that, under the influence of an electric current, the skin, at physiological condition, favors the transport of Na1 over Cl2 because of the cation permselectivity of the SC (Burnette and Ongpipattanakul, 1987). Specifically, SC has an isoelectric point of approximately 4.04.5 (Marro et al., 2001). Above this pH range, the carboxylate groups associated with the amino acid residues present in the skin are ionized. Therefore, the skin is negatively charged when in contact with a solution at physiological pH, which favors cations, and neutral molecules transport and hinder anions transport. The solvent flow can also drag species with high molecular weight (Guy et al., 2000; Marro et al., 2001; Kalia et al., 2004), which may be advantageous for nanoparticles. The total iontophoretic flux is the sum of electromigration and electroosmosis contributions (Eq. (6.2)): Jtotal 5 Jelectromigration 1 Jelectroosmosis zp up cp P 1 v 3 cp Jtotal5 F: z1 u1 c1

(6.2)

where zp, up and cp are drug charge, mobility, and concentration, respectively; I is current density, F is Faraday’s constant (96,500 C/mol); v is the electroosmotic solvent velocity. The relative contribution of electromigration and electroosmosis depends on the physicochemical and electrical characteristics of skin and the permeant. Skin’s negative charge can be also reduced, neutralized, or even reversed by the iontophoresis of certain cationic and lipophilic species. In fact, electroosmosis is significantly reduced by the association of negative compounds to the skin (Guy et al., 2000).

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The contributions of electromigration and the electroosmotic flow on the iontophoresis of DOX loaded in anionic SLNs (175 nm, 231 mV) were investigated. It was shown that, because DOX was positively charged in the formulation (pH 5), the release from SLNs was influenced by electromigration in addition to electroosmosis, increasing penetration into the SC, and then diffusing to epidermis. The amount of drug in different skin layers was distributed differently after passive and iontophoretic administration. The drug retained in the SC decreased from 43% to 26%, whereas accumulation in the viable epidermis enhanced from 26% to 56% after iontophoresis of SLNs. However, the amount of drug found in the receiver solution decreased from 30% to 18% after the application of the electric current. The authors hypothesized that it could be due to the fact that iontophoresis of SLNs pushed more lipids into the SC than passive delivery, which modified the barrier, DOX partition, and diffusion in different skin layers (Taveira et al., 2014). The amount of compound delivered by iontophoresis is directly proportional to the charge, so, in the case of nanocarriers, the determination of the zeta potential is an important step, since contact with the right electrode can maximize the efficiency of iontophoretic penetration. The zeta potential of liposomes shows to play a major role during iontophoresis of insulin. Anodic (from the positive electrode) iontophoresis (0.45 mA/cm2 for 1 h) of cationic liposomes (-N-[1-(2,3Dioleoyloxy)propyl]-N,N,N-trimethylammonium methyl-sulfate (DOTAP) based), with zeta potential ranging from 29 to 52 mV, increased the delivery of insulin into hair follicles to a significantly greater depth than cathodic (from the negative electrode) iontophoresis of anionic liposomes (cholesteryl hemisuccinate-based), with zeta potential varying from 223 to 261 mV (Kajimoto et al., 2011). Iontophoresis also depends on the current applied, the duration of application, and the area of the skin surface in contact with the electrode. The enhancement of the cumulative penetration amount and the steady-state penetration flux of the triamcinolone acetonide acetate entrapped in SLNs were related to the characteristics of the applied pulse electric current, such as density, frequency, and on/off interval ratio (Liu et al., 2008). The density of the electric current, drug concentration, and solvent type influenced in vitro iontophoresis of vancomycin hydrochloride encapsulated in ethosomes, which are phospholipid-based nanovesicles containing a high content of ethanol (Mohammed et al., 2016). The cathodic iontophoresis of negatively charged ethosomes showed maximum transdermal flux (550 μg/cm2/h), compared to free drug solution. The transdermal flux was reduced by altering the current mode from continuous to ON/OFF mode, reducing current density (from 0.5 to 0.25 mA/cm2) and by using normal saline instead of pure water as solvent (due to the competition between the small and mobile chloride ions, and negatively charged ethosomes); conversely, flux was potentiated by increasing drug concentration from 25 to 75 mg/mL (Mohammed et al., 2016).


Some studies have shown that the increase in drug penetration caused by iontophoresis occurs mainly through pre-existing aqueous pathways (paracellular) (Regnier and Preát, 1998; Pliquett et al., 2000). However, for nanostructured systems, which do not have conclusive demonstration that they can penetrate the skin in significantly therapeutically dosage, accumulation in hair follicles seems to be the main route of penetration (Kajimoto et al., 2011; Kigasawa et al., 2012; Huber et al., 2015). The use of anodic iontophoresis in combination with cationic SLN containing DOX (136 nm, 156 mV) showed, in vitro, that the permeation of DOX into viable epidermis was increased by approximately 50-fold, compared to passive delivery from the same SLN, as well as creating high-concentration DOX reservoir zones in the furrows and follicles of the skin (Huber et al., 2015). In vivo, this treatment was effective in inhibiting tumor cell survival and tumor growth, and was accompanied by an increase in keratinization and consequent cell death, indicating a strong and synergic effect of iontophoresis with the DOX-SLN for the topical treatment of skin cancer (Huber et al., 2015). Gene silencing is an effective strategy to limit disease progression by inhibiting the expression of the target protein. A positively charged poly(amidoamine) (PAMAM) dendrimer was employed as a carrier for topical iontophoretic delivery of an antisense oligonucleotide (ASO). In vitro anodic iontophoresis (0.3 mA/cm2 for 4 h) significantly enhanced the penetration of the intact ASO-dendrimer complex into the viable epidermis, benefited by both electromigration and electroosmosis mechanisms. In vivo, the iontophoretic (0.5 mA/cm2 for 2 h) delivery of the ASO-dendrimer complex significantly reduced the tumor volume in mice by 45%, due to gene silencing (Venuganti et al., 2015). The potential of employing iontophoresis for transcutaneous immunization was investigated using a vaccine formulation composed of ovalbumin (OVA) as a model antigen, loaded into negatively charged liposomes containing silver nanoparticles (NPAg) (128 nm). In vitro cathodic iontophoresis (0.5 mA/cm2 for 15 min) increased OVA penetration into the viable epidermis by 92-fold, in comparison to passive delivery (passive penetration: 0.05 6 0.02 μg/cm2, iontophoresis: 4.61 6 0.12 μg/cm2). In vivo, transcutaneous immunization generated a greater response via the Th2 pathway, inducing the production of antibodies, differentiation of immune-competent cells and appeared to present an alternative strategy for needle-free vaccination (Bernardi et al., 2016).

6.5.2 ELECTROPORATION Electroporation is based on the skin application of high voltage pulses, for periods varying from microseconds to milliseconds. The application of the pulses creates aqueous pores in the skin, facilitating the permeation of mainly macromolecules and hydrophilic drugs. The application is made by two electrodes, placed in a chamber containing the solution or formulation, in contact with the skin. The electroporator connected to the electrodes generates the energy dose, creating

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the transient pores. Electroporation can also be used as a pretreatment, before the application of active formulation. For that, the technique is employed using an electrolytic solution, and then replaced by the drug formulation after the end of pulse application (Lopez et al., 2012; Wong, 2014). The drug transport by electroporation can be affected by the voltage, number of pulses, and application duration (Eq. (6.3)): E5

NτðVi2 2 Vf2 Þ 3R 2

(6.3)

where E is the energy of the electric pulse, N is the number of the pulses, τ is the duration of the pulse, Vi and Vf are the initial and final voltages, respectively, and R is the resistance of the diffusion chamber during the pulse. The voltage applied to the skin is much higher than the breakdown threshold of the SC (30100 V), enabling the creation of local transport regions (LTRs). It is suggested that these LTRs act as aqueous pores through the skin, due to the disorganization of the SC lipid matrix. The size of the LTRs is dependent on the density of the voltage applied. After the creation of the LTRs, the drug can diffuse passively through the skin by these routes. Electrophoretic transport can also occur additionally to diffusion when electroporation is applied for longer periods. In this case, charged substances are repelled to the skin when in contact with electrodes with the same polarity, similar to what occurs by electromigration in iontophoresis. In addition, electroosmosis can also contribute to the transport of uncharged hydrophilic substances. Despite the occurrence of these mechanisms, diffusion still plays a major role in the transport of all types of molecules by electroporation (Lopez et al., 2012; Wong, 2014). Skin penetration of insulin after electroporation of insulin-loaded nanoparticle, negatively charged and simple insulin solution were compared in vitro and in vivo (Rastogi et al., 2010). Diffusion of insulin from the solution on passive application was negligible (0.005 6 0.001 μg/cm2). Upon electroporation, the insulin flux changed, linearly proportional to the applied voltage (50200 V), the number of pulses (2040), and the pulse length (515 ms). Passive application of nanoparticles resulted in 1.7-fold increase in cumulative insulin permeation compared to passive insulin solution. Application of electroporation increased insulin skin permeation in both cases. However, electroporation of nanoparticles decreased insulin permeation when compared to insulin in solution, but increased fourfold the insulin accumulated in the skin. In vivo, electroporation of nanoparticles prolonged hypoglycemic levels in diabetic rats from 24 to 36 h, compared to the solution (Rastogi et al., 2010). The effect of electroporation on the human skin penetration of estradiol, a lipophilic drug, from saturated aqueous solution and from ultradeformable liposomes was investigated. Cathodic electroporation (five pulses, 100 V, 100 ms, interval of 1 min) was employed, since liposomes were negatively charged. Skin pulsing considerably increased the estradiol penetration and


skin accumulation from solution, relative to passive delivery. Electroporation of liposomes, however, did not markedly affect estradiol skin penetration. This could be due to energy dissipation by an electrophoresis process, or even due to the ability of phospholipids to actually restore barrier function, acting as retardant and overriding their normal penetration-enhancing action (Essa et al., 2003).

6.5.3 MICRONEEDLES The use of microneedles is a relatively new approach to administrating drugs into the skin. It is based on needles of microdimensions, with a length varying from 50 to 900 μm and an external diameter up to 300 μm, that possess the ability to overcome the SC with a thickness of 1020 μm. The needles create transitory pathways by the mechanical interruption of the SC barrier, promoting an effective and minimally invasive drug delivery. Microneedle arrays can be smaller than 1 mm thickness and an area of 1 cm2 (Van der Maaden et al., 2012; Kim et al., 2012). Once inserted into the skin, microneedles can reach viable epidermis and also the superficial dermis, while the drug can diffuse to act locally or be distributed systemically by the dermal capillaries. Despite the presence of nerves located below the SC, the use of microneedles is painless because of their small diameter and length, also avoiding patient fear and needle-phobia (Kim et al., 2012; Larran˜eta et al., 2016). Microneedles have been studied and developed since 1970, but it was only in the middle of the 1990s that it became possible to effectively produce microneedles adequate for therapeutic purposes, due to the technology development in microfabrication. The increasing industrial knowledge also enabled the association of microsensors, micropumps, and microcircuits into microneedles arrays (Van der Maaden et al., 2012; Kim et al., 2012). There are a wide range of materials that microneedles can be made from, including metals (Tawde et al., 2016), polymers (Wang and Wang, 2015; Zaric et al., 2015; Eltayib et al., 2016; Modepalli et al., 2016), silicone (Gomaa et al., 2012; Deng et al., 2016), and ceramics (Boks et al., 2015; Hartmann et al., 2015). The needles can be cylindrical, rectangular, pyramidal, conic, octagonal and quadrangular. The shape can affect their force, ability to perforate the skin, the flux, and rate of drug release (Kochhar et al., 2013). Finally, microneedles can be solid or hollow (Prausnitz, 2005). Hollow microneedles act by an approach called “poke and flow,” that is closely similar to a conventional hypodermic needle, delivering the agent in bolus, or to a catheter, enabling an infusion therapy (Prausnitz, 2005; Kim et al., 2012). Solid microneedles are more resistant than hollow ones, easily produced, perforate the skin efficiently, enhancing its permeability. They can exert their action through three different strategies: “poke and patch,” “poke and release” and “coat and poke.” With “poke and patch,” the microneedles are applied, removed, and

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replaced by the formulation containing the drug, that can then diffuse through the microchannels. With “poke and release,” microneedles act as drug reservoirs, remaining on the skin during the release. Finally, the “coat and poke” mechanism is based on microneedles coated with small amounts of the active, to avoid interference on perforation ability. Consequently, this method is only suitable for highly potent substances, such as peptides and vaccines (DeMuth et al., 2010; Van der Maaden et al., 2012; Haj-Ahmad et al., 2015). Overall, this technique presents several advantages, such as the facility of the needles to administer the drug, the convenience of transdermal patches, and the possibility to overcome the SC in a minimally invasive way. There are no limitations concerning the type of therapeutic agent to be administered, being suitable for larger molecules, such as peptides, antigens, DNA, and also a variety of nanocarriers (DeMuth et al., 2010; Gomaa et al., 2012; Kumar et al., 2012; Zaric et al., 2015; McCaffrey et al., 2016). The “poke and patch” strategy has been used to investigate the influence of nanoparticles in skin penetration associated with microneedles. The penetration and the distribution of fluorescent PLGA nanoparticles in the human skin treated with silicon microneedles was investigated in vitro across porcine skin. Once the needles were inserted into the skin, the applicator was held in the position for 1 min to let the microneedles create micro conduits. Then, nanoparticles of different sizes (160, 205 and 288 nm) were delivered into the micro conduits created by microneedles, and permeated into the epidermis and the dermis. The penetration was in a particle size-dependent manner and increased with nanoparticle concentration, until a limit value was reached (Zhang et al., 2010). Microneedle-based vaccination system is an emerging modality to induce an immune response through the skin with the least patient inconvenience (Quinn et al., 2014). Tetanus toxoidloaded chitosan nanoparticle (208 nm) distribution was evaluated ex vivo after a pretreatment with a solid microneedle roller system (540 needle, 500 μm length). The nanoparticles were delivered at desirable depth into epidermis and dermis, being accessible to antigen presenting cells, eliciting immune response. Furthermore, the property of nanoparticulate antigen to sustain release leads to more prominent humoral and cellular immune response, in comparison to intramuscular administration of soluble antigen (Siddhapura et al., 2016).

6.6 EXPERIMENTAL TECHNIQUES FOR STUDYING NANOPARTICLE SKIN PENETRATION To evaluate the influence of nanoparticles in skin penetration, their physicochemical characteristics, such as size, shape, and charge need to be known to better understand which of them, and in which intensity, can affect skin penetration. Besides this, it is important to emphasize that nanoparticles characteristics may change when they interact with the skin.

6.6 Studying Nanoparticle Skin Penetration

In silico studies have been developed to predict drug permeability (Ma¨lkia¨ et al., 2004; Ottaviani et al., 2007; Hatanaka et al., 2015) but it is extremely difficult to model percutaneous absorption, since permeation is a multifactorial phenomenon, dependent on several nonlinear chemical, physical, and biological interactions (Yamashita and Hashida, 2003). Therefore, in vitro skin penetration studies, ex vivo, and in vivo experiments are necessary in the development of a nanoparticle formulation to be topically applied. They make it possible to understand the mechanisms of penetration, the interactions of the nanoparticles with the skin, as well as to determine drug levels in the various layers of the skin (Table 6.3).

6.6.1 IN VITRO STUDIES Due to economic and ethical concerns, and to avoid animal testing in the pharmaceutical field, there is a great effort from the scientific community in the development and improvement of in vitro assays that allow good comparisons and correlations with the performance of drug and delivery system profiles in vivo. The use of in vitro assays allows for a large number of different formulations to be evaluated in a reasonable time and cost. Good reproducibility of the results and control of specific parameters are possible, toxic compounds can be evaluated, and there is no need for ethical approval (Bartosova and Bajgar, 2012).

6.6.1.1 In vitro skin permeation In vitro skin permeation studies provide information on the extent to which a drug can be delivered to different layers of the skin when administered from a nanoparticulate delivery system. It is also possible to estimate the flux of drugs through a skin model into a receptor solution (Butani et al., 2016). These studies allow various formulations to be rapidly tested, ensuring that the one that carries the drug in higher concentration to the desired skin layer is evaluated in vivo or in a specific model of disease. They contribute, in this way, to the optimization of the drug delivery system for topical administration (Flaten et al., 2015). To be performed, an in vitro skin permeation study requires a diffusion cell, a skin model, and a selective and sensitive analytical method that allows quantification of the drug in the different layers of the skin and receptor solution. The latter bathes the innermost part of the skin and simulates biological fluids. It is important to keep in mind that the permeation of the drug is monitored in these experiments. The nanoparticle that carries the drug is not quantified or visualized. Accordingly, in order to evaluate the influence of nanoparticles of different compositions or of different natures, it is important to use a control formulation. This should preferably have components similar to those present in the nanoparticle, but not structured like it. Consequently, one can infer the advantages of the use of the nanoparticle in the permeation of the drug and not only of its components.

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Table 6.3 Principal Methods Used to Evaluate the Influence of Nanoparticles on Skin Penetration of Drugs: Advantages and Disadvantages Methods Advantages

Disadvantages

In Vitro Penetration Studies Allows the distinct quantification of the drug released from the nanostructured systems into the stratum corneum, viable epidermis, hair follicle and receptor solution Reduced animal tests and ethics committee approval is not required Allows good correlation with drug penetration profile in vivo Reasonable time and cost effective Useful for optimization of drug delivery systems for topical administration

Skin viability is not maintained It is not possible to evaluate the influence of local metabolism of drugs Need to maintain sink conditions

Electron Microscopy (TEM and SEM) High resolution and fast images SEM is useful for visualization of nanoparticles in different compartments of the skin, such as hair follicles and furrows

Skin should be prepared and fixed TEM is not indicated for metallic or silica nanoparticles

Fluorescent Microscopy (CLSM and TPPL) No need for sample preparation Horizontal optical cuts TPPL affords data with thicker (14 μm) and larger (250 μm) sections of skin High definition images with higher contrast Samples under physiological conditions can be analyzed

Fluorescence is required Skin is autofluorescent

Atomic Force Microscopy (AFM) Samples under physiological conditions can be analyzed Contrast use is not necessary Nondestructive and fast sample analysis

Specific molecules cannot be detected Analysis is limited to the surface of nanomaterials

Raman Spectroscopy Samples under physiological conditions can be analyzed Analysis time is short Good reproducibility

Equipment setup is expensive Thermal decomposition of sample can occur when the excitation intensities are too high High levels of fluorescence (intrinsic or caused by impurities), which overlay the Raman bands (Continued)


Table 6.3 Principal Methods Used to Evaluate the Influence of Nanoparticles on Skin Penetration of Drugs: Advantages and Disadvantages Continued Methods Advantages

Disadvantages

Electron Paramagnetic Resonance (EPR) Spectroscopy High structural resolution One of the most sensible magnetic resonance methods Provides qualitative and quantitative information

Drugs which do not contain unpaired electrons need to be chemically bonded to a stable free radical (spin label) Spin label signal may interfer with antioxidant system of the skin (enzymatic and nonenzymatic) reducing the free radical formation and its intensity

In Vivo Tape Stripping Simple and noninvasive method Quick skin barrier recovery Efficient to remove drugs from SC and alternative routes (e.g. hair follicles) when differential tape stripping is used

Fails to remove strongly coherent keratinocytes Process must be standardized and executed by a single person Ethics committee approval is required Only the quantification on upper skin layers is possible

Microdialysis Continuous measurement of drugs in a complex and dynamic environment Causes minimal interference in the individual or organ physiology Possible to investigate drug pharmacokinetic and therapeutic response

More invasive technique Ethics committee approval is required The size of the drug must be one-quarter the size of the membrane cut off (specially attention for macromolecules studies)

6.6.1.1.1 Diffusion cells The diffusion cells are composed of two compartments separated from one another by a membrane. There are several models: vertical or horizontal, static or continuous, closed or open. In vertical diffusion cells, or Franz-type diffusion cells (as a tribute to its inventor) (Franz, 1978), the skin is placed horizontally between the donor compartment (where the formulation is placed) and the receptor. In horizontal diffusion cells, the skin is placed vertically between these two compartments. Open models do not have the donor compartment occluded; closed systems do. Today, Franz-type diffusion cells are the most used because they better simulate the dynamics of a topical application. They can have the donor compartment closed or not. Although occlusion is not the most common form of substance

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administration, it is often performed to minimize variation between experiments due to undue evaporation of the drug when a very small mass of formulation is applied. However, each nanoparticle should be evaluated separately for the choice of the best diffusion cell to be used. Deformable liposomes, for instance, have penetration reduced when occluded, due to the lower osmotic gradient, which is the driving force of liposome transport (Cevc and Blume, 1992; Honeywell-Nguyen et al., 2003). Static and flow-through diffusion cells models refer to the dynamics of the receiving solution. In the static model, the volume of each sample collected from the receiver compartment each time needs to be replaced with an equal volume of fresh medium; in the flow-through type, a pump is used to perfuse the receiving fluid through the receiver chamber at a constant flow over time. The latter better simulates the blood circulation that perfuses the dermis in vivo, but is less frequently used than the static model. This is because static receptor solution allows greater control of the experiment, as well as the possibility of quantifying a larger mass of drug at each time, which ends up being diluted when the receiving fluid is flowing continuously (Contri et al., 2011). Besides the Franz-type, another type of vertical diffusion cell that is used by some research groups, is the Saarbruecken cell. Saarbruecken-type diffusion cells provide data only for drug penetration into the SC and deep skin layers rather than permeation in the receptor solution, since there is no presence of acceptor medium. A wet filter paper is placed under the skin to prevent it from drying out. This model was developed by H. Loth and colleagues to allow penetration experiments with a long duration without the need of a liquid receptor medium, since the skin itself acts as a receptor chamber and the closed system decreases its deterioration, preventing nonphysiological hydration (Wagner et al., 2001).

6.6.1.1.2 Skin model In vitro skin penetration experiments should be performed using skin, preferably of natural origin, such as human and porcine skin. Synthetic membranes, such as cellulose acetate, silicone, or mixed are not suitable for these studies. However, they are used in in vitro release studies where a formulation should have minimal interaction with the membrane to allow an evaluation only of the influence of the formulation on the release of the drug. Skin permeation studies, converseley, performed with skin as a membrane, aim to verify interactions of the release system with the skin and possible changes caused by it in the skin barrier (Flaten et al., 2015). Tissue culture-based human skin models can be used as a membrane in in vitro permeation studies. They are classified as human reconstructed epidermis models (e.g., EpiSkin, SkinEthic, EpiDerm) and as living skin-equivalent models (GraftSkin, EpiDermFT, Pheninon) (Netzlaff et al., 2005; MacNeil, 2007; Alnasif et al., 2014; Mathes et al., 2014; Planz et al., 2016). The reconstructed epidermis models are constructed from human keratinocytes cultured in cellulose or polycarbonate membranes. The cells are stimulated to differentiate and to form a


multilayer matrix, similar to epidermis. The living skin-equivalent models are constructed from fibroblasts seeding in a de-epidermized dermis (DED) or in a collagen gel. When using DED, later seeding of keratinocytes is necessary, applying air exposure in order to differentiate them. When fibroblasts are seeding in a collagen gel, keratinocytes are seeded on top of the collagen matrix until proliferation and differentiation (Van Gele et al., 2011). Compared to human skin, culture-based human skin has as advantage of less variation in permeability, but presents higher permeability to hydrophobic compounds (Schmook et al., 2001), less resistance when exposed to alcoholic vehicles (Dreher et al., 2002), and is more expensive. Furthermore, the absence of hair follicles in the existing models so far should be considered when studying nanoparticles. As these skin appendages have been prominent in skin penetration of nanoparticles, reconstituted skins should be avoided in these studies. When using nanoparticles to evaluate the difference between the use of reconstituted skin versus in vitro human skin, Labouta et al. (2013) found a sevenfold increase in the number of gold nanoparticles that penetrated the skin when the reconstituted skin model was used. This result reaffirms the structural deficit of the reconstituted skin so far, reinforcing the importance of using natural origin skin models in in vitro skin penetration studies. Several naturally occurring skin models have been used as a membrane for in vitro permeation experiments. Among them is the human skin and pig skin, which are widely used, as well as hairless mouse skin and the shed skin of snake after molting. Today, human skin is considered by regulatory agencies as the best model when the formulation developed is for human use, but pig skin is most used in experiments because of its similarity with human skin (Dick and Scott, 1992) and possibility of obtaining it in large quantities. Another advantage of pig skin for studies of nanoparticulate systems is the greater control of follicle size after excision; hair follicles of human skin contract immediately after excision, whereas in pig skin, its original size is maintained, due to the presence of cartilage (Planz et al., 2016). The skin of various parts of the pig can be used (Bolzinger et al., 2008; Hathout et al., 2010; Nam et al., 2012; Zhai et al., 2014), but the most used part is the ear, because it is not consumed in most countries, avoiding unnecessary death of the animal. When acquired directly from slaughterhouses, care should be taken to remove it prior to any scald procedure of the animal. Because the main barrier to skin permeation is the SC, which is formed by dead cells, it is not necessary to maintain the viability of the tissue. However, it is critical that the SC retains its structure during tests to ensure its barrier function. Two methods are the most used to verify the integrity of the SC before experiments: the measurement of TEWL and the measurement of skin resistivity to the passage of an electric current. The TEWL is measured using by evaporimeters. Typical basal values of TEWL in adults with healthy skin are in the order of 510 g/m2/h2. Physical or

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chemical skin damage could led to an increase in TEWL, proving a loss in the SC barrier function (Kalia et al., 2000; Senyi˘ ¸ git et al., 2016). The electrical resistivity of the skin is another indicator of the structural state of the SC (Tang et al., 2001; Karande et al., 2006). To perform this measurement, the skin must be mounted on the Franz-type diffusion cell and the receiver and donor compartments filled with phosphate buffer saline (PBS). Ag/AgCl electrodes should be inserted into these compartments. The electrode inserted into the receiver compartment is connected to an alternating electric current generator (100 mVRMS AC, 10 Hz). The electric current that goes through the system is measured in the donor compartment through the electrode inserted therein, connected to an ammeter (Fig. 6.3). The total resistance (Rt) of the system to the passage of this electric current is calculated using Ohm’s law (Eq. (6.4)). To determine the resistance imparted by the skin to the passage of the electric current (Rskin), the diffusion cell is assembled without the skin and the voltage generated only by the buffer used for calculating the buffer resistance (RPBS). Considering that Rt, in the presence of skin, is the sum of Rskin and RPBS, Rskin is calculated by the difference between Rt and RPBS (Eq. (6.5)). The resistivity of the skin (r) is determined by the multiplication of the Rskin by the area (A) of the skin available for permeation in the diffusion cell (Eq. (6.6)). R5

U I

FIGURE 6.3 Schematic representation of the electrical resistivity analysis of skin for in vitro skin permeation studies.

(6.4)


where U is the voltage and I is the current. Rskin 5 Rt 2 RPBS

(6.5)

r 5 Rskin 3 A

(6.6)

The skin resistivity measurement should be performed before each experiment involving the skin, to determine if the skin is adequate for the experiments, based on an established resistivity threshold. The limits used in studies range from 35 to 50 kΩ cm2, which were determined by a validation study involving fresh and stored freeze human skin samples, considering its sodium ion permeability and electrical properties. Human skin samples with an initial resistance of 52 kΩ cm2 present a break in the permeability value (Kasting and Bowman, 1990; Tang et al., 2001; Davies et al., 2004). The skin can be used with all its layers (full-thickness skin) or with part of the dermis removed with the aid of a dermatometer (split-thickness skin), in order to have a greater control of skin thickness and less variability between experiments. Split-thickness skin present all the skin layers and a total thickness varying in literature between 200 and 700 μm (Li et al., 2003; Huber et al., 2015; SCCS, 2010). Depending on the purpose of the study, only the epidermis (Frasch et al., 2011) or only the SC (Zellmer et al., 1995; Kong et al., 2011) can be used, provided that the integrity of the SC is ensured.

6.6.1.1.3 Applied dose The dose of the formulation to be tested, which is placed in the donor compartment in contact with the SC, may be finite or infinite. The finite dose (110 mg/ cm2 or #10 μL/cm2) is used for topical, local action products, when simulating the actual conditions of use of the formulation (OECD, 2004). Most often, however, when it is intended to study the influence of various delivery systems on the skin permeation of a drug, the infinite dose, where the concentration of drug in the donor is constant and significantly greater than that found in the receiving solution, is used (Selzer et al., 2013).

6.6.1.1.4 Receptor solution The receptor solution should be hydrophilic in nature and, ideally, at a pH close to the physiological one, in addition to guaranteeing drug stability and sink conditions. In this way, the amount of drug that crosses the skin during the study should not exceed 10% of the solubility of the drug in the receptor medium. High concentrations of drug at the receptor solution can interfere with the rate of drug permeation and result in an erroneous or underestimate permeation profile. To ensure sink conditions, low concentrations of surfactants (Tan et al., 2011; Silva et al., 2012; Song et al., 2012), ethanol (Fang et al., 2008; Senyi˘ ¸ git et al., 2010), or serum albumin (Aungst et al., 1990; Moser et al., 2001; DragicevicCuric et al., 2010) are commonly added in the receiving solution. They should not impair, though, the barrier function of the skin during the experiment period.

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The barrier function of the skin in contact with the receptor solution can be verified by the resistivity measure, for example, in experiments prior to submission of the skin to the formulation under study. Constant exchange of the receptor solution or its continuous flow helps maintain sink conditions, but also impairs drug quantification. When the receiving solution is maintained under static conditions, it should be stirred continuously during the experiment to avoid high local concentration and a defect in sink conditions. The temperature should be kept at 37 C to mimic the internal body temperature and regulated with a water jacket around each diffusion cell or by placing it in a water bath.

6.6.1.1.5 Quantification of the drug on the skin and in the receptor solution Receptor solution The permeation profile of the drug is determined by quantifying the drug in the receptor solution, as a function of the time and skin area available for permeation. Attention should be given to the correction of sample dilution at each sampling time, when the receiving solution is in a static flow regime. In these cases, in order to determine the cumulative amount of drug permeated through the skin with time, it is necessary to perform a calculation according to Eq. (6.7) (Salerno et al., 2010): Cn V 1 Q5

i5n21 P i51

A

Ci Vi (6.7)

Q is cumulative amount of drug in determined time/interval (μg/cm2); Cn and Ci are drug concentrations in receptor cells and concentration of obtained samples (μg/mL), respectively; V and Vi are receptor solution volume and sample volume (mL); and A is the permeation area (cm2). The drug flux (J) can be determined from the permeation profile after suitable mathematical treatment. Normally, to determine the kinetics of drug permeation, three mathematical models are used: zero order, Higuchi, and first order. Zero order model is considered the ideal drug delivery model for controlled release, as it maintains a constant drug supply that is independent of the drug concentration. Zero order model graph involves plotting the cumulative amount of drug released versus time. In the Higuchi model, data are expressed as cumulative percentage drug release versus square root of time; it describes drug release from a matrix system. For first order model, the result is plotted as log cumulative percentage of drug remaining vs. time. The correlation coefficient (r) for each kinetic model is calculated to determine the model of best fit. Skin At the end of the permeation experiment, the skin is removed from the diffusion cell and attached to a smooth surface, with the SC facing upwards (Fig. 6.4B). The part of the skin subjected to permeation is then removed by tape

FIGURE 6.4 Schematic representation of the major steps for in vitro skin penetration studies. (A) Contact of nanoparticles formulation (donor compound) with split or full-thickness skin in a diffusion vertical cell. (B) Tape stripping technique, (1) tape is applied in skin surface, (2, 3) removal of stratum corneum layers. (C) Differential tape stripping, (1) application of cyanoacrylate glue into the hair follicle, (2) adhesive tape is placed on the glue and pressed lightly until its complete drying and polymerization, (3) the adhesive tape is drawn in a single movement to obtain a cyanoacrylate skin surface biopsy containing the follicular casts and corneocytes. (D) Epidermis and dermis are sliced and homogenized for further drug quantification (E) drug quantification from the receiver chamber and extracted from the different skin layers by a analytical method (e.g. HPLC).

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stripping technique. In this technique, SC is removed using about 15 adhesive tapes (for pig and human skin) so that more-or-less complete removal of SC is indicated by the brightness of the exposed skin (viable epidermis). The drug is extracted from the adhesive tapes, together with a suitable solvent, and quantified to determine the amount of drug accumulated in the SC. The remaining skin (skin without SC) should then be chopped and the drug extracted in a suitable solvent. Typically, a tissue homogenizer is used for skin grinding and better extraction of the drug by the solvent. The dispersion should then be filtered and the amount of drug present in these layers of the skin is analyzed by suitable analytical method. This methodology was used by several research groups to evaluate the performance of nanostructured systems when applied to the skin (Arau´jo et al., 2010; Taveira et al., 2012; Vaghasiya et al., 2013; Huber et al., 2015; Zhang et al., 2015b). For instance, Lopes et al. (2006) compared the skin penetration of cyclosporin A from liquid crystalline phases (cubic and hexagonal) and from olive oil. They verified that the hexagonal phase increased the drug penetration in the viable epidermis and also its percutaneous delivery (obtained by receptor medium quantification) but did not alter the retention of the drug in the SC. Conversely, the cubic phase increased drug concentration in the SC and in the viable epidermis, but did not alter the percutaneous delivery. In order to separate epidermis from the dermis, many techniques can be used, such as mechanical scraping (Kilfoyle et al., 2012) and heat or enzymatic digestion (Fort and Mitra, 1994). Mechanical scraping includes the use of a blade to visually remove the dermis. For heating separation, the skin should be heated at 5060 C for 45 s, which facilitates dermis detachment and separation with a blade. The enzymatic digestion process includes sample incubation at 4 C for few minutes in a 0.05%0.2% trypsin solution. Of all these methods, the heating technique is the most used for separating dermis from epidermis (Senyi˘ ¸ git et al., 2010; Tan et al., 2011; Scognamiglio et al., 2013; Tsai et al., 2015). When it is desired to verify the ability of a nanoparticulate delivery system to target hair follicles, immediately after the removal of the SC, but before the skin is chopped, a technique called differential tape stripping can be performed (Teichmann et al., 2005) (Fig. 6.4C). It is a combined method of tape stripping and cyanoacrylate skin surface biopsies that allows the determination of substances selectively present on hair follicles. The follicular content of the permeated skin area, already submitted to tape stripping and therefore without SC, is removed by the application of a drop of cyanoacrylate glue onto the stripped skin surface. A glass slide or an adhesive tape is placed on the glue and pressed lightly until complete drying and polymerization of the glue (approximately after 5 min). Thereafter, the adhesive tape should be drawn in a single movement to obtain a cyanoacrylate skin surface biopsy, containing the follicular casts and corneocytes. The amount of the drug present therein is then extracted with suitable solvent to determine drug accumulation in the follicular region (Lademann et al., 2007).


6.6.1.2 Transmission and scanning electron microscopy Transmission electron microscopy (TEM) and scanning electron microscopy (SEM) are generally used to characterize size and morphology of nanoparticles but they can also be used for assessing nanoparticle distribution inside the skin. Electron microscopy can provide reliable conclusions about the interactions of nanoparticles and skin layers, identifying areas where they can be found. In such techniques, the image is generated by transmitting beams of accelerated electrons through a pretreated sample, in order to enhance the visualization of nanostructures. The electron beam is scattered and some of them collide with the sample atoms, while electromagnetic lenses focus the electrons and magnify them, producing an image (Bibi et al., 2011; Lin et al., 2014). TEM provides high resolution (about 0.2 nm) and fast images (Bibi et al., 2011). However, certain skin compartments have very small dimensions, which are not assessed by TEM (Fernandes et al., 2015). In order to visualize interactions, after the contact with formulation, skin should be carefully split into 6090 nm thickness, avoiding possible contamination of nanoparticles localized in skin surface with those localized in deeper layers (Baroli et al., 2007; Fernandes et al., 2015). After that, skin samples are prepared, including the hydration, fixing, and washing steps, and placed in copper grids. The application of a contrast solution, commonly a hydro alcoholic uranyl acetate solution (3%), is sometimes necessary (Miquel-Jeanjean et al., 2012). However, for metallic or silica nanoparticles in particular, it may be difficult to differentiate electron density of nanoparticles from structural components of skin (Baroli et al., 2007), depending on the gray-scale background obtained in images. In this case, TEM analysis is coupled with other techniques, e.g., with energy dispersive X-ray analysis (TEM-EDX) for titanium dioxide (TiO2) agglomerates in skin layer visualization (Miquel-Jeanjean et al., 2012), or with plasma optical emission spectrometry (ICP-OES) for quantitative evaluation of gold nanoparticle penetration in skin (Fernandes et al., 2015). The SEM technique involves an electron beam scanning over the sample surface, giving information about the topographic details of the skin. Some SEM modes can reach 1 nm of resolution (Lin et al., 2014). In the same way as TEM, SEM could be useful for the visualization of nanoparticle interaction with skin surface (Batheja et al., 2011), specially in hair follicles, furrows (Prow et al., 2012), and sublayers (George et al., 2014). The skin should also be prepared and fixed (Prow et al., 2012). SEM was used, for instance, to visualize the distribution of aggregates of quantum dots (QDs) on the surface of treated skin after an in vitro penetration study. Different modified QDs showed deposition mostly in SC and in skin furrows (Prow et al., 2012). The interactions between SLN and skin surface lipids was also assessed by SEM microscopy (Ku¨chler et al., 2009). The microscopy images showed that, after 2 h of skin penetration study, the number of particles had declined and SLN lost their primary shape, seemingly melted

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and grounding the hypothesis of lipid particle fusion with the skin surface (Ku¨chler et al., 2009).

6.6.1.3 Atomic force microscopy Another valuable tool to characterize nanocarriers and their distribution in the skin is atomic force microscopy (AFM). The potentialities of AFM make it a tool of undeniable value for the study of biologically relevant samples, enabling images of the structure of biomolecules or bio-surfaces, study samples under physiological conditions (analysis of real time dynamics of some biological events), and measure local chemical, physical, and mechanical properties of a sample (Santos and Castanho, 2004). Therefore, it provides a minimal topographic scanning of the skin, in three dimensions, without the need of contrast or electric conductivity. It is nondestructive and samples are fast prepared (Alnasif et al., 2014). AFM consists of a sensor, a micrometric tip (probe) attached to a flexible shaft, named a cantilever, which moves in the three directions: x, y and z, with high precision (about 1 nm). A laser beam falls over the back side of the cantilever while it scans the topographic surface. Thereafter, an image is caught by a sensitive detector light (Sitterberg et al., 2010). The AFM can be classified in different types (contact and intermittent-contact (tapping) mode), based on the interaction between the tip and the sample surface (Ruozi et al., 2005). Generally, after treatment with nanoparticles, different sections of the skin are scanned and the data is plotted for frequency distribution and average force of adhesion, enabling identification of the skin depth of nanostructures. Using this technique, Bianco et al. (2016) evaluated silver nanoparticles in different depths of the SC. The images were obtained both in contact and tapping modes. Different ranges of aggregates and morphologies of silver nanoparticles were observed, according to the SC area analyzed. One of the main technique drawbacks is that AFM is not able to detect or locate specific molecules in skin surface, which limits the analysis of nanoparticle and sample surfaces (Lin et al., 2014).

6.6.1.4 Confocal laser scanning microscopy The confocal laser scanning microscopy (CLSM) is an image technique that produces high definition tridimensional images with higher contrast. It is similar to a SEM, but utilizes a focused laser bean instead of electrons to scan the sample point-by-point, in a three-dimensional way (Caspers et al., 2003; Alvarez-Romań et al., 2004; Zhang and Monteiro-Riviere, 2013). The great attraction of CLSM for the evaluation of drug delivery topical to the skin is the possibility to observe the distribution of the drug or fluorescent marker on the epidermis surface, skin structures, such as hair follicle and sebaceous glands, and also into the deepest skin layers, without physically dissecting the tissue (Alvarez-Romań et al., 2004;


Prow et al., 2012; De Rezende et al., 2014; George et al., 2014; Huber et al., 2015; Guo et al., 2015a). It usually requires the drug or the nanoparticle to emit fluorescence at a given wavelength. If the nanoparticle component does not possess intrinsic fluorescence, fluorescent dyes such as fluorescein 5-isothiocyanate (FITC) can be conjugated to the polymer that compose the nanoparticle before nanoparticle preparation (Zhang and Monteiro-Riviere, 2013). Since the skin is autofluorescent, it is possible to study the penetration of nanocarriers in the tissue, recording two different confocal images in individual excitation/emission wavelengths at the same plane and then superimposing both images (Alvarez-Romań et al., 2004). In order to visualize the distribution and drug penetration pathway of a polymeric micelle nanocarrier for tacrolimus delivery, a micelle copolymer was labeled with a fluorescent dye and the skin was evaluated by CLSM after 24 h of contact with the formulation (Lapteva et al., 2014). Microscopy images showed a deposition of micelles into the hair follicle, as can be seen in Fig. 6.5. CLSM was used to visualize the distribution of nonbiodegradable, polystyrene nanoparticles (20 and 200 nm) labeled with FITC across porcine skin in vitro. At 568 nm, only the autofluorescence from the tissue was observed by the red fluorescence detector. In contrast, 488 nm light was required to obtain an image of the FITC by the green fluorescence detector. The surface images revealed that the nanoparticles accumulated preferentially in the follicular openings in a timedependent manner, and that the follicular localization was favored by the smaller particle size (Alvarez-Romań et al., 2004). A red fluorescent dye, Rhodamine 6G, was added to liposomal formulations, encapsulated into the lipid bilayers of nanovesicles, in order to evaluate the delivery across porcine skin in vitro. The penetration from conventional liposomes was confined only to the upper skin layer with little permeation into the deeper skin layers, whereas ethosomes and deformable liposomes can penetrate into the deeper dermis (Guo et al., 2015a). CLSM studies were performed after passive and iontophoresis permeation to determine in vitro DOX distribution of double-labeled SLN, using DOX and boron-dipyrromethene (BODIPYS) (BOD) (De Rezende et al., 2014) as hydrophilic and lipophilic dyes, respectively. The localization of the drug, DOX, was observed in the red channel, and the nanostructure marked with BOD, in the green channel. Combining both images, it was possible to evaluate the distribution of the entire system into different skin layers and structures, such as the accumulation of SLNs in the furrows and deep in hair follicles after iontophoresis (Huber et al., 2015).

6.6.1.5 Raman spectroscopy The Raman spectroscopy is a vibrational spectroscopy, whereby the sample under investigation is illuminated by monochromatic laser light and the interactions between the incident photons and molecules in the sample result in the scattering

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FIGURE 6.5 Three-dimensional reconstruction of a hair follicle assessed by CLSM. (A) 3D reconstruction side view, and (B) 3D reconstruction view from the top (bar 5 50 μm). Source: Adapted from Lapteva, M., Mondon, K., Mo¨ller, M., Gurny, R., Kalia, Y.N., 2014. Polymeric micelle nanocarriers for the cutaneous delivery of tacrolimus: a targeted approach for the treatment of psoriasis. Mol. Pharm. 11(9), 29893001.

of light. Raman is another technique which provides accurate data about the molecular composition of the skin down to several hundred micrometers below its surface, enabling drug penetration analysis. It also enables the determination of in vivo concentration profiles, as it is completely noninvasive, but the drug must be in sufficient (high) amounts (Caspers et al., 2001; Franzen and Windbergs, 2015; Planz et al., 2016). The combination of CSLM and Raman spectroscopy techniques for penetration skin studies allows the skin morphology to be visualized by CSLM, facilitating the target of (subsurface) structures in the skin for Raman measurements, thus enabling the establishment of relations between skin architecture and its molecular composition. It provides important information for a wide range of applications in fundamental skin research, percutaneous transport, clinical dermatology, and cosmetic research, as well as for noninvasive analysis of blood analytes (Caspers et al., 2003). NCL containing ketoprofen and spantide II for psoriasis topical treatment were surface modified with different cell-penetrating peptides (CPPs). Confocal microscopy and Raman confocal spectroscopy studies were performed using fluorescent dye encapsulated into the NLC to investigate the effect of polyarginine chain length CPPs on dye-skin penetration. It was suggested that the peptide containing 11 arginine (R11) had significantly greater permeation-enhancing ability than other polyarginines and transactivation transcriptional activator (TAT)


peptides. The amount of ketoprofen and spantide II retained in dermis after topical application of NLC-R11 was significantly higher than solution and nonsurface-modified NLC after 24 h of skin permeation (Shah et al., 2012).

6.6.1.6 Electron paramagnetic resonance spectroscopy EPR spectroscopy techniques are used to provide information about the microenvironment of a paramagnetic molecule positions, which are defined as the molecules with unpaired electrons (free radicals) (Coderch et al., 2000). The use of such molecules, called spin labels, is useful for monitoring changes in the spin tropic movement of its unpaired electrons. Drugs or biomolecules which do not contain unpaired electrons can also be studied by this technique, as long as they are chemically bonded to a stable free radical (spin label), such as nitroxide or 3-carboxy-2,2,5,5-tetramethyl-1-pyrrolidinyloxy (PCA). These spin labels can be used as probes, as a fast alternative to fluorescence dyes, providing information about drug penetration in skin layers. One of the disadvantages of EPR is that the spin label signal intensity may interfere with the antioxidant system of the skin (enzymatic and nonenzymatic), leading to the reduction in the unpaired electrons (Haag et al., 2011; Saeidpour et al., 2017; Lohan et al., 2017). The EPR data varies with the microwave frequency applied (Coderch et al., 2000). Low frequency L-band measurements were applied to give information about the ability of a core-multishell (CMS) nanoparticle to promote the penetration of a spin label PCA into the skin. After the in vitro penetration study, skin was subjected to EPR measurement. To evaluate the PCA depth of penetration, seven layers of the SC were removed by tape stripping and the skin was evaluated after each tape removal. The change of EPR intensity of the spin label during the study was due to the PCA penetration in the skin, which could be detected until after six tapes (Haag et al., 2011). In another similar study, dexamethasone (Dx) was labeled with PCA and loaded in CMS. The distribution and penetration into porcine ear skin of DxPCA was assessed by pulsed W-band high frequency wave EPR. The nanocarrier facilitated the drug transport through the skin barrier, increasing the EPR signal intensity after 24 h of contact with the skin. The EPR signal decreases after the removal of three tapes after CMS application, while the control solution strongly decrease after the first tape (Saeidpour et al., 2017), suggesting that CMS increases drug penetration.

6.6.2 EX VIVO SKIN PENETRATION EXPERIMENTS Ex vivo models to mimic nanoparticle skin penetration comprise surgical portions of animal skin, often pig ears, which are perfused with oxygenated blood from the same animal or a tissue preservative solution (Lopez et al., 2012). Surgical pig ears allows studies related to cutaneous pharmacology and toxicology to be conducted in a viable skin preparation that has a normal anatomical structure and a functional microcirculation. Maintaining the skin viability, it is

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possible to evaluate the local metabolism of some drugs, which is not possible in the in vitro studies previously reported. However, ex vivo studies are more laborious and time-consuming, due to the difficulties in maintaining the excised tissue viable for several hours. There are few studies in the literature that actually use ex vivo models to evaluate the performance of drug delivery nanosystems. The use of this type of method was employed to demonstrate the major permeability of natural actives when encapsulated in PLGA nanospheres and the localization and accumulation of these nanoparticles in the hair follicles (Tsujimoto et al., 2007). The penetration profile of gold nanoparticles functionalized with peptides or thiol-containing polyethyelene glycols having either a terminal amine (positively charged particles) or carboxylic group (negatively charged particles), has been demonstrated in viable mouse and human skin models (Fernandes et al., 2015). Most studies described in the literature erroneously apply the term “ex vivo” for experiments that do not use necessary conditions for maintaining skin viability, either human or animal.

6.6.3 IN VIVO SKIN PENETRATION EXPERIMENTS A complete understanding about nanostructure properties, such as size, shape, composition, and aggregation-dependent interactions with biological systems is still lacking (Department for Environment, Food and Rural Affairs (DEFRA, 2005) and it is thus unclear how the exposure of animals or humans to different nanoparticles could produce harmful or appropriate biological responses (Colvin, 2003). For example, assuming that a well characterized 50-nm polymeric nanoparticle is applied in the skin of a patient with squamous cell carcinoma. Although in vitro studies may have shown that the drug it encapsulates penetrates the skin and suggests that nanostructures accumulate in the follicles, the tumor mass may alter the predicted interactions. Consequently, in vivo, the mentioned nanoparticles could become metabolized or altered, changing the expected effects (Fischer and Chan, 2007). Because of this uncertainty about nanosystem effects in biologicals environments, regulatory agencies and general public have raised questions in regards to nanotechnology-based products. Therefore, in vivo studies are an important step in the development of nanoparticles for medical application. The most used in vivo techniques for evaluating nanoparticles intended for topical applications are tape-stripping and microdialysis.

6.6.3.1 Tape stripping technique Tape stripping is a simple and efficient method used to assess drug penetration in the skin after topical application of different formulations, including nanoparticles. It has been used both in vivo (Schwarb et al., 1999; Cambon et al., 2000) and in vitro (Benech-Kieffer et al., 2000; Wagner et al., 2001), in humans (Surber et al., 1993; Wagner et al., 2001) and in animals (Benfeldt, 1998; Wagner et al.,


2001; Darmstadt et al., 2002). Accordingly, tape stripping has become a basic method to study the penetration and the reservoir behavior of topically applied substances (Lademann et al., 2009a), such as nanoparticle delivery systems. Tape stripping is performed by placing adhesive tape onto the skin surface area that had been exposed to a formulation, followed by gentle pressure and subsequent removal by a sharp upward movement, successively removing the cell layers of the SC (Alberti et al., 2001). Drugs or nanoparticles (Jeong et al., 2010; Raber et al., 2014) present in these layers are then also removed. After an adequate treatment of the tapes, the drug is extracted from them to quantify the amount which has permeated into the SC (Russell and Guy, 2009). Because the SC is the primary skin barrier, the amount of drug therein and the depth of its penetration can be used to estimate the drug permeability coefficient. Puglia et al. (2008) evaluated the potential of NLC to deliver antiinflammatory drugs in the skin. In vivo studies were conducted in some human groups, with the objective of assessing the drug permeation through the skin and NLC interactions with the SC. A tape stripping experiment was performed and concluded that NLC significantly enhanced the amount of drugs recovered from the SC and provoke an accumulation of the drugs into the upper skin layers, creating a reservoir which prolonged their release. The approach is classified as noninvasive and enables a fast recovery of skin barrier function (Menon et al. 1992). Due to its advantages and simplicity, the Food and Drug Administration (FDA) proposed a protocol for determining the bioequivalence of topically applied formulations using the tape stripping technique in 1998 (US FDA, 1998). However, the protocol was withdrawn in 2002, due to its inadequacy in predicting the efficacy of formulations that penetrate SC by alternative routes, such as hair follicles, for example (Boix-Montanes, 2011). This is a major disadvantage for the evaluation of rigid nanoparticles, which accumulate at these sites. However, to investigate nanoparticle targeting hair follicles, differential tape stripping can be applied, as described in In vitro skin penetration methods (Section 6.1.1), or selective follicular closing technique (Otberg et al., 2008). Selective follicular closing technique consists of blocking hair follicles and, consequently, the penetration process by this route, with a special wax mixture prior to the application of the formulation under investigation. After the experiment, the obtained penetration results can be compared with results obtained from skin areas with accessible hair follicles (Teichmann et al., 2006). Another optical device, such as CLSM or combined CLSM and confocal Raman spectroscopy, described in Section 6.1.5, are useful optical devices that can be used to investigate follicular penetration in skin biopsies obtained from in vivo experiments (Patzelt and Lademann, 2013). Several intrinsic and extrinsic factors need to be taken into consideration for the evaluation of drug permeability using the tape stripping methods. The amount of SC removed by a single adhesive tape strip depends on the age, the anatomical

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site (Rougier et al., 1987; Iguchi and Tagami, 2000), the number of cell layers (Ya-Xian et al., 1999), and the thickness of the SC (Schwindt et al., 1998). Besides this, the type of adhesive tape (Tsai et al., 1992), the duration of pressure onto the skin (Reed et al., 1995), and the force of removal from the skin (Lo¨ffler et al., 2004) are the main factors that influence the amount of SC removed (Lademann et al., 2009a). Because of this, the entire tape-stripping process must be standardized and executed, preferably by a single person. Moreover, it is important to emphasize that the properties of the topically applied formulation influence the amount of SC removed with every tape strip. When ethanolic or aqueous solutions are applied, adhesion of the horny layer to the tape strips is enhanced while application of an oily formulation, reduces adhesion to the tape strips; consequently, more strips are necessary to remove the same amount of SC (Jacobi et al., 2003). Until now, tape stripping is one of the main techniques used to study drug permeation through the skin. In the recent literature, many studies describe this protocol as an important tool to identify drug skin penetration and retention from nanoparticles administration (Scalia et al., 2012; Iannuccelli et al., 2013, 2014; Bianco et al., 2016).

6.6.3.2 Microdialysis Microdialysis is a sampling technique for continuous measurement of watersoluble molecules in the extracellular space of tissues or organs in vivo. It causes minimal interference in the physiology of the tissue and has been an excellent technique for investigation of pharmacokinetic and therapeutic response to drugs in the skin (Groth and Serup, 1998; Mathy et al., 2003, 2005). It has been used to evaluate the skin penetration of drugs dispersed in topical formulations, such as gels, lotions, creams or ointments (Kreilgaard, 2002). Recently, it has been used to evaluate the penetration of drugs from nanosystems (Kreilgaard, 2001; Desai et al., 2012; Zhang et al., 2014; Guo et al., 2015b; Yong-Tai et al., 2015). A microdialysis system consists of a microdialysis pump, a probe equipped with a semipermeable hollow membrane (dialysis membrane), and a micro-vial in which the dialyzed samples are collected. The probe is implanted in the skin and functions as a blood capillary, which is constantly perfused with a physiological solution at a low flow (0.55 μL/min) (Plock and Kloft, 2005; Li et al., 2006; Herkenne et al., 2008). The formulation is put over the skin and the substances present in the extracellular fluid (Cmedium) are filtered by diffusion into the fluid within the probe, resulting in a concentration within the perfusion medium (Cdialyzed). The samples are then collected and analyzed (Plock and Kloft, 2005; Li et al., 2006; Desai et al., 2012; Lopez et al., 2012). The probe can be implanted under the skin or superficially in the dermis. Studies done in small animal models, such as hairless mice, currently implanted the probe under the skin due to the thickness of their skin (Desai et al., 2012; Zhang et al., 2014; Guo et al., 2015b).


The principle of microdialysis is that a physiological solution pumped through the probe is in equilibrium with molecules in the surrounding tissue. The equilibrium of the drug molecule between the microdialysis membrane, the extracellular medium and the perfusion medium is a dynamic process and reflects only a fraction of the actual concentration present in the environment surrounding the probe (Zhou and Gallo, 2005). Recovery, i.e., the ratio of the dialysate concentration to the actual concentration of the medium, is influenced by numerous factors such as flow, membrane properties, physicochemical characteristics of the substance, geometry and location of the probe, temperature and physiological processes (transport, metabolism). Therefore, it is necessary that the maximum of factors to be controlled and that the recovery is determined for the probe to be used in the microdialysis experiments. The geometry and the size of the probe influence the recovery of the drug in the dialysate, but must be determined according to the target organ (Waga and Ehinger, 1995). They may be linear or concentric, although for performing skin sampling, linear probes are commonly used. They minimize the damage to this tissue compared to the concentric probes because they are smaller in diameter and more flexible, which compensates for the need to perform two perforations in the skin, one inlet and one in the outlet. The pore of the microdialysis membrane that makes up the probe, influences the dialysis process. Only molecules with a molecular weight smaller than the membrane cut off can be collected. However, in order to obtain adequate recovery, the size of the molecule of interest must be about one-quarter the size of the membrane cut off, since this is determined under equilibrium conditions. Experimentally, as there is flow of the perfusate in the interior of the probe, this size may be reduced (Plock and Kloft, 2005). In this way, attention must be given to the study of macromolecules. The perfusion fluid is an aqueous solution that mimics the composition of the surrounding medium, like a buffer solution; this is because it is necessary to prevent the excessive migration of molecules into or out of the probe fluid due to osmotic differences. Besides this, the direction of the diffusion process is dependent on the concentration gradient (Lopez et al., 2012). Perfusion flow is another parameter that must be controlled. In general, low flows result in high recovery and high flows in low recovery, according to the Eq. (6.8). RR 5 1 2 e2rA=F 3 100

(6.8)

where RR is the relative recovery; r is the mass transport coefficient; A is the surface area of the microdialysis membrane; and F is the perfusion flow. However, low perfusion flows are limited by the low volume of the samples collected and the limit of quantification of the analytical method (Plock and Kloft, 2005).

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To characterize the rate of transfer (or exchange) of drug through the microdialysis probe (or membrane), in vitro and in vivo recovery studies should be performed. In vitro recovery is determined by putting the probe in contact with drug solutions at different concentrations, and perfusing it with the same buffer solution used in the in vivo studies. Drug concentrations in the dialysate should be quantified and the in vitro recovery calculated according to Eq. (6.9) (Anand et al., 2004). Rð%Þ 5

Cdialyzed 3 100 Csolution

(6.9)

In vivo recovery is determined by the retrodialysis technique (Sta˚hle et al., 1991; Rittenhouse et al., 1998). In this study, the probe is inserted into the skin and the perfusion liquid contains a certain concentration of the drug. The concentration of drug quantified in the dialysate is evaluated; determination of the in vivo recovery is done according to Eq. (6.10). Rð%Þ 5 100 2

Cdialyzed 3 100 Cperfusate

(6.10)

The in vivo microdialysis data are then corrected by the mean data obtained in the in vivo (retrodialysis) recovery studies, according to Eq. (6.11), to determine the actual drug concentration into the skin or which has crossed the skin (depending on where the probe was inserted). Creal 5

Csample 3 100 Rinvivo

(6.11)

One can then construct a graph, relating drug concentration over time, to determine various kinetic parameters, such as area under the curve, maximum permeate concentration, and time required to reach the maximum concentration. Microdialysis has been explored for evaluation of the skin penetration of different nanoparticles (Michalowski et al., 2004; Desai et al., 2012) because it enables the pharmacokinetic and tissue distribution determination of drugs in a complex and complete environment. For example, the influence of the surface modification of NLCs with cell-penetrating peptides in celecoxibe skin pharmacokinetics and tissue distribution was evaluated by Desai et al. (2012) using microdialysis implanted in rat dermis. Zhang et al. (2014) used microdialysis to study the skin permeation and retention of psoralen from ethosomes. The microdialysis probe was implanted in the dermis of the rats to accurately measure drug molecules in the extracellular fluid of the dermal layer (where the ethosomes accumulated). Therefore, the psoralen released from ethosomes and its pharmacokinetic parameters could be determined. Moreover, Kreilgaard (2001) compared microemulsions containing lidocaine and prilocaine with a conventional oil-inwater emulsion-based cream and hydrogel. The cutaneous concentrations of lidocaine

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6.7 CONCLUSION Nanoparticles have been used for improving topical administration of several drugs. However, SC, composed of corneocytes that are well assembled into a complex lipid matrix, hampers nanoparticle skin penetration. Different nanoparticle characteristics, such as size, type, charge, surface modifications, and shape lead to unique mechanisms that enhance drug permeation, as well as influence nanoparticle ability to interact with skin surface and its appendages. The association of physical methods in topical administration of nanoparticles can further change these mechanisms and interactions with the skin. Therefore, to evaluate the influence of nanoparticles in drug skin penetration, in vitro and in vivo methods are essential. The methods detailed in this chapter are the most frequently applied in investigating the influence of nanoparticles on drug penetration into and through the skin. In vitro skin penetration studies can be performed by using different diffusion cells and skin models, providing information about drugs permeation and accumulation in different skin layers, when nanoparticles are used for their administration. Microscopic techniques, such as TEM and confocal laser microscopy provide better elucidation of nanoparticle localization in the skin structures. In vivo permeation studies could predict the real behavior of nanostructured systems in a complex and dynamic environment. Tape stripping and microdialysis techniques can allow the determination of a drug skin permeation profile and its pharmacokinetics. An initial characterization of nanoparticles is also important to understand their influence on drugs penetration and accumulation in the skin so as to further optimize the nanoparticle drug delivery system for topical administration.

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Yamashita, F., Hashida, M., 2003. Mechanistic and empirical modeling of skin permeation of drugs. Adv. Drug Deliv. Rev. 55 (9), 11851199. Ya-Xian, Z., Suetake, T., Tagami, H., 1999. Number of cell layers of the stratum corneum in normal skinrelationship to the anatomical location on the body, age, sex and physical parameters. Arch. Dermatol. Res. 291 (10), 555559. Yong-Tai, Z., Meng-Qing, H., Li-Na, S., Ji-Hui, Z., Nian-Ping, F., 2015. Solid lipid nanoparticles formulated for transdermal aconitine administration and evaluated in vitro and in vivo. J. Biomed. Nanotech. 11 (2), 351361. Zaric, M., Lyubomska, O., Poux, C., Hanna, M.L., McCrudden, M.T., Malissen, B., et al., 2015. Dissolving microneedle delivery of nanoparticle-encapsulated antigen elicits efficient cross-priming and Th1 immune responses by murine langerhans cells. J. Invest. Dermatol. 135 (2), 425434. Zellmer, S., Pfeil, W., Lasch, J., 1995. Interaction of phosphatidylcholine liposomes with the human stratum corneum. Biochim. Biophys. Acta 1237 (2), 176182. Zhai, H., Maibach, H.I., 2001. Effects of skin occlusion on percutaneous absorption: an overview. Skin Pharmacol. Appl. Skin Physiol. 14 (1), 110. Zhai, Y., Yang, X., Zhao, L., Wang, Z., Zhai, G., 2014. Lipid nanocapsules for transdermal delivery of ropivacaine: in vitro and in vivo evaluation. Int. J. Pharm. 471 (1), 103111. Zhai, Y., Zhai, G., 2014. Advances in lipid-based colloid systems as drug carrier for topic delivery. J. Control. Release 193, 9099. Zhang, N., Said, A., Wischke, C., Kral, V., Brodwolf, R., Volz, P., et al., 2016. Poly [acrylonitrile-co-(N-vinyl pyrrolidone)] nanoparticles—Composition-dependent skin penetration enhancement of a dye probe and biocompatibility. Eur. J. Pharm. Biopharm.pii:S0939-6411(16)30734-2. Zhang, L.W., Monteiro-Riviere, N.A., 2013. Use of confocal microscopy for nanoparticle drug delivery through skin. J. Biomed. Opt. 18 (6), 061214. Zhang, Q., Flach, C.R., Mendelsohn, R., Mao, G., Pappas, A., Mack, M.C., et al., 2015a. Topically applied ceramide accumulates in skin glyphs. Clin. Cosmet. Investig. Dermatol. 8, 329337. Zhang, W., Gao, J., Zhu, Q., Zhang, M., Ding, X., Wang, X., et al., 2010. Penetration and distribution of PLGA nanoparticles in the human skin treated with microneedles. Int. J. Pharm. 402 (1), 205212. Zhang, X., Mao, Z., Chen, S., Chen, S., Wang, L., 2015b. Formulation and evaluation of transdermal drug-delivery system of isosorbide dinitrate. Braz. J. Pharm. Sci. 51 (2), 373382. Zhang, Y., Tekobo, S., Tu, Y., Zhou, Q., Jin, X., Dergunov, S.A., et al., 2012. Permission to enter cell by shape: nanodisk vs nanosphere. ACS Appl. Mater. Interfaces 4 (8), 40994105. Zhang, Y.T., Shen, L.N., Zhao, J.H., Feng, N.P., 2014. Evaluation of psoralen ethosomes for topical delivery in rats by using in vivo microdialysis. Int. J. Nanomed. 9, 669678. Zhou, Q., Gallo, J.M., 2005. In vivo microdialysis for PK and PD studies of anticancer drugs. AAPS J. 7 (3), E659E667.

CHAPTER

DNA aptamer-based molecular nanoconstructions and nanodevices for diagnostics and therapy

7

Elena Zavyalova and Alexey Kopylov Chemistry Department of Lomonosov Moscow State University, Moscow, Russian Federation

CHAPTER OUTLINE 7.1 DNA Aptamers in Diagnostics and Therapy.........................................................250 7.2 Basic Principles of DNA Nanoconstruction Creation ...........................................252 7.2.1 Double Helices...............................................................................252 7.2.2 Triple Helices .................................................................................254 7.2.3 G-Quadruplexes..............................................................................255 7.2.4 i-Motifs .........................................................................................256 7.3 DNA Nanoconstructions for Aptamer Oligomerization ..........................................257 7.3.1 Functional Activity of Aptamer Homodimers......................................258 7.3.2 Oligomeric Aptamers in Sensors ......................................................260 7.3.3 Lifetime of Oligomeric Aptamers In Vivo ...........................................261 7.4 DNA Nanoconstructions With Different Aptamers ................................................262 7.4.1 Thrombin Aptamer Hetero-Oligomers................................................263 7.4.2 Aptamer Heterodimers for Cell Targeting...........................................266 7.5 Designing Extensive DNA Nanoconstructions......................................................269 7.6 Examples of Extensive DNA Nanoconstruction Geometry .....................................271 7.6.1 DNA Tiles.......................................................................................271 7.6.2 Two-Dimensional DNA Origami ........................................................271 7.6.3 Three-Dimensional DNA Origami......................................................272 7.7 Promising Applications of DNA Tiles and DNA Origami........................................273 7.7.1 Drug Delivery Systems ....................................................................273 7.7.2 Membrane-Associated DNA Nanoconstructions .................................275 7.7.3 Spatial Arrangement of the Molecules ..............................................276 7.7.4 Biosensors .....................................................................................279 7.7.5 Molecular Machines........................................................................279


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250

CHAPTER 7 DNA aptamer-based molecular nanoconstructions

7.8 Conclusions.....................................................................................................282 References .............................................................................................................282

7.1 DNA APTAMERS IN DIAGNOSTICS AND THERAPY Elaboration of new therapeutic agents is often complicated with toxicity and side effects that are mostly unpredictable at the research stage of drug development. Natural substances are thought to have inherently low toxicity and side effects. Particularly, natural polymers, proteins, and nucleic acids are of great interest as potential therapeutic agents because of their low toxicity and huge diversity; millions of sequences can be generated simply by rearrangement of the building blocks (amino acids or nucleotides), providing the possibility for specific binding of a variety of targets. For diagnostic applications, proteins and nucleic acids can be readily immobilized or conjugated using well known and rather simple chemical manipulations that can be either sequence-specific or nonspecific. This feature is very attractive, as novel compounds often need special approaches for modification, complicating their adaptation for diagnostics. Peptides, enzymes and antibodies have firmly come into usage in diagnostics and therapy during recent decades. Sales of monoclonal antibody-related products reached about $70 billion in 2013, while production volumes reached about 10,000 kg per year (Ecker et al., 2015; BCC Research, 2015). Values for other recombinant proteins are only slightly lower than those of monoclonal antibodies (Ecker et al., 2015). The nucleic acid biopharmaceutical market is much more modest than that of proteins (the aptamer market was $0.34 billion (BCC Research, 2014), whereas the antisense and RNAi market was $0.88 billion in 2014 (Grand View Research, 2016)), but it is also growing rapidly. Although the market for therapeutic nucleic acids is delayed, compared to the protein one, nucleic acids have substantial advantages as potential drugs; in particular, they are low-immunogenic, ex vivo-produced and generally more stable during storage. One of the promising directions in therapeutic nucleic acids is aptamers having high affinity and high specificity to a target, just like monoclonal antibodies. Aptamers are oligonucleotides with a specific spatial structure stabilized, with double helices and canonical and noncanonical base pairs, G-quadruplexes and i-motifs (Zavyalova and Kopylov, 2016). Aptamer spatial structure is simpler than those of antibodies, so it can be easily assembled and renatured. Target binding proceeds through the formation of noncovalent interactions: electrostatic, H-bonding, and hydrophobic. Although the diversity of nucleic acids is inherently lower than that of proteins, thousands of aptamers to hundreds of targets have been reported to date. The most popular targets for aptamers are proteins, although successful examples of aptamers for inorganic cations, low molecular

7.1 DNA Aptamers in Diagnostics and Therapy

weight organic compounds, viruses, and whole cells have been reported (Bouchard et al., 2010). A specific technique has been developed in vitro to select aptamers for a given target. Systematic Evolution of Ligands by EXponential enrichment (SELEX) is the selection of aptamers from a huge oligonucleotide library. Selection is driven by high affinity of an aptamer to a target, i.e., it complies extraction of the most kinetically stable complexes of oligonucleotides with a target that results from competitive binding of all possible oligonucleotides. To make the selection process successful, the library is subjected to several selection cycles. To generate a huge library, oligonucleotides with partially randomized sequences are synthesized with two defined terminal regions for polymerase chain reaction (PCR) primers. The theoretical capacity of the library is estimated as 4N, where N is the number of randomized nucleotides. Under real experimental conditions with limited oligonucleotide amounts, the library is under-represented to a considerable extent, but even if the best binder is absent, SELEX will result in related sequences that could be further improved with rational design. Several selection approaches have been successfully applied to generate aptamers (Bouchard et al., 2010; Darmostuk et al., 2015; Keefe et al., 2010; Lao et al., 2014). The use of capillary electrophoresis provides high separation efficiency of aptamer-target complexes that decreases the number of selection repeats down to between one and four cycles (Darmostuk et al., 2015; Nitsche et al., 2007). Microfluidics (Ahmad et al., 2011), surface plasmon resonance (Misono and Kumar, 2005), and magnetic beads (Darmostuk et al., 2015) are other useful approaches for selection. SELEX-derived initial aptamers are to be optimized in terms of aptamer size and functional activity. Optimization includes the removal of extra regions, a search for consensus sequences in a pool of selected aptamers, and mapping the aptamer fragments involved in protein binding. The nucleic acid nature of aptamers provides unique properties. For example, aptamers have rational antidotes: a complementary oligonucleotide disrupts the unique aptamer spatial structure, forming a nonfunctional double helix. Aptamers can be readily modified chemically, improving pharmacokinetic and pharmacodynamic properties for therapy, as well as providing immobilization ability for diagnostics (Acquah et al., 2015; Bouchard et al., 2010). Moreover, modifications can considerably enhance aptamer stability (Sagi, 2014) and affinity to the target (Park et al., 2013). For example, aptamer derivatives, slow off-rate modified aptamer (SOMAmers), combine the advantages of classical aptamers and the diversity of amino acid side chains. SOMAmers have heterocyclic bases that are extended with additional hydrophobic groups, thus creating a surface for better target binding (Gold et al., 2012; Jarvis et al., 2015; Kraemer et al., 2011; Park et al., 2013). Fluorescent labels can be easily attached to oligonucleotides to improve the detection limit in diagnostic applications.

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Generally, aptamers are promising molecular recognition elements (MoRE). High affinity and high specificity ensure successful usage in therapy and diagnostics. Dozens of aptamers have already entered clinical trials, yielding only one marketed aptamer so far, pegaptanib. Implementation of this class of novel drug proceeds with caution. Several aspects have a high impact on the advancement of aptamers, but most of them seem to be a consequence of a lack of preclinical and clinical data to date, which must be overcome in the near future (Baird, 2010; Zavyalova, 2015). The diagnostic field is more predictable than therapeutic applications; “aptasensors” have already demonstrated an ability to detect and estimate the content of amino acids, a huge variety of proteins, viruses, bacteria, cancer cells, etc. (Meirinho et al., 2016; Piro et al., 2016; Saberian-Borujeni et al., 2014). Besides practical applications, aptamers are of great interest for fundamental science as small, artificial, three-dimensional structures with high affinity and specificity to a target. This chapter focuses on nanoconstructions and nanodevices that are functionalized with DNA aptamers. By using the term “nanoconstruction”, we mean a complex of several nucleic acid strands that are connected through noncovalent interactions. In speaking of a “nanodevice”, we mean a nanoconstruction that is able to provide a function in response to a signal. Aptamers target nanoconstructions to the desirable location and provide signal-response ability to nanodevices.

7.2 BASIC PRINCIPLES OF DNA NANOCONSTRUCTION CREATION Assembly of DNA nanoconstructions requires locks; specific noncovalent interactions between distinct nucleic acid strands. A few possibilities are known for DNA: formation of double helices, triple helices, G-quadruplexes, and i-motifs (Fig. 7.1). Here, all the types of structure are discussed in terms of possible topology and stability.

7.2.1 DOUBLE HELICES The most obvious and widely used way to assemble distinct DNA strands into a desirable structure is the formation of a double helix between the complementary regions of the strands. This type of lock has several peculiarities to keep in mind: • •

Topology: The two strands in the double helix always have opposite directions (Fig. 7.1A). Interactions: A double helix is formed due to hydrogen bonds; Watson-Crick base pairs are stacked due to the hydrophobic effect. Complementarity can be incomplete with several mismatches or noncanonical base pairs, but the stability of a nonideal double helix is reduced.

7.2 Basic Principles of DNA Nanoconstruction Creation

FIGURE 7.1 Interactions and topologies of DNA locks: (A) double helix lock; (B) triple helix lock; (C) i-motif locks; (D) G-quadruplex locks. 50 and 30 -ends, as well as the PDB ID of the structures, are indicated. The individual strands are shown in different colors; loops are shown in yellow.

•

Conformation uniformity: Interactions are unambiguous if the sequences of the complementary regions in the strands are unique. The availability of similar sequences will result in the formation of several types of structure.

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Table 7.1 A Comparison of the Properties of Different DNA Locks Double Helix

Triple Helix

G-Quadruplex

i-Motif

2

3

2N

24

11 1

11 1

1

1

Opposite None

Opposite H1, Mg21

Any H1

10G/C20 A/T (1)

9Na (2, 3)

Any Cations (K1, NH41, Sr21, etc.) 2 strands: 3G3N3G (6) 4 strands: 3G (7)

Lock length for Tm B60 C

15G/C30 A/T (1)

1011Na (4, 5)

2 strands: 4G4N4G (8) 4 strands: 4G (9)

Coexistence of several locks

11 1

11

Number of DNA strands Conformation uniformity Strand orientation Required stabilizing component Lock length for Tm B37 C

2 strands: 3C3N3C (10) 4 strands: 3C (11, 12) 2 strands: 4C4N4C (13) 4 strands: 5C (14)

References: (1) Kibbe (2007); (2) Mills et al. (1999); (3) Mills et al. (2002); (4) Scaria and Shafer (1996); (5) He et al. (1997); (6) Balagurumoorthy and Brahmachari (1994); (7) Bardin and Leroy (2008); (8) Lu et al. (1993); (9) Hardin et al. (1997); (10) Kaushik et al. (2010); (11) Kanaori et al. (2001); (12) Kaushik et al. (2007); (13) Miyoshi et al. (2003); (14) Mergny et al. (1995). a () Highly sequence specific.

•

•

Structure stability: Double helices are stable in buffers, routinely used in biochemistry. The structure exists between the pH of protonation and deprotonation of nucleotides. Structure stability is increased with helix length and the amount of GC base pairs (Table 7.1). Coexistence of several locks: Due to the high sequence specificity of interactions in a double helix, dozens of different locks can coexist in a single structure. This peculiarity provides the possibility of “DNA origami” construction that will be discussed in detail later on.

7.2.2 TRIPLE HELICES If a double helix is formed with homopurine (A/G) and homopyrimidine (T/C) strands, it can be complemented with a third strand, yielding a triple helix (Fig. 7.1B). The third strand forms hydrogen bonds with the homopurine strand, fitting into the major groove of the double helix (Fox and Brown, 2011; Sugimoto, 2014).

7.2 Basic Principles of DNA Nanoconstruction Creation

•

•

•

•

•

Topology: The two strands in the double helix have the opposite directions, whereas the third strand can be oriented in either a parallel or antiparallel direction, relative to the purine strand. Interactions: The triple helix is formed due to hydrogen bonds; Hoogsteen or reverse Hoogsteen base triplets are stacked due to the hydrophobic effect. In the parallel motif, GC is recognized by protonated cytosine (C1GC) or guanine (GGC), while AT is bound to thymine (TAT). In the antiparallel motif, GC is bound to guanine (GGC), while AT is recognized by adenine (AAT) or thymine (TAT) (Duca et al., 2008). Conformation uniformity: The interactions are unambiguous as they are sequence-specific. The availability of similar purine-rich sequences will result in the formation of several types of structure. Structure stability: The stability of triple helices largely depends on buffer composition: slightly acidic pH, magnesium cations, polyethylene glycol, polyamines, etc. increase stability. Parallel triplexes are usually more stable than antiparallel ones, especially at slightly acidic pH (Duca et al., 2008; Fox and Brown, 2011; Sugimoto, 2014). Coexistence of several locks: Due to the sequence specificity of interactions, different locks can coexist in a single structure.

7.2.3 G-QUADRUPLEXES A guanine quadruplex is a noncanonical type of structure, formed by DNA having guanine blocks. Contrary to double helices, intermolecular G-quadruplexes can be of multiple topologies (Fig. 7.1D). This particular structure is mostly nonunambiguous; intra- and intermolecular topologies can be formed from molecules with the same sequence, depending on conditions such as strand concentration, buffer composition, and temperature profile during structure assembly (Sannohe and Sugiyama, 2010). •

Topology: In a lock, the number of strands can be two, four, eight, or more. For bimolecular structures, five topologies are shown in Fig. 7.1D; they have different locations of the 50 and 30 -ends, providing a variety of possibilities for nanoconstruction design. Tetramolecular structures generate parallel Gquadruplexes with four 50 -ends located on one side of the structure and four 30 -ends located on the opposite side. In octamolecular structures, two parallel tetramolecular G-quadruplexes are stacked, so that eight 50 -ends are placed in the center of the structure, whereas the 30 -ends are at the surface of the construction. One more type of structure is the G-wire, an extensive parallel G-quadruplex formed with multiple strands in an overlapping manner (see scheme in Fig. 7.1D). The molecularity of G-wires is undefined and is largely determined by the sequence and the conditions during structure assembly (temperature profile, buffer composition, concentration of DNA strands).

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256


•

•

•

•

Interactions: G-quadruplexes are formed from several stacked G-quartets, due to the hydrophobic effect. Each G-quartet has four guanines bound together with Hoogsteen hydrogen bonds, so that four oxygen atoms are directed into the center of the structure. The proximity of the oxygen atoms creates a structure tension that can be diminished with a cation being coordinated in the G-quartet center. As most cations are too large for the G-quartet structure and have coordination numbers more than 4, they occupy the central site between two G-quartets, being coordinated with eight oxygen atoms. Conformation uniformity: In most cases, several topologies can be generated from a molecule with the same sequence, so interactions are less specific than double helices. Several factors can shift the equilibrium to the desirable structure. Among them are strand concentration, coordinated cations, and the temperature profile during structure assembly. Structure stability: Generally, G-quadruplexes are more stable than double helices of comparable length. G-quadruplex stability is increased with the number of G-quartets in the structure, the number of intermolecular interactions, optimal loop length that diminishes structure tension, optimal nucleotides that are adjacent to the G-quadruplex (and shield it from the solvent), and good cation coordination inside the structure (K1, Sr21, Ba21, etc.). Coexistence of several locks: As the sequence specificity of interactions is much lower than double helices, the coexistence of different types of lock is generally complicated in a single structure. However, a diversity of lock molecularity and topology provides some preferences, compared to double helix locks.

7.2.4 I-MOTIFS One more interesting noncanonical DNA structure is the i-motif (Fig. 7.1C). It is formed by cytosine blocks at slightly acidic pH. The i-motif represents two intercalated duplexes but its stability and structure assembly are similar to G-quadruplexes. •

•

Topology: In the lock, the number of strands can be two or four. The i-motif is formed by two parallel duplexes with opposite directions and cytosine pairs intercalate with each other, forming a tetra-stranded structure. Both in bimolecular and tetramolecular structures, 50 and 30 -ends are located at the same end of the structure, similar to double helices and contrary to parallel G-quadruplexes. Two bimolecular structures are shown in Fig. 7.1C. In the first structure, two DNA ends are on one side of the structure, and the other two are on the other side. In the second structure, all four ends are located on the same side. In the tetramolecular structure, two 50 and two 30 -ends are located both at the top and at the bottom of the structure. Interactions: i-Motifs consist of two intercalated parallel duplexes, with opposite direction due to the hydrophobic effect. Each parallel duplex is

7.3 DNA Nanoconstructions for Aptamer Oligomerization

•

•

•

stabilized with pairs of cytosine-protonated cytosine, due to hydrogen bonding. The structure formation requires slightly acidic pH. Conformation uniformity: As there are only a small number of possible topologies, conformation uniformity is possible when the conditions for structure assembly are optimized. Structure stability: At slightly acidic pH, the stability of i-motifs is similar to those of G-quadruplexes. Stability has a bell-shaped dependence on pH: the first transition point is B4.3, being the pK of cytosine protonation; the second transition point depends heavily on the exact structure, being in the range of 5.07.5. The most stable structures have five or more cytosine repeats in the block. Besides protons, the structure can be stabilized with Ag1 ions. Stability is higher with an increase of DNA strand concentration, as well as with optimal loop length diminishing structure tension and optimal nucleotides that are stacked with i-motif ends. Coexistence of several locks: Because the sequence specificity of interactions is much lower than double helices, the coexistence of different types of lock is generally complicated in a single structure.

Double helices, triple helices, G-quadruplexes and i-motifs can be used as locks in DNA nanoconstructions. To date, the double helix is the most abundant lock in both small nanoconstructions and large systems like DNA origami. G-quadruplex locks are only used in several small structures, whereas others have not yet been applied. Constructions with double helix and G-quadruplex locks will be discussed in detail below.

7.3 DNA NANOCONSTRUCTIONS FOR APTAMER OLIGOMERIZATION Aptamers are of great promise in therapy and diagnostics. Multiple ways to improve aptamer properties are being developed. Chemical modifications to the DNA backbone improve aptamer stability in vivo; additional functional groups in heterocyclic bases increase affinity to the target. Conjugation with polyethylene glycol, proteins, and other polymers as well as with nanoparticles, liposomes and micelles increases aptamer lifetime in vivo and affects biodistribution within the body (Bruno, 2013; Dickgiesser et al., 2015; Gao et al., 2016; Grijalvo et al., 2014; Heo et al., 2016; Jo and Ban, 2016). DNA nanoconstructions can also provide additional features to aptamers, retaining their chemical nature and thus being fully biocompatible. Among the desirable and achievable features are: enhanced affinity to the target, a high density of recognizing elements in aptasensors, an increase of aptamer stability in biological fluids, and modified biodistribution. Several studies have been made in this direction; we will discuss representative results in this subsection.

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7.3.1 FUNCTIONAL ACTIVITY OF APTAMER HOMODIMERS Linking several aptamers in the entire structure provides a high density of recognizing elements for sensors, together with high affinity. Examples with both covalent and noncovalent linking have been described for several aptamers that allows a direct comparison. Thrombin-binding aptamer, HD1, targets the distal substrate binding site of thrombin and inhibits hydrolysis of several coagulation proteins (Zavyalova et al., 2016a; Zavyalova and Kopylov, 2016). Constructions with two covalently linked aptamers are expected to have 1.22.7-fold higher inhibitory activity, compared to the initial aptamer (Table 7.2): HD1THD1 (Zavyalova et al., 2013), HD1TTTHD1 (Zavyalova and Kopylov, 2016), and circular aptamer cTT2 (Fig. 7.2) (Di Giusto and King, 2004). Noncovalent linking DNA locks provide a similar effect: both in the case of a homodimeric aptamer with a double helix lock (FF aptamer (Neundlinger et al., 2011)), and a G-quadruplex lock (Fig. 7.2) (authors’ unpublished data). The inhibitory activity is twice as high as that of the HD1 aptamer (Table 7.2). A much more exciting effect has been achieved for dimeric aptamers to dimeric targets, due to the cooperative binding mode. A covalently linked dimeric aptamer to vascular endothelial growth factor, VEGF165, has a 20-fold higher affinity for the dimeric protein, compared to the monomeric protein (Hasegawa et al., 2008). Another interesting example is aptamer SL1 to the Met receptor. Met is a phosphokinase receptor protein that forms a dimer after binding to hepatocyte growth factor. Monomeric aptamer, SL1, binds Met but does not induce substantial autophosphorylation, whereas both the covalently linked SL1 dimer and the dimer with a double helix lock (Fig. 7.3) trigger autophosphorylation with comparable efficiency (Ueki et al., 2016).

FIGURE 7.2 Homo-oligomeric DNA aptamers to thrombin with covalent and noncovalent linking.


Table 7.2 Homodimers of Thrombin Aptamers With Covalent Linking and Noncovalent DNA Locks Inhibitory Activity Compared to HD1

Refs (1) (2) (3)

Aptamer

Sequence

HD1 HD1THD1 HD1TTTHD1 cTT2

ggttggtgtggttgg ggttggtgtggttggTggttggtgtggttgg ggttggtgtggttggTTTggttggtgtggttgg

1 2.0 1.2 2.7

FF GL1HD1 GL2HD1

A15T5ggttggtgtggttgg 1 ggttggtgtggttggT5T15 ggttggtgtggttggTTTTGGGTTTTGGG ggttggtgtggttggTTTTGGGGTTTTGGGG

2.0 2.0 2.0

(4) a a

References: (1) Zavyalova et al. (2013); (2) Zavyalova and Kopylov (2016); (3) Di Giusto and King (2004); (4) Poniková et al. (2011). Aptamer sequences are shown in lowercase font; linking sequences are shown in uppercase font. a () Authors’ personal data.

FIGURE 7.3 Dimerization of receptor protein Met with hepatic growth factor (HGF) protein (A) and dimeric DNA aptamer (B).

A similar effect has been shown for a RNA nanoconstruction made up of two aptamers to the OX40 receptor connected with a third strand via two double helix locks. OX40 is expressed on the surface of activated T cells; OX40 activation leads to T cell proliferation and cytokine production. Contrary to the monomeric aptamer, the dimeric aptamer construction is able to bind and activate T cells (Dollins et al., 2008).

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Generally, the functional activity of dimeric aptamers with noncovalent links is similar to those of covalent dimers; however, noncovalent dimers have an obvious advantage in manufacturing, as the yield of the proper aptamer during chemical synthesis is greatly diminished when the length is increased.

