1 Introduction - UEF Electronic Publications

Sami Myllymaa

Novel Micro- and Nano-technological Approaches for Improving the Performance of Implantable Biomedical Devices

Publications of the University of Eastern Finland Dissertations in Forestry and Natural Sciences

SAMI MYLLYMAA

Novel Micro- and Nanotechnological Approaches for Improving the Performance of Implantable Biomedical Devices Publications of the University of Eastern Finland Dissertations in Forestry and Natural Sciences Number 14

Academic Dissertation To be presented by permission of the Faculty on Sciences and Forestry for public examination in the Auditorium, Mediteknia Building at the University of Eastern Finland, Kuopio, on Friday 19th November 2010, at 2 p.m. Department of Physics and Mathematics

Kopijyvä Oy Kuopio, 2010 Editors: Prof. Pertti Pasanen, Prof. Tarja Lehto, Prof. Kai Peiponen Distribution: Eastern Finland University Library/Sales of Publications P.O. Box 107, FI-80101 Joensuu, Finland Tel: +358-50-3058396 http://www.uef.fi/kirjasto

ISBN 978-952-61-0216-0 ISSN 1798-5668 ISSNL 1798-5668 ISBN 978-952-61-0217-7 (PDF) ISSN 1798-5676 (PDF)

Author’s address:

Department of Physics and Mathematics University of Eastern Finland P.O. Box 1627 FI-70211 KUOPIO FINLAND E-mail: [email protected]

Supervisors:

Prof. Reijo Lappalainen, Ph.D. Department of Physics and Mathematics University of Eastern Finland Prof. Juha Töyräs, Ph.D. Department of Physics and Mathematics University of Eastern Finland

Reviewers:

Prof. Timo Jämsä, Ph.D. Department of Medical Technology University of Oulu Prof. Sami Franssila, Ph.D. Department of Material Science and Engineering Aalto University School of Science and Technology

Opponent:

Prof. Jukka Lekkala, Ph.D. Tampere University of Technology Department of Biomedical Engineering

ABSTRACT

Biomaterials are used in a wide variety of in vivo applications, ranging from joint and dental implants to neural prostheses. The ultimate success or failure of implants mainly depends on the biological interactions (molecular, cellular, tissue) at the implant/tissue interface. Recent advances in micro- and nanotechnology offer a great opportunity to develop intelligent biomaterials and the next generation of implantable devices may well be of achieving the desired tissue-implant interaction and resolving various biomedical problems. The main aim of this thesis work was to design, fabricate and evaluate a novel flexible microelectrode array suitable for use in sub- or epidural electroencephalographic recordings. Other aims were to investigate the opportunities to improve the electrochemical and biological properties of neural interfaces using modern micro- and nanotechnology tools as well as to test whether the micropatterning of thin films can be used to guide the cellular response on biomaterial surface. The developed microelectrode array was implemented on polyimide with platinum which achieved both mechanical flexibility and high quality electrochemical characteristics as demonstrated via impedance spectroscopy. Somatosensory and auditory evoked potentials were successfully recorded with epidurally implanted array in rats with excellent signal stability over two weeks. Subsequently, the signal levels declined, most probably due to the thickening of dura and the growth of scar tissue around the electrodes. It was hypothesized that one obvious reason for this limited life-span was the poor biocompatibility of photosensitive polyimide used as an insulation material in these arrays. However, this possibility was excluded by in vitro cytotoxicity studies according to ISO 10993-5 standard. Furthermore, ultra-short pulsed laser deposition was demonstrated to be an effective method to produce nanotextured platinum surfaces as well as ultrasmooth insulators for further development of neural interfaces.

Experiments with osteoblast-like cells and mesenchymal stem cells on micropatterned biomaterial surfaces indicated that even partial coating of silicon with a biocompatible material is an effective way to enhance the cytocompatibility of siliconbased biomedical micro-electromechanical systems. Moreover, it was demonstrated that not only the chemical composition of the materials, but also the shape, edges (height) and size of the features used for surface patterning have a remarkable effect on cell guidance. Overall, the results of the present thesis provide a solid basis for the further development of neural interfaces as well as other types of implantable devices.