7.3.2 OLIGOMERIC APTAMERS IN SENSORS Oligomeric constructions provide a low limit of detection in aptasensors, due to a high density of recognizing elements. The interaction of an oligomeric aptamer with a target is affected by increased effective concentration of the aptamer near the target. The subsequent interaction of the unbound target becomes intramolecular, contributing to the increase in affinity by decreasing the dissociation rates. In addition to increased affinity, higher selectivity in target recognition is provided (Musumeci and Montesarchio, 2012). A variety of approaches have been proposed to achieve the lowest limit of detection (Sassolas et al., 2011). Considering thrombin aptasensors as the most popular conventional systems, the limit of thrombin detection is in the range of 102151028 M (Sassolas et al., 2011); aptasensors with aptamer oligomeric nanoconstructions have an advantage over those with a monomeric aptamer. The limit of detection of an aptasensor with an FF aptamer is 3 3 10210 M, threefold lower than that of aptamer HD1 (Hianik et al., 2008). Similar nanoconstructions with double helix locks were built from three and four strands, with aptamers (Sung et al., 2013). Zhang and Yadavalli (2012) also proposed oligomeric aptamer nanoconstructions: a VEGF-aptamer trimer and a thrombin-aptamer tetramer. The constructions are proposed to be useful in both aptasensors for ligand recognition and protein ordering for reaction catalysis or ligand binding. More extensive structures were made from dozens of repeated, partially complementary elements. DNA lattices with periodically located thrombin aptamers were produced. The structures are coined “DNA tiles,” and they are twodimensional DNA arrays with periodic spacing between neighboring aptamers. Lin et al. (2006) showed a period of 27 nm between aptamers; high capacity for thrombin-binding was also demonstrated. The principles of DNA tile construction will be discussed further. Besides the discussed examples, a variety of sophisticated nanoconstructions have been developed, utilizing DNA and other compounds, to gain high sensitivity (Musumeci and Montesarchio, 2012). Several thorough reviews have been dedicated to aptasensors (Li et al., 2010; Meirinho et al., 2016; Piro et al., 2016; Saberian-Borujeni et al., 2014). As constructions of a mixed chemical nature are out of the scope of this chapter, just a single example is provided as proof of the principle. A thrombin-binding nanoconstruction has been shown to have a limit of detection of 4 3 10211 M (Cheng et al., 2010). The construction was made from magnetic nanobeads functionalized with multiple DNA oligonucleotides, which form double helices with aptamer-containing hairpin oligonucleotides, modified

Current (A)


Incubation

E (V)

Electrochemical signal on

Electrochemical signal off OH

O OH

HO HO HO

H3C

Magnetic bead

OH

OH O OH

OH

O

Carminic acid

CH3

CA dimer

O

Thrombin

CAs-MB

FIGURE 7.4 Aptasensor with a high density of recognizing elements. Magnetic nanobeads are functionalized with thrombin aptamers. Source: Reprinted from Cheng, G., Shen, B., Zhang, F., Wu, J., Xu, Y., He, P., et al., 2010. A new electrochemically active-inactive switching aptamer molecular beacon to detect thrombin directly in solution. Biosens. Bioelectron. 25(10), 22652269, © (2010), with permission from Elsevier.

with an anthraquinone derivative. Thrombin-binding causes dissociation of the anthraquinone dimer and the appearance of electrochemical activity (Fig. 7.4). High sensitivity is achieved, due to the presence of multiple aptamers on the nanobead surface and electrochemical signal registration.

7.3.3 LIFETIME OF OLIGOMERIC APTAMERS IN VIVO Oligomerization of aptamers can be used to affect the pharmacokinetics and pharmacodynamics of the aptamer in vivo, as these properties are strongly affected by the molecular weight of the molecule. The two main reasons for poor pharmacokinetics and pharmacodynamics of nonmodified aptamers are nuclease digestion in biological tissues and fast renal clearance from the bloodstream. Nuclease digestion can be diminished due to extensive nucleotide modifications, whereas conjugation with polyethylene glycol greatly decreases renal clearance (Fishburn, 2008; Grijalvo et al., 2014). Polyethylene glycol lifetime in the bloodstream depends on the molecular weight of the macromolecule; polymers of 20 kDa and higher totally exclude renal excretion (Bailon and Won, 2009; Fishburn, 2008; Grijalvo et al., 2014). Oligomerization of nonmodified aptamers slightly affects the efficiency of nuclease digestion, but can be a good alternative to conjugation

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with polyethylene glycol. Proof of the concept has been demonstrated recently for thrombin aptamers connected with G-quadruplex locks (authors’ unpublished data). Although intermolecular G-quadruplexes have already been thoroughly studied (Hardin et al., 2001; Keniry, 2001), their usage as DNA locks is limited to a few examples. Tetramolecular parallel G-quadruplexes were used to conjugate four peptidomimetics that bind and inhibit chymotrypsin (Cai et al., 2009) and to create two peptide loops that imitate loop positioning in proteins (Ghosh and Hamilton, 2012). Here, we provide the first data for aptamer-based nanoconstructions with G-quadruplex locks. These constructions have been recently proposed by the authors as an alternative way to prolong the aptamer lifetime in vivo. Two HD1 aptamers were linked with dimeric intermolecular G-quadruplexes (Fig. 7.2). The predicted topology of DNA locks is that these nanoconstructions form a dimeric antiparallel G-quadruplex with diagonal loops (topology of 4R45 type, Fig. 7.1D). As both aptamer HD1 and DNA locks are G-quadruplexes, there is an obvious complication in the conformational uniformity of the entire nanoconstruction. Substitutions of thymines with adenines in the lock loops resulted in mixed folding, and gave a decrease in functional activity, due to disruption of the tertiary structure of aptamer HD1. Shortened loops resulted in the formation of G-wires about 400 nm long (data from high-resolution atomic force microscopy, performed by A. Protopopova and D. Klinov from the Scientific Research Institute of Physical-Chemical Medicine, Moscow, Russia; unpublished data). The most successful examples, the dimeric aptamers GL1HD1 and GL2HD1, retain high inhibitory activity (Table 7.2), and have modified pharmacodynamics in rats. Expected molecular masses of these nanoconstructions are about 20 kDa, fourfold higher than that of aptamer HD1, whereas the half-life times of intravenously administered aptamers in the bloodstream are about 4060 min (Fig. 7.5), considerably higher than the half-life of 4 min for aptamer HD1 (Zavyalova et al., 2016b) or unlocked control aptamer (Fig. 7.5) (animal experiments were performed by N. Kust and G. Pavlova from the Institute of Gene Biology RAS, Moscow, Russia; unpublished data). These very first data encourage further search for DNA nanoconstructions that provide increased lifetime and modified biodistribution of aptamers in vivo.

7.4 DNA NANOCONSTRUCTIONS WITH DIFFERENT APTAMERS Hetero-oligomeric aptamer nanoconstructions possess unique properties that cannot be achieved with homo-oligomeric aptamers. The main advantage is an extremely high affinity, due to multicenter binding to the target. In aptasensors,

7.4 DNA Nanoconstructions With Different Aptamers

FIGURE 7.5 Pharmacodynamics of dimeric thrombin aptamers with G-quadruplex locks. The aptamers were administered intravenously in rats in a 2.5 mg/kg dose; a thrombin time test was carried out for plasma blood samples, taken at different time intervals after injection. Control aptamer is monomeric HD1T15; GL2HD1 is ggttggtgtggttggTTTTGGGGTTTTGGGG; GL1HD1 is ggttggtgtggttggTTTTGGGTTTTGGG (authors’ unpublished data).

high affinity implies increased specificity and sensitivity (Musumeci and Montesarchio, 2012). Thrombin aptamers are conventional models in this field, as at least four DNA aptamers and one RNA aptamer target thrombin exosite I, the protein substrate-binding site, whereas at least one DNA aptamer and one RNA aptamer target thrombin exosite II, the heparin-binding site (Zavyalova and Kopylov, 2016). The existence of diverse aptamers provides a possibility to make various bivalent constructions and to easily assess their inhibitory activity by tracking hydrolysis of the enzyme substrates. Another trend, currently under development, is bispecific aptamers that target marker proteins of different cells; these perspectives are also discussed in this subsection.

7.4.1 THROMBIN APTAMER HETERO-OLIGOMERS Nanoconstructions made up of a thrombin aptamer to exosite I and an aptamer to exosite II have superior nonadditive functional activity, when the linker length is comparable with the distance between the two thrombin exosites. Most nanoconstructions contain DNA aptamers HD1 and HD22. Both aptamers have antiparallel G-quadruplex topology, and do not interfere with each other when joined with

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an extended linker. Contrary to aptamer HD1, aptamer HD22 targets thrombin exosite II, and is unable to inhibit the key thrombin functional activities (Lin et al., 2011; Russo Krauss et al., 2013; Tasset et al., 1997). Therefore, the effect of conjugation of aptamers HD1 and HD22 can be easily estimated from the increase of inhibitory activity (Table 7.3). Covalent junction of two aptamers with an oligo-adenosine linker (A15) yielded aptamer HD122 (Muller et al., 2007, 2008). Its affinity to thrombin is very high (Kd 5 0.5 nM) (Muller et al., 2008) and it is substantially higher than the value for the initial aptamer HD1 (Kd 5 716 nM) (Muller et al., 2008; Zavyalova and Kopylov, 2016). The functional activity of aptamer HD122 is also higher than that of aptamer HD1 in coagulation and platelet aggregation tests for plasma and blood samples (Muller et al., 2008). Several similar covalently linked constructions were described. A polythymidine linker (T35) has an affinity of Kd 5 0.4 nM (Rakhmetova et al., 2010); shorter oligothymidine linkers were also described (Hasegawa et al., 2008). A flexible linker, made up of phosphoramidite spacers (Sn), gives maximal, ninefold inhibitory activity for aptamer Bi8S with eight spacers (S8) (Ge et al., 2012; Kim Table 7.3 Heterodimers of Thrombin Aptamers With Covalent Linking and Noncovalent DNA Locks Inhibitory Activity Compared to HD1

Refs

Aptamer

Sequence

HD1 HD22 HD122 A1(35)A2

ggttggtgtggttgg agtccgtggtagggcaggttggggtgact ggttggtgtggttggA15agtccgtggtagggcaggttggggtgact ATGTCTACTggttggtgtggttggGTAGT15 TCagtccgtggtagggcaggttggggtgact ggttggtgtggttggS8gtccgtggtagggcaggttggggtgac ggttggtgtggttgg (Spacer)10gtccgtggtagggcaggttggggtgac CS1ggttggtgaggttggCS10 CS2agtccgtggtagggcaggttggggtgactCS20 A15T5ggttggtgtggttgg 1 ggtagggcaggttggT5T15

1 , 0.2 20 27

(1) (2) (3)

9 130

(1) (4)

300

(5)

3 8

(6) (7)

ggttggtgtggttggTTTTTTTCTACAGGGTA 1 CACAAATTCGGTTTTTgtccgtggtagggcaggttggggtgac 1 TACCCTGTAGAACCGAATTTGTG

(8)

Bi8S A2(24 nm)A1 TBV08 FH 3HTBAAA

DNA150 270 pm

References: (1) Kim et al. (2008); (2) Muller et al. (2007); (3) Rakhmetova et al. (2010); (4) Tian and Heyduk (2009); (5) Ahmad et al. (2012); (6) Poniková et al. (2011); (7) Rangnekar et al. (2012); (8) Chen et al. (2011). Aptamer sequences are shown in lowercase font; linking sequences are shown in uppercase font.


FIGURE 7.6 Hetero-oligomeric DNA aptamers to thrombin with covalent and noncovalent linking.

et al., 2008). A pronounced effect has also been shown for linkers made up of ten spacer 18 units, giving an affinity of Kd 5 0.03 nM for aptamer A2(24 nM)A1 (Tian and Heyduk, 2009). A more rigid, double helix linker was used in aptamer TBV08 (Fig. 7.6); the linker has two connected double helices: CS1ggttggtgaggttggCS10 CS2agtccgtggtagggcaggttggggtgactCS20 , where CSi and CSi’ are complementary sequences. The aptamer has extremely high affinity and inhibitory activity: Kd 5 0.008 nМ and Ki 5 0.03 nМ (Ahmad et al., 2012). Noncovalent nanoconstructions are based on double helix locks. Aptamer FH with an A15T15 double helix lock between aptamer HD1 and the G-quadruplex part of aptamer HD22 has a higher inhibitory activity than its homo-oligomeric counterpart, aptamer FF; aptamer FH is also preferential to aptamer FF for sensor application (Hianik et al., 2009; Ponikova´ et al., 2011). A third DNA strand can be used to connect aptamers HD1 and HD22 (Fig. 7.6), and the construction strongly inhibits thrombin (Chen et al., 2011). All these aptamers represent an alternative to the aptamer TBV08 that comprises double helix covalent linking. Contrary to aptamer TBV08, noncovalent nanoconstructions with double helix locks are much simpler in synthesis but possess lower inhibitory activity. A more sophisticated design of linker type and length for noncovalent heterodimeric aptamers is required to improve functional activity. A more sophisticated nanoconstruction contains a double helix lock of the DNA tile type. Two DNA strands with complementary regions make between one and five double helix elements. Four ends of the strands can be modified with DNA aptamers. One HD22 with one to three HD1 aptamers were joined by a DNA tile (Fig. 7.7). The asymmetric rigid structure of the DNA tile allows placement of the aptamers at programmed positions in the DNA nanoconstruction in

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FIGURE 7.7 Thrombin molecule (A) and a hetero-oligomeric construction with a DNA tile lock (B) proposed by Rangnekar et al. (2012). Source: Reprinted from Rangnekar, A., Zhang, A.M., Li, S.S., Bompiani, K.M., Hansen, M.N., Gothelf, K.V., et al., 2012. Increased anticoagulant activity of thrombin-binding DNA aptamers by nanoscale organization on DNA nanostructures. Nanomedicine 8(5), 673681, © (2012), with permission from Elsevier.

order to optimize spacing and orientation of the aptamers, thereby maximizing anticoagulant activity in functional assays. The most successful example of a DNA nanoconstruction, aptamer 3HTBAAA, has inhibitory activity toward thrombin eightfold higher than aptamer HD1 (Rangnekar et al., 2012).

7.4.2 APTAMER HETERODIMERS FOR CELL TARGETING Besides multicenter binding of a single protein, aptamer heterodimers can be applied for multicenter binding of a particular cell, or for connecting different cells. Aptamers to various protein markers of tumors have been developed and characterized. Several studies have been devoted to recognizing different types of cell with aptamer nanoconstructions. Here, we will discuss the advantages of heterodimeric aptamer nanoconstructions. In addition, approaches for the conjugation of different cells will be discussed. Monomeric aptamers can recognize cancer cells and deliver drugs. However, cancer subtypes have heterogeneous biomarkers and a single aptamer is unable to specifically target tumor cell types. DNA nanoconstructions with bispecific aptamers can address heterogeneity among cancer subtypes for targeted drug delivery (Gopinath et al., 2016; Lyu et al., 2016; Soldevilla et al., 2016b). A drug carrier with extended specificity has DNA aptamers to three different cancer cell lines; aptamers are joined with covalent (polyethylene glycol, oligothymidine) linker or noncovalent linkers (double helix locks). Bispecific and trispecific aptamer nanoconstructions with a double helix lock are interesting as drug carriers, because


hydrophobic drugs, e.g., doxorubicin, intercalate readily between adjacent GC or CG base pairs; this binding is reversible, and the entire construction can be applied to treat cancer cells (Zhu et al., 2012). These structures serve as “nanotrains”, aptamer-based DNA nanostructures with intermolecular double helices with intercalated drugs. Here, the aptamer is a recognizing element, whereas the double helix is a drug-carrying element (Zhu et al., 2013). These nanoconstructions have several significant features: (1) easy assembly from several strands based on complementary interactions; (2) high capacity for drug intercalation; (3) reduction of the cost of DNA synthesis due to the use of short DNA; (4) the aptamer-based structure allows specific targeting in cancer therapy; (5) imaging agents can be coupled to visualize the targeted cells; and (6) alteration of the aptamers provides the possibility of targeting different cell types. Another way to design a drug carrier is to conjugate double helices to nanoparticles, to functionalize them with aptamers using double helix locks, and to load the nanostructure with intercalated drug (Zheng et al., 2013). This oligomeric construction can readily provide multicenter binding of a single cell or different cells. Another direction of aptamer-based cancer treatment is to induce cellular cytotoxicity, promoting the antitumor functional activity of immune cells. The same mechanism plays a pivotal role in antibody-based tumor therapies that recruit natural killer cells to antibody-bound tumor cells through the specific receptor CD16. An aptamer-based analogue, an heterodimeric DNA aptamer, combines the aptamer with CD16α, the receptor on natural killer cells and Met, the membrane phosphokinase that is overexpressed in many tumors. Two aptamers were covalently linked with single-stranded DNA of various length and sequence, providing several constructions with high affinity to both targets. Most constructions enabled both retained affinity, and simultaneously targeted protein binding. The length of linkers approximately corresponded to the distance between complementarity-determining regions and the Fc binding domain in entire antibodies; these aptamers imitate antibodies both in the principle of action and functional activity (Boltz et al., 2011). The heterodimeric DNA aptamer, MRP1CD28, binds to both multidrug resistant-associated protein 1 (MRP1) and CD28 protein. The aptamer was made through covalent linking of two aptamers, with a double helix linker similar to that used in aptamer TBV08. It executes intercellular interactions, having affinity to different types of cell. The aptamer recognizes chemotherapy-resistant MRP1 cancer cells and mediates its binding to T lymphocytes through CD28 protein. As a result, T lymphocytes are stimulated, and immunity to tumor cells is promoted. The functional activity of covalently linked aptamers is 1.52.0-fold higher than that of a mixture of the single aptamers (Soldevilla et al., 2016a). Despite encouraging functional data, this aptamer is more than 200 nucleotides long, too long for efficient chemical synthesis; therefore, PCR was used to produce the nanoconstruction. In contrast, the use of DNA locks instead of covalent linking could be a powerful tool in this case.

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One more example of artificial intercellular interactions involves dimeric, trimeric, and tetrameric DNA aptamer nanoconstructions, where the aptamers are connected with double helix locks. Aptamers to markers of T cells and B cells were used to provide specific cellcell interactions that are significant for the immune response. Several oligomeric structures were designed and tested for functional activity. Double helix locks are rigid structures; single-stranded insertions were shown to increase the flexibility and, thus, affinity of the oligomeric structures. Homodimeric constructions were assembled using a third strand with sequences complementary to the ends of the two aptamers. Homotrimeric constructions utilize a Holliday junction type lock. Homotetrameric constructions were created using a DNA tile (Fig. 7.8). Bispecific aptamers were constructed by connecting two different homotetrameric structures with double helix locks between the ends of DNA tiles. A large number of aptamers were united in a single DNA nanoconstruction. Thirteen aptamers to T cells and 13 aptamers to B cells were conjugated using an extensive DNA tile (Fig. 7.8). Cellcell interactions, mediated with multivalent, bispecific aptamer nanoconstructions, were of comparable efficiency to those achieved with bispecific antibodies (Liu et al., 2011).

FIGURE 7.8 Dimeric (A), trimeric (B), tetrameric (C) and oligomeric aptamer nanoconstructions for binding to T cells and B cells. Source: Reprinted from Liu, X., Yan, H., Liu, Y., Chang, Y., 2011. Targeted cellcell interactions by DNA nanoscaffold-templated multivalent bispecific aptamers. Small 7(12), 16731682, © (2011), with permission from Wiley.

7.5 Designing Extensive DNA Nanoconstructions

7.5 DESIGNING EXTENSIVE DNA NANOCONSTRUCTIONS Large DNA nanoconstructions occupy a particular place in nanobiotechnology. A construction of any topology and size can be generated using a set of DNA strands. DNA tiles and DNA origami can be of any size and shape, being fully programmable. Functionalization of these structures with aptamers provides nanoconstructions with unique properties and capacities. In this subsection we will discuss the principles of DNA origami and DNA tile construction; specific examples will be reviewed later on. DNA origami and DNA tiles are artificial constructions, built with several different DNA strands that have partial complementarity to each other. Because DNA strands are asymmetric and have four types of nucleotide, the variety of possible sequences of N-mer oligonucleotides is 4N. High variability of the strands allows a huge variety of nanostructures to be made. Created double helices are highly ordered, due to junctions between different helices. The existence of various junctions and strict rules for interactions between the strands provides an opportunity for precise design of DNA nanostructures (Jabbari et al., 2015; Kuzuya and Komiyama, 2010). A DNA tile is made of several strands with a length of several hundred nucleotides and partial complementarity. In the DNA tile, the sequences of the strands are designed to be very different from one another so the competitive formation of alternative secondary structures is minimized. A minimal amount of strands are in so-called weave tiles, where two strands are interlaced back and forth into periodic architectures (see an example in Fig. 7.7). DNA tiles are often used as building elements for the next level of hierarchical organization. Sticky ends provide linking of the tiles, resulting in the formation of extended periodic lattices or three-dimensional DNA objects of predetermined geometry (Pfeifer and Sacca`, 2016). Contrary to the DNA tile, DNA origami is made up from a single, very long strand and a number of short, clamping strands. The long strand is usually originated from a viral or phage genome, being several thousand nucleotides long. The same long strand can be used in the construction of various structures, whereas the set of short strands, staples, are unique for a particular structure. The amount of staples can be from dozens to hundreds. Despite the large number of strands, DNA origami is much easier to assemble than a DNA tile of comparable size. This circumstance supports the abundance of DNA origami described and used in various applications (Kuzuya and Komiyama, 2010; Pfeifer and Sacca`, 2016; Tørring et al., 2011). Several commonly used junctions of DNA strands are shown in Fig. 7.9. Threeand four-armed junctions, as well as 4 3 4 DNA tiles, were also used in DNA aptamer oligomers that have been described earlier in this chapter. A DNA triangle is a related structure; it consists of four strands that provide a three-way junction of

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FIGURE 7.9 Major motifs in double helical locks used in DNA tiles and DNA origami. Arrows indicate the 30 -ends of DNA strands. (A) Branched DNA junctions containing three or four strands. (B) 4 3 4 DNA tile of nine unique strands. (C) DNA triangle containing three double helices. (D) DX: double exchange between double helices; TX: two double reciprocal exchanges; PX: two double helices exchange strands at every point where helices come into contact; JX2: PX without the two central crossovers. Source: Reprinted from Tintore´, M., Eritja, R., Fa´brega, C., 2014. DNA nanoarchitectures: steps towards biological applications. Chembiochem 15(10), 13741390, © (2014), with permission from Wiley.

double helices. Other junctions represent a more sophisticated way of connecting two double helices, with crossovers that provide more rigid constructions. A double crossover junction, DX, results from a double exchange between double helices. A triple crossover junction, TX, results from two successive double reciprocal exchanges. Both junctions contain exchanges between strands of opposite polarity. A paranemic crossover junction, PX, is formed when two double helices exchange strands at every point where the helices come into the contact. A JX2 molecule differs from PX in that it lacks the two central crossovers. The exchanges in PX and JX2 are between strands of the same polarity (Tintore´ et al., 2014). Detailed reviews are available for assembly characteristics, including those considered thermodynamic and kinetic (Pfeifer and Sacca`, 2016; Tørring et al., 2011). Briefly, inter-strand recognition should be extended enough to ensure specificity, but sufficiently weak to enable reversibility of assembly, so that incorrect interactions can be removed, providing error check and self-correction of the

7.6 Examples of Extensive DNA Nanoconstruction Geometry

structure and escape from kinetic traps. The energy landscape of both DNA tiles and DNA origami usually has several local minima, so assembly of the structure should proceed under optimal conditions, which must be selected experimentally (Fern et al., 2016). Because edge junctions have lower stability than double helices with crossovers, structures with mixed junctions are generally noncooperative during disassembly (Ramakrishnan et al., 2016).

7.6 EXAMPLES OF EXTENSIVE DNA NANOCONSTRUCTION GEOMETRY 7.6.1 DNA TILES Four-armed junctions and 4 3 4 tile motifs were widely used to produce extensive structures with regular topology. One-dimensional (ribbons) and two-dimensional (canvas) structures of micrometer scale can be readily created (Tintore´ et al., 2014; Zhao et al., 2010). Triangles, double crossovers, and triple crossovers can also be used to construct regular, two-dimensional lattices. Three-dimensional DNA tiles were successfully produced, including those with the topology of different polyhedra. Geometrical figures of high symmetry can be generated using corner junctions and edge-crossover double helices (Tintore´ et al., 2014). As DNA tiles are most usable for regular constructions with strict geometry, sophisticated or asymmetric structures are generated using the origami approach.

7.6.2 TWO-DIMENSIONAL DNA ORIGAMI In planar origami structures, adjacent double helices are interconnected via crossovers. There are several rules for assembling stable planar nanoconstructions. For example, neighboring crossovers should be separated by 16 base pairs, corresponding to 1.5 turns of the double helix. Planar constructions can be equipped with a “landscape,” introducing hairpins in some staples. Each hairpin is placed at a position eight base pairs from the crossover; in this case, all hairpins will be oriented perpendicularly to the origami plane at a distance of approximately 6 nm (Endo and Sugiyama, 2011; Sacca` and Niemeyer, 2012). Hairpins are convenient for functionalization, including the placement of aptamers. The size of the twodimensional structure is determined only by the length of the template strand. To extend the size, a strategy of combining DNA tiles and DNA origami can be of use. Here, DNA tiles substitute conventional staples in planar DNA origami. This technique resulted in micrometer-scale structures with no difficulties (Tørring et al., 2011; Zhao et al., 2010). To produce extended finite structures, several planar DNA origami can be united, e.g., with both complementary interactions and a geometrical clamp, a groove in the first origami, and a bulge of appropriate size in the second origami (Endo and Sugiyama, 2011).

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7.6.3 THREE-DIMENSIONAL DNA ORIGAMI Design of three-dimensional origami structures can be performed from planar origami by applying two approaches: using special crossovers between adjacent double helices, and using staples for DNA origami sheets. The first approach yields tubular and multilayered structures. The geometry is controlled with the number of base pairs between crossovers (Fig. 7.10). Seven base pairs yield 120 between double helices. If the angle is the same in all helices, a six-helix bundled tubular structure (6-HB) is formed. Repeated six-helix bundled tubular structures form a honeycomb lattice. Eight base pairs between crossovers yield 90 between double helices, which corresponds to a cubic lattice. Other alterations in the distance between crossovers yield twisted and bent honeycomb origami structures (Endo and Sugiyama, 2011; Sacca` and Niemeyer, 2012; Tørring et al., 2011). The complexity of the structure and the high density of crossovers leads to week-long folding times, whereas planar structures are formed in several hours. More complex structures have been designed, including those with high curvature: spherical shells, ellipsoidal shells, a nanoflask, and chiral structures, e.g., Mo¨bius band (Kumar et al., 2016; Tørring et al., 2011).

FIGURE 7.10 Schematic representation of crossover patterns and generated lattices used for the construction of 2D and 3D DNA origami. Double helices are indicated by circles and are viewed along their central axis. The crossovers between helices are indicated with black arrows. The number of base pairs between consecutive crossovers is given, as well as the resulting arrangement of the helices. Source: Reprinted from Sacca`, B., Niemeyer, C.M., 2012. DNA origami: the art of folding DNA. Angew. Chem. Int. Ed. Engl. 51(1), 5866, © (2012), with permission from Wiley.

7.7 Promising Applications of DNA Tiles and DNA Origami

The second approach comprises assembly of DNA origami sheets in threedimensional hollow structures, such as boxes, prisms, tetrahedrons, etc. Structures are created by folding multiple two-dimensional origami domains, with staples at the edges of the figure (Endo and Sugiyama, 2011; Tørring et al., 2011). Hollow structures can be used as nanocarriers, with programmable release of the content. Examples of such drug delivery constructions will be discussed further.

7.7 PROMISING APPLICATIONS OF DNA TILES AND DNA ORIGAMI The programmable size and shape of DNA nanoconstructions provides a variety of potential applications. Among the most promising ones are drug delivery, biosensors, channels in membranes, precise arrangement of molecules, and molecular machines (Chandrasekaran, 2016b; Kearney et al., 2016; Kumar et al., 2016). The majority of the applications require conjugation of the nanoconstructions with molecular recognizing elements, e.g., aptamers. Several impressive examples of successful functionalization of DNA tiles and DNA origami will be discussed further, giving special attention to aptamer-based nanoconstructions and nanodevices.

7.7.1 DRUG DELIVERY SYSTEMS Hollow, three-dimensional origami can be used as nanocarriers. The most promising constructions release the content in response to a signal. Here, the benefit of the aptamers is obvious, as they act as recognizing elements, targeting the nanocarrier to the required place, e.g., to the definite type of cell in the body. Moreover, a target binding-induced conformational rearrangement triggers the programmable release of the content. One of the significant problems of DNA nanocarriers is stability, as content is strictly to be released after binding to the targeted molecule. The overall stability of typical DNA origami is sufficient for use under physiological conditions. However, the structure is susceptible to hydrolysis with nucleases, although the structure folding considerably affects the rate of hydrolysis (Kumar et al., 2016; Tintore´ et al., 2014). DNA origami, as large, highly charged biomolecules, can be nonspecifically internalized inside cells; they appear in lysosomes during the first 12 h, and then degrade in about 60 h (Shen et al., 2012b). The cell uptake of DNA nanoconstructions can be enhanced with viral envelope proteins or intercalators that modify the surface properties of DNA origami (Kumar et al., 2016). Andersen et al. (2009) have designed a hollow box, with dimensions of 42 nm 3 36 nm 3 36 nm, from a single genome of bacteriophage, M13. 220 staple strands were used to assemble six sheets that are linked together due to the

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common template strand. The sheets were assembled with 59 staple strands, yielding a closed box. This algorithm provides unimolecular cubes with sufficient yield. In addition, opened boxes were also assembled. The structure was equipped with a programmable mechanism of opening, using special keys. A dual lock-key system was composed of DNA duplexes with sticky-end extensions to provide a “toehold” for the displacement by externally added key oligonucleotides (Fig. 7.11). Box-opening was detected with fluorescence resonance energy transfer (FRET), due to the increase of the distance between two fluorescent dyes placed on the lid and the edge of the box. Opening of the box demands two oligonucleotide keys; therefore, the algorithm of a logic AND gate is realized, as at least two external signals are required. The box could potentially be designed to close again in the presence of specific signals, representing a logic “not” gate. Designing two potential lids on opposite sides of the box would give the possibility of opening with different keys, representing a logic “or” gate. Accordingly, a DNA box is a nanodevice with programmable behavior, operated by combinations of oligonucleotides.

FIGURE 7.11 Hollow DNA origami box with “keys” for its programmable opening: (A) planar scheme, (B) spatial structure of the box and (C) opening of the box with the keys. Source: Reprinted by permission from Andersen, E.S., Dong, M., Nielsen, M.M., Jahn, K., Subramani, R., Mamdouh, W., et al. 2009. Self-assembly of a nanoscale DNA box with a controllable lid. Nature 459(7243), 7376, © (2009), with permission by Macmillan Publishers Ltd.


Aptamers can be the locks, and aptamer targets can be the keys. A successful example was provided by Douglas et al. (2012). A DNA origami nanoconstruction had the form of a hexagonal barrel, with dimensions of 35 nm 3 35 nm 3 45 nm. The barrel has two parts, joined together in a seashell mode. These parts are connected by staples modified with DNA aptamer-based locks. These locks can be opened after the target protein binding with the aptamer. The idea is that target protein-induced disassembly of a double helix, formed by an aptamer strand and a complementary oligonucleotide strand. To ensure that the structure is properly closed, two aptamer-based locks were introduced. The length of the double helix of the lock is a critical parameter for the nanoconstruction. Shorter double helices give better sensitivity and faster unlocking rates, but the frequency of spontaneous activation is also increased. Double helix locks of 16 and 23 base pairs provide similar sensitivity for the target protein (10 pM), while double helices longer than 30 base pairs require 1 nM of protein. Several aptamers were tested in homo- and heterofunctional constructions. Proof of the principle was demonstrated for a nanodevice with two identical aptamers to platelet-derived growth factor (PDGF) in the locks. The barrel was loaded with antibodies to human CD33 and CDw328 Fab. Both antibodies were covalently linked to the inner surface of the barrel. The nanodevice was added to natural killer cell leukemia (NKL) cells, and functional activity was detected. The nanodevice opened after targeted cell binding, inhibiting growth of the cell population, and suppressing several phosphokinase cascades. Similar results were obtained for T cells that were activated with the nanodevice. These results demonstrate that DNA aptamer-based nanodevices can induce a variety of tunable changes in cell behavior. Several other examples have been reviewed in detail, including approaches for cancer treatment, immunity activation, etc. (Kearney et al., 2016; Li et al., 2013; Mohri et al., 2014; Zhan et al., 2014).

7.7.2 MEMBRANE-ASSOCIATED DNA NANOCONSTRUCTIONS The unique, programmable, three-dimensional structure of DNA origami provides a possibility to construct analogues of proteins. Hollow DNA origami structures can imitate membrane channel proteins, whereas two-dimensional DNA origami or tiles can serve instead of receptor proteins, for vesicle incorporation. Hydrophobic functional groups are incorporated to anchor highly charged DNA into the hydrophobic membrane. Several studies have been conducted and reviewed in detail (Czogalla et al., 2016; Hernańdez-Ainsa and Keyser, 2014; Langecker et al., 2014). A channel in the lipid membrane was constructed using the DNA origami technique. The channel consists of two modules: a hollow tube that penetrates the lipid bilayer and a barrel-shaped envelope that adheres to the membrane. The barrel has 48 double helices. Interaction with the lipid membrane is mediated by 26 cholesterol moieties attached to the outer surface of the barrel. The hollow tube

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has a structure topology of 6-HB origami (six double helices), having a cavity diameter of 2 nm and a length of 42 nm. The whole structure acts like an ion channel, providing conductivity under applied voltage; the channel displays a current gating, as do many natural ion channels. The nature of the gating consists of fluctuations of the structure, demonstrated by several modified constructions (Langecker et al., 2012). Aptamers can be of use for the channel origami as a lock. In the absence of the aptamer target, a double helix is formed between the aptamer and the complementary oligonucleotide. In the presence of the target, the double helix is untwined, the lock is opened, and the pore performs a function. One more application for aptamers is targeting the origami structure to specific cells that yields a modulation of ion exchange and cell death (Hernańdez-Ainsa and Keyser, 2014). DNA nanoconstructions can serve as analogues of receptor proteins, implementing vesicle internalization. The simplest system was described by Hadorn and Eggenberger Hotz (2010). Streptavidin was anchored into the membrane and bound to biotinylated oligonucleotides. Two vesicles with complementary oligonucleotides interact with each other, due to the formation of double helices. Potential applications of the systems are: studying cellcell complexes and creating vesicles with diverse content. The applications can be extended by introducing aptamers as recognizing elements for targeted drug delivery.

7.7.3 SPATIAL ARRANGEMENT OF THE MOLECULES A variety of modifications can be introduced into DNA tiles and DNA origami in a sequence-specific way; this allows conjugation of almost any molecule to a desirable site within the nanoconstruction. A bottom-up self-assembly technique provides high spatial resolution and precision for the arrangement of various components in microscopic dimensions, whereas a top-down technique provides visual control at the macroscopic scale. Combining the two techniques results in extended highly oriented constructions. The size of DNA origami structures is about 100 nm, which matches to the size that can be routinely manipulated by lithographic technology. This means that DNA nanoconstructions can bridge the microscale-to-macroscale gap. It is believed that the integration of top-down fabrication and bottom-up self-assembly will advance nanodevice fabrication significantly (Liu et al., 2014). Metal nanoparticles are worthy objects for demonstrating the advantages of DNA nanoconstructions as an organizing template, as metal particles are easily detected with a variety of techniques, mainly various kinds of microscopy and plasmonic resonance. In this field, several key features have been demonstrated. Nanoparticles were placed in certain geometry into two-dimensional origami and DNA tiles, with high load and yields (Liu et al., 2014). The model object, aurum nanoparticle, interacts with thiol-modified DNA; the techniques for DNA modification with thiol groups in a sequence-specific way have been elaborated.


The modification density of nanoparticles is limited by particle size, only as the capacity for DNA nanoconstruction modification is extremely high (Liu et al., 2014; Samanta et al., 2015). Based on two-dimensional origami structures, a three-dimensional helical arrangement of metal nanoparticles was created. A rectangular DNA origami structure with fifteen thiol groups, arranged in two parallel strands, was assembled. Aurum particles with a diameter of 10 nm were attached to the planar DNA nanoconstruction. The staples were then added, joining the two opposite sides of the rectangular DNA origami. The resulting three-dimensional structure represents a nanorod, with a helical arrangement of aurum nanoparticles; the structure has a chiral arrangement of metal nanoparticles (Shen et al., 2012a). The optical properties can be readily modified by changing the parameters of the helix. A more sophisticated example has been reported recently by Urban et al. (2016). Two enantiomeric toroidal structures are constructed from aurum particles and DNA origami, using hierarchical assembly. A DNA origami ring is constructed of four parts. Each part consists of 24 curved double helices arranged in a honeycomb lattice. Four monomers are connected together using staples, forming a DNA origami ring, with an inner diameter of 120 nm. Aurum nanoparticles are arranged into either left-handed or right-handed helices (Fig. 7.12A). These two examples highlight the ability to create any geometrical arrangement, due to sequence-specific modification of DNA nanoconstructions. Other molecules and nanoparticles can be arranged instead of metal nanoparticles, e.g.,

FIGURE 7.12 (A) Spatial arrangement of aurum particles by a DNA origami toroidal structure. (B) Spatial arrangement of two different protein types by DNA tiles modified with DNA aptamers. Source: (A) Adapted with permission from Urban, M.J., Dutta, P.K., Wang, P., Duan, X., Shen, X., Ding, B., et al., 2007. Plasmonic toroidal metamolecules assembled by DNA origami. J. Am. Chem. Soc. 138(17), 54955498, © (2007), American Chemical Society. (B) Adapted with permission from Chhabra, R., Sharma, J., Ke, Y., Liu, Y., Rinker, S., Lindsay, S., et al., 2007. Spatially addressable multiprotein nanoarrays templated by aptamer-tagged DNA nanoarchitectures. J. Am. Chem. Soc., 129(34), 1030410305, © (2007), American Chemical Society.

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proteins, magnetic nanoparticles, and quantum dots. Protein ordered arrays are of special interest, as they can be used in sensors, biocatalysis, and ordered enzyme cascades. The arrangement of proteins in DNA nanoconstructions can be achieved with covalent linking, ligand binding, DNA aptamer binding, Holliday junction binding, nucleic acid hybridization of DNA-tagged proteins, etc. The resulting nanoconstructions have the functional activity of settled proteins (Godonoga et al., 2016), leading to a variety of applications, such as bioreactors (immobilizing one enzyme) and enzyme cascades (co-localizing several enzymes that catalyze subsequent reactions) (Simmel, 2012). Ligand-assessed arrangement of proteins has been demonstrated for streptavidin binding, with a sequence-specific biotinylated DNA nanoconstruction. A variety of DNA tiles and DNA origami, functionalized with streptavidin, have been reported. Similar to metal nanoparticles, the possibility of making a desirable distribution of streptavidin has been demonstrated (Chandrasekaran, 2016b). DNA aptamers are artificial ligands for targeted proteins, and they can be easily introduced into DNA nanoconstructions. Similarly to the biotinstreptavidin complex or thiolaurum binding, aptamers provide a fine arrangement of proteins with high load and yield. Several corresponding examples are further discussed in detail. DNA tiles were used to arrange several types of protein into ordered extended two-dimensional nanoarrays. Two aptamers, a thrombin aptamer and a PDGF aptamer, were introduced in double crossover DNA tiles. The four-tile set assembles into two-dimensional arrays, displaying parallel and alternate lines of thrombin and PDGF aptamers. The period between identical aptamer lines is 64 nm, whereas the distance between different aptamer lines is 32 nm. The addition of either thrombin or PDGF results in 64 nm-periodic structures, whereas addition of both proteins results in 32 nm-periodic structures (Fig. 7.12B). Consequently, aptamers allow arrangement of different proteins simultaneously, with high precision and without unwanted interference between proteins (Chhabra et al., 2007). The next example overlaps with hetero-oligomeric nanoconstructions of aptamers. Thrombin aptamers, HD1 and HD22, are applied to arrange thrombin molecules using two-site binding. As the DNA origami nanoconstruction is a rather rigid structure, the effect of the proximity of two aptamers on the thrombinbinding ability can be readily estimated and visualized. DNA tiles made up of five adjacent double helices were applied. DNA aptamers were placed at the neighboring double helices and, for other variants, skipping one, two, or three double helices; therefore the distance between two aptamers was 2, 3.5, 5.3, and 6.9 nm, respectively. For these constructs, the thrombin-binding efficiency varied from 10% to 45% depending on the distance between the aptamers. The highest thrombin-binding turned out to be in the case of 5.3 nm distance, which is expected as the distance between two targeted thrombin sites is about 4 nm (Rinker et al., 2008). This study reveals the possibility of manipulating DNA


nanoconstruction affinity to a protein by applying a geometric factor. Medium affinity can be useful in those cases where protein binding is to be reversible, e.g., for highly specific protein extraction. One more way to achieve reversible protein binding is the toehold-mediated displacement of the protein-tagged strand (Chandrasekaran, 2016b).

7.7.4 BIOSENSORS Sensor application is one of the most promising fields for DNA nanoconstructions. A variety of examples have been described and recently reviewed (Chandrasekaran et al., 2016a; Kearney et al., 2016). Among the directions to explore are: (1) pHdependent i-motif-based nanoconstructions, where DNA strands with cytosine blocks lock a structure at slightly acidic pH and unlock a structure at alkaline pH; (2) sequence-specific detection of DNA and RNA strands, where formation of a double helix with analyte drastically affects the conformation of the nanoconstruction; (3) sequence-specific detection of DNA and RNA strands, where an analyte displaces the reporter strand in the DNA nanoconstruction, leading to the disappearance of the structure labeling; and (4) detection of aptamer binding with a protein target, using physical methods. The last example overlaps with protein arrangement using aptamer-based DNA origami and tiles. Oligomeric aptamer nanoconstructions are of great promise for application in sensors. A representative sensor, made up of a non-DNA nanoconstruction, modified with multiple aptamers, has been recently reported (Li et al., 2016). Aptamers are coupled to polymer nanochannels to mimic the function of adenosine receptors. A number of aptamers to adenosine are immobilized into conical nanochannels, made up of solid polymer membrane; the channel input and output diameters are 740 and 20 nm, respectively. Adenosine-binding produces a conformation change within the aptamer, from a flexible structure to a rigid structure, and the electric conductivity of the channel varies upon target binding. The initial aptamer has an affinity to the target of Kd 5 6 μM, whereas the sensor operating concentration is about 0.1 μM. These concentrations are considerably lower than both the affinity range and the limit of detection of enzyme-linked aptasensors (18 μM, Xiang and Lu, 2011; and 12 μM, Liu et al., 2012). A possible reason for that could be due to the large number of aptamers settled within a small volume, drastically increasing the effective local concentration of the aptamer and, consequently, the probability of complex formation. This idea could be implemented into DNA nanoconstructions, as they have a wide variety of possible DNA origami structures and can be readily functionalized with DNA aptamers.

7.7.5 MOLECULAR MACHINES A number of aptamer-based nanodevices have been considered in the previous subsections, such as drug delivery systems with aptamer locks. One more superb

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application field is DNA origami-based mimetics of biological molecular machines, e.g., kinesins, the motor proteins that move along microtubule filaments using adenosine triphosphate hydrolysis as a driving force. Goodman and Reck-Peterson (2014) described a DNA origami nanoconstruction that can be modified with several motor proteins, allowing movement along a microtubule filament. Twelve double helices are organized in a cylinder with six helices inside and six helices outside. The structure is supplied with special single-stranded handles for functionalization with motor proteins, labels, etc. The handles are located every 14 nm on the six outer helices, resulting in 15 handles per helix and 90 handles in the whole structure (Fig. 7.13). This arrangement allows for functionalization with numerous different molecules. For structure functionalization with proteins, Goodman and Reck-Peterson used proteins covalently linked to oligonucleotides; these oligonucleotides are complementary to the handles. Dynein and kinesin were attached to the DNA origami using photocleavable oligonucleotides, and the motion of the entire construction was studied. When dynein-associated oligonucleotides are cleaved, the construction moves to the “plus” end of the microtubule, due to the driving motion of kinesin. On the contrary, the cleavage of kinesin-associated oligonucleotides leads to movement in the opposite direction, due to the driving motion of dynein. Interestingly, the number of motor proteins has very little impact on the movement velocity.

FIGURE 7.13 A nanodevice with controllable movement in the cell: (A) DNA origami with periodically placed handles; (B) DNA origami functionalized with motor proteins that move the construction along the microtubule. Source: Reprinted from Goodman, B.S., Reck-Peterson, S.L., 2014. Engineering defined motor ensembles with DNA origami. Methods Enzymol. 540, 169188, © (2014), with permission from Elsevier.


In addition to photocleavable groups, other adjustable linkers can be used in nanoconstructions. For example, cistrans transitions can be useful for creation of UV radiation control and, contrary to photocleavage, cistrans isomerization is a reversible event. Another example is disulfide bridges that can be manipulated with redox potential. In another, DNA aptamers can be switched off with complementary oligonucleotides and switched on with target binding; both capabilities can provide operated release of the cargo. Another exciting direction is that of DNA nanoconstructions with several communicating modules. Interesting examples of controllable DNA tweezers have been reported recently (Elbaz et al., 2009a,b; Song et al., 2013; Wang et al., 2010). The basic principle of tweezer action is that the choice of DNA structure from several possible conformations is directed by minimization of free energy. This can support a competitive assembly, like the transfer of a DNA strand from one duplex to another because of the difference in stability. Employment of energy-related intermolecular or intramolecular constructions in the machinery system can activate several mechanical nanodevices simultaneously, allowing the devices to communicate with each other (Song et al., 2013). One of the examples involves two interconnected DNA tweezers. Tweezer A includes an aptamer to adenosine in the crossbar, while the arms are also connected with a linker strand. Tweezer B includes arms that can be connected with the same linker strand. The sophisticated design of the tweezer sequences provides a possibility of communicating between these two tweezers. The linker strand hybridizes with the arms of tweezer A, having a lower free energy than with tweezer B. Binding of adenosine unhybridizes the linker strand and it dissociates from tweezer A; the linker strand hybridizes with tweezer B. As a result, in the absence of adenosine, tweezer A is closed, while tweezer B is empty and open. In the presence of adenosine, tweezer A forms a complex with adenosine, while the linker strand closes the arms of tweezer B (Elbaz et al., 2009a; Song et al., 2013). A similar nanoconstruction was designed with a pH-sensitive sequence in the tweezer crossbar. The sequence folds into an i-motif conformation at slightly acidic pH and unfolds at neutral pH. As a result, coherent activation of the two tweezers occurs due to pH cycling (Elbaz et al., 2009b; Song et al., 2013). An even more sophisticated system has three tweezers that are regulated with pH, Hg21 ions, cysteine, and complementary strands to two different linker strands. Hg21 ions and cysteine control the opening and closing of tweezer A, pH controls the conformation of tweezer B and the conformation of tweezer C is controlled with tweezers A and B. The complementary strands open all three tweezers simultaneously. Different combinations of ligands result in 16 different reversible configurations of the whole system or eight different states of the tweezers. The interplay in tweezer states can be used for construction of a complicated communication network (Song et al., 2013; Wang et al., 2010). This tweezer system could be used in complex nanoconstructions as a regulator of moving parts.

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7.8 CONCLUSIONS The unique properties of DNA aptamers can be supplemented with DNA nanoconstructions, providing higher affinity and specificity to the target, as well as modified biodistribution in vivo. Strict and definite rules of DNA structure assembly allow the creation of a huge variety of two- and three-dimensional nanostructures with predictable geometry and variable stability. Functionalization with aptamers provides unique opportunities, such as controllable conformational switching, programmable decoration with proteins, and specific cell targeting. All these possibilities allow the construction of exciting artificial molecular machines with great promise in nanoscience and nanomedicine.

ACKNOWLEDGMENTS The work was supported by the Russian Foundation for Basic Research (grant no. 16-03-00136).

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CHAPTER

Nanobiodevices for electrochemical biosensing of pharmaceuticals

8

Sevinc Kurbanoglu, Bengi Uslu and Sibel A. Ozkan Ankara University, Ankara, Turkey

CHAPTER OUTLINE 8.1 Nanobiodevices ...............................................................................................292 8.1.1 Calibration .....................................................................................294 8.1.2 High Sensitivity..............................................................................294 8.1.3 Wide Measurement Range ...............................................................295 8.1.4 Stability.........................................................................................295 8.1.5 Lifetime.........................................................................................295 8.1.6 Response Time ...............................................................................296 8.1.7 Selectivity......................................................................................296 8.1.8 Rapid Response Time .....................................................................296 8.2 Nanobiodevices Based on Bio-Constituents........................................................297 8.2.1 Biomaterial Immobilization in Nanobiodevices ..................................298 8.2.1.1 Adsorption .............................................................................. 300 8.2.1.2 Covalent Bonding .................................................................... 300 8.2.1.3 Cross-Linking .......................................................................... 301 8.2.1.4 Entrapment Method................................................................. 301 8.2.2 Enzyme-Based Nanobiodevices........................................................302 8.2.3 Microbial Nanobiodevices ...............................................................306 8.2.4 Immunosensors ..............................................................................310 8.2.5 DNA-Based Nanobiodevices ............................................................312 8.2.6 Tissue-Based Nanobiodevices ..........................................................318 8.3 Electrochemical Methods in Biosensing ............................................................318 8.4 Conclusion ......................................................................................................322 References .............................................................................................................323


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CHAPTER 8 Nanobiodevices for electrochemical biosensing

8.1 NANOBIODEVICES Nanobiodevices are well suited for studying the interaction of an immobilized biological element. They provide a different theoretical aspect, with a biocomponent closely linked to a physical transducer, sensitive to the physicochemical interaction between the biocomponent and the analyte of interest. Electrochemical methods are valuable in investigating the biosensing responses where the electrode can convert the biological recognition event into a useful electrical signal, electron transfer, potential, or impedance change at the electrodesolution interface (Buerk, 1995; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000; Pohanka and Skla´dal, 2008; Scheller and Schubert, 1992). Starting from the 1980s, modification of the surface of the biosensors have gained importance. Currently, modified electrodes coupled with different techniques are commonly used and nanobiodevices such as lab-on-a-chip (LOC) platforms are launched with the developments in biosensor technology. Since the first LOC system was launched in 1999, more developments occur in the nanobiodevices area (Figeys and Pinto, 2000). It is believed that the future of biosensors will depend on LOC systems, due to their miniaturized biochemical analysis systems and their ability to give miniaturized skills to biochemical analysis (Buerk, 1995; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000; Pohanka and Skla´dal, 2008; Scheller and Schubert, 1992). Nanobiodevices, named biosensors, are commonly analytical devices that can recognize an analyte of interest and convert the analytical signal into a measurable electrical signal. Since the first biosensor was developed by Clark and Lyons in 1962, nanobiodevices have emerged as promising tools for both laboratories and industry (Buerk, 1995; Clark and Pazdernik, 2013; Miller and Nagarajan, 2000; Scheller and Schubert, 1992). Classification of biosensors can depend on the type of biorecognition and transducer. Depending on the biomaterial, biosensors can be named as biocatalysis-based biosensors, where enzymes are the biosensing elements. When whole cells or microorganisms are used as the biosensing material, the biosensor is called a cell-based biosensor. In bioaffinity-based biosensors, biosensing materials can be antibodies or oligonucleotides. When it is antibodies, immunosensor is the specific name of the biosensor where the antibodyantigen interaction is the main concern. When nucleotide structures are the biosensing element, they are oligonucleotide sensors or deoxyribonucleic acid (DNA) sensors. Generally, biorecognition parts can be produced from enzymes, DNA, tissues, bacteria, yeast, antibodies, antigens, liposomes, and organelles. An immobilization method that provides a stable biomolecule layer and a favorable environment for maintaining the biological activity is required in designing biorecognition parts. Suitable immobilization techniques, depending on the surface of a transducer, is the

8.1 Nanobiodevices

main challenge in biosensor technology. Related to their transducer parts, electrochemical biosensors can be classified as potentiometric, amperometric, conductimetric, impedimetric, etc. (Gerard et al., 2002; Gronow, 1991; Janata et al., 1994; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000) (Fig. 8.1). The transducer part of biosensors is necessary to determine if the analyte may consist of electrochemical (amperometric, potentiometric, and conductometric), optical, thermometric, piezo-electrical, or magnetic groups (Sharma et al., 2003; Velasco-Garcia and Mottram, 2003). The function of a biosensor depends on the biochemical specificity of the biologically active substance. The selection of biological substances depends on many factors such as specificity, storage, and operational stability. Moreover, determination of which parameters, such as chemical compound, antigens, microbes, hormones, nucleic acids, or taste and smell parameters, influences the selection of biological material (D’Souza, 2001; Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000). Several types of biosensors have been developed and used in many fields, such as clinical, medical, peripheral monitoring, and control of industrial processes for numerous analyses. In parallel with developments in microsensor and biotechnology fields, research and development studies in biosensors have sharply increased. The most important advantage for using biological materials in sensors is the increase in selectivity to certain compounds. Another benefit of biological materials is to recognize biologically-related molecules with them (Karube and Nakanishi, 1994). Most conventional methods require pretreatments to detect biomolecules; however, pretreatments are mostly not necessary in biosensors.

FIGURE 8.1 Schematic representation of electrochemical nanobiodevices. Reprinted from Kurbanoglu, S., Ozkan, S.A., Merkoçi, A., 2017a. Nanomaterials-based enzyme electrochemical biosensors operating through inhibition for biosensing applications. Biosens. Bioelectron., in press, with permission from Elsevier.

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Sensitivity of conventional methods is higher but miniaturization of them is difficult. Electronic or electrochemical detection is more advantageous when considering these aspects (Balasubramanian and Burghard, 2006). Biosensors are appropriate for real-time analyses, and such a feature is significant in industry, for instance, for rapid measurements in process visualization and controls. Besides this, biosensors have broad applications, such as pharmaceutics industry, medical diagnostic, food and drug tests, and environmental analyses (Karube and Nakanishi, 1994; Balasubramanian and Burghard, 2006). Transducers convert physicochemical signals, developed as a result of bioagent analyte interaction, into electrical signals and then convert it into readable and recordable data. Biosensors are divided into four sub-groups, depending on the desired signal type developed as a result of biological reactions: electrochemical, optical, piezo-electrical, and thermal biosensors. To properly design electrochemical, optical, or other biosensor types, there are basic physical properties that the measuring system must have. Some significant characteristics and properties of an ideal biosensor are as follows (Karube and Nakanishi, 1994; Balasubramanian and Burghard, 2006).

8.1.1 CALIBRATION A calibration graph is drawn according to the concentration of the analyte or logarithm after measuring corrected steady state responses, based on a balance current. In the calibration graph, the concentration interval—where there is a linear relationship between the concentration of the substrate and the response of biosensor—is called linear working interval. It is desired that an ideal sensor should not require one calibration or at least one. However, this feature was not being able to be applied in applications. It should not be forgotten that parameters affecting the response of biosensor will also affect the calibration of sensor. Biosensors must be calibrated during their lifetimes (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.2 HIGH SENSITIVITY Sensitivity is defined as the gradient in the response of a biosensor, depending on the change in the concentration of a special chemical. In most cases, while the biosensor is measuring, it detects the concentration change of the reactant or product. The sensitivity of biosensor with the analyte is related to the species directly determined, in accordance with the stoichiometry of chemical reaction. In some biosensor types, measurements are obtained through dynamic responses of biosensor. In this case, sensitivity can be defined as the change in signal with changing concentration (Chaubey and Malhotra, 2002; Eggins, 2002;

8.1 Nanobiodevices

Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000). There are many parameters determining effective sensitivity of the biosensor prepared for the target analyte: physical dimension of the sensor, the thickness of membrane, mass transfer of chemical species from the sample to the detector zone, and various operations decreasing the activity of biosensor. Ideally, sensitivity of a biosensor must not change during its lifetime and be high enough to determine output signal of the transducer (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.3 WIDE MEASUREMENT RANGE The response of a biosensor decreases due to reasons such as: the deterioration of stability after adding a standard solution at a certain concentration, the decrease in enzyme activity, the accumulation of some radicals on the surface of electrode, and no response of the biosensor is observed at higher concentration levels. In the calibration graph, the zone where the relation between the analyte concentration and the response of sensor is linear, is called “linear range.” For accurate measurements, the linear zone where the response of biosensor reveals a linear tendency must be defined; that is the linear interval where the measurement range is defined for biosensor. It is advantageous for a biosensor to have a wide linear range (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.4 STABILITY One of the performance factors indicating the lifetime of biosensors is stability. It is measured by the fact of how many measurements are possible by using the same biosensor. Greater stability of a biosensor provides significant advantages, in terms of labor and economy. The purity level, source, and immobilization method of the enzyme substantially influence the stability of enzyme sensor (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.5 LIFETIME Lifetime is the factor defined as the change in the activity of biomolecule as a result of measurements done by biosensors. Creation of necessary working and environmental conditions will show a preservative effect to the activity of

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biosensor. Accordingly, the lifetime of a biosensor will be higher (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.6 RESPONSE TIME Response time is defined as the time interval passing from the signal change, due to adding the analyte into the measuring media until it is stable again. The response time of a biosensor can be understood from its currenttime curves. Response time of a biosensor depends on the electrode material, the type of supporting material, the structure of biomolecule, the affinity of the analyte to biomolecule, and the redox potential of the analyte on the surface of electrode (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.7 SELECTIVITY Selectivity is an important parameter of an ideal biosensor, because this feature enables a biosensor to be used in a complex matrix as analytical measuring equipment. The selectivity of a biosensor is defined as the sensitivity of the biomolecule to the target analyte, and not being affected by other substances in the measuring media. An ideal biosensor should only respond to the change in the concentration of the target analyte and it should not be influenced by the existence of other chemical species. Otherwise, some problems occur in determining the concentration of the target analyte (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.1.8 RAPID RESPONSE TIME The time passing from the addition of the analyte into the reaction media and the change in the current until the stability of the current is called response time. The response time of a biosensor can be understood as its currenttime curves. In stairshaped curves, if the shape of steps is broad and wide, then response time is long (slow), otherwise, it is short (fast). Before adding the substance to be analyzed in amperometric applications into the working media, the current generated on the surface of the electrode must be stable (static state). After the stability of the current, the analyte is added into the media and the change in the current is recorded. The current must be stable before the second addition (Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000).

8.2 Nanobiodevices Based on Bio-Constituents

8.2 NANOBIODEVICES BASED ON BIO-CONSTITUENTS Nanobiodevices include biological and transporter systems that transduce data coming from biological systems into useful analytical signals. They generate signals depending on the concentration of a specific analyte or a group of analytes. The biological constituents of biosensors are categorized into two groups: biomaterials with catalytic properties and biomaterials without catalytic properties. The first group includes enzymes, microorganisms and tissues; the second group is antibodies, receptors, and nucleic acids (Sharma et al., 2003; Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000). Biosensors can be divided into groups related to the working principles of the biomaterial used: catalytic or affinity. In catalytic biosensors, oxidoreductases and hydrolases are generally used, which are called enzyme biosensors. Depending on the transducer type, formation or consuming of the species can be followed in the medium. In catalytic biosensors, inhibition processes can also be followed. In a catalytic biosensor, the enzyme can be obtained from an organism, extracted from tissues or organs, and whole cells. In that case, they can be called microbial sensors. As the analyte is absorbed by the microorganism, the increase in or decrease in respiratory activity can be observed. As products such as protons or ammonia are released from the reaction medium, they can be detected by the biosensor (Upadhyay and Verma, 2013; Ziolkowski and Gorski, 2014; Sharma et al., 2003; Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Toone, 2006; Illanes, 2008; Updike and Hicks, 1967). In affinity biosensors, there exists an affinity of the molecule for a target molecule. This leads to identification and quantification of the analyte of interest. Nucleic acids, which have affinity to their complementary nucleic acid, aptamers which have affinity to target molecules, antibodies which have affinities to their corresponding antigens, receptors or membranes which have affinity to specific molecules, belong to this type of biorecognition group (Lim et al., 2010; Costa et al., 2014; Bahadır and Sezgintu¨rk, 2015a,b; Pei et al., 2013; Lie´bana et al., 2014; Justino et al., 2015). In an affinity-based biosensor, the reaction between the analyte and the bioreceptor does not always result in an analytical signal, therefore labels are generally used in this type of biosensor. Consequently, the affinity based biosensors can also be classified in terms of labeled or label-free. In DNA sensors, aptasensors and immunosensors, labels are used that provide electrochemical signal, fluorescent signals, or dyes that can absorb radiation. In some cases, enzymes can also be used as labels (Park et al., 2008; Newman and Setford, 2006; Upadhyay and Verma, 2013; Ziolkowski and Gorski, 2014; Sharma et al., 2003; Chaubey and Malhotra, 2002; Eggins, 2002; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005). Examples of the operation mechanism of immunosensors, enzymatic sensors, microbial sensors, DNA sensors, and aptasensors are represented in Fig. 8.2.

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FIGURE 8.2 Examples of operation mechanism of (A) immunosensors, (B) enzymatic sensors, (C) microbial sensors, (D) DNA sensors, and (E) aptasensors. Reprinted from Pietrzak, M., 2015. Sensors and bioselective reagents, reference module in chemistry. Mol. Sci. Chem. Eng., 14, with permission from Elsevier.

8.2.1 BIOMATERIAL IMMOBILIZATION IN NANOBIODEVICES Good interaction between biological system and transporter system, i.e., the use of a successful immobilization technique, is of great importance in preparing an effective biosensor. Biomaterial immobilization can be defined as the physical or chemical binding of biomolecules to a supporting material in order to provide continuous uses by protecting the catalytic/affinity activity of biomaterials. Although the utilization of living organisms as raw materials is economically a limiting situation, such as enzyme production, this problem seems to be solved, due to microbial sources (Cosnier, 1999; Costa et al., 2005; Datta et al., 2013;


Guisan, 2006; Jesionowski et al., 2014; Kaur et al., 2016; Kennedy and White, 1985; Sassolas et al., 2012; Malhotra et al., 2005; Miller and Nagarajan, 2000; Carr and Bowers et al., 1980; Arroyo, 1998). Nevertheless, the purification of enzymes from microbial sources is an expensive process. The recovery of the activity of free biomaterial without losing its activity is very difficult in industrial applications. Control of the reaction is difficult, because free biomaterial cannot be removed from the medium of reaction at a desired moment. It can be considered that inhibitors can be added to the reaction to stop it at a desired moment. However, new impurities will be introduced to the reaction products that have already been contaminated by free biomaterials. The purification of products from these impurities substantially increases the cost. Because the extraction of free biomaterial used as catalyst from the reaction medium without losing its activity is impossible, reusability of the biomaterial is not possible, either. That is an important factor that increases the cost, because biomaterials are specific but also expensive catalysts. Moreover, free biomaterials cannot be continuously applied to production systems. Consequently, immobilization techniques have been developed to benefit from purified biomaterials as much as possible (Cosnier, 1999; Costa et al., 2005; Datta et al., 2013; Guisan, 2006; Jesionowski et al., 2014; Kaur et al., 2016; Kennedy and White, 1985; Sassolas et al., 2012; Malhotra et al., 2005; Miller and Nagarajan, 2000). Advantages of immobilized biomaterial compared to free biomaterial are: •

• • • • • • • •

It can be removed from the medium at the end of the reaction (filtration, centrifuging, etc.) and does not cause contamination of products by biomaterial. It is resistant to environmental conditions (pH, temperature, etc.). It can be used many times for a long time. It can be applied to continuous processes. It is more stable, compared to natural biomaterial. The formation of products can be kept under control. It is appropriate for stepwise reactions. It can reveal higher activity than free biomaterial, in some cases. The possibility of degradation of biomaterial by itself decreases.

While selecting the method to be used in biomaterial immobilization, attention should be paid so that the active center must not be damaged during or after immobilization. The biomaterial structure should be well known during making such a choice. If there is any bonding between biomaterial and carrier, carriers in which this bonding will not occur over the active center should be chosen, or the active center should be protected during the immobilization process (Cosnier, 1999; Costa et al., 2005; Datta et al., 2013; Guisan, 2006; Jesionowski et al., 2014; Kaur et al., 2016; Kennedy and White, 1985; Sassolas et al., 2012).

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8.2.1.1 Adsorption This is the oldest and simplest method used for biomaterial immobilization. It is based on the principle of mixing a surface-active, non-soluble adsorbent (active carbon, porous glass, ash, silica-gel, calcium carbonate (CaCO3), starch and gluten) with a biomaterial solution and removing the excess amount of biomaterial by washing (Nelson and Griffin, 1916; Klibanov, 2001). Biomaterial immobilization process by the adsorption method is implemented under moderate conditions; the activation of the carrier used in this way is not necessary. It is economically advantageous and considerably protects the biomaterial activity at the end of immobilization process. Although this method seems to be easy, the determination of optimum conditions is difficult. If the non-covalent interaction between carrier and biomaterial is not tight enough, biomaterial can pass to the medium through desorption, which causes the product to become contaminated (Nelson and Griffin, 1916; Klibanov, 2001).

8.2.1.2 Covalent Bonding Protein immobilization by covalent bonding method is the most used among all immobilization methods. One of the advantages of using this method is that the bond between biomaterial and matrix is well balanced and strong. Consequently, the dissolution of biomaterial back to the solution is prevented. However, covalent bonding of basic amino acid residues to the carrier should be prevented, to increase the bonding activity percentage. This phenomenon shows very difficult conditions are required for this application. Formation of covalent bonds is usually seen on the side chains of amino acids shown in the biomaterial. Covalent bonding of a biomaterial is formed, due to the interaction of the polymer with agents containing reactive groups (copolymerization of ethylene, anhydride of maleic acid) or the interaction of two functional agents to be served as a bridge between polymer and biomaterial. Several constituents including functional groups, such as sulfide, sulfhydryl, oxide, amino, carboxyl, hydroxyl, ammonium, amino, amide, methyl thiol, guanidyl, and phenol ring participate effectively in chemical bonds. There are several advantages of covalent bonding immobilization technique. The adsorption of biomaterials to the carrier matrix is appropriate, therefore, it is widely used. Bonding reactions over wide ranges of spectrum, the formation of a strong bond with carriers having various functional groups, and continuity of the biomaterial activity after bonding (Cosnier, 1999; Costa et al., 2005; Datta et al., 2013; Guisan, 2006; Jesionowski et al., 2014; Kaur et al., 2016; Kennedy and White, 1985; Sassolas et al., 2012). No reverse reaction of covalent binding as a result of some factors, such as pH, ionic intensity and substrate, is main advantage of covalent bonding (Fig. 8.3A).


FIGURE 8.3 Common methods of immobilization used in biosensing: (A) covalent binding; (B) adsorption; (C) cross-linking; (D) encapsulation; and (E) entrapment.

8.2.1.3 Cross-Linking In cross-linking biomaterial immobilization methods, one or multi-functional reactants have small molecules making bonds with biomaterial molecules, forming non-soluble complexes (Fig. 8.3C). The degree of cross-bonding and immobilization substantially depends on the concentration of protein and reactive, pH, and the biomaterial to be immobilized. The most-used cross-linking reactants are glutaraldehyde, chloroformate and carbonyl imidazole, heterocycles halogens, bioxiranes, divinylsulfones, p-benzoquinone, transition metal ions, and epichlorhydrins. Generally, glutaraldehyde is used as a cross-bonding reactant due to being cheap and commercially available (Kurbanoglu and Toppare, 2015; Schoevaart et al., 2004).

8.2.1.4 Entrapment Method The entrapment method, forces the biomaterial to stay in a specific place. The biomaterial cannot escape from where it is trapped. This method can be carried out in the cages of a polymer matrix, as well as in semipermeable membranes via microcapsules and micelles. The most important difference of this method from

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covalent and cross-bonding immobilizations is that the biomaterial molecule is not physically or chemically bound to any carrier. The entrapment method is divided into three groups: entrapment in polymer matrix, microencapsulation, and liposome technique. Entrapment in polymer matrix method is based on the principle of forming a high degree of cross-linked polymer in a biomaterial solution. Biomaterial molecules are entrapped between cross-linking networks as a result of polymerization, so their diffusion to the mother liquor is inhibited. The most used polymer is polyacrylamide, cross-linked by N, N0 -methylenebisacrylamide. The encapsulation method is based on the principle of entrapping biomaterial in a semipermeable membrane. The size of the microcapsules changes between 1 and 100 μm. Pore diameters of semipermeable membranes should be in an appropriate size so that they can allow the substrate molecules to enter into and product molecules to exit from the capsule. The smaller the substrate molecules are, the higher the productivity of the biomaterial immobilized by this method will be. Lastly, the liposome method is based on the liquid-surface constructor membrane. The most important priority of this method is that it is discontinuous, alternate, and entirely physical. It enables many biomaterials to be immobilized simultaneously in one step and has a large contact surface area. Its important disadvantages are that the diffusion of substrate and biomaterial through the membrane depends on solubility, the biomaterial is inactive during the process, and can escape from the liquid membrane (Brady and Jordaan, 2009; Brandi et al., 2006; Carr and Bowers, 1980).

8.2.2 ENZYME-BASED NANOBIODEVICES Many chemical reactions occur spontaneously. Some reactions need to be catalyzed to occur at a certain rate. Catalysts are molecules that decrease the activation energy necessary to chemically transform the substrate into another material. Thermodynamically, this amount of energy can be stated as the change in free energy. Catalysts are not consumed or subject to change during reactions. Catalysts that are used many times in reactions restrict or continue the transformation of the substrate into the product. This is related to the deterioration of their structure. Enzymes are purified from living organisms or living-originated sources, such as tissue, blood, and microorganisms for industrial and analytical purposes, because they are produced by living organisms. Many biochemical reactions occurring in living organisms are catalyzed by enzymes that are proteinstructured specific biocatalysts (Blackstock, 1989; Clark and Pazdernik, 2013; Holmquist, 2000; Liu, 2013; Palmer, 1991; Perumal and Hashim, 2014). The most commonly used biomolecules as bio-receptors are enzymes. The originality of the substrate to the analyte defined as the differences of biosensors from each other is actually the originality of bio receptor to the analyte. Enzymes have considerably high originality and affinity to their substrates, such that they choose their related substrate among many chemicals and trigger the reaction. Enzymes


do not change the equilibrium of a reversible reaction; they just make it easier to reach its equilibrium state. Enzymes can also bind to specific molecules, such as inhibitor, activator, or allosteric activator. Coenzymes that can be inorganic compounds, like metal ions or organic molecules, play important roles because they make enzymes catalyzed the reactions. Non-covalent interactions that stabilize the enzymesubstrate complex are the same as the interactions that stabilize the protein structure (Blackstock, 1989; Clark and Pazdernik, 2013; Holmquist, 2000; Liu, 2013; Palmer, 1991). Characteristics of enzymes come from the differences of their molecular structures. Enzymes are proteins composed of hundreds of amino acids. Amino acids are covalently connected to each other with a bond between one amino acid’s carboxylic carbon atom and the next one’s nitrogen atom in the α-amino group. Amino acids can be hydrophilic or hydrophobic, based on characteristics of their radical groups. Enzymes are usually colorless and can be dissolved in water or salt solution (Blackstock, 1989; Bowers and Carr, 1980; Liu, 2013; Palmer, 1991). The potency of an enzyme is expressed in units. The unit activity is the amount of change or destruction of a given quantity of a substrate under certain defined conditions. Often, the number of units is defined as micro moles (μmoles) of substrate, transformed to product/min per mg of total enzyme-polymer conjugate. Another way of expressing the activity of an enzyme is by specific activity, the number of μmoles of substrate transformed to product/min/mg of protein, under certain specified conditions of temperature, pH, etc. All factors (the concentration of enzyme, the concentration of substrate, temperature, pH, and the existence of inhibitors) affecting the rate of the reactions catalyzed by an enzyme are substantially important (Blackstock, 1989; Holmquist, 2000; Illanes, 2008; Perumal and Hashim, 2014). Biochemical reactions occur at this site by imparting catalyzing and specificity characteristics to enzymes. Side chains of amino acids form a three-dimensional (3D) structure, compatible with the substrate; the active site constitutes an enzyme substrate complex by binding to the substrate. Enzymes separate from the product after the transformation of the complex into the product (Blackstock, 1989; Bowers and Carr, 1980; Holmquist, 2000; Illanes, 2008; Liu 2013; Palmer, 1991; Perumal and Hashim, 2014). In enzyme-based nanobiodevice design, enzyme inhibition phenomena is the main interest for pharmaceutical analyses (Kurbanoglu et al., 2015, 2017a,b; Mayorga-Martinez et al., 2014; Perumal and Hashim, 2014). Generally, amperometric techniques are used in enzyme-based sensing. Scientific research about first-generation enzyme electrodes has started as a result of the increase in necessity of rapid and high sensitive analyses of several compounds. As a consequence of these progressions, various types of biosensors have been developed with combinations of different biological materials and transporter systems. Electrochemical systems have the highest usage rate among all these systems. Similar to biosensors, enzyme sensors have bio-receptors, carriers, and

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measurement systems (Kurbanoglu et al., 2015, 2017a,b; Mayorga-Martinez et al., 2014; Perumal and Hashim, 2014). Electrodes modified by enzymes are the most important applications of amperometric biosensors. These systems show high sensitivity in biological-based substrate uses, due to the biological function of the enzyme. The kinetics of an enzyme reaction is monitored by the rate of the product or the rate of the decrease in the amount of the reactant. If the product in enzyme sensors is electroactive, the reaction between them can be examined with amperometric techniques. Amperometric biosensors are formed as a result of these reactions. These electrochemical reactions occur on the surface of electrode and they are used to generate current. Electrons generated as a result of the enzymatic oxidation in the reaction are carried to the electron donor by being transferred to the anode (Kurbanoglu et al., 2015, 2017a,b; Mayorga-Martinez et al., 2014; Perumal and Hashim, 2014). Selection of the electrode to be used in enzyme biosensors is done based on electrical conductivity. Solid-supporting electrodes are usually used, such as gold, platinum, or carbon (amorphous carbon, carbon fiber, or nanotube film). Enzymes to be used for the composition of enzyme electrodes are selected based on the reaction type of the analyte. Redox-enzymes (oxidoreductases) are the most commonly used ones. Hydrolysis enzymes, such as lipases and esterases, can be also used with redox enzymes. The most serious problems encountered during the use of enzyme electrodes are oxygen requirements and hydrogen peroxide requirements of reactions while being monitored by amperometric techniques. These problems can be overcome by using appropriate mediators. Conductive and functionalized polymers, composite materials, solgels and nanomaterials can be used as immobilization materials in enzyme sensors (Kurbanoglu et al., 2015, 2017a,b; Mayorga-Martinez et al., 2014; Merkoçi, 2007; Perumal and Hashim, 2014). In their research, Kurbanoglu et al. developed a nanodevice based on iridium oxide nanoparticles, conjugated with magnetic nanoparticles for the inhibitionbased for the analysis of methimazol drug using tyrosinase enzyme. LOC devices and batch devices are designed as very sensitive, with a low limit of detection of 0.006 and 0.004 μM for batch and LOC nanobiodevices (Fig. 8.4). In another study for the detection and determination of pharmaceuticals, Turan et al. developed a nanobiodevice for the detection of antidementia drugs, neostigmine and donepezil. For this purpose, acetylcholinesterase and choline oxidase are coimmobilized in the matrix of electrocopolymerized 5,6-bis(octyloxy)-4,7-di (thiophen-2-yl)benzo[c][1,2,5]oxadiazole (BODT) with (2-(((9H-fluoren-9-yl) methoxy)carbonylamino) acetic acid on graphite electrode. Using the designed nanobiodevice, very low detection limits of 0.027 μg/L donepezil and 0.559 μg/L neostigmine were achieved (Fig. 8.5). In our last study, captopril drug was analyzed through inhibition of tyrosinase. The biosensor was constructed by iridium oxide nanoparticles, and electrochemically reduced graphene oxide modified screen-printed electrode.


FIGURE 8.4 SEM images of MNPs and Tyr (A), IrOx and Tyr (B) and IrOx NPsTyr-MNPs nanocomposite (C). (D) Schematic representation of the proposed detection system displaying tyrosinase (Tyr) and the reaction involved in the catechol detection. (E) Lab-on-a-chip design. Reprinted from Kurbanoglu, S., Toppare, L., 2015. Ethanol biosensor based on immobilization of alcohol oxidase in a conducting polymer matrix via crosslinking with glutaraldehyde. Revue Roumaine Chim., 60 (56), 453460, with permission from Elsevier.

Tyrosinase is immobilized onto a screen-printed electrode by the chemistry between 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide and N-hydroxysuccinimide (Fig. 8.6). The captopril detections were achieved by two inhibition pathways, chelating copper at the active site of tyrosinase and thioquinone formation, with low limits of detection values of 0.019 and 0.008 μM, respectively (Kurbanoglu et al., 2017a,b). Agricultural pharmaceuticals, pesticides, and insecticides are generally analyzed by enzyme inhibition phenomena using enzymatic biosensors. To illustrate: Haddaoui and Raouafi developed a biosensor, based on tyrosinase and ZnO nanoparticles for the detection of chlortoluron (Fig. 8.7). The designed biosensor can

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FIGURE 8.5 Schematic representation of the proposed biosensing system. Calibration curves for (A) DON and (B) NEO with the new designed biosensor (in 50 mM PBS, pH 7.0, 25 C, 20.7 V, 5 min incubation time). Reprinted from Turan, J., Kesik, M., Soylemez, S., Goker, S., Kolb, M., Bahadir, M., et al., 2014. Development of an amperometric biosensor based on a novel conducting copolymer for detection of antidementia drugs. J. Electroanal. Chem., 735, 4350, with permission from Elsevier.

detect phenol between 0.1 and 14 μM, and can be used to detect herbicide between 1 and 100 nM chlortoluron, with a limit of detection of 0.1 ppb (Haddaoui and Raouafi, 2015).

8.2.3 MICROBIAL NANOBIODEVICES In microbial biosensors, the microorganism that specifically recognizes the related material is used as the biocomponent. Biomaterial is immobilized into the appropriate transporter system. The transporter system transforms the biochemical signal into the electrical signal. Obtained signals change, depending on the concentration of analyzed compound. Transporter systems used in microbial


FIGURE 8.6 Schematic representation of proposed detection system for Captopril detection through (A) thioquinone formation (B) chelating copper at the active site of tyrosinase. Reprinted from Kurbanoglu, S., Rivas, L., Ozkan, S.A., Merkoçi, A., 2017b. Electrochemically reduced graphene and iridium oxide nanoparticles for inhibition-based angiotensin-converting enzyme inhibitor detection. Biosens. Bioelectron., in press, with permission from Elsevier.

biosensors can be optical, potentiometric, amperometric, or calorimetric (D’Souza, 2001; Tkac et al., 2005; Karube and Nakanishi, 1994). Microbial biosensors can be categorized into two groups, based on measurement principle: those that measure the metabolic activity and those that measure the electrochemical active metabolite. Electrochemical microbial biosensors can be amperometric, potentiometric, and conductometric, based or in the form of a microbial fuel cell (Lei et al., 2006). Luminescence-based physical transporter systems are used for the preparation of optical microbial biosensors. Luminescence densities of photo bacteria are based on their metabolic activities. High sensitive microbial biosensors can be prepared by combining photo bacteria with optical systems (D’Souza, 2001; Tkac et al., 2005; Karube and Nakanishi, 1994). It is possible to prepare thermal-based microbial biosensors by considering temperature changes due to metabolic activities of living cells. For example, a connection can be generated between the temperature change due to the injection and the concentration of the analyte. It is possible to measure the temperature change by thermistors (D’Souza, 2001; Tkac et al., 2005; Karube and Nakanishi, 1994). To illustrate this, in their study, Bayram and Akyilmaz developed a new microbial biosensor based on Bacillus subtilis using carboxylated multiwalled carbon nanotube (CNT) and conductive polyaniline in the presence of glutaraldehyde

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FIGURE 8.7 (A) Schematic representation of (i) electrochemical deposition of ZnONPs using cyclic voltammetry, (ii) functionalization of ZnO/SPCE electrode using tyrosinase/glutaraldehyde, (iii) (a) detection of phenol and (b) inhibition-based detection of chlortoluron. (B) Histogram representing the inhibition levels depending on the incubation time and for a set of concentration. (C) Chronoamperograms showing the reversibility of inhibition process: control (i), before incubation (ii), after the incubation in 100 nM solution of chlortoluron 30 min (iii), and after rinsing of electrode with PBS (iv). All the measurements were recorded in PBS in presence of before 5 μM of phenol. Reprinted from Haddaoui, M., Raouafi, N., 2015. Chlortoluron-induced enzymatic activity inhibition in tyrosinase/ZnONPs/SPCE biosensor for the detection of ppb levels of herbicide. Sens. Actuators B 219, 171178, with permission from Elsevier.

on a gold working electrode for determination of paracetamol. Amperometric detection of paracetamol was achieved in the linear range of 5630 μM, with a limit of detection value of 2.9 μM (Fig. 8.8).


FIGURE 8.8 Cyclic voltammograms (A) in the absence and in the presence of 100, 250, 1000 μM (inner to outer) acetaminophen. Differential pulse voltammograms (B) in the absence and in the presence of 100, 250, 1000 μM (inner to outer) acetaminophen. Effect of applied potential for amperometric method (C) at different applied potentials from 0.3 to 0.6 V. Measurements were done in phosphate buffer; pH 7.0, 50.0 mM containing 0.1 mM KCl; T: 30 C. Scan rate is 10 mV/s; Basillus sp.: 10 mg/mL; aniline: 0.05 M; cMWCNT: 1 mg/mL. Reprinted from Bayram, E., Akyilmaz, E., 2016. Development of a new microbial biosensor based on conductive polymer/multiwalled carbon nanotube and its application to paracetamol determination. Sens. Actuators B: Chem., 233, 409418, with permission from Elsevier.

In another study, Karasinski et al. developed a multi-array dissolved oxygen multi-electrode sensor for the detection of Escherichia coli, Escherichia adecarboxylata, Comamonas acidovorans, Corynebacterium glutamicum and Staphylococcus epidermidis. The authors reported that the upper concentration limit to produce a DOX fingerprint falls between 11 and 56 3 106 cfu/mL, while the lower limit must be above 1 3 106 cfu/mL if one wishes to obtain results in under 8 h (Fig. 8.9).

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FIGURE 8.9 Schematic of the DOXPCA concept. (AC) represent DOX responses for high medium and low cell concentrations, respectively. Reprinted from Karasinski, J., White, L., Zhang, Y., Wang, E., Andreescu, S., Sadik, O.A., et al., 2007. Detection and identification of bacteria using antibiotic susceptibility and a multi-array electrochemical sensor with pattern recognition. Biosens. Bioelectron. 22(11), 26432649, with permission from Elsevier.

8.2.4 IMMUNOSENSORS The specificity of the molecular identification of antigens by antibodies forming a stable complex is based in both analytical immunoassays and those formed by a solid-supporting material. Immunoassay technology is significantly important for clinical laboratories, and research about this topic has continued intensively. In addition to clinical studies, quality controls in the food industry and the determination of trace amounts of substances in environmental analyses requires the use of immunoassays. The requirements in these applications have increased the attention to immunosensors that enable continuous analyses (Luppa et al., 2001;


FIGURE 8.10 Schematic representation of electrochemical immunosensor. Reprinted from Bahadır, E.B., Sezgintu¨rk, M.K., 2015a. Applications of electrochemical immunosensors for early clinical diagnostics. Talanta 132, 162174; Bahadır, E.B., Sezgintu¨rk, M.K., 2015b. Electrochemical biosensors for hormone analyses. Biosens. Bioelectron. 68, 6271, with permission from Elsevier.

Bahadır and Sezginturk, 2015a,b; Wu et al., 2007; Donahue and Albitar, 2010; Perumal and Hashim, 2014). In Fig. 8.10, schematic representation of electrochemical immunosensor was shown. In the early 1950s, the first antibody-based biosensor was designed, which started the immunosensor phenomenon. Since antibodies—proteins produced by the immune system—have an extreme affinity to their antigens, these type of biosensors generally have higher specificity and lower limit of detection values. Immunosensors are generally used for the detection of proteins, such as alphaprotein and protein biomarkers (Bahadır and Sezginturk, 2015a,b; Wu et al., 2007; Donahue and Albitar, 2010; Perumal and Hashim, 2014). In their work, a new and label-free electrochemical immunosensor for sensitive detection of bisphenol A was reported by Huang and Zheng. The authors reported an immunosensor for bisphenol A using multiwalled carbon nanotubes (MWCNTs) and gold nanoparticles (AuNPs) modified on glassy carbon electrode (Huang et al., 2016). They used Rutin as the redox probe for the first time. Moreover, cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) were utilized to monitor the electrochemical immunosensor (Fig. 8.11). The method was optimized in terms of incubation time and pH of phosphate buffer solution. A linear range was obtained between 1.0 3 1028 and 1.0 3 1026 M bisphenol A, with a limit of detection value of 8.7 3 1029 M. The designed biosensor was also applied to real samples to determine bisphenol A (Huang et al., 2016).

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FIGURE 8.11 (A) The schematic illustration for fabrication of the electrochemical immunosensor (B) CVs of BSA/anti-BPA/AuNPs/MWCNTs/GCE immunosensor at different scan rates (fromatog): 10, 30, 50, 70, 90, 100, 110 mV/sin 2.5 3 1025 M Rutin, 0.1 M KCl 0.1 M pH 7.0 PBS solution. Inset: the plots of peak current vs. v and ν 1/2, respectively. Reprinted from Huang, Y., Li, X., Zheng, S., 2016. A novel and label-free immunosensor for bisphenol A using rutin as the redox probe. Talanta 160, 241246, with permission from Elsevier.

8.2.5 DNA-BASED NANOBIODEVICES DNA is the biomaterial that is generally used in production of new devices in nanotechnology and biosensor technology. DNA-based biosensors found a place in nanobiodevices field, related to their numerus applications such as DNAdrug interaction mechanisms, detection of DNA damage, detection of DNA sequences, etc. DNA is an essential molecule to determine hereditary characteristics and transfer the data needed for replications. DNA, because of its chemical and biophysical properties, has been an important topic for the analysis and discovery of genetic code and the genome sequence (Diculescu et al., 2005; Hahn et al., 2005;


Mikkelsen, 1996; Teles and Fonseca, 2008; Fojta et al., 2016). DNA biosensors are used to determine contaminant agents or metals, genetically-modified foods or organisms, DNAdrug interaction mechanisms, bioterrorism, diagnosis of diseases, damage to DNA-base, or specific DNA sequences of human, virus, or bacteria (Diculescu et al., 2005; Hahn et al., 2005; Mikkelsen, 1996; Teles and Fonseca, 2008; Fojta et al., 2016; Rauf et al., 2005). Aptamers are small nucleic acid sequences (DNA or RNA) selected in vitro from large combinatorial pools to bind to specific targets. Due to their high affinity for a series of biomolecules, they are largely used in biosensor development. To illustrate aptamer-based biosensors, selective electrochemical aptasensor for the ultrasensitive detection of an antiinflammatory drug, ibuprofen, has been developed by Roushani and Shahdost-fard. The biosensor was constructed by the modification of a glassy carbon electrode with MWCNTs/ionic liquid/chitosan (IL/Chit). Moreover, immobilization of ibuprofen-specific aptamer in a designed biosensor matrix was achieved by covalently using a redox marker, methylene blue. Methylene blue was intercalated onto the aptamer as the electrochemical redox marker. Using differential pulse voltammetry (DPV) experiments, a linear range was obtained between 70 pM and 6 μM ibuprofen, with the detection limit value as 20 pM (Roushani and Shahdost-fard, 2016) (Fig. 8.12). In another work, Derikvand et al. developed a new aptasensor, based on platinum nanoparticles on CNTs, functionalized with polyethyleneimine for the immobilization of diclofenac (DIF) aptamer. The authors used EIS for the detection of diclofenac. After adding diclofenac as the target of the biosensors to the system, the aptamer specifically binds to DIF. After that, the diclofenac aptamer end folds into a DIF-binding junction, retarding the interfacial electron transfer of the probe at the surface of modified electrode. Sensitive quantitative detection of DIF was carried out by monitoring the increase of charge transfer resistance (Rct) by increasing the DIF concentration. The designed aptasensor presented a linear range from 10 to 200 nM, with a detection limit of 2.7 nM (Derikvand et al., 2016). In their work, Ilkhani et al. developed nanostructured surface-enhanced Raman scattering (SERS)-electrochemical biosensors for the DNA interaction of doxorubicin (DOX). The biosensor was analyzed through SERS spectroscopy and electrochemical methods. Moreover, these techniques were used to screen biosensor modification and observe the drug-DNA reactivity. The self-assembled, monolayer-protected, gold-disk electrode (AuDE) was covered with a reduced graphene oxide (rGO), plasmonic gold-coated Fe2Ni@Au magnetic nanoparticles, functionalized with double-stranded DNA, and a sequence of the breast cancer gene. The nanobiosensors AuDE/SAM/rGO/Fe2Ni@Au/dsDNA were then subjected to the action of a model DOX, to assess the DNA modification and its dose-dependence as represented in Fig. 8.13. The SERS measurements have corroborated the DOX intercalation into the DNA duplex, whereas the electrochemical scans have indicated that the DNA modification by DOX proceeds in a

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Nanocomposite

Terephthalaldehyde

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GCE

MB

GCE

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O

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O

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H

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FIGURE 8.12 The schematic illustration for the detailed measurement process of the electrochemical aptasensor for IBP. Reprinted from Roushani, M., Shahdost-fard, F., 2016. Covalent attachment of aptamer onto nanocomposite as a high performance electrochemical sensing platform: fabrication of an ultra-sensitive ibuprofen electrochemical aptasensor. Mater. Sci. Eng. C 68, 128135, with permission from Elsevier.

concentration-dependent manner, with limit of detection of 8 mg/mL (Ilkhani et al., 2016). In their work, Shi et al. developed an extremely sensitive and selective “sandwich”-type electrochemical biosensor for the detection of the sequence-specific target DNA. The biosensor was constructed by immobilization of DNA in gold nanorods (Au NRs), reduced graphene oxide (rGO) sheets, and reporter probelabeled gold nanoparticle (AuNPs) matrix, which resulted in enhancing the sensing performance. In this work, the authors used the drug adriamycin as an electrochemical indicator, due to its electrostatically bonding properties to the anionic phosphate of DNA strands. The peak current of the Adriamycin indicator was followed as it changed after hybridization, using DPV. Under optimal conditions, the peak currents of adriamycin in DPV were linear, with the logarithm of target DNA concentration in the range of 1.0 3 10216 to 1.0 3 1029 M and a detection limit of 3.5 3 10217 M (Shi et al., 2014) (Fig. 8.14).


FIGURE 8.13 Construction of SERS/electrochemical nanobiosensors for testing of anti-cancer drug interactions with DNA: (A) SERS biosensor with Au(111) substrate coated with a selfassembled monolayer (SAM) of hexanedithiol (HDT) octanethiol (OT) with covalently attached gold nanoparticles (AuNPs) or magnetic Fe2Ni@AuNPs functionalized with ssDNA probe; (B) SERS/ electrochemical nanobiosensor with gold disk electrode (AuDE) substrate protected with a SAM of CYS and reduced graphene oxide (rGO) with covalently attached AuNPs or Fe2Ni@AuNPs functionalized with dsDNA. (C) Cyclic voltammograms for a AuDE electrode in 0.5 mM uncomplexed Fe2redox probe in 50mM PBS. (D) Differential pulse voltammograms. Reprinted from Ilkhani, H., Hughes, T., Li, J., Zhong, C.J., Hepel, M., 2016. Nanostructured SERSelectrochemical biosensors for testing of anticancer drug interactions with DNA. Biosens. Bioelectron. 80, 257264, with permission from Elsevier.

In recent studies, LOC devices are also found a place in DNA biosensors. Kahanda et al. developed a DNA device for following the repair of DNA damage, produced by a redox-cycling anticancer drug, beta-lapachone (β-lap). Using square wave voltammetry, designed DNA chips were used as drug-free control (Fig. 8.15).

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FIGURE 8.14 (A) Schematic represents the fabrication procedure of DNA biosensor. (B) DPV curves of the biosensor hybridized with increasing concentrations of complementary DNA: 0 M (a); 1.0 3 10216 M (b); 1.0 3 10215 M (c); 1.0 3 10214 M (d); 1.0 3 10213 M (e); 1.0 3 10212 M (f); 1.0 3 10211 M (g); 1.0 3 10210 M (h); 1.0 3 1029 M (i). (C) Logarithmic plot of peak current of adriamycin (I) vs. the concentration of complementary DNA. Reprinted from Shi, A., Wang, J., Han, X., Fang, X., Zhang, Y., 2014. A sensitive electrochemical DNA biosensor based on gold nanomaterial and graphene amplified signal. Sens. Actuators B: Chem. 200, 206212, with permission from Elsevier.

In another study conducted by Shen et al., based on the binding of anticancer drug dacarbazine to DNA and DNA bases, the author used gold nanoparticles, to enhance the biosensing response. The interaction between dacarbazine and DNA was employed by cyclic voltammetric technique. Adenine and guanine bindings were followed and the potential valuable applications of these gold nanoparticles


FIGURE 8.15 (A) The futile redox cycle associated with β-lap in the presence of high levels of NQO1 and NADH. (B) Illustration of tracking β-lapachone (ß-lap) drug activity with DNA chips against a control with no β-lapachone. Experimental electrodes are treated with all of the components to generate ß-lap-induced DNA damage and subsequent repair activity, leading to lower signals. Reprinted from Kahanda, D., Chakrabarti, G., Mcwilliams, M.A., Boothman, D.A., Slinker, J.D., 2016. Using DNA devices to track anticancer drug activity. Biosens. Bioelectron. 80, 647653, with permission from Elsevier.

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in the relative biomedical and bioengineering areas were discussed (Shen et al., 2008). In their latest study, Kurbanoglu et al. also developed a biosensor for the following of interaction between dsDNA and the antidepressant drug, aripiprazole, with original and damaged dsDNA. The binding constant between dsDNA and aripiprazole was calculated with voltammetric and spectrophotometric techniques (Kurbanoglu et al., 2015).

8.2.6 TISSUE-BASED NANOBIODEVICES Biosensors composed of tissue cross-sections instead of isolated enzymes have been encountered in several studies. Tissue-based nanobiodevices have many advantages, such as enzymes’ being more stable, having high enzyme activity, being cheap, able to be prepared easily, and not requiring the cofactor of enzymes (Filho and Vieira, 2000). The signal is obtained by monitoring hydrogen peroxide or oxygen, especially in bio-electrodes composed of plant tissue. Analysis of H2O2 by coconut tissue (Kozan et al., 2007), diamine by pea tissue (Mei et al., 2007), ethanol by mushroom tissue (Huang and Wu, 2006), flavonol by apple tissue (Cummings et al., 1998), paracetamol by avocado tissue (Filho et al., 2001), and glycolic acid by spinach tissue (Zhu et al., 2004) have been successfully carried out.

8.3 ELECTROCHEMICAL METHODS IN BIOSENSING Nanobiodevices can consist of two parts: biorecognition and transducer. The transducer part is directly responsible for the obtained signal from biochemical reactions that took place in biorecognition part (Thevenot et al., 2001a,b). The working principle of the transducer is to transfer the signal from the output domain of the biorecognition system and convert a response to a signal (Thevenot et al., 2001a,b). According to transducer type, biosensors can be classified as, electrochemical (amperometric, impedimetric, potentiometric, conductimetric) optical, acoustic, thermal, and piezoelectric (Ozkan et al., 2015; Ramanathan and Danielsson, 2001; Thevenot et al., 2001a,b). In nanobiodevices, electrochemical methods are generally used: among electrochemical methods such as amperometry, voltammetry, and potentiometry, voltammetry is a commonly used method, due to its easily applicable properties, high sensitivity, selectivity, inherent simplicity, and ease of mass production (Pohanka and Skla´dal, 2008; Ronkainen et al., 2010). Moreover, related to the low cost and availability of instrumentation, electrochemical biosensors have been improved and now better in quality. When the electrochemical species is consumed or produced during the biological reaction, the electrochemical signal can be recorded using an electrochemical detector. The working principle of

8.3 Electrochemical Methods in Biosensing

electrochemical biosensors is based on the electrochemistry of the substrate or product. An effective, simple to handle, low cost, and fast detection in electrochemical biosensor is related to the analytical power of electrochemical techniques and the process for a specified biorecognition element (Freire et al., 2003; Pohanka and Skla´dal, 2008). In voltammetry, the current is recorded as a function of applied potential (Bockris and Khan, 1993; Brett and Oliveira-Brett, 1993, 1998; Gileadi et al., 1975; Gosser, 1988). Voltammetry can be divided into CV, step and pulse voltammetry, alternating current, and stripping voltammetry. Linear sweep voltammetry (LSV) and CV are the most commonly used, and initial experimental methods in biosensing applications. CV can give initial information about the mechanism of the drug and the interaction with biomaterial. In enzyme-based modified biosensors, the modification character can be followed by the power of the CV in the enzymatic responses. In drug-DNA analyses, the decrease in the CV can give a first impression and information about the DNAdrug interaction (Bockris and Khan, 1993; Brett and Oliveira-Brett, 1993, 1998; Christian et al., 2013; Gileadi et al., 1975; Gosser, 1988; Wang, 2006). In summary, LSV and CV depend on the application of a potential that is decreasing/increasing with time. The degree of peak currents, Ip,a and Ip,c, and the potentials where the peaks are observed, Ep,a and Ep,c, are important parameters in CV and LSV (Brett and Oliveira-Brett, 1998; Gosser). Pulse voltammetric techniques can be classified as normal-pulse, alternating current, differential pulse, square wave, and staircase. These voltammetric techniques, are directly based on the application of pulse changes of potential. The related current response is measured, depending on the timing of the pulse (Barker and Jenkins, 1952; Barker and Gardner, 1960; Gosser, 1988; Christian et al., 2013; Ozkan, 2012). Moreover, these techniques are better in sensitivity, due to measuring the current at the end of the pulse resulting in a difference in capacitive and faradaic current (Barker and Jenkins, 1952; Barker and Gardner, 1960; Gosser, 1988; Christian et al., 2013; Ozkan, 2012; Wang, 2006; Brett and Oliveira-Brett, 1998). In normal-pulse voltammetry, which is another type of voltammetric technique, charging current effect is minimized by applying increased amplitude of the pulses. In this technique, applied pulses generally start from the initial potential, and the current response is measured at the end of each pulse (Wang, 2006; Gosser, 1988; Bard and Faulkner, 2001; Brett and Oliveira-Brett, 1993; Ozkan et al., 2015). DPV, which is the most common voltammetric technique, measures the difference between the current just before the end of the pulse and just before application of the pulse. DPV is the one of the most sensitive and applicable electrochemical methods, due to the large ratio between faradaic current and charging current (Wang, 2006; Gosser, 1988; Brett and Oliveira-Brett, 1993; Bard and Faulkner, 2001). Square wave voltammetry, where the net current is measured, is faster than other methods, due to the shorter pulses and higher frequencies. The main difference is that the difference between the current at the end

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of the forward and reverse pulses is measured (Brett and Oliveira-Brett, 1998; Kissinger and Heineman, 1996; Smyth and Vas, 1992). Alternating current voltammetry is generally used to gain information about the kinetics. In this technique, sinusoidal, triangular, or square wavebased alternating voltage is applied with a gradually and linearly change in positive or negative potentials. As a result, voltage alternating current occurs, which is directly related to the composition and the concentration of the electroactive species in the solution (Brett and OliveiraBrett, 1998; Kissinger and Heineman, 1996; Smyth and Vas, 1992; Wang, 2006; Trasatti, 1977). Preconcentration, accumulation of the electroactive species before applying an electrochemical technique, can increase the sensitivity of the analytical method. This procedure is called stripping method and can be coupled with other electrochemical techniques. As they are coupled, the technique can be named with two techniques, such as adsorptive stripping voltammetry, adsorptive stripping DPV, and adsorptive stripping square wave voltammetry. The stripping step can be achieved by controlling the potential, current or open circuit, and, depending on the accumulation step, accumulation potential, or stirring conditions, the quality of the analyses can be different (Brett and Oliveira-Brett, 1998; Kissinger and Heineman, 1996; Smyth and Vas, 1992; Wang, 2006; Trasatti, 1977; Wang, 1985; Ozkan, 2012; Wang, 1988). The working principle of potentiometric biosensors is based on quantitative analyses of species in the measuring media, after monitoring potential changes of a biochemical reaction media designed by both reference and working electrodes. There is a logarithmic relation between the concentration of the species to be analyzed and potential value occurring between working and reference electrodes in the measuring media (Brett and Oliveira-Brett, 1998; Kissinger and Heineman, 1996; Pohanka and Skla´dal, 2008; Ronkainen et al., 2010; Smyth and Vas, 1992). Potentiometric biosensors include glass electrodes sensitive to pH or univalent ions, ion-selective electrodes sensitive to anions and cations, gas-sensitive electrodes to carbon dioxide, or ammonium. In potentiometric biosensors, defined as ion-selective, electrodes that give a response to changes in ion activity and complex biological matrix including ions, such as Na1, K1, Ca21, H1 or NH41, ions are determined from sensitive changes in electrode potentials generated, due to connecting ions to appropriate ion exchange membranes (Wang, 2006; Gosser, 1988; Brett and Oliveira-Brett, 1993, 1998; Bard and Faulkner, 2001; Kissinger and Heineman, 1996; Smyth and Vas, 1992; Berezhetskyy et al., 2008; Wang, 1988). The working principle of conductometric biosensors is based on the fact that concentrations of some ions after biochemical reaction and the conductivity change occurring in the reaction media are monitored by measuring the conductivity between two metal electrode pairs (Pohanka and Skla´dal, 2008; Ronkainen et al., 2010). For field-effect transistor (FET)-based sensors, one of electrochemical transducers mentioned previously, can be designed on a silicon-chip based FET with smaller dimensions. This method is usually used in potentiometric

8.3 Electrochemical Methods in Biosensing

sensors, as well as conductometric or voltammetric sensors (Wang, 2006; Gosser, 1988; Brett and Oliveira-Brett, 1993; Bard and Faulkner, 2001). As one type of promising and inexpensive photoelectric devices, photo electrochemical biosensors are an alternative to conventional analytical methods, due to their high sensitivity and potential in array analysis. In this method, conductive polymers, transition metal complexes, semiconductor nanoparticles, or other semiconductor nanostructures are widely used as photosensitizers on conducting electrodes (Lassalle et al., 2001; Liang et al., 2008; Stoll et al., 2008; Wang et al., 2006; Zhou et al., 2009). The electrons of photosensitizers are excited from their ground state to the excited state to produce electronhole pairs after absorbing photon energy. If an electrode with an appropriate energy level close to either the conductance band or valence band of the photosensitizers is used, the photoexcited electrons or holes transfer to the electrode and produce a photocurrent (Zhou et al., 2009). Because of different forms of energy for excitation (light) and detection (current), this method is very sensitive, with low background signals (Liang et al., 2008; Lassalle et al., 2001; Stoll et al., 2008; Wang et al., 2006; Zhou et al., 2009). Although possessing remarkable advantages, photo electrochemistry is mainly used in the field of solar cells. So far, the photo electrochemical method has been used in the detection of DNA damage, DNA, proteins, and small molecules (Liang and Guo, 2007; Liu et al., 2006; Yildiz et al., 2008; Pardo-Yissar et al., 2003). Increased attention has thus been given to photo electrochemical procedures with electronic detection that are simpler in instrumentation and easier to miniaturize than conventional optical methods. Therefore, it is desirable to improve the utilization of photo electrochemistry in the area of biotechnology (Dzyadevych et al., 2008; Zhang et al., 2010). Electrodes modified by enzyme are constituents of amperometric biosensors and bio fuel-cells. Enzyme electrodes are miniature transporter systems that combine electrochemical processes with immobilized enzyme activity (Dzyadevych et al., 2008; Zhang et al., 2010). These systems show high selectivity to biologically originated substrate molecules, due to the biological function of the enzyme. Therefore, that property of enzymes is an advantage for preparing enzyme sensors. Typically, the kinetics of an enzyme reaction are monitored by the rate of the product or reactant decrease in the amount of the reactant. If the product or reactant is electroactive, the reaction can be directly monitored by amperometry. These types of electrochemical reactions can be used to prepare amperometric biosensors (Liang and Guo, 2007; Liu et al., 2006; Yildiz et al., 2008; Pardo-Yissar et al., 2003). Large amounts of electrochemical biosensors include oxidoreductase enzymes that use or generate electrons during catalytic reaction. These electrochemical reactions occur on the surface of electrodes and are used to generate current. Electrons generated due to enzymatic oxidation are transferred to anode and there they are carried to a higher electron donor, like oxygen atom. Major (and constant) constituents of enzyme electrodes are enzymes and electrodes (Blackstock,

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1989; Holmquist, 2000; Illanes, 2008; Kosland, 1960; Sarma et al., 2009). Significant parameters to select the type of electrode are electrical conductivity and hardness of the material. Therefore, solid-supporting electrodes are used, such as gold (foil or rod), platinum (foil or rod), or carbon (paper, rod, paste, metallized carbon, amorphous carbon, carbon fiber, or nanotube film). Enzymes used for enzyme electrodes are selected according to the reaction type of the analyte to be analyzed. Redox (oxidoreductase) enzymes are the most commonly used. Furthermore, hydrolysis enzymes, such as lipases and esterases, can be used with redox enzymes. The most important problem of enzyme electrodes is that they require oxygen as a cosubstrate and monitoring of the reaction by amperometric method is based on hydrogen peroxide (Blackstock, 1989; Clark and Pazdernik, 2013; Holmquist, 2000; Liu, 2013; Palmer, 1991). This problem can be eliminated by using mediators (Sarma et al., 2009). While preparing enzyme sensors, conductive and functionalized polymers, composite materials, solgels, and nanomaterials can be used as immobilization materials (Blackstock, 1989; Holmquist, 2000; Illanes, 2008; Kosland, 1960; Sarma et al., 2009). Enzyme electrodes can also be used in bio fuel-cells (Sarma et al., 2009). Fuelcells are systems that convert chemical energy into electrical energy. Bio fuelcells are named according to the type of the catalyst, while enzymes catalyzing the oxidation of fuels are called enzymatic bio fuel-cells (Davis and Higson, 2007). Anode and cathode are separated from each other by using a semipermeable membrane; however, structures with single-division are also available. Biological species can be present at the anodic compartment or immobilized on the surface of electrode. When appropriate fuel is added, it is partially or completely oxidized at the anode, and oxygen is reduced by carrying electrons to the cathode (Davis and Higson, 2007; Pohanka and Skla´dal, 2008; Gerard et al., 2002; Gronow, 1991; Kuila et al., 2011; Malhotra et al., 2005; Miller and Nagarajan, 2000; Scheller and Schubert, 1992). Enzymatic bio fuel-cells have higher power density (the power generated on the surface of electrode, W/cm2) than microbial fuel-cells, but their lives are shorter (typically 710 days). If enzymes with appropriate selectivity are used at anode and cathode, the necessity of membrane to separate anode and cathode is eliminated, which is an important advantage of enzymatic bio fuel-cells (Blackstock, 1989; Davis and Higson, 2007; Holmquist, 2000; Illanes, 2008; Minteer et al., 2007).

8.4 CONCLUSION Nanobiodevices have gained a great deal of importance, due to their advantages in terms of response speed, low cost, portability, small sample and reagent consumption, and low energy requirements. The combination of sensing molecules that are in biorecognition and transducer parts is the main aspect of biosensors.

References

The choice of suitable matrix, monitoring/quantifying the interactions between the analytes and the biorecognition part is very important in designing biosensors. This chapter it is tried to summarize and collect together, all nanobiodevices for electrochemical biosensing of pharmaceuticals. Different biorecognition parts, such as enzymes, DNA, tissues, bacteria, yeast, antibodies, antigens, liposomes, and organelles were discussed. Nanobiodevices, such as LOC platforms in biosensors and their applications in pharmaceutical analysis are also discussed. Recent studies about DNA biosensors, enzyme-based nanobiodevices, immunosensing applications were also shared.

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CHAPTER

Imprinted polymeric nanoparticles as nanodevices, biosensors and biolabels

9

´ Monika Sobiech and Piotr Lulinski Medical University of Warsaw, Warsaw, Poland

CHAPTER OUTLINE 9.1 Introduction .....................................................................................................331 9.2 Principles of Imprinting Process .......................................................................332 9.3 Overview of Formats of Imprinted Nanomaterials ................................................334 9.4 Imprinted Drug Delivery Nanodevices ................................................................335 9.5 Imprinted Nanosorbents for Sample Preparation.................................................344 9.6 Imprinted Nanocomposites for Biosensors .........................................................352 9.7 Biolabeling With Imprinted Polymeric Nanostructures ........................................363 9.8 Conclusions.....................................................................................................363 References .............................................................................................................364 Further Reading ......................................................................................................374

9.1 INTRODUCTION During the lecture presented by honored Nobel Prize laureate Richard Phillip Feynman, the concept of a new field of science, currently known as nanoscience, was revealed: “. . . the principles of physics, as far as I can see, do not speak against the possibility to maneuvering things atom by atom. It is not an attempt to violate any laws; it is something in principle, that can be done; but, in practice, it has not been done because we are too big” (Feynman, 1960). Since that time, intensive studies that aimed to explore the field of nanoscience were carried out. Those fundamental investigations disclosed the unique physical and chemical properties of nanomaterials (Granqvist, 1977; Comsa, 1977), and revealed their practical utility (Kreuter, 2007). Nanomaterials are defined as materials with at least one dimension ranging between 1 and 100 nm. This extremely small size provides extraordinary properties, such as: an increase of its catalytic activity due


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CHAPTER 9 Imprinted polymeric nanoparticles as nanodevices

to their huge surface area, high ratio of surface atoms to total number of atoms, and increased surface energy. Those features result in enormous utilization of nanomaterials in life sciences, especially in pharmaceutical field of knowledge, as advanced drug carriers. Molecular imprinting is a technique for the manufacturing of synthetic polymeric materials with high recognition ability for a particular target analyte. The high specificity of the imprinted polymer is a result of the fabrication process, based on the template-tailored polymerization of functional monomers. The imprinted materials have a vast application in separation techniques, catalysis, drug delivery, sensing, and antibody replacement in diagnostics (Alvarez-Lorenzo and Concheiro, 2013). However, the elicitation of all benefits of molecular imprinting technology imposes the quest for new formats of polymers and shifts the material seizure to nanoscale. Here, the significant progress in material science and engineering has allowed development of a variety of formats of nanomaterials that can be compatible with molecular imprinting technology, offering considerable advantages in physicochemical characterization of sorption processes on imprinted polymers (higher surface-to-volume area, better access to the higher number of three-dimensional cavities, lower diffusion times). In this chapter, the recent advances in synthetic approaches for the fabrication of imprinted nanomaterials, together with diversity of formats for possible applications in pharmaceutical science will be outlined. The physicochemical behavior of imprinted nanostructures will be discussed in the context of their practical utility. Finally, the current limits and future prospects for the imprinted nanomaterials will be pointed out.

9.2 PRINCIPLES OF IMPRINTING PROCESS The molecularly imprinted polymers are characterized by the high level of selectivity, due to the presence of specific recognition sites formed in the polymer network by the template-tailored synthesis. The synthesis of imprinted materials consists of three steps: the formation of prepolymerization structure, the polymerization reaction, and the template removal (Fig. 9.1). The formation of the prepolymerization structure could be obtained by covalent or noncovalent strategies. The covalent approach assumes the chemical reaction between template molecule and functional monomer in order to form functionalized compound prior to the polymerization (Wulff et al., 1972). The noncovalent approach utilizes the range of weak intermolecular interactions, such as ionic forces, hydrogen bonds, van der Waals forces, or ππ interactions that can exist between the template molecule and functional monomer in order to form the prepolymerization complex prior to the polymerization (Arshady and Mosbach, 1981). There are a few critical moments during the synthesis of imprinted polymers that can strongly affect the property of imprinted materials and effectiveness of the

9.2 Principles of Imprinting Process

FIGURE 9.1 The schematic idea of the imprinting process.

imprinting process. The most important factors are as follows: the choice of template, the preselection of functional monomer, and effective template removal. The choice of template is one of the crucial moments of the synthesis of imprinted materials. The selected structural elements of the template molecule are required to facilitate the imprinting process. The presence of functional groups or heteroatoms that could form covalent bonds or that could noncovalently interact with functional monomers is the main factor for the stabilization of the prepolymerization structure. The aromatic, heteroaromatic or polyaromatic rings increase the stabilization and, together with cyclic systems as well as aliphatic chains, are responsible for the formation of steric effects inside the polymer network. In principle, the template and the target molecule ought to be the same (the target molecule is defined as the bioanalyte or pharmacologically active compound to which the imprinted polymer is dedicated). However, a majority of biocompounds or drugs possess physicochemical properties that could limit their application in the

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synthesis of molecularly imprinted polymers, due, for instance, to insufficient solubility in organic solvents. Simultaneously, the economic aspects ought to be taken into account because a lot of biocompounds or drugs are expensive, or rarely available. Consequently, the structural analogues of target molecule as the template during the imprinting process are widely applied (Xu et al., 2012). An important parameter in the design of imprinted materials is the appropriate selection of functional monomers that should interact with the template molecule via complementary functionalities. It could be a time-consuming process because the validation of the proper monomer can proceed after the synthesis and analysis of binding properties of the imprinted polymer. Here, theoretical analyses have been employed as a powerful tool in the preselection of functional monomers (Cowen et al., 2016). Finally, the template molecule must be removed from the polymer matrix to leave three-dimensional cavities empty (Lorenzo et al., 2011). This stage is very important during synthesis, because only total removing of the template molecule allows for obtaining specific binding sites in the imprinted polymer network. This step is more effective when nanoscale imprinted particles are formed, due to facile access of removal solvent to templated sites. Consequently, the imprinting process is considered to be the process that allows modification of the surface of polymer by the template molecule. The main parameter describing the efficacy of the imprinting process is the imprinting factor. In the simplest way, the imprinting factor is defined as ratio of the binding capacity of the template on the imprinted polymer to the binding capacity of the template on the reference nonimprinted polymer. Hence, the synthesis of nonimprinted polymer has to be carried out in the same conditions omitting the addition of template molecule (Alvarez-Lorenzo and Concheiro, 2013). Above presented brief overview of the synthetic aspects of imprinted polymers outlined the most important moments of the imprinting process. The imprinting technique has gained widespread attention, mainly due to its application in the fabrication of molecularly imprinted sorbents for the purpose of selective separation of compounds, in biosensing, as catalysts, or as potential drug delivery vehicles. There are some excellent books and reviews that provide more detail information regarding the synthesis and application of imprinted polymers (Wulff, 1995, 2002; Alexander et al., 2006; Piletsky and Turner, 2006; AlvarezLorenzo and Concheiro, 2013; Luliński, 2013; Whitcombe et al., 2014; Zaidi and Shin, 2014; Li et al., 2015; Sharma et al., 2015; Cieplak and Kutner, 2016; Iskierko et al., 2016; Wackerlig and Schirhagl, 2016; Zaidi, 2016).

9.3 OVERVIEW OF FORMATS OF IMPRINTED NANOMATERIALS Apart from the widespread application of molecularly imprinted materials in separation science and commercialization of selected imprinted sorbents, the main

9.4 Imprinted Drug Delivery Nanodevices

drawback of imprinted materials is related to the manufacturing process and limited scale-up capabilities. Accordingly, significant efforts were taken up to implement current advances of nanoscience into imprinting technology. Merging both nanoscience and imprinting techniques, new formats of imprinted nanomaterials were synthesized, aiming to improve inconvenient scale-up production, as well as to overcome problems relating to traditional imprinted materials, such as low binding capacity, troublesome template removal, insufficient accessibility of binding sites, high mass-transfer resistance, difficult control over size and shape of particles, and low effectiveness of imprinting process (Poma et al., 2010). Consequently, for last few years, new formats of imprinted nanomaterial were worked out. Amongst them, we can find imprinted nanobeads, imprinted layers on core-shell magnetic nanoparticles, quantum dots (QDs), carbon nanotubes or nanofibers blended with imprinted nanoparticles, imprinted nanogels, nanowires, nanofilaments, nanosponges, as well as nanocomposites and imprinted photonic crystals (Wulff et al., 2006; Flavin and Resmini, 2009; Tokonami et al., 2009; Kim and Chang, 2011; Linares et al., 2011; Servant et al., 2011; Li et al., 2013; Lv et al., 2013; Suedee, 2013; Meshur et al., 2014; Deshmukh et al., 2015; Ma et al., 2015; Poma et al., 2015; Chen et al., 2016; Hamdan et al., 2016; Wang et al., 2016a). Those formats are characterized by higher surface-to-volume area, providing better kinetics of sorption processes. As a result, higher availability of imprinting sites for analytes is observed and the effectiveness of imprinting process has increased. The micrographs of exemplary imprinted nanomaterial formats obtained by scanning or transmission electron microscopy are presented in Fig. 9.2. The diversity of formats of imprinted nanomaterials prepared in the last few years reveals enormous progress in this field. Selected examples of advanced research, presenting application of imprinted nanomaterials for drug delivery, separation, sensing and bioimaging, are discussed below. The literature survey shows no evidence today for scale-up manufacturing of imprinted nanoparticles and commercial application is still immature. However, the examples mentioned further on can serve as a starting point for a new generation of high affinity imprinted nanomaterials that could be commercialized in the near future.

9.4 IMPRINTED DRUG DELIVERY NANODEVICES Imprinted drug delivery nanodevices can be a promising alternative for traditional pharmaceutical formulations which do not fulfill the demands of modern pharmacotherapy. These materials could offer improved transport properties, provide optimized pharmacokinetic profiles, control the drug release rate and maintain the drug concentration within its therapeutic window, as well as enhance delivery efficacy by increasing diffusivity and biodistribution. Moreover, optimum drug delivery carriers should be synchronized with the physiological status of the

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FIGURE 9.2 The exemplary formats of imprinted nanomaterials: (A) transmission, (B) scanning electron microscopy images of imprinted multiwalled carbon nanotubes as novel artificial antibody for troponin T (Moreira et al. 2011), (C) inverse opal of molecularly imprinted photonic crystals as colorimetric detectors for enantioresolution of pyroglutamic acid (Zhang et al. 2013a), (D) molecularly imprinted membrane on multiwalled carbon nanotubes for recognition of bovine serum albumin (Zhang et al. 2010), (E) molecularly imprinted poly(ethylene-co-vinyl alcohol)/quantum dot composite nanoparticles for urinalysis (Lin et al. 2009). Source: Reproduction of (A)(E) with permission of Elsevier.

patient and should provide a drug in response to the changing intracorporeal environment. The imprinted drug delivery nanodevices also have great potential because of releasing drugs in the feedback regulated way, which is extremely important in a modern model of pharmacotherapy, oriented towards the delivery of the best suited drug to a single patient in the right place and at the right time. Additionally, the imprinted nanodevices are stereoselective materials, capable of maximizing the delivery of a given eutomer, the isomer of interest, and of reducing or even eliminating the delivery of the distomer, the undesirable isomer. These properties make imprinted nanodevices very promising drug delivery carriers. Nevertheless, aspects of biocompatibility and toxicity must be considered before final implementation into pharmacotherapy (Luliński, 2013; Suedee, 2013; Zaidi, 2016). In one advanced paper, an anticancer agent, 5-fluorouracil, imprinted nanodevice for drug delivery was worked out for the intravenous route of administration. The imprinted drug delivery nanocarrier was designed to overcome existing limitations: submicroscale particles are not appropriate for intravenous administration, the precise targeting of cancer cells is impossible, and the polymer matrix is not biodegradable. The treatment with 5-fluorouracil is troublesome because


this drug is quickly metabolized, providing insufficient level in serum and decreasing the therapeutic activity. Conversely, continuous drug administration provoked severe toxic effects. Consequently, new magnetic nanoparticles were fabricated for control delivery of 5-fluorouracil (Hashemi-Moghaddam et al., 2016). The magnetic nanoparticles have important advantages as drug delivery carriers because their size can be controlled to match the system of interest, their delivery can be manipulated by an external magnetic field, and they can be easily modified and functionalized for different therapies. However, the composition of magnetic nanocarriers is critical to ensure appropriate therapeutic effects when targeting a specific site. These entities must be coated to prevent drug conjugation and particle agglomeration, as well as to limit the interactions with nontargeted cell sites. As it was mention previously, the control of particle size is crucial for an effective delivery carrier. It is well known that nanoparticles below diameter of 100 nm can be smoothly phagocytized by the liver cell. Unfortunately, it is very difficult to obtain such nanoparticles while typical composition of imprinted drug delivery hydrogels is applied (viz. acrylic or methacrylic acid as the functional monomers, 2-hydroxyethyl methacrylate as the backbone monomer, and N,N-methylenebisacrylamide as the cross-linker). Accordingly, in the construction process, the Fe3O4 core of nanoparticles was coated with a thin layer of polydopamine, imprinted by the 5-fluorouracil template. The core provides magnetic properties to the nanoparticle, which can be used to guide the nanodevice to a disease site by applying an external magnetic field. Such a property of the imprinted nanocarrier is very promising in targeted delivery of a drug. Moreover, the size of particles is appropriate for intravenous administration, with an average diameter of 80 nm. Another advantage is that the polydopamine layer can be easily modified and functionalized for different therapies. The release pattern of 5-fluorouracil from imprinted nanodevice revealed its continuous character for a period of 48 h and nearly 80% of drug release was noted within the first four hours of experiment. The release of 5-fluorouracil reached its maximum in 24 h and remained constant to 48 h. The high initial ratio of release of the drug can be explained by the surface distribution of 5-fluorouracil binding sites, which facilitates discharge. In the in vivo experiments, the mouse breast cancer model was applied with murine mammary adenocarcinoma cells that were transplanted into the flank region of female Balb/C mice. The 5-fluorouracil imprinted nanoparticles were dispersed in distilled water and were administrated to tumor-bearing mice via the tail vein. Five test groups were investigated: control without treatment, sham with implanted magnet without treatment, a group treated with only 5-fluorouracil, a group treated with the imprinted nanodevice soaked with 5-fluorouracil without external magnet use, and a group treated with the imprinted nanodevice soaked with 5-fluorouracil with external magnet use. The histopathological analyses revealed that the total score of malignancy in the last group was reduced. It was concluded that the new imprinted nanodevice increased the local release of 5-fluorouracil, enhanced the control of

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tumor growth, and extended the life span of the examined mice. However, during in vitro release experiments, a major burst effect was observed, which was an important drawback. In order to improve biocompatibility of the imprinted nanocarriers for targeted delivery of 5-fluorouracil, the synthetic reagents were modified. Here, a new cross-linker, tannic acid, was applied to provide the biodegradable matrix of magnetic nanoparticles (Asadi et al., 2016a). The core-shell magnetic nanodevice was prepared in three steps. First, Fe3O4 nanoparticles were synthesized. Then, the magnetic nanoparticles were modified by a silica layer and fluorescein isothiocyanate was incorporated into the structure of nanoparticles to facilitate their location and distribution control inside the body. Finally, the nanoparticles were modified by 3-(trimethoxysilyl)propyl methacrylate prior to imprinting process, including 5-fluorouracil (template) and tannic acid (cross-linker). A comprehensive physicochemical analysis of the new material was carried out, together with in vitro degradation tests, in vitro cytotoxicity, and in vivo release experiments. The images from transmission electron microscopy, as well as dynamic light scattering analysis, revealed the diameter of imprinted nanoparticles between 10 and 30 nm. The degradation studies, which are very important considering biomedical application, were performed at various pH 3, 7.4 and 11 (for intravenous administration, pH 7.4 corresponds to pH of blood and pH 11 to a basic kidney environment). The results revealed faster degradation at pH 3 or pH 11 than at pH 7.4. The results confirmed the utility of the innovative biodegradable material, which, after intravenous administration, could survive in the physiological blood environment and deliver the drug to the target site, followed by removal to kidneys where it could be degraded. The cytotoxicity tests of the imprinted material (without the drug) were performed on NIH/3T3 cell lines and the cell viability did not show any significant difference for 7 days. The biocompatibility was also evaluated on human embryonic kidney (HEK293) cell line, ensuring no changes in morphology or cell death symptoms over a period of 3 days. This means that the material is not toxic for living cells. In vivo experiments were performed on healthy rats, to which a certain amount of imprinted drug-loaded nanodevices was injected intravenously. After applying an external magnet into the region of the kidney, fluorescent images were performed after 48 h, revealing a high concentration of nanoparticles in this organ. A similar approach was proposed in another paper (Asadi et al., 2016b). Here, olanzapine, the second generation of antipsychotic drug, was used as the template. The poor bioavailability and low permeability into the brain constrains the clinical application of olanzapine. Conversely, high and frequent dosing promotes extra-pyramidal side-effects, tremors, and somnolence. Consequently, new imprinted nanodevice was developed. Here, the methacryloyl derivative of fructose was applied as a new, biocompatible cross-linker. The applied strategy involved fabrication of magnetic fluorescent multi core-shell structure of imprinted material, which facilitated the aggregation of the carrier near the target


FIGURE 9.3 A schematic illustration of the preparation procedure of biodegradable magnetic fluorescent molecularly imprinted polymer for targeting drug delivery of olanzapine under external magnet field (OZ, olanzepine; TEOS, tetraethoxysilane; MPS, 3-(trimethoxysilyl) propyl methacrylate; APS, (3-aminopropyl)trimethoxysilane; FITC, fluorescein thiocyanate). Source: Reproduction with permission of Elsevier.

tissue under external magnetic field. The schematic illustration of the preparation procedure of biodegradable magnetic imprinted nanodevice for targeting drug delivery of olanzapine is presented in Fig. 9.3. Morphological characterization was carried out. The surface morphology and crystalline size growth of composite nanoparticles were studied by transmission electron microscope, revealing spherical entities with an average diameter of 20 nm. The narrow size distribution of particles could be related to the effect of solvent used in the fabrication step. The polydispersity index and z-average for moleculary imprinted polymer (MIPs) were 0.14 and 58 nm, respectively. Atomic

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force microscopy was also employed in morphological analysis. The cytotoxicity of imprinted nanoparticles was measured on NIH/3T3 cell-line using 3-(4,5dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) viability assay. The results showed no dramatic change in cell viability during 5 days of incubation, proving the nontoxic nature of imprinted material. The fluorescent images of rat following targeting olanzapine delivery are presented in Fig. 9.4AE. The pharmacokinetics profiles of control and test groups are presented in Fig. 9.4F. In another excellent paper, the temperature and magnetism bi-responsive molecularly imprinted polymer was proposed for the purpose of 5-fluorouracil delivery (Li et al., 2016). Here, the low critical solution temperature and reversible solubility in aqueous solution at 32 C of N-isopropylacrylamide were gained in the synthesis of imprinted drug delivery nanodevice. In the fabrication process,

FIGURE 9.4 Fluorescent images of rat to follow targeting drug delivery to brain under magnetic field, (A) before injection of olanzapine carrier, (B) 24 h after injection, (C) 48 h after injection, (D) 72 h after injection, (E) the tested rat under magnetic field, (F) plasma concentration of olanzapine loaded on imprinted nanoparticles (test group) in comparison with alone olanzapine administration (control group). Source: Reproduction with permission of Elsevier.


the surface grafting copolymerization method was employed. First, Fe3O4-encapsulating carbon nanospheres (Fe3O4@C) were prepared, following silanization with an activated surface by modification of 3-(trimethoxysilyl)propyl methacrylate (Fe3O4@CSi). Then, the surface of Fe3O4@CSi was grafted by N-isopropylacrylamide functional monomer to form Fe3O4@CSi@PNIPAM. Finally, Fe3O4@CSi@PNIPAM was imprinted on the surface using 5-fluorouracil (template) and N,N-methylenebisacrylamide (cross-linker). After synthesis, the template was removed, and careful characterization of particles was carried out, including scanning and transmission electron microscopy, infra-red spectroscopy, thermogravimetry, dynamic light scattering, and vibrating sample magnetometry analyses. Highly monodispersed, uniform and spherical nanoparticles possessed average diameter of 152 nm. The transmission electron microscopy analysis revealed typical core-shell structure, with an amorphous coating layer covering the inner magnetite core with thickness of 50 nm. The advantage of imprinted magnetic nanoparticles of Fe3O4@CSi@PNIPAM can be related to targeting of the specific site of the body via external magnet, as well as to its potential to realize magnetic hyperthermia that could induce the phase transition of temperaturestimuli polymer, resulting in controlled release of the drug. The loading capacity of imprinted Fe3O4@CSi@PNIPAM and its nonimprinted counterpart were equal to 94.54 and 61.77 mg/g, respectively. In vitro release studies revealed that at 25 C, nearly 70% and 84% of 5-fluorouracil adsorbed on imprinted and nonimprinted Fe3O4@CSi@PNIPAM was released, respectively. In the elevated temperature, the release from imprinted Fe3O4@CSi@PNIPAM was equal to 91% because the shrinking of polymer disrupted the interactions between the template and the monomer. The additional value of this paper derived from molecular modeling that was used for the prediction of the strength of interactions between 5-fluorouracil and N-isopropylacrylamide. The computationally designed nanoparticles of molecularly imprinted polymers as drug delivery systems for release of naltrexone, an efficient drug used in treatment of alcohol and narcotic addiction, were described by Rostamizadeh and coworkers (2012). Here, the density functional theory method at the Becke, 3-parameter, Lee-Yang-Parr (B3LYP) level of theory, in conjugate with 631 1 G(d) basis, was employed to estimate the interactions between the drug and the monomers. In an efficient pharmacotherapy, two important factors must to be considered: targeting the specific site in the body as well as ensuring the interaction of drug with receptor and penetration across the biological membranes. A new, advanced idea was proposed by Esfandyari-Manesh and coworkers (2016). In order to facilitate the transport of a carrier with drug into the cell, the molecularly imprinted polymer conjugated with a polyethylene glycolfolic acid layer was synthesized. The imprinted nanoparticles were obtained by miniemulsion polymerization technique and the antineoplastic agent, paclitaxel, was used as the template, together with methacrylic acid, methyl methacrylate, and ethylene glycol dimethacrylate as functional monomers and cross-linker, respectively. A carbodiimide-mediated

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reaction was employed to conjugate the polyethylene glycolfolic acid layer. The 1 H NMR spectroscopy was used to confirm the structure of nanoparticles. The release study revealed significantly greater initial release, due to weakly adsorbed drug on the surface of nanoparticles. The pH-dependence of release was also noted. It was concluded that paclitaxel resolved faster in acidic condition and the release rate was faster in lower pH, which could be beneficial in cancer treatment. In vitro cytotoxicity tests were carried out on MDA-MB-231 (folate positive) and A549 (folate negative) cancer cell lines. The paclitaxel imprinted polymer conjugated with polyethylene glycolfolic acid layer possessed higher anticancer activity when compare to reference paclitaxel imprinted polymer (without layer) and paclitaxel itself (IC50 equal to 4.9, 7.4, and 32.8 nM, respectively). Consequently, the superiority of the new drug delivery carrier was confirmed. Conjugated folic acid fragment enhanced interactions of drug nanocarrier with the folate receptor to facilitate the penetration into the cell. A very similar idea was proposed by Zhang and coworkers (2015). In their study, a disulfide-linked α-helix-containing peptide (a hybrid apamin-p32 polypeptide) was used in the epitope approach to mimic the N-terminal part of the membrane protein p32. The nanoparticles, of average diameter of 37 nm, were formed with capability of potent adsorption to the target protein, recognizing p32 positive tumor cells. The fluorescence polarization experiments were carried out to quantify the interactions of imprinted nanoparticles, with p32 revealing successfully mediated targeted photodynamic therapy. Very low solubility in water as well as low bioavailability after oral administration, fast metabolism, and rapid excretion were the reasons for designing a new delivery nanodevice for curcumin (Zhang et al., 2014a). The imprinted-like biopolymeric micelles for curcumin delivery were synthesized, based on gelatindextran conjugate support cross-linked by genipin in the presence of 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide acting as a template molecule, as well as a structural analog of the pharmacologically active curcumin target. The strategy involving structural analogs as the templates during the imprinting stage is widely used to improve the efficiency of imprinting or to reduce costs of fabrication when a MIP is dedicated to a highly labile or costly drug. After synthesis, physicochemical analysis and curcumin-loading to the nanodevice, the in vitro release studies were performed in media at various pH: 2, 5, 6.8, which corresponded to pH of various parts of the gastrointestinal tract or intracellular components. The results revealed a sustained release of curcumin from the imprinted-like biopolymeric micelles with desorption of 54%, 47%, and 60% of curcumin loaded at pH 2, 5, and 6.8, respectively after 72 h. In addition, the curcumin-imprinted vehicle was highly dispersible, expecting that the bioavailability of curcumin could be improved due to high dispersity. The cytotoxicity tests of the curcumin-free imprinted vehicle were performed towards L-929 cells, showing its low impact on cell viability. Finally, in vivo oral bioavailability tests were carried out. The plasma pharmacokinetics after oral intake were studied


on the rat model involving male Sprague-Dawley rats. The free curcumin suspension was used as a control. The results revealed that the curcumin concentration in plasma after intake of 100 mg/kg of the curcumin-loaded imprinted nanodevice was higher than that after intake of 500 mg/kg of free curcumin suspension. However, a similar phenomenon was observed in other curcumin nanoformulations. Accordingly, it seems that the nanoscale of the device, and not the imprinting process itself, is responsible for better pharmacokinetic characteristics. Considering the transdermal route of administration, a propranolol-imprinted polymer nanoparticle-on-microsphere porous cellulose membrane for the enantioselective delivery of racemic propranolol was investigated (Jantarat et al., 2008). The enantioselective release of S-propranolol (eutomer), its diffusion, and transit across the excised rat skin in in vivo experiments, were investigated using confocal laser scanning microscopy. The mechanism underlying the release of S-propranolol from imprinted composite membrane was found to involve specific adsorption and mobility of this enantiomer at the binding site in imprinted composite membrane during its transition from dry to wet state. It was concluded that the proposed imprinted composite membrane could be used for self-controllable permeability systems, responding to the presence of target solutes. Predominantly imprinted drug delivery devices reveal quite selective binding, only at low concentrations. The result is a low capacity of the imprinted polymer. This problem is very important, considering its application of the drug dosage forms due to considerably limited utility. Sufficient loading capacity is necessary to ensure a prolonged release of the drug. Typically, the imprinted polymer is drug-loaded by soaking in an equilibrium-dependent procedure. Otherwise, the loading can be performed in situ during the imprinting process, introducing the drug as a part of the prepolymerization mixture. The latter procedure increases the capacity of imprinted drug delivery devices but is limited to drugs which are stable during the polymerization process. Here, an improvement of the binding capacity of drug nanocarrier was investigated in a pH/glutathione double-responsive drug delivery system for doxorubicin (Zhang et al., 2016). The results indicated that loading drug during the imprinting process could lead to multiple release ways, with efficient pHand glutathione-controlled manner increasing the capacity of formulation. However, the precise determination of the amount of the drug after the polymerization and postpolymerization preparation may be difficult to perform and there is no guarantee that the whole amount of the drug was introduced to the polymer network during the preparation process (Byrne et al., 2008). The examples mentioned previously revealed the significant progress in the field of molecularly imprinted nanocarriers. Moreover, a lot of recent works show the results of in vivo experiments which are an extremely important part of the way to the implementation of such drug vehicles in modern pharmacotherapy. There is still a lot of investigation ahead that must be carried out, but the future of imprinted nanocarriers is very promising.

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9.5 IMPRINTED NANOSORBENTS FOR SAMPLE PREPARATION The major technological development of recent decades introduced innovative instrumental techniques into pharmaceutical analysis. Moreover, miniaturization in separation techniques has gained wide attention in pharmaceutical analysis, because it offers the reduction of costs and minimal use of toxic and environmentally unsafe solvents. The current analytical protocols were simplified after the introduction of advanced analytical instruments. The time of analysis and the error-prone parts of analytical schemes were reduced. The novel methods for separation and analysis of compounds, such ultraperformance liquid chromatography combined with chemiluminescence detection or mass spectrometry, have become widely applied. Current instrumental analysis determines compounds which occur at very low concentrations. However, the accuracy of analytical methods is limited, due to the presence of interfering compounds in highly complicated biological or environmental samples. Consequently, the clean-up step is mandatory to separate the selected analytes. Liquidliquid extraction or solid phase extraction are most frequently used techniques of separation and enrichment of analytes. Liquidliquid extraction suffers from the use of vast amounts of toxic solvents and is considered today as not environmentally benign. The low selectivity of commercial stationary phases hampered the practical application of solid phase extraction. Therefore, novel and advanced sorbents are required. Among the promising materials used for selective separation are molecularly imprinted polymers. Nevertheless, up to date, a few limitations of imprinted sorbents can be identified, such as template bleeding, restricted mass transfer, high or instable pressure, and irregular flow time while use in solid phase extraction cartridges. Those drawbacks are mainly related to bulk format of material and classic extraction procedures, but they can be overcome using new formats of imprinted sorbents and applying dispersive extraction procedure. Imprinted nanomaterials are a promising solution because they can be easily used in dispersive solid phase extraction (for instance with additional magnetic properties), they can provide facile access of analytes to adsorption sites and they can be easily washed out from the template. The simple approach to fabricate imprinted nanoparticles is to employ the precipitation polymerization technique with a high volume of solvent. In such way, Ebrahimzadeh and coworkers obtained molecularly imprinted nanoparticles (MINs) for the selective extraction of loratadine (Ebrahimzadeh et al., 2013a). The prepolymerization system consisted of methacrylic acid and ethylene glycol dimethacrylate; loratadine was used as the template molecule. The precipitation polymerization was carried out in acetonitrile. The optimization of stoichiometry between functional monomer and cross-linker revealed that the ratio 1:5 produced nanoparticles with the average diameter of 100 nm. The fabricated imprinted

9.5 Imprinted Nanosorbents for Sample Preparation

nanoparticles were used as sorbents in the solid phase extraction of loratadine from plasma and urine samples. After the extraction and centrifugation, the subsequent step of miniaturized liquidliquid extraction was employed to enhance analytical performance. Finally, loratadine was determined in high performance liquid chromatography coupled with diode array detector. The proposed methodology was effective for monitoring of the drug concentration in urine of healthy volunteers who consumed a single oral dose of 10 mg tablet of loratadine. The limit of detection and recovery were 0.2 μg/L and 90%, respectively. In summary, it was noted that this procedure had several advantages, such as lack of matrix effect, selectivity, simplicity, low consumption of organic solvents, and safety, as well as applicability to clinical laboratories. The same group (Ebrahimzadeh et al., 2013b) used imprinted nanoparticles as sorbents in the dispersive solid phase extraction of haloperidol from spiked plasma and urine samples. The separation step was combined with analysis of haloperidol by high performance liquid chromatography, coupled with a diode array detector. The precipitation polymerization technique was used during synthesis of imprinted nanosorbent. The dispersive solid phase extraction protocol was optimized. The effects of pH and contact time were considered for loading, and the effects of composition, volume, and time of elution were investigated. The optimized protocol consisted of loading with an alkaline sample solution of volume of 5 mL (pH 5 10) for 90 min at room temperature, washing (after sedimentation) with 1.5 mL of methanolacetic acidtrifluoroacetic acid (79.9:20:0.2 v/v/v) and elution with the same system for 75 min at 50 C. The novel analytical method was validated for determination of haloperidol in spiked urine, as well as plasma, and the following digits of merit were obtained: limits of detection and quantification were equal to 0.3 and 2 μg/L, respectively (linear range between 2 and 10,000 μg/L), intra- and interday precision between 3.8% and 6.9%, and recoveries between 97.3% and 98.2%. In conclusion, it was stated that imprinted nanoparticles used in the separation step of new analytical method allowed for very high recovery of haloperidol from spiked urine and plasma samples, with good precision, high preconcentration factor, low matrix effect, and low consumption of organic solvents. In another approach, Chauhan and coworkers applied nanosized multitemplate imprinted polymer for the simultaneous extraction of polycyclic aromatic hydrocarbon metabolites from urine samples (Chauhan et al., 2015). Here, the precipitation polymerization was employed and three templates were simultaneously used in the prepolymerization mixture: 1-naphtol, 9-hydroxyfluorene, and 9-phenanthrol in equimolar stoichiometry. Methacrylic acid and ethylene glycol dimethacrylate were used as the functional monomer and cross-linker, respectively. After the synthesis and postsynthetic treatment, the nanoparticles were packed into the solid phase extraction cartridges and a new extraction protocol was proposed. It involved conditioning with 2 mL of water, followed by 2 mL of acetonitrile, loading with 2 mL of sample, washing with 1 mL of water and

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elution with 5 mL of methanolacetic acid (80:20 v/v). The analytes were determined by ultrahigh performance liquid chromatography, coupled with fluorescence detection. The validation of the analytical method was carried out for five polycyclic aromatic hydrocarbons, viz. 1-hydroxypyrene, 2-hydroxyfluorene, 9-hydroxyfluorene, 9-phenanthrol, and 1-naphtol, with limits of detection equal to 2.6, 0.9, 0.62, 0.4, 0.33 μg/L, respectively and limits of quantification equal to 8, 2.7, 2, 1.2, and 0.99 μg/L, respectively. The intra- and interday precision for all tested compounds varied between 3.78% and 9.75%, and recoveries were between 78% and 95%. The application of the analytical method, based on the separation of new nanosorbent, was confirmed in real urine sample analysis. The samples were collected from male kitchen workers in commercial hotels and smokers, revealing the presence of 1-naphtol (1125 μg/L), 9-phenanthrol (14380 μg/L), and 1-hydroxypyrene (23169 μg/L). The superiority of the innovative analytical method was showed in comparison to liquidliquid extraction, which was employed in a separation step instead of solid phase extraction. However, the comprehensive analyses of interactions in the prepolymerization mixture was omitted. The highest specificity was obtained for 1-naphtol (imprinting factor equal to 3.7), which could be related to its low molecular volume. Investigations aimed to obtain the imprinted polymers with high affinity to specific enantiomer were forwarded by combination of imprinted polymers with modern nanostructure material. In a very interesting example, Vandevelde and coworkers described the fabrication of hierarchically nanostructured polymer films, based on molecularly imprinted surface-bound nanofilaments (Vandevelde et al., 2007). In the synthesis, porous alumina was used as a sacrificial nanoporous template material, followed by nanomolding that produced surface-bound nanofilaments. The size and shape of nanofilaments differed due to the adjustment of the pore structure of alumina. This process was controlled by optimized electrooxidation and postoxidation treatment. In the next step, the glass slides were modified to allow for covalent bonding of polymer layer prior to applying of copolymerization system, containing trimethylolpropane trimethacrylate, methacrylic acid, and S-propranolol. Finally, the cover layer of alumina was provided. The nanostructured imprinted material was characterized as highly selective during the resolution of S and R enantiomers of propranolol. The material possessed high binding capacity and superb accessibility of binding sites. The practical application of the material was not shown, but it could be expected that further experiments will provide more information. Yoshimatsu and coworkers described the synthesis of MINs that were encapsulated using electrospinning method, to form nanofiber affinity membranes. The material was used in dispersive solid phase extraction process for separation of propranolol from spiked tap water in environmental analysis (Yoshimatsu et al., 2008). Methacrylic acid served as a functional monomer and racemic propranolol was used as the template in the precipitation polymerization. In optimization of the synthetic process, two cross-linkers were tested, viz. trimethylolpropane


trimethacrylate and divinylbenzene, but only the latter provided imprinted nanoparticles with an average diameter of 130 nm. Next, the encapsulation of imprinted nanoparticles was carried out into poly(ethylene terephtalate) nanofibres by electrospinning. It was noted that only nanoparticles were effectively encapsulated into nanofibers. The dispersive solid phase extraction of propranolol was performed in the following steps: conditioning with 1 mL of methanol and 1 mL of water (each time for at least 1 h), loading with 1 mL of water sample solution, washing with 1 mL of acetonitrileammonium acetate buffer (80:20 v/v) for 1 h, elution with 1 mL of acetonitrile-wateracetic acid (40:30:30 v/v/v) for 1 h. However, careful optimization of the washing step was necessary, because nearly total adsorption was observed after loading step on both imprinted and nonimprinted sorbents. Amongst other tested systems, acetonitrileammonium acetate buffer (80:20 v/v) was the most effective in the removal of nonspecific propranolol from nonimprinted sorbent, but not disrupting specific interactions in imprinted one. The total recovery of the extraction process was equal to 64% on imprinted and only 16% on nonimprinted nanofibers. Finally, the analytes were detected with high performance liquid chromatography, coupled with a tandem mass spectrometry detector. The new analytical method was characterized for analysis of propranolol in spiked tap water samples. A large volume of sample, up to 100 mL, was extracted with a mass of 5 mg of imprinted nanofibers providing 21 ng of propranolol, revealing its capability to trace environmental analysis. In another example, MINs were used in dispersive solid phase extraction of fluoroquinolone antibiotics from sea water (Tan et al., 2013). The imprinted nanosorbent consisted of mesoporous carbon nanoparticles coated with an imprinted layer (nanoMCN@MIP). Mesoporous carbon nanoparticles were obtained from mesoporous silica particles by functionalization with (3-aminopropyl)triethoxysilane, dispersion in glucose solution, hydrothermal treatment, and carbonization. The surface of nanoMCN was then functionalized with vinyl groups, followed by copolymerization of methacrylic acid (functional monomer), trimethylolpropane trimethacrylate (cross-linker), and ofloxacin (template) yielding nanoMCN@MIP. The geometrical mean diameter of nanoMCN was B200 nm, with the imprinted thin layer of B1015 nm. Those nanoparticles were used in dispersive solid phase extraction of six antibiotics, viz. ofloxacin, gatofloxacin, balofloxacin, enrofloxacin, norfloxacin, and sarafloxacin from spiked sea water, in the presence of oxolinic acid acting as interfering compound. The extraction protocol consisted of: loading with 3 mL of water solution for 90 min and elution with 1 mL of methanolacetic acid (90:10 v/v), with removal efficiency of around 90% for all tested compounds except norfloxacin and oxolinic acid. The determination of analytes was carried out by high performance liquid chromatography coupled with UV detection but the analytical method was not validated. The reusability of new nanosorbent was analyzed, revealing that, after five consecutive extraction processes and regeneration cycles, it still possessed high extraction capability.

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A very interesting approach was presented by Liu and coworkers (2013). In their investigation, low cross-linked liquid crystal-based imprinted nanosorbent for zopiclone was used in capillary electrochromatography. The main disadvantage of highly cross-linked imprinted sorbents is the low availability of imprinting site embedded inside the polymeric network and significant reduction of binding capacity. In low cross-linked liquid-crystalline imprinted materials, the imprinting sites are preserved by orientation of mesogenic side groups of monomers and polymer backbone. The mesomorphic order enhances the interactions between the template and liquid crystal network, which reinforces the shape of the imprinted cavity. Consequently, low cross-linked liquid crystalline imprinted polymeric networks are able to possess well defined cavities, similarly to those in highly cross-linked systems, simultaneously assuring sufficient access to cavities embedded inside the net. The effectiveness of adsorption can be increased further using nanosized material. Here, four various liquid crystalline backbone monomers were used in the synthesis: 4-(4-methylphenyl)-40 -vinyl-1,10 -bicyclohexyl, 4-(4-vinylcyclohexyl)benzonitrile, 4-(4-methylphenyl)-40 -allyl-1,10 -bicyclohexyl, 4-(3,4-difluorophenyl)-40 allyl-1,10 -bicyclohexyl, together with methacrylic acid (functional monomer) and ethylene glycol dimethacrylate (cross-linker). The obtained particles possessed a diameter of 150180 nm and were monodispersed and uniform. They were used as the sorbents in capillary electrochromatography. The separation parameters were optimized but careful choice of mobile phase was necessary to ensure effective resolution. The highest column efficiency was noted for enantioseparation of racemic zopiclone, with a pH of 3.6. It was summarized that low cross-linked imprinted nanoparticles based on liquid crystalline monomers could be used in effective enantioseparation of zopiclone, with sufficient symmetric peaks in capillary electrochromatograms and with very good reproducibility. In another valuable approach, Huamin and coworkers employed graphene oxide sheets to react with chitosan, since epoxy groups of graphene oxide easily undergo addition with primary amine groups (Huamin et al., 2013). Graphene attracted attention for its extraordinary properties and potential utility in the synthesis of nanomaterials. A graphene oxidechitosan product was added to prepolymerization system that consisted of acrylamide, ethylene glycol dimethacrylate, and sulfamethoxazol (template) to create the imprinted nanostructure composite. The analysis of transmission electron microscopy images revealed that graphene oxide sheets were decorated with imprinted polymer and chitosan films. The imprinting sites were generated in the hybrid material. The following advantages of imprinted composite used as sorbent of sulfamethoxazol were highlighted: high adsorption capacity provided by chitosan, high specific surface area provided by graphene oxide, and high selectivity provided by imprinted polymer. Moreover, employment of chemiluminescence detection of sulfamethoxazol resulted in very low limit of detection (29 nmol/L). The proposed analytical method was verified in the analysis of sulfamethoxazol in drugs samples.


The facile synthesis of magnetic MINs for separation of perphenazine was proposed by Safdarian and coworkers (2016). Here, methacrylic acidcoated magnetite (Fe3O4@MAA) was obtained before the precipitation polymerization of ethylene glycol dimethacrylate, in the presence of the template molecule perphenazine (Fe3O4@MAA@MIP). The obtained nanoparticles possessed a diameter between 120 and 160 nm. They were used as nanosorbents in magnetic solid phase extraction with following protocol: conditioning with 2 mL of water, loading with 5 mL of water standard solution at pH 4.5 for 1 min, washing with 2 mL of water, elution with 0.2 mL of methanolacetic acid (90:10 v/v) for 1 min. The external magnetic field was applied after each step to facilitate the separation. The analysis of perphenazine was carried out by high performance liquid chromatography, coupled with diode array detector. The new analytical method was validated for analysis of perphenazine in spiked human urine and plasma samples with the following digits of merit: limit of detection and quantification equal to 5 and 18 μg/L, respectively (linear range between 20 and 5000 μg/L). Interday precision was between 4.3% and 5.7%, and total recoveries were between 92% and 102% from plasma as well as 95% and 103% from urine samples. In conclusion, the simplicity, considerable reduction of time, and economic aspects were emphasized as main advantages of this method. The molecularly imprinted nanocomposite sorbent was used in the dispersive magnetic microsolid phase extraction of sexual pheromones of olive fruit fly (Bactrocera oleae Gmelin) from various olive oil samples as its contamination (Alcudia-Leon et al., 2016). The detection of 1,7-dioxaspiro-[5,5]-undecane, the olive fruit fly pheromone, is very important, since it affects the quality of olive oils, mainly due to significant reduction of antioxidant properties. The magnetic imprinted nanoparticles (Fe3O4@SiO2@MIP) were prepared in multistep synthesis. Firstly, the coprecipitation and surface pregrafting allowed for the obtaining of aminofunctionalized nanoparticles with magnetic core (Fe3O4@SiO2@NH2). Secondly, the polymerization of cross-linker, ethylene glycol dimethacrylate, in the presence of template molecule of 1,7-dioxaspiro-[5,5]-undecane was carried out to obtain imprinted material (Fe3O4@SiO2@MIP), used as the nanosorbent in solid phase extraction. The spherical particles were an average diameter of 20 nm. The separation of 1,7-dioxaspiro-[5,5]-undecane, the pheromone compound, from oil was carried out for 40 min, followed by washing (1 mL of hexane) and eluting (0.5 mL of methanol). The optimization of extraction procedure was necessary, because the adsorption from aqueous solution was nonspecific. In contrast, the adsorption from hexane or acetonitrile resulted in high specificity (imprinting factor between 4 and 5). The compound was analyzed by gas chromatography, coupled with mass spectrometry. The analytical method was validated on spiked refined sunflower oil sample. The limit of detection was equal to 3.2 μg/L (in calibration range 101000 μg/L) and the total recoveries were between 95% and 99% (for comparison, the total recoveries for nonimprinted counterpart were between 13% and 23%). The reusability tests revealed that, even after nine consecutive extraction and elution cycles, the efficiency of magnetic

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nanosorbent for extraction of 1,7-dioxaspiro-[5,5]-undecane was practically unchanged. The real sample analysis showed the presence of 1,7-dioxaspiro-[5,5]undecane in two tested samples: lampante virgin olive oil (2335 μg/L) and virgin olive oil (102 μg/L). In conclusion, the effectiveness and simplicity of the new method was emphasized, allowing for separation of compounds of interest from complex natural samples such as olive oil. In order to improve the stability and biocompatibility of magnetic imprinted nanoparticles, Zhang and coworkers developed chitosan-Fe3O4 support (Zhang et al., 2013b). The support also acted as the functional monomer in copolymerization with methacrylic acid and ethylene glycol dimethacrylate. The composite magnetic molecularly imprinted polymer was dedicated to the separation of carbamazepine. The implementation of chitosan was responsible for excellent binding capacity and high specificity of sorbent, probably due to strength of the intermolecular interactions between functional groups of support and analyte, making the material promising candidate for pharmaceutical analysis. A more sophisticated method was proposed by Xiao and coworkers (2013). The aim of the innovation was to improve the surface-to-volume ratio of prepared material. In this example, the employment of magnetic carbon nanotubes combined with molecularly imprinted polymer (MCNTs@MIP) was presented. The composite material was fabricated by carboxylation of carbon nanotubes, followed by treatment with FeCl3, ethylene glycol, diethylene glycol, and sodium acrylate in stainless-steel autoclave at 200 C to obtain magnetic carbon nanotubes (MCNTs). Next, the imprinting process was carried out in dimethyl sulfoxide with prepolymerization system of methacrylic acid, ethylene glycol dimethacrylate, and gatifloxacin acting as the template. The incorporation proceeded in the presence of MCNTs and polyvinylpirrolidone. The composite imprinted material was used in the magnetic solid phase extraction of gatifloxacin from human serum. The results revealed that MCNTs@MIP displayed rapid dynamic adsorption with high binding capacity. The method of extraction coupled with high performance liquid chromatography was validated. High recoveries were obtained (79%85%) and the limit of detection was equal to 6 μg/L. The reusability test revealed that, even after four consecutive extraction cycles and regeneration processes, fifth recovery of gatiflaxocin from serum sample was still excellent. In concluding remarks, it was emphasized that MCNTs@MIP could be promising candidates for fast and selective extraction of therapeutic agents from biological fluids. Finally, an interesting example for preparation of a novel nanobiosorbent was presented by Basak and coworkers (2014). In this paper, the dead biomass of Candida rugosa in the presence of zinc nitrate and chitosan was cross-linked by epichlorhydrin to produce zinc imprinted nanobiosorbent. After ion removal, this material could be used as a natural yeast nanobiosorbent for zinc removal from wastewaters or sewer systems and also as a antimicrobial agent. Table 9.1 summarizes selected papers devoted to the analytical methods, based on imprinted nanosorbents.

Table 9.1 Summary of Selected Papers Devoted to the Separation Techniques and Analytical Methods Based on Imprinted Nanosorbents Separation Technique

Instrumental Method

Nanomaterials

Analyte

LOD

References

MISPE

UHPLC-FL

MIP NPs

0.332.6 μg/L

Chauhan et al. (2015)

GC-FID

MIP NPs

Polycyclic aromatic hydrocarbon metabolites Amitriptyline

0.71.2 μg/mL

HPLC-DAD HPLC-UV

MIP NPs MWCNTs/MIP

Khanahmadzadeh and Tarigh (2014) Gao et al. (2016) Gao et al. (2011)

HPLC-UV

MIP NPs

HPLC-UV HPLC-UV HPLC-UV HPLC-DAD HPLC-MS/MS

MIP NPs SiO2/MIP NPs CaSiO3/MIP NPs MIP NPs MIP/PET nanofibers nanoMCN/MIP

d-SPE

HPLC-UV CEC-UV Magnetic d-μSPE

GC-MS

liquid crystal /MIP NPs Fe3O4/SiO2/MIP

UHPLC-MS/MS HPLC-UV HPLC-DAD

Fe3O4/SiO2/MIP Fe3O4/SiO2/MIP Fe3O4/MAA/MIP

Luteolin Estrogenic compounds Salbutamol Efavirenz Sulfonamides Salicylic acid Haloperidol Propranolol

10.216.1 μg/L 1.41.7 μg/mL 10.617.3 μg/L 2.8114.6 μg/L 0.20.35 μg/L 24 μg/L

Fluoroquinolone antibiotics Zopiclone Sexual pheromone of fly (1,7-dioxaspiro [5,5] -undecane) Aflatoxins Carvedilol Perphenazine

Alizadeh and Shamkhali (2016) Pourfarzib et al. (2015) Gao et al. (2010) Meng et al. (2014) Ebrahimzadeh et al. (2013b) Yoshimatsu et al. (2008) Tan et al. (2013) Liu et al. (2013)

10 μg/L

Alcudia-Leon et al. (2016)

0.050.1 μg/kg 0.13 μg/L 5.3 μg/L

Tan et al. (2016) Azodi-Deilami et al. (2014) Safdarian et al. (2016)

LOD, Limit of detection; MISPE, Molecularly imprinted solid phase extraction; d-SPE, Dispersive solid phase extraction; μ-SPE, Microsolid phase extraction; UHPLC-FL, Ultrahigh performance liquid chromatography coupled with fluorescent detector; GC-FID, Gas chromatography-flame detection; HPLC-DAD, High performance liquid chromatography-diode array detection; HPLC-UV, High performance liquid chromatography-ultraviolet detection; HPLC-MS/MS, High performance liquid chromatography coupled with tandem mass spectrometry; CEC-UV, Capillary electrophoresis-ultraviolet detection; GC-MS, Gas chromatography-mass spectrometry.

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In conclusion, it should be emphasized that novel imprinted nanosorbents can be effectively used in various separation techniques, mainly in dispersive magnetic solid phase extraction with high selectivity and satisfactory recoveries of target analytes from various complex biological samples. Moreover, the proposed new analytical methods that include a separation step on imprinted nanosorbents are characterized by low limits of detection, good linearity, and suitable precision.

9.6 IMPRINTED NANOCOMPOSITES FOR BIOSENSORS According to the definition from International Union of Pure and Applied Chemistry, a chemical sensor is a device that transforms chemical information, ranging from the concentration of a specific sample component to total composition analysis, into an analytically useful signal (Hulanicki et al., 1991). MINs are attractive as recognition elements, together with different transducers used during construction of chemical sensors. Apart from good selectivity and specificity, MINs’ attraction in the sensing area is connected with properties such as stability, mechanical and thermal resistance, ease of dispersion, uniform spherical geometry and facility, and low cost of manufacturing. MINs could be used for detection of wide range of analytes, starting from small molecules (drugs, sugars, peptides), through to big organic compounds (proteins) finishing on whole cells or viruses. Moreover, MINs could be applied when biological receptors are not available (Wackerlig and Schirhagl, 2016). Several examples of MIN application in detection of different analytes were found in the pharmaceutical area. Despite MINs many advantages, some limitations should be noted: lack of reproducibility, and commercial exploitation of molecular imprinting is still at the beginning (Poma et al., 2010). In the chemical sensor area associated with MINs, the binding event is usually converted into a detectable signal. Different types of transducers were connected with imprinted materials for developing chemical sensors. The transducer type used in sensor depends on the properties of analytes and MINs form. The main transducer techniques combined with MINs are connected with optical, electrochemical, and mass sensitive detectors. An optical transduction is relatively unsophisticated and does not need many requirements for imprinted material. Thereupon, optical techniques are commonly used in the fabrication of biosensors. The variety of optical detectors enables the appropriate choice of transducer conformed to MINs form and analyte properties. Fluorescence, absorption, reflectometric interference, Raman, and surface plasmon resonance (SPR) spectroscopies are applied as signal transduction techniques in chemical sensors. Fluorescence is often chosen in cases of MINs optical sensing, because of low detection limits, attraction for different analytes in challenging matrices, high

9.6 Imprinted Nanocomposites for Biosensors

signal outcome, and manageable measurements (Guan et al., 2008; Wackerlig and Lieberzeit, 2015). However fluorescence is observed only for a few molecules. Given that the number of fluorescent compounds is small, luminescent MINs are produced. The way of obtaining the fluorescent MINs is as follows: (1) using fluorescent monomers; (2) copolymerization of luminescent material with functional monomers, cross linkers and template; (3) using encapsulation method where presynthetized polymers together with the template and luminescent nanoparticles are mixed (Ma et al., 2015). Fluorescent properties of the template were utilized during fluorescence polarization measurements of enrofloxacin in milk and tap water samples (Ton et al., 2012). Enrofloxacin-imprinted nanoparticles were used during analysis because the template interaction with MINs induced an increase of polarization. Firstly, samples were diluted with acetonitrile to precipitate proteins, then the definite amounts of MINs were added, and fluorescence polarization measurements were conducted. A low limit of detection of 0.1 nmol/L was achieved at MINs concentration of 5 μg/mL. Affinity studies towards different antibiotics confirmed MINs selectivity for fluoroquinolones. A fluorescent monomer (2-hydroxyethyl anthrancene-9-carboxylate) was used during one-pot synthesis of MINs for tetracycline quantification in real biological samples (Niu et al., 2015). Obtained nanoparticles had surface-grafted hydrophilic polymer brushes that made them compatible with biological samples. Fluorescence quenching, with maximum reached around 25 min was observed in undiluted bovine serumcontained tetracycline during analysis of MINs optosensing properties. Selectivity studies measuring MINs binding capacities toward template and its analogue chlorotetracycline in mixed solution were conducted and showed obvious selectivity toward tetracycline in different solvents (organic, pure water, bovine serum). The usefulness of tetracycline imprinted nanoparticles as optical sensor for template determination in undiluted serums, with a detection limit of 0.26 μmol/L and recoveries of 98%102% (even in the presence of interfering drugs) was proved. QDs are the most common luminescent nanomaterial used during copolymerization with monomers for obtaining MINs. As an example, the nanosensor for glycoproteins, based on molecularly imprinted 3D structure and boronate affinity, was described (Zhang et al., 2014b). During the sensor synthesis process, cadmium telluride (CdTe) QDs were coated with polymerizable surfactant (octadecyl-p-vinylbenzyldimethylammonium chloride, OVDAC). Next, the copolymerization of OVDAC-coated QDs was performed with monomers N-isopropylacrylamide and 4-vinylphenylboronic acid in the presence of the template glycoprotein (horseradish peroxidase or ovoalbumin). Molecularly imprinted polymer spatial structure and boronic acidbased recognition elements led to higher specific binding capacity. N-isopropylacrylamide was responsible for the temperature-sensitivity of MINs, and pH changes regulated template interactions with 4-vinylphenylboronic acid. Template-nanosensor interactions were observed

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as fluorescence quenching, increasing with higher template concentration. The nanosensor obtained was selective, sensitive (with limit of detection of 0.24 μmol/ L), and gave the possibility of glycoprotein detection in urine. Li an coworkers used QDs, silica nanoparticles, and imprinted technique to form a hybrid structure with the ability of fluorescence detection of bovine hemoglobin in blood and urine samples (Li et al., 2013). In the first step of this work, carboxyl-terminated CdTe QDs were covalently attached to the surface of aminofunctionalized silica nanoparticles. Next, the imprinted thin layer was fabricated. As a result, a selective and sensitive fluorescent sensor was obtained (limit of detection of 9.4 nmol/ L and recoveries of 97.3%104%). Another type of luminescent material used for fabrication of fluorescent MINs is called upconverting nanoparticles (UCNPs). They have many advantages, such as low cytotoxicity, narrow emission bandwith, high photostability, and a large anti-Stokes shift. Beyazit and coworkers obtained trypsin MINs with UCNPs, a visible light emitter (NaYF4:Yb31, Er31) (Beyazit et al., 2014). In the first step, core-shell UCNPs were fabricated. N,N0 -ethylenebisacrylamide, 2-hydroxyethyl methacrylate were used as monomers, N-acryloyl-p-aminobenzamidine was applied as anchoring monomer, eosin Y and trimethylamine played the role of initiating system. Polymerization was initiated by irradiation at 980 nm with a LED light source. After that, the second shell, composed of imprinted polymer, was produced. Authors composed imprinted prepolymerization mixture containing hydrophilic monomers N,N0 -ethylenebisacrylamide and 2-hydroxyethyl methacrylate, template: trypsin, eosin with triethylamine. Recognition properties of MINs were investigated using labeled trypsin and fluorescence measurements. Selectivity of obtained MINs toward trypsin was proved in competitive studies. This method gives the possibility of producing MINs with UCNPs that can be applied in biosensing or optical imaging. Ivanova-Mitseva and coworkers, used cubic fluorescent core in manufacturing of imprinted materials for the first time (Ivanova-Mitseva et al., 2012). They applied polyamidoamine dendrimers with fluorescent label (dansyl chloride) and S-(carboxypropyl)-N,N-diethyldithiocarbamic acid (iniferter) attached to the peripheral amino groups of dendrimers as a cubic core. The prepolymerization mixture consisted of: the template 2,4-diamino-6-methyl-1,3,5-triazine (acetoguanamine), structural analogue of melamine, the functional monomer: methacrylic acid, the cross-linkers trimethylolpropane trimethacrylate and ethylene dimethacrylate, and the core. An increase in fluorescent intensity was observed when the template was binding to the imprinted polymer during MIN property analysis. Almost no change in fluorescence was noted in the presence of the template structural analogues. Good selectivity and affinity with a fast response of MINs was observed. The proposed method was sensitive, with a limit of detection of 3 3 1028 mol/L. Another example of interesting luminescent core was used for enrofloxacin MIN fabrication (Descalzo et al., 2013). The authors applied Ru(phen)321


(phen is a 1,10-phenantroline) complex encapsulated into silica nanoparticles as a core for Fo¨rster resonance energy transfer (FRET) sensor preparation. The sensor was built from the core and the MIP thin shell, localized on the surface of the core. In this work, Ru(phen)321 provided long-lived donor emission, and cyaninelabeled enrofloxacin was an acceptor in the FRET process. Competitive FRET assays were performed using simultaneous addition labeled, the excess of nonlabeled template, or the interferent molecule to the MINs suspension. As a result, red emission quenching from the luminescent core was observed. Developed FRET biomimetic assay was sensitive (limit of detection of 2 μmol/L) for target analytes. Noble metal nanoparticles are applied during construction of fluorescence sensor because they could change fluorescence signals. Gu¨ltekin and coworkers applied gold nanoparticles and gold-silver nanoclusters for the preparation of fluorescent sensors for cholic acid determination in blood and urine (Gu¨ltekin et al., 2012a,b). Methacryloylamidohistidine-Pt(II) was used as a monomer, which made metal-complexing polymeric receptors for selective binding of cholic acid. Metal nanoparticles or nanoclusters were methacryloyl-activated to form molecularly imprinted polymer layer, reminiscent of a self-assembled monolayer. Binding property analysis of the obtained material was evaluated using fluorescence spectroscopy. When cholic acid was bound to polymeric recognition sites, the photoluminescence emission from gold nanoparticles was observed and a decrease of fluorescence intensity was noted. The sensor was selective, rapid as well as sensitive and cheap, with a limit of detection of 0.1 μmol/L. Encapsulation method with QDs was used for the preparation of composite MIP/QD nanoparticles for one-pot urinalysis (Lin et al., 2009). Firstly, the imprinted nanoparticles were formed from poly(ethylene-co-vinyl alcohol) and the template molecules: creatinine, albumin, and lysozyme (biomarkers in urine). To the final MIN microemulsion, QDs were then added, to obtain the composite imprinted nanoparticles/QDs. The mixture of three MINs/QDs nanoparticles improved sensitivity of all target molecules, except albumin, in their reference concentration in urine. This kind of material could help in early diagnosis of kidney diseases and inflammation. Another optical transduction technique used with imprinted nanoparticles is reflectometric interference spectroscopy (RIfS). This method is based on the multiple reflexion of white light at thin interfaces. Delmitz and Uslu constructed RIfS nanosensor for sarcosine (N-methylglycine), a potential biomarker of prostate cancer (Diltemiz and Uslu, 2015). Here, sarcosine imprinted nanoparticles were formed by emulsion polymerization, followed by the attachment of MINs to transducer slides, to form an RIfS sensor. The obtained nanosensor showed quick responses (20 min), low limit of detection of 45 nmol/L, good selectivity. and provided the possibility of sarcosine detection in urine. Resonance light scattering (RLS) was also used as a transduction method in connection with MINs. Ahmadi and coworkers prepared mefenamic

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acidimprinted polymer layer-coated magnetide nanoparticles, used as RLS nanosensor (Ahmadi et al., 2014). Fe3O4 nanoparticles were firstly coated by silica layer and next by molecularly imprinted polymer layer. As well as selectivity studies of the obtained sensor, various factors which are important during RLS analysis were optimized using a “one-at-a-time” method. The limit of detection of proposed method was 50 μg/L. Many examples of combining MINs with SPR method are noted. SPR sensors measure the change of refractive index by using the surface coated with a thin layer of a noble metal, which is excited by p-polarized light. Different types of molecularly imprinted polymer nanostructures are used for creation of MIP-SPR nanosensors. Erythromycin-imprinted nanoparticles were deposited on the gold surface of SPR chips to fabricate a nanosensor (Sari et al., 2016). The developed sensor showed high selectivity, with limit of detection of 0.4 μmol/L. Myoglobinimprinted nanofilm was synthetized on the gold surface of an SPR sensor (Osman et al., 2013). The nanosensor enabled high selectivity and sensitivity, with limit of detection of 26.3 μg/L, label-free detection, real-time monitoring, and could be used in myocardial infarction diagnosis. Another interesting example of SPR sensors was constructed from molecularly imprinted polymer, gold nanoparticles (GNPs), and reduced graphene oxide (rGO) for rapid detection of ractopamine (Yao et al., 2016). Firstly, the MINs were obtained, then they were mixed with GNPs and rGO, and deposited on an SPR sensor surface. The fabricated sensor showed high selectivity and sensitivity, with limit of detection of 33 μg/L, short response time, and good stability. Surface-enhanced Raman scattering (SERS) sensors are the next possible detectors where MINs could be employed as a part of the sensing system. SERS is a technique that enhances Raman scattering by molecules being adsorbed on metal surfaces. Kamara and coworkers developed molecularly imprinted polymer-based SERS sensor for label-free detection of propranolol (Kamara et al., 2016). During the creation process, core-shell propranolol-imprinted nanoparticles were obtained. In the first step, the imprinted polymer core was fabricated and then the shell was produced for aminofunctionalization of imprinted polymer. After that, MINs were attached to the carboxy-functionalized gold substrate by carbodiimide coupling. Next, amino groups on the MINs were used to connect gold nanoparticles and produce a Raman active surface. Another interesting example of molecularly imprinted nanosenor for SERS detection was dedicated for transferrin (the biomarker that controls the level of free iron in biological fluids) determination in human serum (Lv et al., 2016). The molecularly imprinted nanosensor was prepared by self-polymerization of dopamine on the surface of gold nanorods, with a template protein attached. Competitive binding studies with interfering proteins improved the selectivity of the sensor. The limit of detection of the obtained nanosensor was 1028 mol/L. MINs were employed in the construction of an optical waveguide spectroscopy (OWS) sensor for determination of t-Boc-phenylalanine anilide (Sharma et al., 2014). The OWS method allows measurement of the refractive index changes that are


associated with molecular binding events. MINs were embedded in a poly(N-isopropylacrylamide) hydrogel structure, then the obtained composite was attached to the surface of the SPR sensor. The composite could swell and shrink dependent on template binding by imprinted polymer. The change of hydrogel volume imprinted material caused the refractive index changes and could be observed by OWS. The method allowed for template detection with limit of detection of 2 3 1026 mol/L. Electrochemical sensors are attractive in quantification analysis because of their flexibility in functional design, high sensitivity, low cost, and ease of operation. However, it should be remembered that electrochemical read out methods are useful for charged molecules (Liu et al., 2014c). Different electrochemical techniques are used in connection with MINs in pharmaceutical analysis. Examples are voltammetry, potentiometry, or impedance spectroscopy. Many examples of electrochemical sensors where a molecularly imprinted polymer plays a role of recognition element are produced by the incorporation of MINs into paste electrodes or membranes. Alizadeh and Allahyari constructed a voltammetric sensor for the determination of N-nitrosopropranolol (carcinogenic metabolite of propranolol) in biological samples using MINs (Alizadeh and Allahyari, 2013). During MIN synthesis, they used propranolol as the dummy template. Electrochemical impedance spectroscopy was applied to characterize the obtained sensor. During property analysis, authors improved selectivity toward target analyte, even in the presence of the propranolol, and obtained sensitivity with limit of detection of 0.08 μmol/L. In the second example, a potentiometric sensor for histamine was fabricated with MINs incorporated within a membrane, which was used in ion-selective electrode formation (Basozabal et al., 2014). The resulting senor was used in fish and wine analysis. Another method of obtaining an electrochemical sensor was proposed by Alizadeh and Akbari (Alizadeh and Akbari, 2013). They coated graphite electrode surface with nanosized urea imprinted polymer. High sensitivity with a limit of detection of 3.7 pmol/L was achieved, and the possibility of real sample (blood, urine) analysis was improved. Another type of molecularly imprinted nanostructures used in electrochemical sensor manufacturing are core-shell nanoparticles with different cores and imprinted polymer shells. As an example, there is a magnetic core made from Fe3O4@SiO2 particles, used for electrochemical sensor preparation (Jiang et al., 2016). The sensor was used for the measurement of gram-negative bacterial quorum signaling molecules (N-acylhomoserine-lactones), with limit of detection of 8 3 10210 mol/L. In the next case, the composite of imprinted polymer and mesoporous carbon nanoparticles (MCNs) was applied for preparation electrode sensing material for selective detection of ofloxacin, with limit of detection of 80 nmol/L (Tan et al., 2014). Yu and coworkers constructed selective and sensitive (limit of detection of 2 3 1028 mol/L) nanocomposite material with gold nanoparticles, SiO2 core and molecularly imprinted polymer shell for electrochemical determination of dopamine in blood and urine samples (Yu et al., 2012).

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Mass sensitive devices are the third most popular transducers used in connection with MINs. The most popular piezoelectric device is quartz crystal microbalance (QCM). QCM is a sensor which shows the changes of mass on quartz crystal surface through measuring its difference of frequency in real time (Gupta et al., 2014). Mass sensitive devices are marked with good detection limits, low cost of application, and ease of assembling (Irshad et al., 2013). The typical scheme of construction of a molecularly imprinted QCM sensor is based on deposition of thin prepolymerization mixture film on chip surface and then the polymerization process. An example of MIN-QCM nanosensor is the selective and sensitive QCM nanosensor for the determination of tobramycin in pharmaceuticals, milk, and egg white samples, with limit of detection of 6 pmol/L (Yola et al., 2014a). In the second method of MIN-QCM sensor preparation, imprinted nanoparticles are firstly synthetized and then attached to the sensor surface. This method was applied to obtain a folic acid QCM sensor with limit of detection in the low ppm range (Hussain et al., 2013). The simple and straightforward method for the fabrication of microplate wells coated with imprinted nanoparticles was proposed by Chianella and coworkers (2013). The microplate wells were used in an assay for enzyme-like immunosorbent for vancomycin. The preparation consisted of immobilization of vancomycin on the surface of glass beads, followed by the solid-phase polymerization in the presence of acrylic acid, N-isopropylacrylamide, N,N0 -methylenebisacrylamide, and N-tert-butylacrylamide. After the polymerization was completed, the material was treated with water in elevated temperature. The process allowed for the disruption of the interactions between vancomycin attached to solid support prior to formation of high affinity nanoparticles. Next, the imprinted nanoparticles were dispensed into the microplate wells. The resulting material was characterized by the very low limit of detection equal to 2.5 pmol/L, which was 25 times lower than the limits of detection of reference immunoassay. The assay could be used to determine vancomycin in plasma at clinically relevant levels, with very low cross-reactivity. The performance of imprinted nanoparticles was comparable with high quality monoclonal antibodies. Moreover, the microplate well test was resistant to exposure of elevated temperature, suggesting that the assay did not require refrigeration during transport and storage. Finally, another example of a sensor where MINs were used is a chemiresistor sensor for selective and sensitive determination of ethanol vapor (Alizadeh and Rezaloo, 2013). An ethanol sensing element was prepared by mixing of multiwalled carbon nanotubes, MINs, and poly(methyl methacrylate). Table 9.2 summarized selected papers devoted to the biosensors based on imprinted nanomaterials. Summing up, many examples of MINs applications in the sensing area, connected with pharmaceutical analysis, give an outlook on the important role of imprinted material in modern analytical trends. MINs could possibly be used as biosensing or bioimaging materials in widely understood pharmaceutical analysis, starting from drug analysis through to biological sample analysis and in vivo diagnostics.

Table 9.2 Summary of Selected Papers Devoted to Biosensors Based on Imprinted Nanomaterials Sensor

Transducer

Nanomaterial

Analyte

LOD

References

Optical

FL

MIP NPs MIP NPs CdTe QDs/MIP CdSe/ZnS QDs/ MIP CdSe QDs/MIP CdTe QDs/MIP CdTe QDs/MIP

Enrofloxacin Tetracycline Horseradish peroxidase (glycoprotein) Tocopherol

5 μg/mL 0.26 μmol/L 0.24 μmol/L 5.80 3 1028 mol/L

Ton et al. (2012) Niu et al. (2015) Zhang et al. (2014b) Liu et al. (2014a)

Ractopamine Cysteine Clenbuterol Melamine Bovine serum albumin DNA Caffeine Uric acid Cysteine Estriol Zearalenone (mycotoxin)

7.57 3 10210 mol/L 0.85 μmol/L 120 μg/L 75 μg/L 1.1 3 1027 mol/L

Liu et al. (2014b) Chao et al. (2014) Huy et al. (2014)

0.002 μmol/L

Fang et al. (2014)

Amylase (salivary proteins) Lipase Lysozyme Creatinine (urinary protein) Albumin Lysozyme Bovine hemoglobin

0.1 μg/mL 0.1 μg/mL 0.013 μg/mL 0.635 mg/mL 0.000898 mg/mL 0.00021 mg/mL 9.4 nmol/L

Lee et al. (2010)

Bacillus cereus spores (dipicolinic acid)

0.1 μmol/L (2.4 3 105 CFU/mL)

Gültekin et al. (2009b)

CdTe QDs/MIP CdS QDs/MIP CdSe/ZnS QDs/ MIP

CdSe/ZnS QDs/ MIP QDs/MIP

QDs/MIP

Silica/CdTe QDs/ MIP Au-Ag NPs/MIP

Yang et al. (2014) Diltemiz et al. (2008) Lin et al. (2004)

Lin et al. (2009)

Li et al. (2013)

(Continued)

Table 9.2 Summary of Selected Papers Devoted to Biosensors Based on Imprinted Nanomaterials Continued Sensor

Transducer

PhR LM RIfS

RLS SPR

Nanomaterial

Analyte

LOD

References

UCNPs/MIP Cubic polyamidiamine dendrimers/MIP Au-Ag NPs/MIP Au NPs/MIP Au-Ag NPs/MIP Au NPs/MIP

Trypsin Acetoguanamine

3 3 1028 mol/L

Beyazit et al. (2014) Ivanova-Mitseva et al. (2012)

Biotin Cholic acid Cholic acid Bacillus cereus spores (dipicolinic acid) Domoic acid (shellfish toxin)

15 nmol/L 0.1 μmol/L 0.05 μmol/L 0.1 μmol/L (3.2 3 105 CFU/mL) 67 nmol/L

Özcan et al. (2012) Gültekin et al. (2012a) Gültekin et al. (2012b) Gültekin et al. (2009a)

Enrofloxacin

2 μmol/L

Descalzo et al. (2013)

Sarcosine

45 nmol/L

t-Boc-phenylalanine anilide Mefenamic acid

60 μmol/L 50 ng/L

Diltemiz and Uslu (2015) Kolarov et al. (2012) Ahmadi et al (2014)

E. coli endotoxin Erythromycin Diclofenac Chloramphenicol Lysozyme Hepcidin-25 Glutamic acid Myoglobin

0.44 μg/L 0.29 ppm 18 μg/L 40 ng/kg 84 pmol/L 5 pmol/L 2 nmol/L 87.6 μg/L

Altintas et al. (2016) Sari et al. (2016) Altintas et al. (2015) Kara et al. (2013) Sener et al. (2011) Cenci et al. (2015) Riskin et al. (2010) Osman et al. (2013)

Mn-doped ZnS QDs/MIP Ru(phen)321/ SiO2/MIP MIP NPs MIP NPs Fe3O4/SiO2/MIP NPs MIP NPs MIP NPs MIP NPs MIP NPs MIP NPs MIP NPs Au/MIP NPs MIP nanofilm

Dan and Wang (2013)

SERS

Electrochemical

OWS DPV

CV

SWV

PM

0.022 μg/L 0.0025 μg/mL (4.3 3 1029 mol/L) 0.017 μg/L 5 μg/L

MIP nanofilm MIP nanofilm

Amoxicillin Amikacin

MIP nanofilm MIP/GNPs/rGO nanofilm Au nanodisk/MIP MIP NPs Au nanorod/MIP composite MIP NPs MIP NPs MIP NPs

Triclosan Ractopamine Amylase Propranolol Transferrin

1028 mol/L

t-Boc-phenylalanine anilide Paracetamol Salbutamol

2 3 1026 mol/L 50 nmol/L 6 3 1028 mol/L

MIP NPs

Fluoxetine

2.8 3 1029 mol/L

Fe3O4/SiO2/MIP NPs

N-acyl-homoserine-lactones (gramnegative bacterial quorum signaling molecules) Dopamine Ofloxacin

8 3 10210 mol/L

Sharma et al. (2014) Luo et al. (2016) Alizadeh, and Fard (2013) Alizadeh and Azizi (2016) Jiang et al. (2016)

2 3 1028 mol/L 80 nmol/L

Yu et al. (2012) Tan et al. (2014)

Theophylline Promethazine Tramadol

2.8 3 10212 mol/L 0.004 μmol/L

Kan et al. (2010) Alizadeh et al. (2012) Afkhami et al. (2013)

Histamine

1.12 3 1026 mol/L

Au/SiO2/MIP NPs Mesoporous carbon NPs/MIP Au NPs/MIP MIP NPs Fe3O4/SiO2/MIP NPs MIN NPs

Yola et al. (2014b) Yola et al. (2014b) Atar et al. (2015) Yao et al. (2016) Kamara et al. (2016) Guerreiro et al. (2016) Lv et al. (2016)

Basozabal et al. (2014) (Continued)

Table 9.2 Summary of Selected Papers Devoted to Biosensors Based on Imprinted Nanomaterials Continued Sensor

Mass-sensitive

Pseudoimmunoassay

Transducer

Nanomaterial

Analyte

LOD

EIS

MIP NPs MIP NPs

Promethazine N-nitrosopropranolol

8 3 1028 mol/L

MIP NPs

Urea

3.7 pmol/L

CR

MIP NPs

Ethanol

0.5 ppm

QCM

MIP NPs MIP NPs MIP NPs MIP NPs MIP nanofilm MIP nanofilm MIP nanofilm MIP nanofilm MIP nanofilm

Cholic acid Propranolol Lysozyme Folic acid Kaempferol Tobramycin Caffeic acid Lovastatin Bupivacaine

0.0065 μmol/L 10 μmol/L 1.2 μg/L

MIP NPs

Vancomycin

2.5 pmol/L

ELISA

6 3 10211 mol/L 5.7 3 10212 mol/L 7.8 nmol/L 0.03 nmol/L 30 μg/L

References Alizadeh (2012) Alizadeh and Allahyari (2013) Alizadeh and Akbari (2013) Alizadeh and Rezaloo (2013) Gültekin et al. (2014a) Reimhult et al. (2008) Sener et al. (2010) Hussain et al. (2013) Gupta et al. (2014) Yola et al. (2014c) Gültekin et al. (2014b) Eren et al. (2015) Suriyanarayanan et al. (2014) Chianella et al. (2013)

LOD, Limit of detection; FL, Fluorescence; PhR, Phosphorescence; LM, Luminescent; RIfS, Reflectometric interference spectroscopy; RLS, Resonance light scattering; SPR, Surface plasmon resonance; SERS, Surface-enhanced Raman scattering; OWS, Optical waveguide spectroscopy; DPV, Differential pulse voltammetry; CV, Cyclic voltammetry; SWV, Square wave voltammetry; PM, Potentiometry; EIS, Electrochemical impedance spectroscopy; CR, Chemiresistor; QCM, Quartz crystal microbalance; ELISA, Enzyme-like immunosorbent assay.

9.8 Conclusions

9.7 BIOLABELING WITH IMPRINTED POLYMERIC NANOSTRUCTURES An advanced tool for cell imaging attracted attention with hope for better localization and quantification of molecular biomarkers in cancer or infection. Here, a novel photopolymerization method was described, to coat QDs with imprinted shell by using the visible light emitted from QDs excited by UV, and to obtain specific fluorescent core-shell particles imprinted towards glucuronic acid or N-acetylneuraminic acid (Panagiotopoulou et al., 2016). These particles were used for the recognition of hyaluronic acid, as well as sialylated glycoproteins and glycolipids, on human keratinocytes (HaCaT cells) as an example of the multiplexed detection of glycosylations in living cells. In conclusion, it was stated that use of QDs coated with an imprinted polymer layer as artificial receptors and imaging agents for glycosylation sites gave the possibility for application in diagnostics, theranostics, and therapeutics. In another example, Wang and coworkers (2016b) described monosaccharide-imprinted fluorescent nanoparticles for the targeting and imaging of cancer cells. Fluorescence imaging of human hepatoma carcinoma cells (HepG-2) over normal hepatic cells (L-02), and mammary cancer cells (MCF-7) over normal mammary epithelial cells (MCF-10A) were presented. It was emphasized that the aforementioned nanomaterial could be promising in many applications, such as, in tissue imaging for pathological investigations or targeted photothermal therapy.

9.8 CONCLUSIONS It is only more than 10 years since papers related to imprinted nanomaterial have been published regularly. However, during this time an enormous progress was observed. The new formats of imprinted materials such as nanobeads, imprinted layers on core-shell magnetic nanoparticles, QDs, carbon nanotubes or nanofibers, nanogels, nanowires, nanofilaments, nanosponges, nanocomposites and imprinted photonic crystals were worked out in various aspects of pharmaceutical science, as drug delivery nanocarries, as nanosorbent for separation purposes, in biosensing and bioimaging in diagnostics. Nanoscale format of imprinted materials significantly contributed to increase of the effectiveness of imprinted process when compare to common bulk material. As a result, more satisfactory characterization of imprinted nanomaterials was noted. It is still a long way to overcome some existing problems such as scale-up synthesis that prevents the commercialization of imprinted materials. However, the forthcoming days for MINs as nanodevices, biosensors and biolabels are very promising and it could be expected that in near future the innovations developed in the laboratory and described in this chapter make their way to market.

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FURTHER READING Shutov, R.V., Guerreiro, A., Moczko, E., de Vargas-Sansalvador, I.P., Chianella, I., Whitcombe, M.J., Piletsky, S.A., 2014. Introducing MINA the molecularly imprinted nanoparticle assay. Small 10, 10861089.

CHAPTER

10

Poly(lactic-co-glycolic acid) (PLGA) matrix implants

Joana A.D. Sequeira1, Ana C. Santos1,2, Joaõ Serra3, Francisco Veiga1 and Antońio J. Ribeiro1,2 1

University of Coimbra, Coimbra, Portugal Institute for Molecular and Cell Biology, Porto, Portugal 3 Tecnimede Group SA, Dois Portos, Portugal

2

CHAPTER OUTLINE 10.1 10.2 10.3 10.4 10.5 10.6 10.7

Introduction .................................................................................................376 Implantable Drug Delivery Systems................................................................376 Ability to Sustain and to Control Drug Delivery ...............................................377 The Issue of Biocompatibility ........................................................................379 Poly(Lactide-co-Glycolic Acid) (PLGA) ...........................................................380 Biodegradability...........................................................................................381 PLGA Matrix Implants ...................................................................................381 10.7.1 PLGA Matrices as Sustained Drug Delivery Systems .................... 382 10.7.2 Manufacturing Techniques ........................................................ 384 10.7.3 Drug Release ........................................................................... 387 10.7.4 Factors Affecting Degradation and Also Drug Release From Degradable PLGA Matrices................................................ 389 10.7.5 Therapeutic Peptides and Proteins Incorporated in PLGA Matrix Implants............................................................... 391 10.8 Successful Case Studies ..............................................................................393 10.8.1 Zoladex ................................................................................... 393 10.8.2 Suprefact Depot ....................................................................... 394 10.8.3 Ozurdex ................................................................................... 395 10.9 Problems to Overcome and Opportunities .......................................................395 10.10 Conclusions .................................................................................................397 References .............................................................................................................398


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CHAPTER 10 Poly(lactic-co-glycolic acid) (PLGA) matrix implants

10.1 INTRODUCTION Matrix implants are solid, rigid structures, usually cylindrical rods measuring millimeters in length, and with a diameter of less than a millimeter made essentially from polymers that allow a wide variety of drugs to be incorporated, usually entrapped, in the polymer matrix. They are dosage forms that allow controlled release of a drug for an extended period of time of months or even years, depending on the type of polymer used in their production (Markland and Yang, 2002). They are usually are placed subcutaneously in the patient, requiring a hypodermic needle or an easy surgery. If the polymer chosen is biodegradable, no additional surgery is required to remove the implant at the end of the therapeutic period (Iyer et al., 2006). For the past four decades, there has been a proliferation of academic and industrial research on implantable delivery systems (Langer and Folkman, 1976; Furra and Hutchinsonb, 1992; Kleiner et al., 2014; Kamaly et al., 2016). Injectable implants from poly(lactic-co-glycolic acid) (PLGA) have been successfully prepared to deliver small drugs and therapeutic peptides, but there has been also substantial research to deliver therapeutic proteins and vaccines over a period of days, weeks, and even months at a constant release rate, depending on the degradation behavior of the PLGA employed. With the topic being so vast in scope, the authors felt the need to address a chapter only on PLGA matrix implants and leave behind other modalities often included when addressing the topic of biodegradable implantable drug delivery systems (IDDS).

10.2 IMPLANTABLE DRUG DELIVERY SYSTEMS IDDS, using polymers as delivery vehicles, started being developed in the 1960s, introducing a new concept in drug delivery (Kleiner et al., 2014; Kamaly et al., 2016). Since then, major efforts have been made to improve their formulation and release characteristics. Using these systems, drugs could be delivered in a sustained, continuous, and predictable way, especially useful to control chronic diseases that require repeated treatments (Fung and Saltzman, 1997). The systems were based on biocompatible polymers with appropriate physical properties; in the beginning they were developed to deliver hormones, among other therapeutics, (Fung and Saltzman, 1997; Kamaly et al., 2016) using nondegradable polymers. The term IDDS refers to both nondegradable and degradable systems. The first controlled-release systems were based on nondegradable polymeric materials such as silicone elastomers (Fung and Saltzman, 1997; Folkman and Long, 1964). This led to the development of reservoir drug delivery systems that released the drug by controlled diffusion, through the polymer wall of the delivery device (Brown et al., 1986). The release was exceptionally controlled; however, after releasing the drug at the end of the therapeutic period, the empty reservoir needed to be extracted from the body by a health professional. Following the reservoir IDDS,

10.3 Ability to Sustain and to Control Drug Delivery

solid matrices of nondegradable polymers appeared and were used for long-term drug release. They differ from the previous ones by being simple in production and in safety. Being homogeneous matrices, they were also potentially safer because, unlike a reservoir device, a mechanical defect in a matrix has a reduced possibility of dose dumping and thereby to produce adverse effects or even druginduced toxicity. Unfortunately, solid matrices cant achieve the idealized zero order constant drug release, like nondegradable reservoir ones (Fung and Saltzman, 1997) and require challenging formulation study. There has been a major stimulus to develop biodegradable IDDS that can degrade in biologically compatible components under physiological conditions (Kleiner et al., 2014; Markland and Yang, 2002). Among the advantages is the unnecessary removal of the implant at the end of the therapeutic period, which improves patient compliance, especially when managing chronic illnesses that require long therapeutic periods and repeated treatments to control the disease (usually by frequent injections). These systems are made of natural or synthetic biodegradable materials, usually polymers that can be degraded in vivo, enzymatically, nonenzymatically, or both, to produce biocompatible, endogenous metabolites that can be eliminated by the normal metabolic pathways of the body without associated toxicity (Makadia and Siegel, 2011). In the case of matrix implants, biodegradable polymers are generally used to form delivery devices by physically entrapping drug molecules into matrices. Because biodegradable polymers dissolve after implantation and drug release, they undergo degradation at the same time as they are releasing the drug, allowing certain aspects of device degradation and erosion to be controlled by careful selection of the appropriate polymer properties (Fung and Saltzman, 1997). The main advantage of this approach over micro and nanoparticulates is related to the ability to be extracted when in the presence of undesirable adverse events, since the matrix implants retain a degree of reversibility which is not available in depot injections (Makadia and Siegel, 2011; Rabin et al., 2008; Siegel et al., 2006).

10.3 ABILITY TO SUSTAIN AND TO CONTROL DRUG DELIVERY By definition, a controlled release system provides slow release of a drug over an extended period of time. This is usually achieved by using polymeric materials where drugs are incorporated. During this period of time, the system is capable of providing control at a constant drug level, also known, as zero-order drug release (Stevenson et al., 2012; Alexis, 2005). A sustained release system is able to prolong the release of a drug over a period of time, but unlike the former, not necessarily at a constant drug level (Stevenson et al., 2012). Nevertheless, both systems provide long-term delivery. The effective half-lives of drugs administered by implants, where they are entrapped in a polymer, is much longer than when compared to the free drug

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administered by the same route. This leads to a potentially reducing of the amount of drug and longer periods of drug release (Fung and Saltzman, 1997; Desai et al., 2010). Keeping constant therapeutic drug blood levels is important, especially for drugs with a low therapeutic window that need maintenance of the effective dose at precise therapeutic levels for the desired duration of the treatment, usually extensively prolonged. This prolonged systemic delivery approach is particularly useful for hormonal anticancer therapies that require therapeutic peptides as drugs. Many of the developed therapeutic agents to combat cancer are proteins that have short half-lives, and cross biological barriers poorly, which makes them almost ineffective when delivered as free drugs in a bolus injection (Fung and Saltzman, 1997). Hormone-responsive cancers, such as carcinomas of the breast, prostate, and endometrium, require frequent injections of hormone agonists, usually synthetic agonistic analogs of luteinizing hormonereleasing hormone (LH-RH), as conventional treatments to inhibit tumor growth by lowering plasma testosterone levels in the body. However, these agonists have short plasma half-lives. Treatment with these therapeutic peptides requires daily injections, necessary to maintain therapeutic concentrations. This can be a burden for patients, due to their inconvenience and due to their concentration fluctuation, which can produce undesirable testosterone release responses in patients, leading sometimes to the withdrawal of therapy (Fung and Saltzman, 1997). Polymer drug delivery systems provide an opportunity to deliver high, localized doses of chemotherapy for a prolonged period, as shown in Fig. 10.1.

FIGURE 10.1 Drug concentration in plasma following subcutaneous injection of a drug-loaded matrix implant (solid lines) or daily subcutaneous injections of free drug (dotted line). When the implant is administered, drug therapeutic concentrations are maintained for over a longer period. In contrast, following an injection of free drug, most of the drug is eliminated in 1 day, and repeated injections are necessary to maintain the desire drug concentration along the therapeutic period of time.

10.4 The Issue of Biocompatibility

One of the advantages of controlled delivery system specifically IDDS is the possibility of protecting drugs that are inherently unstable, such as therapeutic proteins and peptides, once stabilized, allowing their continuous delivery at consistent rates for several months. (Langer, 1998).

10.4 THE ISSUE OF BIOCOMPATIBILITY When an implant is administered, scarring of the tissue usually occurs to a greater or lesser extent even if a needle is used to made the insertion, instead of a surgical procedure. Once inside, it is considered a foreign body by the organism and the immunological system is going to react to it; after implantation a response to injury is initiated (Anderson and Shive, 1997). The foreign body physiological response is initiated, not so much due to the nature of the material, which in the case of PLGA is considered well tolerated and with time is metabolized, but due to the size and necessity of scarring tissue after implantation (Morais et al., 2010) Scarring of the tissue will originate a tissue response, such as inflammation, followed usually by fibrous encapsulation isolating the implant (Oviedo Socarra´s et al., 2014; Anderson et al., 1993). Table 10.1 shows the sequence of events that usually take place. Each event leads to the next. In a sequence, from injury to acute inflammation, to chronic inflammation, granulation tissue formation, foreign body reaction, and fibrous encapsulation (Anderson and Shive, 1997). The body tends to completely isolate foreign implants like a foreign body by forming a sheath-like, relatively avascular, densely fibrous membrane capsule around the implant within a few weeks after implantation that effectively walls off the implant from its environment (A. Rothen-Weinhold et al., 1999a,b; Iyer et al., 2006). It is surrounded by fatty tissue, mainly composed of lipid-laden macrophages, which in a long-acting device, can lead to overall variability of drug release (Anderson et al., 1993). This could impair the function of the matrix implant and, in the end, lose the capacity of control-releasing the concentration of drug necessary to treat or manage the disease. Table 10.1 Sequence of Tissue Response to an Implantable Delivery System Tissue injury: Implantation Acute inflammation: Polymorphonuclear leukocytes Chronic inflammation: Monocytes and lymphocytes Granulation tissue: Fibroblasts and new blood capillaries Foreign body reaction: Macrophages and foreign body giant cells at Tissue: implant interface Fibrosis: Fibrous capsule Adapted with permission from Anderson, J.M., Shive, M.S., 1997. Biodegradation and biocompatibility of PLA and PLGA microspheres. Adv. Drug Delivery Rev. 28(1), 524.

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10.5 POLY(LACTIDE-CO-GLYCOLIC ACID) (PLGA) Aliphatic poly(esters) are the best-characterized and most widely studied biodegradable synthetic materials (Uhrich et al., 1999). Polymers such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA) and their copolymer PLGA, poly(ε-caprolactone), poly(3-Hydroxybutyrates) (Uhrich et al., 1999) and poly(alkyl cyanoacrylates) (Kamaly et al., 2016) are included in this family of polymers. Among them, polyesters based upon lactic and glycolic acids are currently the most used biodegradable synthetic polymer for controlled release applications in research academia and in industrial research. They have been in focus since the 1970s, with regulatory approval first as sutures, then as excipients for drug delivery, and finally as IDDS (Kamaly et al., 2016). They are currently the most investigated with available clinical and toxicological data (Witt et al., 2000). Their chemical structure is represented in Fig. 10.2. PLA can exist in an optically active stereoregular semicrystaline form (LPLA), due to the high regularity of the polymer chain, and in an optically inactive racemic amorphous form (D,L-PLA), due to the irregularities of the polymer chain structure. PGA exists as a high crystalline polymer, due to the lack of methyl side groups that are present in PLA (Jain, 2000). Poly(lactic-co-glycolic acids), which contain L-PLA and PGA, are crystalline copolymers, while those with (D,L-PLA) and PGA are amorphous (Jain, 2000). Commercially available PLGAs are usually

FIGURE 10.2 Chemical structures of: (A) copolymer of lactide (x) and glycolide (y);( B) homopolymer of lactide, and (C) homopolymer of glycolide; “x” and “y” refer to the relative amounts of lactide and glycolide units.

10.7 PLGA Matrix Implants

presented with intrinsic viscosity as a form of characterization, which is directly related to their molecular weights helping when choosing the PLGA of interest (Yeo and Park, 2004; Kapoor et al., 2015). PLGAs are glassy, but have a fairly rigid chain structure, with considerable mechanical strength that allow them to be formulated as drug delivery devices. Their Tg is usually above 37 C, but it decreases with the decrease of lactide content and molecular weight (Jain, 2000). Furthermore, mechanical strength is also a reflection of the molecular weight, copolymer composition in terms of PLA-PGA ratio, crystallinity, and geometric regularity of the polymer chains (Jain, 2000). Lactic acid is more hydrophobic than glycolic acid. PLGA copolymers with high ratio of lactic acid in their composition are expected to be less hydrophilic, absorb less water and degrade slower (Jain, 2000; Kapoor et al., 2015). As one can already notice, PLGAs with a wide range of physicochemical properties, so a wide range of erosion times, are commercially available. It is possible to tailor the release profile of a drug by the choice of the PLGA with the right physical chemical characteristics or blend different PLGAs to achieve the best characteristics of both (Fredenberg et al., 2011; Hines and Kaplan, 2013).

10.6 BIODEGRADABILITY PLGAs are relatively stable when in a dry state but degrade by hydrolysis of the polymer backbone when exposed to moisture or an aqueous environment, such as the in vivo environment (Stevenson et al., 2012; Jain, 2000). They degrade by a process of hydrolytic chain scission, where polymer chains are first broken down to small oligomers and then to monomers (Alexis, 2005; Lao et al., 2011). The scission of the chains occurs through cleavage of the backbone ester linkages and can be seen in Fig. 10.3. The two monomers lactic acid and glycolic acid are endogenous acid metabolites, they are going to enter the Krebs cycle and be cleared from the body as carbon dioxide and water (Houchin and Topp, 2008; Danhier et al., 2012) with only minimal systemic toxicity associated leaving no trace in the end and only increased local acidity due to the degradation can lead to irritation at the site of where the polymer is implanted (Uhrich et al., 1999).

10.7 PLGA MATRIX IMPLANTS A monolithic implant, also known as millirod or injectable monolith (Mitragotri et al., 2014), is a unitary structure, typically having dimensions greater than about 0.5 mm, more preferably from 1 to 30 mm (Booth et al.,

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FIGURE 10.3 PLGA hydrolysis when in moisture or aqueous environment. PLGA is biologically hydrolyzed into lactic and glycolic acid. These hydrolysis products are metabolized via endogenous body mechanisms with very minimal systemic toxicity associated.

2003). PLGA matrix implants can be considered monolith implants, where a drug is homogeneously distributed throughout the PLGA matrix (Markland and Yang, 2002). A blend or composite material is formed between a drug and the polymer, or even with more components, and the drug is entrapped mainly by the polymer chains (Weinberg et al., 2007). The system does not require extraction because of PLGA biodegradability.

10.7.1 PLGA MATRICES AS SUSTAINED DRUG DELIVERY SYSTEMS As previously mentioned, PLGA degrades through hydrolytic cleavage of its poly (ester) backbone (Houchin and Topp, 2008), as seen in Fig. 10.3. Chains of


polymer are cleaved into shorter chains or oligomers (Lao et al., 2011). Then, due to the transformation of the latter in CO2 and water, the mass of the matrix is reduced, this mass loss process being erosion (Lao et al., 2011). PLGA is considered to be a bulk-eroding polymer (Lao et al., 2011; Kapoor et al., 2015). Bulk degradation is a homogenous process in which degradation occurs throughout the polymer matrix and can be described as consisting of four consecutive steps (Lao et al., 2011). First, a polymer absorbs water and undergoes some swelling. The water penetrates, and secondary or tertiary structures, previously stabilized by van der Waals forces and hydrogen bonds, are broken. Second, hydrolysis of the covalent ester bonds in the polymer backbone begins generating more and more carboxylic end groups, which can autocatalyze the hydrolysis, improving the breakdown rate of the polymer backbone. This step marks the beginning of a molecular mass decrease and loss of mechanical strength. Third, massive cleavage of the backbone covalent bonds continues. At some critical value of molecular weight, significant mass loss begins to occur. Loss of physical and mechanical integrity occurs at the same time this process is taking place. Fourth, the polymer loses substantial mass, due to solubilization of oligomers into the surrounding medium. The polymer breaks down to many small fragments, which will be further hydrolyzed into free acids (Lao et al., 2011; Hines and Kaplan, 2013). In spite of degradation by bulk-eroding being a characteristic of the PLGAtype polymers, and being that this process is what allows a drug entrapped in a matrix to be released, drugs are usually released from macroscale PLGA-matrix implants via a combination of three mechanisms: diffusion-controlled release, drug-carrier affinity, and degradation of the matrix material (Kamaly et al., 2016; Kearney and Mooney, 2013). The drug release is diffusion-driven and can be affected by concentration gradients, matrix swelling, and diffusion distance, related to the shape of the implant (Kamaly et al., 2016). That is why the size of the matrix also plays an important role in the release. The presence of pores on the matrix could lead to differences in the release. Other processes may also contribute to the release mechanism, including: water penetration and solubilization after the device is first submerged in an aqueous environment, erosion and diffusion of PLGA polymer fragments, and the rate of diffusion of the releasing drug. Because these processes can occur at the same time, the release mechanism may be complex (Hines and Kaplan, 2013). The release profile for a macroscopic PLGA matrix implant usually has three different phases, due to the heterogeneous degradation that usually takes place. It is often called tri-phasic profile. Phase I is usually described as a burst release and has been attributed to drug particles on the surface of the matrix that are easily solubilized by the exterior aqueous environment on the implantation site. Phase II is often a slow-release phase, usually referred to as lag-phase. During this phase, the drug diffuses slowly, either through the few existing pores, or through the relatively dense polymer. Polymer chains still have enough length and still entrap the drug, while polymer degradation and hydration is already

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taking place. The lag-phase may not necessarily be caused by a dense polymer with low porosity, but by pore closure, polymerdrug interactions, or drugdrug interactions that inhibit the release of the drug. Phase III is usually a period of faster release, often attributed to the onset of erosion. This phase is sometimes called the second burst and usually, most of the drug is released rapidly in this phase. It can also be due to crack formation in the matrix. Most of the time, the initial burst release is not desired and formulation optimization can allow changes to the classic tri-phasic release profile. PLGA degradation is a dynamic process; the properties and behavior change while the degradation is ongoing. As the pH and other characteristics inside the matrix are changing during degradation, the conditions that caused slow release are also changing; a pore formation process can dominate over pore closure. The choice of type of PLGA used will also influence drug rapid release. This rapid release could start in Phase II instead of at Phase III, since the onset of rapid drug release was found to be correlated with swelling, erosion, and deformation of the device, and, accordingly, the easier hydration of the PLGA matrix. Phases may also be superimposed and one cannot draw conclusions and predict the drug release by only analyzing the release profile (Fredenberg et al., 2011). Many mathematical models have been developed to try and describe drug release from PLGA matrix implants (Siepmann and Siepmann, 2012). Mathematical modeling is a useful tool that can be used to identify release mechanisms, characterize the significant transport processes involved, estimate unknown parameters such as the diffusion coefficient when diffusion is involved, reduce experimentation, and provide predictive capabilities (Hines and Kaplan, 2013). These models help to elucidate the governing release mechanisms and provide predictive power on the release behavior of a particular formulation. There are comprehensive reviews and several excellent articles that outline mathematical models developed to predict the release kinetics of drug controlled release from bulk-degrading systems (Lao et al., 2011; Siepmann and Siepmann, 2012).

10.7.2 MANUFACTURING TECHNIQUES PLGAs have characteristics of low-melting thermoplastics that soften and melt when heated above their Tg. They can be shaped as macroscopic (or even microscopic) millirods relatively easily with the aid of pressure and mild to high temperature, allowing the manufacturing of devices of several morphologies (Hines and Kaplan, 2013; Jain, 2000). Different implant manufacturing techniques leads to different PLGA processing conditions. Differences in the molding operation, in generated shearing forces, and thermal treatments in the end may change the final molecular weight, crystallinity, or microporous structures of the polymer, leading to differences in degradation of the final implant (a. Rothen-Weinhold et al., 1999a,b). Those macro or microscopic differences make it essential to individually evaluate,


in vitro and in vivo, each implant produced (a. Rothen-Weinhold et al., 1999a,b). The degradation properties of the polymers, the manufacturing technology, as well as the relative drug-loading of the device determine the in vivo performance of the PLGA matrix implant delivery system (a. Rothen-Weinhold et al., 1999a,b; Rothen-Weinhold et al., 1997). Most of the properties that characterize the final matrix implant are a consequence of the manufacturing technique chosen to produce it. Common manufacturing techniques for the preparation of biodegradable implants are listed below: Solvent casting: Solvent casting is used to fabricate large sized, macroscopic matrix formulations with millimeter size which can be implanted or inserted subcutaneously in the body. When applied to PLGA, this technique depends on the dissolution of the polymer in an appropriate organic volatile solvent such as acetone. The polymer solution is then cast in a mold with the desired size and shape, and the solvent allowed to evaporate, resulting in a composite material (Rabin et al., 2008; Makadia and Siegel, 2011). Usually, this method is chosen to produce films and laminar implants as simple monolithic discs or even multilayer discs (Umeki et al., 2011; Dorta et al., 2002). It is not the ideal technique for industrial scale-up; it requires large amounts of organic solvent to dissolve PLGA and the drug, in order to blend them. The use of solvents can introduce the risk of denaturation of drugs, especially if they are proteins. After the blending step, the solvent need to be removed, which requires a very long time to completely remove all solvents from the resulting material. Lastly, it is not a continuous process. The fact that a technique is not a continuous process leads to possible increases in batch-to-batch variation in the composition of the implants, as well as cost increase of manufacturing (Widmer et al., 1998; Makadia and Siegel, 2011). Compression molding: This is a technique with similarities to solvent casting. After a previous solvent casting step, the mixture material is also compressed in a mold, using both high temperature and pressure, producing an implant with higher density (Makadia and Siegel, 2011). With this technique, the use of heat and solvents is avoided, and compression force is used instead. A mixture of the powders is compressed using punches on a punch press at room temperature, in a similar process to the one used to produce tablets (Santovenã et al., 2009). Implants prepared by this technique are reported to release the drug quickly and in a short time (Negrin et al., 2004; Onishi et al., 2005). To delay the release, using a coating layer is a useful approach, however, here the use of solvents is introduced (Schliecker et al., 2004). Ram extrusion: An easy manufacturing technique that could or not use solvents and offers some advantages, such as requiring only a small amount of raw material and low temperatures for preparation. There are reports for the production of implants by ram extrusion that result in almost no peptide drug

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degradation (a. Rothen-Weinhold et al., 1999a,b). A previously made mixture of the chosen polymer and drug is made requiring or not a solvent, then the blend is introduced into a barrel, where an inserted piston rod will move the blend under appropriate pressure and temperature. Briefly, the blend is forced through a small orifice and, because it has viscoelastic properties, an extrudate is produced. Then the extrudate will pass in a die creating long rods with the desired width. As a final step, the produced rods are cut into implants with the desire length (a. Rothen-Weinhold et al., 1999a,b). This technique requires the use of force, high pressure, and temperatures above the Tg of the chosen PLGA to made it soft enough to be forced through the die. Again, stability of the drug needs to be accessed. Because this technique requires the use of high temperature, a decrease in Mw of the PLGA is often reported (a. RothenWeinhold et al., 1999a,b). There are also reports that relate an increase of impurities formed during this process, when high temperatures and long extrusion times are applied (a. Rothen-Weinhold et al., 1999a,b; RothenWeinhold et al., 1999a,b; Rothen-Weinhold et al., 2000). However, extrusion is a continuous process, able to allow scale-up suitable for industrial production. Hot melt extrusion: Following its introduction in the plastic industry, hot-melt extrusion has now been applied in the pharmaceutical field with success. It is a continuous process, where a mixture of polymer and drug circulate through a die to create implants with fixed diameter, without the use of solvents (Makadia and Siegel, 2011; Wang et al., 2010). The blend is melted above the Tg and forced through a die at that temperature. If necessary, previous premixing can be made to help achieve a homogenous extrudate. It is the most appropriate technique to produce matrix systems, with the drug uniformly distributed throughout the implant. A change in the used temperature could have consequences on the drug-loading and MW of the polymer. Therefore, the extrusion process possess a limitation on the drugs that can be used, based on their melting point, polymorph stability, and chemical interactions with PLGA (Makadia and Siegel, 2011). Surprisingly, negligible loss of biological activity has been reported for some peptides (Ghalanbor et al., 2010; Bhardwaj and Blanchard, 1997). Injection-molding: Injection-molding is a technique also more appropriate for large scale industrial production. This manufacturing technique ensures a good mixing treatment between polymer and drug, and allows the manufacture of implants of various shapes. A special adapted injection molding machine is usually used, with a mold of the desired shape and size for the final desire form. The mixture of the polymer and drug is previously subjected to an appropriate temperature to plasticize the polymeric blend, then to an injection temperature and pressure to mold the blend to the desire form. The cast implants are allowed to cool at a much lower temperature (a. RothenWeinhold et al., 1999a,b). Besides being a continuous process, it can be


automatized and is very reproducible. Again, this technique requires the use of high temperature and a decrease in Mw of the PLGA is often reported. As previously explained, this might lead to some degradation of the active compound (a. Rothen-Weinhold et al., 1999a,b). A partial material sterilization is also possible in some cases, with the disadvantages of the higher temperatures used for preparation being deleterious when formulating peptides and proteins. Radiation sterilization seems to be the method of choice to implant matrices for clinical uses (Rothen-Weinhold et al., 1999a,b; Alexis, 2005). There are reports that compare implants produced with different manufacturing techniques (a. Rothen-Weinhold et al., 1999a,b). According to the production technique, authors reported differences related to appearance, color uniformity, smoothness, and small differences in diameter even if the die of the extruder has the same diameter as the mold of an injection-molder. Extruded implants are shown to acquire a slightly bigger diameter than the diameter of the die (a. Rothen-Weinhold et al., 1999a,b). There are also reported differences in the brittle characteristics, in the density of the matrix, resulting higher for injection-molded implants when compared to implants with the same formulation produced by extrusion, and in the weight average molecular weight (Mw) of the polymer. A decrease in Mw and polydispersity of the polymer has also been reported after melt manufacturing techniques, where relative high temperatures are used such as injection molding and extrusion (a. Rothen-Weinhold et al., 1999a,b; RothenWeinhold et al., 1997); a more pronounced Mw drops when the implant is produced by injection molding. In injection-molding, the polymer is exposed to higher temperatures, to high pressures, and to greater shearing forces. In contrast, during the extrusion process, the matrix is exposed to lower temperatures but also to high pressures. Rothen-Weinhold et al. compared PLA implants made by extrusion and injection-molding and noticed that implants extruded were extremely porous and extensively fragmented, whereas the injection-molded ones showed numerous cracks throughout the matrix, but no porosity. Injection-molded samples appear to be denser than the extruded ones, which may lead to increased water uptake in the extruded implants, which may consequently accelerate the degradation process. They concluded that elevated temperatures and a slow rate of cooling allows polymer chains to move and to realign themselves in a more regular structure, becoming the more crystalline polymer, depending on the cooling rate during solidification after melting (a. Rothen-Weinhold et al., 1999a,b).

10.7.3 DRUG RELEASE 10.7.3.1 Mechanism of Drug Release From the Implant One of the most important steps when developing PLGA formulations is foreshadowing the release process, which could be achieved by understanding the erosion

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process that is going to take place (Alexis, 2005). Several authors state that the release of the drug confined in a PLGA matrix is complex (Fredenberg et al., 2011) and mainly a direct consequence of the erosion process that takes place when exposed to hydrolytic chain scission (Alexis, 2005). PLGA matrix implants are usually administrated subcutaneously or placed directly in specific areas of the body. Following implantation, the drug is expected to slowly be released from the implant to the immediate vicinity tissue fluids via complex release mechanisms that involve diffusion, polymer erosion, or a combination of both (Alexis, 2005). The drug is subsequently transported into the systemic blood circulation via diffusion and/or convective processes (Iyer et al., 2006; Larsen et al., 2009; Shen and Burgess, 2015). The degradation properties of the polymers, the manufacturing technology, as well as the relative drug-loading of the device, determine the in vivo performance of the delivery system, as stated previously.

10.7.3.2 Biodistribution After subcutaneous implantation, the drug is released in the immediate vicinity of the implant, comprising cells and interstitial fluid of the extracellular space (ECS) (Iyer et al., 2006). The fluid, with the consistency of a gel, consists of fibrous collagen, proteoglycans, salt solution at pH 7.4, and proteins derived from plasma in less concentration (Iyer et al., 2006). After being exposed to this aqueous environment, the drug released from the implant migrates away from the polymer/tissue site interface via the ECS. It moves by passive diffusion and fluid convection or facilitated transport to reach cells, blood capillary, and the lymphatic vessels to be subsequently transported into the systemic blood circulation via diffusion and/ or convective processes (Fung and Saltzman, 1997; Iyer et al., 2006; Larsen et al., 2009). If the drug is sufficiently lipophilic, it may penetrate cell membranes rapidly enough to move through a transcellular path (Fung and Saltzman, 1997). The size exclusionlike properties of the ECS significantly reduce diffusion of plasma proteins and other macromolecules. Drug molecules with molecular masses below 2 kDa after leaving the implant can enter blood and lymph capillaries and will be cleared predominantly by the blood vessels (Porter and Charman, 2000; Larsen et al., 2009). The degree of uptake by the lymphatics increase directly with the size of the drug molecule (Porter and Charman, 2000). If the drugs are macromolecules above 16 kDa instead of small peptides and particulates of PLGA from the implant, they will be preferentially removed from the tissue by the lymphatic system (Iyer et al., 2006). The presence of blood and lymphatic capillary is the reason for observed variability in absorption rates following subcutaneous administration of therapeutic proteins at different administration sited in the body (Larsen et al., 2009). As previously mentioned, following implantation, normal subcutaneous wound healing is going to take place, due to the macroscopic size of these types of millirods and the necessity to scar in more or less extent the tissue. This can smooth


the implant surface and chemically inert the implant, due to fibrous encapsulation. This tissue scarring and wound healing process is dependent on the size and shape of the device, as well as the biocompatibility of the PLGA (Larsen et al., 2009). The formation of the capsule could limit drug diffusion and perfusion transport, leading to a less effective implant. Contributing to the reduced effectiveness is also the infiltration of scleroprotein, formed as a result of the tissue reaction to the presence of a foreign body in the pores of the surface of the matrix, also known as ghost formation (Iyer et al., 2006). The fate of drug molecules released and delivered to tissues in the end is a combination of the rates of transport (via diffusion and fluid convection), elimination (by degradation, metabolism, transcapillary permeation and internalization (Fung and Saltzman, 1997), maintenance of implant integrity, and administration location in the body. However, the rate-limiting step in the absorption of a drug released by a PLGA matrix implant will be the controlled release, provided by the device (Larsen et al., 2009).

10.7.4 FACTORS AFFECTING DEGRADATION AND ALSO DRUG RELEASE FROM DEGRADABLE PLGA MATRICES Being an IDDS, a major factor affecting the degradation and, consequently, the drug-release mechanism of a matrix implant is the type of polymer used, among other important factors (Fredenberg et al., 2011). Some polymer properties could be tailored to achieve a specific drug release behaviour (Hines and Kaplan, 2013). Polymer composition: Polymer composition, in terms of PLA-PGA ratio is recognized as the most important factor that influences the degradation rate (Alexis, 2005; Washington et al., 2017). A ratio increase in more hydrophilic glycolic acid in PLGA composition accelerates weight loss, due to specific chain scission of glycolic linkages, leading to preferential degradation of those units instead of lactic acid units (Makadia and Siegel, 2011). The higher the amount of PGA in the copolymer, the higher the hydrophilicity of the PLGA matrix, increasing the rate of degradation of PLGA backbone due to hydrolytic scission and subsequently faster drug release (Makadia and Siegel, 2011). Conversely, a ratio increase in lactic acid residues increases the crystallinity of PLGA, resulting in higher degradation rate of the backbone. Reports suggest that the rate of decrease in Mw, when subject to hydrolysis, was higher for PLGA with high initial crystallinity, meaning a high content of L-PLA not subject to any process that could result in loss of crystallinity like compressing molding or quenching processes (Alexis, 2005). Crystallinity: Crystallinity can be correlated with Tg. It is affected by the percentage of PGA and PLA units (Makadia and Siegel, 2011). If the PLGA has more L-PLA or PGA, it is more crystalline.

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Weight average molecular weight (Mw): PLGAs with higher Mw have bigger polymer chains, therefore require more time to undergo full hydrolysis and degrade (Makadia and Siegel, 2011). The opposite occurs with L-PLA, due to their high crystallinity that increases with Mw (Makadia and Siegel, 2011). PLGA chains with low Mw are known to have higher biodegradation rates and to release drugs faster than PLGA chains with high Mw (Alexis, 2005). This behavior is accompanied by a reduction of the glass transition temperature. The opposite behavior is seen in L-PLA, due to the high degree of crystallinity that is present as the Mw decreases. Low Mw, low PLA-PGA ratio and uncapped end groups result in a less hydrophobic PLGA, with increased rates of water absorption, hydrolysis and erosion (Tracy, 1999; Fredenberg et al., 2011). However, as earlier revealed, those properties change along the degradation and drug release period. During these processes, PLGA has a dynamic nature, as its properties and behavior change with degradation. Hydrophobic, high Mw and slow-degrading PLGA will eventually become more-hydrophilic, low Mw and fast-degrading PLGA (Fredenberg et al., 2011). Morphology of the matrix: Especially for large devices, the higher the ratio of surface area-to-volume, the higher the degradation of the matrix (Grizzi et al., 1995; Kapoor et al., 2015). Bulk degradation will be faster and subsequently faster drug release (Makadia and Siegel, 2011). The presence of pores in the matrix can anticipate drug release if the initial pores had sufficient size to allow drug molecules to be released by diffusion when water fills the pores after implantation. Manufacturing technique: As discussed earlier, the choice of technique to manufacture the implant will have consequences in the final PLGA characteristics, due to the processing variables used and this will have an impact on the degradation (Hines and Kaplan, 2013). Drug type: The type of drug entrapped could change the mechanism of the matrix degradation and affect the rate of degradation (Siegel et al., 2006; Makadia and Siegel, 2011). The choice of the drug, due to its chemical properties, can tremendously affect the release mechanism. The behavior of the release mechanism would vary greatly, depending on the chemical properties of the drug of choice. In general, lower drug loadings, salts, and drugs that aided hydrolysis process led to faster degradation of the matrix. Interactions between the polymer and drug may occur, affecting the rate of release (Hines and Kaplan, 2013). When dispersed in a PLGA matrix, some drugs can also decrease the Tg (such as caffeine and salicylic acid) and others can interact with the carboxyl groups generated during the process of PLGA degradation itself (as reported for quinidine) (Alexis, 2005). When the drug entrapped is a protein or peptide, additives like protein stabilizers which are salts (basic anions and divalent cations) are commonly found in the formulations (Milacic and Schwendeman, 2014). They can neutralize acids but also be pore forming, catalyzing even more the hydrolysis and having the opposite effect (Fredenberg et al., 2011).


pH: Alkaline and strongly acidic media accelerate PLGA degradation. When the degradation of PLGA is taking place, the accumulation of carboxylic end groups is going to autocatalyze the degradation by hydrolysis of the ester linkages (Makadia and Siegel, 2011). Drug load: When entrapping drugs in a PLGA matrix, polymer-to-drug ratio has an influence on the rate and duration of release. Higher polymer-to-drug ratio leads to higher rate and duration of release; however, this will lead to a higher burst release if present in the mechanism of release (Makadia and Siegel, 2011). When the drug is released, the space left will be probably a pore in the matrix. As the medium enters the pore, it will facilitate more drug to be released by diffusion (Fredenberg et al., 2011). Other factors: Other factors affecting the degradation and, as a consequence, the drug-release mechanism of PLGA also include: flow rate, sterilization of the final pharmaceutical form, strain, presence of plasticizers, and the presence of enzymes, although there are contradictory reports linking a correlation between the role of body enzymes and the degradation of PLGA (Alexis, 2005).

10.7.5 THERAPEUTIC PEPTIDES AND PROTEINS INCORPORATED IN PLGA MATRIX IMPLANTS When the drug is a therapeutic peptide or protein, which are extremely unstable molecules, the difficulty to formulate and to deliver them without some loss of purity is usually reported (Frokjaer and Otzen, 2005). Proteins are highly organized, complex and unstable macromolecules, most of them with short halflives (Vaishya et al., 2015). Maintaining their three-dimensional structure with their chemical integrity during and after manufacturing of the implant is a major concern (Fu et al., 2000). Once degraded, they are likely to cause immunogenicity and an impact on its release behaviour and bioactivity. Their complexity and fragile nature imposes certain restrictions on the process parameters that may be employed to formulate these drugs in PLGA matrix implants. Depending on the manufacturing technique chosen, if the drug is a peptide or protein the exposure to the use of organic solvents, elevated temperature shear, and high pressure in processing might jeopardize their function, due to a loss of stability (Ghalanbor et al., 2010). Each protein is unique and so is likely to have a unique set of conditions that would maintain their stability during the manufacturing technique of implant production. Nonetheless, the possibility of denaturation needs to be consider when developing these types of pharmaceutical forms. When producing implants with those type of drugs, peptide purity should to be evaluated right after the implant manufacturing, to assess the feasibility of the technique employed. There are strategies to reduce peptide degradation, which are based on an understanding of the effect of the changes in the manufacturing chosen conditions (Rothen-Weinhold et al., 1999a,b).

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PLGA can also contribute to physical and chemical protein degradation (Zhu et al., 2000). After implant administration, the protein entrapped in PLGA comes in contact with moisture and the ambient temperature of the body for the total degradation period of the implant, which is often months. It is known that water plays a role in protein degradation (Fu et al., 2000). PLGA degradation will produce an acidic micro environment inside the implant and in the surroundings of the implant location. The products of implant degradation are going to contribute to jeopardizing the stability of the protein released. Several reactions between chemical groups of the protein entrapped and the groups of PLGA can take place after the implantation on the body, as soon as the degradation reaction of the matrix implant starts (Fu et al., 2000). Production of acylation products are dependent on the type of protein entrapped. Proteins like calcitonin and parathyroid analog have been reported to undergo this kind of acylation reaction, whereas leuprolide peptides are known to maintain stability in this situation (Brown, 2005). This usually has to do with the peptide primary chemical sequence. It is a common practice to add salts, such as basic anions and divalent cations like Mg(OH)2, to reverse the acidity. These are usually referred to as stabilizers or additives (Fredenberg et al., 2011). Acid labile proteins, such as bovine serum albumin (BSA), fibroblast growth hormone, and bone morphogenic protein-2, can be stabilized by adding base salts that retard acidity. The common degradative pathways are mostly identified; several processing and formulation strategies, namely the use of certain excipients, can prevent or reduce deleterious chemical reactions (Fu et al., 2000; Manning et al., 2010). During implant manufacturing, impurities may form. A thorough study on the formation of impurities in the protein during implant manufacturing is fundamental to try to figure out how the polymer matrix affects the stability of the incorporated protein. Implants of PLA and the peptide vapreotide were extruded under harsh conditions of high temperature and shear, deliberately, in order to identify the impurities formed and access the influence of the PLA matrix on their formation (Rothen-Weinhold et al., 2000). A high percentage of an impurity resulting from the addition of lactide from the PLA on the peptide phe-terminal amino group was found (Rothen-Weinhold et al., 2000). In another study, BSA-PLGA implants were produced by hot-melt extrusion. After being submitted to an in vitro release study, it was noticed that the protein was not completely released from the implant, due to the formation of insoluble covalent adducts that BSA formed with PLGA by thioester linkages (Ghalanbor et al., 2012). Even if the manufacturing technique chosen is less deleterious to their stability, the biggest issue with formulations for proteins controlled-release systems is the instability of proteins themselves (Fu et al., 2000). A rigorous analysis of protein conformation using techniques like circular dichroism, enzyme-linked immunosorbent assays (ELISAs), and chemical state using high-performance liquid chromatography (HPLC) should be performed over the entire life of the implant, starting in the formulation phase, release in vivo, and during stability studies.

10.8 Successful Case Studies

Table 10.2 Commercial Controlled-Released PLGA Matrix Implants Drug

Trade Name

PLGA Type /Technology

Company

References

Goserelin

Zoladex

PLGA 50:50

AstraZeneca

Furr and Hutchinson (1992)

Buserelin

Suprefact Depot

PLGA 75:25

Sanofi-Aventis

products.sanofi. ca (2015)

Dexamethasone

Ozurdex

Novadur technology Blend of PLGA 50:50 uncapped and capped

Allergan, Inc.

EMA.europa.eu (2015)

10.8 SUCCESSFUL CASE STUDIES There are several patents which indicate the potential and amount of work put into this type of formulation, especially for the delivery of peptide and protein drugs (Booth et al., 2003; Deghenghi, 2000). However, only three examples are present in the market currently sold for use in humans (Table 10.2). They include two formulations based on LH-RH agonists, which are peptides with a large therapeutic window, and dexamethasone, a corticosteroid.

10.8.1 ZOLADEX In order to increase the circulation half-life of LH-RH from minutes to hours, several LH-RH analogs (LH-RHa) have been developed. However, there was not enough half-live duration to treat chronic conditions, such as hormone-responsive cancers. Long-acting injections with a duration of several months was required (Stevenson et al., 2012). To fulfill that need, formulations that can last for months, taking advantage of PLGA-based systems, have been present in the market since the late 1980s; Zoladex being the first product to appear in the form of a matrix implant (Brown, 2005; Stevenson et al., 2012). Zoladex, marketed by AstraZeneca, provides a subcutaneous depot formulation of LH-RH analog goserelin. Goserelin, a synthetic peptide blocker of LH-RH receptor, works by reducing testosterone or oestradiol production (Kleiner et al., 2014). It is administered in a relatively large 1.5 mm subcutaneous implant (Brown, 2005) millirod, in which the drug is dispersed in a PLGA matrix, using a preloaded single-use syringe device and packaged in sealed, light, and moistureprotected aluminium foil pouch. After subcutaneous implantation into the skin of

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the patient’s upper stomach, the encapsulated goserelin is released by a combination of a controlled diffusion and erosion mechanisms (Kleiner et al., 2014). The Zoladex product line consists of implants of bioerodible PLGA rods for 1 month or 3 month delivery of 3.6 and 10.8 mg of goserelin, respectively. They can be used for the palliative treatment of prostate and breast cancer, as hormone ablation therapy due to advanced carcinoma (Stevenson et al., 2012; Kleiner et al., 2014), and treating endometriosis, fibroid tumors, and precocious puberty (Brown, 2005). First launched in 1987, Zoladex is now available and widely prescribed in more than 100 countries, with indications in prostate and breast cancer. According to AstraZeneca, it is an important oncology product, with annual sales of approximately $1 billion and one of AstraZeneca’s leading cancer drugs, offering therapy for prostate cancer, with proven survival data, and similar overall survival benefits and improvements in quality of life, compared to surgical castration (orchidectomy) (AstraZeneca.com, 2013).

10.8.2 SUPREFACT DEPOT Another LH-RHa is buserelin acetate. It has a greatly enhanced LH-RH effect and longer duration of action than natural LH-RH. When administered for periods greater than 13 months, it results in clinical inhibition of gonadotropin release, and when used in large pharmacologic doses of 50500 μg SC/day or 3001200 μg IN/day, induces reduction of serum testosterone or estradiol to castration levels. Unfortunately, buserelin is rapidly inactivated by enzymes from liver, kidney, and anterior pituitary. The main buserelin-degrading enzyme is pyroglutamyl-amino-peptidase (PGP) found in the mammalian liver and anterior pituitary, and chymotrypsin-like enzymes, such as neutral endopeptidase from the pituitary (products.sanofi.ca, 2015). Suprefact Depot, also known as Profact Depot and Trigonist, marketed by Sanofi-Aventis, are implants that contain 6.6 mg (buserelin acetate implant equivalent to 6.3 mg of buserelin base) or 9.9 mg of buserelin acetate (buserelin acetate implant equivalent to 9.45 mg of buserelin base) in a matrix of 26.4 or 39.4 mg PLGA 75:25 molar ratio. They are administered every 2 or 3 months as a subcutaneous injection in the abdomen to help diminish the size of cancer in patients with advanced prostate cancer, by reducing the levels of testosterone to the desired level for managing this condition (products.sanofi.ca, 2015). Suprefact Depot comes in a prefilled disposable ready-to-use sterile syringe/ applicator with an integrated safety-engineered needle (internal needle diameter of 1.4 mm). It contains two or three cream-color implants, biodegradable and biocompatible rods that deliver 6.3 or 9.45 mg of buserelin in total per administration (Schliecker et al., 2004). The administration needs to be made by a doctor or nurse (products.sanofi.ca, 2015). However, chronic administration of the implant every 2 months or 3 months by a healthcare professional ensures continuous

10.9 Problems to Overcome and Opportunities

suppression of testosterone secretion with no accumulation of buserelin release after repeated dosing, ensuring therapeutically effective systemic concentrations and palliative treatment for hormone-dependent advanced carcinoma of the prostate gland.

10.8.3 OZURDEX Allergan, Inc. developed Ozurdex, an intravitreal biodegradable implant containing 0.7 mg of dexamethasone for the treatment of macular edema related to branch retinal vein occlusion (BRVO), central retinal vein occlusion (CRVO), and noninfectious uveitis. This implant uses the Novadur technology and has been approved by Food and Drug Administration since 2009 (Kleiner et al., 2014). Novadur is a solid, cilindrical implant sustained-release drug delivery system that used ester-terminated 50:50 and acid-terminated 50:50 PLGA. Ozurdex is used for the controlled delivery of 700 micrograms of dexamethasone. Ozurdex is presented in a specially designed injector with a 22-gauge needle sealed in a foil pouch containing desiccant. It should be administered by an health professional in the vitreous cavity of the eye in a sterile ambient (EMA.europa.eu, 2015; allergan.com, 2014). A 12-month clinical study to evaluate the safety and efficacy of one or two treatments showed that reinjection with Ozurdex implant was safe and well tolerated over the entire period of the study; improvements in BCVA and central retinal thickness were also observed after the second treatment, without serious treatment-related incidence of adverse events (Haller et al., 2011).

10.9 PROBLEMS TO OVERCOME AND OPPORTUNITIES Biocompatible and biodegradable PLGA matrix implants possess the capability of delivering a variety of drugs in a controlled manner over periods of weeks to several months taking advantage of the physical chemical characteristics of the copolymer. They have been known for 40 years and they present a huge market potential, if several important issues are resolved. Among them are the assurance of product performance and safety through in vitro quality control tests design specifically for the product in question that, unfortunately, are lacking (Shen and Burgess, 2015). To obtain a reliable, robust correlation of both the rate of drug release obtained in a laboratory using in vitro assays and the real rate of drug release in the body is still difficult. A faster drug release in vivo, when comparing to in vitro release, suggest differences in drug absorption and drug dissolution. The acidic products that result from PLGA degradation process may accumulate at the local sites, lowering the pH in the interstitial space immediately surrounding the implant, and could

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accelerate polymer degradation with subsequent increased drug release in vivo when compared to in vitro (Shen and Burgess, 2015). Furthermore, chronic inflammation in response to the presence of the implant as a foreign body in the interstitial site could lead to fibrosis isolating the implant, slowing drug absorption or drug release (Anderson et al., 1993). Observed differences are likely to be due to local transport and metabolism of drugs in tissues, as well as to differences in the rate of polymer degradation after implantation (Fung and Saltzman, 1997) as previously explained. An in vitroin vivo correlation (IVIVC) is a predictive mathematical model describing the relationship between an in vitro property and a relevant in vivo response of a drug product, one of the most important issues in the field of drug development. IVIVCs are categorized into five different levels: A, B, C, D, and multiple Level C. Level A IVIVC being the most desired, since the in vitro release method can be used as a surrogate for bioequivalence studies, if preapproval and postapproval changes are required throughout the process of development of the drug (Shen and Burgess, 2015). The successful development and application of a meaningful IVIVC allows for an accurate prediction of the in vivo performance of a developed drug product and could minimize human or animal studies, also reducing the regulatory burden with the final goal of less money spend in research and development (Shen and Burgess, 2015). A meaningful IVIVC can be used to guide formulation and/or process development changes in the various stages of drug product development and can be used to support and/or validate the use of an in vitro dissolution method, helping set clinically relevant dissolution specifications to ensure product quality (Shen and Burgess, 2015). Establishing an IVIVC for nonoral dosage forms, such as parenteral polymeric implants, has remained extremely challenging, due to their complex characteristics and the lack of standardized, compendial in vitro release testing methods, capable of mimicking in vivo drug release conditions (Shen and Burgess, 2015). Besides the lack of suitable in vitro release testing methods, the development of IVIVCs for nonoral drug products is even more complicated, due to their complex multiphasic release. This issue has been thoroughly addressed by Burgess and colleagues since the beginning of 2000s (Burgess et al., 2004; Martinez et al., 2008, 2010; Shen and Burgess, 2015). Regulatory guidance to aid stablishing an IVIVC is only available for extendedrelease oral dosage forms and, to date, this guidance is being adapted to the development of many parenteral polymeric microspheres and implants, transdermal patches/gels, as well as ocular inserts (Shen and Burgess, 2015). During the last two decades, considerable progress has been achieved with “proof-of-concept” research that could only demonstrate the possibility of developing point-to-point linear correlations, or Level B correlations, based on one formulation. Although a Level A IVIVC is the most informative and recommended, other levels of IVIVC, such as multiple Level C and Level B, can be helpful to assure product quality and to assist in formulation development (Schliecker et al.,

10.10 Conclusions

2004; Shen and Burgess, 2015). Designing in vitro release studies that reflect, as much as possible, the in vivo behavior of these products is still a challenge to overcome. Even now, one could see in the available literature several in vitro release methods used, such as sample-and-separate (Schliecker et al., 2004), membrane dialysis, and flow through to determine in vitro drug release characteristics and to develop IVIVCs. Those methods are not able to adequately mimic different in vivo drug release conditions and so there is still a dearth of biorelevant in vitro dissolution methods that are capable of reflecting the complex and dynamic in vivo environment these dosage forms encounter (Shen and Burgess, 2015). In vivo drug release can usually be determined through analyzing drug content in the plasma or blood fluid. Noninvasive techniques, such as micropositron emission tomography (PET/Micro-PET) imaging are reported as a possible useful technique to determine in vivo drug release and facilitate the development of IVIVCs (Hu¨hn et al., 2010) for those kind of formulations. This will allow a reduction in animal testing, due to unnecessary drug content in the plasma or blood fluid analysis. This requires frequent sampling, followed by complicated procedures to extract the drug from biological samples, high sensitive analytical instruments, as well as extensive animal/human experiments. New nondestructive biomedical imaging techniques that can be used to study implants in situ are highly desirable in the evaluation of IDDS (Zhou et al., 2015) and could help achieving relevant data. Recent data suggest the evaluation of formulations that have a dissolution rate of the drug, on a lower, medium and higher scale, in order to obtain the best possible correlation, A, B or C, which decreases the number of extremely expensive in vivo studies (Shen and Burgess, 2015).

10.10 CONCLUSIONS Despite the continuous improvements in the field of controlled drug release with the trend to go nanoscale instead of staying macro, matrix implants are still playmakers in the controlled-delivering arena. Taking into account that nanoscale manufacturing involves high costs, PLGA matrix implants should be consider. The added value of PLGA in terms of how well characterized it is, biocompatible, clinical tested and approved for human use and the use of drugs in solid state important when dealing with proteins, as well as, the straight forward implant manufacturing techniques without complicated apparatus and mostly suited for continuous processes easy to scale up for industrial production are desirable when developing and producing a new pharmaceutical product. More emphasis needs to be placed on the problem of IVIVC for non-oral drug products, such as polymeric implants, and on development of innovative in vitro models that allow accurate prediction of in vivo drug release as well as mathematical methods and simulation techniques. Regulatory approval, high cost of

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preclinical to clinical translation of these systems (in particular for biological drugs) and patient compliance are among the matters still need to be improved. Nevertheless, those drawbacks are the same that arise when developing a formulation based in micro- or nanoparticles so PLGA matrix implants came with advantages as a controlled release form even nowadays.

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CHAPTER

Hydrogels for biomedical applications

11

Luciane R. Feksa1,2, Eduardo A. Troian1, Cristina D. Muller1, Fabian Viegas1, Aline B. Machado1 and Virgıńia C. Rech2,3 1

Feevale University, Novo Hamburgo, RS, Brazil 2Federal University of Rio Grande do Sul, Porto Alegre, RS, Brazil 3Franciscan University, Santa Maria, RS, Brazil

CHAPTER OUTLINE 11.1 Hydrogels: Concepts and Definitions ...............................................................404 11.2 Classification of Hydrogels .............................................................................407 11.2.1 Classification Based on Response ................................................407 11.2.2 Classification Based on Type of Cross-Linking ...............................409 11.2.3 Classification According to the Method of Preparation ...................409 11.2.4 Classification Based on Source ....................................................409 11.2.5 Intelligent Hydrogels...................................................................412 11.3 Applications of Hydrogels...............................................................................414 11.3.1 Tissue Regeneration....................................................................414 11.3.2 Tissue Dressing ..........................................................................417 11.3.3 Contact Lenses...........................................................................419 11.3.4 Drug Delivery System..................................................................420 11.3.5 Hygiene Products .......................................................................421 11.3.6 Others Applications ....................................................................423 11.4 Hydrogel Technical Features ..........................................................................424 11.5 Metabolism and Hydrogels .............................................................................424 11.6 Regulation of Hydrogels .................................................................................425 11.7 Potential Risks of Hydrogels ...........................................................................426 11.8 Nanoparticle Biosafety ...................................................................................427 11.9 Final Remarks................................................................................................428 References .............................................................................................................429 Further Reading ......................................................................................................438


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CHAPTER 11 Hydrogels for biomedical applications

11.1 HYDROGELS: CONCEPTS AND DEFINITIONS Nanotechnology is the next revolution in the history of scientific innovations and advancements; it is an interdisciplinary field which bridges various disciplines, from physics, chemistry to biology, and from engineering to medicine, thereby establishing new trends of technological applications. This upcoming technology is still a virgin area to be explored by researchers, but it has already been proved to be a potential field in physiology and allied bioscience subjects. Nanoparticles, owing to their unique physic-chemical properties, have found their application in various fields, starting from electronics to biomedical science and therapeutics (Gomes et al., 2016). Nanotechnology is considered a key to the 21st century and holds promises in sectors like electronics, chemistry, biotechnology and medicine. In fact, nanoparticles and nanobased products are considered prodigious tools, participating in various areas of biological applications. Nanoparticles are known to exist in diverse shapes, such as spherical, triangular, cubical, pentagonal, rodshaped, shells, ellipsoidal, and so forth. Nanoparticles, by themselves and when used as building blocks to construct complex nanostructures such as nanochains, nanowires, nanofibers, nanoclusters and nanoaggregates, find use in a wide variety of applications in different fields (Gomes et al., 2016). Throughout the years, scientists classified hydrogels in different ways. The main concept is that hydrogels are a water-swollen, cross-linked, polymeric network, produced by reacting one or more monomers (Ahmed, 2015). Hydrogels have received considerable attention in the past 50 years, due to their exceptional promise in a wide range of applications (Buchholz and Graham, 1998; Li et al., 2013; Ahmed, 2015). They also demonstrate a high degree of flexibility that can be compared to biological tissues, due to their large water content (Ahmed, 2015). Nanocomposite hydrogels are biomaterials that have potential to be used in biomedical and pharmaceutical industry, although not restricted to them (Motealleh and Kehr, 2017). In the 1950s, a US company developed the first hydrogel using polyacrylamide as base. In the 1970s, a British company improved its water retention efficiency up to 400 times its own capacity (Wofford Jr and Koski, 1990). The first hydrogel designed for medical use was synthesized by Wichterle and Lim from the copolymer 2-hydroxyethyl methacrylate with ethylene dimethacrylate, initially used as contact lenses (Kopeˇcek, 2002; Motta, 2009). Hydrogels consists in natural or synthetic materials that can absorb several times their own dry weight (Kondiah et al., 2016; Park et al., 2017). Hydrogel composition consists in a three-dimensional (3D) highly hydrated polymeric network, and can hold 20- to 40-fold more water, compared with their dry weight. Developing biomaterials with controlled physical, chemical, electrical, and biological properties will be helpful to form functional tissues (Fisher et al., 2010; Slaughter et al., 2009). Among different biomaterials, hydrogels are the most promising candidates to

11.1 Hydrogels: Concepts and Definitions

FIGURE 11.1 An example of a hydrogel after its preparation.

mimic most biological tissues (Fisher et al., 2010; Kloxin et al., 2010; Slaughter et al., 2009). Due to their unique physical properties, these networks can be shaped or cast into various sizes and shapes. This feature is possible, due to its hydrophilic backbone structure. Its backbone can also prevent hydrogels undergoing dissolution process, although some hydrogels are designed to disintegrate and dissolve after absorbing a specific amount of water or chemical molecule (Calo´ and Khutoryanskiy, 2015; Galante et al., 2016; Lai and He, 2016) (Fig. 11.1). To prepare hydrogels, it is required to crosslink hydrophilic polymers to form a three-dimensional structure (3D), as shown in Fig. 11.2. The cross-linking process can be done through the use of chemical or physical interactions, or radiation (Gehrke and Lee, 1990; Amaral, 2009). Because of hydrogels’ biocompatibility, they can be used in different biomedical applications, such as drug delivery systems, wound dressings, contact lenses orthodontic applications, stem cells engineering, immunomodulation, cellular, and molecular therapies (Hoffman, 2002; Ahmed, 2015; Discher et al., 2009; Tibbitt and Anseth, 2009). It is important to observe that hydrogels absorb water and physiological fluids, which leads to the increase of the biodegradability and further biocompatibility (Motta, 2009) (Fig. 11.3). In addition to maintain a moist environment, hydrogels can also be used to stimulate angiogenesis, autolytic debridement, and neuroprotection (Bradbury et al., 2008; Sackheim et al., 2006). The elastic nature of hydrated hydrogels allows minimizing of the irritation of surrounding tissue after an implantation;

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FIGURE 11.2 Chemical hydrogel preparation scheme by a cross-linking agent. Notice that when the hydrophilic polymer chains are put together, H2O molecules enter the structure, and this feature is responsible for making it flexible.

FIGURE 11.3 Chemical and physical hydrogels preparation. Source: Adapted by Hoffman, A.S., 2002. Hydrogels for biomedical applications. Adv. Drug Deliv. Rev. 54(1), pp.312. Available at: http://linkinghub.elsevier.com/retrieve/pii/S0169409X01002393.

11.2 Classification of Hydrogels

lowering interfacial tension between the surface of the hydrogel and body fluid, reducing the likelihood of a negative immune reaction (Bhattarai et al., 2010). Hydrogels are relatively deformable and can adapt to the shape of the surface to which they are applied; this feature makes them a promising mucoadhesive dressing (Hoare and Kohane, 2008). Regarding industrial use, hydrogels can be implemented in agriculture and have an indirect impact on human health, despite the lack of studies (Venturoli and Venturoli, 2011). Despite many other uses for biomaterials, this chapter will be specifically focused on biomedical applications in regenerative medicine, drug delivery systems, wound dressings, and contact lenses, among others.

11.2 CLASSIFICATION OF HYDROGELS Hydrogels can be classified in different ways, according to their characteristics (Dumitriu, 2001). It can be according to their source, as natural hydrogels (obtained naturally without chemical reaction), synthetic hydrogels (manufactured obtained from cross-linking process) or hybrid hydrogels (by adding a hybrid nanocomposite cross-linked with a natural hydrogel). Preparation method, degradability, cross-linking, physical properties, ionic charge and response can also be used to classify, as shown in Fig. 11.4 (Ullah et al., 2015; Ahmed, 2015; Pal et al., 2009; Calo´ and Khutoryanskiy, 2015).

11.2.1 CLASSIFICATION BASED ON RESPONSE Reversible or physical hydrogels: If molecular entanglements and/or secondary forces, such as ionic, H- bonding or hydrophobic forces, play the main role in forming the network. Physical gels are often reversible and it is possible to dissolve them by changing environmental conditions, such as pH, and the ionic strength of solution or temperature.

FIGURE 11.4 Hydrogel classification key. Source: Image by Author.

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Table 11.1 Hydrogel Response Profile Chemically responsive

Physically responsive

Biochemically responsive

• • • • • • • • • • •

pH-responsive Glucose-responsive Oxidant-responsive Temperature-responsive Pressure-responsive Light-responsive Electric field Magnetic field Antigens-responsive Enzymes-responsive Ligands-responsive

Adapted by author.

Permanent or chemical hydrogels: the network of covalent bonds joining different macromolecular chains can be achieved by cross-linking polymers in the dry state or in solution. These gels may be charged or noncharged depending on the nature of functional groups present in their structure. The charged hydrogels usually exhibit changes in swelling upon variations in pH, and it is known that they can undergo changes in shape when exposed to an electric field. Biochemical hydrogels are responsive to antigens, enzymes and a variety of ligands. As seen in Table 11.1 (Calo´ and Khutoryanskiy, 2015).

11.2.1.1 Stimuli-responsive hydrogels Stimuli-responsive hydrogels respond to environmental stimuli and experience unexpected changes in their growth actions, network structure, mechanical strength and permeability, so consequently called environmentally sensitive, smart hydrogels (Kashyap et al., 2005; Gil and Hudson, 2004; Ullah et al., 2015). Physical stimuli include temperature, pressure, light, electric fields, mechanical stress, magnetic fields, and the intensity of various energy sources, which change molecular interactions at critical onset points. Chemical stimuli include pH, chemical agents, and ionic factors, which change the interactions between polymer chains and solvents, and between polymer chains at the molecular level. Another class, which are called dual responsive hydrogels, results from a combination of two stimuli-responsive mechanisms in one hydrogel system. Polyacrylicacid-copolyvinyl sulfonic acid is an example of a dual responsive polymer system (Kang and Bae, 2003; Ullah et al., 2015). A biochemical stimulus involves the responses to ligand, enzyme, antigen, and other biochemical agents (Gil and Hudson, 2004; Ullah et al., 2015). Accordingly, stimuli-responsive hydrogels are attractive biomaterials for biomedical, pharmaceutical, and biotechnology applications (Kashyap et al., 2005; Ullah et al., 2015).


11.2.2 CLASSIFICATION BASED ON TYPE OF CROSS-LINKING Hydrogels are classified into two different categories, according to the chemical or physical nature of the cross-link reactions. Chemically cross-linked networks have permanent junctions. Conversely, physical networks have temporary junctions that arise from either polymer chain entanglements or physical interactions (ionic and/ or hydrogen bonds or hydrophobic interactions) (Hacker and Mikos, 2011).

11.2.3 CLASSIFICATION ACCORDING TO THE METHOD OF PREPARATION Homopolymeric hydrogels: polymeric networks obtained from a single monomer, in which the basic structural unit is any polymer network (Takashi et al., 2007; Ahmed, 2015). Homopolymers can have cross-linked skeletal structure, according to the nature of the monomer and polymerization protocol. Copolymeric hydrogels: are polymeric networks structured with two or more different monomers, one of them necessarily hydrophilic, configured as random, block, or alternating along the polymeric chain (Yang et al., 2002; Ahmed, 2015). Multipolymer interpenetrating polymeric hydrogel (IPN): consists of two independent cross-linked synthetic and/or natural polymers in network-like shape. In semi-IPN hydrogels, one component is a cross-linked polymer and another component is a noncross-linked polymer (Maolin et al., 2000; Hacker and Mikos, 2011; Ahmed, 2015).

11.2.4 CLASSIFICATION BASED ON SOURCE Hydrogels can also be classified into two groups, based on their origins (Wen et al., 2013; Ahmed, 2015). Over the years, synthetic hydrogels replaced natural hydrogels because they offer long service life, improved water absorption capacity, and higher gel strength. Synthetic polymers also have well-defined structures to improve degradability and functionality degrees, according to their use. Hydrogels can be obtained from synthetic components and demonstrate thermostability (Ahmed, 2015). However, there is a third group, the hybrid hydrogels (known as nanocomposites hydrogels). These can be exemplified by the following:

11.2.4.1 Nanocomposite hydrogels Nanocomposite (NC) hydrogels demonstrate organic-inorganic features and are of great interest as artificial three-dimensional biomaterials for biomedical applications. Nanocomposite hydrogels are prepared in water by chemically or physically crosslinking organic polymers with nanomaterials (NMs). The incorporation of hard inorganic NMs into the soft organic polymer matrix enhances properties of NC hydrogels. They are also promising candidates for artificial 3D biomaterials, especially in tissue engineering, in which they can simulate chemical, mechanical, electrical, and biological properties of biological tissues (Motealleh and Kehr, 2017).

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Nanoparticles, such as carbon-based nanomaterials (carbon nanotubes (CNTs), graphene, nanodiamonds), polymeric nanoparticles, inorganic/ceramic nanoparticles (hydroxyapatite, silica, silicate) recombined with the polymeric network to obtain nanocomposite hydrogels (Fig. 11.5). These nanoparticles physically or

FIGURE 11.5 Nanocomposite (hybrids) hydrogels for biomedical applications. Engineered nanocomposite hydrogels. A range of nanoparticles, such as carbon-based nanomaterials, polymeric nanoparticles, inorganic nanoparticles, and metal/metal-oxide nanoparticles, are combined with the synthetic or natural polymers to obtain nanocomposites hydrogels with desired property combinations. These nanocomposites networks are their physically or chemically cross-linked. By controlling the polymerpolymer or polymer-nanoparticle interactions, the physical, chemical, and biological properties of the nanocomposite hydrogels can be tailored. Source: Adapted by Gaharwar, A.K, Peppas, N.A., Khademhosseini, A., 2014. Nanocomposite hydrogels for biomedical applications. Biotechnol. Bioeng. 111 (3), 441453.


covalently interact with the polymeric chains and result in novel properties of the nanocomposite network (Goenka et al., 2014; Schexnailder and Schmidt, 2009). The development of NC hydrogels with tailored functionality opened up new possibilities into developing advanced biomaterials for various applications (Peppas et al., 2007).

11.2.4.1.1 Nanocomposite hydrogels from carbon-based nanomaterials Carbon-based nanomaterials such as graphene and CNTs, are widely used to incorporate new features, such as high mechanical, electrical conductivity, and optical properties to the synthetic or natural polymers (Cha et al., 2013). For example, both CNT- or graphene-based NC hydrogels can be applied as actuators, conductive tapes, biosensors, tissue engineering scaffolds, drug delivery systems, and biomedical devices (Goenka et al., 2014; Kuilla et al., 2010).

11.2.4.1.2 Nanocomposite hydrogels from polymeric nanoparticles Polymeric nanoparticles, such as dendrimers, nanogels, liposomes, hyperbranched polymers, polymeric micelles, and core-shell polymeric particles gained attention with to their use in drug delivery systems, due to their ability to entrap hydrophobic or hydrophilic drugs (Joshi and Grinstaff, 2008; Seidlits and Peppas, 2007). However, further studies are required in both in vitro and in vivo animal experiments to examine their behavior in vivo, considering nanoparticles properties (Paul et al., 2016).

11.2.4.1.3 Nanocomposite hydrogels from inorganic nanoparticles The new generation of advanced biomaterials, by combining inorganic ceramic nanoparticles with natural or synthetic polymers, are the most promising systems. A range of bioactive nanoparticles have been reported, including hydroxyapatite (nHA), calcium phosphate, synthetic silicate nanoparticles, bioactive glasses, silica, glass ceramic, and b-wollastonite for biomedical applications (Hench and Polak, 2002). Incorporating inorganic nanoparticles within polymeric hydrogels may induce bioactive characteristics to the network. Many NC hydrogels have been synthesized by incorporating ceramic nanoparticles within the polymer matrix. Silicon, another ceramic nanoparticle mineral component, plays a key role in backbone development. It is known that silicon stimulates the osteogenic differentiation in human stem cells. It also promotes collagen type I synthesis. The addition of nHA to the polymeric network provided elastomeric properties, enhanced mechanical strength, and improved the physiological stability of the nanocomposite networks. Moreover, the addition of nHA resulted in enhanced cell adhesion, compared with poly(ehtilene glycol) (PEG) hydrogels. Consequently, cross-linking inorganic nanoparticles within polymeric hydrogels can induce bioactive characteristics to the network (Gaharwar et al., 2011, 2014).

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11.2.4.1.4 Nanocomposite hydrogels from metal and metal-oxide nanoparticles Most inorganic nanoparticles consist of minerals already in the body, playing a key role in most physiological chemical reactions in order to maintain homeostasis (Hoppe et al., 2011). For example, calcium is one of the most important components of bone and plays a vital role in bone development and maintenance. Its intracellular presence, along with phosphate, in osteoblasts can mineralize bone matrix and prevent tissue loss. Degrading products of ceramic nanoparticles results in a promising biological response and represents a potential use in different biomedical applications. Metal and metal-oxide nanoparticles have shown antimicrobial properties (silver (Ag) nanoparticles) and can thus be used as drug delivery systems (Schexnailder and Schmidt, 2009). Gold (Au), silver (Ag), and other noble metal nanoparticles can be used to prepare NC hydrogels, whereas metal-oxide nanoparticles include iron oxide (Fe3O4, Fe2O3), alumina, zirconia, and titania (TiO2), (Schexnailder and Schmidt, 2009). Paul et al. (2016) showed that microfabricated nanoengineered microgels can be used to pattern and control cellular behavior, and showed that NC hydrogels have high biocompatibility, as demonstrated by in vivo experiments. Specifically, hydrogels exhibited minimum immune responses, showing their ability for tissue engineering applications due to their biocompatibility.

11.2.4.1.5 Next generation of nanocomposite hydrogels Nanocomposite hydrogels for various biomedical applications is an arising field of study. The enhanced surface interactions between the nanoparticles and the polymer chains result in material properties that may be useful for various biomedical applications. However, further studies regarding NC biodegradation are required to evaluate their behavior after long periods of exposition.

11.2.5 INTELLIGENT HYDROGELS Intelligent hydrogels are polymeric materials that possess all the characteristics typical of a hydrogel, but which have different responses, according to external stimuli (Irie, 1993). That is, they vary shape and structure, depending on the environment where they are placed and their variations of some features of the same (Dusek, 1993). Consequently, there are hydrogels that respond to changes in pH, such as the acrylate-based hydrogels (Qiu and Park, 2001); temperature variations, such as N-isopropilacilamide-based hydrogels; and ionic strength of the medium changes, among others. The great advantage of these systems, when compared with the conventional hydrogels, is that it is possible to control several features, adjusting the media where they will be applied. For example, a system can be synthesized that is dependent on pH for therapeutic administration in the intestine (Middle tended


FIGURE 11.6 Representative phase diagram of a system with LCST and another with UCST. Source: Adapted by Almeida, J.F.D.S.L., 2010. Tese sobre Preparaçaõ e Caracterizaçaõ de Hidrogeís para Aplicaço˜es Biome´dicas. Faculdade de Cieˆncias e Tecnologia da Universidade de Coimbra. Departamento de Engenharia Quı´mica (Almeida et al., 2010).

basics) without the material and the drug being broken down in the stomach (acid medium). This type of hydrogels are often found as controlled-release systems of several substances in various media. For these type of intelligent hydrogels in particular, the base of poly(acrylic acid), a dissociation in aqueous solution with a pH greater than 5.5, causes the electrostatic repulsion of the polymer chain and, consequently, its expansion. For heat-sensitive gels, one of its components must be insoluble above or below a certain temperature, i.e., expose the gel to a lower temperature than the low critical solution temperature (LCST), or a higher temperature than the upper critical solution temperature (UCST) (Bromberg and Ron, 1998). In the first case, the gels reduce their volume with increasing temperature, creating a phase separation at that time. Conversely, gels with UCST contract with decreasing temperature. The following image presents the phase diagrams for systems with LCST and UCST (Fig. 11.6). The first polymer described in the literature as showing thermo-reversible characteristics was gelatin, a protein prepared by partial hydrolysis of collagen, containing the main aminoacids proline and glycine (Choi et al., 1999). Subsequently, several studies were presented on polysaccharides with similar properties, including materials with agarose, amylose, and amylopectin. In these cases, for high temperatures, polymers assume a random conformation. With the decrease of temperature, the formation of double and triple propellers (for the polysaccharides and gelatin, respectively) formed between the various join points chains, leading to the phenomenon of gelation of each compound. Fig. 11.2 shows

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a schematic of the process of gel formation in the cases described. All the changes described previously are reversible, i.e., the gel can repeat the transition/ gel and gel solution/solution without any limit in time of gel life. Considering the degree of expansion of a gel is related to its physical properties, the variation in volume as a response to any external stimulus causes changes in these properties, such as the kinetics of water absorption, permeability of polymer network, pore size and consequent drug release. It is known that some diseases change biochemical parameters of the organism such as pH, temperature, or concentration of substances. This type of material might demonstrate an advantage of its application compared to conventional hydrogels due to its ability to adapt to metabolic fluctuations of the body (Langer and Peppas, 2003; Zhang and Wu, 2004). Beyond the capacity of these gels to respond differently to external stimuli, there are other properties that favor the use of them as biomaterials. In this way, their elastic nature, similar to that of natural fabrics, or their biocompatibility, have led to greater use of them. The use of heat-sensitive gels in medicine has been vast. Implementation of these systems is in the release of insulin in diabetic patients, using the degradation of the system to allow the release of the compound in the sick body (Ratner et al., 1996; Kost and Langer, 2001). Another example was achieved by introducing to the polymeric matrix an enzyme, in this case glucose oxidase, which, in the presence of a particular sugar concentration, reacts, forming gluconic acid. This lowers the pH in the location, leading (in the presence of a gel responsive to pH) to its expansion and the release of the drug (Ratner et al., 1996).

11.3 APPLICATIONS OF HYDROGELS 11.3.1 TISSUE REGENERATION Over the years, the shortage of organs for transplants has brought about the need to develop new techniques to aid tissue regeneration or even replace them (Ersumo et al., 2016). Therefore, tissue engineering is growing through the application of multipotent stem cells, the extracellular membrane (ECM) in cell behavior, and development of biomaterials. The development of therapeutic measures for tissue regeneration is possible through the knowledge of the characteristics of mesenchymal stem cells (MSC), where their proliferative capacity and differentiation allow precursor structures able to incorporate the damaged areas of the patient’s anatomical and physiological point of view (Ahsan and Nerem, 2005; Elisseeff et al., 2005). That is, the MSCs are capable of differentiating into multiple cell types, depending on the setting, including neurons, myoblasts, and osteoblasts (Engler et al., 2006). The ECM of cells and tissues plays a complex role in cellular behavior, acting on the stability and shape of organs and tissues as well

11.3 Applications of Hydrogels

as actions such as adhesion, migration, proliferation, differentiation, and apoptosis (Abbott, 2003). The main goal of tissue engineering is based on the development of biocompatible materials, equivalent to the ECM, capable of providing supports, which serve as scaffolds for cells using a three-dimensional interaction (Joddar et al., 2016). Hydrogels are useful biomaterials for biomedical applications, because they mimic the mechanical behavior of the ECM of native tissue (Ersumo et al., 2016). These biomaterials containing 3D cells encapsulated within it, associated or not with growth factors, are an important technique that enables cell culture and tissue development, such as avascular tissues, such as bone and cartilage, as well as thin membranes such as skin (Va´zquez-Portalatıń et al., 2016; Wang et al., 2006; Joddar et al., 2016). The interior of these structures are more like cells in vivo when compared to standard cell culture. Hydrogels are biomaterials composed of synthetic polymers, such as those synthetically formed (poly(ethylene glycol), poly(hydroxyethyl methacrylate), poly lactic acid, poly co-glycolic acid, oly-hydroxybutyrate, poly vinyl alcohol, poly(N-isopropylacrylamide-coacrylic acid)), or naturally occurring polymers (e.g., collagen, hyaluronan, heparin, chitosan, alginate, alginic acid, pectin, dextran, cellulose, chondrin sulfate, agarose, pullulan (Padhi et al., 2016; Joddar et al., 2016). These polymers have as an essential characteristic: the ability to swell by absorption of water and shrinking as a result of fluid flow, in response to osmotic pressure imbalances, electrostatic forces, and elastic restoring forces. Their use can mimic the environment of different tissues such as, e.g., cartilage (Ganji et al., 2010; Koetting et al., 2015; Spiller et al., 2008; Evmenenko and Budtova, 2000). This feature is also useful for the development of stimuli-responsive drug delivery networks, which can be biodegradable or biocompatible (Burdick and Vunjak-Novakovic, 2009). The advantage of using hydrogels is that they are adaptable to different forms of tissue damage, adapting to the structural holder. In this way, the cells are distributed homogeneously in their midst. Another advantage is that hydrogels have characteristics similar to the natural microenvironment and surrounding cells, allowing the diffusion of nutrients to the encapsulated cells. The nanoporous structure of the polymer hydrogels are determined by the size and degree of cross-linking between the fibrils (Drury and Mooney, 2003; Trevors and Pollack, 2005; Elisseeff et al., 2005). Furthermore, the use of synthetic polymers rather than the natural is most favorable, since it enables a change of the physical properties of polymers with more accuracy and repeatability. Consequently, it is possible to control issues such as adhesion surface cells or exposure of the same nutrients. In addition, additive techniques allow for precise spatial control in the manufacture of a hydrogel (Mandrycky et al., 2016). Hydrogels are among the leading biomaterials used analogously to the ECM, nevertheless most of them lack an essential property for the proper functioning cell: cell adhesion. With the exception of collagen, synthetic polymers have no biological ligands to interact directly with the cell surface receptors.

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Consequently, the cells have adhesion difficulties in structural support (Kim et al., 2000; Drury and Mooney, 2003). The integrins and growth factor receptors present on the cell membrane are responsible for the contact between the MEC and the cells, where they subsequently interact with matrix proteins, triggering different pathways that are vital to the cell (Jia et al., 2016). Accordingly, the literature suggests some techniques that allow for cell adhesion to the polymer hydrogel base (Elisseeff et al., 2005; Kim et al., 2000) or the development of substrates capable of interacting with the cell: commonly used functional fragments of ECM proteins (Koepsel et al., 2012; Jia et al., 2016). The tripeptide sequence Arg-Gly-Asp (RGD) is the primary recognition domain for cell adhesion by interacting with integrins, promoting cell adhesion. Some authors have reported the importance of adhesion to the polymer by means of covalent bonds in various types of hydrogel (DeLong et al., 2005; Jia et al., 2016; Drury and Mooney, 2003). When the ligands are immobilized, they effectively promote cell adhesion, allowing cell survival. However this does not occur when cells undergo apoptosis (Lopes et al., 2005; Zahir and Weaver, 2004). In this RGD sequence, many authors have associated other elements (degradable peptides, ionic charges, growth factors, inorganic mineral) for the stimulation of cellular interaction with specific tissues with the materials to hydrogels base. However, RGD is further applied in the preparation of matrices (Trojani et al., 2006; Elisseeff et al., 2005; Nuttelman et al., 2005). The regeneration of bone is based on approximations of mesenchymal stem cells (MSCs), based on photo-cross-linked alginate hydrogels (PAHs) used for repair of fractures, where a greater cellular retention time results in higher functional benefit (Ho et al., 2016). The morphogenetic protein-2 (BMP-2) is a potent bone osteoinductive for osteogenic differentiation. However, their supraphysiological concentrations can damage cells and result in apoptosis. Ho et al. (2016) examined the effects of PAHs with MSCs associated with low concentrations of BMP-2 and concluded that this combination results in a reduced rate of apoptosis, promoting bone formation, compared with only PAHs and MSCs. The challenge for bone tissue engineering is that tissue integration occurs prior to bacterial adhesion, preventing adhesion for certain bacterial species to human tissue and implanted biomaterials. Some bacterial strains are able to form biofilms. Consequently, these bacteria can often withstand host immune response and antibiotic therapies (Ribeiro et al., 2016; Arciola et al., 2012). However, the increasing number of antibiotic-resistant bacterial strains is also an important factor when considering the development of new antimicrobial approaches. In this way, metallic elements forms of nanoparticles, such as silver nanoparticles and gold nanoparticles are excellent for antimicrobial applications, due to their chemical stability, long life, and heat resistance. The antibacterial activity of this type of nanoparticle has been related to inhibition of enzymatic activities, prevention of DNA replication, and bacterial cell membrane disruption (Prabhu and Poulose, 2012; Ribeiro et al., 2016).


For cartilage repair, many biomaterials are studied. Alginate hydrogels (type-I collagen base) incorporated with MSC chondrocytes promotes differentiation and repair of the cartilaginous tissue. The addition of type-II collagen to these biomaterials is being studied aiming uniting its biological activity with the superior mechanical properties of type-I collagen (Va´zquez-Portalatıń et al., 2016). According to Joddar et al. (2016), the alginate hydrogel, a commonly used material for cell culture through ionic cross-linking, is mechanically unstable, limiting its usage in bone, heart, or tumor tissue engineering. Therefore, the authors proposed a strengthening base of these structures from multiwalled carbon nanotubes (MWCNT) and found that this association presented more stable and stiff hydrogels. Xiao et al. (2016) investigated the glutamine, histidine, arginine-glutamic acid-aspartic acid-glycine-serine peptide (QHREDGS) immobilized to the chitosan-collagen hydrogel, to accelerate wound healing in diabetic rats. They concluded that this hydrogel, used as a delivery vehicle associated with the QHREDGS, showed good results by accelerating the healing of wounds by reepithelialization and granulation tissue formation. Mahdieh et al. (2016) synthesized nanocomposite biomaterial consisting of a blend of thermoplastic starch and ethylene vinyl alcohol as the polymer matrix, and nanostructured forsterite as the ceramic reinforcing phase for bone tissue engineering applications. Moreover, this addition modified the pH in the methyl thiazolyl tetrazolium (MTT) assay and stimulated cell proliferation. Cell adhesion assays indicated a favorable interaction between cells and the biomaterial. The proposed nanocomposite has appropriate biocompatibility, as well as mechanical properties, in order to be used in bone tissue engineering.

11.3.2 TISSUE DRESSING Taking care of wounds is a dynamic and complex process, especially when they become chronic. Chronic wounds affect the quality of life of patients, besides generating great cost by prolonged hospitalization (Lima and Guerra, 2011). They are rapidly evolving, resistant to some types of treatments, and depend on some conditions that preclude their recovery (Candido, 2001). The initial aim in the treatment of wounds is promoting healing, with rapid functional recovery, and esthetics, through tissue growth and regeneration. The healing process involves a complex biological process which involves the balance of vascular, inflammatory activities associated with connective tissue cells and epithelials (Kokabi et al., 2007). Bacterial resistance to commercially available antibiotics is another serious problem in the management of wounds. Infection of serious wounds needs care, because these infections result in formation of exudates, delay healing and facilitate unfit collagen deposition (Jayakumar et al., 2010; Madhumathi et al., 2010).

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Formerly, the materials used for dressings had the sole purpose of inhibiting the bleeding and protecting against physical, mechanical, and biological external agents, having a minimal role in the tissue regeneration process and subsequent healing (Kokabi et al., 2007). There are currently three types of dressings used in the clinic: biological, synthetic, and biological-synthetic. Most biological dressings such as pigskin are immunogenic, have a high risk of contamination and present poor adherence. Synthetic bandages minimizes the inflammatory reaction but might be not biocompatible. The use of biologic-synthetic dressings consists of polymer and biological materials patched in bilayers, pottentially less immunogenic and more biocompatible (Jayakumar et al., 2010). Conventional dressings have little flexibility, low mechanical strength, lack of porosity, and lack of antimicrobial activity. Moreover, they usually adhere to the wounds which worsens the situation (Sudheesh Kumar et al., 2012). Dressings may not exhibit promising toxicity and allergic reactions, and cannot be adherent enough to permit its removal without trauma to the treated area. They should also provide a moist environment, protection against secondary infections, absorb the formation of exudate fluids from the wound, reduce necrosis of the wound surface, avoid dissecting this wound, and stimulate growth factors (Paul and Sharma, 2004; Higa et al., 1999; Lin et al., 2001; Jayakumar et al., 2010). Hydrogel dressings are receiving attention due to its advantages, such the creation of a moist environment, immediate pain control, gas exchange necessary for tissue recovery, barrier to microorganisms, easy replacement, high hemostatic potential, higher biocompatibility and biodegradability (Jayakumar et al., 2010). Additionally, these dressings are usually transparent, allowing monitoring of wound healing and favoring healing up to two times faster than traditional therapies (Cabrales Vega, 2012). In patients with burns, the fluid balance is also essential for the maintenance of body temperature and metabolic rate. Dressings for these patients should have ease of application, removal, and homogeneous adhesion in order to prevent the formation of bags satisfaction liquid, which promotes bacterial proliferation. In situforming hydrogels are being studied for the preparation of dressings that enable a mold of the wound dressing, allowing a conformation free of wrinkles and fluting (Balakrishnan et al., 2005). Balakrishnan et al. (2005) developed an oxidized alginate/gelatin system as an in situforming wound dressing material. Consequently, the healing properties of the alginate through added proinflammatory stimuli to induce hemostasis caused by gelatin sponge resulted in a synergy of desirable components in a dressing. Added to this was a borax matrix, having antiseptic effects, which prevented bacterial colonization in wounds induced in rats. Dressings with a basis of silver or silver particlecoated medical devices have been used for quite a while as promising antimicrobial agents (Rai et al., 2009).


These silver nanoparticles, Ag-loaded chitosan nanoparticles, are being added to an improved version of its antimicrobial properties for the prevention of wound infection (An et al., 2010). However, this biomaterial increased concern of the potential toxicity of silver. The asymmetric chitosan membrane is like a sponge and has favorable requisites for wound healing, while incorporated silver sulfadiazine, a potent antimicrobial, acts on the controlled and sustained release of its ions (Mi et al., 2003; Jayakumar et al., 2011). Mi et al. (2003) tested their asymmetric chitosan membrane incorporated into silver sulfadiazine, which showed a potent antimicrobial activity and a decreased toxicity of silver. Chitosan polymers provide a nonprotein 3D matrix stimulating the proliferation of fibroblasts and contribute to collagen deposition and hyaluronic acid synthesis (Paul and Sharma, 2004). Kumar et al. (2012) developed compound chitosan hydrogel/nano zinc oxide composite bandages with the objective of being used in chronic wounds, burns, and diabetic foot ulcers. The authors noticed several benefits, including faster reepithelialization, improved water retention, uptake of wound exudate, antimicrobial activity, and an increase of blood coagulation in the studied areas. This formulation did not show cellular toxicity. Ribeiro et al. (2009) examined the cytotoxicity of chitosan hydrogel dressing in skin fibroblasts isolated from rats. It demonstrated the absence of toxicity in the cells. Furthermore, histological analysis revealed no inflammatory reaction and no pathological changes. The authors also showed several advantages, including the type growth factors and cell proliferation. Chitosan hydrogels have also shown to encourage the chemotaxis of fibroblasts in cell culture when supplemented with 5% fetal bovine serum, demonstrating its ability to act as an occlusive dressing (Ishihara et al., 2002). Kweon et al. (2003) evaluated the effect of healing wounds from a water-soluble chitosan/heparin complex in rodents, showing almost complete regeneration within 15 days. Occlusive chitin dressings are semipermeable. Its advantages are its gas permeability, transmembrane fluid transporte becoming flexible, soft and transparent. Also, they lack of cytotoxicity and allergenicity (Yusof et al., 2003). Mi et al. (2003) prepared a novel asymmetric chitosan membrane and reported promising results for this membrane preventing the penetration of bacteria and dehydration of the wound surface, while allowing drainage of exudate and increased tissue regeneration A new biomaterial hydrogel, based on a natural polysaccharide, pullulan, through chemical cross-linking was studied by Li et al. (2013), which showed water absorption properties up to 4000%, preventing exudate accumulation in the wound bed, due to rapid hemostatic capacity, and showing no cytotoxicity.

11.3.3 CONTACT LENSES Silicone hydrogels (Si-Hy) can be used as promising material to prepare contact lenses, due to the presence of siloxane (Si-O) bonds, which increases gas

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permeability, and consequently biocompatibility. Silicone hydrogels’ most relevant aspect might be its drug delivery capacity, maintaining the drug integrity by preventing its oxidation and chemical structure changes (Xu et al., 2011; Helaly et al., 2017; Childs et al., 2016). Peng et al. (2010) and Kim et al. (2010) found that Si-Hy lenses improved drug release time, compared to conventional lenses. Hsu et al. (2013) performed an experiment that demonstrated the same therapeutic effect of cysteamine, by wearing Si-Hy lenses for 2 h, as by pouring eye drops for hours. Formulating the lenses with vitamin E also offers effective protection against UV radiation, preventing corneal damage, and facilitates drug diffusion by soothing its barrier. Because of the hydrophobic character of the lens surface, it is important to secure the wettability profile to prevent mechanical eye damage or biofilm growth (Gavara & Compan˜, 2016; Peng et al., 2010). Garcıá-Porta et al. (2016) suggested that preventing the Si-Hy contact lens dehydration is a way to avoid eye damage. Moreover, it is important to keep the eye surface in homeostasis to prevent any ocular disease complication. The use of proper solutions to make the lens surface wet is highly recommended, especially when it will be worn for a long period of time. It is found that the proper hydration is a parameter to keep both eye and lens integrity (Lau et al., 2016). Hyaluronic acid inclusion is the most promising wetting technique because of its natural presence in biological systems and extreme biocompatibility, without compromising the overall Si-Hy lens integrity (Stach et al., 2016; Hook and Wygladacz, 2016). It is known that this hydrogel preparation is susceptible to bacterial growth that will cause further biofilm formation and a potential biosystem failure. Riber et al. (2016) found that silicone hydrogel preparation embedded with irgasan significantly reduced bacterial growth, suggesting high drug delivery potential. However, further studies are required to assess toxicity and body fluid interaction. Another drawback is protein adsorption, which may interfere in light transmission and alter its surface shape, damaging the ocular globe. The use of proper cleaning solutions is the key to reducing protein adsorption and avoid dryness for both eyes and Si-Hy contact lenses (Helaly et al., 2017; Childs et al., 2016).

11.3.4 DRUG DELIVERY SYSTEM Drug delivery systems are methods in which a drug will be administered to an animal or human and will increase its efficacy and safety profile (Safari and Zarnegar, 2014; Jain, 2008; Tiwari et al., 2012). Oral and intravenous administrations are the most frequent systems used for pharmacological therapy. The use of nanotechnology as drug carriers is becoming a rising field of study, due to its high specificity and effectiveness (Tam et al., 2016; Tan et al., 2017). Mucoadhesive hydrogels are an innovation for drug delivery in mucous body areas that have lower adhesion of other molecules. Different drug approaches are tested to evaluate their efficiency, compared to the conventional suppository or


oral administration pathway (Lo et al., 2013). Xu et al. (2016) compared the efficacy of sulfasalazine-loaded rectal hydrogel injections, with oral administration, for colorectal cancer treatment. The gel preparation was based on catecholmodified chitosan cross-linked with genipin, and then loaded with sulfasalazine. The experiment consisted of three gel injections using a syringe in which the gel was prepared. The authors found that, despite the fact that the drug concentration delivered by the hydrogel was half of the oral dose, lower plasmatic toxic byproduct levels were found, characterizing a safer profile and demonstrating a more therapeutic pathway. They concluded that the rectal hydrogel injection was more effective and had a safer drug profile, compared with conventional oral administration. They also found an intriguing event in which the hydrogel preparation without sulfasalazine demonstrated a discreet but significant healing potential, suggesting that the hydrogel itself may have antiinflammatory potential. This finding was sustained by the histological evaluation. Hydrogels can also take advantage of temperature, light, pH, redox potential, magnetic field and ultrasound to undergo conformational changes (e.g., swelling or shrinking) from its original state or degradation, in order to keep drugs plasmatic rates constant (Thambi et al., 2016). Chen et al. (2016) developed a new strategy for type-2 diabetes control. They prepared three thermosensitive poly(ε-caprolactone-co-glycolic acid)-poly(ethylene glycol)-poly(ε-caprolactone-co-glycolic acid) (PCGA-PEG-PCGA) triblock copolymers with similar molecule weights but different ε-caprolactone/glycolide (CL/GA) ratios, and then loaded them with Lira (an antidiabetic polypeptide that can reduce blood glucose and decrease insulin resistance). As a result, they found that a single injection of this preparation caused a notable hypoglycemic efficacy up to 1 week. Three successive injections of the gel within 1 month significantly lowered glycosylated hemoglobin, indicating a promising new treatment with better patient compliance (Table 11.2).

11.3.5 HYGIENE PRODUCTS Superabsorbent polymers have been introduced into agriculture and the diaper industry about 30 years ago. Since then, their uses have been extended to several other applications, due to their excellent water retention. Superabsorbent polymers were firstly commercially produced in Japan in 1978 for use in feminine napkins, and this early material was represented by a cross-linked starch-g-polyacrylate (Masuda, 1994) At the end of the 1990s, superporous hydrogels were introduced and presented as a different type of water-absorbent polymer system. As superabsorbent polymers, superporous hydrogels (SPH) are formed by covalently crosslinked hydrophilic polymers, but, unlike superabsorbent polymers, they show exceptional size-independent fast-swelling kinetics. The first generation of superporous hydrogels was generally made from highly hydrophilic acrylamide, salts of acrylic acid, and sulfopropyl acrylate. Later generation of SPHs are represented by hybrid superporous hydrogels, produced by adding a so-called hybrid agent

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Table 11.2 Short List of Hydrogel Applications Kang-Mieler et al. (2016)

Poly lactic-co-glycolic acid (PLGA)

Li et al. (2013)

Pullulan hydrogel

Balakrishnan et al. (2005)

Oxidized alginate/gelatin 1 borax

Mi et al. (2003)

Asymmetric chitosan 1 silver sulfadiazine

Ribeiro et al. (2009)

Chitosan hydrogel

Yusof et al. (2003)

Chitin

Mi et al. (2003)

Asymmetric chitosan membrane

Paulo et al. (2010)

Hydrogel poly 2-hidroxietil dimetacrilate (PoliHEMA) 1 polypropylene

• Advantages for drug delivery in ocular structures • Disadvantages • Difficult sterilization • Limited shelf • Damage to biopharmaceuticals in the cross-linking process • Unwieldy • Increase of 4000% water absorption • Rapid hemostatic activity without cytotoxicity • Improved healing and hemostasis • Prevention of bacterial colonization • Potent antimicrobial action • Low toxicity • Type of growth factors • Cell proliferation • Inflammatory reaction and no toxicity • High permeability to vapors and gases • Flexibility • Transparency • Absence of cytotoxicity and inflammatory reactions • Drainage of exudate • Increased tissue regeneration • Correction meshes for tissue defects and hernias • Reduction in the formation of visceral adhesions in rats. • Increased inflammatory reaction (Continued)


Table 11.2 Short List of Hydrogel Applications Continued Bhowmick et al. (2016)

PVA 1 silver nanoparticles

Varaprasad et al. (2010)

Hydrogel AgNPs-curcumina

García-Porta et al. (2016)

Silicon hydrogel

Xu et al. (2016)

Sulfasalazine loaded mucoadhesive hydrogel Thermosensitive poly (ε-caprolactone-co-glycolic acid)poly(ethylene glycol)- poly (ε-caprolactone-co-glycolic acid) (PCGA-PEG-PCGA)

Chen et al. (2016)

• Maintaining the sterile environment in dressings • Dressings antimicrobial properties • Increased biocompatibility • Improved drug delivey • Safer drug profile • Efficient glycemic control • Lower glycosylated hemoglobin

Adapted by authors.

(natural or synthetic water-soluble or dispersible polymer capable of chemical or physical cross-linking) to the previously made superporous hydrogels (Masuda, 1994; Calo´ and Khutoryanskiy, 2015). Superabsorbent hydrogels, in particular the acrylate-based materials, are extensively used in hygiene products to absorb fluids. They are able to hold moisture away from the skin, promoting skin health, preventing diaper rash, and providing comfortable use. Parents in all the industrialized countries, as well as hospitals around the world, employ disposable diapers containing superabsorbent polymers (Calo´ and Khutoryanskiy, 2015). A further increase in the use of these materials is observed in training pants and adult incontinence product markets. Superabsorbent polymers can also prevent the colonization of germs, reducing the risk of fecal contaminations, and potential spread of gastrointestinal infections. The first use of superabsorbent polymers in the diaper industry was in 1982 in Japan, with its subsequent use in sanitary napkins. With the technological advance, diapers became thinner and improved water retention performance, up to 50% its size (Calo´ and Khutoryanskiy, 2015).

11.3.6 OTHERS APPLICATIONS We currently have several applications of hydrogels, as well as the establishment of the first synthetic hydrogels by Wichterle and Lim in 1954 (Wichterle, 1960; Ahmed,2015). Hydrogel technologies may be applied to hygienic products (Singh et al., 2010), agriculture (Amulya, 2010), sealing (Singh et al., 2010), coal dewatering (Sun et al., 2002), artificial snow (Singh et al., 2010), food additives (Chen et al., 1995), diagnostics (Van der et al., 2003), wound dressing

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(Sikareepaisan et al., 2011), separation of biomolecules or cells (Wen et al., 2013) and barrier materials to regulate biological adhesions (Roy et al., 2010), and biosensors (Krsko et al., 2009).

11.4 HYDROGEL TECHNICAL FEATURES The functional features of an ideal hydrogel material can be listed as follows (Zohuriaan-Mehr, 2006; apud Ahmed, 2015): • • • • • • • • • • •

the highest absorption capacity in saline desired rate of absorption (preferred particle size and porosity) depending on the application requirement. the lowest price the highest durability and stability in the swelling environment and during the storage the highest biodegradability without formation of toxic species following the degradation the highest absorbency under load pH-neutrality after swelling in water photo stability colorlessness, odorlessness, and absolutely nontoxic the lowest soluble content and residual monomer rewetting capability (if required) the hydrogel has to be able to give back the imbibed solution or to maintain it, depending on the application requirement (e.g., in agricultural or hygienic applications).

Of course, it is impossible that a hydrogel sample would simultaneously fulfill all the above mentioned required features. For example, a hygienic product of hydrogels must possess the highest absorption rate and the lowest rewetting, and the hydrogels used in drug delivery must be porous and respond to either pH or temperature.

11.5 METABOLISM AND HYDROGELS Nanoparticles, owing to their unique physico-chemical properties, have found application in various biological processes, including metabolic pathways taking place within the body (Gomes et al., 2016). Nanoparticles in metabolic pathways and their influence in the energy metabolism, fundamental criteria for the survival, and physiological activity of living beings. The human body utilizes energy (ATP) derived from food resources through a series of biochemical reactions involving several enzymes, cofactors (metals, nonmetals, vitamins etc.) through the metabolic pathways (glycolysis, tri

11.6 Regulation of Hydrogels

carboxylic acid cycle, oxidative phosphorylation, electron transport chain, etc.) in cellular system (Gomes et al., 2016). The use of tolterodine is recommended for the treatment of urinary incontinence and other overactive bladder (OAB) symptoms (Andersson et al., 1998). After oral administration, it is rapidly absorbed through the gastrointestinal tract and metabolized in the liver from 5-methyl group to 5-hydroxymethyl tolterodine (5-HMT), under the oxidation action of CYP2D6 in extensive metabolizers (EMs) (Dmochowski and Appell, 2000). 5-HMT is the major active metabolite of tolterodine and demonstrates an antimuscarinic property comparable to tolterodine (Abrams et al., 1998). 5-HMT hydrogels significantly improved the therapeutic effect, avoided the metabolism difference of enzymes in bodies and provided a single active compound in plasma. Compared with tolterodine hydrogels in early reports (Sun et al., 2013), it realized effective dose control and reduced the potential adverse effect from two active compounds. 5-HMT hydrogels appeared to be a promising new therapeutic for the treatment of overactive bladder (Liu et al., 2017). Dutta et al. (2016) showed that the network structure of the hydrogel collapsed at 37 C and thus the residual lysozyme might be entrapped in the collapsed matrix, thereby retarding the lysozyme release. At higher temperature, hydrophobic interactions between the protein and hydrogel could also decrease the release rate. These types of hydrogels may be used in various applications, e.g., adsorbents or carriers of other drug/biomolecules. Thomas et al. (2007) concluded that silver nanoparticles can be produced within the swollen polymer network. The developed silver-hydrogel nanocomposite demonstrates optimum antibacterial activity against Escherichia coli. Their bacterial action depends on size of the particles, amount of silver nanoparticles within the hydrogel, and amount of monomer acid in the feed mixture. The formation of silver nanoparticles within the gel takes place due to the entrapment of Ag1 ions into the swollen hydrogel network via Ag1H1 ion-exchange mechanism, followed by citrate reduction. Goetsch et al. (2015) demonstrated that degradable hydrogels are widely used in different tissue regeneration approaches. A notable feature of these hydrogels is that, according to cell invasion rate and cellular enzymatic activity, the peptidebased cross-linking of the polymeric hydrogel might change. Future studies should lay emphasis on the formation mechanisms, with indepth insights into interfacial interactions, the tactical functionalization of hydrogels for desired properties, and expanding of their applications.

11.6 REGULATION OF HYDROGELS Materials at nanoscale have different properties, compared with the same material in bulk state. The explanation of this effect is related to the increased surface area and the domain of quantum size effects (Dhawan et al., 2011). The spectrum of

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possible applications of nanomaterials has reached a magnitude and versatility which initiates a real technological revolution. Stimuli-responsive, three-dimensional cross-linked hydrogels have gained tremendous research interest for the past few decades, for their numerous biomedical applications, including materials for drug delivery. Such hydrogels can swell/deswell drastically in response to a small change in the external conditions such as temperature, pH, ionic strength, solvent composition, and light intensity (Kawasaki et al., 1997; Hirotsu et al., 1987; Tanaka, 1981; Mukae et al., 1993). Stimuli-responsive hydrogels, which can change their properties in response to external stimuli such as temperature and pH have increasingly important applications in various fields. Hydrothermal (HT) treatment of hydrogels and the resulting changes in chemical structure. The volume expansion, rapid swelling, and increased lower critical solution temperature (LCST) of the treated hydrogels may result from the decross-linking of their network and the generation of carboxyl groups from the hydrothermal hydrolysis of amide bonds. First, the hydrogel deswells because of its temperature responsiveness, during which water is expelled and hydrophobic forces become dominant. The different swelling behaviors of the hydrogels are due to their different crosslinking densities, where the highly cross-linked core has less chain mobility and therefore swells less and more slowly than the shell. The pH-responsiveness of the hydrogel core and shell (Rongcong and Chia-Hung, 2012; Matsuo and Tanaka, 1988; Kali et al., 2013). For example, poly(nisopropylacrylamide) (PNIPAM), a temperature-responsive polymer,is widely used in the fabrication of responsive hydrogels. PNIPAM hydrogel can undergo a reversible volume phase transition near its LCST, which is close to physiological temperature (Matsuo and Tanaka, 1988; Kali et al., 2013). Because of this unique property, PNIPAM hydrogels are attractive for various applications, such as smart actuators, microfluidic valves, microlens optical systems, soft biomimetic machines, and drug delivery vehicles. Moreover, long-term cell culture experiments have demonstrated that PNIPAM hydrogels are biocompatible to several cell types, and can be potentially used in the field of tissue engineering. Despite their application potential, PNIPAM hydrogels prepared using the conventional redox initiation method exhibit uncontrollable swelling and an untunable LCST, which limits their application. To synthesize core-shell hydrogels, a partially cross-linked PNIPAM hydrogel was first synthesized by interrupting the redox initiation polymerization process (Rongcong and Chia-Hung, 2012).

11.7 POTENTIAL RISKS OF HYDROGELS Toxicological studies are important to avoid potential adverse effects caused by nanomaterials on humans and the environment, avaliating the characteristics of health hazard, in vitro and in vivo dose-response process, the exposure time of

11.8 Nanoparticle Biosafety

the organism to the nanoparticle, and the hazard identification (Kuempel et al., 2012; Pate´-Cornell, 1996). Nanoparticles hold great promise for in vivo diagnostic or therapeutic applications. Significant efforts have been made in the field of nanotoxicology, concurrent with the development of new types of engineered nanoparticles. Studies with synthetic poly(ethylene glycol) (PEG) hydrogels are investigated to provide the reproducibility in structure and cell support required for toxicology studies. Cellular proliferation was monitored over several weeks to demonstrate the applicability of this new measurement system. It is expected that this system could be adaptable to the needs of the toxicologist (Mansfield et al., 2014). At this moment, there are a few studies demonstrating hydrogel toxicity. The cream type of denture adhesives cannot be easily removed from oral mucosa after use and have the potential risk to change the oral flora. The effects of the temperature-responsive hydrogel Pluronic F-127 (PF) on the complex viscosity of denture adhesives were evaluated. Carboxy methylcellulose (CMC) mass fractions (1%, 2%, 3%, and 4%) were added to 20% and 25% PF hydrogels. Complex viscosity was measured over a temperature cycle (40 C-10 C-40 C) and fixed temperature points (23 C and 37 C). Adhesive strength tests were performed with 2 resin plates at 23 C and 37 C. One commercial cream-type denture adhesive, New Poligrip (NP), was evaluated as a control. Complex viscosity values for PF 20% groups at 23 C were lower than those for NP at 37 C. Adhesive strength of PF 20% with CMC 2%, was higher at 23 C when compared to NP at 37 C, which suggests that PF 20% CMC 2% is an effective adhesive and is easily removed after mouth rinsing (Zhao et al., 2016). Contact lens-induced papillary conjunctivitis (CLPC) continues to be a major cause of dropout during extended wear of contact lenses. This retrospective study explores risk factors for the development of CLPC during extended wear of silicone hydrogel lenses (Tagliaferri et al., 2014).

11.8 NANOPARTICLE BIOSAFETY Nanoparticles and biointeraction, in terms of metabolic pathways and immune function of the body, has been overviewed in the last two subheadings. The issue which popped up every now and then is regarding these particle’s fate which is an integral part of “Biosafety and Risk” issues associated with these particle. Risk assessment is of priority, since it yields many benefits, including early identification and prioritization of health, biosafety, and environmental concerns. Before implementing these particles in various arenas of biological science, its very necessary to have a complete understanding of these engineered particles within biological systems. A full understanding of the effects imposed by these nanoparticles will make a major contribution to the risk assessment, which is urgently needed to ensure that

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Table 11.3 Relationship Between the Size of Silver Nanoparticle and Reduced Contact Microorganism. Font: Quator Nanotechnology Particle Size (nm) 30 70 150 300

Microorganism Reduction (%) 99.9 99.4 97.7 97.5

Adapted by author.

products that utilize nanoparticles are made safely, are exploited to their full potential, and disposed of safely. Nanoparticles can have potential effects on organs, tissue, cellular, subcellular, and protein levels, due to their unusual physicochemical properties (e.g., small size, high surface area-to-volume ratio, chemical composition, crystalline aspect, electronic properties, surface structure reactivity and functional groups, inorganic or organic coatings, solubility, shape, and aggregation behavior). From the experimental evidences available, it is clear that particles engineered at nanometer scale can efficiently interact and influence biological systems. These particles come in contact with our body through skin, nasal cavity or via food or water into the gastrointestinal tract. Human skin, lungs, and the gastrointestinal tract are in constant contact with the external environment. These three routes are probable points of entry for natural nanoparticles. Injections and implants are other possible routes of exposure, but are mostly limited to engineered nanomaterials for biomedical applications. Owing to their miniature size, these particles have the capacity of translocation from their entry portals into the circulatory and lymphatic systems, and gradually to body tissues and organs, thereby producing irreversible detrimental damage to cells by oxidative stress and/or organelle injury. After administration of a nanoparticle into a biological system, how these particles penetrate the cell and goes to various organs and tissues is of concern to toxicologists; from the route of entry to its passage to different body compartments, depends on the type, size and functionalization of the nanoparticles (Table 11.3).

11.9 FINAL REMARKS Nanotechnology is the next revolution in the history of scientific innovations and advancements. It is an interdisciplinary field which bridges various disciplines, from physics, chemistry to biology, from engineering to medicine, thereby establishing new trends of technological applications. This privileged situation has led to the development of new products containing new features that, despite legislation currently under construction, already participates in our daily lives.

References

Nanoparticles, owing to their unique physic-chemical properties, have found their application in various fields, starting from electronics to biomedical science and therapeutics. Nevertheless, many important challenges still need to be overcome, such as the industrial-scale processes, understanding the interactions with biosystems (impact against humans and environment), and regulation issues. Nanotechnology is considered a key to the 21st century and holds promises in sectors like electronics, chemistry, biotechnology and medicine. In fact, nanoparticles and nanobased products are considered to be prodigious tools participating in various areas of biological applications, being considered one of the enormous scientific and technological challenges that nanoscale offers the world. Recently, many hydrogel networks have been designed to address different applications. When compared to polymeric hydrogels, NC hydrogels demonstrate superior physical, chemical, electrical, and biological properties. The favorable property of these hydrogels is their ability to swell when put in contact with an aqueous solution. Their use as sensors, drug delivery, stem cell engineering, and regenerative medicine might represent a huge advance for biomedical science, although further studies are needed to better elucidate their action mechanisms. Nanotechnology is growing as innovative area, bringing benefits to health. However, we must not forget about the possibility of toxic effects.

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CHAPTER

Silk-based matrices for bone tissue engineering applications

12

Promita Bhattacharjee1, Prerak Gupta2, M. Joseph Christakiran2, Samit K. Nandi3 and Biman B. Mandal2 1

Indian Institute of Technology, Kharagpur, West Bengal, India 2Indian Institute of Technology Guwahati, Guwahati, Assam, India 3West Bengal University of Animal and Fishery Sciences, Kolkata, West Bengal, India

CHAPTER OUTLINE 12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10 12.11 12.12

Introduction .................................................................................................440 Global Perspective on Orthopaedic Trauma Management ................................441 Rising Needs for Bone Grafts ........................................................................441 Ideal Scaffold for Bone Tissue Engineering....................................................442 Lacunae of Current Materials and Practices for Bone Regeneration.................443 Bioderived and Synthetic Materials for Bone Tissue Engineering ....................444 Context and Structure of the Chapter .............................................................445 Various Sources of Silk ................................................................................445 Silk From Silkworms ....................................................................................447 Spider Silk ..................................................................................................448 Benign Aspects of Silk for Bone Tissue Engineering .......................................448 Processing Silk Into Various Formats ............................................................450 12.12.1 Particulate Leaching............................................................... 450 12.12.2 Nanofibrous Scaffolds Using Electrospinning ............................ 451 12.12.3 Biopatterning ......................................................................... 452 12.12.4 Knitted Scaffolds ................................................................... 452 12.12.5 Freeze-Drying......................................................................... 452 12.12.6 Silk Microparticles and Microfibers as Reinforcements .............. 453 12.12.7 Hydrogels .............................................................................. 453 12.12.8 Computer-Aided Fabrication of SF Scaffolds (3D Printing) ......... 453 12.13 Silk Composites for Bone Tissue Engineering.................................................454 12.13.1 Hydroxyapatite ....................................................................... 454 12.13.2 Clay- or Silica-Based Additives ................................................ 456


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440

CHAPTER 12 Silk-based matrices for bone tissue engineering

12.13.3 Bioactive Glasses ................................................................... 456 12.13.4 Silk Inclusion in Other Substrates............................................ 457 12.13.5 Other Miscellaneous Additives for Silk Scaffolds ....................... 458 12.14 Beyond the Classic Mulberry Silk Fibroin ......................................................458 12.14.1 Nonmulberry Silk ................................................................... 458 12.14.2 Spider Silk............................................................................. 459 12.14.3 Silk Sericin............................................................................ 460 12.15 Recent Trends in Bone Tissue Engineering ....................................................460 12.15.1 Use of Bioreactors .................................................................. 460 12.15.2 Cocultures of Multiple Cells on Silk Scaffolds ........................... 461 12.15.3 Growth Factors Delivery Through Silk Scaffolds......................... 462 12.15.4 Gene Therapies ...................................................................... 462 12.15.5 Conclusion and Future Perspectives ......................................... 463 References .............................................................................................................464

12.1 INTRODUCTION Natural bone is a composite of 70 wt% inorganic phase (hydroxyapatite (HAp) nanocrystals) and 30 wt% protein (B90% collagen fibrils), with a hierarchical structure ranging from nano- to macroscale (Fan et al., 2010; Wang et al., 2011; Holzwarth and Ma, 2011). This composite structure of collagen fibers, reinforced by mineral nanoparticles, gives bone its remarkable strength and toughness (Swetha et al., 2010). The HAp nanoparticles in bone tissue are about 50 3 25 3 3 nm3 in size (Venugopal et al., 2010). The HAp in bone tissue has a Ca to P ratio of 1.67. Natural bone has an aspect of self-assembly to it, due to biomineralization, with the organic/inorganic interface introducing certain unique material properties (Wang et al., 2011; Vepari and Kaplan, 2007). Apart from type I collagen, bone tissue also contains certain other noncollagenous proteins, such as osteopontin, bone sialoprotein (BSP), and osteocalcin, etc. (Venugopal et al., 2010). These noncollageneous proteins help in modulating the apatite crystal growth and serve as key markers to assess the state of osteogenically primed progenitor cells (Midha et al., 2016). Osteopontin, an early stage marker, helps in initiating the nucleation of apatite, while osteocalcin and BSP help in regressing the crystal growth (Gupta et al., 2016a). Depending on its location and biological functionality, natural bone can adjust the required mechanical properties by finetuning the proportion of components, crystallinity, and porosity (Gu et al., 2015). Long bones consist of diaphysis, epiphysis, and epiphyseal plates or growth plates (Olszta et al., 2007). Diaphysis, made primarily by cortical or compact bone, is a tightly knit, organized architecture consisting of osteons and harvesian canals that exhibit limited porosity (Porter et al., 2009). Small osteonic canals radiate parallel to the long axis of the bone, which constitutes the vasculature. Epiphysis contains

12.3 Rising Needs for Bone Grafts

spongy or trabecular bone, having pores of variable sizes. The trabculae exhibits haphazard arrangement, unlike the cortial bone. This enables the spongy bone to realign in the direction of stress changes (Boyle and Kim, 2011).

12.2 GLOBAL PERSPECTIVE ON ORTHOPAEDIC TRAUMA MANAGEMENT Since the Second World War, world population has been steadily rising, along with average life expectancy. This has led to an increase in the percentage of population vulnerable to musculoskeletal disorders, which in turn, implies an increasing number of contenders for orthopaedic tissue repair and/or transplants. Apart from ageing, the demand for bone grafts has also gone up in lieu of increasing number of traumas. Now, as a case in point, WHO predicts that by 2020, worldwide fatalities and injuries derived from road accidents would rise by 114% in 15 Asia-Pacific countries, including India (Cameron, 2004). In terms of tissue transplants undertaken globally, bone transplants rank second after blood transfusions. According to a survey conducted by Frost and Sullivan, the global orthopaedic market is expected to grow at a compound annual growth rate of 26.7% (Morgan, 2004). The number of orthopaedic surgeries is on the rise and it is predicted to double globally (Browne et al., 2009). The global market for orthopaedic implant product segments in a study done with respect to regional markets in United States, Europe, Asia-Pacific and rest of the world is projected to reach US$ 46.5 billion by 2017 from an estimated US$ 21.1 billion in 2007, growing by a compound annual growth rate of 8.2% between 2007 and 2017 (Iordache et al., 2013). This is indicative of the growing increase in focus towards bone tissue engineering (BTE) globally for the quest of finding better materials and strategies to aid in orthopaedic care and management.

12.3 RISING NEEDS FOR BONE GRAFTS Major trauma to bone tissue, genetic defects, age-related defects, such as osteoporosis, osteotomies, nonunion during fracture setting, and other critical conditions are usually addressed using bone grafts. In the US alone, over half a million patients are treated with orthopaedic grafts annually. Globally, the market for bone graft substitutes was valued at $1.9 billion in 2010, with expectations that it will reach $3.3 billion by 2017 (Che Senik et al., 2013). Some commonly used bone graft substitutes include allografts, coralline, ceramics, growth factorloaded scaffolds, tissue-engineered grafts, and processed allogenic and xenogenic grafts.

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12.4 IDEAL SCAFFOLD FOR BONE TISSUE ENGINEERING The ideal scaffold for BTE acts as a template for replicating tissue structure, while supporting, organizing, and reinforcing tissue regeneration and/or restoration (Swetha et al., 2010). The ideal scaffold should be able to flawlessly mimic natural extracellular matrix (ECM) at both micro and macro levels (Holzwarth and Ma, 2011; Kasoju and Bora, 2012). Numerous researchers have investigated the desirable properties required to design an optimal scaffold for BTE (Rockwood et al., 2011; Holzwarth and Ma, 2011; Farokhi et al., 2013; Mitra et al., 2013). Some of the crucial criteria prescribed are: • • • • • • • • •

biocompatibility and biodegradability bioresorbability commensurate mechanical strength osteoconductivity and osteoinductivity osteointegrative porosity and interconnectivity ability to carry and deliver osteoprogenitor cells ability to carry and deliver growth factors in a controlled manner decomposition products of scaffolds should not be toxic or cause any immunoreactivity and safely eliminated or assimilated by the body (bioresorbability).

Osteoconductivity is the ability of a scaffold to support osteoblast attachment and growth, while osteoinductivity is the ability of a scaffold to induce differentiation of stem cells into osteoblasts (He et al., 2013). Additional characteristics that need consideration include mechanical properties (like strength), surface properties such as roughness and cell attachment sites, hydrophilicity, surface stability (Rockwood et al., 2011; Wu et al., 2010). Porosity and pore interconnectivity of a scaffold allows cells penetration into the structure in order to develop threedimensional (3D) tissue constructs, along with unhindered transport of oxygen, nutrients, body fluid and blood to support cell metabolism and growth. It also aids in removal of metabolic wastes (Wenk et al., 2009; Mou et al., 2013). One of the important prerequisites of an ideal bone graft is its load-bearing ability. The cortical bone, owing to its compact nature, has been found to have a tensile yield strength of 78151 MPa and compressive yield strength of 131224 MPa, exhibiting Young’s modulus of 1720 GPa along the longitudinal axis and 613 GPa along the transverse axis (Porter et al., 2009). The trabecular bone, due to its anisotropic nature, exhibits Young’s moduli about 50100 MPa (Porter et al., 2009). The high degree of variance is attributed to the divergent bone mineral density and porosity, noticed in different people (Burghardt et al., 2010). Consequently, mechanical resilience becomes the foremost aspect for any biomaterial when probing its suitability for use in BTE applications.

12.5 Lacunae of Current Materials and Practices

By controlling the scaffold’s microstructure, various properties like degradation time, scaffold strength, and ability to differentiate stem cells in to specific lineage, can be influenced (Makaya et al., 2009). Wang et al. (2008) studied the in vivo degradation of silk fibroin (SF) in rats, processed by either aqueous or organic solvents. This study concluded that SF processed in water degrades faster (26 months) compared to SF processed in organic solvents (up to 1 year). Tissue invasion was more abundant and uniform in the aqueous scaffolds. Macroporous 3D scaffolds of SF, with pore sizes over 900 μm, prepared by aqueous method, produces scaffolds with better performance in vitro than scaffolds made by processing in an organic solvent (Meinel et al., 2004). This could be due to the difference in pore surface topology produced by the two different solvents. Accordingly, choice of solvent can be made based upon whether the scaffold needs to serve a short term drug release usage or as the support for a long term bone regeneration process. Meinel et al. (2006a) used an interesting approach to restore a bone defect, using a stack of five silk scaffolds with slightly different properties in order to better approximate the bone structure. Of these scaffolds, a layer of SF scaffold had mesenchymal stem cells (MSCs) that had been differentiated to osteogenic lineage with 5-week culture in spinner flasks; a layer of SF scaffold had undifferentiated MSCs. This integrated approach produced better results than using just SF scaffolds without stem cells. Porosity may be introduced in a scaffold by using particulate leaching, gas foaming, or lyophilization. Particulate leaching is, of course, one of the simplest methods and NaCl is the most widely used porogen. Freeze-drying with porogens added is often used to form porous SF scaffolds. By managing the porogen particle size, a relatively tight control may be exercised over pore size. With pore size, the mechanical strength of the scaffold can be controlled. Partial solubulization of the particulates can produce rougher internal surface of pores that improves cell attachment (Correia et al., 2012).

12.5 LACUNAE OF CURRENT MATERIALS AND PRACTICES FOR BONE REGENERATION Grafts required during bone repair are often taken from another location of the same patient (autograft) or from a donor (allograft) (Bhumiratana et al., 2011). Autografts carry a risk of morbidity at the donor site, a secondary surgery, and have limited availability; allografts are susceptible to immunoreaction and infection (Diab et al., 2012). Overall, allogeneic bone grafts are faced with high failure rates: 16%50% (Nisbet et al., 2008). Introduction of growth factors can remarkably improve the bone reconstruction process. However, delivery of such growth factors possesses a challenge. Systemic or direct injection of such growth factors is constrained, due to slow penetration

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from tissue sites, short half-life, and rapid dissemination from the injection site (Diab et al., 2012). First generation tissue-engineered constructs currently available for clinical use consist of a combination of bone morphogenic proteins (BMPs) embedded in a matrix of collageneous material. However, collagen is costly, does not have enough mechanical strength, is limited in supply, and can add a risk of infection (Mou et al., 2013). Titanium and its alloys are also often used as orthopedic (including dental) implants, due to their strength and biocompatibility (Cao and Mao, 2007). Additionally ß-type titanium has better stress-shielding behaviors that matches the Young’s moduli of native bone, prevents bone atrophy and enhances bone remodeling (Niinomi and Nakai, 2011). However, such metallic implants are plagued by lack of cell adhesion, immunoreactivity, limited osteoconductivity, and infections, leading to eventual implant failure (Naskar et al., 2014). Failure would mean subsequent exacerbated surgical interventions. Attempted bioengineered solutions may often lack the required strength to take the place of a skeletal tissue. There is also the issue of sustaining angiogenesis along with osteogenesis, without which typical tissue-engineered constructs are not able to get bone thickness more than 1 mm (Meinel et al., 2006a). Scaffolds made of certain polymers that have been proposed for BTE have the complication of acidic hydrolysis products that can cause inflammation (Tungtasana et al., 2010). While diverse materials such as ceramics, bioglasses, polymers, etc., have been tried for BTE, none of the currently available solutions have all the required positive attributes of a scaffold. Drawbacks of current practices refocus the attention on tissue engineering. The ambition is that tissue engineering can provide a solution that overcomes the limitations of current solutions by balancing a composition of scaffold material, osteoprogenitor cells, and required growth factors.

12.6 BIODERIVED AND SYNTHETIC MATERIALS FOR BONE TISSUE ENGINEERING An innate advantage of natural materials is their biocompatibility and biodegradability. Several bioderived materials are biopolymers: lignocellulose, polysaccharides (alginate, chitosan, hylauronic acid derivatives, starch), and proteins (collagen, fibrin, soy, silk) (Swetha et al., 2010; Zhu et al., 2012a). With proper selection of material properties, such biodegradable polymers could be engineered to match their degradation rate with neo-tissue formation. Apart from bioderived polymers, certain synthetic biodegradable synthetic polymers have also been widely used, e.g., poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and their copolymer (PLGA) (Zhu et al., 2012a). Moreover, US Food and Drug Administration (FDA)- approval has also been granted to a few synthetic polymers, including polycaprolactone (PCL) (Hardy and Scheibel, 2010; Roohani-Esfahani et al., 2012). In order to combine the advantageous properties of both synthetic and

12.8 Various Sources of Silk

natural biopolymers, researchers have also focused on hybrid scaffolds, having beneficial properties of both (Holzwarth and Ma, 2011). A hybrid structure could use the coating of collagen (a natural ECM protein) over synthetic materials to improve cell viability, proliferation, and differentiation (Thibault et al., 2013). For BTE purposes, a mineral component is further added to the above composite to improve mechanical properties. The chosen mineral is often HAp (Ca10(PO4)6(OH)2), due to its close resemblance to the natural mineral phase of bones and osteoconductive nature (Li et al., 2006). Combining with polymers compensates for HAp’s lack of toughness, brittle nature, and tendency to drift from implant sites (Li et al., 2008). Other biocompatible options, such as calcium phosphate variants (tri-calcium phosphate) or bioglass have also been used (Holzwarth and Ma, 2011).

12.7 CONTEXT AND STRUCTURE OF THE CHAPTER Tissue engineering is a vibrant, multidisciplinary, and fast-growing field of research. Every year, thousands of research works are published in this field. Using Google Scholar, .6000 search results were obtained for “BTE” and more than 1000 results were obtained for “BTE 1 silk,” just for 2015. Periodic reviews of works can have an immense importance in such a vigorously growing field. As discussed in Section 12.5, materials with bioderived sources are showing considerable promise for tissue engineering research of the future. In this regard, silk is an abundantly available biomaterial and its reputation for remarkable strength transcends the scientific domain into popular culture. Consequently, it is not surprising that researchers perceive silk to have a significant role to play in BTE. This work looks at a few of the more interesting studies in the domain of skeletal tissue engineering from the past decade and half. We try to organize the chapter around a viewpoint of materials, additives, and fabrication approaches. Fig. 12.1 presents an overview of this chapter’s architecture; along with inter linkages between the different sections. The Introduction focuses on the need for BTE, what could be considered ideal scaffolds, and how bioderived sources may have the answer to the lacunae of present day materials. Section 12.3 briefly discusses antecedents and favorable characteristics of silk sources. Section 12.4 summarizes fabrication of various silk scaffolds and their relevance with BTE. Section 12.5 illustrates certain recent trends being followed in this particular field. We conclude with a view towards how interdisciplinary cooperation should help shape the future of BTE.

12.8 VARIOUS SOURCES OF SILK Silk is a naturally occurring structural protein of high molecular weight (B300 kDa) (Kim et al., 2010). In a generic sense, silks are spun in the form of

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Bioderived sources to overcome current limitations

Rising needs for bone tissue engineering

Silk: antecedents and benignant aspects

Future perspective

Silk as a biomaterial for bone tissue engineering

Additives in silk scaffolds Biomineralization

Beyond mulberry silk

The spectrum of silk use in bone tissue engineering

Fabricating silk scaffolds

Fabrication techniques Scaffold architecture

FIGURE 12.1 A representation of the article architecture and the interconnection between different sections.

fibers by larvae of certain lepidoptera (like silkworms), by certain arachnids (like spiders and some scorpions), and by some flies and mites (Macintosh et al., 2008; Altman et al., 2003). Silk fibers have a standing reputation for their impressive mechanical strength and exceptional strength to weight ratio, superior to most man-made polymers (Kundu et al., 2013). Being a naturally occurring polymer of high structural strength, silks become an obvious candidate for BTE. As an added advantage, SF can also guide calcium-phosphate formation (Marelli et al., 2012). Commercially produced silk fibers come from Lepidopteron insects of Bombycidae or Saturniidae families. Bombyx mori, commonly known as mulberry silkworm, is the most important member of Bombycidae family and has been domesticated by human beings for a long time. From the Saturniidae family, two species are important: the Indian tropical tasar silkworm Antheraea mylitta, Antheraea assamensis (endemic to Assam, India) and the Chinese temperate tasar silkworm, Antheraea pernyi. Spider silk, though not of any current commercial significance, is also lightweight and has superior mechanical properties. Silk from B. mori, among silkworms, and the genus Nephila (golden orb weavers) among spiders (Macintosh et al., 2008) are most widely explored for biomedical research and tissue engineering applications.

12.9 Silk From Silkworms

12.9 SILK FROM SILKWORMS Silkworm silk is broadly classified into two categories, based on the eating habits of species: mulberry silk and nonmulberry silk. B. mori is the source for mulberry silk, while nonmulberry silk is obtained from multiple species like A. mylitta, A. assamensis, A. pernyi, and Philosamia ricini or Samia cynthia ricini. The silk from silkworms has two major components, fibroin and sericin, which are made of 18 different amino acids. Because the amino acid arrangement varies depending on the silkworm species (Andiappan et al., 2013), the different silk fibers have diverse properties. SF structure resembles that of type I collagen in some manners (Gu et al., 2011). Fibroin is fiber-like protein and constitutes the core of silk fibers, whereas sericin is a water soluble, glue-like class of protein, that coats the fibroin fibers (Mobini et al., 2013; Li et al., 2014). SF from mulberry silk worm (B. mori) have a heavy chain component (B350 kDa) and a light chain component (B25 kDa). These two chains are connected via a disulfide link (Ming et al., 2014). The heavy chain has a hydrophobic nature, whereas the light chain is more hydrophilic (Hardy and Scheibel, 2010). Silk protein may exist in the following arrangements: random coil, α helix (also called Silk I), and β sheet (also called Silk II). Fibroin of silkworm silk is built of 17 amino acids. Repeating sequences of alanine and glycine (nonpolar amino acids) in the fibroin are responsible for forming the antiparallel β sheet crystal. This crystalline structure is the primary reason for silk’s mechanical strength (Li et al., 2014). The β sheets generate cross-linking between the protein chains via strong hydrogen bonds and van der Waals interaction, thus dramatically improving the fiber strength (Li et al., 2014). Crystallinity of the β sheet structure can be induced and modulated through chemical or physical treatments. The US FDA, approved silk for use in “soft tissue repair,” classifying silk as nonabsorbable (Chao et al., 2010; Kim et al., 2010), and the US Pharmacopoeia classifies silk as a nondegradable implant (Kim et al., 2010; Li et al., 2014). Considering their favorable record when used in biological environments, silkbased systems become an obvious choice for use as drug delivery platforms (Farokhi et al., 2014). Although the silk varieties of tasar, eri, muga etc. are referred to as wild silk, they have been used for fabric production, for a considerable period, in certain regions of the world. However, none of these varieties are farmed in the same scale as mulberry silkworms. Most of the research conducted with silkworm silk has used mulberry silk in recent years, a growing trend has been seen regarding studies conducted with nonmulberry silk varieties (Kasoju and Bora, 2012; Gupta et al., 2016b). Since the silk of different silkworms differs, due to the sequencing of amino acids, it may be expected that this leads to certain differences in biocompatibility features. The rise of interest may be ascribed to certain superior

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characteristics of the nonmulberry varieties. Unlike mulberry silk, the SF from nonmulberry silk lacks the light chain and is mainly homodimers composed of heavy chain (B220240 KDa) (Kundu et al., 2012). Moreover, the mulberry SF H-chain consists of poly(glycine-alanine) repeats, whereas the nonmulberry counterpart consists of poly(alanine) repeats. This has causes remarkable difference in the tensile properties noticed between the silk varieties (Malay et al., 2016; Gupta et al., 2015). The major difference between silk from B. mori and the silk of nonmulberry varieties, like A. mylietta., A. assamensis and A. pernyi, is the presence of the tripeptide sequence, arginine-glycine-aspartic acid (RGD) in nonmulberry silks (Vepari and Kaplan, 2007; Naskar et al., 2014; Gupta et al., 2015). This sequence provides recognition sites for integrin-mediated cell adhesion, with several cell-types showing a preference for surfaces coated with RGD peptides (Maghdouri-White et al., 2016). Since B. mori silk does not have this inherent sequence, attempts have been made to artificially coat B. mori SF with RGD peptides (Macintosh et al., 2008). However, silk varieties with inherent RGD sequence have an obvious advantage over mulberry silk when it comes to promoting cell attachment and proliferation (He et al., 2012a). RGD sequence may also aid differentiation of cells into skeletal tissue components (Murphy and Kaplan, 2009).

12.10 SPIDER SILK Spider silk, like the silk from silkworm, primarily has glycine and alanine, but does not have a sericin coat (Macintosh et al., 2008; Kundu et al., 2014). Different spider species have evolved a variety of silks suitable for diverse functionalities—from building a web to capturing prey—by altering content and sequencing of amino acids (Altman et al., 2003). Like SF, spidroin is also endowed with impressive strength. One of the most extensively studied spider silks is the major ampullate spidroin protein (dragline silk) of the species Nephila clavipes, with molar mass of B250350 kDa (Gomes et al., 2011). Its tensile strength is in line with that of Kevlar (Hardy and Scheibel, 2010). Collecting spidroin at large scale for biomedical purposes is a difficulty, unless a method to artificially mimic and fabricate spidroin turned up recently. Most of the subsequent discussion focuses on silkworm fibroins, unless specified otherwise.

12.11 BENIGN ASPECTS OF SILK FOR BONE TISSUE ENGINEERING The natural strength of silk fibers makes them of obvious interest for BTE scaffolds. Additionally, silk-based scaffolds have shown to be biocompatible in different forms of use and with different tissues, without any haemotoxic effect

12.11 Benign Aspects of Silk for Bone Tissue Engineering

(PA¸SCu et al., 2013). Silk scaffolds also offer good permeability to oxygen and other nutrients (Lin et al., 2008). Because of the well-established silk industry, sourcing of silkworm fibroin for biomedical research and use is also economical. Mechanical properties of silk are a comprehensive mixture of high modulus, breaking strength, elongation at break, which together make it tough and ductile, apart from just having a high strength. The strength-to-mass ratio of silk could be up to 10 times that of steel (Kundu et al., 2013). The ultimate tensile strength of silk was estimated to be between 61 and 690 MPa, far superior than the existing natural polymers (collagen B0.97.4 MPa) and synthetic polymers (PLA B 2850 MPa) used for BTE application (Mandal et al., 2012). While silk normally takes a year or more to become absorbed in vivo, modifications to processing methods can cut down the degradation time by half or more (Kim et al., 2010; Li et al., 2014). Slow degradation of silk scaffolds can be helpful in bone tissue repair so the scaffold can support the structure until native tissue grows strong enough to take its place (Mandal and Kundu, 2009). B. mori SF can be dissolved in aqueous solutions of lithium thioisocyanate, lithium bromide etc., at temperatures of B70 C and then dialyzed into water, to obtain the regenerated form of SF (Murphy and Kaplan, 2009). Regenerated SF offers better moldability and surface manipulation options and, accordingly, most use of SF is in the regenerated form (Kundu et al., 2013). Regenerated SF is also stronger and tougher (Jiang et al., 2013). Hydrophilic pendent carboxyl groups on regenerated SF also serve as suitable nucleation sites for HAp during biomimetic mineralization (Cao et al., 2013). Silk can also be processed into various formats: hydrogels, coatings, films, porous structures, nano and micro fibrous matrices etc. Owing to its ability to be moulded into diverse forms, silk has been used successfully used in bone, cartilage, tendon, and ligament tissue engineering research (Yan et al., 2012). As the silk protein structure is better understood, fabrication methodology would also improve and become fine-tuned to specific tissue engineering applications. Silk topology also makes it suitable to serve as the substrate for mineralization process. The β sheet content of SF can be controlled by manual tuning of induction process (chemical and/or physical), thus giving a means to control crystallinity of the fiber. By controlling crystallinity, mechanical properties and rate of degradation can be manipulated (Wray et al., 2012). Advantageous properties of SF have been further improved using chemical modifications (Vepari and Kaplan, 2007). The diverse amino-acid side chains of silk polymer can be chemically altered, using reactions such as carbadiimide coupling, to facilitate surface modifications such as coupling of growth factors, RGD sequences, other polymers, etc. (Murphy and Kaplan, 2009). Such modifications can alter features of a silk scaffold: its hydrophilicity, β sheet structure and content, surface morphology, and cell binding sites. It has often been noticed that scaffolds made from silk do not have strength in the same league as native silk fibers. However, this may be overcome by

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modifying the silk regeneration process (Kundu et al., 2013). Chemical modifications can also be used to improve these properties of silk as extension, elasticity, and strain hardening (Maghdouri-White et al., 2016). The ability of silk to promote apatite deposition, forming a silk-apatite composite, is one of the beneficial aspects of using SF for BTE (Jang et al., 2009). Still, a few lacunae exist that limit its application for BTE application. Silk in itself does not have optimal osteoconductivity (Wu et al., 2011) and B. mori silk lacks any inherent bioactive peptides. If SF is accidentally used with sericin residue, it may trigger antigenic reaction (Meinel et al., 2005b). Low molecular weight degradation products of silk scaffolds may not be completely congenial to surrounding tissue (Kim et al., 2010). These are a few minor drawbacks of native silk fiber and, as further discussions will show, most of these have been overcome by researchers, with relatively straightforward stratagems.

12.12 PROCESSING SILK INTO VARIOUS FORMATS As discussed earlier, a major attraction of using silk for BTE is the ease and flexibility in processing of silk-based scaffolds. A variety of silk-based scaffolds have been investigated for their use in BTE, spanning from micro to nanoscale. Scaffold properties, such as elasticity, toughness, etc., can be modified to obtain osteogenic differentiation of stem cells (Kundu and Kundu, 2010). Scaffold degradation rates can be modulated by following methods: aqueous processing and ethanol treatment for fast degrading and hexafluoro isopropanol processing, and autoclaving for slow degradation. In this section, we discuss some of the major processing strategies for silk-based scaffolds and their desired effects.

12.12.1 PARTICULATE LEACHING Particulate leaching method is widely used to form and control the porous structure of a scaffold with good interconnectivity (Mou et al., 2013). Particulate leaching can be used as a one-step procedure to fabricate scaffolds with just silk or composites of SF with other polymers and/or minerals (Mou et al., 2013). Porogen size has a direct relation with the pore sizes formed in the scaffold; 250500 μm porogens are often used (Makaya et al., 2009). Using NaCl as porogen, 3D porous SF scaffolds with controllable porosity and pore size can be prepared via an all-aqueous process (Kim et al., 2005). Makaya et al. (2009) used particulate leaching to produce 3D porous scaffolds for cartilage tissue engineering, by using salt and sugar as porogens. Salt-leached scaffolds produced harder structureS that were digested quickly by a protease enzyme, while the sugar-leached scaffolds produced much more stable scaffolds.

12.12 Processing Silk Into Various Formats

12.12.2 NANOFIBROUS SCAFFOLDS USING ELECTROSPINNING Nanofibrous scaffolds, due to their nanoscale architecture, have good mechanical properties, can provide large surface area-to-volume ratio, and have good porosity and interconnected pores. Such scaffolds can make a convincing reproduction of natural ECM (Kim et al., 2014a). The surface features of nanofibrous scaffolds can also aid and advance cell attachment and proliferation (Nisbet et al., 2008). From its initial use for textile industry in the early 1900s (Holzwarth and Ma, 2011), electrospinning has become the most extensively studied technique for producing nanofibrous scaffolds intended for tissue engineering use (Farokhi et al., 2013). Electrospun nanofibrous mats of SF have shown good biocompatibility, while showing no evidence of any inflammatory properties (Bhardwaj and Kundu, 2010). Electrospun scaffolds can have nano- to submicron scale fibers, interspersed with microscale pores, in close resemblance to bone ECM structure (Bhardwaj and Kundu, 2010; Wei et al., 2011; Shanmugavel et al., 2014). As per need, a wide variety of polymers—natural, synthetic, or recombinant proteins— can be electrospun (Shanmugavel et al., 2014; Li et al., 2006). Electrospinning also conveniently produces composite scaffolds from a mix of different polymers or nanoscale mineral particles, cytokines mixed with the polymer solution (Li et al., 2006; Sheikh et al., 2013). The as-spun SF nanofibrous scaffolds primarily contain Silk I, and consequently are treated with methanol to convert Silk I into the crystalline Silk II structure. This conversion proceeds from the surface of fibers to their core and so; the fibers have a hard crust and a soft core (Wang et al., 2004). Electrospun scaffolds of just-regenerated SF have also been shown to maintain growth and differentiation of bone marrow stromal cells (BMSCs) in vitro (Bhardwaj and Kundu, 2010). Ming and Zuo (2012) prepared electrospun composite fibers of SF and HAp by spinning the SF with varying proportions of HAp nanoparticles (upto 30%). These matrices showed an increase in fiber diameter from 92 to 582 nm, and an increasing trend of mechanical strength as HAp percentage increased from 0 to 30. This work also showed that cospinning of HAp and SF does not alter crystal structure of SF. Alternate soaking in calcium and phosphate solutions have also been used to obtain uniform distribution of HAp upon the nanofibrous SF scaffolds (Wei et al., 2011). A major obstacle for wider adoption of electrospinning in BTE is that electrospinning is often able to produce only a fibrous mat, while BTE applications require a larger 3D scaffold (Holzwarth and Ma, 2011). Three-dimensional matrices can be obtained by dramatically increasing the electrospinning time interval. This is one aspect of electrospun scaffolds that needs deeper examination, and here electrospinning could benefit through hybridization with some other processing methods. Three-dimensional nanofibrous spongy scaffolds have been prepared by combining electrospinning with particulate leaching (NaCl porogen)

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and these scaffolds performed better in terms of cell adhesion and growth compared to 2D electrospun mat (Park et al., 2010). These 3D scaffolds can provide an alternative for larger bone defects.

12.12.3 BIOPATTERNING Several plants have structures that resemble the morphology of bone, i.e., the anisotropy, micro- and macroporous architecture, etc. Qian et al. (2014) used this idea to create negative NaCl moulds of cane, which they then later used to prepare porous SF scaffolds. These scaffolds were able to replicate the original porous structure of the cane. Wray et al. (2012) designed and produced linear wire arrays around which a silk scaffold could be built, thus providing precisely spaced and dimensioncontrolled hollow channels in the scaffold. These scaffolds were successfully used to localize MSCs in the bulk of the scaffold and human amniotic epithelial cells in the hollow channels. Such scaffolds would be very conducive to vascularization of tissue.

12.12.4 KNITTED SCAFFOLDS Knitting (instead of braiding) SF fibers to form a scaffold means that the porous architecture of the scaffold would sustain any mechanical tension, promising better nutrient and cell diffusion in vivo. He et al. (2013) combined knitted SF with Hap-coated silk sponge to obtain such a hybrid scaffold which supports and propagates osteogenic differentiation of BMSCs.

12.12.5 FREEZE-DRYING Freeze-drying (lyophilization) is a widely used technique to fabricate porous scaffolds. It is often combined with some form of particulate-leaching for ensuring porosity. Freeze-drying preserves the β sheet structure of SF (Yan et al., 2012). The percentage of SF can be varied to control pore size, interconnectivity of pores, and strength of scaffolds. Higher SF concentration results in reduced pore size and superior mechanical strength. Lin et al. (2008) reduced the processing steps to produce porous 3D scaffolds by freeze-drying SF solution with methanol, followed by glutaraldehyde cross-linking. An aqueous solution of SF containing phosphate was freeze-dried followed by slow infiltration of a warm (37 C) aqueous calcium chloride solution into the frozen sample. Freeze-drying can also yield a composite scaffold of SF and natural silk fibers, where the silk fibers serve the role of a reinforcing agent and the scaffolds have enhanced

12.12 Processing Silk Into Various Formats

compressive strength (Mobini et al., 2013). These reinforced scaffolds successfully sustained stem cell growth and osteogenic differentiation.

12.12.6 SILK MICROPARTICLES AND MICROFIBERS AS REINFORCEMENTS Silk, used in the form of microparticles, can act as a strengthening element for a polymer scaffold, specifically for silk scaffolds too (Kundu et al., 2013). Milling the silk particles can improve their ability as a reinforcing agent (Mandal et al., 2012). While silk fibers possess a remarkable tensile strength, reinforcing with silk microparticles can improve the compressive strength of SF scaffolds. The strength and calcium adsorption proportionately increases as the quantity of microparticles is increased (Rockwood et al., 2011). Similar reinforcement behavior has also been observed from short silk fibers (of 2 mm dimension), when used with a SF/HAp composite (Jin et al., 2014). These reinforced scaffolds acted as a suitable medium for delivering human placentaderived mesenchymal stem cells in vivo, with resulting good repair of bone defects. Alkali treatment of SF particles can release surface fibrils and thus enhance interaction at the surface with mineral particles, like HAp. This leads to improved hardness of the composite, as well as uniform pore distribution and sizing (Wang et al., 2005). The use of alkali can turn the composite into a gel form, which could be suitably injected at bone defect sites. Silk microparticles can be used as a controlled delivery medium for growth factor proteins, such as BMP-2, embedded to support and enhance osteogenic differentiation of stem cells and consequential bone repair (Bessa et al., 2010).

12.12.7 HYDROGELS Hydrogels have a 3D polymeric network of interconnected pores that retains a significant volume of water. When forming SF-HAp hydrogels, the proportion of HAp influences β sheet formation in SF and a greater quantity of HAp leads to smaller pore size, but higher pore density (Kim et al., 2014b).

12.12.8 COMPUTER-AIDED FABRICATION OF SF SCAFFOLDS (3D PRINTING) The precise control of computer-aided design (CAD) and manufacturing (CAM) holds an obvious attraction for tissue engineering. These processes could lead to scaffolds designed fit for their purpose, while also saving time. Sun et al. (2012)

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used 3D printing to print a scaffold with graded pore spacing between 200 and 700 μm, by utilizing an ink of SF and HAp particles.

12.13 SILK COMPOSITES FOR BONE TISSUE ENGINEERING Polymers display rapid strength deterioration and, consequently, they are composited with additives such as ceramics to enhance the mechanical properties. Silk is amenable to such compositing; high-strength silk-based composites have been explored in this regard. Modifying just the surface of a scaffold has certain advantages, since it interacts directly with cells and surrounding tissue. For example, plasma treatment can effect changes in the depth range of a few nanometers. Baek et al. (2008) observed that plasma treatment slightly increased roughness and pore size of a 3D electrospun, porous, nanofibrous scaffold of PCL and Hap, while significantly improving the ability of human articular chondrocytes to attach and proliferate on these surfaces. Nonmulberry SF-grafted PCL nanofibrous scaffolds (Bhattacharjee et al., 2015) can be coated with HAp by electrodeposition to augment the osteoconductivity by inclusion of bioceramics. The coated scaffold retains the mechanical resilience of silk and is not impacted by the brittle nature of the ceramic. Such scaffolds, combined with growth factors, can enhance bone formation in vitro as the mineralized surface of the scaffolds provides an osteoconductive environment for the osteoblast (Bhattacharjee et al., 2016b).

12.13.1 HYDROXYAPATITE HAp is a favored choice when it comes to embedding a mineral component in silk-based scaffolds. HAp that has a calcium deficiency (Ca9(HPO4)(PO4)5OH, CDHA), in stoichiometric terms, serves a closer resemblance to biological apatite in terms of structural and chemical properties; as such, it has been used as a successful implant in orthopaedic and dental surgeries (Choi et al., 2012). HAp crystals in natural bone have their c-axis parallel to the organic substrate’s long axis. That is how HAp crystals formed on SF, by biomimetic mineralization, also orient themselves (Niu et al., 2010). HAp in natural bone also has a low crystallinity and that feature has also been replicated for forming better HAp/SF composite scaffolds (Shi et al., 2013). HAp composites with SF have been prepared by biomimetic mineralization, mechano-chemical processing, coprecipitation, freezedrying, etc. (Niu et al., 2010; Chen et al., 2013, 2014). The use of chemically disintegrated SF in HAp coprecipitation processes can improve HAp crystal alignment and reduce the size of crystals formed (Nemoto et al., 2004). HAp has been included in freeze-dried composite hydrogels of SF and PVA (Zhang et al., 2012).

12.13 Silk Composites for Bone Tissue Engineering

Sponge-like matrices of SF with embedded HAp have been used for growing MSCs (Bhumiratana et al., 2011). An attempt towards fabricating a multilayered matrix has been made by compressing layers of SF and Hap-deposited SF films together to create a matrix of improved Young’s modulus (Kino et al., 2007). Na2SiO3 has also been used with SF to aid formation of HAp particles. It was found that Na2SiO3 concentration over 0.007% (w/v) accelerates HAp deposit. It was also observed that lower concentrations of Na2SiO3 yield a sheet-like deposit of Hap, whereas increasing concentration leads to rod-like bundles of HAp (Li et al., 2008). A preparation using PCL film, coated with biphasic calcium phosphate as aggregation site for nanofibrous silk, results in a porous, tough, and nonbrittle scaffold, with high compressive strength (RoohaniEsfahani et al., 2012). The hydrophobic nature of PCL meant that the scaffold better preserved its strength in an aqueous medium. Most bone cements in current use are composites of HAp and wollastonite. Therefore, Zhao et al. (2011) formed a composite of HAp, SF, and wollastonite for bone repair that had higher compressive strength (85 MPa) and hardness than a standard SF/HAp composite, in addition to having good bioactivity. HAp coating on SF scaffolds usually do not affect the existing pore architecture of the scaffold (Jiang et al., 2013). Bonding of HAp onto SF surfaces may be improved via ionic interactions by using a grafted layer of poly[4-methacryloyloxyethyl trimellitate anhydride] (4-META); the resultant scaffold is suitable in terms of cell adhesion and proliferation (Bhattacharjee et al., 2016b). HAp/SF/collagen composites have been formed by in situ precipitation to form a matrix with uniform HAp distribution. Addition of collagen meant that these matrices supported cell proliferation better than matrices with only SF and HAp (Chen et al., 2014). Three-dimensional SF scaffolds, modified with HAp and gelatin from collagen, improved the surface of scaffolds in a manner so as to increase cell adhesion and proliferation (Moisenovich et al., 2014) and the porosity could be adjusted by modulating HAp levels (Mou et al., 2013). By mineralizing a net of SF, instead of the usual SF films, a porous composite of SF and HAp is formed with open pores. HAp inclusion improves hydrophilicity and cytocompatibility of the scaffold and provides a viable environment for attachment and growth of osteoblast. Embedded nanoceramics in a freeze-dried scaffold of SF improves bioactivity, strength, and surface roughness of the scaffold, while maintaining a desired level of porosity without the need for a porogen (Ghorbanian et al., 2013). Nanoparticles of calcium phosphate may also be included on an aqueous solution of SF to fabricate composite scaffolds, following salt leaching and lyophilization. The degree of porosity has an inverse relation with the proportion of HAp used (Yan et al., 2013). SF particles, prepared with the aid of polyethylene oxide, can be coated with calcium-deficient HAp crystals. This coating is better at a higher pH, due to more negative charge on the surface of silk particles (Choi et al., 2012). HAp-coated SF structures have also been tried

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for other cutaneous use and have proven to be suitable for adhesion and growth of osteoblast (Bhattacharjee et al., 2016a). One of the reasons for the success of SF/HAp composite scaffolds could be the strong bonding (due to molecular interactions or chemical bonds) between SF and HAp. The surface of silk polymer provides several groups that could serve as the nucleation and bonding site for HAp: amine (NH), carboxyl (COOH), carbonyl (CO), and hydroxyl (OH) (Wang et al., 2010b). Use of mineralized silk scaffolds may not be sufficient in itself for all sizes and nature of bone defects. The osteoconductive nature of HAp means that such mineralized scaffolds form an ideal substrate for delivery of stem cells and growth factors. It also enhances the osteogenic differentiation of stem cells, thus extending its applicability (Jiang et al., 2013) and new bone formation in vivo (Niu et al., 2010). HAp addition also tends to raise the surface roughness of scaffolds and improve mechanical stiffness and compressive strength (Bhumiratana et al., 2011; Wang et al., 2010a). Mineral deposition from stem cells can increase with the initial quantity of HAp present in an implant (Bhumiratana et al., 2011). Bone regeneration taking place via a scaffold of SF and HAp may be expected to regain the levels of original bone strength (Wang et al., 2010b; Shi et al., 2013). A pictorial representation of HAp deposition on SF is given in Fig. 12.2.

12.13.2 CLAY- OR SILICA-BASED ADDITIVES Silica has an osteoinductive nature. It can lead to enhancement of bone formation and may be sourced from clay montmorillonite or sodium silicate (Mieszawska et al., 2011). Since silica has a long history of use in fabrication processes other than those involving biomaterials, a considerable degree of control is available when it comes to morphology and size of silica particles (Mieszawska et al., 2010b).

12.13.3 BIOACTIVE GLASSES Glasses with a silica basis have good bioactivity and strong binding ability to bone (Mieszawska et al., 2010a). Moreover, bioactive glasses are reported to be the most osteoconductive material among the class of bioceramics. The soluble leachates from these bioactive glasses regulate the osteogenic differentiation of progenitor cells and stimulate faster bone regeneration. The biodegradation rate of bioactive glass (BG) can be controlled by altering its composition ratio, and thus be made suitable for different tissue engineering applications. Composite films of SF and nano-BG, prepared by solvent casting, have been tested in vitro for their compatibility with osteoblast growth and proliferation. Results suggested that nano-BG improves mineralization on SF scaffolds (Zhu et al., 2012a).

12.13 Silk Composites for Bone Tissue Engineering

FIGURE 12.2 Use of a bioreactor in fabricating in vivo implants.

Mesoporous bioactive glass (MBG) is a comparatively new introduction in the field of biomaterials and, due to its ordered nanochannel structure, it shows better bioactivity than standard BG in vitro (Wu et al., 2011). Silk-modified MBG scaffolds show uniform and well-connected pores, good ability for mineralization and drug delivery, excellent cell attachment, proliferation, and differentiation of MSCs. Silk introduction into MBG scaffolds enhances their mechanical strength (Wu et al., 2010). In turn, MBG inclusion in SF scaffolds improves their strength and mineralization potential, although .10 wt% inclusion of MBG in SF scaffolds leads to nonuniform distribution of the particles (Wu et al., 2011).

12.13.4 SILK INCLUSION IN OTHER SUBSTRATES Inclusions of SF can be used to improve the performance of substrates of other materials. Inclusion of silk microparticles can enhance both surface roughness and strength of a scaffold to the level of natural bone (Rockwood et al., 2011). Micropatterning of a composite of nano-HAp and SF on implants of titanium or titanium alloys can significantly improve their biocompatibility. Zhu et al.

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(2012b) deposited nanocrystals of HAp/SF onto Ti alloy scaffold using electrostatic spray deposition, guided by precisely designed microtemplates. Silk coatings have also been used on biphasic calcium phosphate scaffolds to improve the mechanical properties and osteogenic response of cultured MSCs (Li et al., 2013). Sericin from A. mylitta, both singularly, and in combination with RGD sequence, has been used to functionalize the surface of titanium implants and this functionalization improved adhesion and proliferation of osteoblasts in vitro (Nayak et al., 2013). This is due to the bioactivity and biocompatibility of silk sericin. When fibroin from B. mori and A. mylitta were both used for surface modification of titanium scaffolds, performance from A. mylitta SF-coated substrates was found to be better (Naskar et al., 2014).

12.13.5 OTHER MISCELLANEOUS ADDITIVES FOR SILK SCAFFOLDS A subsidiary advantage of silk as a material, is its ability to be blended with other materials of biological or synthetic origin. Although silk in itself has several inherent advantageous properties, combining it with other materials can improve some of these advantageous properties. Biodegradable polymers used in combination with SF for BTE (synthetic in origin or natural in origin), have often been combined with SF to form a composite organic matrix. Such a composite matrix could serve as the template for HAp, with the calcium ions forming strong bonds with the amino group of chitosan or the amide of SF, and enhance compressive strength of the ceramic composite formed. The biological molecules most often combined with silk matrices are proteins, such as growth factors/cytokines, or growth factorcoding genes (Ji et al., 2011). Strong bonding between SF and HAp can improve the ductility of ceramic composite, and chitosan may be added to take advantage of its adhesive nature (Lima et al., 2013). Choukroun platelet-rich fibrin, prepared from the patient, has also been combined with acid pretreated SF to improve in vivo bone healing (Lee et al., 2010).

12.14 BEYOND THE CLASSIC MULBERRY SILK FIBROIN 12.14.1 NONMULBERRY SILK As discussed in Section 3.1, SF from nonmulberry silkworms (NSF) has certain inherent advantages for tissue engineering applications due to differences in its structure. However, mulberry silk worm has been domesticated by human beings for a few millennia and, as such, is widely spread across the world and available to different researchers. Conversely, nonmulberry species are mostly regarded as wild varieties, although cultivation of these species is being carried out in certain

12.14 Beyond the Classic Mulberry Silk Fibroin

countries like China, India, and Thailand. Due to their limited availability, and in spite of their inane advantages, they have been explored in only a limited number of works. We discuss some such studies here. Studies by Mandal and Kundu (2009) showed that 3D scaffolds of SF from A. mylitta had better porosity and similar inflammatory or immunogenic response, compared to 3D SF scaffolds of B. mori. Greater mechanical strength of A. mylitta SF meant that these scaffolds would induce osteogenesis while B. mori SF scaffolds supported adipogenesis. Better porosity led to greater cell penetration and attachment into the scaffolds of nonmulberry SF. Ren et al. (2007) successfully used SF from A. pernyi as a regulator for HAp mineralization, creating composites with similarity of microstructure to natural bone. He et al. (2012a) fabricated a composite of HAp, chitosan, and SF from A. pernyi, which presented optimal properties for BTE. The HAp formed by coprecipitation method using the NSF had low crystallinity, was calcium deficient, with orientation along c-axis, as has also been observed when using mulberry SF. Favorable results have also been observed for mineralized, electrospun matrices of P. ricini (eri) SF, with MSCs culture and osteogenic differentiation of the stem cells (Andiappan et al., 2013). Tungtasana et al. (2010) used scaffolds made from fibroin of Thai silk worm, for a 12-week duration in vivo study in rats, without any significant irritation or inflammation. Because it lacks an inherent RGD sequence, SF from B. mori has been chemically modified with RGD cell-binding domains to improve cell attachment, osteogenesis of MSCs, and eventually, bone formation (Vepari and Kaplan, 2007). Conversely, Nagano et al. (2012) expressed a recombinant, silk-like protein in Escherichia coli that was designed for its ability to bind calcium ions, with a view to acting as an aid during HAp nucleation.

12.14.2 SPIDER SILK A relatively fewer number of works use spider silk, compared to SF from silkworms. We discuss a few such attempts here. Gomes et al. (2013) used a composite of spider silk, from N. clavipes, with BSP to produce films. These films supported cellular functions in vitro and caused no adverse reactions in vivo. Spider dragline silk, in itself, has also been used as a template for HAp mineralization and the crystals formed were mostly oriented along the c-axis (Cao and Mao, 2007). A combination of self-assembling domains from dragline silk of N. clavipes and a silica mineralizing peptide from Cylindrotheca fusiformis proved suitable for mineralization via silica precipitation This combination improved strength, due to the self-assembling of the spidroin segments, and was biodegradable (Mieszawska et al., 2010b). These silk films covered with silica showed favorable results for MSC proliferation and osteogenesis.

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12.14.3 SILK SERICIN As opposed to SF, silk sericin has not seen much widespread use for tissue engineering applications. This could be due to the immunogenic reaction that is caused from any trace sericin with fibroin, meaning that sericin would not be an obvious choice. It has been shown that sericin, by itself, does not trigger any immunoreaction and does so only in the presence of fibroin (Macintosh et al., 2008). Silk sericin has been used as modifier in the process of fabricating HAp microspheres, (Liu et al., 2013) and as an initiator for HAp nucleation in biomimetic metastable calcium phosphate solution, especially with high molecular weight (Cai et al., 2009). Aimed towards improving cell-substrate interaction (proliferation and adhesion) A. mylitta sericin has been employed to functionalize the surface of titanium substrates (Nayak et al., 2013). Greater concentration of silk sericin increases the size and crystallinity of HAp crystals formed. The form could be matched to that of natural enamel by optimizing silk sericin concentration and mineralization time (Cai et al., 2009).

12.15 RECENT TRENDS IN BONE TISSUE ENGINEERING 12.15.1 USE OF BIOREACTORS In vitro cultures can be carried out on scaffolds for a long duration by using bioreactors, until bone-like tissue forms on the scaffolds. BMSCs cultured on 3D porous SF scaffolds in osteogenic conditions yield trabecular bone-like structures (Meinel et al., 2004, 2005a) up to 1.2 mm thick (Meinel et al., 2004). Hofmann et al. (2007) prepared 3D SF scaffolds that had different-sized pores, distributed on different sides. Culturing MSC on these scaffolds in a spinner flask, maintaining osteogenic growth media, produced cortical bone tissue on the side, with larger pores and trabecular structure on the smaller pore side. These SF scaffolds consequently showed the flexibility of developing different bone tissue types on the same scaffold, to match the structural requirements at certain implantation sites. In vitro culture in spinner flasks and other bioreactors means that scaffolds with already developed bone tissue can be implanted at reconstruction sites, dramatically improving their ability for bone regeneration in vivo. In the previously cited works, credit has often been given to the slower degradation of SF scaffolds as opposed to, say, collagen scaffolds. Bioreactors offer much closer control on culture conditions than static culture; managing fluid flow, substrate stress, and any mechanical (or other) stimuli (Kundu and Kundu, 2010). Bioreactors can improve the supply of oxygen, nutrients, and regulatory proteins to every part of a scaffold by overcoming the limitations of a static culture, and

12.15 Recent Trends in Bone Tissue Engineering

FIGURE 12.3 A pictoral representation of HAp deposition on silk fibroin substrate.

accordingly make it possible to tissue engineer larger bone fragments (Meinel et al., 2004). The use of a bioreactor in developing silk-based implants for use in vivo is illustrated in Fig. 12.3. However, bioreactors are complex devices and require considerable upfront investment, maintenance, and operational knowledge.

12.15.2 COCULTURES OF MULTIPLE CELLS ON SILK SCAFFOLDS The skeletal system is a combination of multiple tissues, including bone, ligament, cartilage, etc. and it also needs appropriate vascularization for a competent level of reconstruction. Consequently, several studies of SF scaffolds for BTE applications have considered and successfully implemented cocultures of different tissue progenitor cells to verify applicability over a range of bone defects. At a basic level, SF scaffolds and SF-HAp composite scaffolds have been used for long term cocultures of osteoblast and osteoclasts with good results. In their experiments, Hayden et al. (2014) found that it was more advantageous to use nitrogenous bisphosphonate than nonnitrogenous bisphosphonate, as the later required large dosage levels to aid osteoblast activity, unlike the former, which can diminish osteoblast activity. Sun et al. (2012) used the direct wire assembly method to fabricate precisely designed, 3D SF-HAp composite scaffolds with filamentous structure, to aid vascularization. Similarly, SF fibers have been used to improve poly(D-Lactic acid) scaffold stiffness so the scaffold could support endothelial cells and promote vascularization when implanted in vivo (Stoppato et al., 2013). He et al. (2012b) designed a hybrid SF scaffold with three layers combined, with one layer being suitable for osteoblasts (due to HAp inclusion) while another

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being suitable for fibroblasts. When BMSCs were cocultured in between these two layers; each layer is seeded with osteoblasts and fibroblasts respectively, the stem cells differentiated into fibro-cartilaginous tissues.

12.15.3 GROWTH FACTORS DELIVERY THROUGH SILK SCAFFOLDS Growth factors can form a vital part of a scaffold implanted for bone regeneration. Accordingly, silk scaffolds in different configurations have been tried with different growth factors in vitro and in vivo trials. As was expected, silkbased scaffolds with growth factor inclusion do produce superior performance in terms of osteogenesis and new bone formation, compared to pure silk scaffolds. If certain growth factors, such as BMP-2, are not delivered precisely to the defect site, they can lead to inflammation in soft tissue or formation of ectopic bone (Diab et al., 2012). Methods of successfully implemented delivery include: coating growth factor on the silk nanofibers or films, as part of an electrospun nanofibrous mesh (Diab et al., 2012); a combination of rhBMP-2 and MSCs, as part of a calcium phosphate cement/SF injectable system (Gu et al., 2011); embedded in SF microspheres (Bessa et al., 2010); and adenovirus with growth factor encoding embedded in SF scaffolds (Zhang et al., 2011). For coupling growth factors such as BMPs onto surface of silk scaffolds, a commonly used process is covalent bonding via carbadimide reaction; such a linkage may be more conducive to osteogenesis than just physically trapping BMP molecules in SF scaffolds (Bhattacharjee et al., 2016b). Porous 3D scaffolds of SF with PLGA microparticle-loading, has been used as a delivery medium for insulin-like growth factor I (IGF-I), with the PLGA microparticles serving as moderating agents for growth factor release rate (Wenk et al., 2009). Fibroin microparticles could be used in a similar role, for controlling growth factor release, as part of a scaffold or as an injectable system (Bessa et al., 2010). Microparticles also enhance the surface area available for embedding the protein molecules. In favor of an injectable system, such a composite of calcium phosphate cement, SF, and rhBMP-2 produced repaired bones similar in stiffness to those produced by autografts in an in vivo trial (Gu et al., 2011).

12.15.4 GENE THERAPIES Gene therapy involves introducing specific genes in an individual’s cells— directly or via a scaffold—that can promote specific biological action. With its potential to deliver growth factor at defect sites in a sustained fashion, gene therapy holds a promise to add a new dimension to BTE research. Engineered adenovirus has been used to transfect cells and elevate osteogenic growth factor expression (Zhang et al., 2012). Transfected cells secreting growth factors offer a

12.15 Recent Trends in Bone Tissue Engineering

much better solution than the external introduction of growth factors. Adenovirus can also be delivered by combining it with SF scaffolds. As shown by Zhang et al. (2012), such a scaffold can give a sustained generation of BMP-7, thus supporting cell proliferation an osteogenic differentiation, for bone formation in vivo. Adenovirus with a human BMP-2 gene can be used to transfect MSCs, cultured in vitro, and then used for bone reconstruction in vivo to reliably serve as a source for secreting biologically active BMP-2 (Meinel et al., 2006b). Additionally, nanoparticles hold great potential in targeted delivery of DNA and silk offers the possibility to achieve this. SF-oligochitosan nanoparticles were used for delivery of siRNA (Shahbazi et al., 2015) for gene silencing applications, where a particular pathway could be targeted to mediate faster bone regeneration or target osteosarcomas.

12.15.5 CONCLUSION AND FUTURE PERSPECTIVES Two major trends were focussed in this chapter. One was the growth of mineralized silk scaffolds and the other is a move towards nanostructured scaffolds, attempting to better replicate the basic structure of native bone. Beyond these trends, though, several avenues of BTE using different forms of silk, and silk blends have been attempted with mostly successful outcomes. Silkworm silk is likely to continue as the dominant material source, considering spiders are not quite suitable for domestication. It is not unlikely that exploration of the varied domain of spider silks could turn up a form immensely useful for BTE. As of now, however, it does not look like the status quo would change. On a positive note, silk as a material source is more economic than many artificial polymers. A major lacuna that needs to be addressed is that, in spite of several successful in vitro and in vivo trials of silk-based scaffolds, none have proceeded to the phase of human trials. Consequently, the benefits of silk-based tissue engineering still remains distant for patients and victims. Knowledge of cellmaterial interaction, the subsequent reverberations that it would have on differentiation potential and compliance of silk scaffolds must be assessed thoroughly. The challenge lies in implanting this knowledge in the context of clinical translation. A systems biology approach is needed to understand these complications, and to study and implement the suitability of cells, the pathway they trigger, and comprehensively validate the compliance of these scaffolds. This is true for BTE, as well as all fields of tissue engineering. Accordingly, this element is of immediate importance for consideration by the whole community working on silk-based scaffolds. Collaborative action for remedying this situation would be one of the important developments we may aim for in the future.

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CHAPTER

Implantable drug delivery systems: An overview

13

Anoop Kumar1,2 and Jonathan Pillai1 1

Translational Health Science & Technology Institute (THSTI), Faridabad, India 2 Indo-Soviet Friendship College of Pharmacy (ISFCP), Moga, Punjab

CHAPTER OUTLINE 13.1 Introduction ...................................................................................................473 13.2 Classification of Implantable Drug Delivery Systems.........................................479 13.2.1 Passive Implants ........................................................................479 13.2.2 Dynamic Implants ......................................................................484 13.2.3 Electromechanical Systems .........................................................487 13.3 Design Approaches ........................................................................................489 13.3.1 Implant Material Selection ..........................................................489 13.3.2 Mechanisms of Drug Release From Implantable Drug Delivery Systems.....................................................................................490 13.4 Current Therapeutic Applications ....................................................................492 13.4.1 Women’s Health .........................................................................492 13.4.2 Chronic Diseases ........................................................................492 13.4.3 Infectious Diseases (Tuberculosis) ...............................................495 13.4.4 Neurology and Central Nervous System Health ..............................496 13.5 Current Challenges and Future Perspectives ....................................................496 13.5.1 Biocompatibility-Related Issues...................................................496 13.5.2 Patient Compliance ....................................................................498 13.5.3 Regulatory Aspects .....................................................................498 13.5.4 Cost-Effectiveness ......................................................................499 13.5.5 Future Perspectives ....................................................................499 13.6 Conclusion ....................................................................................................500 References .............................................................................................................501

13.1 INTRODUCTION The concept of implantable drug delivery systems (IDDSs) in modern medicine may be traced to Deansby and Parkes who, in 1938, subcutaneously (SC) implanted compressed pellets of crystalline estrone to study their effect upon Nanostructures for the Engineering of Cells, Tissues and Organs. DOI: http://dx.doi.org/10.1016/B978-0-12-813665-2.00013-2 © 2018 Elsevier Inc. All rights reserved.

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CHAPTER 13 Implantable drug delivery systems: An overview

castrated male chickens (Kleiner et al., 2014). Folkman and Long pioneered implantable formulations, with drug release rates controlled by a polymeric membrane, in the 1960s. They investigated the use of silicone rubber (Silastic) for long-term drug delivery at a systemic level (Folkman and Long, 1964). From these early beginnings, the potential of this mode of delivery in overcoming problems associated with oral administration, such as drug bioavailability, stability, toxicity, and duration of release, was recognized. Implant delivery systems have been subsequently designed to reduce the frequency of dosing, prolong duration of action, increase the patient compliance, and reduce the systemic side effects (Liu et al., 2015; Martıńdel et al., 2009). In certain areas, like interventional cardiology and contraception, IDDSs have revolutionized clinical management practices and proven themselves as Industry standard and/or excellent therapeutics. Progress in other areas, such as oncology has been slower. Nonetheless, the promise of personalized and precision medicine delivery is driving a recent resurgence in this field. The advantages of IDDSs have been summarized in Table 13.1. IDDSs are very attractive for a number of classes of drugs, particularly those that cannot be delivered via the oral route, are irregularly absorbed via the gastrointestinal tract, or that benefit from site-specific dosing. Examples include steroids, chemotherapeutics, antibiotics, analgesics and contraceptives, and biologics Table 13.1 Summary of Design Features of Idealized Implantable Drug Delivery Systems Design Feature

Summary of Potential Advantages

Localized delivery

Drug(s) are released in immediate vicinity of implant. Action may be diffusion, limited to the specific location of implantation Patient does not need to comply with repeated and timely intake of medication throughout the implantation period. Compliance is limited to one-time implantation (and potential removal in the case of nonbiodegradable implants) Controlled release for extended periods of time and localized dosing possible with at site of action; adverse effects away from site of action are minimized; peaks and valleys in plasma drug concentration from repeated intermediate release dosing are avoided Localized implantation of site specific drugs can avoid firstpass hepatic effects, thereby reducing dose required to ensure systemic bioavailability Protection of drug undergoing rapid degradation in the gastrointestinal and hepatobiliary system Hospital stay or continuous monitoring by healthcare staff may not be required for chronic illnesses If allergic or other adverse reaction to drug is experienced, discontinuation of therapy by implant removal is possible

Improved patient compliance

Minimized systemic side effects

Lower dose

Improved drug stability Suitability over direct administration Facile termination of drug delivery

13.1 Introduction

such as insulin or heparin (Amory et al., 2006; Forwith et al., 2011; Grossman et al., 1999; Matheny et al., 2014; Mohamed et al., 2012; Schmidmaier et al., 2006; Schaepelynck et al., 2011; Yang et al., 2010). Implant morphology is typically cylindrical, with monolithic devices at the millimeter or centimeter scale being most commonly employed. Implantation is typically done in subcutaneous or intramuscular tissue, with the aid of special implantation devices, needles, or the use of surgery (Wong et al., 2012). Subcutaneous tissue or intramuscular tissue are ideal locations for implantation of drug-depot devices, due to high fat content that facilitates slow drug absorption, minimal innervation, good hemoperfusion, and a lower possibility of localized inflammation (low reactivity to the insertion of foreign materials) (Fumimoto et al., 2009). In addition to subcutaneous implantation, various other body regions have also successfully served as implantation sites, particularly for delivery to localized tissue such as intravaginal, intravascular, intraocular (Birch et al., 2013), intrathecal (Brand et al., 2007), intracranial, and peritoneal (Grabowska-Derlatka et al., 2016). Perhaps the most common clinical application to date targets cardiac or carotid arteries as sites for drug-eluting stents (DES), delivering therapy to intravascular locations. Implants can be used as delivery systems for either systemic or local therapeutic effects. For systemic therapeutic effects, implants are typically administered SC, intramuscularly, or intravenously, whereby the incorporated drug is delivered from the implant and absorbed into the blood circulation. Implants for local effects are placed into specific body sites, where the drug acts locally, with relatively negligible absorption into the systemic circulation (De Oliveira et al., 2015). Pathology-targeting implants aim to release a drug and limit the therapeutic effect at the site of implantation (Lueshen et al., 2015). Implants are typically designed to release the incorporated drug in a controlled manner, allowing the adjustment of release rates over extended periods of time, ranging from several days to years. An ideal IDDS should be designed to substantially reduce the need for frequent drug administration over the prescribed treatment duration. As such, it should be environmentally stable, biocompatible, sterile, and be readily implantable and retrievable by medical personnel to initiate or terminate therapy. Additionally, it must enable rate-controlled drug release at an optimal dose, be easy to manufacture and provide cost-effective therapy over the treatment duration (Rajgor et al., 2011; Ronak et al., 2012). Beginning with Gliadel (Westphal et al., 2003), initial progress in commercializing safe and effective implants has been slow, despite considerable effort. Performance issues, such as lack of compatibility between drug and carrier leading to burst release or shutdown, and concerns regarding potential toxicity and carcinogenicity, stability and reproducibility, had to be overcome. Furthermore, acceptance by patients and physicians alike was initially limited (Loudon, 2007; McGregor et al., 2000; Takamura et al., 1994). However, the approval by the FDA in 1990 (US Food and Drug Administration, 1990) of Norplant, a

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silicone-based device, was a breakthrough event in establishing IDDS as an acceptable method of drug delivery. By the time Boston Scientific Corporation got FDA approval of the PROMUS Element stent on June 01, 2015 (US Food and Drug Administration, 2015), IDDS had become truly mainstream products. The PROMUS Element stent has a metal scaffold, with the drug everolimus contained in a thin coating on the stent’s surface (Boston Scientific News Releases, 2011). Recently, Boston Scientific Corporation also obtained approval for the Innova Vascular Self-Expanding Stent System, which is used to re-open narrowed regions of an artery that supplies blood to the leg (Boston Scientific, 2015). These implants have been presented, along with a selection of other notable commercialized implants, in Table 13.2. This chapter reviews various types of IDDSs, from biomaterial-based to electromechanical systems. Design approaches to optimal drug delivery including methods to tailor drug release profiles and the mechanism of release kinetics are briefly reviewed. Potential therapeutic applications and biocompatibility-related issues are discussed. Finally, this chapter concludes with a summary of future perspectives of IDDSs, particularly in their applicability to precision and personalized medicine. It may be noted that, while numerous other implants also serving a therapeutic function exist on the market, this chapter focuses exclusively on those that specialize in drug delivery. As such, dental, orthopedic, cardiovascular, and gastric implants often serve to provide structural support, supplement the function of localized tissue or promote homeostasis and wound healing. However, they do not modulate localized delivery of therapeutic molecules. Common examples include heart valves, bare-metal or nonresorbable stents, dental and bone implants, artificial joints, sutures, ocular lenses, etc. Similarly, devices such as cardiac pacemakers or deep-brain stimulation electrodes deliver controlled quantities of electrical signals to localized tissue, thus also providing a therapeutic effect. Mechanical ports and catheters provide sustained flow paths for long-term intravenous infusion of fluids. However, none of these may be classified by the traditional definition of a drug delivery device, and are therefore excluded from this discussion and review. Similarly, a number of biosensors may be implanted for various durations, but are restricted to only a sensing function. While this sensory information may be fed back to a control loop that eventually modulates drug delivery, the sensor itself is not considered to be an IDDS as per our definition. More recently, many excellent examples of micro-needle patches employing both passive and active modes of drug delivery have been developed. However, as these micro-needles have a relatively minimal portion of the total device penetrating the skin, often on a transient basis, they cannot be classically defined as “implants,” but fall under the class of trans-dermal delivery systems. These too have been excluded from the current discussion, even though they represent an exciting and promising mode of drug delivery for the future.

Table 13.2 Summary of Major Implantable Systems in Chronological Order of Approval by FDA PMA Applicant

Approval Date

FDA Current Status

AKORN

1974

Discontinued

Astrazeneca

1989

Population Council

1990

Available in market Discontinued

Ganciclovir

Bausch and Lomb

1996

Discontinued

Cytomegalovirus (CMV) retinitis

Gliadel

Carmustine

Arbor Pharms LLC

1996

Available in market

Brain tumors

6.

Jadelle

Bayer Pharmaceuticals

1996

Marketed

Contraceptive system

7.

Marketed

1998 2000

Discontinued Discontinued

Symptoms associated with menopause Contraceptive system Prostate cancer

10.

NuvaRing

Merck & Co.’s

2001

Marketed

Contraceptive system

Mascarenhas et al. (1998) Fowler et al. (2000), Wright et al. (2001) Friend (2011)

11.

Cypher stent

Etonogestrel Leuprolide acetate Etonogestrel and ethinyl estradiol Sirolimus

Pharmacia and Upjohn Organon DUROS

1996

8. 9.

Estring silicone intravaginal ring (IVR) Implanon Viadur

Levonorgestrel (modern version of the Norplant) Estradiol

2003

Approved

Antiproliferative effects

Tian et al. (2015)

12.

Taxus stent

Paclitaxel

Johnson & Johnson/Cordis Boston Scientific

2004

Approved

Tian et al. (2015)

13.

Taxus Endeavor Xience V

Everolimus

2004

Approved

To prevent cell proliferation and neointimal For improving coronary luminal diameter in patients with symptomatic heart disease due to de novo native coronary artery lesions

Specific Drugs Being Delivered

S. No.

Product Name

1.

Ocusert System

2.

Zoladex

3.

Norplant

Pilocarpine and alginic acid Goserelin acetate Levonorgestrel

4.

Vitrasert

5.

Abbott Vascular

Indication

References

Glaucoma and xerostomia Advanced prostate cancer and advanced Contraceptive system

Karthikeyan et al. (2014) Hutchinson and Furr (1987), Shi and Li (2005) Rademacher et al. (2014), Glasier (2002), Johansson, (2000), Sivin (2003) Choonara et al. (2010), Silva et al. (2010), Wong et al. (2013) Aoki et al. (2014), Attenello et al. (2008), Dörner et al. (2011), La Rocca and Mehdorn (2009) Brache (2014)

Baum et al. (2012)

Drugs@FDA

(Continued)

Table 13.2 Summary of Major Implantable Systems in Chronological Order of Approval by FDA Continued Specific Drugs Being Delivered

S. No.

Product Name

14.

Vantas

15.

Nexplanon

16.

ITCA 650

Histrelin acetate Etonogestrel (a new version of Implanon) Exenatide

17.

Cypher

Zotarolimus

18.

Everolimus

19.

PROMUS Element Plus EverolimusEluting Platinum Chromium Coronary Stent System Nesterone

20.

Capronor

Nestorone and estradiol Levonorgestrel

21.

Profact or Suprefact Depot

Buserelin acetate

22.

Synchromed

Infumorph, lioresal, prialt (ziconotide), floxuridine, methotrexate and baclofen

PMA Applicant

Approval Date

FDA Current Status

Endo Pharm

2004

Marketed

Merck & Co.

2006

Marketed

ALZA Corporation Medtronic

2007 2008

Under development Marketed

Boston Scientific Corporation

2015

Approved

Population Council Research Triangle Institute Sanofi-aventis Canada Inc.

Not approved Not approved

Phase III clinical trials Clinical trials

Medtronic

Recalled on June 03, 2013

Indication

References

Prostate cancer symptoms Contraceptive system

Schlegel (2009)

Type II diabetes

Henry et al. (2013)

Patients who have a narrowing in their coronary arteries caused by atherosclerosis Narrowing of coronary arteries caused by coronary artery disease

Drugs@FDA

Uhm et al. (2016)

Drugs@FDA

Contraceptive system Contraceptive system Hormone-responsive cancers such as prostate cancer or breast cancer and in assisted reproduction Insulin delivery, anticoagulant therapy and cancer chemotherapy

Oliveira-Ribeiro et al. (2006), Kleiner et al. (2014)

Michael et al. (2012), Yue et al. (2016)

13.2 Classification of Implantable Drug Delivery Systems

13.2 CLASSIFICATION OF IMPLANTABLE DRUG DELIVERY SYSTEMS Classification of IDDSs is difficult, given that there are numerous exceptions and hybrids that may be listed under multiple categories. However, drug implants can be broadly subdivided into passive and active systems. Passive systems can be further classified into nondegradable and degradable implants, that typically have no moving parts or mechanisms. Active systems employ some energy-dependent method for providing a positive driving force to modulate drug release. These energy sources may be as diverse as osmotic pressure gradients or electromechanical drives.

13.2.1 PASSIVE IMPLANTS Passive implants tend to be relatively simple, homogenous and singular devices, typically comprising the simple packaging of drugs in a biocompatible material or matrix. By definition, they do not contain any moving parts, and depend on a passive, diffusion-mediated phenomenon to modulate drug release. Delivery kinetics are partially tunable by the choice of drug, its concentration, overall implant morphology, matrix material and surface properties.

13.2.1.1 Nondegradable implantable drug delivery systems While membrane-enclosed reservoirs and matrix-controlled systems are by far the most common, several other variants of nondegradable implants are commercially available. The matrix materials used in all these systems are typically polymers, with a documented history of both pre-clinical and clinical evaluation. Commonly used polymers include elastomers such as silicones and urethanes, acrylates and their copolymers, and copolymers vinylidenefluoride and polyethylene vinyl acetate (PEVA) (Colaris et al., 2016; DeWitt et al., 2005; Nunes-Pereira et al., 2015; Shastri, 2003; Zur et al., 2011). Within the polymeric matrices forming most passive monolithic implants, the drug is typically dispersed homogeneously throughout the matrix material. Alternatively, reservoir-type systems are characterized by a compact drug core, surrounded by a permeable nondegradable membrane, the permeability and thickness of which controls the diffusion of the drug into the body (Martıńez-Rus et al., 2012). One of the earliest, widely deployed, nondegradable reservoir implants is Norplant. This IDDS was developed and trademarked by the Population Council in 1980, introduced worldwide in 1983. As stated earlier, it was approved by the US FDA in December 1990, following which marketing in the United States was initiated in February 1991 (Spicehandler, 1989). This contraceptive system consists of six thin, flexible silicone capsules (silastic tubing), each loaded with 36 mg of the hormone levonorgestrel. When implanted SC, typically on the inside

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FIGURE 13.1 Nonbiodegradable implants (A) norplant and (B) implanon.

upper arm of female users (Rademacher et al., 2014), it is capable of offering contraceptive protection for up to 5 years (Fig. 13.1). Its effectiveness and popularity may be gauged by the fact of its approval in over 60 countries. While Norplant ceased to be marketed in the United States in 2002, it is still available in other countries and has been successfully used by over 60 million women (Glasier, 2002; Johansson, 2000). Another FDA-approved IDDS contraceptive implant is Implanon, which was launched on the international market in 1998, eventually entering the United States in 2006. It is a single-rod implant (length 4 cm, width 2 mm) and consists of a PEVA core (reservoir) that encapsulates 68 mg of etonogestrel and releases the drug over 3 years (Fig. 13.1). The rate of drug release is controlled by a PEVA membrane covering the rod (Bennink, 2000; Zheng et al., 1999). Protection from pregnancy can be extended beyond the initial 3 years upon removal and immediate replacement with a fresh implant. Designed for easier subcutaneous insertion and removal than Norplant, Implanon has found just great acceptance by patients and providers alike (Mascarenhas et al., 1998). Other notable IDDS for the delivery of different progestins include Capronor and Nesterone, which are still in development by the Research Triangle Institute and the Population Council, respectively. These are being designed in various form factors, including rods, pellets, and microcapsules. Nesterone implant is a


single rod, 2-year silasticsub-dermal implant containing 93 mg of Nestorone. All newer designs are designed for the following features: they are highly efficacious, cost effective, safe, rapidly reversible, show extended duration, require no daily effort, stabilize hormone levels, contain no estrogen, are usable during lactation, are discreet (nearly invisible) and comfortable during use (Oliveira-Ribeiro et al., 2006; Kleiner et al., 2014). In more recent times, the most important and pervasive example of nondegradable IDDS is the DES. This mode of delivery has revolutionized the treatment of vascular disease. Specifically, in the case of coronary artery disease (CAD), DES may reduce restenosis typically seen in bare-metal stents by 60%75%. DESs are fast becoming the standard of care in the treatment of stenotic CAD, frequently deployed for opening blockages and maintaining patency in a coronary artery. A DES is normally a three-component system, comprising a scaffold (or stent) for ensuring vascular luminal patency, a matrix or coating (polymer) to control drug release, and a drug to inhibit neointimal restenosis. Release of drugs from these coatings is typically diffusion-controlled (Acharya and Park, 2006; Bønaa et al., 2016). A good example of an implant delivering antiviral drugs is Vitrasert, which was developed and commercialized for the treatment of cytomegalovirus (CMV) retinitis. The system releases ganciclovir, following an intravitreal implantation of a compressed tablet of the drug. The tablet is coated with polyvinyl alcohol (PVA), then partially over-coated with PEVA, and finally affixed to a PVA suture stub (Choonara et al., 2010; Silva et al., 2010). While nonbiodegradable systems are typically robust and structurally resilient over their intended lifetime, their obvious drawback is the need for extraction after depletion of the drug cargo. While their biocompatibility is typically proven for the duration of proposed use, the risk of infection and cosmetic defacement at the site of subcutaneous implantation following long-term in situ placement are both undesirable. In many instances, it is therefore expedient to extract the implant following expenditure of its drug cargo to avoid adverse clinical sequelae. However, explantation itself carries its own risks of infection, tissue damage, or other serious adverse effects, particularly when carried out by poorly trained healthcare staff, or in unsanitary conditions. In the case of Norplant deployed in the United States, for instance, the implant was associated with explantation difficulties, due to poor initial insertion technique. This problem was attributed to operator inexperience, eventually contributing to its removal from the US market (Dunson et al., 1995; Sivin, 2003). Consequently, design of the implant form factor, i.e., shape and size, as well as the availability of specialized devices for easy implantation and extraction, are particularly important in this case. Alternatively, inclusion of an x-ray visible tracer material may be necessary for detection of the embedded implant prior to extraction by imaging (Brache, 2014).

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13.2.1.2 Biodegradable implants To overcome the drawbacks of nonbiodegradable implants, biodegradable systems, based on polymers such as poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA), poly(caprolactone) (PCL) or their block copolymer variants with other polymers have been developed. A major advantage of biodegradable systems is that the biocompatible polymers used for fabricating these delivery systems are eventually broken down into safe metabolites and absorbed or excreted by the body (Claes and Ignatius, 2002; Tian et al., 2012). Labile bonds that are prone to degradation by hydrolysis or enzymes, such as ester, amide, and anhydride bonds, are characteristic of the backbone of biodegradable polymers. Complete degradation of the implant postdrug release makes surgical removal of the implant after the conclusion of therapy unnecessary, thereby reducing potential complications with explantation and increasing patient acceptance and compliance. Among the many biodegradable polymers that have been explored as a component of an IDDS, aliphatic polyesters of lactic and glycolic acids are perhaps the most widely investigated in terms of availability of clinical data. Their desirable features include a long and well-documented history of biocompatibility (including acceptability of degradation products) since their first use in 1970, and facile tuning of degradation kinetics. An additional benefit is that they been approved by the FDA for use in several pharmaceutical products, including hormones, anticancer drugs and antibiotics. In general, pre-approval of the implant material or biodegradable matrix helps in reducing the regulatory burden for proving the matrix safety for such implants. The ease of fabrication by several techniques is an additional reason why these polyesters have been adopted substantially more than other polymers. For example, polymers such as PCL lend themselves to injectable formulations that subsequently form depot-like implants in situ at the location in tissue (Rahimi et al., 2016). Thermoforming by melt extrusion (Bernasconi et al., 2014; De Nijs, 1990), solvent evaporation (Dong et al., 2014; Wong and Visionex, 1989), and compression molding from powder or pellet form (Gangadharam et al., 1999) are other reported methods for simple implant fabrication. Drugs of various types and properties, ranging from small molecules to peptides and proteins, antibiotics (Gad et al., 2008; Tamaddon et al., 2015), antiviral drugs (Kunou et al., 2000), anticancer drugs (Brem et al., 1991; Ranganath et al., 2010), analgesics (Berde et al., 2001) and steroids (Sanders et al., 1986), have reported to be incorporated into PLGA and PLA polymers for the preparation of implants. Although the acidic byproducts of polyester degradation can cause instability of proteins and localized inflammation, some therapeutic proteins, such as recombinant human growth hormone and insulin, have been evaluated for delivery. Other important biodegradable polymers include polyanhydrides, poly (orthoesters), and poly(phosphoesters) (Parent et al., 2013).


In the context of peptide and protein drugs that are highly susceptible to acidic and enzymatic degradation, lipid implants may prove to be particularly attractive, given their good tissue biocompatibility and ability to protect the drug cargo. Acid-stable drugs can also be incorporated within lipid implants and be effectively protected against physical and/or chemical aggressions (e.g., enzymatic degradation) within the human body. Lipid implants can easily be produced by compression (Kreye et al., 2011), or by the twin-screw extrusion process (Schulze and Winter, 2009). However, mechanisms of drug release can be very complex and are not yet fully understood. Kreye et al. (2011) observed that diffusion with constant diffusivities was the dominant mass transport mechanism controlling drug release from compressed lipid implants. In the future, more studies are needed, which will significantly help to facilitate the optimization of this type of advanced drug delivery system (DDS). There are various commercial biodegradable implants available in the drug market. Selected examples include the Gliadel wafer for the treatment of brain tumors, Zoladex for treatment of advanced prostate cancer and advanced breast cancer, and Profact or Suprefact Depot for treatment of hormone-responsive cancers, such as prostate cancer or breast cancer and in assisted reproduction. Each of these is briefly reviewed below. The Gliadel wafer is one of the earliest examples of a biodegradable IDDS, approved by the FDA in 1996. It consists of biodegradable polyanhydride disks (1.45 cm in diameter and 1.0 mm thick), designed to deliver the chemotherapeutic drug, bis-chloroethylnitrosourea (BCNU) or carmustine, directly into the cavity created after surgical resection of the tumor (high-grade malignant glioma). The biodegradable polyanhydride copolymer in a 20:80 molar ratio of poly[bis (p-carboxyphenoxy)propane:sebacic acid], has been used to control local delivery of carmustine (Aoki et al., 2014; Attenello et al., 2008; Do¨rner et al., 2011; La Rocca and Mehdorn, 2009). Zoladex is a biodegradable implant containing goserelin acetate, which is a decapeptide analogue of lutinizing hormone-releasing hormone (LH-RH). It uses PLGA or PLA as a carrier, in which the drug is dispersed in the polymer matrix using hot-melt extrusion method. For commercial use, the implant is distributed in the form of a prefilled syringe. The drug is continuously released over a period of 1 or 3 months (Hutchinson and Furr, 1987; Shi and Li, 2005; Product Monograph of Zoladex, 2016). Profact Depot or Suprefact Depot contain buserelin acetate (gonadotropinreleasing hormone agonist) and PLGA (75:25 molar ratio) is used as a drug carrier. The implant is designed for 2- and 3-month drug release, where the duration of action depends upon the relative ratio of drug and PLGA in the implants (Product Monograph of Suprefact, 2016). Some disadvantages of biodegradable IDDS are: their higher complexity, development cost, regulatory requirements, and the lack of availability of polymers with the exact physical properties needed, including mechanical strength

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and tunable degradation kinetics. In general, the development of biodegradable systems is a more complicated task than formulating nondegradable systems. When fabricating new biodegradable systems, variables to be taken into consideration include the in vivo degradation kinetics of the polymer, which must ideally remain constant to maintain sustained release of the drug. Unfortunately, these can be highly variable, depending upon patient age and disease state. In comparison to metal implants, biodegradable implants may also be more expensive, due to additional costs associated with specialized materials and regulatory approval (Iyer et al., 2006; Zhao et al., 2017).

13.2.2 DYNAMIC IMPLANTS As defined earlier, dynamic implant systems harness a positive driving force to enable and control drug release. As a result, these are typically able to modulate drug doses and delivery rates much more precisely than passive systems. However, this comes at a higher cost, both in terms of complexity and actual device price.

13.2.2.1 Implantable pump systems External control of dosing is a requirement for many drugs, a feature that is difficult to obtain when using biodegradable or nondegradable delivery systems. Pump systems have been used to provide the higher precision and remote control needed in these situations. Additionally, they offer a number of advantages, such as evasion of the GI tract, avoidance of repeated injections, and improved release rates (faster than diffusion-limited systems). With advances in microelectronics since the 1970s, remote control over delivery rates or integration of implantable sensors to create feedback-controlled drug delivery is now feasible. Implantable pumps primarily utilize osmosis, propellant-driven fluids, or electromechanical drives to generate pressure gradients and enable controlled drug release as described below.

13.2.2.2 Osmotic pumps Osmotic pumps have found wide acceptability among all active IDDSs (Chen et al., 2016). The first osmotic pump was devised by Australian pharmacologists, Rose and Nelson, who developed an implantable osmotic pump in 1955, named the Rose and Nelson osmotic pump (Jain, 2001). Higuchi and Leeper made a few modifications to this design and introduced it to the pharmaceutical world in the year 1973 (Higuchi and Leeper, 1973). In the same year, a dispensing device with a means of filling, containing a powdered agent capable of generating osmotic pressure, was also designed (Theeuwes, 1973). In 1975, a design known as the elementary osmotic pump came into existence and was patented by the Alza Corporation in 1976 (Theeuwes, 1975). The design comprises a drug reservoir surrounded by a semipermeable membrane, which allows a steady inflow of


surrounding fluids into the reservoir through osmosis. A steady efflux of the drug then ensues via the drug portal, an opening in the membrane, as a result of the hydrostatic pressure built on the drug reservoir. Nearly constant or zero-order drug release is maintained until complete depletion of the drug packaged in the reservoir (Gong et al., 2015; Pan et al., 2016). The DUROS leuprolide implant, named Viadur, was approved as the first implantable osmotic pump for humans in the United States in the year 2000 (Fowler et al., 2000). DUROS devices can be designed to deliver therapy for lengths of time ranging from several weeks to as long as 1 year. While the device can potentially deliver a broad array of therapeutic molecules, it is particularly suitable for potent peptides and proteins that require chronic dosing. Specifically, the Viadur system has been marketed for the palliative treatment of prostate cancer via the delivery of the GnRH analog leuprolide (Wright et al., 2001). The cylindrically shaped device consists of a reservoir made of an inert titanium alloy, capped at one end by a water-permeable membrane. At the other end, a diffusion moderator, through which drug formulation is released from the drug reservoir, caps the reservoir. The cylinder diameter ranges from 4 to 10 mm and the length is typically 45 mm, although smaller or larger systems may be designed, based on requirements of drug loading and the implantation site (Wright et al., 2001). The drug formulation, piston, and osmotic engine are contained inside the cylinder. The piston separates the drug formulation from the osmotic engine and seals the osmotic engine compartment from the drug reservoir. The diffusion moderator is designed, in conjunction with the drug formulation, to prevent body fluid from entering the drug reservoir through the orifice (Fig. 13.2). The specially designed, semipermeable polyurethane polymer membrane enables control of delivery rate, and is chosen for specific water permeability and antifouling properties during in vivo operation. Drug may be loaded into the reservoir in either dissolved or suspended formulations in suitable and approved solvents. Since formulation stability is of utmost importance, the choice of solvents and other excipients is critical and specific to the properties of the drug being delivered. The system is under further development for the delivery of exenatide and has also been investigated for the delivery of other drugs, including interferon (OMEGA DUROS system), sufentanil (Chronogesic system), and other opioids. Another example of an osmotic pump for continuously delivery of drugs, hormones, and other test agents is the ALZET pump. This device operates by osmotic displacement and is able to deliver its cargo at controlled rates over delivery windows between 24 h and 6 weeks. The drug to be delivered is filled into a core reservoir, which is isolated from a chamber, containing an osmotic salt, by a semipermeable membrane. Due to the presence of a high concentration of salt in a chamber surrounding the reservoir, water enters the pump through the semipermeable layer. The entry of water increases the volume in the salt chamber, causing compression of the flexible reservoir and delivery of the drug solution into the host via the exit port (Hill et al., 2013). Grossman et al. (1999)

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FIGURE 13.2 Schematic of osmotic pumps for drug delivery (A) DUROS and (B) ALZET.

have shown that the ALZET pumps were able to release 262 mg/h of the opioid hydromorphone over 2 weeks, to produce stable plasma concentrations of approximately 3040 mg/mL. This type of delivery system successfully mitigates the “initial burst effect” that is inherent in other systems, particularly in degradable matrices. This level of control over a sustained period can be particularly important in the delivery of critical medication for chronic pain management, thereby justifying the additional cost, complexity, and need for implantable systems.

13.2.2.3 Propellant infusion pumps While osmotic pumps offer a higher level of control and zero-order release, compared to biodegradable systems, the volume of drug that they can release limits them. To counter this deficit, an alternative design utilizes propellant gas instead of an osmotic agent to generate a constant positive pressure for zero-order release. The use of a compressible medium, such as gas, allows for a larger volume of drug to be stored and released. A good example of this technology is the Infusaid, which is a fully implantable fixed-rate pump. In this design, the drug formulation chamber is separated from a chamber for a fluorocarbon propellant by a metal bellows encased in disk-shaped titanium housing. The system contains a flow restrictor to ensure a constant vapor pressure from the propellant at operating temperatures. The overall package is approximately 9 cm in diameter by 3 cm in height and has been utilized for insulin delivery, anticoagulant therapy, and cancer chemotherapy (Gill et al., 1986; Haq et al., 1986; Spa¨th et al., 1985).


13.2.3 ELECTROMECHANICAL SYSTEMS While osmotic and propellant-driven constant pressure pumps work well for small volumes of medication, this may be a severe limitation for certain chronic diseases requiring daily infusion of medication, precluding their use over long timespans. In such cases, it may be necessary to consider larger implants, wherein the storage capacity of the pump may be replenished from time to time, while the pumping mechanisms stay implanted. By necessity, this implies the use of electrically powered mechanical pumps, typically with moving parts and advanced control systems. The Synchromed pump, developed by Medtronic Inc., is an example of a peristaltic pump implant, featuring external micro-electronic control of the delivery rate. It has been employed for pain management using intrathecal delivery of opioids (Michael et al., 2012), treatment of severe spasticity using baclofen, and for other indications (Yue et al., 2016). The pump consists of an outer titanium shell that encases the pump mechanism and controller, a reservoir holding the drug solution, and a battery. It can be conveniently refilled with a needle and syringe via a silicone rubber septum on the system. The implant dimensions are 8.8 cm in diameter and 2.5 cm thickness. The system is typically implanted in the abdominal cavity; delivery of baclofen is enabled by a silicone catheter tunneled to the intrathecal space (Yue et al., 2016). Electromechanical infusion technology has been rapidly growing and offers some unique solutions for biomedical applications, particularly to address unmet medical needs related to precision dosing. An implantable medical DDS in the micronanometer range can be fabricated to accurately and conveniently administer very small amounts of medication. Micro electromechanical systems (MEMS) technology enables the manufacture of small devices using microfabrication techniques, similar to that used to fabricate silicon-based computer chips. MEMS technology has been used to construct micro-reservoirs, micropumps, nano-porous membranes, nanoparticles, valves, sensors, micro-catheters, and other structures using biocompatible materials appropriate for drug administration (Brady et al., 2014; Elman et al., 2010). Numerous drug delivery implants based on micro- or nano-fabrication technologies are being developed. Implantable medical devices (IMDs) are electronic devices implanted within the body to treat a medical condition, monitor the state, or improve the functioning of a body part. Current examples of IMDs include pacemakers and defibrillators to monitor and treat cardiac conditions (Altieri et al., 2016), neuro-stimulators for deep-brain stimulation in cases such as epilepsy or Parkinson’s disease (Jonas et al., 2016), DDSs in the form of infusion pumps, and a variety of biosensors to acquire and process different bio-signals. As noted earlier, we restrict the current discussion only to IMDs designed for drug delivery.

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An ideal IMD would protect the drug from the body until needed, allow continuous or pulsatile delivery of both liquid and solid drug formulations, and be controllable by the physician or patient. A device meeting these criteria may typically include an array of individually sealed reservoirs that could be opened on command to expose their contents to the body. One or more drug formulations could be sealed in the reservoirs, protecting them from the environment until the reservoir is opened and the drug is released. One of the important advantages of implantable delivery systems with individually addressable drug-containing reservoirs is the ability to totally control drug delivery amount and timing; continuous or pulsatile delivery could be accommodated. There is great flexibility in tailoring these systems for specific applications because the release characteristics can be governed independently by release mechanism, reservoir geometry, or drug formulation. ChipRx (Lexington, KY, USA) has proposed an implantable, single-reservoir device (Fig. 13.3) that, in theory, could be adjusted to deliver drugs with targeted pharmacokinetics and bioavailability. The release mechanism employs polymeric artificial muscles that surround and control micrometer-sized holes, and that open to release drug. The polymeric ring expands or contracts in response to an electrical signal transmitted through a conducting polymer that contacts a swellable hydrogel (Staples et al., 2006). Some of the latest IMDs have begun to incorporate numerous communication and networking functions, usually known as telemetry, as well as increasingly sophisticated computing capabilities (Cordes et al., 2016; Jiang et al., 2016). This has resulted in implants with more intelligence and provided patients with more autonomy, as medical personnel can access data and re-configure the implant remotely, i.e., without the patient being physically present in a medical facility. Apart from a significant cost reduction, telemetry and computing capabilities also allow healthcare providers to constantly monitor the patient’s condition and to

FIGURE 13.3 Small pill implant from ChipRx.

13.3 Design Approaches

develop new diagnostic techniques based on the IntraBody Network (IBN) of medical devices.

13.3 DESIGN APPROACHES The design approaches for IDDSs depend upon drug properties, side effect of drugs, targeted site, and disease. Selection of materials plays a major role in designing of implants, as does form factor and design for manufacture and assembly. In the case of all dynamic systems, internal packaging of components as well as design of drug storage volumes may be critically important. While detailed exploration of all relevant design parameters is beyond the scope of this chapter, we explore a few critical parameters, particularly materials for biocompatibility and mechanisms for drug release.

13.3.1 IMPLANT MATERIAL SELECTION The primary design consideration modulating both drug release and eventual implant resorbtion for biodegradable, passive systems is the degradation rate of the matrix polymer. The rate of degradation of the polymer in the body is dependent on many factors. For example, variations in body pH or temperature due to inflammation or other causes can result in a transient fluctuation in the degradation rate. Similarly, the surface area of the delivery system is a key variable affecting degradation. If the system is not restricted to surface erosion and is degraded throughout its volume, the overall ratio of surface area to volume of the implantable system typically increases. This, in turn, may accelerate the degradation rate as the monolithic implant degrades into smaller sub-units. However, if the erosion is strictly limited to the surface of the implant without any nonhomogeneous decomposition or breakdown into smaller units, then the total surface area available for erosion will gradually decrease, thereby resulting in a corresponding deceleration of the degradation rate. Consequently, the transformation in shape of the implant resulting from in vivo degradation should be precisely accounted for in the design. While targeting uniform and constant release, geometric shapes whose surface area does not change substantially as a function of time during erosion are preferable. A flattened slab-like shape that has no edge erosion may most closely approximate a zero-order release kinetic profile. Alternative designs employ an inert, biodegradable core, coated with the active drug matrix, to alleviate the change in surface area problem encountered during erosion (Tapolsky et al., 2000). Slow diffusion of the drug from the polymer matrix is another limitation of bioerodable systems. The rate of drug diffusion is usually slower than the bioerosion of the system, and is dependent upon the physicochemical properties of the polymeric matrix utilized in the formulation of the IDDS. This becomes a major challenge for extended-release applications or for drugs with a narrow therapeutic index.

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13.3.2 MECHANISMS OF DRUG RELEASE FROM IMPLANTABLE DRUG DELIVERY SYSTEMS Most passive implanted DDSs are based on four basic delivery mechanisms: matrix degradation, controlled swelling, osmotic pumping, and passive diffusion. Of these, the latter two, i.e., osmotic pumping and diffusion, have been the most successful in delivering drug in a linear process, wherein the drug dosage released is proportional to square root of the release time. In DDSs based on swelling control, solvent penetration into the matrix of the drug device is normally much slower than diffusion of the drugs, which then results in a lowered release rate. Delivery systems based on the osmotic principle can provide a constant release rate, as seen with the simple oral tablets commonly used (Verma et al., 2002). The release kinetics of drugs from systems mediated by osmotic pressure, swelling, and passive diffusion depend on the some key factors. The solubility and diffusion coefficient of the drug in the polymer, the drug load, and the in vivo degradation rate of the polymer are all important variables affecting release (Gulati et al., 2016). This is especially true for biodegradable systems, which may require advanced mathematical modeling to accurately predict drug release rates (King and McGinty, 2016).

13.3.2.1 Mechanism of drug release from nondegradable polymeric matrices As indicated earlier, passive diffusion is the main driver of solute transport from nondegradable systems of both reservoir- and matrix- type designs. Reservoir systems have the advantage of maintaining a relatively constant release rate, independent of the concentration gradient. This is likely to be mediated by the thickness and permeability of the rate-controlling polymeric membrane, and zero-order release kinetics may potentially be achieved. This is because, unlike direct diffusion, the driving force for release of the agent across the membrane is constant, assuming that concentration of the drug within the reservoir constantly equilibrates with the inner surface of the enclosed membrane (Fu and Kao, 2010). In contrast, drug release for matrix-type devices is more likely to be driven by Fickian diffusion. In this case, solute movement is directly driven by the concentration gradient, and is mediated by diffusion lengths and the degree of swelling. In general, nonerodible, diffusion-controlled DDSs work best for drugs with molecular weight of 1000 Da or less (Siepmann and Siepmann, 2012). Matrix systems offer slow diffusion of the drug through the polymeric material, thereby providing sustained release of the drug from the delivery system, as shown in Fig. 13.4. However, release kinetics of such systems are typically variable, and is temporally dependent on the volume fraction of the agent in the matrix. In general, release from the system is directly proportional to the concentration of the encapsulated drug within the matrix.

13.3 Design Approaches

FIGURE 13.4 Cross sectional view of idealized reservoir system and matrix system, showing diffusion of drug across the polymer.

13.3.2.2 Mechanism of drug release from biodegradable implants The drug release from biodegradable polymeric systems is controlled either by diffusion, degradation or a combination of both (Lance et al., 2015). The degradation-controlled mechanism happens when the diffusion rate of a drug is less than the degradation or erosion rate of a polymer carrier. The drug is released concurrently with the polymer degradation; sigmoidal release profiles are usually observed. As with the mechanisms of polymer erosion, drug release based on the degradation-controlled mechanism can also be divided into surface-degrading approach and bulk-degrading approach (Shuwisitkul, 2011; Wu and Chu, 2008). The main parameters influencing drug release profiles in implants modulated by the surface-erosion mechanism are the surface-to-volume ratio and the geometry of implants. In such polymers, degradation is restricted to the outer surface of the device. In contrast, degradation is nearly homogeneous throughout the material in a bulk-degrading polymer. For either case, particularly for polymers susceptible to hydrolytic degradation, access of water to the material is an important factor. Accordingly, water plays a critical role in modulating degradation as well as release kinetics. The degradation of semi-crystalline polymers occurs in two stages. The first stage is initiated by the infusion of water into the amorphous regions, leading to random hydrolytic cleavage of labile ester bonds at the boundaries. The second stage is characterized by the degradation of entire domains within the amorphous

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regions. The breakdown fragments of the polymer chain are significantly shorter than the original backbone. This resultant decrease in the average molecular weight of the polymer may be used as a metric to quantify the extent of degradation. Since PLGA or PLA polymer undergoes bulk degradation, drug release from such carriers is controlled only by bulk degradationcontrolled release mechanism (Jiang et al., 2011). The physicochemical properties of drugs also play a role in drug release from hydrolytically degraded PLGA- or PLA-based products. Implants containing highly water soluble drugs showed a large initial burst release, followed by a rapid release. Finally, the drug loading has also an impact on the release rates, with fast release being observed from implants fabricated with high drug-loading (Rahimi et al., 2016).

13.4 CURRENT THERAPEUTIC APPLICATIONS IDDSs are finding increasing applications in the areas of women’s health, chronic diseases (including lifestyle diseases), oncology, pain management, and neurology. In the following section, we review some of the major clinical areas in which implants have made a significant contribution to therapy and disease management.

13.4.1 WOMEN’S HEALTH In the area of women’s health, IDDSs plays a major role, especially in field of contraception (El Ayadi et al., 2016; Richters et al., 2016). As described earlier, Norplant was a popular 5-year nonbiodegradable implant (Affandi et al., 1987). A modern version of the Norplant, called Jadelle, is currently marketed by Bayer Pharmaceuticals (Brache, 2014). Pfizer has developed a silicone intravaginal ring (IVR), marketed under the brand name Estring. The device is designed to treat symptoms associated with menopause and releases 2 mg of estradiol for 90 days (Baum et al., 2012). The NuvaRing IVR from Merck offers combined delivery of 120 μg of etonogestrel and 15 μg of ethinyl estradiol per day over a 3 week period (Friend, 2011). Nexplanon is a new version of the Implanon implant also from Merck that is capable of delivering 68 mg of etonogestrel for up to 3 years. Like Norplant and Implanon, Nexplanon also has a cylindrical rod shape for subcutaneous implantation in the arm (Uhm et al., 2016).

13.4.2 CHRONIC DISEASES Of all clinical use scenarios, IDDS perhaps find their best applicability in the treatment of chronic diseases. As such, devices for a wide variety of clinical applications against chronic illnesses have been developed, as described further.

13.4 Current Therapeutic Applications

13.4.2.1 Cardiovascular disease DES are the best example of IDDSs for vascular disease treatment. DES were developed to specifically address the problems of restenosis encountered with bare-metal stents (BMS), and typically consist of a BMS, coated with a polymer which gradually releases a drug to inhibit cell proliferation that causes restenosis. The first DES to be launched was the Cypher stent in 2003, followed by the Taxus stent in 2004 (Tian et al., 2015). However, by the late 2000s, a specific problem started to emerge with DES that had not been seen with BMS: very late stent thrombosis. This is the phenomenon of a thrombus developing inside the stent more than a year after insertion, far later than usually observed with a BMS. While it is not yet fully clear what causes this pathology, there are two main potential culprits. The first is the antirestenotic drug itself that can delay endothelialization, while the second is a hypersensitivity reaction to the polymer, which remains on the surface of the stent structure once it has released its entire drug (Choi et al., 2014; Ma et al., 2010). For this reason patients, are currently advised to take anticlotting drugs for at least a year after they receive a DES, longer than with a BMS. The latest developments in DES technology are, therefore, understandably focused on overcoming the problem of late stent thrombosis, while retaining a superior clinical profile. Leading the field is the development of stents with biodegradable polymers, where the polymer breaks down once it has finished releasing the drug, essentially leaving a BMS. A number of companies have stents with biodegradable polymers that are FDA-approved (Table 13.2). Other notable current developments in stent technology are polymer-free stents, where the drug is released directly from a metal film coating on the stent, and completely bioresorbable stents, where the stent platform, along with its polymer coating, degrades over time. However, neither of these stent developments is yet commercially available, and their long-term safety and efficacy are yet to be proven in a sufficiently broad range of patient types (Harada et al., 2016; Otto et al., 2016; Sim et al., 2016).

13.4.2.2 Cancer The major challenge in anticancer therapy is to develop DDSs to deliver chemotherapeutic drugs safely and effectively without side effects. Therefore, IDDSs have a great potential to deliver chemotherapeutic drugs in a more effective and safe manner. Brain, prostate, and bladder cancer are a few examples, for which attempts have been made to enable treatment with a drug delivery implant (Barros et al., 2016; Exner and Saidel, 2008). The Gliadel Wafer described earlier has been approved as one of the first implantable brain cancer treatment to deliver chemotherapy directly to the tumor site. Another example is the Zoladex biodegradable implantable rod, delivering goserelin acetate for treating prostate cancer (Goldspiel and Kohler, 1991).

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Similarly, Endo Pharmaceutical has developed a hydrogel subcutaneous implant, called Vantas, also targeting prostate cancer. This product is capable of delivering 50 mg of histrelin acetate over a 12-month period (Schlegel, 2009). Also notable is a nonbiodegradable drug-eluting product from TARIS Biomedical, designed to provide relief for nonmuscle invasive bladder cancer (Burt and Hunter, 2006).

13.4.2.3 Diabetes Diabetes, a condition already at the level of global epidemic affecting 371 million patients, is a chronic disease state where implantable systems have the potential to transform the current standard of both diagnosis and treatment. On the diagnostic end, continuous glucose monitoring (CGM) may be achieved by SC implanted sensors for blood glucose measurement. Sensors like Medtronic’s Enlite, Dexcom’s G4 Platinum, and GlySens ICGMTM are already available commercially (Bergenstal et al., 2010). However, as these are mostly restricted to a sensing function without subsequent insulin delivery, they are only briefly mentioned here for reference. Of interest in the therapeutic area is US-based Intarcia Therapeutics Inc., whose product is designed to help treat type II diabetes. The ITCA 650 is essentially a DUROS implant delivery technology that Intarcia licensed from the ALZA Corporation in 2007. As such, it has the same cylindrical reservoir osmotic pump, enclosed by a titanium alloy package approximately the size of a matchstick. When implanted SC, it delivers over 1 year exenatide, a glucagon-like peptide-1 receptor agonist (Henry et al., 2013, 2014). This technology has shown promising results in its Phase IV clinical trial and has the potential to become a disruptive innovation, but it remains to be seen how successful it is commercially.

13.4.2.4 Ocular therapy Several different implantable systems have been evaluated to provide prolonged ocular delivery. These include membrane-controlled devices, implantable silicone devices, and implantable infusion systems (Barar et al., 2016; Yasukawa et al., 2006). An example of the membrane-controlled system is the Ocusert device, containing pilocarpine and alginic acid in a drug reservoir, surrounded by a release ratecontrolling ethylene-vinyl acetate membrane. The release kinetics of the Ocusert system for delivery of pilocarpine at 20 or 40 μg/h show initial burst release, followed by a near zero-order over 1 week. Device performance in adults appears to be satisfactory, providing control of intraocular pressure with minimal side effects (Karthikeyan et al., 2014). Another example of an implant evaluated for ocular cancer treatment is a silicone rubber balloon containing an anti neo-plastic agent, BCNU. The device consists of two sheets of silicon rubber glued at the edges with silicon adhesive to form a balloon like sac through which a silicone tube is inserted. A diffusional

13.4 Current Therapeutic Applications

process slowly releases the BCNU solution through the silicone tube. Upon completion of delivery, the device is refilled with drug solution. The Vitrasert implant is yet another interesting example, delivering the antiviral drug, ganciclovir, and indicated for the treatment of CMV retinitis in patients with acquired immunodeficiency syndrome (AIDS) (Wong et al., 2013).

13.4.2.5 Pain management Chronic pain is a particularly challenging disease state, because of the need for repeated dosage, morbidity and mortality from overdosing, and the high risk for addiction to oral and parenteral medications. Data from the CDC between 1999 and 2010 shows a reported rise in deaths from prescription pain drug overdoses of 400% among women and a 265% in men (CDC, 2013). IDDSs therefore offer some unique and potent solutions for managing chronic pain. Axxia Pharmaceuticals has a patented design for a subcutaneous, drug-eluting device with zero-order kinetics for continuous 3090 day delivery of hydromorphine. The device is aimed at treating patients suffering from cancer or HIV/ AIDS-induced neuropathic chronic pain. Similarly, the LiRIS program from TARIS Biomedical offers direct and prolonged delivery of lidocaine for treatment of interstitial cystitis/bladder pain syndrome (IC/BPS). The implantable Synchromed Infusion Pump System for delivering baclofen for muscle spasticity from Medtronic may be also classified in this category. A specific response to the problem of opioid abuse is found in the Propuphine subcutaneous drug-eluting device, developed by Titan Pharmaceuticals (Titan, 2016). The implant is intended to deliver buprenorphine hydrochloride from an ethylene vinyl acetate matrix, for 6 months following implantation. Titan commercialized the product in 2012, as a result of an exclusive licensing agreement with Braeburn Pharmaceuticals.

13.4.3 INFECTIOUS DISEASES (TUBERCULOSIS) A fundamental problem in the treatment of tuberculosis (TB) is the long duration of therapy and side effects of drugs, which can hamper patient lifestyle and induce patient noncompliance, treatment failure, and development of drug resistant strains (Fears et al., 2010; Young et al., 2008). In this context, IDDSs are a promising approach for the development of more effective and more compliant DDSs for TB treatment (Dong et al., 2014). Pre-clinical studies have demonstrated that a single implant of Isoniazid (INH) in PLGA copolymer could ensure sustained levels of free INH for a period of up to 8 weeks, following implantation in rabbits (Kailasam et al., 1994). Gangadharam et al. (1999) have also investigated the chemotherapy of TB in mice, using single implants of INH and Pyrazinamide (PYZ). Such devices, however, inherently suffer from the disadvantages of immobilization at the implantation site and surgical requirements for implantation.

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13.4.4 NEUROLOGY AND CENTRAL NERVOUS SYSTEM HEALTH Schizophrenia is a mental health condition wherein approximately 50% of patients under treatment are noncompliant (Llorca, 2008; Ward et al., 2006). For such patients, systems capable of controlled and sustained delivery antipsychotics are of great value (Rabin and Siegel, 2012). One such system available commercially is the subcutaneous MedLaunch Implant Program developed by Endo Pharmaceuticals, for delivery of risperidone (Schwarz et al., 2012). Braeburn is the commercialization partner for this technology as well, similar to the Probuphine device mentioned earlier.

13.5 CURRENT CHALLENGES AND FUTURE PERSPECTIVES 13.5.1 BIOCOMPATIBILITY-RELATED ISSUES Biocompatibility is a critical design parameter for successful IDDS performance, given the need for a substantial time period of implantation. Since most IDDS are in close proximity to critical internal organs and in intimate contact with tissue, biocompatibility with the human environment is essential and indispensible. Unfortunately, it is very unlikely that any synthetic material will be completely inert or harmonious with the living environment. For a substance to be considered as biocompatible, it must fulfill certain requirements. All implant materials, not including the drug cargo, must be chemically inert, noncarcinogenic, hypoallergenic, and mechanically stable at the implant site (Williams, 2015). The material should also not be physically or chemically modified by local tissue, and the implant should not cause any inflammatory response at the site of implantation. As per WHO, all implantable materials should pass different tests for stability and biocompatibility (Guidance for Industry and Food and Drug Administration, WHO, 2016). Specific issues associated with post-implantation stability and reactivity need to be investigated, such as the formation of fibrous capsules around the implant. Similarly, erosion-based devices need to be thoroughly evaluated for possible toxicity or immunogenicity of the byproducts of polymer degradation. If acceptable biocompatibility is not achieved, many adverse effects may occur. These may range in severity and intensity from benign consequences, such as capsular contracture, to relatively serious complications, such as unexpected release of the drug, platelet adhesion, inflammation, fibrosis, thrombosis, tissue damage, or infection of the area surrounding the implant (Flamant et al., 2016; Nuss and Von Rechenberg, 2008; Park and Park, 1996; Pavithra and Doble, 2008). Biomaterial implantation is always accompanied by some unavoidable injury as a result of the implantation procedure. This initiates various responses at the tissue and molecular levels (Anderson et al., 2008), as shown in Fig. 13.5. After the first contact with tissue, proteins from blood and interstitial fluids adsorb to

13.5 Current Challenges and Future Perspectives

FIGURE 13.5 Various biological responses against implants.

the biomaterial surface and form the layer of proteins which determines the activation of coagulation cascade, complements system, platelets, and immune cells (Franz et al., 2011; Williams, 2008). Indeed, auto-activation of Factor XII (FXII) is catalyzed by surface contact with biomaterials (De Maat and Maas, 2016). Although auto-activated coagulation Factor XII on biomaterials initiates the generation of thrombin, the amount produced is not sufficient to induce clot formation (Gorbet and Sefton, 2004). Blood coagulation on biomaterials has been recently demonstrated to require the combination of both contact activation and platelet adhesion/activation (Sperling et al., 2009). Thrombin is one of the main activators of platelets. Low concentrations of thrombin released by FXII on biomaterials are suggested to activate platelets, which, in turn, release mediators of the coagulation process and expose negatively charged phospholipids, thus providing the supposed catalytic surface for the coagulation cascade (Franz et al., 2011; Vieira et al., 2007). Furthermore, thrombin initiates the cleavage of fibrinogen to fibrin, forming the primary fibrous mesh around the biomaterial (Fig. 13.5). Tissue response to synthetic implants can vary from material to material. Degradation products and surface changes of the biomaterial that occur due to the degradation process, additionally affect the immune response for biodegradable IDDS deployed in vivo (Anderson et al., 2008). There are many factors to consider during the development of an IDDS which influence implant biocompatibility, such as implant size, shape, material composition, and surface wettability, roughness, and charge. Both the physical and chemical properties of the bulk

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material may be involved in provoking a particular tissue reaction, and soluble breakdown products may induce their own local tissue response. Implant size also has marked effect on tissue response (Williams, 2008). It is also important to remember that the drug itself can have important effects on the biocompatibility of a DDS, particularly for formulations that involve a stationary depot. This is illustrated very well in the case of the sustained release of local anesthetics. For example, a wide range of DDSs, including polymeric particles, spray-dried lipid-protein-sugar particles, liposomes, cross-linked hyaluronic acid gels, and rheological blends of polysaccharides have been employed to provide sustained release of bupivacaine. In the absence of drug, these formulations cause varying degrees of inflammation, but minimal or no actual tissue injury. However, when loaded with bupivacaine, all have caused varying degrees of muscle injury (Padera et al., 2008, 2006). Finally, testing for biocompatibility often relies upon animal models with imperfect correlation to human beings. A material cannot, therefore, be conclusively shown to have the same tissue reactions in humans as those demonstrated in animals. In addition, genetic differences within human populations may account for the fact that some patients will have adverse reactions even to materials that have no adverse effects on the vast majority of patients. Modern implant design is therefore directed towards coopting this immune response to improve implant integration, while avoiding its perpetuation leading to chronic inflammation and foreign body reactions, and thus loss of the intended function.

13.5.2 PATIENT COMPLIANCE Some IDDSs require minor surgery for implantation and extraction, which lowers patient acceptance and demands a less invasive alternative. Implanting (and potential explanting) may require specialized training, in most cases where implants are designed for administration by lower level healthcare workers instead of surgeons. Patient acceptance is typically lowered if the IDDS exert pain and discomfort while implanted in the body (Patel and Brandstetter, 2016; Zernotti et al., 2012). However, in most cases where implants are indicated, the potential cumulative benefits of the prolonged therapy, minimal ongoing compliance burden, and optimized or localized delivery far outweigh the potential risks. As such, most patients in need of such therapy may be suitably counseled to comply with the temporary morbidity associated with implantation and extraction, if needed.

13.5.3 REGULATORY ASPECTS With respect to IMDs, the evolution of technologies from mere electromechanical devices to ones with more advanced computing and communication capabilities have many benefits, but also entail numerous security and privacy risks for the

13.5 Current Challenges and Future Perspectives

patient. The majority of such risks are relatively well known in classic computing scenarios, although, in many respects, their repercussions are far more critical in the case of implants. Remote cyber attacks against a telemetrically connected IMD can put at risk the safety of the patient who carries it, with fatal consequences in certain cases. Causing an intentional malfunction of an implant can lead to death. As recognized by the U.S. Food and Drug Administration (FDA), such deliberate attacks could be far more difficult to detect than accidental ones (FDA, 2013). Furthermore, these devices store and transmit very sensitive medical information that requires protection, as dictated by European and US Directives (e.g., Directive 95/46/ECC; CFR 164.312).

13.5.4 COST-EFFECTIVENESS The inherent complexity of implantable systems, particularly in comparison with simpler oral dosage forms, potentially makes the vital and necessary regulatory path for approval longer and prohibitively expensive for some companies to pursue. The cost-benefit ratio is a vital parameter influencing patient acceptance and commercial success. It is possible that, after considering the costs for development, manufacturing, and regulatory approval, the IDDSs may not be cost effective enough for reimbursement, insurance coverage, or out-of-pocket purchase in uninsured markets. This, in combination with the explosion of new and potential treatments resulting from recent discoveries of bioactives, gene therapies, etc., has led to the challenging problem of selection of the mode in therapy and costeffective product development. Major pharmaceutical companies will usually not undertake the cost involved, unless reimbursement is highly probable or is negotiated beforehand. This leaves smaller companies to accept the risk and expenses incurred during the development and approval process. Nonetheless, as reviewed earlier, there are multiple clinical situations where implantable systems are not only cost-effective, but may substantially lower the overall cost of treatment when subjective factors like patient compliance and satisfaction are accounted for. This is perhaps a significant reason why both large medtech multinational companies, as well as small startups operating in niche areas, find it profitable to compete for the same market.

13.5.5 FUTURE PERSPECTIVES Overall, there is an ongoing a drive to make IDDSs more cost-effective and patient-friendly. This trend indicates that future devices will probably be smaller, less invasive, and more site-specific. All these features will need to be accomplished while maintaining the dose at precise therapeutic levels for the desired duration. Micro/nano-fabricated devices lend themselves precisely to this task, since manufacturability is well established by borrowing concepts from the semiconductor industry. Some IDDSs may even be hybrids that contain a targeting

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therapeutic molecule and include a diagnostic feature, creating a new class of so-called next generation theranostics. Future versions of IDDSs may additionally have a remotely controllable feature, whereby a physician located at considerable distance may precisely control their operation. This may enable them to become even more patient-friendly and increase their long-term cost-effectiveness. However, it certainly appears that the key drivers in the future will be IDDS applications that are more cost-effective while improving patient compliance and acceptance. Most of the electromechanical IDDSs use batteries, which must be changed frequently by surgical procedures, or require the power source to remain externally placed outside the body. To overcome these problems, Dagdeviren (2016) recently developed a flexible micro-generator that can harvest energy from the natural movements (contraction and relaxation) of organs to power implantable devices. Consequently, it is possible that, in the near future, we may see IDDSs with micro-bionic dynamos to power the next generation of dynamic drug delivery implants. However, further investigation into the safety profile of these types of systems is needed to evaluate whether they could be safely integrated into drug delivery implants. Environmentally sensitive polymeric delivery systems are currently designed to achieve targeted and controlled in vivo delivery in response to specific stimuli, such as pH, ionic strength, enzyme-substrate, magnetic, thermal, electrical, ultrasound, etc. (Fu and Kao, 2010). It is expected that implantable system will leverage such materials to generate a feedback-controlled release mechanism to modulate zero-order or nonzero-order release profiles. It is hoped that, in the future, development of new implantable systems will help reduce the cost of drug therapy, increase the efficacy of drugs, and enhance patient compliance.

13.6 CONCLUSION Development of new drug candidates is expensive and time consuming. Improving the safety-efficacy ratio of “old” drugs has been attempted, using different methods such as individualizing drug therapy, dose titration, and therapeutic drug monitoring. Delivering drug at controlled rate, slow delivery, targeted delivery are other very attractive methods and have also been pursued vigorously. IDDSs have seen reasonable clinical and commercial success as a mode of enhanced drug therapy. However, optimization of performance characteristics, including long-term biocompatibility and drug release kinetics is critical. Furthermore, clinical validation of current systems under development is essential for regulatory approval and their commercial success. However, as reviewed here, numerous commercial systems are able to attain nearly ideal zero-order release

References

kinetics profiles in vivo, over extended time periods. IDDSs therefore present a viable, cost-effective and clinically acceptable alternative route of sustained drug delivery for chronically ill patients.

ACKNOWLEDGMENTS Authors are thankful to Dr. Subham Banerjee, Young Scientist, THSTI, for his expert comments and valuable suggestion regarding the drafting of the manuscript. First Author, Anoop Kumar is also thankful to Shri. Parveen Garg, Chairman, ISFCP, for providing research facility.

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14

Nanobionics and nanoengineered prosthetics

Hemant K.S. Yadav, Ghufran A. Alsalloum and Noor A. Al Halabi RAK Medical and Health Sciences University, Ras al Khaimah, United Arab Emirates

CHAPTER OUTLINE 14.1 14.2 14.3 14.4

Introduction .................................................................................................514 History ........................................................................................................517 Definition.....................................................................................................520 Types and Classifications .............................................................................523 14.4.1 Orthopedic Prostheses .............................................................. 524 14.4.2 Plastic and Reconstructive Prostheses........................................ 526 14.4.3 Neuroprostheses....................................................................... 526 14.4.4 Cerebrospinal Fluid Drainage Systems........................................ 527 14.4.5 Ophthalmic Prostheses ............................................................. 527 14.4.6 Cardiovascular Prostheses ......................................................... 527 14.4.7 Myoelectric Prostheses ............................................................. 528 14.4.8 Dental Prostheses..................................................................... 528 14.5 Manufacture ................................................................................................529 14.5.1 Lithography.............................................................................. 529 14.5.2 Photolithography ...................................................................... 530 14.5.3 Beam Lithography .................................................................... 530 14.5.4 Micro and Nano Contact Printing ............................................... 530 14.5.5 Jet Printing .............................................................................. 531 14.5.6 Scan Probe Lithography ............................................................ 531 14.5.7 Dip-Pen Nanolithography .......................................................... 532 14.6 Nanobiomaterials .........................................................................................532 14.6.1 Polymeric Materials .................................................................. 533 14.6.2 Nanotitanium (NanoTi).............................................................. 535 14.6.3 Carbon Nanotubes .................................................................... 536 14.6.4 Nanodiamonds ......................................................................... 539 14.6.5 Nanobioceramic ....................................................................... 540 14.6.6 Nanocomposite ........................................................................ 543 14.6.7 Peekpolymer ............................................................................ 549 14.6.8 Hydrogel .................................................................................. 549


513

514

CHAPTER 14 Nanobionics and nanoengineered prosthetics

14.7 Applications ................................................................................................550 14.7.1 Orthopedic Prostheses .............................................................. 550 14.7.2 Neuroprostheses....................................................................... 566 14.7.3 Cardiovascular Prostheses ......................................................... 571 14.7.4 Cerebrospinal Fluid Drainage Systems........................................ 572 14.7.5 Plastic and Reconstructive Prostheses........................................ 573 14.8 Ethical Issues ..............................................................................................574 14.8.1 Safe use: benefits Versus Risks.................................................. 575 14.8.2 Justice .................................................................................... 575 14.8.3 Identity, Privacy, and Accountability........................................... 575 14.8.4 Autonomy ................................................................................ 576 14.8.5 Validity of Informed Consent...................................................... 576 14.8.6 Problems of Ambition: Treatment Versus Enhancement................ 577 14.9 Safety Issues Pertinent to Nanobionics and Prosthetics..................................577 14.10 Conclusion ..................................................................................................579 References .............................................................................................................580 Further Reading ......................................................................................................587

14.1 INTRODUCTION Since the beginning of humankind, nature has grabbed the attention of human minds with its intelligent systems and perfect balance. Therefore, scientists have been thoroughly exploring the biosystems of nature and its mechanical and dynamic mechanisms in order to find a way of mimicking those systems. One the most popular examples of a scientist and an artist who tried to mimic nature in his inventions was Leonardo Da Vinci, whose investigations were mainly focused on the mechanism of movement of birds and their way of flying. Even though it is not clear whether those investigations resulted in Da Vinci actually taking flight or not, it is certain that he managed to successfully produce numerous sketches and notes on the anatomy of birds, the biophysics of flying, and the structural design of flying machines. With a lifetime lasting 67 years, starting from 1452 to 1519, Leonardo Da Vinci is considered one of the earliest founders of bionic research (Dickinson, 1999). It can be inferred from this, that the science of bionics is the science of studying the mechanical aspects of biosystems in nature and applying them to many fields of modern technology. In addition, it is considered a combination of biology and electronics that involves engineering devices that mimic the function of biological systems. When speaking of biology, the first structure that comes to mind is the human body, as it is one of the most magnificent creations in nature. Therefore, bionics are regularly linked with artificial organs and engineered body parts. This branch of bionics can be termed as medical bionics. Merging man

14.1 Introduction

with machine has produced stunning outcomes in curing diseases and mending defects. For instance, cochlear implants have been successful in restoring hearing by stimulating the auditory nerve; similarly, the bionic eye can give vision to the blind by stimulating the optic nerve. Furthermore, neuroscientists believe that inserting tiny bionic implants would allow them to detect abnormal neural activity and then correct it by electronic stimulation, thus producing a solution for mental disorders like epilepsy and Parkinson’s disease. Another very popular and outstanding achievement of medical bionic engineering is the manufacture of fully functional artificial limbs (Conroy, 2011). Since ancient times, handicapped soldiers have replaced their amputated extremities with artificial metal ones to help them fight in wars. Artificial extremities have come a long way since then, particularly since nanotechnology has been incorporated into the manufacturing process (Norton, 2007). Nanotechnology, a novel concept originally introduced by Richard Feynman in his lecture There is plenty of room at the bottom in 1959 (Feynman, 1992), is the branch of science that deals with material at the molecular level; or in other words, at the nano scale, which is one billionth of a meter. Nanotechnology has found its way into a variety of sciences such as biophysics, molecular biology, biomedical engineering and, most importantly, medicine. When the dimensions of a specific moiety are reduced to less than a micro, they exhibit exceptional characteristics and properties that facilitate their utilization in a variety of applications. Some of these applications have already reached the market and have achieved huge profit and great benefit for the public (Jain, 2007). Some of the most successful implementations of nanotechnology are: •

• • • • • • •

soft tissue repair and healing using wound dressings that are manufactured from nanofibers, as well as topical skin care products, like sunscreens, that contain lipid nanoparticles implants and prostheses, including plastic surgery procedures tissue engineering nerve and bone regeneration drug delivery cancer treatment (Ver Halen et al., 2014) gene delivery oral vaccine formulations.

Nanotechnology is utilized in the medical and pharmaceutical fields via two basic nanotools: nanodevices and nanomaterials. Nanodevices, including nanoelectromechanical and microelectromechanical systems, microarrays, and microfluidics, are used in biosensors, bioactuators and detectors. Conversely, nanomaterials, which are subclassified into nanocrystalline and nanostructured materials, are employed in tissue engineering, drug encapsulation, bone replacement, implants, and prostheses like neuroprosthetics, artificial organs, and maxillofacial prosthetics (Fig. 14.1; Jain, 2007).

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FIGURE 14.1 Classifications of nanotools and their applications.

The greatest ambition of bionics researchers and scientists is to engineer artificial appendages with the same anthropomorphic and mechanical properties as that of the human body. Even though medical bionic engineering has made tremendous improvement since its wooden peg-leg days, it still endures a lot of obstacles and difficulties. However, with the recent advances in nanotechnology and its application in organ and tissue engineering, as well as orthotics and prosthetics, bionic engineers have found many answers to their queries. Nanobionics is the term used when nanotechnology is utilized in the engineering and the manufacture of bionics. Extensive research is being undertaken to produce prosthetics with human-like performance and good stability by utilizing human-machine neural interfaces, muscle-like actuators, and biomimetic humanoid control schemes (Herr et al., 2003). Aside from mechanical stability, nanotechnology provides fast integration, high fatigue resistance, enhanced durability, great reactivity and optimal design to the prostheses. These beneficial qualities occur due to increased miniaturization of the building components of the prostheses, reduction of limb weight, and improved biocompatibility of implant. Some examples of successful nanobionic products are: • • • •

artificial digestive tract organs including artificial sphincter, artificial esophagus, peristalsis stent, etc. artificial myocardium. control units of central nervous system function. detection of baroreflex reactivity in blood vessels (Yambe, 2009).

Tissue engineering for the purpose of manufacturing prostheses or implants is drawing the attention of scientists and engineers, as it holds endless possibilities. Nanostructures are used to make scaffolds that are structurally similar to living

14.2 History

tissue, which are then used in designing and producing implants or prostheses exhibiting ideal characteristics of the model organ, consequently, allowing the newly implanted organ to function in a safe and natural way inside the body. Different types of nanocomposites are used, depending on the system or the tissue that is to be replaced (SalaheldeenElnashaie et al., 2015). In this chapter, we will focus on various aspects of nanobionics in general, with a focus on nanoengineered prosthetics in particular. These aspects involve manufacture, design, evaluation, safety, and ethical issues as well as many others. Moreover, distinct applications will be discussed including: • • • • • • •

orthopedic prostheses and ligament prostheses cardiovascular implants neural implants and cerebrospinal fluid (CSF) drainage plastic and reconstructive implants dental implants ophthalmic systems in addition, a variety of other modern utilizations.

14.2 HISTORY Many of the most successful human inventions are based on knowledge derived from the natural world. Even nowadays, whenever engineering problems occur, people return to nature for inspiration and guidance. The history of bionics starts with the ancient warriors who used to replace any amputated limbs with artificial ones, made of metal, to help them fight in wars (Kakade, 2006). However, progress in this field has been slow and gradual. Early attempts at bionic engineering started with Leonardo Da Vinci’s sketches and designs of flying machines inspired by the anatomy of birds. Then, almost 300 years later, a German scientist, Otto Lilienthal, managed to fly his gliding machines, which were patterned after birds (Dickinson, 1999). Lilienthal introduced the term biomimetics, which is synonymous with the term bionics, meaning the imitation of natural mechanisms and processes to improve human technology (Schmitt, 1969). Apart from aerodynamics, attempts at creating machines for the medical purposes, that mimic humans instead of animals started as early as the 16th century. When it comes to medical bionics, prosthetic limbs are the earliest attempts to create a biomimetic artefact. Even though hooks and pegs were the only prosthesis used in the Middle Ages, the Renaissance inspired more sophisticated and natural-looking hands, made of wood or metal. In 1504, a German knight called Go¨tz von Berlichingen used a pair of iron prosthetic hands during a battle, which enabled him to continue fighting in wars until he reached the age of 64. What was special about this pair is that it was the first pair of hands to ever have

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flexible finger joints (Frumento et al., 2010). Advancements in medical bionics completely ceased until the 19th century, when contact lenses were first developed for vision correction (Siviglia, 2010). After less than 20 years, early attempts by the French Judet brothers, Robert and Jean, were made to invent a functional artificial hip replacement. However, it failed due to its exceptional susceptibility to corrosion (Gomez and Morcuende, 2005). Fast forward a few years, an artificial iron lung was built for the treatment of polio victims. The iron lung helped in saving thousands of polio sufferers from respiratory paralysis (Emerson, 1958). Later on, in the early 20th century, the first successful kidney dialysis machine was developed for the treatment of renal failure (Blagg, 2007). The years between the 1950s and the 1990s were years of firsts, and years of great importance for medical bionic advancements; a lot of artefacts were created, including the first artificial heart valve, the heart-lung machine, the first cochlear implant, the first auditory brainstem implant, the first successful single channel cochlear implant in a child, the first permanent total artificial heart (Jarvik-7), and the first clinical application of a bioartificial liver device. Moreover, with the start of the 21st century, bionic engineers succeeded in implanting a prototype artificial pancreas, as well as a permanent self-contained total heart replacement (AbioCor) (Historical Highlights in Bionics and Related Medicine, 2002). The science and engineering of prosthetics was, and still is, continuously progressing alongside the aforementioned improvements in medical bionics. An old woman from Cairo was the person that held the first functional prosthesis, which was discovered in a tomb found in the vicinity of the ancient city Thebes, in 2000. The prosthesis was a toe attached to a mummy of a 5060-year-old woman, having three joints, and displaying signs of damage (Frumento et al., 2010). In addition, an artificial below-knee prosthesis was found in Capua, Italy (Norton, 2007). The artificial leg had a bronze and iron body with a wooden core. The engineering of prosthetics continued to improve using the same basic material, but more evolved springs and releases instead of joints. This advancement is depicted in the iron hands of Go¨tz von Berlichingen. After losing his right arm in the Battle of Landshut, the knight attached an artificial hand that could be moved by a system of releases and springs, and could remain suspended with leather straps. A well-known surgeon, called Ambroise Pare, was capable of creating similarly functional hands. However, lower limb prosthesis was not developed until the end of the 17th century, the year of 1696, when Pieter Verduyn, whose designs became the blueprint for joint and corset devices, made the first nonlocking below-knee prosthesis. No major developments in this field occurred until the late 1800s, when Gustav Hermann used aluminum instead of steel in formulating an artificial limb that was more lightweight and practical (Frumento et al., 2010). Throughout the history of prosthetics, battles have been the major stimulant of progress, with most of the developments occurring at times directly after or during wars. Step after step, a compilation of techniques and principles led to the modern technology we currently have. Thanks to all these advancements in

14.2 History

prosthetics, people can finally have aesthetic-looking artificial limbs that don’t set them apart from others, and give them a sense of fullness and improved their quality of life (Harvey et al., 2012). A popular example of a successful double leg amputee, who has a pair of artificial legs that allowed him to participate in Olympics, is Oscar Pistorius. Known as Blade Runner, Oscar Pistorius is a world record holder in 100, 200, and 400 m Paralympic events and the first amputee to ever participate in nondisabled Olympic games and win. Modern technology has caused great improvement in prosthesis functionality and durability. These technologies include microprocessors, computer chips, robotics, and nanotechnology. Moreover, the materials used in manufacturing artifacts are no longer limited to metals; silicone is now used to provide realistic, natural-looking covers, in addition to various other materials (Norton, 2007). Ceramics are used for manufacturing prostheses for dental, orthopedic, and musculoskeletal applications. Metals are still used to sustain the artifacts with endurance and durability. Metallic-based materials are mainly used for orthopedic prostheses, in order to allow them to endure rigorous activities (SalaheldeenElnashaie et al., 2015). Nanotechnology is a multidisciplinary concept that started less than 60 years ago, but has found its way into every aspect of daily life. Applications of nanotechnology are found in everything, from food industry to medicine, including bionics and biomedical engineering. The early days of nanotechnology involved direct manipulation of atoms and molecules in order to synthesize specific materials. The United States National Nanotechnology Initiative classified nanotechnology into four generations (Fig. 14.2):

FIGURE 14.2 Generations of nanotechnology developments with examples.

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• • • •

First generation (B2000): is the formulation of passive nanostructures that can only perform a single task Second generation (B2005): is the formulation of active nanostructures that can perform multiple functions Third generation (B2010): involves thousands of components that interact together in a network or nanosystem Fourth generation (B2015): involves molecular nanosystems.

Nanotechnology achieved huge advancements in the biomedical field, especially in medical implants and prosthetics, because it utilizes particles having the size of nanometer that are similar to that of living cells. While designing and manufacturing prostheses, the main aim is to create scaffolds identical to natural tissue and to produce prostheses capable of functioning naturally in the body environment (SalaheldeenElnashaie et al., 2015). Apart from artificial limbs, nanotechnology is now being applied in the development of bionic eyes, bionic ears, neuroprosthetics, as well as cardiovascular, dental, and orthopedic prosthetics (Fig. 14.3; Baumann, 2013).

14.3 DEFINITION In order to understand the term prosthetics, we need to elaborate on a broader term: medical devices. A medical device is an article, that could be an instrument, apparatus, material, or an appliance, which is used for the purpose of diagnosis, prevention, treatment, alleviation and/or monitoring of a disease, compensation

FIGURE 14.3 Applications of nanotechnology in prosthetics.

14.3 Definition

for a handicap, control of conception, and modification or replacement of anatomy or physiology. Medical devices, typically, achieve their intended function by physical means, such as mechanical or physical action and replacement of or support to bodily organs or functions. In order to decide whether an implant is considered a medical device or not, different criteria should be explored, such as the time or duration of contact with the patient, the location affected by the use of the device, and the degree of invasiveness. An invasive device is defined as a device that, completely or partially, penetrates the body, either through a natural orifice or through the surface of the body. Implants are considered medical devices because they are surgically invasive, and also because they remain inside the body long term or permanently. Implantable devices must reside in the patient for a specific duration, not less than 30 days, after a procedure. For example, a nontunneled catheter used for temporary vascular access that is intended to be used not more than 710 days is not considered a long-term implantable device. Nor is a surgical suture that is taken out before 30 days considered an implant. Implantable devices serve a number of purposes, according to which they can be classified into different groups. Implants are imported into the body for different purposes, to serve as: • • • •

prosthetics, to replace missing body parts orthotics, to support weak or ineffective joints or muscles monitors, to keep track of bodily functions drug delivery systems, to administer medication (European Commission, 2001).

In simple terms, prosthetics is the science of developing artificial body parts and then surgically replacing the amputated body parts with the artificial ones. Prosthesis is the artifact used to replace the missing parts and restore bodily functions. It should be noted that prosthetics is a different field from orthotics, although they are related. Orthotics is a branch of medicine concerned with supporting weak joints and muscles using special artifacts like braces and splints. When an individual is fitted with an artificial prosthesis, he or she is labeled with the name “Cyborg,” meaning part-human part-machine. Since the start of the 21st century, a large number of implants and devices were developed to restore the function of many human biological organs or systems. These prostheses include: • • • • • • •

artificial extremities, both robotic ones and ones with sensory abilities artificial polymer-made muscles artificial skin, with healing promotion abilities artificial hips, joints, and vertebrae artificial bone, intended to help in healing fractures and replacing defects artificial teeth and dental implants cervical implants and bracing systems, intended to provide support to the spine

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• • • • • • •

silicone or plastic implants for cosmetic and maxillofacial reconstruction artificial larynxes for speech restoration and speech processor wired to the nervous system artificial blood vessels retinal implants, intraocular lenses, and artificial cornea for vision restoration cochlear implants neuroprosthetics and spinal-neuro implants many others that will be discussed later on in this chapter (Brey, 2005).

In a broad sense, nanotechnology is a general-purpose technology, based on reducing the particle size of materials to the size of the nano in order to produce new and improved properties and characteristics. The ultimate goal of prostheses is to ensure biocompatibility and functionality, both of which can be enhanced by the utilization of nanostructures and nanodevices in the manufacture of the implant. In addition, multiple benefits can be gained by using nanotechnology in prosthetics, including enhanced mechanical stability and durability, fast integration and low roughness of nanoengineered surfaces, as well as greater strength and high resistance to wear and fatigue (Fig. 14.4). Furthermore, the incorporation of nanotechnology-based tissue engineering techniques can greatly improve the biocompatibility and stability of implantable devices and prostheses (Torrecillas et al., 2009).

FIGURE 14.4 Beneficial effects of nanotechnology on prosthetics.

14.4 Types and Classifications

The application of nanotechnology in implantable devices is considered a subdivision of Nanomedicine. Nanomedicine is basically the branch of nanoscience concerned with the utilization of nanotechnology in medical devices and for medical purposes (SalaheldeenElnashaie et al., 2015). Relatively, nanobionics is the field that integrates both nanoscience, or more particularly nanomedicine, and biomedical engineering for the purpose of creating fully functional bionic implantable devices. The term bionic comprises bio- from biology and -nic from electronic. Nanobionic devices are electronic artifacts that are manufactured to imitate, mimic, and restore bodily functions (Herr et al., 2003). In summary, this chapter focuses on nanoengineered prosthetics, which are implantable devices that are intended to replace a missing body organ and that utilize nanotools and nanostructures in their constitution to gain exceptional and superior properties.

14.4 TYPES AND CLASSIFICATIONS According to the United States Food, Drug and Cosmetics Act (FD&CA), implantable medical devices are classified into three classes: •

•

•

Class I: General controls must be applied to this class and they state that a medical device from this class cannot be adulterated, misbranded, subject to recall, or a banned device. Moreover, the firm must be registered with the Federal Drug Administration (FDA), must list its devices with the FDA, must maintain the required reports and records, and must apply good manufacturing practices. Class II Special controls, like performance standards, postmarket surveillance, patient registries and guideline, are the control procedures applied when general controls are not sufficient to assure the safety and effectiveness of a device. Class III When neither general controls nor special controls are sufficient to ensure the safety of a medical device, premarket approval is the type of control applied. A premarket approval application must contain information about the safety and effectiveness of the device, information about the components, ingredients, and properties and of the principles of operation, of the device, information about the manufacture, processing, packaging, and installation of the device, and references, labels and clinical trial certification, as well as other relevant information.

According to their purpose, implantable devices are divided into four groups. Some implants are prosthetics, introduced into the body to replace missing body parts, while some are orthotics, intended to support weak or ineffective joints or

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FIGURE 14.5 Types of implantable devices.

muscles. Some implants are inserted into the body to monitor the internal environment and to keep track of bodily functions. One more type of implants is implantable drug delivery systems; these are used to deliver drugs or hormones (Fig. 14.5; European Commission, 2001). Since the start of the cyborg era, many prostheses were invented for cardiac, cosmetic, orthopedic and other applications. However, only some artifacts have implemented nanotechnology into their manufacture and design. There is no specific system that can be followed to classify nanoengineered prostheses, therefore, they are classified according to the organ they replace or support (Fig. 14.6; Baumann, 2013).

14.4.1 ORTHOPEDIC PROSTHESES In the market of prosthetics, orthopedic implants take up the largest segment of market value. This is a reflection of the unhealthy sedentary lifestyle of a large portion of the society and the rapidly growing prevalence of degenerative musculoskeletal disorders. Major disorders of the bones and joints include osteoarthritis and rheumatoid arthritis, both of which can cause considerable damage to the synovial joints like the hips, knees, shoulders, ankles, and elbows. Damage to these joints, especially the weight-bearing ones like the hips and knees, elicits excruciating pain. Accordingly, when the possibility of replacing the damaged joints with prostheses became a reality, it was a relief for patients all over the world. Arthroplasty, the process of inserting an artificial joint for the purpose of treating joint defects, relieving pain, and restoring movement, is a major advancement in the field of orthopedic surgery. It is based on steps like the excision,


FIGURE 14.6 Types of nanoengineered prostheses.

interposition, and the replacement of diseased bone or cartilage. For example, a total knee arthroplasty involves replacing the ailing cartilage of the femur, the tibia, and the patella with a metallic or polymeric prosthesis (Khan et al., 2013). Orthopedic implants have gained huge benefit from nanotechnological advancements. Applications of nanotechnology in this field involve using nanograined materials, including metal alloys, polymers, and nanoceramics like alumina, titania, and hydroxyapatite. Nanoceramics demonstrate better mechanical stability, enhanced osteoblast adhesion, increased calcium deposition, and improved in vitro osteoblast proliferation, when compared to their microsized counterparts (Teoh et al., 2014). In addition, coating the implants with crystalline calcium phosphate nanoparticles exhibited increased contact between bone and implant (Deshmukh et al., 2010). Since most of the cellular activities involved in implanting a prostheses, such as cell attachment, locomotion, growth, gene expression, and stem cell differentiation, occur at the nano scale, incorporating nanodevices and nanotools in the process of implantation will increase the chances of prostheses success. Examples of nanoengineered orthopedic prostheses are knee and hip joint prostheses, spinal implants, bone fixators, and tendon and ligament prostheses.

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14.4.2 PLASTIC AND RECONSTRUCTIVE PROSTHESES Another field that has benefitted greatly from nanotechnological advancements is plastic and reconstructive surgery. It is a very diverse field, involving aesthetic surgery and injectable collagen for soft tissue augmentation, oncologic and congenital reconstruction such as breast augmentation, and craniofacial or maxillofacial reconstruction, burn wounds and trauma care, and various other surgical and nonsurgical procedures. Miscellaneous nanotools are used in the field of plastic surgery, such as nanofibers, nanoparticles, nanocomposites, in addition to other nanostructures. Nanofiber matrices have been successfully used for constructing and repairing a variety of tissues, such as muscle, cartilage, bone, and skin, which are needed for plastic surgery procedures in both in vitro and in vivo conditions. Currently, reconstructive aesthetic surgeons are using nanoengineered cartilage tissue for ear reconstruction or nasal reconstruction, as well as bone and skin tissue for craniofacial reconstruction after congenital defects, trauma, and cancer. Nanofibrous scaffolds, made of polylactic and polyglycolic acid and ingrained with growth factor, are used to formulate artificial skin for the treatment of skin defects by stimulating and enhancing the skin healing process (Ver Halen et al., 2014).

14.4.3 NEUROPROSTHESES Neuroprosthetics is a rapidly expanding field that is greatly dependent on computer technology and consequently, it will surely be broadened with the application of nanotechnology. In simple terms, neural prostheses are artifacts that replace or repair neural function through electronic interfaces such as neuromuscular electrical stimulation for improving strength and fatigue resistance, and functional electrical stimulation for stimulating paralyzed limbs to execute motor functions. In other words, neuroprostheses are electronic devices used for stimulating nerves to do bodily functions lost due to damage or trauma (Prochazka, 2009). With the application of nanotechnology to the cell-electrode interface in implantable devices, human cognitive function can be improved in patients with neurological disease like Parkinson’s and Alzheimer’s disease, both of which are caused by damage to neural and cognitive function. Nanotools also increase the lifetime of the implant, thereby enhancing its quality and avoiding additional surgeries and procedures for replacing it. In addition, restoration of hearing via cochlear implants, and sight via retina implants are huge applications of neuroprosthetics that benefit greatly from nanotechnology (Wolpe and Wu, 2006).


14.4.4 CEREBROSPINAL FLUID DRAINAGE SYSTEMS The CSF flows in the space between the brain and the inner lining of the skull, then it is drained to the blood via sinuses. When the outflow of the CSF is obstructed, often by a tumor, pressure inside the skull increases due to fluid build-up. This is a condition called hydrocephalus, in which the increased pressure leads to compression of brain tissue, resulting in a number of symptoms like headaches, drowsiness, and fainting. The solution to this problem is to implant a CSF drainage system, or shunt. However, conventional shunts have a limitation of being prone to infections and blockages, which will lead to the patient requiring another surgery to replace the dysfunctional shunt. In order to make these shunts more functional, carbon nanotubes (CNTs) can be used in bundles to form a new catheter that acts as filter of the CSF to prevent bacteria and other macromolecules that can cause blockage of the shunt (Spiers et al., 2010).

14.4.5 OPHTHALMIC PROSTHESES Vision impairment due to ocular disease, such as diabetic retinopathy and macular degeneration, are serious problems that permanently damage the photoreceptor cells in the eyes. Retinal implants are a modern approach to fix these problems and restore vision by stimulating the retina to produce visual percepts. Nanophotonic devices are currently being used to provide vision with high resolution, along with improved biocompatibility and reduced power consumption, leading to reduced tissue damage. In addition, utilizing nanowires and nanosized surface modification techniques results in enhanced tissue integration (Hamsika et al., 2016) Another utilization of nanotechnology in ophthalmic devices like intraocular lenses, keratoprostheses, ophthalmic lenses, contact lenses, and drainages for glaucoma, is the manufacture of a special coating made of metabolically active material, such as platinum nanocoating, which aids in bioadhesion as well as providing a variety of unique properties (Babizhayev, 2013).

14.4.6 CARDIOVASCULAR PROSTHESES Various branches of cardiology have benefited from nanotechnology since the manufacture of the first pacemaker. Such applications include intravascular stents, transjugular intrahepatic stents, and portocaval stents, all of which are currently being evaluated for the use of nanobiomaterials to prevent in-stent restenosis, bioincompatibility, and other complications. Stents are used as standard procedure of percutaneous coronary intervention, also known as coronary angioplasty, for the management of atherosclerosis (Kong et al., 2006). Polymeric nanocoatings, absorbable stents, and nanocoated drug-eluting stents are also extensively studied and have been in use for several years (Keyhanvar et al., 2015).

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Another example of a major heart disease is valvular heart disease, in which a heart valve needs to be repaired or, in severe cases completely replaced. The replacement can be a prosthetic heart valve, which is mechanical, made of stainless steel or silicon, and has great durability, lasting a lifetime; or a bioprosthetic heart valve, which is extracted from humans or animals and requires replacement after a certain period. Nanotechnology offers the chance to bioengineer a more compatible and durable heart valve that can be even better than the normal valves. One more brilliant prosthetic device is the pacemaker, which is an electrical device intended to function like the sinoatrial node to control the speed of impulse generation, thereby controlling the rhythm of the heart. Utilizing nanodevices in pacemakers technology can result in more biocompatible devices with less chance of rejection or displacement, allowing the patients more freedom in their daily life and activities. Furthermore, the use of nanowires and nanochips in engineering artificial pacemakers is currently being investigated for better conduction and control of electrical impulses (Iqbal and Mohan, 2010).

14.4.7 MYOELECTRIC PROSTHESES After upper or lower limb amputation, a patient has the opportunity to choose between two types of prosthetic limbs. The first is passive prosthesis, generally used to serve cosmetic purposes, and is usually used for upper limb amputations. The other type is functional prosthesis, which allows the user to perform simple tasks and is divided into two types: body-powered and myoelectric-powered. Body-powered prostheses have limited functionality and are moved by the action of nearby muscles; however, myoelectric prostheses provide a functionality almost identical to that of the real limb. They act by amplifying the electronic signals coming from the muscles of the residual limb (Patel, 2012). The greatest limitation of functional prostheses is that they require a lot of energy to function, specifically, to drive its motor into motion. Chemically fueled prostheses make a better replacement for conventional body-powered ones, as they are capable of imitating the action of the actual muscle fibers. CNTs, when filled with hydrogen sulfate, have the ability to move in a fashion similar to that of muscle tissue. Because of a complicated mechanism of action, the nanotubes require only water and air as fuel for the artificial limb. Moreover, the CNTs provide the limb with a high degree of stress tolerance, even more than that of the normal skeletal muscle, indicating that utilizing nanotechnology in myoelectric prostheses can actually give the user higher strength and better abilities than other people (Ebron, 2006).

14.4.8 DENTAL PROSTHESES When it comes to the dental field, nanotechnology has many applications, including: • •

local nanoanesthesia hypersensitivity cure

14.5 Manufacture

• • • • • • • • •

tooth repositioning nanorobotic dentifrice dental cosmetics nanodiagnostics nanofillers nanoadhesives bone replacement dentition replacement therapy prosthetic implants.

Dental prosthetic implants are used for replacing missing teeth or root systems and for any defective soft or bony structures of the jaw and palate. Dental prosthetics or prosthodontics utilize nanotechniques in the development of unique surface characteristics with a precise topography and chemical structure that mimics the surface topography of the extracellular matrix, improving the chances of successful cell attachment, proliferation, and differentiation. Moreover, special agents like antibiotics and growth factors are incorporated into the implant to reduce the chances of infections or rejection of prostheses (Bhardwaj et al., 2013; Tomsia et al., 2011). Since tissue integration is the primary determinant of implant success, it was essential to develop modern techniques that promote osseointegration, mainly by surface modification and the application of bioactive nanocrystalline coating made of nanomaterials such as calcium phosphate. In addition, nanoceramics such as silicon carbide, alumina, and zirconia, are used in the formulation of dental implants (Pai et al., 2015).

14.5 MANUFACTURE Manufacturing of nanobionics and nanoprosthetics is a process of high precision to serve a specified application and stimulate a very particular area. Fabricating nanoscaled devices and gadgets with electrical circuits needs a flawlessly peculiar technique that can be applicable to wide range of materials; organic, inorganic or biomolecules. Before Dip-Pen nanolithography, many approaches were implemented to construct nanoengineered bionics, including multiple types of lithography and printing, explained as follows:

14.5.1 LITHOGRAPHY Lithography comes from the Greek words (lithos) which means stone and (graphein) which means write. In other words, it means stone-writing which designates the ancient technique of using stones after chemical treatment and modification as stamps. Therefore, we can define it as a method in which a particular pattern is transferred onto a solid surface in a miniaturized manner (Madou, 2012). It can be easily comprehended if we compare it to printing a picture on

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your computer. You design this picture and specify the measurements, dimensions and details, then give the order of printing. Then, an ink jet printer transfers the design and its specifications and copies it on a paper. In that case, the predesigned picture resembles the specific detailed pattern desired, and the technique of transferring the picture from computers to papers is lithography. The ink printed can refer to various means of transfer upon which lithography can be classified into photolithography, scanning beam lithography and other classes.

14.5.2 PHOTOLITHOGRAPHY This class indicates the use of light in lithography. Here, the substrate is covered with a thin layer of light-sensitive liquid, known as photoresist, and a patterned mask is applied. Afterwards, it is exposed to UV light, which alters the solubility of the substrate. When rinsed with a suitable solvent, the exposed area dissolves, the unexposed remains and subsequent chemical treatments and processing follows (Madou, 2012; O’Connell et al., 2015). Much advancement was made to generate specialized photolithographic techniques, like extreme-UV lithography and X-ray lithography, applying limited wavelengths. In spite of the high costs required and the difficulty in tailoring soft organic materials, it was considered the best technique of lithography prior to the invention of dip-pen nanolithography.

14.5.3 BEAM LITHOGRAPHY This type of lithography can be considered as an upgraded extension to photolithography, in which a fixated beam generates the pattern, without a mask or a photoresist film. The beam can exert two actions. It selectively either removes or deposits a material. The beams can be charged or uncharged. More commonly used are electron beams (e-beam) and ion beams (i-beam), which introduce high current density influencing electron-sensitive or ion-sensitive resists (Madou, 2012). The process is very slow and tedious, it take 24 h/cm2 to synthesize 20 nm structures. Yet, the outcomes possess a distinguished resolution and pattern integrity which suitably qualifies it for production of nanowires of neuron interfacing transistors, among other beneficial applications (O’Connell et al., 2015).

14.5.4 MICRO AND NANO CONTACT PRINTING This procedure can be described as an actual chemical stamping. A stamp (conventionally polydimethylsiloxane (PDMS)) is carved with intended pattern, and then an organic dye or metallic ink transfers this pattern on a solid surface, where they form covalent bonding. Generally, it produces microformats but it is

14.5 Manufacture

also able to produce them in nano ranges to a limited extent (around sub100 nm). They feature stamp size-controlled resolution, which means that the size of the stamp influences the resolution of the printing. However, the size is determined by (1) the accuracy of the stamp mold when carved, (2) the nano properties of the stamp material, (3) stamp distortion upon surface contact. Being relatively inexpensive has facilitated its use in much biomedical research (O’Connell et al., 2015).

14.5.5 JET PRINTING The term jet printing is very popular with documentations and publications, but due to its functional flexibility it was endorsed for biofabrication. It has the ability to directly deposit a unique ink convoluting living cells or subcellular components. What makes them flexible and versatile is the fact that the droplet size can range from microns (μL) to picoliters (pL). This contributes to printing of high resolution textures, but the resolution can be still altered by (1) properties of liquid material, (2) wettability of substrates, (3) jet nozzle diameter. It could be used widely if it wasn’t for the specific requirement of conductive substrates (O’Connell et al., 2015).

14.5.6 SCAN PROBE LITHOGRAPHY This is a group of techniques that uses a probe or a small tool that can measure, test, or obtain information, which is the reason why it was considered a good characterization tool. However, when customized for fabricating miniaturized structures, it is adjusted to have a very sharp nanosized tip. It is capable of scanning the surface of the substrate in an alternating manner (forwards and backwards), while printing images of the desired nanostructures. Although scanning probe lithography can be used, atomic force microscope (AFM) is popularly utilized for this purpose. Applying this technique offers unique precision in positioning atoms individually, as well as fabricating larger devices, such as transistors. While depositing the materials, AFM delivers the correct and exact mechanical forces, heat energy, and electrical characteristics of voltage and currents needed. The working principle is that the AFM tip scratches or carves into the substrate, yielding the pattern wanted. If a nanograft is desired, a subsequent step can follow, whereby other materials are adsorbed to fill the voids formed by the AFM nanotip. Applying electrical bias can also be used to guide the deposition of charged particles to the voids and is preferred in the case of synthesizing metal oxide semiconductors. Many applications relied on this innovation, from printing DNA and proteins to fabricating metal electrodes (O’Connell et al., 2015).

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14.5.7 DIP-PEN NANOLITHOGRAPHY If we scrutinize AFM techniques, we can find out that both scanning probe lithography and dip pen nanolithography are actually subtypes of AFM and, accordingly, perceived as characterization methods. The main difference between the two is the working principle, as it no longer uses the sharp tip to shave a substrate surface. In fact, it treats the tip as a quill that is dipped into an inkwell then used to write or draw respective pattern on any surface. Consequently, it offers superb attributes over the aforementioned techniques, such as: 1. Flexibility: the fact that it isn’t a destructive technique (shoveling and cutting through substrates is no longer necessary) opens up a wide range of choices for substrates of various natures (soft and hard); 2. Tremendous resolution: it allows accurate and precise control of the pen movement, position and spacing; 3. Scalability: the ability of scaling up this method from small laboratory scale with a single pen to large production scale with multiple cantilever arrays that can print up to 11 million patterns simultaneously in an area of few cm2; 4. Relatively economical: providing all these merits it’s still considered cost-effective, when compared to other methods (Eby and Leckenby, 2004; O’Connell et al., 2015). There are two main modes to apply dip-pen nanolithography, which are called meniscus transport and liquid ink deposition. Meniscus transports adopt a kind of ink that was formulated by the self-assembly of molecules in monolayers. The AFM tip is dipped into the ink and allowed to dry. Then, it is put in touch with the surface of the substrate, where it forms a meniscus or a crescent of water as the capillary condensation phenomenon takes place. The ink dissolves when in contact with the water meniscus, shifts to the surface and forms covalent bonds, which allow it to stick. Liquid deposition uses capillary action to transfer and deposit the ink with a solvent carrier. Therefore, the solvent carrier permits a large range of choices of ink to include proteins, metal nanoparticles, and conducting polymers. The modes previously mentioned are the most basic and common of approaches. Nonetheless, some ink types or substrates may require some alterations or modifications. There may be similar approaches, but under different names, such as: nanografting, polymer pen nanolithography, electrostatic transport, and electrochemical dip-pen nanotechnology technique (DPN) (O’Connell et al., 2015).

14.6 NANOBIOMATERIALS With the surge of implantable devices for biomedical applications, plenty of nanobiomaterials have been explored to employ only those with desired and suitable features for nanobionic gadgets. Nanomaterials vary in chemical

14.6 Nanobiomaterials

composition, size, shape, intrinsic properties, and functional groups attached, along with optical and magnetic characteristics, rendering them highly versatile. Therefore, it is necessary to confirm that any material to be implanted possesses biocompatibility, in terms of: (1) safety, that it is nontoxic when interacting with living tissues. It shouldn’t cause any thrombosis in blood vessels or provoke tumor formation; (2) nonimmunogenic, which means it doesn’t trigger immune responses, and has less chance of rejection; (3) chemical inertness, to ensure that it won’t have any side reactions with enzymes or ions available in the surrounding environment (Wang et al., 2015). Biomaterials can be classified into three categories: bioinert, bioresorbable, and bioactive. Bioinert materials are well accepted by biological systems, they don’t cause any interaction or elicit any response. Bioresorbable materials possess only surface activity; they dissolve in the body and are replaced by soft or hard tissues. Bioactive materials make obvious interactions and chemical bonds (Ben-Nissan, 2004). There are various types of nanobiomaterials that can be involved in nanoprostheses, including: polymers, nanotitanium, CNTs, nanodiamonds, nanoceramics and nanocomposites. The classification of these materials is illustrated in Fig. 14.7. Each type of these materials can be generally implanted in human body. Few of which are implanted as prosthetics, and even fewer are involved in nanobionics and nanoprosthetics. The different characteristics each type possesses, qualifies them for a specialized use (Aguilar, 2012).

14.6.1 POLYMERIC MATERIALS Polymers are involved in many industrial fields, and biomedical devices have their share of polymeric materials of various qualities. Despite exploiting about 1020 types of polymers in nanobionic widgets, the diverse characteristics and traits each polymer provides contribute to the wide applications in which they can be found. Here, we will discuss different types of polymeric materials that have served great purposes in numerous classes of nanoprostheses, e.g., polyimide, silicone, parylene, and liquid crystal polymers (LCPs), which are mainly used in neural prostheses. Other polymers are available, but their use is limited for specialized applications.

FIGURE 14.7 Types of implantable biomaterials.

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14.6.1.1 Polyimide A macromolecule of imide units, or \ monomer that has been mainly and widely used in insulating metal electrodes against many environmental factors such as moisture, ions, corrosion, and physical damage by forming a passivation layer. They display remarkable advantages, such as stability in wide range of temperatures and in the presence of oxidizing agents. They also demonstrate good mechanical strength that qualifies them to act as a buffer for mechanical stress applied on the electrode. Their chemical resistance is admirable and they also feature a unique ability to absorb the α particles that can be emitted by ceramics. In addition to that, they offer a possibility to deposit metals after insulation of the electrode with polyimides. They can be followed with further patterning to form a wafer-shaped device that can be easily removed or peeled off with aid of tweezers. Many neural prostheses were implemented in the peripheral and central nervous systems using a polyimide, the majority of which revealed biosafety, biostability, biocompatibility, and long-term effective functioning. Minor cases reported mild immune response to polyimide-coated electrodes (Hassler et al., 2010).

14.6.1.2 Silicone (PDMS) Silicone is used commercially in abundance. The most suitable silicone for medical applications is silicone rubber, which is referred to as polydimethylsiloxane (PDMS) in chemistry. Ordinarily, it presents a low molecule weight and viscosity, yet, cross-linking can alter the molecular weight, permitting it to have a rubberlike nature. Other ameliorations are made to render a grade of silicone suitable for introduction to biosystems, such as incorporating stannous octate catalyst and base polymer, upon production (Lee et al., 2006). Having exquisite endurance to biodegradation and ageing along, with impressive biocompatibility has resulted in prolonged stability in vivo. It can serve as a highly permeable membrane for gases and vapor which are used as ion barriers. They are mainly used for electronic insulation, semiconductor synthesis, and as sealants or adhesives for construction purposes. A good example is Norplant, which is a popular silicone-based device used to deliver luteinizing hormone-releasing hormone in patients suffering from tumors in the male reproductive system. It can also be involved in pace makers, cochlear implants, as well as electrodes in CNS and PNS (Hassler et al., 2010).

14.6.1.3 Parylene Parylene is the commonly used term referring to polyparaxylylene, which is one among several thermoplastic polymers that have a semicrystalline, noncrosslinked morphology. It was introduced in the mid 20th century and manufactured in different types with few variances in the properties. For biomedical purposes, parylene C was found more convenient, as it offers a pleasant electrical/


barrier combination, in addition to its biocompatibility and inertness. Therefore, it played a prominent role in neural prostheses. Another subtype, known as parylene HT, is emerging in this field as it possesses a better tolerance to elevated temperatures. The fragility and lack of robustness in the thin sheets of parylenes compels manufacturers to adopt delicate measures when handling them, and it is the only noteworthy drawback pertaining to their employment in neural implants (Hassler et al., 2010).

14.6.1.4 Liquid crystal polymers This is a distinctive class of thermoplastic polymers, portrayed as linked and aligned molecules of rigid and flexible monomers that gives resemblance to the spatial arrangement within crystals. This class is characterized by a remarkable mechanical strength when under high temperature, drastic resistance to chemical degradation, low hygroscopicity and permeability, and tends to form a barrier for gases with good qualities. These traits qualify LCPs to be used in circuit boards and semiconductors. They aren’t considered of high popularity in biomedical applications; yet, some researchers began to experimentally utilize them in flexible electrode arrays for neural interfaces, with optimistic perspectives regarding their performance (Hassler et al., 2010).

14.6.2 NANOTITANIUM (NANOTI) Titanium alloy is a good example of an allotrope, which means it possesses two structures or forms which are: (α), a closely packed hexagon and (β), with a cubic centered architecture. It can exist, accordingly, as (α), (near- α), (α 1 β), metastable (β), or stable (β), and by modifying the alloy composition and manipulating temperature, the desired structure can be established. The main components of Ti alloys, other than titanium, are aluminum and vanadium; each of them can stabilize one of the allotropic structures by influencing the transformation temperature. For example, the presence of aluminum, oxygen, nitrogen and carbon stabilizes (α) forms. Conversely, molybdenum, vanadium, niobium, and tantalum favor the isomorphous (β) form, and iron, tungsten, chromium, and cobalt yield (β) eutectoids. If neutral alloys are to be obtained, zirconium is the element of choice (Liu et al, 2005). Titanium alloys display outstanding features of lightness, flexibility, admirable tensile strength and resistant to both corrosion and biological fluid interactions. Their corrosion resistance can be attributed to the fact that titanium tends to form a thin layer of titanium oxide upon exposure to air or any oxygen-containing medium. However, this layer deactivates any titanium prosthesis in biological environments. Therefore, many means of surface modification were adopted to improve the bioactivity and biocompatibility, especially with bone tissues where it is are often employed. Initially, scientists explored sand blasting and acid etching. These mechanical and chemical methods enhanced the surface properties and roughness, but failed to control the topography as well as

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residuals remaining with these kinds of treatment. The next candidate they tried was plasma spray-coating with hydroxyapatite and calcium phosphates, since they are an integral component of bones. They showed good outcomes short-term, but weren’t successful in long-term use, due to the formation of weak coating-tometal adhesions and ability to dissolve when implanted. A concept of altering surface to acquire nanometer features emerged to resemble the nanometer characteristics of natural bones. Anodization or anodic oxidation is the electrochemical method used to attain the most suitable surface properties of titanium. This process involves three major events: alkaline cleaning, followed by acid activation to remove the TiO2 layer and other residues, and electrolyte anodization in a threeelectrode setup of electrochemical cells including a Ti anode, Pt cathode and Ag/AgCl reference electrode. The applied voltage leads to deposition of an oxide layer on Ti anode in specified thickness, which controls the morphology and nanoscale roughness of the surface. Anodized titanium alloys showed irregular porous surface structures, with greater resistance to corrosion, due to increased layer thickness of TiO2 and optimistic outcomes in terms of cytocompatibility (Jackson and Ahmed, 2007). Another alloy of Ti, known as titanium-nickel (TiNi) alloy, exhibited extraordinary elastic behavior, referred to as shape memory effect or SME. We can clearly portray it in a manner, where the metal is disfigured and twisted, but soon it rebounds and reverts to the shape previously retained. This superelasticity against stress and strain, which can be attained by temperature rising, along with corrosion resistance, led to its employment as orthodontic dental archwire, catheter guide wires, intracranial aneurysm clip, filters of vena cava, and contractile artificial muscles. Another implantable titanium form are titania nanotubes (TNT), which have successfully replaced stainless steel wires in bone fixation, due to the superior advantages they display in terms of biocompatibility, osseointegration, and mechanical properties. The most prominent innovation here lies in minimizing the incidence of infections that may accompany any bone fixation. Stainless steel wires, also known as Kirschner wires and K-wires, used to introduce a passage for microorganisms as they are inserted transversely through the skin. TNT have an the capacity of drug loading, in which antibiotics can be loaded in the wires and released at the site of implantation to fight the microorganisms imported inside. Beside antibiotics, proteins and growth factors can also be delivered to improve bone healing. Consequently, they ensure long term survival of the implant and decrease the chances of failure of implantation (Gulati et al., 2011).

14.6.3 CARBON NANOTUBES CNTs are one of the novel innovations that acquired a great position in multiple fields of application. Yet, it is the so-called one-dimensional construction that entitles them for employment in the design and manufacture of nanobionic devices. Prior to the discovery of CNT attributes in this field, silicone-based


devices were the primary trend. However, some limitations pertaining to scalability, electron tunneling, leakage currents, and other variables related to the device structure and activation, motivated research to find a material that can bypass and overcome these limitations and provide better performance outcomes. Accordingly, CNTs have impressed them with their electron transport properties and motivated them to scrutinize the possibility of utilizing them in nanoelectronic devices (Avouris and Chen, 2006; Abuelma’atti, 2013). In essence, CNTs are hollow nanofibers of carbon. They are assembled as one or more layers of graphite that are folded or wrapped as a pipe, known as SWCNT and MWCNT. SWCNT refers to single-walled carbon nanotube, where only one sheet of graphite is rolled, displaying a diameter ranging from 0.4 to 3 nm. Alternatively, MWCNT refers to multiwalled carbon nanotube, that incorporates multiple concentric layers that present an outer diameter up to 100 nm and an interlayer gap of 0.34 nm (Scarselli et al., 2012; Fukuda et al., 2003). The obtainable length can be in micro-, milli-, and even centimeters. According to the geometries that act as repetitive two-dimensional (2D) building blocks perpendicular to the axis of the nanotube, which are controlled by the equation (C 5 na1 1 na2 5 (n,m)), SWCNTs can be classified into three categories, as mentioned in Table 14.1, and shapes illustrated in Fig. 14.8. CNTs have the ability to behave as metallic and semiconductor materials, as per the diameter and chirality. What determines its behavior is the (n, m) rule, in which n & m represent the atoms that roll together to form the cylindrical shape. Table 14.1 Class, Type and Shapes of Nanotubes Class Type Shape

Armchair Shaped Nanotubes

Zigzag Shaped Nanotubes

Chiral Nanotubes

Type (n, n)

Type (n, 0)

Any other shape

FIGURE 14.8 Carbon nanotubes geometries.

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If (n 5 m), which is the case of an armchair structure, the metallic properties dominate and their electrical conductivity outpaces the metals themselves. For zigzag geometry, (nm 5 3p) where p is an integer and 6¼ 0, it renders them metallic. If (nm 5 3p 6 1) an energy leak occurs and it results in semiconducting transduction, but still way better than ordinary semiconductors. Therefore, we can conclude that SWCNTs can be either metallic or semiconducting materials, but MWCNTs always possess metallic conductivity (Avouris and Chen, 2006; Abuelma’atti, 2013; Scarselli et al., 2012). The underlying cause of such behavior can be comprehended when noticing the types of electrons available in nanotubes. All graphite-containing systems consist of a valance system of Sigma (δ) and Pi (π) electrons. (π) Electrons are distinguished by their delocalization and free mobility, which makes them easily polarized and transported. Accordingly, (π) electrons generate sensitive electronic structures. (δ) Electrons are localized and contribute to good mechanical properties, including high tensile strength and elasticity as well as stability, in terms of chemicals and temperature (Rotkin, 2004; Scarselli et al., 2012). The early approaches to fabricating CNTs, by either arc discharge or laser ablation, have impressive significance in formulating good quality SWCNTs and MWCNTs, that resulted in them being adopted for large scale production. The arc discharge method is described as passing a high current through a gaseous nonconductive medium (commonly helium) between carbon anode and cathode to form gas plasma. It helps evaporating carbon, which will be collected separately and deposited on a substrate in a particular pattern, constructing the desired nanodevices. Laser ablation employs a laser beam to scan a composite of graphite and different compositions of metals or metal oxides in a closely planned style. This procedure takes place in temperatures of 1200 C in a tube furnace. Then, the soot formed is transferred by argon gas flow from the hot zone in the furnace to a cooled zone outside, where a copper collector awaits deposition of soot. It is noticed that both techniques result in impure CNTs, either in their main composition or overcoated with byproducts like fullerenes and graphitic polyhedrons. Therefore, subsequent purification is necessary to render the nanodevices safe and effective. However, it is very expensive, in a way that exceeds the costs of the entire manufacturing process. In addition, if purification was performed, confirming purity is difficult, since there are no conventional tests to serve this purpose; many defects in the geometry of the nanotubes arise, rendering them undesirable for human use or requiring further modifications. Therefore, as always, a new approach was developed to overcome these drawbacks, referred to as chemical vapor deposition. Chemical vapor deposition, or CVP, is considered the technique of choice for vast production of CNTs. This approach is based on utilizing a metallic nanoparticle as a catalyst, an inert substrate like quartz, silicon, or alumina, and an exposure to hydrocarbon flow. The metallic catalyst provides a surface to grow CNTs and initially form a hemispherical cap, followed by selfassembled growth. Numerous metals like iron, cobalt, nickel, copper, aluminum,


gold and others can be used to serve this purpose. The difference in type, size, and thermodynamic properties in each of them helps building variable structures (amorphous, single-walled or multiwalled). The process describes how each metal provides a special solubility of carbon when under high temperature, and how they segregate as they loom on the surface. In other words, they guide the carbon within the particles to attach and adhere as graphene, then start to coil around the metallic catalyst, to give the final tubular shape. This method offers many advantages in CNT fabrication, such as the ability to fabricate ultralong, superaligned CNTs with uniform electrical characteristics. Developments have been made to offer continuous production. This can be obtained if the substrate is treated to act as a platform to support the CNTs and a catalyst at the same time. A good example of that is stainless steel; it possesses the appropriate nanoscale roughness, as well as the possibility of extracting iron exclusively to be the sole catalyst in the reaction. Moreover, it is a reproducible and reusable substrate for subsequent production, but care should be taken to critically remove all the synthesized CNTs to preserve their reusability (Scarselli et al., 2012).

14.6.4 NANODIAMONDS This type of carbon-based material was discovered in the 1960s but didn’t come under spotlight until the 1990s breakthroughs. They are miniaturized diamond particles with a diameter in the nanometer scale. The persistent research done in this field led to the revelation of outstanding characteristics of nanodiamonds, including exceptional hardness, biocompatibility, electrical resistivity, and chemical stability. What is noteworthy is their relatively lower toxicity, compared to carbon nanoparticles, which makes them safer for biomedical applications. In addition, nanodiamonds display a significant tendency to self-assemble and a capability to bind distinctive molecules on their surface that contributes to the increased attention in their application. It is scientifically acknowledged that diamonds are metastable in nature, due to the sp3 cluster surface which necessitates stabilization prior to any application. This can be achieved by either surface reconstruction into sp2 or functionalization. Controlling the size, morphology, and surface terminations can somewhat guarantee the stability. Plenty of methods were adopted to synthesize nanodiamonds, but the most commercially applicable, producing 210 nm diamonds, are detonation, laser ablation, and high-energy ball milling of high-pressure high-temperature (HPHT) diamond microcrystals. Detonation relies on explosive molecules, since they can give a suitable combination of carbon source and conversion energy together. Detonation is carried out in a closed chamber. It can take two courses of synthesis: dry, where an inert gas fills the chamber, or wet, where water coolant or ice is used. In the end, soot is formed, consisting of nanodimonds (45 nm) along with some carbon allotropes and impurities. The applied temperature and pressure are adequate to not liquefy the carbon in bulk, but to create a liquefied state at nanoscale. In other words, the

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temperature and pressure causes condensation and crystallization of the liquid carbon by provoking a homogeneous nucleation within a volume of supersaturated carbon vapor. If shock waves were used to initiate the explosion, the synthesized nanodiamonds will have a larger diameter (.10 nm). The impurities within the soot need to be eliminated and a subsequent purification of extracted nanodiamonds is necessary to keep them stable. These impurities can develop as a result of the azides use as igniters, or the steel wall of the chamber, and they have the tendency to be superficially attached or internally trapped. Purification can be conducted by liquid oxidants or any other chemical treatment, but it was found hazardous and expensive. The ecofriendly and economical approach depends on oxidizing the noncarbon impurities with air or ozone-enriched air at high temperatures (Mochalin et al., 2011). Nanodiamonds are one of the recent nanobiomaterials that are used in visual prosthetics. They are used for two purposes: to insulate the conductive channels in the electrode array, and to decrease the space between the artificial stimulation and the target neuron. They were employed in much research as either a capsule covering a compartment or as a complete coating of microdevices. The addition of these synthetic diamonds improved the charge injection limit. Other recent works involved the production of electrodes from nanocrystalline diamonds by codeposition of nitrogen as a dopant, or a substance that produce the desired electrical features. Another modification was then made to alter the charge injection limit in a way to improve the neural stimulation without need for water hydrolysis. It was attained by anodization of the electrode in an electrochemical cell to generate an iridium oxide layer on the surface. Advancements are further explored to enhance and widen their field of application (Ghaffari et al., 2016).

14.6.5 NANOBIOCERAMIC Ceramic, in essence, is comprised of polycrystalline compounds that generally include silicates, metallic oxides, carbides and numerous refractory hydrides, sulfides, and selenides. Some special types of ceramics may have covalent bonds to diamonds and other carbon-based compounds (Lee et al., 2006). They were found very beneficial due to their exquisite. Bioceramics refers to the class of ceramics that is suitable for biomedical purposes, especially in replacing diseased or damaged bones. They are most commonly exploited in dental and orthopedic implants, due to their low thermal and electrical conductivity and the satisfactory color and translucency, as well as their hardness, rigidity, resistance to abrasion, and low density (Wang et al., 2015; Khalil, 2012). They can be classified into crystalline, bioglass, alumina, and zirconia ceramics, as described in Fig. 14.9.

14.6.5.1 Crystalline ceramics These are structurally and chemically analogous to human bone mineral component, which is displayed as hexagonal symmetry. To achieve such morphology in


FIGURE 14.9 Classes of bioceramics.

submicrometric level, several methods of production are employed, such as precipitation of aqueous solutions, liquid mix technique, and aerosol synthetic technique. Being analogous to bone tissues doesn’t mean they are exactly the same. Crystalline ceramics show a deficiency in incorporating carbonate ions to the same extent as bone tissues. Conversely, they are capable of substituting other ions such as Na1, K1, Mg21, Sr21, Cl2, F2, and [HPO4]22 which augment the growth of tissue in a faster manner. Biphasic mixtures are synthesized to generate some of the mineral bone material. It has two phases: stable hydroxyapatite, and resorbable beta tricalcium phosphate (β-TCP), that can form carbonate hydroxyapatite and facilitate the regrowth of bones. Cements are also included in this category; they contain mostly calcium, phosphate, sulfate, or carbonate salts that are biocompatible and excellent for bone repair. Cements acquire a powder phase (of Ca or PO4 or both) with an aqueous medium that crystallize in room or body temperature to form calcium-phosphate crystals that can be directly deposited into bones and hasten repair (Tiwari and Tiwari, 2014).

14.6.5.2 Bioglass ceramics Bioactive glass is a polycrystalline ceramic made by controlled crystallization and is composed of SiO2, Na2O, CaO, and P2O5. It used to be produced through UV irradiation-driven precipitation of metals, which induced nucleation and crystallization of glass into fine grains with thermal and mechanical properties (Lee et al., 2006). Currently, a sol-gel process is adopted to control the necessary composition and architecture. They can exist in binary state of CaO-SiO2, or ternary state of CaO-P2O5-SiO2. Both systems form a Si-OH bond that hinders the formation of calcium-phosphate in amorphous forms and delays the crystallization, but ternary systems takes longer delays in comparison. Generally, this type is used as an

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aid or assistance for regeneration and growth of bones, since they have the power to accelerate mineral apatite action (Tiwari and Tiwari, 2014).

14.6.5.3 Alumina and zirconia ceramics Alumina or aluminum oxide (Al2O3) is mainly available in bauxite and corundum and can be synthetically prepared by calcination of aluminum trihydrate. Synthetic alumina (α) exists as a single crystal or polycrystalline with a rhombohedral shape. Alumina for implantation is required to be constituted of 99.5% pure alumina and 0.1% or less of silicon and alkali oxides. Zirconia, in its pure form, can be acquired from zircon (ZrSiO4). It undergoes phase transition in high temperatures, which is why a dopant oxide is always required to stabilize the cubic phase; Y2O3 is usually used for this purpose (Lee et al., 2006). Both alumina and zirconia are used in dental dentures and ceramic crowns. Alumina is distinguished by its aesthetic suitability, admirable hardness, bioinertness, chemical stability, and abstinence from stimulating any allergic response. The only demerit pertaining to its use is brittleness and possibility of cracking. Zirconia is well-known for its appropriate resistance to abrasion and physiological corrosion, high modulus of elasticity, flexural strength, and hardness. However, it is still considered to be inferior to that of alumina (Wang et al., 2015).

14.6.5.4 Nano versus traditional ceramics Nanoceramics are ceramic structures at the nano scale. They preview outstanding and very distinct properties, compared to traditional ceramics, that can solve problems of low ductility and brittleness. Table 14.2 shows the differences and the changes that occur when handling both traditional and nanoceramics. Atoms in nanoceramics can easily migrate by application of any force of deformation, which can explain the ductility they exhibit. The mechanical strength and hardness of nanocermics is more than traditional ceramics by four- to fivefold. Reinforcement of nanoceramic with CNTs in the form of composites, which will be further discussed in the following section, showed enhanced electrical and mechanical features (Wang et al., 2015). Table 14.2 Comparison Between Traditional and Nanoceramics Plasticity Brittleness Ductility Atom arrangement Mechanical properties

Traditional Ceramic

Nanoceramic

Nonplastic Brittle Poor Fixed in crystals Good

Superplastic Tough Good Loose and indefinite Superior


14.6.6 NANOCOMPOSITE Composites are heterogeneous, engineered materials that are assembled in two distinctive phases, which are physically and chemically different. Each phase preserve its own properties and forms an interface between the two phases. Two phases or more can be involved in composites; what is essential for composite fabrication is the existence of a matrix phase and a dispersed phase. Generally, engineering such material is recommended to establish a particular or a synergetic property that cannot be achieved or attained by any phase on its own. This can be explained by understanding the function of each phase. The matrix phase transfers the stress and distributes it between phases, provides a protection from the environment, and supports the dispersed phase. The dispersed phase enhances the matrix characteristics like tensile strength, fracture toughness, and creep resistance; this is usually referred to as reinforcement. There are a variety of factors that can influence the composite features, including properties, volume fracture, and the homogeneity and orientation of the dispersed phase. Classification of composites can be done on three bases, as summarized in Fig. 14.10. Simple composites are the outcome of dispersing a single homogenous dispersion through the matrix phase. If more than one homogenous dispersion is used, it will be called a complex composite. When one or more dispersions are intentionally kept inhomogeneous and dispersed through the matrix, it joins the graded composite family. Hierarchal composites are considered to be of higher level, since they initially form a simple or complex composite with fine entities, agglomerate it to increase the particle size, then disperse it in another matrix material in a hierarchal manner. Particulate composites are those containing particles or flakes, fibrous ones enclose fibers, and laminar incorporates phases as

FIGURE 14.10 Classification of composites.

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laminates or panels within the matrix. A hybrid composite represents a combination of any of the previous three types. Designing a composite requires taking three determinants into consideration: (1) proper selection of both matrix and dispersed phase, (2) the methods of preparation and processing, and (3) the internal and external structural design of the composite (Dorozhkin, 2015). Composites have been very advantageous and involved in variety of fields, but scientists always look for more and for better. They have questioned whether miniaturizing composites to nanoscale can have any impact on their properties, or if it can yield better outcomes in biomedical applications. Nanocomposites can be properly defined as a material of multiple phases, where at least one of which displays a component in 100 nm or less. The first nanocomposite produced was a polyamide nanocomposite in 1950, but the first one to be applied was a polymer/ layered silicate clay mineral composite by Toyota researchers in 1976. Since then, more work was done across many fields. Ordinarily, they are nanoparticle building blocks of clay, polymer, carbon or a combination thereof. When comparing composites to nanocomposites, significant changes are noticed in terms of properties. Their properties don’t merely rely on the properties of the parent material, they also depend on the morphology and surface characteristics (Okpala, 2013). Many alterations can be obtained when the matrix phase is in nanometer scale. The loading, degree of dispersion, size, shape, and orientation of the nanoscale second phase, and interactions between the matrix and the second phase, are determined as per the dimensions of the matrix. Other changes can be attributed to the increased surface-to-volume ratio of the reinforcing material, which is responsible for the improvements in mechanical strength, toughness, and electrical and thermal conductivity of the matrix material. In special materials, it can enhance the chemical resistance, flame retardancy, and decrease the permeability to gases, water, and hydrocarbon (Okpala, 2014). Nanocomposites shouldn’t be confused with composites containing nanoparticles. The former describes when the system of two phases is nanodimensional in total, whereas the latter carries at least one nanoparticulate phase. They share some of the properties related to the surface. However, the difference should be established since their mechanical and optical properties are unalike (Dorozhkin, 2015). Nanocomposites are categorized into two main types: inorganic/organic nanocomposites and lamellar nanocomposites. The lamellar class can be further subdivided into intercalated and exfoliated. Intercalated subtype is when alternating layers of inorganic material and polymer exist in a restricted composition and a limited number of layers and interlamellar spaces. They demonstrate good charge transport, which makes them favorable for use in electronics. Exfoliated possess a ˚ variable number of polymer chains between layers that are separated by .100 A space and are known to have better mechanical properties (Okpala, 2013). Fabrication and processing of nanocomposites is carried out by numerous techniques. Choice of suitable technique depends on the type of materials to be incorporated, the desired properties, and the purpose of their manufacture. Polymeric


nanocomposites are very popular for production of nanoprostheses, and they are commonly synthesized by one of these methods: In situ polymerization: In situ is the Latin phrase for in position. Accordingly, in situ polymerization means the polymerization that takes place within the reaction mixture and cannot be carried out separately. Here, the nanoparticles of the dispersed phase are dispersed in the matrix material, which can be composed of liquid monomers or of low molecular weight precursor in its solution form. A homogeneous mixture is formed, and an initiator is added. Initiators can be free radicals that break the double bonds. Eventually, the mixture is subjected to heat or light to yield the nanocomposites in their final form. Polymers can be thermoset and thermoplastic in nature. Thermoset polymers are synthesized by this method; they are covalently cross-linked which means they lack the elasticity to reshape like nylon-6 and phenolic polymers. A good example of utilizing this method is the fabrication of CNT- PMMA nanocomposites (Lee et al., 2006). Solvent casting: It is a simple approach that doesn’t require expensive equipment or high temperature for processing. In addition, it offers the merit of drugloading, which can be very beneficial for specific applications. The process starts as the polymer is dissolved in a solvent and cast on a smooth surface such as a Petri dish. Then, it is exposed to air to evaporate in room temperatures or to heat in hot ovens. At the end, films or membranes are formed and detached from the casting surface, as illustrated in Fig. 14.11 (Moura et al., 2016). Layer-by-Layer (LbL) assembly: Layer-by-Layer assembly is known to be a method of surface modification, but it also fabricates layered and ordered nanocomposites in an efficient manner. It is frequently described as simple, reproducible, and flexible. The procedure starts with adsorption of different macromolecules in a sequential fashion to create layered forms. As layers are formed, various intermolecular forces are created, such as electrostatic interactions, van der Waals forces, and hydrogen bonding. The deposition of layers can be made through three modes; dip, spin, and spray-coating. Layer-by-Layer assembly can control the size, geometry, and chemical composition of the layers,

FIGURE 14.11 Solvent casting to synthesize nanocomposites.

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which can make this method quiet versatile and feasible in many applications. It can organize the building blocks into flat templates that can produce multilayered films and free standing membranes. It can also arrange them into threedimensional (3D) templates to form capsules, tubular, and porous structures, and hierarchical reactors (Moura et al., 2016). Sol-Gel method: Sol-gel is a solution that converts slowly and evenly to a gel-like diphasic system. The end product is diphasic, which suits the concept of nanocomposite matrix and dispersed phases. This process takes three main steps, as follows: •

•

•

Phase separation: This step is necessary when the particle density is low; the solution can be left for some time until sedimentation takes place and then the liquid will be poured off. If waiting cannot be afforded, centrifugation can be used. Drying: The remaining liquid is removed by any appropriate approach. The system starts to shrink and densify. The drying rate is highly influenced by the porosity and its distribution. Firing: This refers to thermal treatment, which is crucial for condensation and mechanical property enhancement. It is followed by sintering, to cause further densification and grain growth, without need of application of higher temperatures (Hench and West, 1990).

At the end of this process, a matrix and dispersed system is formed that can be further processed to construct nanodevices. Electrospinning: It is one of the impressive techniques exploited in synthesizing nanostructures, due to the high control of composition, morphology, and porosity it provides. The working principle can be explained as: the polymer mixture is injected from a syringe towards a collecting area in the presence of an electrical field. The applied electrical field provokes internal electrical repulsion that induces and maintains a fibrous texture. There are three modes that can be pursued to form nanocomposites through this technique: (1) wet-dry electrospinning, which utilizes a volatile solvent that evaporates as the fibers are formed, (2) wet-wet electrospinning, where the formed fibers are nonvolatile solvents spun with other solvents, and (3) coaxial electrospinning, which allows simultaneous spinning of two components in a core-sheath fibrous style (Moura et al., 2016). Many examples of nanocomposites are available for biomedical applications, some of which are extensively employed in nanobionics and nanoprosthetics.

14.6.6.1 Cellulose nanocomposite Cellulose ranks first of most abundant polymers in the biosphere. It is composed of β 1-4 glycosidic linked glucose units in a crystalline form. It can be obtained by extraction from natural sources or prepared by means of biotechnology. Nanocellulose has captured attention lately; it was manufactured by different methods to suit particular applications. Mechanical treatment can yield nanofibrous


cellulose, where acid hydrolysis can generate nancrystalline or “nanowhiskers” forms. The other type of cellulose is bacterial cellulose (BC) that is synthesized by bacteria at the nano scale. Nanocellulose proved to be great as a reinforcement in nanocomposites; it increases the strength, flexibility, and biocompatibility. However, some drawbacks are reported pertaining to hygroscopicity, inconsistent thermal stability, and high hydrophilicity that render dispersions in polymer matrix poor. Functionalization becomes a necessity in such conditions to attain better homogeneity in cellulose-polymer nanocomposites. It is accomplished in two ways: (1) a blending process, where nanocellulose plays the part of either a matrix or a nanofiller, or (2) chemical modification of superficial hydroxyl groups via esterification, etherification, or oxidation. BC is considered to be the purest form of nanocellulose and it has been an active topic of recent research. This goes back to the biocompatibility and biodegradability, fine fibrous network, good water-holding capacity, as well as a remarkable strength-to-weight ratio. Many nanocomposites containing BC are prepared and continuously explored for various medical purposes, such as wound healing, bone tissue engineering. A few were efficient enough to synthesize nanobionics that replace damaged organs, such as synthetic blood vessels, and even the replacement of dura mater, which is the tissue surrounding the brain. A few examples of cellulose-containing nanocomposites are: BC-hydroxyapatite nanocomposites, aliphatic polystersBC nanocomposites, poly(lactic acid)-BC nanocomposite, and many more. Recent work on shape memory polymers pointed out the poor mechanical properties of such polymers, which results in low recovery stress as well as slow recovery speeds that limit their application. Therefore, nanocellulose is incorporated to improve the mechanical strength and widen the range of applications (Olalla et al., 2016; Panaitescu et al., 2016).

14.6.6.2 Chitin Chitin is the second most available polymer in nature. It is a polysaccharide which is biocompatible, bioabsorbable, nontoxic, and nonantigenic. It can be found in many organisms like insects, fungi, and mollusks. It has a strong crystalline structure, with an acetamide group and two hydroxyl groups, which can form strong hydrogen bonds. It is an anisotropic material, which means it has variable magnitudes, depending on the direction of measurement. For example, strength and Young’s modulus are high when measured on the axes, but differ when measured in another direction. In biosystems, it can be found in three semicrystalline structures: α, β, and γ. When it undergoes acid hydrolysis, it becomes highly crystalline and forms chitin whiskers, nanocrystals, or nanofibrils. It also possesses characteristics of self- assembly when present in suspension form. So, nanofibrils start clustering into fiber bundles and then start to self-assemble into tissues. The tissues are cross-linked and calcified, increasing their rigidity. Owing to the resemblance in structure between chitin and cellulose, they are prepared or produced in similar methods. Chitin nanostructures are mainly utilized in

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nanocomposites as reinforcing nanofiller to induce positive impact on the overall features and performance (Joaõ et al., 2016).

14.6.6.3 Chitosan Chitosan is a polysaccharide derived from chitin by alkaline or enzymatic deacetylation, or just deacetylated chitin. They almost have the same properties and methods of production. It is an excellent candidate for nanocomposite fabrication. It is occasionally prepared in two classes: chitosan-inorganic material and chitosan-polyanion complex nanocomposites. There are plenty of materials that are compatible with chitosan to form very effective nanoengineered composites. Layered silicate, metallic and ceramic nanoparticles, CNTs, and graphene-based materials are used to reinforce chitosan (Cui et al., 2014). Clay refers to hydrous silicates or aluminosilicates; their main structure consists of Si, Al or Mg, O, and (OH), accompanied by diverse cations. They are available in three types: kaolins, smectites, and layered silicate acids. The reason why they need reinforcement is hydrophilicity. To synthesize proper nanocomposite, homogeneity between silicates and chitosan should be confirmed. Accordingly, there are three possible dispersions conformations: (1) tactoid structure characterized by nonexpendable interlayer space, which results in polymer intercalation failure. This makes this conformation undesirable; (2) intercalated structure where the layers are well-maintained but the space between them can be manipulated to permit polymer penetration through layers; and (3) exfoliated structures that contain well-separated clay layers, which allows efficient mixing of both phases that yields appropriate dispersion. Aluminosilicates, known as montmorillonite, present admirable characteristics such as appearance of negative charge, that give them a weak acid behavior. They also attract cations and forms electrostatic or hydrogen bonding with chitosan, forming very popularly used nanocomposites. Nanoparticles are also other nanofillers that can be used. Their properties should be highly controlled, to tune the overall qualities and behavior of nanocomposites. This includes: particle size, shape, crystallinity, and chemical composition. There are three basic categories of nanoparticle that are extensively utilized, especially in biosensors, e.g.: metallic nanoparticles of silver, gold or zinc oxide. Also, bioactive glass has been intercalated in various polymeric matrixes, due to its surface activity and ability to attach to bones and other physiological structures. Another type is used is ceramic nanoparticles, and the most famous in this category is hydroxyapatite nanoparticles which are known for its osteoconductivity, osteoinductivity, biodegradability and mechanical strength. CNTs and graphene based material are also involved. They increase the elastic modulus and improve thermal and electrical conductivity. However, the surface chemical inertia stands as an obstacle in the way of forming appropriately blended nanocomposites (Moura et al., 2016).


14.6.7 PEEKPOLYMER PEEK is an acronym for a polymer called polyetheretherketone. It is employed in polymer-engineered spinal cages, usually referred to as a Brantigan cage. This is attributed to its biocompatibility and elastic modulus close to that of cortical bones. The small differences in elastic modulus can be compensated via carbon fiber reinforcements. Determining the required fiber amounts and orientation is essential to render the nanocomposites of desired features. PEEK can be involved in two types of nanocomposites and coatings: HA-PEEK and TI-PEEK combinations. HA-PEEK nanocomposites: Hydroxyapatite (HA) was initially added as an attempt to make this combination resemble natural bone, which is a collagenreinforced HA composite. This exploratory move led to the discovery of an impressive alternative for osseointegration. Many approaches were followed to establish this result. For instance, one of the researchers created a strontiumcontaining HA-PEEK composite that mimicked the elasticity of cortical bones and had better bioactivity, as per in vitro testing. Other than HA, some ceramics were experimented like calcium silicates, bioglass and β-TCP. Ti-PEEK nanocomposites: Titanium was used alone instead of PEEK prior to its discovery. It had its merits and demerits comparatively and so does PEEK. A brilliant idea came into light by combining both Ti and PEEK in one nanocomposite, where they can have a positive synergistic effect in some areas and can cover some of the drawbacks that appeared when used separately. Many researchers noted the impact as improved cell attachment, elevated levels of growth factors vital for osteogenesis and maturation, and augmented bone-implant fusion. They are very promising but still not clinically ready to be adopted within therapeutic options (Rao et al., 2014).

14.6.8 HYDROGEL Hydrogels are one of the most vast and versatile biomaterials, they can be manipulated to possess specified qualities physically, chemically, electrically, and biologically. They can be formulated using a wide range of precursors and materials that have even wider spectrum of features. Therefore, we can find hydrogel composites containing carbon-based materials, polymeric, inorganic, and metallic nanoparticles. Hydrogel nanocomposites are utilized in biosensors and biomedical devices. They offer many superior qualities physically, chemically, and biologically, based on altered nanoparticle-polymer chain interactions. Newer generations are being developed to synthesize multicomponent networks that can incorporate many functional groups that help tailor nanocomposites to fit the purpose (Gaharwar et al., 2013).

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14.7 APPLICATIONS Since the manufacture of the very first prosthesis, scientists and engineers have been trying to produce artificial replacements with better functionality, durability, and appearance to suit the demands and needs of users all over the world. Engineering a prosthetic device that is identical to the organ it is trying to replace is the primary goal of all efforts made in this field. In order to produce artifacts that closely mimic the function of living organs of the body, the mechanisms of these organs must be understood thoroughly and perfectly. In addition to the mechanisms of action of both living and prosthetic organs, the utilization of nanotechnology in different types of prosthetic devices is also discussed herein.

14.7.1 ORTHOPEDIC PROSTHESES 14.7.1.1 Artificial joints Arthroplasty, one of the major advancements of the 20th century in the field of orthopedic surgery, is the process of resurfacing or completely replacing a damaged joint with an artificial one made of low friction metal or other material. Damage of joints can be caused by a number of conditions such as inflammatory arthritis, osteoarthritis, avascular necrosis, rheumatoid arthritis, fractures, developmental dysplasia, tumors, road accidents, or soldier’s injuries; all of which lead to pain, stiffness, and deformity, resulting in deteriorated joint movement and stability. As a solution to these problems, total joint replacement procedures were introduced to improve the quality of life for all patients suffering from joint disease (Pramanik et al., 2005). Arthroplasty is classified, according to the location of the joint, into three major types: total hip replacement, total knee replacement, and total shoulder replacement. In addition, other joints that may undergo arthroplasty include ankle, elbow, wrist, foot, and hand joints. In order to understand how any joint replacement prostheses works, the mechanism of action of the normal joint should be completely understood (Janeway and Janeway, 2007). The hip joint is a ball-and-socket joint that allows the leg to move in many directions. The ball of the joint is the femoral head that fits into the acetabulum, which is the socket of the pelvis. A smooth layer of cartilage is pressed between the ends of the acetabulum and the femoral head to prevent friction and provide cushion for constrained motion. Hip joint disease occurs when the cartilage wears away and the bones rub against each other, which results in inflammation and pain (Pramanik et al., 2005; Understanding Total Joint Replacement Surgery, 2016). The operation of total hip arthroplasty is a major surgical procedure that involves distinct risks, as well as hospital admission, anesthesia, and rehabilitation (Neil, 2010). It involves replacing both the acetabulum cup and femoral head

14.7 Applications

with a smooth, freely moving prosthesis that resembles the shape of the actual joint. It consists of two compartments: at the head of the thighbone, a metal or ceramic ball with a long stem is inserted into the thighbone to anchor the prosthetic joint in place. On the other side, a socket that is made of press-fit titanium or cemented plastic is implanted into the pelvic bone (Pramanik et al., 2005; Neil, 2010). According to the materials used, the bearing surfaces of artificial hip joints are either ceramic on plastic, ceramic on ceramic, metal on plastic, or metal on metal, all of which have been approved by the US FDA. They can be cemented or pressfit (Janeway and Janeway, 2007). Although cemented prostheses have been clinically used for a long time as the most common choice of surgeons, noncemented prostheses exhibit better biological compatibility and enhanced osseointegration, due to their rough surfaces (Understanding Total Joint Replacement Surgery, 2016). The choice of prosthesis depends on the patient’s age, their degree of activity, and the quality of the bone (Foran, 2016; Neil, 2010). The knee joint, the largest joint in the body, is made up of three parts: the lower end of the femur, the upper end of the tibia, and the patella. The surfaces of the knee joint are covered with articular cartilage where the three bones come in contact, which allows them to move smoothly. Synovial fluid covers the rest of the knee and produces a lubricating fluid the keeps the friction of the bones to a minimum. Pain, stiffness and muscle weakness result when the cartilage or the bone is damaged, which leads to deformity and the development of a limp. This happens in a number of diseases like osteoarthritis, osteonecrosis, and rheumatoid arthritis, all of which necessitate the performance of total knee replacement surgery. Total knee arthroplasty is contraindicated in tuberculosis and purulent arthritis. During a total knee arthroplasty, the orthopedic surgeon removes the rundown cartilage and bone, then inserts the artificial knee prostheses in position to recover the function and the alignment of the knee. The artificial knee joint consists of four components: a femoral component, made of polished metal that wraps around the lower end of the thighbone; a tibial component, made of either metal or plastic, and used to cover the surface of the upper end of the tibia; a patellar component, made of plastic and used for resurfacing the underside of the patella; and a spacer component, made of plastic and placed on top of the tibial component to form a plastic liner to aid in the smooth movement of the artificial joint (Tateishi, 2001; Understanding Total Joint Replacement Surgery, 2016). There are three major types of knee prostheses: the constrained, the semicostrained, and the unconstrained type. However, only the constrained type is used for total knee arthroplasty. Also termed hinged prostheses, the early constrained prostheses consisted of a tibial and femoral component, connected together with a hinged mechanism and hence exhibited limited freedom of movement. These did not allow for any rotational movement of the joint, therefore extra pressure was applied on the prosthesis when the knee was in motion, which often led to

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destruction of the bone in that area and the wear of the implant. As a result, this type of prosthesis is not commonly used and, in its place, there is a new design of hinged prosthesis called the rotatory prosthesis, reported to exhibit less complication (Tateishi, 2001). The rotatory platform design displays a reduction in the contact stress, shear force, and fatigue, due to a large contact area between separate articulations and separated rotational movements (Haider and Garvin, 2008; Ranawat et al., 2004). One method to fix the knee prosthesis in place is to use methyl methacrylate bone cement, resulting in what are termed cemented prostheses, which are needed for patients with fragile bones, to ensure the stability of the joint and prevent its loosening. The other method, mainly performed in young patients with strong bone, is to use screws, metal beads, and/or hydroxyapatite granules for an uncemented prosthesis. Another method is to mix both these techniques to perform a hybrid fixation of the prosthesis in which one of the femoral or tibial components, usually the tibial, is fixed in a cemented fashion while the other one, usually the femoral, is left uncemented (Tateishi, 2001). The third main type of arthroplasty is total shoulder replacement. The shoulder, much like the hip joint, is a ball-and-socket joint. The upper end of the humerus form has the shape of a ball that allows it to glide against the glenoid that forms a socket in the scapula. The surfaces of theses bones are covered with smooth cartilage that ensures the effortless movement of the joint (Elbaz). A total shoulder arthroplasty involves the replacement of the damaged head of the humerus with a metal ball, which is attached to a stem inserted into the humerus to secure the ball in place. On the other side of the joint, the damaged glenoid must be replaced with a smooth polyethylene socket, which is fixed in place using bone cement. Another technique, which can be completely cementless, involves attaching the metal ball to the socket and the plastic socket to the upper head of the humerus, this is termed a reverse shoulder arthroplasty (Craig, 2013). Hip, knee, and shoulder are the main, and most common, joints that may undergo arthroplasty procedures. However, other joints such as ankle, elbow, wrist, and finger joints may undergo total replacement, when required. Arthritis and degeneration of cartilage in ankle joints causes severe pain as well as limited range of motion. Therefore, artificial ankle joints become an excellent solution that returns the ankle to its normal movement and reduces pain without placing pressure on the other joints of the limb. The early designs of ankle prostheses contained a tibial component and a talar component; however, these joints were constrained and did not allow a wide range of motion. Unconstrained, mobile bearing, three-component artificial ankles were developed later on and had an additional polyethylene meniscus component, along with the metal tibial and talar components (Wang and Brown, 2016). Much like the knee joint, the elbow is a hinge joint. A prosthetic elbow has two components, a humerus component and an ulnar component. In addition,

14.7 Applications

some designs of a prosthetic elbow have a pivot or hinge, while other designs rely on ligamentous and muscular power to hold the joint together. The linked prostheses have an advantage of having a better range of motion and better joint stability, while the unlinked type has a lower risk of wear and less body invasion (Sanchez-Sotelo, 2011). The wrist is a very complicated and flexible joint, constituting carpal bones, radius, ulna, and articular cartilage. A prosthetic wrist helps in rebuilding balance and providing motion (Ma and Xu, 2016). The most common design of wrist implant has two components: the first part is inserted into the radius and has a curved end where it faces the wrist joint. The carpal component is inserted into the hand bone and has a flat surface facing the first component. A polyethylene spacer is fitted between the two metal components. The spacer is flat on the side of the carpal component and round on the side of the radial component to ensure a perfect fit and to enable a natural wrist motion (Carlson and Simmons, 1998). Hands have a number of joints functioning together to achieve a desired motion. Among these joints, the metacarpophalangeal joint (MCPJ), the proximal interphalangeal joint (PIPJ), and the first carpometacarpal joint (CMCJ) undergo total replacement procedures if damaged. The trapeziometacarpal joint, in particular, has a unique anatomic structure that grants it a special range of motion over three different planes, which makes a total arthroplasty procedure quite difficult (Watts and Trail, 2011) Early finger implants were of the constrained hinge type, however, they faced a lot of complications and problems including loosening and fractures. Conversely, the unconstrained prostheses exhibited more favorable outcomes. Silicone and pyrocarbon, among other materials, are used in the manufacture of finger implants (Badia and Sambandam, 2006). In any total arthroplasty procedure, there are a number of complications and problems that need to be overcome in order for the patient to have a completely functional prosthetic joint. These complications include infection, nerve injury, venous thrombosis, as well as loosening, dislocation, and rejection of the prosthesis. Scientists have experimented with many potential solutions, such as the use of new materials in the manufacture of a prosthetic implant for better biocompatibility, or new structural designs for better movement. The utilization of nanotechnology in total joint replacement prostheses is a major approach to improve the outcome of an arthroplasty procedure for many reasons. Nanotechnology provides a solution for plenty of the problems that occur with the use of an artificial joint. Conventional artificial joints are made of titanium, having microsurface features. This results in the body recognizing the artifact as foreign material and immediately initiating a rejection response. Even the smallest rejection response can lead to painful loosening or weakening of the implant. Since osteoblasts function at the nanoscopic scale, it would be easier for the implant to merge with the bone at the prosthesis-bone interface, if it can interact at the same basic level as the osteoblasts. With this idea in mind, scientists utilized nanotechnology in manufacturing artificial joints. The application of nanomaterials in a prosthetic

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implant can greatly reduce the chances of implant loosening by improving the osseointegration (Sullivan et al., 2014). For the purpose of ensuring implant success, a number of issues must be addressed, including the fixation of the implant at the interface between prosthesis and bone. Modification of implant surface properties is essential for improved osteoblast response and osseointegration. This is achieved by designing the surface in a fashion that mimics the natural environment of biological tissue and also incorporating adhesive factors into the interface, thus boosting osteoblast-implant integration. Since all these interactions occur at the nanoscale, it is befitting to use nanotechniques to control the surface topography of the prosthesis, by forming a nanotextured surface pattern that allows easy and efficient osseointegration (Torrecillas et al., 2009). The increased activity of osteoblasts leads to increased bone growth on the nanotextured surface of the prostheses, especially the nanoroughened hydroxyapatite-coated ones (McHale et al., 2013). Therefore, nanotubes became of great benefit in the field of prosthetics. These nanotubes are usually made of metals, metal oxides, metallic alloys, polymers, and ceramics (Ganguly et al., 2014). Titanium prostheses coated with nanotubes have shown better osteoblast attachment. This is because nanotubes have the same chemistry as DNA, which makes it easy for protein components to attach to the surface of these nanotubes, thereby reducing the rejection response or even completely preventing it. Even though the conventional artificial joints served well, they had a short lifetime (of under 20 years) as well as high wear rates. Self-assembling CNTs were introduced into the implant design, resulting in high quality implants that last longer and are better accepted by the body (Chun et al., 2004). The basic concept of self-assembling nanotubes is that they have the same chemistry as DNA which allows them to assemble themselves in the form of tiny rosette-shaped rings, made of base pairs of guanine and cytosine on the surface of the prostheses, which promotes the adhesion of osteoblasts to the titanium joint. Many rings then join together to create rod-like nanotubes having the width of only 3.5 nm. It was found that not only bone cells but also cells from other parts of the body adhere better to foreign bodies exhibiting surface bumps as wide as 100 nm. These bumps mimic the surface features of natural tissue, which promotes cell adhesion as well as cell growth, leading to the development of longlasting natural implants. In addition to their potential in future material design and drug delivery systems, nanotubes can be tailord to suit specific body parts by adding special amino acid sequences or growth factors that signal the surrounding cells to attach to the implant. This is of crucial importance in transplantation of artificial body parts like blood vessels or even the brain (Chun et al., 2004). Furthermore, incorporating nanotubes into the implant is effective and also very economical, as researchers have found that even low concentrations of nanotubes can provide the same results as high concentrations. In other words, you can achieve high osteoblast attachment using a very little amount of nanotubes.

14.7 Applications

Another useful characteristic of these rosette nanotubes is that they automatically arrange themselves into a web on the surface of the implant, which bears a perfect resemblance to the pattern of natural collagen fibers in bones (Chun et al., 2004). After implanting a prosthetic joint, the body reacts to the synthetic material with an inflammatory rejection response that limits the biological integration and activity of the prosthesis, leading to dislocation, loosening, and failure of the implant. Protective nanotextured coating of the prosthesis protects it from this reaction and ensures better mechanical stability and improved long-term results (Torrecillas et al., 2009). Some materials used for coating prosthetic joints are nanoceramics, polymer nanocomposites and nanocrystalline diamonds (McHale et al., 2013). A number of useful properties can be added to the prostheses by coating their surfaces with nanostructured materials such as nanoengineered hydroxyapatite, titanium, and cobalt-chromium-molybdenum, all of which have shown better adhesion of osteoblasts to the implant when compared to their microscopic or macroscopic counterparts. Hydroxyapatite, in particular, has been widely used in its nanophase for orthopedic purposes. A paste made of nanocrystalline hydroxyapatite showed incredible results in healing bone defects and fractures. The paste is used as a filler in between bone cracks and as a substitute to bone grafts for treatment of bone defects (Sullivan et al., 2014). Additionally, nanotextured hydroxyapatite is used in conjugation with type 1 collagen to produce nanocomposite scaffolds, which are arranged in layers to form prosthetic implants that are intended for fixing osteochondral defects. These implants are made of specific and accurate ratios of collagen and nanohydroxyapatite. For instance, the layer of the implant that is inserted into the cartilage region is made of 100% type 1 collagen, whereas a 70% nanohydroxyapatite and 30% type 1 collagen ratio is for the layer to be inserted into the bone region. This results in less morbid and more compatible prostheses (Sullivan et al., 2014). To avoid malignant growth of osteoblasts around the implant, nanoengineered selenium is applied to titanium prostheses to inhibit tumor growth at the interface between implant and tissue. Nanoengineered silver, however, is effective in preventing wound infections and stimulating healing. Therefore, when applied to titanium prostheses, nanoscopic silver prevents acute infections after a total joint replacement surgery, due to its bactericidal and antiadhesive properties (Sullivan et al., 2014). A modern modification to total joint prostheses is the incorporation of nanophase drug delivery systems, by coating the artificial joint with a biodegradable polypeptide nanostructured film of a particular drug, usually antibacterial. For example, when a cefazolin nanofilm was applied to total joint replacement prosthesis, a reduction of bacterial load was observed, as well as an improvement in osteoblast integration. Moreover, the nanofilm resulted in enhanced osteoblast adherence and proliferation, even without the incorporation of cefazolin into the film (Sullivan et al., 2014).

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Nanosized crystallites in diamond coatings are a new trend in the manufacture of artificial joints. Nanocrystalline diamond has great potential in medical implants, as it combines surface smoothness with high corrosion resistance, resulting in reduced amounts of wear debris and prolonged lifetime of prostheses. The coating exhibited the typical morphological patterns of bone cell growth, indicating a lack of cytotoxicity. In fact, most of the wear debris of this type of coating is made up of completely harmless and inert diamond particles that elicit no or little allergic reactions. The very low roughness of the nanocrystalline diamond surface contributes to the biotribological and biomechanical behavior at the sliding contact surface of the prosthesis at the bone-prosthesis interface. In addition to antioxidant and anticancer properties, nanocrystalline diamond also has protective properties that allow it to form a selective protective barrier between the implant and the biological environment. Moreover, it exhibits the highest resistance to bacteria in comparison to steel or titanium (Amaral et al., 2008). The function of nanocrystalline diamond can be improved further by implementing a number of modifications. One example is linking human immunoglobulin G antibody to the surface of nanocrystalline diamonds, consequently providing it with the capability of biomolecular recognition. Another example is functionalizing nanocrystalline diamond coating with bone morphogenetic protein-2, making it more biomimetic and improving its osseointegration (Amaral et al., 2008).

14.7.1.2 Bone fixators Artificial joints are not the only orthopedic prostheses utilized in our current time; bone fixation devices have also become commonly used tools in medical orthopedic procedures. Bone fixators are the medical tools used for osteofixation of bones in the case of orthopedic trauma therapies (Giannoudis et al., 2007). Bone is one of the biological tissues that retains the capability to regenerate throughout a person’s life. The process of bone regeneration is a complicated and well-orchestrated formation process that takes place in normal fracture repair. Although bone tissue has a large capacity for self-healing, intervention is required in many clinical conditions. Such conditions include cases that require a large quantity of bone regeneration, like large-scale skeleton reconstruction in cases of trauma, tumors, and skeleton abnormalities, or in other conditions where the regeneration is hindered by a number of factors like osteoporosis, atrophic nonunions, and avascular necrosis (Dimitriou et al., 2011). Various techniques were developed in an effort to accelerate or compensate for insufficient regeneration and to produce substitutes that are identical to natural bone. These intervention methods include the use of autologous bone grafts, which is considered the so-called “gold standard” for fracture repair; use of osteoconductive scaffolds; and allograft implantation. In addition to the methods that promote the regenerative process, there are other techniques that mainly enhance the mechanical stability and mechanical stimulation of the fractured bone, thereby

14.7 Applications

ensuring optimal bone healing, including the use of bone fixation devices (Dimitriou et al., 2011). Bone fixation devices are systems of internal or external stabilization designed to enhance the mechanical stability and the bone repair process with the help of surgical intervention. Mechanical properties of bones vary according to the interaction of bone with the applied stress, which is described by Wolff’s law. Scientists are attempting to decipher the phenomena of bone fracture repair by applying Wolff’s law, along with variations in other parameters such as implant rigidity, fracture stability, fracture gap size, and interfragmentary strain (Giannoudis et al., 2007). Treatment of fractures mainly aims to restore the anatomical alignment of broken bone fragments in order to allow the bones to heal in the correct position, and to relieve pain. For inherently stable fractures, a minimal amount of effort, using casts or braces, is needed to minimize the interfragmentary movement, as these fractures selfheal by intramembranous and endochondral ossification. When it comes to complex fractures, with severe soft tissue damage or with infections, fixation systems and implants are utilized. The purpose of any fixation system is to provide the stability by reducing external loading and muscle activity to an extent that brings the interfragmentary movement to a minimum. Interfragmentary compression is achieved through devices like compression plates and tension bands, which bring the fracture fragments together via an external application of compression force. Bridging plates are fixation devices that bridge the fracture fragments with a number of screws anchored to the intact parts of the fractured bone, thus decreasing the stress load and increasing vascular supply. Accordingly, bone fixators can be classified into two broad groups (Fig. 14.12; Krishnakanth, 2012).

FIGURE 14.12 Classifications of bone fixation devices.

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External bone fixators are indicated in cases of open fractures having substantial soft tissue injuries that need vascular procedures, fasciotomy, polytrauma, fractures in children to avoid fixation of pins or screws through the growth plate, temporary joint bridging before later open reduction internal fixation (ORIF), and arthrodesis of the ankle, elbow, or knee. Other conditions include corrective osteotomies and limb-lengthening procedures (Taljanovic et al., 2003). Healing of fractures under external fixation is dependent on factors that include the rigidity of the fixator, the configuration of the fracture, the accuracy of fracture reduction, and the amount of physiologic stresses (Aro and Chao, 1993). External fixators have a number of advantages over ORIF and intramedullary nailing, such as the adjustable construction, the simple application, and the increased accessibility for wound care (Moss and Tejwani, 2007). External fixation devices are composed of pins and wires such as Schanz screws, Steinman pins, and Kirschner wires. They are positioned percutaneously on both sides of a fracture. Various clamps connect the pins and wires to external fixation rods made of stainless steel or carbon fibers. Krischner wires, having the widest indications and easiest method of application, have been the most commonly used fixators due to their simplicity, versatility, and economy (Egger, 1992). External fixators are classified into three basic types: standard pin fixator, ring fixator, and hybrid fixator. As the name indicates, a standard pin fixator consists of pins, percutaneously placed into the bone, and externally connected to a rod. Pin fixators are also termed as standard uniplanar fixators. They are indicated in cases of long bone fractures, except those of the proximal femur or humerus, or in cases of complex distal radius fractures. The second type of external fixators is called a ring fixator. Ring fixators, frequently termed Ilizarov fixators, consist of thin wires that are attached to external rings or frames. They were first applied in limb-lengthening procedures by Russian surgeon Ilizarov but are currently used in other applications. Hybrid fixators combine both standard pin fixators and ring fixators together by attaching Krischner wires to a proximal ring that is connected to a unilateral external rod, which is, in turn, connected to the distal bone shaft via Schanz screws. This set up is most commonly used for the mending of proximal and distal tibial fractures that are close to the joint. Some external fixation devices are pinless fixators, in which the clamps are anchored onto the cortex without penetrating the medullary canal. These fixators are commonly used for tibial fractures. Although they are not highly stable, they allow later safe intramedullary nailing (Taljanovic et al., 2003). One of the earliest methods to treat fractures without completely immobilizing the limb with a cast or skeletal traction is ORIF. Internal fixation allows prompt function of the injured bone while it heals. Internal fixators, usually made of stainless steel or titanium, are classified into wires, plates, pins, and screws, and intramedullary nails or rods. Staples and clamps are also used along with plates and wires. Wires, used alone or in conjugation with other fixators, are commonly utilized to reattach osteotomized bone fragments. Used together with pins and

14.7 Applications

screws, wires can create a compression band on the fractured bone. Wires of various diameters are braided together to suture bone and soft tissue (Taljanovic et al., 2003). Another internal fixation device is the compression plate or neutralization plate, which is a stainless steel or titanium plate with screw holes on its length. Compression plates apply pressure at the fracture ends; however, in cases of severe fractures or bone loss where compression is not possible, the plate is applied as neutralizing plating to keep the bone fragments in place as the fracture heals. Therefore, plates are usually used for spinal and long bone fractures. Bone fixation plates are divided into a number of types: dynamic compression plate (DCP), lowcontact DCP, tubular plates, blade plates, reconstruction plates, bridge plates, as well as newer types like point-contact fixator, and less invasive stabilization system plate. Some plates are designed to fit a specific anatomical position such as condylar plate, angled blade plate, condylar buttress plate, T-plate, cobra head, obliqueangled T-plate, and spider plate. Other than compression plating and neutralization plating, percutaneous plating is a new technique for internal fixation and is considered an evolution of plating techniques (Taljanovic et al., 2003). Pins and screws are manufactured in a large variety of sizes and are frequently used in orthopedic practice. They serve a number of purposes, including temporary fixation of fracture fragments, accurate placement of larger screws, and attachment of skeletal traction devices. Krischner wires and Steinman pins are among the most regularly used pins. Screws, usually used with plates, nails or rods, are divided into two classes, according to the Association for the Study of Internal Fixation: cortical screws, which are usually threaded and used in the diaphysis, and cancellous screws, which are designed to cross long portions of cancellous bone. Cortical screws often have smaller thread diameter and less pitch than cancellous screws. According to function, screws are divided into: interfragmentary screws, crossing the fracture line and providing compression between fragments; cannulated screws, having a hollow shank and used for fixation of subcapital hip fractures; syndesmotic screws, used to stabilize distal tibiofibular syndesmosis; and dynamic compression screws, used in the treatment of intertrochanteric proximal femur fractures. One more type is the Herbert screw, which is used for fixation of scaphoid fractures. Anchor screws are used for capsular, tendinous, and ligamentous repairs. Another type is the Kurosaka screw that is an interface screw. Interface screws can be metallic, or bioabsorbable and radiolucent, and are used as fixation devices to anchor bone grafts (Taljanovic et al., 2003). Intramedullary nails and rods are mainly used for the treatment of femoral and tibial diaphyseal fractures, as well as humeral shaft fractures. They have the advantages of allowing early weight bearing, providing optimal biomechanical positioning, resisting torsion and bending, as well as minimizing soft tissue exposure. Various designs of intramedullary nails and rods are available. For example, femoral nails are anteriorly bowed to fit the anatomical shape and position of

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the bone, and reconstruction nails have proximal locking holes designed to accommodate screw insertion into the femoral head and neck. In addition, flexible intramedullary rods, like the Ender nail, Lottes nail, and Rush pin, have small diameters and great flexibility to accommodate long bone anatomy (Taljanovic et al., 2003). Bone fixation devices can be incorporated with an antibacterial property through the utilization of nanoadditives. Metal-based nanoparticles are added to the outer surface of the fixators, inhibiting the growth of a number of pathogens that are present on human skin, thereby decreasing the incidence of infections. This antibacterial protection against bacteria, mold, and mildew, along with regular cleaning of fixation devices prevents cross contamination and prolongs the product life (Frydry´sˇek et al., 2011). Another method to prevent infections of osteofixators is surface modification of titanium devices. Certain nanosized topographies of titanium and its alloys have been proven to greatly reduce the bacterial adhesion to the fixator, while simultaneously enhancing bone repair. Nanoroughened titanium surfaces displayed reduced adhesion of the bacteria that are commonly associated with orthopedic implants infection: Staphylococcus aureus, Staphylococcus epidermidis, and Pseudomonas aeruginosa (Yang et al., 2015). Another useful application of nanoadditives is the incorporation of bionanocomposites on CNT. Bionanocomposites are nanocomposites of natural organic or inorganic polymer matrices, having remarkable properties of biocompatibility, biodegradability, and stability. CNTs exhibit unique characteristics that make them ideal for the use in fixation devices. In addition to their ability to align themselves naturally into ropes, they display great strength, special electrical properties and high thermal conductivity (Frydry´sˇek et al., 2011). While using conventional metallic fixators, the surgeon must repeat X-ray diagnostics a number of times from different angles to visualize bone fractures properly during a bone fixation operation. Consequently, the need for X-ray invisible fixators has arisen, in order to make the operation easier and shorter, and to reduce the radiation exposure of both the patient and the surgeon (Frydry´sˇek et al., 2012). Incorporation of polyurethanes into the nanotubes have proven to significantly enhance the mechanical properties and reduce X-ray absorption. A great advantage of using CNTs in the manufacture of bone fixation devices is their high X-ray invisibility. This is important as it helps in the clear visualizing of the bone fractures (Frydry´sˇek et al., 2011). Nanotechnology has also aided in improving the fixation efficacy of different internal fixators. Different techniques involve the adjustment of thread design and shape of screws, as well as surface modifications. Surface modifications include nanostructured topographies and nanomaterial coating. Nanostructured topography of titanium enhances the osseointegration of internal fixation devices. Coating internal bone fixators with nanosized HA coating enhances their fixation efficacy, biodegradability, stability, and osseointegration (Yang et al., 2015).

14.7 Applications

Another method to improve osseointegration of titanium implants is anodization. Studies have proven that vitronectin and fibronectin, proteins known to promote cell adhesion, are highly adsorbed on the surface of anodized nanotubular titanium, thus promoting the adhesion and proliferation of osteoblasts. In addition, anodized titanium screws and pins have displayed better skin growth and resistance to infections, when compared to conventional titanium implants. Moreover, it was reported that bone cells deposited high levels of calcium on anodized titanium, resulting in enhanced osteocalcin synthesis, which is important for bone synthesis (Yang et al., 2015). TNTs, in conjugation with titanium wires are currently being used to design new bone fixation tools with drug delivery properties. Electrochemical anodization is used to produce TNT arrays on the surface of titanium implants for the purpose of carrying growth factors, proteins or drugs, including antibiotics and lipophilic drugs. Drug release from antibiotic-loaded TNTs is controlled by a number of factors such as the structure of the nanotubes, their surface topography, and polymeric coating. Poly(lactic-co-glycolytic acid) and chitosan coatings provide extended release of lipophilic drugs, as well as drug-loading nanocarriers like polymeric micelles. Incorporating the antibiotic into the osteofixation device provides the advantage of directly releasing the drug from the implant into the infected area around it, which enhances the antibacterial action and reduces the incidence of infections. In addition to their ability to resist microbial infections, TNTs have the merit of excellent biomechanical compatibility with natural bone, as they have matching elastic modulus. This also results in an even distribution of skeletal load over both the bone and the implant, thereby shielding the bone from stress-induced bone degradation (Gulati et al., 2011).

14.7.1.3 Bone grafts Since the start of this century, orthopedic surgeons have been trying to utilize bone grafts for treatment of fractures as an alternative to internal and external bone fixators. The difference lies in that bone grafts are used to fill in the bone defect, whereas bone fixators just stabilize it mechanically in order to allow it to properly selfheal. There are a number of ways to classify bone grafts: autografts and allografts, or cancellous and cortical bone grafts. The gold standard of bone grafts is autografts. An autograft or autologous bone graft is a specific portion of osseous matter that is extracted from an anatomical site then transplanted into another site in the same person. It has the advantage of complete histocompatibility, along with osteogenic healing properties. Donor site pain, excessive blood loss, infections, and long operation time are the major limitations of autografts. In addition, there is a very limited supply of grafts in the case of pediatric patients, which creates a big obstacle in the application of autologous bone-grafting procedures (Roberts and Rosenbaum, 2012). Another limitation is the inability to customize the graft tissue into the required form (Taljanovic et al., 2003). Allopathic bone grafts, i.e., allografts, are osseous

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matter that is extracted from a different genotype from the same species, sterilized, and then transplanted to the recipient. Allografts have the merits of short surgical time and lack of donor site morbidity. However, they have the limitations of high costs and risk of viral transmission (Roberts and Rosenbaum, 2012). Other disadvantages include loss of biologic and mechanical properties secondary to its processing, as well as immunogenicity (Taljanovic et al., 2003). Bone grafts have various forms, such as osteochondral, cancellous, and cortical grafts, as well as demineralized bone matrix. A cancellous bone graft is the most common form of autograft. It possesses a high concentration of osteoblasts and osteocytes, providing it with great osteogenic property. It also has a large trabecular surface area, which encourages the revascularization at the recipient site and leading to ease of incorporation of the bone graft into the natural bone. Contrary to cancellous bone grafts, cortical bone grafts have great structural integrity, which provides excellent mechanical support to the bone. Conversely, cortical bone has a comparatively small supply of osteocytes and osteoblasts; therefore, its osteogenic properties are very limited. Moreover, most osteocytes in cortical bone grafts perish after transplantation, further deteriorating the osteogenic function and hindering the incorporation process (Roberts and Rosenbaum, 2012). The various complications of allografts and autografts have encouraged scientists to search for alternative bone graft substitutes; accordingly, synthetic bone grafts came into application. Bone graft substitutes are indicated for vertebroplasty, osteomyelitis, fracture augmentation, fracture nonunion, and augmentation of defects occurring due to benign bone lesions. Many materials are commercially available for the fabrication of bone graft substitutes. For instance, ceramics and ceramic composites, collagen and mineral composites, demineralized allograft bone matrix, coralline hydroxyapatite, calcium sulfate and calcium phosphate cement, in addition to bioactive glass (Taljanovic et al., 2003). Bone substitutes have numerous forms such as powder, putty, pellets, and coatings on implants (Harvey et al., 2010). Bone graft substitutes are also engineered in the form of synthetic scaffolds that provide optimal conditions for cell adhesion, proliferation, and growth. They can incorporate diverse polymers, like hydroxyapatite, calcium phosphate, bioactive glasses, and glass-ceramics in various combinations and designs. New methods of manufacture have succeeded in the fabrication of nanostructured bone scaffolds that display excellent biomechanical and biocompatible properties, along with unique characteristics that allow for bone ingrowth, cell migration, vascularization, and extensive fluid transport. CNTs, helical rosette nanotubes, silver nanoparticles, as well as other nanotechnologies have been incorporated into nanocomposite scaffolds in pursuit of optimal bone graft substitution (Harvey et al., 2010). The promising aspects of bionanocomposites include the resemblance between their structure and natural bone geometry, as well as the incorporation of new

14.7 Applications

artificial materials that combine bone minerals with biocompatible materials. In addition, the nanoscale dimensions grant the biocomposites a high surface reactivity, strong interfacial bonds, large surface area, flexibility of design, and good mechanical properties. Integrating bionanocomposites with tissue engineering techniques involving cells and scaffolds, provides the potential of creating authentic bone grafts (Zhao et al., 2010). For a scaffold to be considered successful, it should be osteoconductive with a porous 3D structure that supports formation of new bone, diffusion of nutrients, and neovascularization (Patrascu et al., 2015). Nanocomposite scaffolds can incorporate single-walled CNTs to enhance the mechanical properties (Harvey et al., 2010). CNTs display exceptional toughness and flexural strength, hence their use as reinforcement material. They are also excellent for proliferation of cells (Ravichandran and Rajendiran, 2015). Helical rosette nanotubes (HRN) also have the potential to be used for manufacturing scaffolds and coating implants. They display a self-assembly that can modulate direct bone growth and prevent osteomyelitis. Helical rosette nanotubes are made of DNA base pairs that assemble themselves into stable nanotubes of 3.5 nm diameter in biological solutions. They are linked by hydrogen bonds, hydrophobic interactions, and base-stacking interactions. Helical rosette nanotubes have an helical geometry that mimics the structure of collagen in bone tissue. Incorporating different peptides, such as arginine-glycine-aspartic acid (RGD) and lysine, into the HRNs improves the function of osteoblasts. For instance, HRN coating on titanium implants enhanced the adhesion of osteoblast into the implant surface (Webster, 2008). In addition, the characteristics of the graft-bone contact surface have a crucial role in determining the success of graft-bone integration. The surface topography of the bone graft should be of similar roughness to the fracture surface in order to promote bone growth. Nanotextured surfaces have been proven to improve adhesion, induce metabolism, and release osteoinductive factors like growth factors and matrix proteins (Harvey et al., 2010). Another surface modification technique that improves graft-bone integration is the functionalizing of graft surface through the covalent immobilization of bioactive agents. A special coating, called biocatalytic latex, is 50% by volume microorganisms that form a film on graft surface. The latex emulsion preserves the cell by partial desiccation and forms pores as it dries. As a result, the living cell is entrapped and encircled by nanopores. These latex coatings have shown a great potential in osseous stimulation and antimicrobial action (Harvey et al., 2010). Other nanomaterial used to improve bone graft scaffolds includes silver nanoparticles as antimicrobial agents. Conventional implants exhibit a high risk of infections, which eventually leads to implant failure, even mortality. Therefore, it would be of great benefit if the scaffold biomaterial had an antimicrobial action. Chitosan, a polymer often used in bone grafts, has wound-healing and antimicrobial properties, as well as high metal-binding efficacy, especially to zinc, copper, and silver. Nanosilver has great antibacterial action against harmful pathogenic

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microorganisms. However, metallic copper has a better antimicrobial action than silver. Copper kills bacteria rapidly by severely damaging the bacterial cell membrane. Silver nanoparticles have a wide spectrum action against both grampositive and gram-negative bacteria. The actual mechanism of action is still unclear but the accumulation of silver in the tissues is known to cause disruption of bacterial cells. Silver can kill bacteria through a number of mechanisms; one is through the binding of silver ions to bacterial DNA preventing their replication, or to the sulfhydryl groups of their enzymes disabling cell respiration and the transport if vital nutrients. This usually produces reactive oxygen species that damage the cell structure (Saravanan et al., 2011). Other than nanosilver, nanoparticles incorporated in bone graft scaffolds include biologics and genes. Polymeric spheres are used to entrap growth factors that are released after the degradation of the polymer. In addition to polymers, hydroxyapatite, carbonate apatite, collagen, chitosan, alginate, and hyaluronic acid are used for encapsulation of biologics. The advantage of nanoparticles over microparticles is their ability to penetrate the tissues and cells in order to achieve accurate targeted delivery. This might aid the bone healing process (Harvey et al., 2010).

14.7.1.4 Spinal implants The purpose of a spinal prosthesis is to stabilize and treat an ailing spine. The success of the implant is greatly dependent on the material and, to a lesser extent, the mechanics. Osteoconductive properties are necessary to ensure proper fusion of an implant, which is crucial to achieve the needed clinical outcome of a spinal fusion (Martz et al., 1997). Titanium is used to manufacture most spinal implants like PEEK cages, orthopedic screws, and plates. PEEK devices have also been used during the last two decades; however, much like titanium, PEEK had low bioactivity. PEEK is a polymer of high mechanical strength, excellent imaging compatibility, good biological compatibility, chemical inertness, low toxicity, and stiffness closely matching natural bone. To overcome the problem of inadequate osseointegration and achieve high rates of bioactivity and bone infusion, osteoinductive and osteoconductive agents can be incorporated (Yang et al., 2015). Surface modification of spinal prostheses improves implant stability and osseointegration by interacting with the body to generate a bioactive layer at the implant-bone interface (Yang et al., 2015). For instance, surface modification of titanium implants increases both the on-growth and in-growth of bone. Initially, cells attach to the implant, adhere, and then spread through the material. Cell attachment can be improved by manipulating the surface roughness. Nanoscale surface modification increases the porosity and produces surface patterns that resemble natural bone, thus increasing the cell adhesion and differentiation. Increasing the porosity highly enhances the in-growth of cells (Rao et al., 2014). As mentioned earlier, the incorporation of bioactive agents like bioceramics can improve the chances of implant success. Hydroxyapatite is the most commonly

14.7 Applications

used bioceramic in orthopedic implants as it is very similar to natural bone. However, the mechanical properties of hydroxyapatite are not ideal for spinal prosthesis. This problem can be overcome by coating titanium implants with hydroxyapatite, thereby achieving high strength due to the titanium body, and good osseointegration, due to the hydroxyapatite layer on the surface. To achieve ideal results, hydroxyapatite nanoparticles are used in these coatings (Rao et al., 2014). Hydroxyapatite nanoparticles are also incorporated into PEEK implants to improve their biomechanical behavior, as well as their bioadhesion. In addition, nanophase titanium oxide can be added to PEEK implants to achieve enhanced osteoblast attachment and spreading (Yang et al., 2015).

14.7.1.5 Tendon and ligament prostheses Tendons and ligaments play a crucial role in joint stability and movement; therefore, any damage to them can affect the function of the joint and lead to degenerative disease. Tendon and ligament tissue is a dense, fibrous connective tissue connecting muscles to bones and bones to bones. There are two approaches for the management of tendon and ligament injury, be it chronic or acute. The first is a conservative approach. It provides pain relief through rest, injection of antiinflammatory agents, and orthotics. The surgical approach is based on surgically repairing damaged tissue or excision of inflamed areas. Tissue grafts are currently being used to restore the function of damaged tendons and ligaments; however, they may lead to mechanical mismatch, insufficient tissue integration, and laxity, along with the risk of rejection in the case of allografts (Rodrigues et al., 2012). The physiological environment of tendons and ligaments make it hard for a prosthetic device to fit in their position and mimic their action. Nevertheless, a number of prosthetic devices and tissue substitutes have been intended for use as replacement of damaged tendons and ligaments. For example, polytetrafluoroethylene (Gore-TexW), terephthalic polyethylene polyester (LarsW ligament), and polyester ethylene terephthalate (Leeds-KeioW). These products have displayed good results over short-term use, but the results of using them for a long period are still unclear and ambiguous, exhibiting a variety of complications. Although ligament prostheses displayed mechanical characteristics similar to those of natural human tissue, they were not adequate, as they exhibited complications of wear and degeneration. The mechanical properties of tendon and ligament prostheses are highly dependent on the mechanism of surgical fixation. Surgical fixation devices like sutures, staples, screws, and washers can decrease the stiffness and modulus of prosthetic devices. To date, no prosthesis has shown satisfactory results; therefore, autografts remain the first choice for primary tendon and ligament reconstruction (Rodrigues et al., 2012). Currently, the applications of nanotechnology in this field are limited to tissue engineering of biomimetic scaffolds (Chen et al., 2009). Some studies propose the use of polylactide/glycolidenanofiberous scaffolds to engineer tendon and

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ligament prostheses, with an improved collagen deposition and increased tenogenesis (Yang et al., 2013).

14.7.1.6 Dental prostheses In the dental field, prosthetic devices are used to replace missing teeth. The main requirements in dental prostheses are osseointegration and integrity of epithelial junction. Osseointegration is influenced by the surface characteristics of the implant, along with its chemical composition, and wettability. Nanotechnology has opened the chance to modify the surface features to control biological interactions and tissue integration (Y Pai et al., 2016). Nanoscale surface modification and calcium phosphate coating are applied to dental implants to improve their biocompatibility and stability. Like orthopedic prostheses, nanohydroxyappatite coatings were used to provide titanium implants with osteoconductive surfaces. Calcium phosphate coating was used to precipitate biological apatite nanocrystals, which promoted the incorporation of proteins and the bioadhesion of osteoprogenitor cells. This led to the production of the extracellular matrix of bones (Y Pai et al., 2016). Nanoroughness of the implant surface enhances the bone-to-implant contact and improves the healing of bone around the implant, leading to better osseointegration. In addition, titanium oxide nanotubes improved the production of alkaline phosphatase by osteoblastic cells on the surface of titanium implants. This led to better osteogenic differentiation and bone tissue integration (Y Pai et al., 2016). Prosthodontics, the branch of dentistry concerned with implantable prostheses, utilizes various types of nanomaterials to improve the properties of implant material, such as ceramics and cements, and enhance their durability and efficiency. For example, silver and platinum nanoparticles are added to polymethyl methacrylate as antimicrobial agents. Moreover, the addition of metal nanoparticles, such as titanium oxide and ferric oxide to PMMA materials increases the hydrophobicity and decreases bimolecular adherence (Y Pai et al., 2016). Nanoceramics have been used to revolutionize dental crowns. Conventional dental crowns are often made of alumina ceramic and zirconia ceramic. Traditionally, ceramics were made of clay and natural materials, whereas modern ceramics are made of silicon carbide, alumina, and zirconia. Nanoceramics made the dental crown tougher and more ductile than conventional materials. They also have excellent mechanical properties (Y Pai et al., 2016).

14.7.2 NEUROPROSTHESES Neuroprostheses are based on the principle of brain and machine communication, termed brainmachine interface (BMI). It involves processes like signal recording, interpretation, sensory feedback, and adaptation (Greevenbroek, 2011). Currently, only a limited number of neuroprosthetic devices are being commercially used, such as cochlear implants (Chhatbar, 2009).

14.7 Applications

Learning the mechanisms of the brain is crucial for the engineering of intelligent neuroprosthetic devices. However, even after all the research done on the functioning of brain, it still holds so many mysteries that scientists could not decipher to date. Another cause of difficulty in understanding this magnificent creation is the great variations between individuals, which is to be expected when observing how personalities and convictions differ from one person to another. These variations include structural differences in the sizes and connections of the different regions of the brain. Conversely, scientists have discovered that single neurons exchange information via electrical discharges that travel between neurons and accumulate, creating an action potential that leads to the firing of more signals (Greevenbroek, 2011). A BMI functions as a technological replacement of a biological signal modality. For instance, if there is a failure in communication between the brain and an extremity, a BMI neuroprostheses can connect them by reading the signals from the motor cortex, using them to control a robotic limb, then providing feedback to the sensory cortex in a closed-loop control system. Theoretically, this method allows thoughts to be converted into actions seamlessly and subconsciously, which indicates that any activity of the nervous system can be implemented by neuroprostheses. With this idea in mind, scientists have managed to apply this concept to cochlear implants, for the treatment of deafness; and retinal implants, for the treatment of blindness (Greevenbroek, 2011).

14.7.2.1 Cochlear prostheses Cochlear stimulation was first discovered by Allesandro Volta, who also invented the electric battery. The first cochlear implant that succeeded in electrical stimulation of the auditory nerve was created by Djourno and Eyrie´s. A modern cochlear implant is composed of a microphone, a speech processor, and a battery pack fitted in the ear shell. The sound is recorded, processed, and then communicated to an external transmitter, which, in turn, sends the signals into a receiver implanted in the skull. The receiver generates a current between the intracochlear electrodes and the reference electrode (Greevenbroek, 2011). Nanofibers and nanowires have been successfully used to improve the stimulation of auditory nerves in a cochlear implant. Creating electrodes that are made of nanofibers and nanowires provides the implant with an elasticity that allows it to take a curving shape. Moreover, nanofibrous electrodes exhibit an increased number of nanoelectrode brush-ends, thereby increasing the number of stimulating sites. Consequently, the electrode brush-ends are placed closer to their target nerve endings; this solves any problems that may arise from an electrode-nerve gap. In addition, CNTs and polymer composites are used to coat the electrode at the nerve endingelectrode interface to improve the conductivity (Aurora et al., 2010).

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14.7.2.2 Retinal prostheses A fully functional retina consists of photoreceptors that detect the light via photosensitive molecules at their outer segments. When subjected to light, the photosensitive molecules trigger a series of neurochemical events that stimulate ganglion cells to send signals to the visual centers of the brain. An inactive photoreceptor can cause the retina to lose its sensitivity to light; however, the remaining neurons can still be electrically stimulated (Weiland and Humayun, 2014). A basic retinal prosthesis is composed of a number of functional compartments, including an imager that converts light to electrical impulse, an electronic unit that processes the image and generates stimulating signals, and an array of microelectrodes that excite the retina. Wireless data and power transmission are also included in some retinal implants. Retinal prostheses are divided into three classes, according to the location of the functional electrode-nerve interface (Weiland and Humayun, 2014). Epiretinal prostheses, placed on the upper surface of the retina, include Argus I, GmbH (IMI) and the Epi-Ret Consortium (Epi-Ret). All these have wireless data and power transmission to avoid transcutaneous wires (Weiland and Humayun, 2014). Since the prosthesis is implanted on top of the ganglion cells, direct electrical stimulation is possible. An external camera is mounted on glasses to transfer image information wirelessly to an intraocular electrode (Lin et al., 2015). Subretinal prostheses are placed under the retina in the vacancy of degenerated photoreceptors. They can be either passive or active implants. An active system has an external power supply and has a subretinal chip that combines the photoreceptor and electrode array together. It is mainly driven by the surrounding light of the actual image. The incident light is communicated as an electrical impulse by an amplifier, which creates an electronic picture transmitted to an electronic cell that perceives the image in shades of gray (Lin et al., 2015). In a passive system, a subretinal prosthesis is activated by the incident light activating the photodiode arrays on a silicon disk, resulting in an electrical impulse that excites the retina (Weiland and Humayun, 2014). Passive implants are not currently used, as they do not provide a meaningful perception. Generally, a subretinal implant processes information in the inner retina; this allows a natural feel to the perception experience. In addition, they allow natural eye movement, unlike the systems that need a camera, where head movement is required to look at objects. So far, Alpha-IMS is the only subretinal prosthesis that has managed to reach human clinical trials (Lin et al., 2015). The third type of retinal implant is suprachoroidal prosthesis. These are placed between the sclera and choroid, by making an incision in the sclera, and placing an electrode in a position that keeps the exposed electrode contacts directed towards the retina. Another electrode is situated in the vitreous cavity. However, this type of prosthesis requires the aid of a camera to guide the patients (Lin et al., 2015).

14.7 Applications

Many nanostructures are used for the improvement of retinal prostheses. Generally, they showed good cell-implant interaction, improved structural integrity, and high biocompatibility and stability. Decreasing the size of prosthesis by using nanoscale materials will improve the biological compatibility of the implant; reducing tissue injury and inflammatory reaction. Nanoscale resolution can be guaranteed by the application of nanostructures including nanoflakes, nanocrystals, nanoparticles, and nanowires. Good communication between electrode and neuron is crucial for high-resolution perception. Neural electrodes are usually made of gold, platinum, and titanium compounds. Nanoporous electrodes, made by electrochemical methods to form a porous platinum film, displayed an increased specific surface area and lower impedance. Moreover, nanopores increase the mechanical stability for long-term implants. Gold nanostructured electrodes have also shown better neural signaling and electrical performance. In addition, microelectrodes can be modified to design a flake nanopattern on the surface of gold electrodes in order to increase the effective surface area, decrease electrode-electrolyte interface impedance, and improve clarity of vision (Ghaffari et al., 2016). Nanoparticles, both semiconductors and metals, showed good optical sensitivity, a broad-spectrum absorption, and narrow excitation spectra. Optically excited nanoparticles can influence voltage-gated ion channels in neurons; this leads to either the initiation or the suppression of an action potential. Cadmium sulfide quantum dots produced an electrical field of sufficient strength to excite ion channels (Ghaffari et al., 2016). Nanowires have proven to be useful for neural stimulation purposes. Using nanowire microelectrodes in retinal implants enables precise control over the nature of stimulation as they provide high aspect ratio as well as high surface area. They also integrate well with biological tissue, as their dimensions are similar to those of natural nanostructures (Ghaffari et al., 2016). Coating metal electrodes with nanomaterials such as CNTs, polymers, and hydrogels, greatly improves biocompatibility and reduces electrode impedance. Polymers exhibiting loose structure provide high mechanical modulus match, thereby decreasing tissue injury and inflammation. Examples of these polymers are: polypyrroles, poly(ethylenedioxythiophene), and parylene. These are often made of nanoparticles, nanowires, nanotubes, and nanofibers (Ghaffari et al., 2016). Nanostructures having a carbon base, such as CNTs and graphene, have the potential to improve neural electrodes. They display a large surface area, excellent conductance, and the ability to decrease electrode impedance. They also have a flexibility that can aid in engineering a seamless integration circuit. Furthermore, nanocrystalline diamonds, as well as ultra nanocrystalline diamonds, were used as an electrically insulating sheath over the electrode to keep it close to the retina. The space between electrode and neuron should be kept to a minimum, in order to achieve sufficient electrical stimulation of ganglion cells (Ghaffari et al., 2016).

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14.7.2.3 Motor neuroprostheses The purpose of motor neuroprostheses is to achieve cognitive control over an artificial limb. The purpose of a motor neuroprosthetic device is quite different from neuromuscular stimulation device, as the latter stimulates actual muscle while the former links the brain to an external prosthetic device (Turner, et al., 2005). This can only be achieved by creating a connection between the central nervous system and the prostheses, and effectively integrating sensory and motor impulses. This concept was practically realized through brainmachine interfaces (Carmena, 2013). Brainmachine interface neuroprosthetic devices function through a series of steps, starting with real time signal detection via a sensor, then signal transmission to an actuator that will perform the task intended by the first signal. The sensor detects the signal by measuring a certain physiological variable, e.g., the electrical signals and action potential of neurons. It could be achieved through signal detection of individual neurons, known as single-unit recording; or multiple neurons, known as multiunit recording; local field potentials; or several square centimeters of cortex, as in electrocorticography and electroencephalography. Signal detection on both low and high levels is necessary for the construction of a clear, well-detailed control signal. After the sensor detects the signal, it interprets it according to the nature of the signal, its location, and the time of recording, relative to the action time. The next step is to send command signals to nerves, muscles, or even a computer interface of a robotic extremity. A feedback signal is then sent to the brain to fine-tune the action and compensate for any errors that may occur during the execution of the command (Greevenbroek, 2011). To understand the flow of electrical signals, principles of neural decoding and interfacing must be explained. Decoding is a crucial step for a prosthetic device in order to understand and translate a brain signal. A decoder is a device that records neural signals via electrodes and recognizes signal patterns that are then interpreted using multiple mathematical functions. Decoding of neural signals gives rise to cognitive control over paralyzed limbs or prosthetic devices, although gaining control over paralyzed limbs is still far from reality (Vidal et al., 2016; Pedreira et al., 2009). Many upper limb prostheses have been successfully manufactured and utilized, including Luke-Arm, developed by Deka; Modular Prosthetic Limb, developed by John Hopkins Applied Physics Laboratory; and the German Prosthetic Arm, developed by Otto Brock Healthcare. However, lower limb prostheses still do not have adequate evidence of their safety and effectiveness. Examples of lower limb prostheses include Power Knee and Foothill Ranch (Greevenbroek, 2011). To achieve a seamless integration of prostheses and the human nervous system, nanoscale techniques must be applied. For example, since neurons function at the nano level, the application of nanomaterials is essential to enable the

14.7 Applications

designing of a fully functional brainmachine interface. Nanotechnology is incorporated into neuroprosthetic devices through various nanostructures such as nanoparticles, nanowires, CNTs, nanobiomaterial coatings, and nanodevices (Kotov et al., 2009).

14.7.3 CARDIOVASCULAR PROSTHESES 14.7.3.1 Artificial heart valves Prosthetic heart valves are of two types: mechanical and biological. Both types function passively, opening and closing in response to pressure difference and flow changes in the heart chambers. The general appearance and action of prosthetic heart valves is quite similar to natural heart valves. Mechanical heart valves are engineered from rigid nonphysiological material, whereas biological heart valves are made of flexible tissue and synthetic material. Currently, biological heart valves can be made of porcine aortic valves or bovine pericardium (Butany, 2005). Many types of cardiac valvular replacement devices are used currently, including caged-ball devices like Starr-Edwards ball-in-cage prosthesis, caged disc prosthesis (Beall valve), tilting disc valves (Bjork-Shiley valves), and bileaflet tilting disc valves (St Jude Medical valves). Other types include tissue valves (Medtronic or Hancock porcine valves), porcine, or pericardial valves (Carpentier-Edwards valves) and many other devices (Butany, 2005). Mechanical heart valves are made of synthetic material, which tends to be highly thrombogenic. Pure titanium, chromium cobalt, and graphite are used for the manufacture of mechanical heart valves. They are composed of three major compartments. The first is the occluder, which could be a ball or a disc. The second part is the superstructure that keeps the occluder in position. The valve base is the third part and it has a fabric-sewing ring, used to hold the valve in place (Butany, 2005). Biological heart valves, also termed bioprosthetic heart valves, tissue valves, or xenografts, have a structure and shape similar to that of a native aortic valve. The only difference lies in that a biological heart valve is attached to a prosthetic frame. Bioprosthetic heart valves are divided into two classes: heterografts or xenografts, such as porcine aortic valves or bovine pericardial valves. The second class is homografts or allografts, e.g., the aortic or pulmonic valves extracted from human cadavers. Autografts are a new alternative to these options, whereby the patient’s own pulmonary valve is used. The process involves excising the pulmonary valve then grafting it into the aortic root, and then placing a homograft in the pulmonary site (Butany, 2005). When a heart valve becomes too rigid, it becomes difficult for the heart to pump blood sufficiently, thus leading to high blood pressure and possible myocardial infarction. Utilizing nanorods to alter the structure of a valve has proven to

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be useful to solve this problem, as well as many complications of prosthetic heart valves (Duckworth and Kirkham, 2010).

14.7.3.2 Artificial myocardium A new application of nanotechnology is the manufacture of a nanosized artificial myocardium. An artificial myocardium is used to assist the pulsation of the heart and is placed on its external surface. It involve the use of nanosized sensor, control chip and an actuator. A nanofilm sensor that utilizes diamond-like carbon was successfully developed. In addition, another nanosensor that uses optical fibers was designed and passed the animal experiment. A nanosensor controls the hemodynamics of an artificial heart, thereby increasing the life expectancy of patients. Nano actuators are not yet successfully developed but once they are, they can be applied to many organs (Yambe, 2008).

14.7.4 CEREBROSPINAL FLUID DRAINAGE SYSTEMS CSF, having a volume of 150 mL, is encompassed in the brain ventricles (25 mL) and subarachnoid spaces (125 mL). The CSF circulates from its secretion sites to its absorption sites under the influence of arterial pulse waves, respiratory waves, posture, and jugular venous pressure. The drainage of CSF is mainly achieved through the arachnoid villi, cranial, and spinal nerve sheaths, the cribriform plate, as well as cerebral arteries adventitia. Normally, the CSF is renewed about four times in a day, however, the rate of turnover decreases as the person ages, leading to an increase in the levels of catabolites and waste products in the brain. This is usually the basis of a number of neurodegenerative diseases. The CSF serves as the hydromechanical protection of the brain and spinal cord. It also influences the development and regulation of the central nervous system and the activity of neurons, as well as the homeostasis of brain interstitial fluid (Sakka et al., 2011). The circulation of CSF is based on the balance between secretion and absorption. The choroid plexus is mainly responsible for the secretion of the CSF. The two major routes for drainage of the CSF are the arachnoid villi route, taking place in the wall of venous sinuses; and the lymphatic drainage route, mediated by olfactory mucosa and cranial nerve sheaths, including optic, trigeminal, facial, and vestibulocochlear nerves. The epidural venous plexus and spinal nerve sheaths mediate the drainage of the CSF into the lymphatic system in the spinal subarachnoid space (Sakka et al., 2011; Pollay, 2010). Any disturbance in the balance between secretion and drainage results in disturbances in the cerebral physiology and disorders of CSF hydrodynamics, resulting in dementia and hydrocephalus. Hydrocephalus is a condition of high intracranial fluid volume, caused by obstruction of drainage system (Sakka et al., 2011). The treatment of hydrocephalus involves the implantation of prosthetic shunt systems. A simple shunt system, like the ventriculoperitoneal shunt, consists of

14.7 Applications

three compartments: a ventricular catheter, made of silicon rubber and placed into the ventricle; a valve, placed between the skull and scalp and used to regulate the CSF flow and pressure; and a peritoneal catheter, used to transport the CSF into the abdominal cavity where it is reabsorbed. A prosthetic shunt remains inside the body of the patient for a lifetime. This gives rise to a variety of complications like blockade and infection (Spiers et al., 2010). The more modern form of a prosthetic shunt is the Programmable Automatic Shunt System. This shunt system is made of a valve, a number of microelectromechanical system (MEMS) sensors, and a microelectronic signal-processing unit. The structural design consists of an implanted coil antenna that transmits information about pressure and flow to a unit that, in turn, adjusts the CSF drainage valve. Control of the valve can be via the input of a neurosurgeon or the internal feedback of a smart system, which is the case in smart shunts (Ferrari et al., 2007). Nanotechnology is involved in the improvement of all compartments of a shunt system. The catheter can be improved by the incorporation of nanotube bundles. The CNTs will work as a filter for the CSF as it enters the shunt through an adjustable valve, keeping out any bacteria or proteins that might block the catheter. This will prevent the complications of infections and blockade; therefore, the risk of shunt malfunction is reduced and the need for additional surgery to replace a faulty shunt is avoided. The overall shunt system would still be the same; therefore, the surgeons will not need additional training (Spiers et al., 2010). Another possible modification to the shunt system is designing an entire catheter made of nanotubes. The downside of this approach is that each nanotube will need a valve of its own to control the drainage of CSF. This problem can be overcome by the manufacture of novel nanovalve systems. However, nanovalves operate by responding to alteration in the pH of the surrounding medium. Since the local pH of the CSF remains constant, nanovalves are not effective in shunt systems at the current time (Spiers et al., 2010). A Programmable Automatic Shunt System can be modified by utilizing (bioNEMS) nanoelectromechanical system instead of its MEMS sensors. Nanoelectromechanical system sensors have demonstrated single molecule sensitivity in biomedical studies. BioNEMS are capable of interacting with an extremely low portion of analyte and generating a response in a short time. Given their nanoscale size, bioNEMS can perform force measurements locally and with very small samples (Roukes, 2000).

14.7.5 PLASTIC AND RECONSTRUCTIVE PROSTHESES 14.7.5.1 Breast augmentation The treatment of breast cancer usually involves partial or complete removal of the breast. This exerts harmful physical and psychological effects on the patient.

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Therefore, reconstructive breast augmentation has become a common procedure to improve the aesthetic and emotional condition of patients. This is usually done through implantation of a synthetic prosthesis, because it requires less recovery time. Breast implants are of three types: silicone gel-filled, inflatable saline-filled, and dual lumen gel/saline-filled, all of which have been approved by the FDA. The implant is often covered by a polyurethane coating layer to improve its appearance and anchor it in place (Gurgacz et al., 2013). Studies have proved that designing nanotopographical features on the surface of the implant improves the biointegration by suppressing the immune response towards a prosthetic breast implant. Furthermore, surface nanotopography helps in anchoring the implant to its place (Salehahmadi and Hajiliasgari, 2013). Artificial skin is also used to enhance the appearance of the breast prosthesis. Nanofibrous scaffolds are used in the engineering of artificial skin, to provide an aesthetic appearance after breast augmentation (Ver Halen et al., 2014). In addition, a common problem that occurred with conventional breast implants is leakage and rupture, which can be overcome by incorporating nanocomposite polymers that will improve the mechanical stability and biocompatibility of implants (Tan et al., 2016).

14.7.5.2 Craniofacial reconstruction Prosthetic facial rehabilitation is usually required after tumor ablative surgery for craniofacial cancers. In ancient times, wax was used in designing artificial facial prostheses. Currently, craniofacial prosthetic devices are often made from polymethyl methacrylate and urethane-backed medical-grade silicone and backed up with skin adhesives, skin loops, or even glasses and magnets (Lemon et al., 2005; Dostalova et al., 2011). Nanomaterials are used in engineering craniofacial replacement grafts. In addition to nanomaterials, nanorobots are used during craniofacial surgery to perform precise processes at the nanoscale and aid in maxillofacial hard tissue repositioning. Furthermore, quantum dots are used for imaging of tumors (Ver Halen et al., 2014).

14.8 ETHICAL ISSUES Since the beginning of civilizations, laws and regulations were made to govern most aspects of life, but there are always few foggy areas of ethical dilemma, which cannot be covered by these rules. Accordingly, moral and ethical obligations surfaced to guarantee the welfare of people, and where would it be most needed, if not in patient healthcare. With the high hopes and prospects for reinforcing the involvement of nanotechnology in bionics and prosthetics for a wide range of biomedical applications,

14.8 Ethical Issues

comes many ifs. Questions such as, the potential to extend their use beyond patients of neural dysfunctions and amputees; or, are there any enfolded facts about these devices that are unethically withheld from the free willing, fully informed, consenting patients? Or the question of would it be morally unacceptable to deprive the patient his right to choose the course of treatment he finds suitable if clinical judgment of the physician declare them undesirable? (Gilbert, 2013). Additionally, since cost and accessibility influence the chances of receiving these implants and prosthetics, can we ensure justice and equal opportunities? (Brey, 2005; Frumento et al., 2010). In order to explore this further, we need to address the following concerns.

14.8.1 SAFE USE: BENEFITS VERSUS RISKS Since safety is a prerequisite in any healthcare services, considering safe implantation and use of nanobionics and prosthetics becomes ethically essential. In spite of all the in vitro, animal tests, and clinical trials conducted and confirming safe use, the risk accompanied with it remains real and individual. There is a constant possibility of them being unsuitable, causing side-effects, or even complete rejection of implantation (Brey, 2005). With that being so, how do we proceed from here? Should we ignore those revolutionary advancements out of fear, or do we go ahead regardless of risk?

14.8.2 JUSTICE There are plenty of legislations and guidelines constituting the criteria for a candidate to receive an implant or a prosthetic. Most of which are medically based on the condition of the patient, which sounds fair. This criteria neglects important factors that establish a barrier to patient care, such as cost, accessibility, adaptability, and training (Frumento et al., 2010). Is it equitable that only few can get the best quality devices and services just because they are financially able or have direct and easy access to centers that sell, set up, and provides training? Should patients pay more in case of immune system reactions against the implantable device, and more for higher quality ones that don’t generate such reactions? And are manufacturers supposed to produce devices for individualized specifications? (Brey, 2005).

14.8.3 IDENTITY, PRIVACY, AND ACCOUNTABILITY When prosthetics are introduced to people’s life, everybody is concerned with the amelioration they provide and overlook the part where people need to accept that prosthetic as a new piece of them. They must live with the fact that they depend on an artificial technology for the rest of their life. This is most likely to be noticed with neural prosthetics, due to the cognitive function and personality

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changes it can produce (Gilbert, 2013). Some neural implants can improve memory and others may control mood; they might change the identity of the patient. This raises important questions: Is it tolerable to have access to one’s memories and emotions? Is it alright to share data obtained from your body with other third parties? Does such access violate people’s right to privacy? Under these circumstances, if a patient with cognitive prostheses committed a crime or felony, will he/she be accountable before the law? (Brey, 2005; Erden, 2016).

14.8.4 AUTONOMY Self-government, or the freedom to one’s own actions, is how the dictionaries define autonomy. Yet, being autonomous among the norms implies being able to take care of oneself; nanobionics and prosthetics offer great promise in this area. When deaf people use cochlear implants, they restore the sense of hearing, thus becoming capable of taking care of themselves. The same goes with amputees that utilize a prosthetic limb, but with neural implants comes the concern. If thoughts and perception are affected by the employment of such devices, can a person still be considered autonomous? Which is more ethically valuable, the person’s autonomy or human race transformation? (Brey, 2005; Erden, 2016).

14.8.5 VALIDITY OF INFORMED CONSENT Technically, there are pillars of informed consent that must be available prior to any experiment or clinical trial to legally approve utilization of data obtained. They include being a noncoerced volunteer with physical, mental, and emotional competency to evaluate merits and demerits of the given procedure. Yet, are these pillars sufficient to accept the informed consent? Let us observe the following scenarios: (1) Many people went through plastic surgeries that enhanced their appearance using what is known as Botox, consciously disregarding the fact that this material is a bacterial toxin that may result in several side effects and dangers when injected into human body; (2) Large groups of people are also encouraged, through media and advertisements, to favor special techniques over others. Directing the conception of an audience toward an area of interest is something that isn’t new to media, but at the same time it can’t be taken as irrefutable evidence. These kind of behaviors in which people are running after quick positive effects, regardless of the long-term consequences; or being already preconditioned to endorse a procedure under influence of media, makes us question the validity and legality of a given informed consent and the need of expanding the borders of approval in this field (Gilbert, 2013).

14.9 Safety Issues Pertinent to Nanobionics and Prosthetics

14.8.6 PROBLEMS OF AMBITION: TREATMENT VERSUS ENHANCEMENT Are nanobionics a course of treatment or a form of enhancement? Answering this question determines the need for employment and development of this new technique, but it also requires an insightful moral reasoning. We can begin with differentiating between both concepts. Treatment merely describes the actions taken to recover health and prevent illness, or any anticipated complication. Defining enhancement, on the other hand, is a bit more complicated, as it can be viewed in both philosophical, ethics-oriented terms and scientific outcomesoriented terms, which discard ethical values (Gilbert, 2013). Estimating the requirement of upgrades in nanobionic devices might become confusing when we differentiate between treatment and enhancement; this confusion ends by asking the right question. We cannot deny the fact that any kind of treatment always strives to enhance a certain circumstance. The real question lies in: Can we consider all enhancements as a form of treatment? In cases of amputees using prosthetic limbs or deaf people using cochlear implants, it is an enhancement of patient’s quality of life which clearly complies with treatment goals. Whilst in other cases, like using neural implants to alter cognitive and physical functions without being therapeutically indicated, they do not constitute a form of treatment. A good example addressing this issue is the military programs for human augmentation. The idea of creating super invincible soldiers with implants, drugs, genetically designed muscles, and generating cyborgs is brilliant, but it isn’t treatment, which reveals that all treatment portrays enhancement but not necessarily vice versa (Brey, 2005; Gilbert, 2013). Another problematic issue linked to neural implants is the optimistic speculations neuroscientists have in this regard. With the rise of attention on the modifications of brain functionality, neuroscientists become more fascinated with idea of developing cognitive traits rather than understanding the brain in essence. Their area of interest is drawn away from changing the brain from an unwanted state and more towards creating a supernatural ideal brain. Do they have the right to be ambitious in changing the human limited potentials and transforming us into more capable, superior organisms in this way? Do such enhancements comply with rules of nature? And are there any consequences to these advancements? (Gilbert, 2013).

14.9 SAFETY ISSUES PERTINENT TO NANOBIONICS AND PROSTHETICS Generally, medical devices should be designed and manufactured in ways that ensure patient safety and well-maintained clinical condition, when operating

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normally for their intended purposes. However, safety is such a relative term. Every device contains a percentage of risk that ought to be assessed in preliminary stages. Devices on nanoscale are the novel technology in focus, promising great advancement in the treatment of sensitive cases and replacement of function losses. Regardless, the risks that can be associated with these tiny structures cannot be controverted. In general terms, risks can be categorized into four classes of risks (Renn and Roco, 2006): 1. Simple risks: characterized by direct, obvious, causal relationship between cause and effect. In other words, the effect is merely attributed to the suspected cause. 2. Complex risks: in this situation, the cause-effect relevance is present, yet, it is difficult to measure and quantify. The difficulty can refer to a host of reasons, like interactive impacts, long lag time between cause and effect, interpersonal variations, or foreign variables. 3. Uncertain risks: uncertainty of the risk points to the involvement of human knowledge, which is always limited and picky, and leans towards predictions and assumptions. If any cause-effect connection is established, the human knowledge intervention will only weaken the confidence initially demonstrated. 4. Ambiguous risks: risks under this category are vague and doubtful, due to two reasons: either there are numerous rational interpretations based on observation and results; or there are variable standards of evaluation to take effects into account, and those are based on the investigator’s opinion. Therefore, when we look into sources of risks using nanobased implantable medical devices, we can find several safety-compromising factors at levels of development, manufacture, packaging and labeling, advertisement, sale, use, and disposal (WHO, 2003). Here, we are more concerned with the risk corresponding to the actual application of the nanobionic device in the human body, which can be traced to either the device itself, or the device as a nanosized entity. The device is constructed from electrodes, a circuit, and other electrically wired components. Unlike large prosthetics, these nanodevices don’t possess external control. Ventimiglia mentioned in his report that one of the design requirements is to ensure safety during conditions of working and failure of the device. He stated that the movement of actuators must be paused in case of control signal loss; that when a battery’s power is low, an LED light notifies the consumer; and, finally, that there must be software and mechanical limits preventing undesired joint movement (Ventimiglia, 2012). In Weir’s chapter on Design of Artificial Arms And Hands for Prosthetic Applications, he highlighted the vital role of manual or current sensing switches, supported with mechanical stops to protect from any unintended motion or any drive burns following (Weir, 2004). This privilege isn’t granted in nanodevices, since they are implanted inside, with no chance of external manual control.

14.10 Conclusion

Consequentially, they become prone to software and hardware malfunction, technical deficiency in cases of battery power exhaustion, and therapeutic misconfiguration (Gupta). During clinical trials for cochlear implants, it was reported that deep insertion of the electrodes can raise complications of destroying the remaining hair cells. It was salvaged later when combined with drugs as a vigilant measure, but that still leaves significant concerns regarding the implantation of other nanobionics and prosthetics to take place in the future (Liu et al., 2013). Another problem acknowledged is the threat to unsecured privacy of related data. In neural cognitive implants, there are considerable data actions that must remain within control of two parties: the patient and the healthcare professional. Data access, accuracy, update, and device identification and settings must be controlled by the aforementioned parties if not by the monitoring healthcare professional alone. Any leakage of data jeopardizes the safety of the patient (Halperin et al., 2008). As a nano entity, the major peril lies within their capability to penetrate biological membranes and interactions with subcellular moieties. Multiple studies confirmed a link between nanosystems and DNA impairment, with further aggravations in embryos that were frequently fatal (Castanõ et al., 2014). Eventually, the question of whether nanobionics and nanoprosthetics are safely used for their intended purposes is merely relative when employing such delicate technologies for replacement or enhancement of a body function. Weighing the risks and balance becomes essential in these cases. To fully depend on them, such judgments should be made by higher authorities and regulations should be clearly stated once they are officially approved.

14.10 CONCLUSION Through the mid-20th century, nanotechnology built a bridge towards the world of prosthetics to synthesize nanoengineered prosthetics, which revolutionized the way prosthetics can compensate for lost body functions or damaged organs. It is remarkable how marvelous advancements can be made by few manipulations at nanoscale. Nanoprostheses have proven to offer better biocompatibility, durability, strength, integration, wear resistance, and other outstanding qualities. The broad spectrum of nanobiomaterials with their diverse natures, chemical compositions, morphology, thermal, and electrical conductivity can be invested to serve multiple purposes. Various biomedical applications require various types of nanobionics and prostheses. They are readily available in the market as orthopedic, dental, cardiovascular, neural, ophthalmic, and cochlear implants. Continuous and persistent efforts are maintained to keep exploring modern and advanced techniques, and materials that can help reach the ultimate goal of human wellness.

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Further Reading

World Health Organization, 2003. Medical Device Regulations: Global Overview and Guiding Principles. WHO Library Cataloguing-in-Publication Data, Geneva, pp. 38. Yambe, T., 2008. Artificial organs with nanotechnology and development of the new diagnosis tool. Annals NanoBME 2, 167176. Yambe, T., 2009. Development of the various kinds of artificial organs and clinical application of the new diagnosis tool. In: Yamaguchi, T. (Ed.), Nanobiomedical Engineering. Imperial College Press, London, pp. 373385. Yang, G., Rothrauff, B., Tuan, R., 2013. Tendon and ligament regeneration and repair: clinical relevance and developmental paradigm. Birth Defects Res. Part C: Embryo Today: Rev. 99, 203222. Yang, L., Gao, C., Wei, D., Yang, H., Chen, T., 2015. Nanotechnology for treating osteoporotic vertebral fractures. Int. J. Nanomed. 10, 51395157. Zhao, H., Biswas, A., Bernstein, G., Porod, W., 2010. Design, synthesis and characterization of all-bone minerals multicomponent bionanocomposites for bone grafts using bone-tissue engineering. ND Science and Engineering Summer Research Symposium and MIND Annual Workshop, University of Notre Dame.

FURTHER READING Good Samaritan Hospital orthopaedic center of excellence, 2009. Guidelines For Total Joint Replacement Patients. Gupta, S. Implantable Medical Devices Cyber Risks and Mitigation Approaches. http://csrc.nist.gov/news_events/cps-workshop/cps-workshop_abstract-1_gupta.pdf (accessed 15.07.16.).

587


Index Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A Absorption, distribution, metabolism, and excretion (ADME), 2627 Acetylcholinesterase, 304 Acid labile proteins, 392 Acid sphingomyelinase (ASM), 118 Acquired immunodeficiency syndrome (AIDS), 495 Activatable cell penetrating peptide (ACP), 74 Active drug targeting, 56 Additives, 392 for silk scaffolds, 458 Adsorption, 300 Affinity-based biosensor, 292293, 297 Alginates, 171 Aliphatic poly(esters), 380 Alkanesulphoic acid, 168 Allopathic bone grafts, 561562 Alternating current voltammetry, 319320 Alumina and zirconia ceramics, 542 Aluminosilicates, 548 Aluminum oxide, 542 ALZET pump, 485486 Amino acids, 111, 303 Amino groups, 171, 356 Amino-terminated ligands, 168 Amperometric biosensors, 321 Amperometric techniques, 303304 Amphipathic and hydrophobic CPPs, 6970 Amphipathic cell-penetrating peptides, 6970 Angiogenesis, 112 Anodization, 535536, 539540, 561 Antidementia drugs, 304 Antisense oligonucleotide (ASO), 205 Appendage route of penetration, 195 Aptamer heterodimers for cell targeting, 266268 Aptamer oligomerization, DNA nanoconstructions for, 257262 functional activity of aptamer homodimers, 258260 lifetime of oligomeric aptamers in vivo, 261262 oligomeric aptamers in sensors, 260261 Aptamers, 250252, 279, 313 oligomerization of, 261262 Aptasensors, 252, 260261, 261f

Aqueous porous regions, 189191 Arginine-glycine-aspartic acid (RGD), 447448 Arginine-rich CPPs, 68, 81 Arthroplasty, 524525, 550 Artificial heart valves, 571572 Artificial joints, 550556 Artificial myocardium, 572 As-spun SF nanofibrous scaffolds, 451 AstraZeneca, 393394 Atomic force microscope (AFM), 220, 531 Autograft, 443, 561562 Autologous bone graft, 556557, 561562

B Bacillus subtilis, 307308 Baclofen, 495 Bacterial cellulose (BC), 546547 Bare-metal stents (BMS), 493 Beam lithography, 530 Betacelluin, 114115 Beta-lapachone (β-lap), 315 Bioactive glass (BG), 456457, 548 Bioceramics, 540, 541f Biocompatibility, 2324, 338, 379, 496 Biodegradability, 381 Biodegradable implants, 480f, 482484, 491492 Biodegradable polymers, as drug delivery vehicles, 1617 Biodegradable synthetic polymers, 380, 444445 Biodistribution, 388389 Bio fuel-cells, 322 Bioglass ceramics, 541542 Bioinert materials, 533 Biolabeling with imprinted polymeric nanostructures, 363 Biological dressing, 418 Biological heart valves, 571 Biological-synthetic dressing, 418 Biomarkers, 161162 Biomaterial immobilization by adsorption method, 300 defined, 298299 Bionanocomposites, 560 Biopatterning, 452 Bioprosthetic heart valves, 571 Bioreactors, use of, 460461

589

590

Index

Bioreactors, use of (Continued) in fabricating in vivo implants, 457f Biorecognition, 318 Biosensors, 279, 292295, 297, 318. See also Nanobiodevices imprinted nanocomposites for, 352362, 359t Bis-chloroethylnitrosourea (BCNU), 483 5,6-Bis(octyloxy)-4,7-di(thiophen-2-yl)benzo[c] [1,2,5]oxadiazole (BODT), 304 Blood-aqueous barrier (BAB), 35 Bloodbrain barrier (BBB) model, 3132, 35, 174177 strategies to overcome, 32 Bombesin (BBN), 88 Bombyx mori, 446, 449451 Bone fixation devices, 556557, 557f, 560 Bone fixators, 556561 Bone grafts, 441, 561564 Bone morphogenic proteins (BMPs), 443444 Bone tissue engineering (BTE), 417 bioderived and synthetic materials for, 444445 future perspectives, 463 global perspective, 441 ideal scaffold for, 442443 lacunae of current materials and practices for, 443444 recent trends in, 460463 cocultures of multiple cells on silk scaffolds, 461462 gene therapies, 462463 growth factors delivery through silk scaffolds, 462 use of bioreactors, 460461 rising needs for bone grafts, 441 silk-based matrices for. See Silk Botox, 576 Brain, drug delivery to, 3132 Brain capillary endothelial cells (BCECs), 31 Brainmachine interface (BMI), 566567, 570 Brain targeting of payload using mild magnetic field, 167 magnetic nanoparticles, 167178 feasibility of superparamagnetic iron oxide nanoparticles for gene delivery systems, 177178 in vitro characterization of, 176177 in vivo distribution of, 173 passive and active targeting, 174 physical characterization of, 171173 polymers used in magnetic targeting, 167171 targeting brain delivery using, 174175

targeting brain tumors with, 175176 magnetic nanoparticles, diagnostic applications using, 178182 cancer theranostics, 179 challenges and future directions, 181182 hyperthermia, 180181 magnetic resonance imaging and other applications, 178179 magneto acoustic tomography (MAT), 180 Branch retinal vein occlusion (BRVO), 395 Brantigan cage, 549 Breast augmentation, 573574 Brownian rotation, 73 Buserelin, 393t, 394

C c-5-epimer α-L-guluronic acid, 171 Cadmium sulfide quantum dots, 569 Calcium phosphate coating, 566 Cancellous bone graft, 562 Cancer, 3738, 158, 160 chemotherapy treatment, 38 drug delivery to, 3839 effect of nanoparticles in, 128129 implantable drug delivery systems in, 493494 Cancer theranostics, 179 Candida rugosa, 350 Captopril drug, 304305 Carbon-based nanomaterials, 410411 nanocomposite hydrogels from, 411 Carbon nanoparticles, 125 Carbon nanotubes (CNTs), 527, 536539 3-Carboxy-2,2,5,5-tetramethyl-1-pyrrolidinyloxy (PCA), 223 5(6)-Carboxyfluorescein, 85 Carboxylates, 167168 Carboxy methylcellulose (CMC), 427 Cardiovascular disease, implantable drug delivery systems in, 493 Cardiovascular prostheses, 527528, 571572 artificial heart valves, 571572 artificial myocardium, 572 Carpometacarpal joint (CMCJ), 553 Carrier-based targeting, 8 Catalysts, 302303 Catalytic biosensors, 297 Cathodic electroporation, 206207 Cationic amphiphilic drugs (CADs), 118 Caveolae proteins, 110 Caveolin-mediated endocytosis, 121 CD28 protein, 267

Index

Cell-penetrating peptides (CPP), 67, 119, 222223 amphipathic and hydrophobic CPPs, 6970 cationic CPPs, 6869 cell-penetrating bacterial effector proteins as novel tools, 8990 CPP-conjugated nanoparticles, 73, 92 cyclic CPPs, 8384 delivery of drugs and proteins, 7076 delivery of nucleic acids, 7683 plasmid DNA, 7880 RNA interference molecules, 8083 discovery and design of, 8388 future respects, 9192 nonviral gene delivery, 76, 80 recombinant production of, 9091 Cell-penetrating poly(disulfide)s (CPDs), 86 Cell targeting, aptamer heterodimers for, 266268 Cellular targeting, 30 Cellulose nanocomposite, 546547 Central nervous system (CNS) disorders, 31 Central retinal vein occlusion (CRVO), 395 Ceramic nanoparticles, 540, 548 Cerebrospinal fluid drainage systems, 527, 572573 Chemical hydrogel preparation, 406f, 408 Chemical vapor deposition (CVP), 538539 Chemotherapy, 3 ChipRx, 488, 488f Chitin films, 419, 547548 Chitosan, 171, 548, 563564 Chitosan hydrogels, 419 Choline oxidase, 304 Chronic diseases, implantable drug delivery systems in, 492495 cancer, 493494 cardiovascular disease, 493 diabetes, 494 ocular therapy, 494495 pain management, 495 Chymotrypsin, 124125 Citric acid, 170 Clathrin assembly lymphoid myeloid leukemia (CALM) protein, 109 Clathrin-dependent endocytosis (CDE), 109 Clathrin-independent endocytosis (CIE), 109110 Clathrin protein, 109 Clay, 548 Clay- or silica-based additives, 456 Clusters, 194195 Coagulation Factor XII, 496497 “Coat and poke” mechanism, 207208 Cochlear prostheses, 567

Coenzymes, 302303 Collagen, 191 Colloidal carriers, 179 Comamonas acidovorans, 309 Compression molding, 385 Compression plate, 559 Computer-aided fabrication of SF scaffolds, 453454 Conductometric biosensors, 320321 Confocal laser scanning microscopy (CLSM), 220221 three-dimensional reconstruction of a hair follicle assessed by, 222f Contact lenses, hydrogels in, 419420 Contact lens-induced papillary conjunctivitis (CLPC), 427 Contraception field, implantable drug delivery systems in, 492 Controlled drug release, 397 Copolymeric hydrogels, 409 Copper, 125126, 563564 Core-multishell (CMS) nanoparticle, 223 Corneocytes, 189191, 193195 Coronary angioplasty, 527 Cortical screws, 559 Corynebacterium glutamicum, 309 Covalent bonding, 300 Craniofacial reconstruction, 574 CRISPR-Cas systems, genome editing by, 80 Crystalline ceramics, 540541 Cyclic voltammetry (CV), 311, 319 Cysteine, 168 Cysteine-rich CPP (CyLoP-1), 88 Cytomegalovirus (CMV), 481 Cytoplasm-targeted action of nanoparticles, 119

D Da Vinci, Leonardo, 514 Decoding, 570 De-epidermized dermis (DED), 212213 Deformable liposomes, 196197 Dendrimer nanocarriers, 1720, 19f Dental prostheses, 528529, 566 Deoxyribonucleic acid (DNA) sensors, 292293 Dermis, 191 Designing extensive DNA nanoconstructions, 269271 Dexamethasone, 223, 393t Dextran, 170 Diabetes, implantable drug delivery systems in, 494

591

592

Index

Diamond coatings, nanosized crystallites in, 556 Diaphysis, 440441 Differential pulse voltammetry (DPV) experiments, 313, 319320 Diffusion cells, 211212 Dipalmitoylphosphatidylcholine (DPPC), 123 Dip-pen nanolithography, 532 DNA aptamers in diagnostics and therapy, 250252 DNA-based nanobiodevices, 312318 DNA locks, 254t, 262, 267 interactions and topologies of, 253f noncovalent linking, 258 DNA nanoconstruction creation, basic principles of, 252257 double helices, 252254 G-quadruplexes, 255256 i-motifs, 256257 triple helices, 254255 DNA nanoconstructions for aptamer oligomerization, 257262 functional activity of aptamer homodimers, 258260 lifetime of oligomeric aptamers in vivo, 261262 oligomeric aptamers in sensors, 260261 with different aptamers, 262268 aptamer heterodimers for cell targeting, 266268 thrombin aptamer hetero-oligomers, 263266 membrane-associated, 275276 DNA origami, 269 three-dimensional, 272273 two-dimensional, 271 DNA tiles, 260, 269, 271 DNA tiles and DNA origami, applications of, 273281 biosensors, 279 drug delivery systems, 273275 membrane-associated DNA nanoconstructions, 275276 molecular machines, 279281 spatial arrangement of the molecules, 276279 DNA tweezers, 281 Docetaxel (DTX), 74 Donepezil, 304 Dopamine, 168 Double helices, 252254 Doxorubicin, 70, 73, 113114, 193, 204, 221, 313314 Dressing hydrogels, 418

Drug delivery devices, for pulmonary delivery, 3435 Drug delivery systems, 420421 hydrogels in, 420421 Drug-eluting stents (DES), 481 Drug flux, 216 Drug-loaded nanoparticles. See Nanomedicine Drugs skin penetration, nanoparticles for, 187 atomic force microscopy (AFM), 220 characteristics of, 196201 charge, 200201 shape, 201 size and surface area, 199200 type of nanoparticle, 196199 confocal laser scanning microscopy (CLSM), 220221 electron paramagnetic resonance spectroscopy, 223 ex vivo skin penetration experiments, 223224 in vitro skin permeation, 209218 applied dose, 215 diffusion cells, 211212 quantification of the drug on the skin and in the receptor solution, 215 receptor solution, 215216 skin model, 212215 in vivo skin penetration experiments, 224229 microdialysis, 226229 tape stripping technique, 224226 mechanisms and routes of, 192195 physical methods to enhance, 201208 electroporation, 202t, 205207 iontophoresis, 201205, 202t microneedles, 202t, 207208 Raman spectroscopy, 221223 skin structure, 189192 skin appendages, 192 transmission and scanning electron microscopy, 219220 Drug targeting, 230 definition and reasons for, 23, 3f properties influencing, 2030 aggregation and concentration, 2728 carrier properties, 2228 drug properties, 2122 functionalization and surface coating, , 2627 physicochemical properties, 2226 physiological and anatomical properties, 2830 strategies, 420

Index

active drug targeting, 56 combination targeting, 67 common approaches, 48 drug targeting systems, 820 lipid-based nanosystems, 915 passive drug targeting, 45 physical targeting, 78 polymer-based nanosystems, 1520 Dry powder inhalers (DPIs), 35 d-SPE (dispersive solid phase extraction), 351t Dual responsive hydrogels, 408 DUROS leuprolide implant, 485 Dynamic implants, 484486 electromechanical systems, 487489 implantable pump systems, 484 osmotic pumps, 484486 propellant infusion pumps, 486 Dynamic light scattering (DLS) measurements, 172173

E Elastic/ultraflexible liposomes, 196197 Elastin network, 191 Electrochemical impedance spectroscopy (EIS), 311, 357 Electrochemical methods, 292 in biosensing, 318322 Electrochemical microbial biosensors, 307 Electrochemical nanobiodevices, 293f Electrochemical sensors, 357, 359t Electromechanical systems, 476, 487489 Electromigration, 203 Electron beams (e-beam), 530 Electron microscopy, 219 Electron paramagnetic resonance (EPR) spectroscopy, 223 Electroosmosis, 203, 206 Electroporation, 202t, 205207 Electrospinning, 346347, 546 nanofibrous scaffolds using, 451452 Electrospun scaffolds, 451 Ellipticine (EPT), 81 Endocytosis, 107111 biological pathway of, 111 Endoplasmic reticulum and Golgi apparatustargeted action of nanoparticles, 119 Endosome/lysosome-targeted action of nanoparticles, 117118 Enhanced GFP (EGFP), 91 Enhanced permeation and retention (EPR) effect, 46, 174, 178179

Enrofloxacin-imprinted nanoparticles, 353 Entrapment method, 301302 Enzymatic bio fuel-cells, 322 Enzyme-based nanobiodevices, 302306 Enzyme biosensors, 297 Enzyme electrodes, 321322 Enzymes, 302304 Epidermal growth factor (EGF), 179 Epidermal growth factor receptor (EGFR)mediated action of nanoparticles, 114115 Epidermis, 189, 191 Epiphysis, 440441 Epiregulin, 114115 Epiretinal prostheses, 568 Epithelial cells, 191 Escherichia adecarboxylata, 309 Escherichia coli, 309, 425 Ethical issues, 574577 autonomy, 576 benefits versus risks, 575 identity, privacy, and accountability, 575576 informed consent, validity of, 576 justice, 575 treatment versus enhancement, 577 Exocytosis, 107108 Extensive DNA nanoconstruction geometry, examples of, 271273 DNA tiles, 271 three-dimensional DNA origami, 272273 two-dimensional DNA origami, 271 Extracellular membrane (ECM), 414415 Extracellular space (ECS), 388 Ex vivo skin penetration experiments, 223224 Eye anatomy and physiology of, 36 drug delivery to. See Ocular drug delivery systems

F Ferimoxtran-10, 170 Ferrofluids, 170171 Feruglose, 170 Feynman, Richard, 515 Fibroin, 447, 462 Fick’s equation, 193 Field-effect transistor (FET)-based sensors, 320321 First order targeting. See Organ targeting Fluorescein 5-isothiocyanate (FITC), 221 Fluorescence, 352353 5-Fluorouracil, 336338 Folate receptor-mediated action of nanoparticles, 113

593

594

Index

Fo¨rster resonance energy transfer (FRET) sensor preparation, 354355 Franz-type diffusion cells, 211212, 214215 Freeze-drying, 452453 FTIR, 173 Fuel-cells, 322

G Gastrin releasing peptide receptor (GRPR), 88 Gene delivery systems feasibility of superparamagnetic iron oxide nanoparticles for, 177178 Gene silencing, 205 Gene therapy, 76, 462463 Genomic action of nanoparticles, 126127 Gliadel wafer, 483, 493494 GLRKRLRKFRNK peptide sequence, 84 Glutaraldehyde, 301 Glutathione S-transferase (GST), 8990 Glycolic acid, 381 Gold-disk electrode (AuDE), 313314 Gold nanoparticles, 7072, 81, 124125, 200, 311, 356 Goserelin, 393394, 393t G-quadruplexes, 255256 Graphene oxide sheets, 348 Green fluorescent protein (GFP), 74 Growth factors delivery through silk scaffolds, 462 Guanidinylated dendrimers, 123124

H 1 H NMR, 173 Hands, 553 Helical rosette nanotubes (HRN), 563 Heparin-binding EGF-like growth factor, 114115 Herceptin, 114115 Hetero-oligomeric aptamer nanoconstructions, 262263 High-pressure high-temperature (HPHT) diamond microcrystals, 539540 High sensitivity microbial biosensors, 307 Hollow microneedles, 207 Homodimeric constructions, 268 Homo-oligomeric DNA aptamers, 258f Homopolymeric hydrogels, 409 Homotetrameric constructions, 268 Homotrimeric constructions, 268 Hot melt extrusion, 386 Human serum albumin (HSA), 124125 Hyaluronic acid, 419420

Hybrid fixators, 558 Hybrid hydrogels, 407 Hydrocephalus, 527 Hydrogels, 403, 453, 549 applications of, 414424, 422t contact lenses, 419420 drug delivery system, 420421 hygiene products, 421423 tissue dressing, 417419 tissue regeneration, 414417 chemical and physical preparation of, 406f classification, 407414, 407f according to method of preparation, 409 based on response, 407408 based on source, 409412 based on type of cross-linking, 409 intelligent hydrogels, 412414 concepts and definitions, 404407 metabolism and hydrogels, 424425 nanoparticle biosafety, 427428 potential risks of, 426427 regulation of, 425426 technical features, 424 Hydrolases, 297 Hydrolysis enzymes, 304, 321322 Hydrophobic cell-penetrating peptides, 6970 Hydroxyapatite (HA), 440441, 451, 454456, 548, 555, 564565 HA-PEEK nanocomposites, 549 5-Hydroxymethyl tolterodine (5-HMT) hydrogels, 425 Hygiene products, hydrogels in, 421423 Hyperthermia, 180181 use of magnetic nanoparticles as, 181

I Ilizarov fixators. See Ring fixators Immobilized metal ion affinity chromatography (IMAC), 9091 Immune cells, 191 Immunoglobulin (IgG), 115116 Immunosensors, 310311 i-motifs, 256257 Implanon, 480, 480f Implantable devices, types of, 524f Implantable drug delivery systems (IDDSs), 376377, 380, 473, 477t classification of, 479489 dynamic implants, 484486 electromechanical systems, 487489 passive implants, 479484 current challenges, 496500

Index

biocompatibility-related issues, 496498 cost-effectiveness, 499 patient compliance, 498 regulatory aspects, 498499 current therapeutic applications, 492496 chronic diseases, 492495 infectious diseases, 495 neurology and central nervous system health, 496 women’s health, 492 design approaches, 489492 implant material selection, 489 mechanisms of drug release, 490492 design features of, 474t future perspectives, 499500 Implantable medical devices (IMDs), 487489 Implantable pump systems, 484 Imprinted polymeric nanoparticles, 331 biolabeling with imprinted polymeric nanostructures, 363 for biosensors, 352362 formats of, 334335 imprinted drug delivery nanodevices, 335343 principles of, 332334 for sample preparation, 344352 Infectious diseases, 495 Infectious disorders, effect of nanoparticles in, 130131 Infusaid, 486 Injection-molding, 386387 Inner limiting membrane (ILM), 36 Inorganic nanoparticles, 544545 nanocomposite hydrogels from, 411 Insulin, skin penetration of, 206 Insulin-like growth factor I (IGF-I), 462 Integrins, 115 Integrins receptors (InRs), 115 Intelligent hydrogels, 412414 Intercellular route, 193194 Interpenetrating polymeric hydrogel (IPN), 409 Intracellular trafficking, 118 Intravaginal ring (IVR), 492 Intravitreal drug delivery, 37 Intravitreal implants, 37 In vitro bloodbrain barrier model, 176177 In vitroin vivo correlation (IVIVC), 396 In vitro skin permeation, 209218, 217f applied dose, 215 diffusion cells, 211212 experiments, 224229 microdialysis, 226229 tape stripping technique, 224226

quantification of the drug, 215 in receptor solution, 216 on skin, 216218 receptor solution, 215216 skin model, 212215 Ion beams (i-beam), 530 Iontophoresis, 201205, 202t Iron oxide nanoparticles, 178

J Jet printing, 531 Jet-type nebulizer, 35

K Keratinocytes, 191, 212213 Kirschner wires (K-wires), 535536 KLA peptide, 7475, 79 Knitted scaffolds, 452 Krischner wires, 558559 K-wires, 535536

L Lab-on-a-chip (LOC) platforms, 292 Lactic acid, 381 Lamellar nanocomposites, 544545 Langerhans cells, 191 Lifetime, defined, 295296 Ligand-based targeting, 8 Linear chitosan (LCO), 110111 Linear range, 295 Linear sweep voltammetry (LSV), 319 Linear working interval, 294 Lipid-based nanosystems, 915 Lipid drug conjugates (LDCs), 1415 Lipid implants, 483 Lipid nanoparticles, 196 Lipids, interaction of nanoparticle with, 123124 Lipoplexes, 11 Liposomes, 912, 129130, 196, 198, 301302 functionalization of, 1112 zeta potential of, 204 Liquid crystal polymers (LCPs), 533, 535 Lithography, 529530 Local transport regions (LTRs), 206 Low molecular weight protamine (LMWP), 7273 Luminescence-based physical transporter systems, 307 Lung anatomy and physiology of, 33

595

596

Index

Lung (Continued) drug delivery to. See Pulmonary drug delivery Luteinizing hormonereleasing hormone (LH-RH), 377378 Lysosomal membrane permeabilization (LMP), 118 Lysosomes, 117118

M Macropinocytosis, 110111 Macropinocytosis-associated endocytosis (MAE), 110111 Macropinosomes, 110111 Magic bullets, 162 Magnetic carbon nanotubes (MCNTs), 350 Magnetic d-μSPE, 351t Magnetic force, 174 Magnetic nanoparticles, 167178 applications of, 172f diagnostic applications using, 178182 cancer theranostics, 179 challenges and future directions, 181182 hyperthermia, 180181 magnetic resonance imaging and other applications, 178179 magneto acoustic tomography (MAT), 180 different functional groups found on ligands used in synthesis of, 169f feasibility of superparamagnetic iron oxide nanoparticles, 177178 in vitro characterization of, 176177 in vivo distribution of, 173 passive and active targeting, 174 physical characterization of, 171173 polymers used in magnetic targeting, 167171 targeting brain delivery using, 174175 targeting brain tumors with, 175176 widely used ligands/polymers for synthesis of, 169f Magnetic resonance imaging, 178179 Magneto acoustic tomography (MAT), 180 Magnetofection, 177178 Magnetolipososmes, 175 Major histocompatibility complex II (MHCII), 110111 Mass-sensitive sensors, 359t Mechanical heart valves, 571 Medical bionics, 514518 Membrane-associated DNA nanoconstructions, 275276 Mesenchymal stem cells (MSCs), 414416, 443, 455, 459460, 462

Mesoporous bioactive glass (MBG), 457 Mesoporous carbon nanoparticles (MCNs), 347, 350, 357 Metabonomic action of nanoparticle, 127128 Metacarpophalangeal joint (MCPJ), 553 Metal and metal-oxide nanoparticles nanocomposite hydrogels from, 412 Metered dose inhalers (MDIs), 34 Methacrylic acid, 346347 Methacryloylamidohistidine-Pt(II), 355 Methyl methacrylate bone cement, 552 Met receptor, 258, 259f Micro and nano contact printing, 530531 Microbial biosensors, 306307 Microbial nanobiodevices, 306309 Microdialysis, 226229 Micro electromechanical systems (MEMS) technology, 487 Microneedle-based vaccination system, 208 Microneedles, 202t, 207208 Millirod/injectable monolith, 381382 MISPE (molecularly imprinted solid phase extraction), 351t Mitochondria-targeted action of nanoparticles, 119120 Molecular imprinting, 332, 333f Molecularly imprinted nanoparticles (MINs), 344347, 349, 352358 Molecular machines, 279281 Molecular recognition elements (MoRE), 252 Moleculary imprinted polymer (MIP), 339340 Monolithic implant, 381382, 489 Monomeric aptamers, 266267 Mononuclear phagocyte system (MPS) cells, 5 Montmorillonite, 548 morphogenetic protein-2 (BMP-2), 416, 462 Motor neuroprostheses, 570571 Mucoadhesive hydrogels, 420421 Mulberry silkworms, 446448 Multilamellar vesicles, 10 Multipolymer interpenetrating polymeric hydrogel, 409 Multipotent stem cells, 414415 Multiwalled carbon nanotubes (MWCNTs), 311, 417, 536539 Myoelectric prostheses, 528

N Nanobioceramic, 540542 alumina and zirconia ceramics, 542 bioglass ceramics, 541542 crystalline ceramics, 540541

Index

nano versus traditional ceramics, 542 Nanobiodevices, 292318 adsorption, 300 biomaterial immobilization, 298302 calibration, 294 covalent bonding, 300 cross-linking, 301 DNA-based, 312318 entrapment method, 301302 enzyme-based, 302306 high sensitivity, 294295 immunosensors, 310311 lifetime, 295296 microbial, 306309 rapid response time, 296 response time, 296 selectivity, 296 stability, 295 tissue-based, 318 wide measurement range, 295 Nanobiomagnetism, 180181 Nanobiomaterials, 532549 carbon nanotubes, 536539 hydrogel, 549 nanobioceramic, 540542 alumina and zirconia ceramics, 542 bioglass ceramics, 541542 crystalline ceramics, 540541 nano versus traditional ceramics, 542 nanocomposite, 543548 cellulose nanocomposite, 546547 chitin, 547548 chitosan, 548 nanodiamonds, 539540 nanotitanium (NanoTi), 535536 PEEK polymer, 549 polymeric materials, 533535 liquid crystal polymers, 535 parylene, 534535 polyimide, 534 silicone, 534 Nanobionics and nanoengineered prosthetics, 513 applications, 550574 cardiovascular prostheses, 571572 cerebrospinal fluid drainage systems, 572573 neuroprostheses, 566571 orthopedic prostheses, 550566 plastic and reconstructive prostheses, 573574 definition, 520523 ethical issues, 574577

autonomy, 576 benefits versus risks, 575 identity, privacy, and accountability, 575576 informed consent, validity of, 576 justice, 575 treatment versus enhancement, 577 history, 517520 manufacture, 529532 beam lithography, 530 dip-pen nanolithography, 532 jet printing, 531 lithography, 529530 micro and nano contact printing, 530531 photolithography, 530 scan probe lithography, 531 nanobiomaterials, 532549 carbon nanotubes, 536539 hydrogel, 549 nanobioceramic, 540542 nanocomposite, 543548 nanodiamonds, 539540 nanotitanium (NanoTi), 535536 PEEK polymer, 549 polymeric materials, 533535 nanotools classifications and their applications, 516f risks, classes of, 578 safety issues pertinent to nanobionics and prosthetics, 577579 types and classifications, 523529 cardiovascular prostheses, 527528 cerebrospinal fluid drainage systems, 527 dental prostheses, 528529 myoelectric prostheses, 528 neuroprostheses, 526 ophthalmic prostheses, 527 orthopedic prostheses, 524525 plastic and reconstructive prostheses, 526 Nanobubbles (NBs), 73 Nanocarriers, 2, 89 Nanoceramics, 525, 542, 566 Nanocomposite (NC) hydrogels, 409412 for biomedical applications, 410f from carbon-based nanomaterials, 411 from inorganic nanoparticles, 411 from metal and metal-oxide nanoparticles, 412 next generation of, 412 from polymeric nanoparticles, 411 Nanocomposites, 543548 cellulose, 546547 chitin, 547548

597

598

Index

Nanocomposites (Continued) chitosan, 548 Nanoconstruction, 252 Nanodevice, 252, 275 Nanodiamonds, 539540 Nanoemulsion, 198, 200201 Nanoengineered prostheses, types of, 525f Nanofiber matrices, 526 Nanofibrous scaffolds, 526 using electrospinning, 451452 “Nanomaterial surface,” concept of, 24 Nanomedicine, 2021, 106107, 131132, 159, 523 endocytosis mechanism of, 107111, 112f Nanoparticle action in biological system, 107111 limitation of, 117 in cellular and subcellular system, 111116 antigen-specific action, 112 epidermal growth factor receptor-mediated action of nanoparticles, 114115 folate receptor-mediated action, 113 integrins receptor-mediated action, 115 neonatal Fc-receptor-mediated action, 115116 receptor-mediated action, 113 transferrin receptor-mediated action of nanoparticles, 114 endocytosis mechanism of nanomedicine, 107111 biological pathway, 111 macropinocytosis, 110111 phagocytosis, 108 pinocytosis, 109110 future scopes, 132 interaction of nanoparticle in biological system, 122126 interaction with DNA, 125126 interaction with lipids, 123124 interaction with proteins, 124125 interaction with smaller biomolecules, 126 intracellular and subcellular targeted action, 117122 cytoplasm-targeted action, 119 endoplasmic reticulum and Golgi apparatustargeted action, 119 endosome/lysosome-targeted action, 117118 mitochondria-targeted action, 119120 nucleus-targeted action, 121122 in pathophysiological condition, 116 pharmacological action, 126128 genomic action of nanoparticles, 126127

metabonomic action of nanoparticle, 127128 proteomic action of nanoparticle, 127 therapeutic application, 128132 in cancer, 128129 in infectious disorders, 130131, 131f in neurological disorders, 129130 in vascular disorders, 129 Nanoparticle biosafety, 427428 Nanophotonic devices, 527 Nanoscience, 331332 Nanosilver, 563564 Nanosized crystallites in diamond coatings, 556 Nanostructured lipid carriers (NLCs), 1415, 197198, 225 Nanotechnology, 12, 404, 515, 519520, 520f Nanotextured hydroxyapatite, 555 Nanotitanium (NanoTi), 535536 Nanotrains, 266267 Nanotubes, 126127, 350, 554 Nano versus traditional ceramics, 542 Naphthofluorescein (NF), 8384 Natural hydrogels, 407 N-E5L peptide, 84 Nebulizers, 35 Neonatal Fc-receptor (FcRn), 115116 Neostigmine, 304 Neovasculature, 112 Nesterone implant, 480481 Neurological disorders effect of nanoparticles in, 129130 Neurology and central nervous system health implantable drug delivery systems in, 496 Neuroprostheses, 526, 566571 cochlear prostheses, 567 motor neuroprostheses, 570571 retinal prostheses, 568569 Neutralization plate, 559 Nee´l relaxation, 73 NickFect (NF), 8182 N-isopropylacrylamide, 340341, 353354 Nitrogenous bisphosphonate, 461 NMR, 173 Nonaarginine, 81 Nondegradable implants, 479481, 480f Nonmulberry SF-grafted PCL nanofibrous scaffolds, 454 Nonmulberry silkworms (NSF), 458459 Normal cells and abnormal cells, 30 Normal-pulse voltammetry, 319320 Norplant, 479480, 480f, 492, 534 Novadur, 395

Index

Nuclear pore complexes (NPCs), 121 Nucleic acids, delivery of, 7683 CPP-conjugated nanostructures for, 77t intracellular delivery of, 69f plasmid DNA, 7880 RNA interference molecules, 8083 Nucleus-targeted nanomedicines, 121122

O Octadecyl-p-vinylbenzyldimethylammonium chloride (OVDAC), 353354 Ocular drug delivery systems, 3537 classification of, 36 route of ocular administration, 3637 Ocular therapy, implantable drug delivery systems in, 494495 Ohm’s law, 214215 Oleic acid, 168 Oligomeric aptamers in vivo, lifetime of, 261262 nanoconstructions, 279 in sensors, 260261 Oligomerization of aptamers, 261262 Oligonucleotide sensors, 292293 Onconase, 9091 “One-pot” synthesis, 170 Open reduction internal fixation (ORIF), 558559 Ophthalmic prostheses, 527 Opsonization, 108 Optically excited nanoparticles, 569 Optical sensor, 359t Optical waveguide spectroscopy (OWS) sensor, 356357 Oral delivery systems, 28 Organic nanocomposites Organ targeting, 3039 drug delivery to brain, 3132 neoplastic disease, drug delivery to, 3739 ocular drug delivery systems, 3537 classification of, 36 route of ocular administration, 3637 pulmonary drug delivery, 3235 advantages of, 3233 disadvantages of, 33 drug delivery devices, 3435 mechanism involved in deposition of particles in lung, 3334 Orthopaedic trauma management, 441 Orthopedic prostheses, 524525, 550566 artificial joints, 550556 bone fixators, 556561 bone grafts, 561564

dental prostheses, 566 spinal implants, 564565 tendon and ligament prostheses, 565566 Orthotics, 521524 Oscar Pistorius, 518519 Osmotic pumps, 484486, 486f Osteoconductivity, 442 Osteopontin, 440441 Ovalbumin (OVA), 205 OX40 receptor, 259 Oxidoreductases, 297, 304, 321322 Ozurdex, 393t, 395

P Paclitaxel-loaded nanoparticles, 129130 Pain management, implantable drug delivery systems in, 495 Pancreatic tumor cells, 111 Particulate leaching, 443, 450 Parylene, 534535 Passive drug targeting, 45 Passive implants, 479484 biodegradable implants, 482484 nondegradable implants, 479481, 480f PEEK polymer, 549 PEGylation, 5, 27 PEP-1 peptide, 84 PepFect (PF), 8182 Perfusion flow, 227 Perfusion fluid, 227 Peripheral nervous system (PNS), 130131 Permanent/chemical hydrogels, 408 Personalized medicine, 156157 Phage-display, 91 Phagocytosis, 108 Phosphates, 167168 Phospholiposomes, 175 Photo-cross-linked alginate hydrogels (PAHs), 416 Photo electrochemical biosensors, 321 Photo electrochemistry, 321 Photolithography, 530 Photosensitizers, electrons of, 321 Physical hydrogel preparation, 406f, 407 Physical targeting, 78 Pilosebaceous unit, 192 Pin fixators, 558 Pinocytosis, 109110 Plasmid DNA, delivery of, 7880 Plasmid encoding green fluorescent protein (pGFP), 8586 Plastic and reconstructive prostheses, 526, 573574

599

600

Index

Plastic and reconstructive prostheses (Continued) breast augmentation, 573574 craniofacial reconstruction, 574 Platelet-derived growth factor (PDGF), 275, 278 “Poke and patch” strategy, 207208 “Poke and release” strategy, 207208 Poly(3-Hydroxybutyrates), 380 Poly(alkyl cyanoacrylates), 380 Poly(amidoamine) (PAMAM) dendrimer, 20, 110, 205 Poly(aminoethyl ethylene phosphate)/poly(Llactide), 179 Poly(caprolactone) (PCL), 482 Poly(ε-caprolactone), 380 Poly(ethylene glycol) (PEG), 7980, 427 Poly(glycolic acid) (PGA), 380 Poly(lactic acid) (PLA), 380381, 482 Poly(lactic-co-glycolic acid) (PLGA), 7374, 199200, 224, 375376, 380381, 482 ability to sustain and to control drug delivery, 377379 biocompatibility, issue of, 379 biodegradability, 381 case studies, 393395 Ozurdex, 393t, 395 Suprefact Depot, 393t, 394395 Zoladex, 393394, 393t drug release, 387389 biodistribution, 388389 mechanism of drug release from the implant, 387388 factors affecting degradation and drug release from, 389391 crystallinity, 389 drug load, 391 drug type, 390 manufacturing technique, 390 morphology of the matrix, 390 pH, 391 polymer composition, 389 weight average molecular weight, 390 implantable drug delivery systems (IDDS), 376377 manufacturing techniques, 384387 compression molding, 385 hot melt extrusion, 386 injection-molding, 386387 ram extrusion, 385386 solvent casting, 385 problems to overcome and opportunities, 395397 as sustained drug delivery systems, 382384

therapeutic peptides and proteins incorporated in, 391392 Poly(nisopropylacrylamide) (PNIPAM), 426 Poly(propylene imine) (PPI), 20 Polyacrylamide, 301302 Polycaprolactone (PCL), 444445 Polycarboxybetaine (PCB), 107 Polycitric acid, 170 Polydimethylsiloxane (PDMS), 534 Polydispersity index, 172 Polyester ethylene terephthalate, 565 Polyetheretherketone, 549 Polyethylene glycol (PEG), 5, 107, 168, 170 Polyethyleneoxide, 171 Polyethylene vinyl acetate (PEVA), 479 Polyimide, 534 Polylactic acid, 171 Polylactide-co-polyethylene glycol (PLA-PEG), 109 Polylactide/glycolidenanofiberous scaffolds, 565566 Poly-L-lysine (PLL), 119120 Polymerase chain reaction (PCR) primers, 251 Polymeric materials, 533535 liquid crystal polymers, 535 parylene, 534535 polyimide, 534 silicone, 534 Polymeric nanoparticles (PNPs), 1517, 130131, 179 nanocomposite hydrogels from, 411 Polymers used in magnetic targeting, 167171 Polymethacrylic acid, 171 Polyparaxylylene, 534535 Poly-R8 peptides, 79 Polytetrafluoroethylene, 565 Polyvinyl alcohol (PVA), 168, 170171, 481 Polyvinylpyrrolidone, 171 Potentiometric biosensors, 320 Precision drugs (PDs), 157158 progress towards, 158162 Precision medicine (PM), 155157 Profact Depot. See Suprefact Depot Programmable Automatic Shunt System, 573 Propellant infusion pumps, 486 Prostate carcinomas cells, 112 Prostate-specific membrane antigen (PSMA), 86 Prosthetic heart valves, 528, 571 Prosthetics, 520522 nanotechnology applications in, 520f Prosthodontics, 566 Protein databank, 127

Index

Protein immobilization by covalent bonding method, 300 Proteins/drugs, delivery of, 7076, 71t Proteomic action of nanoparticle, 127 Proximal interphalangeal joint (PIPJ), 553 Pseudoimmunoassay, 359t Pseudomonas aeruginosa, 560 Pseudomonas exotoxin A (PE), 8990 Pulmonary drug delivery, 3235 advantages of, 3233 disadvantages of, 33 drug delivery devices, 3435 dry powder inhalers (DPIs), 35 metered dose inhalers (MDIs), 34 nebulizers, 35 mechanism involved in deposition of particles in lung, 3334 Pulse voltammetric techniques, 319320 Pyrocarbon, 553 Pyrogenicity testing, 171173 Pyroglutamyl-amino-peptidase (PGP), 394

Q Quantum dots (QDs), 7273, 123125, 353355, 363 Quartz crystal microbalance (QCM), 358

R R8-hybrid peptides, 83 Rabies virus glycoprotein 29-amino-acid peptide (RVG29), 110 Raman spectroscopy, 221223 Ram extrusion, 385386 Receptor solution, quantification of the drug in, 215 Redox (oxidoreductase) enzymes, 304, 321322 Reduced graphene oxide (rGO), 356 Reflectometric interference spectroscopy (RIfS), 355 Reinforcement, 453, 543 Remote cyber attacks, 498499 Resonance light scattering (RLS), 355356 Response time, defined, 296 Response time of a biosensor, 296 Reticulo-endothelial system (RES), 5, 173 Retinal prostheses, 568569 Reversible/physical hydrogels, 407 rhBMP-2, 462 Rhodamine 6G, 205206 Ring fixators, 558 RNA interference molecules, delivery of, 8083

Ropivacain-loaded NLC, 198 Rotatory prosthesis, 551552

S Saarbruecken-type diffusion cells, 212 Sample preparation, imprinted nanosorbents for, 344352 “Sandwich”-type electrochemical biosensor, 314 Sanofi-Aventis, 394 Scanning electron microscopy (SEM), 219220 Scan probe lithography, 531 Schanz screws, 558 Scherrer equation, 173 Schizophrenia, 496 Scleral drug delivery, 37 Sebaceous glands, 192, 195 Second order targeting. See Cellular targeting Selectivity of a biosensor, defined, 296 Self-targeting carriers, 158 Sensitivity, defined, 294295 Sericin, 458, 460 Shape memory effect (SME), 535536 Shoulder replacement, 552 Silica-based additives, 456 Silica nanoparticles, 199200 Silicone, 534, 553 Silicone hydrogels, 419420 Silicone rubber, 473474, 494495 Silk, 457458, 463 benign aspects of, for bone tissue engineering, 448450 cocultures of multiple cells on silk scaffolds, 461462 composites, for bone tissue engineering, 454458 additives for silk scaffolds, 458 bioactive glasses, 456457 clay- or silica-based additives, 456 hydroxyapatite (HAp), 454456 silk inclusion in other substrates, 457458 growth factors delivery through silk scaffolds, 462 mechanical properties of, 449 nonmulberry silk, 458459 processing, into various formats, 450454 biopatterning, 452 computer-aided fabrication of SF scaffolds, 453454 freeze-drying, 452453 hydrogels, 453 knitted scaffolds, 452

601

602

Index

Silk (Continued) nanofibrous scaffolds using electrospinning, 451452 particulate leaching, 450 silk microparticles and microfibers as reinforcements, 453 sericin, 458, 460 silk fibers, 446 from silkworms, 447448 sources of, 445446 spider silk, 448, 459 Silk fibroin (SF), 443, 447, 455, 457458 computer-aided fabrication of, 453454 Silk microparticles and microfibers as reinforcements, 453 Silver nanoparticles, 563564 Single-photon emission computed tomography (SPECT), 160 Single-walled carbon nanotubes (SWCNTs), 7980, 536538 SiRNA (short interfering RNA), 8081 Skin, 187188 anatomy, 190f appendages, 192 model, 212215 quantification of the drug on, 215 structure, 189192 Skin penetration of drugs, nanoparticles influence in. See Drugs skin penetration, nanoparticles for Smart polymers, 180181 Sodium hydrogen exchangers (NHEs), 118 Sol-gel method, 546 Solid lipid nanoparticles (SLNs), 1215, 13t, 193, 197198, 200 Solid microneedles, 207208 Solid-supporting electrodes, 304, 321322 Solvent casting, 385 Spider silk, 448, 459 Spidroin, 448 Spinal implants, 564565 Square wave voltammetry, 319320 Stabilizers, 392 Stainless steel wires, 535536 Staphylococcus aureus, 560 Staphylococcus epidermidis, 309, 560 Static and flow-through diffusion cells models, 212 “Stealth” liposomes, 1011 Steinman pins, 559 Stereochemistry, 86 Stimuli-responsive hydrogels, 408, 426

Stratum corneum (SC), 188, 218 Streptavidin, 276, 278 Stripping method, 319320 Subcellular targeting, 30 Subconjunctival administration, 37 Subcutaneous MedLaunch Implant Program, 496 Subretinal prostheses, 568 Superabsorbent polymers, 421423 Superparamagnetic iron oxide (SPIO), 128129 Superparamagnetic iron oxide nanoparticle (SPION) system, 175, 178179 for gene delivery systems, 177178 Superparamagnetic nanoparticles, 175, 180 Suprachoroidal prosthesis, 568 Suprefact Depot, 393t, 394395 Surface coating, 27, 173 Surface-enhanced Raman scattering (SERS) sensors, 313314, 356 Surface plasmon resonance (SPR) spectroscopy, 352, 356 Sustained drug delivery systems, PLGA matrices as, 382384 Sweat glands, 192 Synchromed pump, 487 Synthetic dressing, 418 Synthetic hydrogels, 407 Systematic Evolution of Ligands by EXponential enrichment (SELEX), 251 Szeto-Schiller peptide, 85

T Tape stripping, 217f, 218, 223226 Targeted cancer therapies, 158 Targeted drug delivery system (TDDS), 17, 2122, 2829. See also Drug targeting Targeted therapies, 155156 “Targeting/targeted drug”, 158 Tat peptide, 6768, 7273, 8586 Tear film, 3536 Tendon and ligament prostheses, 565566 Terephthalic polyethylene polyester, 565 Tetanus toxoidloaded chitosan nanoparticle distribution, 208 Theranostics, 179 Therapeutic application of nanoparticle, 128132 in cancer, 128129 in infectious disorders, 130131 in neurological disorders, 129130 in vascular disorders, 129 Therapeutic molecules, applications of CPPs in delivery of, 7083

Index

delivery of drugs and proteins, 7076 delivery of nucleic acids, 7683 Therapeutic peptides, 391392 Thiol groups, 168 Third order targeting. See Subcellular targeting Three-dimensional DNA origami, 272273 Thrombin aptamer hetero-oligomers, 263266 Thrombin-binding aptamer, 258 Thrombin-binding nanoconstruction, 260261 Tight junctions, 31 Tissue-based nanobiodevices, 318 Tissue culture-based human skin models, 212213 Tissue dressing, hydrogels in, 417419 Tissue regeneration, hydrogels in, 414417 Titania nanotubes (TNT), 535536, 561 Titanium, 564 Ti-PEEK nanocomposites, 549 Titanium-di-oxide (TiO2) nanoparticles, 123, 126127 Titanium oxide nanoparticles, 125126 Titanium/titanium alloys, 443444, 457458 Topical drug delivery, 37 Total shoulder replacement, 552 Transcellular route, 193194 Transducer, 292294, 318 Transepidermal water loss (TEWL), 198, 213214 Transepithelial electrical resistance (TEER), 176177 Transferring receptor (TfR), 114 Transforming growth factor-α (TGF-α), 114115 Translocation, 107108 Transmission electron microscopy (TEM), 172173, 219220 Trapeziometacarpal joint, 553 Trigonist. See Suprefact Depot Triphenylphosphonium (TPP), 119120 Triple helices, 254255 Triple negative breast cancer (TNBC) cells, 74 Trisaccharide-substituted chitosan oligomers (TSCO), 110111 Tuberculosis (TB), 495

Tweezer action, 281 Two-dimensional DNA origami, 271 Two-dimensional gel electrophoresis, 127 Tyrosinase, 304305 Tyrosine kinase receptor (TRK) pathway, 110111

U Ultraflexible liposomes, 196197 Ultrasonic nebulizer, 35 Uni-lamellar vesicles, 10 Upconverting nanoparticles (UCNPs), 354

V Vascular disorders, effect of nanoparticles in, 129 Vertical diffusion cells, 211212, 225 Vesicular DDSs. See Liposomes Viadur, 485 Vinylidenefluoride, 479 Viral vectors, in gene therapy, 76 Virus-like particle (VLP), 82 Voltammetry, 318319

W Watson-Crick base pairs, 252 Women’s health, implantable drug delivery systems in, 492 Wrist, 553

X X-ray diffraction measurements, 173 X-ray photoelectron spectroscopy, 173

Z Zero-order drug release, 377, 484485 Zero order models, 216 Zeta potential of liposomes, 204 Zirconia, 542 Zoladex, 393394, 393t, 483, 493494

603