National Library of Medicine Classification: QT 36, QT 37, WL 102, WV 270 INSPEC Thesaurus: biomedical materials; biomedical electrodes; thin films; microelectrodes; electroencephalography; bioelectric potentials; polymer films; silicon; platinum; metals; adhesion; surface texture; surface topography; surface energy; nanopatterning; nanotechnology; microfabrication; vapour deposited coatings; pulsed laser deposition Yleinen suomalainen asiasanasto: biomateriaalit; implantit; polymeerit; pii; platina; pinnoitus; pinnoitteet; pintarakenteet; pintailmiöt; kuviot; mikrotekniikka; nanotekniikka; mikroelektrodit; EEG

Preface This study was carried out during the years 2005-2010 in the Department of Physics and Mathematics, the Department of Neurobiology and BioMater Centre, University of Eastern Finland, Microsensor Laboratory, Savonia University of Applied Sciences, Kuopio, and the Department of Medicine, Helsinki University Central Hospital. I wish to express my deepest gratitude to everyone who has contributed to the studies included in this thesis. Especially, I wish to mention the following persons. First of all I would like to thank my principal supervisor, Professor Reijo Lappalainen for giving me the opportunity to work in his research group and providing great facilities and scientific guidance for research work. I am also grateful to my second supervisor, Professor Juha Töyräs for his professional guidance and never-ending enthusiasm and optimism during this thesis work. I want to thank my official reviewers Professor Timo Jämsä and Professor Sami Franssila for their constructive criticism and valuable suggestions. I am also grateful to Ewen MacDonald for linguistic review of this thesis. The studies included in this interdisciplinary thesis would not have been possible without the excellent collaboration with colleagues in the several research units. Professor Heikki Tanila is gratefully acknowledged for his valuable ideas and guidance related to neuroscience. Kaj Djupsund is gratefully thanked for his patient testing of microelectrode array prototypes in the animal model. Professor Yrjö T. Konttinen, Emilia Kaivosoja and Vesa-Petteri Kouri are sincerely thanked for great co-operation in studies aimed to clarify the interaction phenomena between cells and engineered biomaterial surfaces. Professor Mikko Lammi, Virpi Tiitu, Sanna Miettinen and Aila Seppänen are gratefully acknowledged for their contributions in cytotoxicity testing. Moreover, I would like to thank all other co-authors in

the original publications of this thesis including Professor Juha E. Jääskeläinen, Irina Gureviciene and Tarvo Sillat. I sincerely thank all of the members in the Biomaterial Technology Research Group and BioMater Centre for providing supportive working atmosphere. In particular, the contributions of Hannu Korhonen, Markku Tiitta, and Juhani Hakala have been valuable. Aimo Tiihonen is sincerely thanked for technical assistance in electronics. The former and current staff of Microsensor Laboratory, including Matti Sipilä, Pasi Kivinen, Mikko Laasanen and Ari Halvari is greatly acknowledged for providing excellent microfabrication facilities and giving their contributions to my work. Picodeon Ltd. Oy is acknowledged for providing nanotextured depositions. I am greatly indebted to my parents Seija and Heimo for their love and support during my life. I want to thank my parents-inlaw, Anneli and Jukka, for their help and encouragements. I am also grateful to my friends and relatives for their support. Finally, I owe my deepest thanks to my beloved wife Katja, for her irreplaceable love, support, and understanding. As a research colleague, your contribution to my research has been invaluable. Words are not enough for me to describe how important you are as a wife and as the mother of our two little daughters. Dearest Lumia and Neelia, you are the sunshine of my life and have helped me to remember that there are more important and valuable things in life than work. This study was financially supported by Finnish Funding Agency for Technology and Innovation (TEKES), the National Graduate School of Musculoskeletal Diseases and Biomaterials, Otto A. Malm Foundation, Ulla Tuominen Foundation, Foundation for Advanced Technology of Eastern Finland and COST, European Cooperation in Science and Technology.

Kuopio, October 2010

Sami Myllymaa

ABBREVIATIONS AND SYMBOLS

µCP AEP AFM Ag Ag-AgCl Al Al2O3 ALP ANOVA AP Au BHK-21 Bio-MEMS Ca Cl CN CNS Cr CVD CZ DBS DC DLC DMEM DNA ECG EEG EIS EMG ESEM EP FCS FDA FZ HMDS

microcontact printing auditory evoked potential atomic force microscopy silver silver-silver chloride aluminium alumina (aluminium oxide) alkaline phosphatase analysis of variance action potential gold baby hamster kidney fibroblast biomedical micro-electromechanical systems calcium chlorine carbon nitride central nervous system chromium chemical vapour deposition Czochralski process for silicon crystallization deep brain stimulation direct current diamond-like carbon Dulbecco’s modified Eagle’s medium deoxyribonucleic acid electrocardiography electroencephalography electrical impedance spectroscopy electromyography environmental scanning electron microscope evoked potential fetal calf serum Food and Drug Administration float-zone process for silicon crystallization hexamethyldisilazane

hMSC Ir ISO ITO K KOH MEA MSC MEMS MS MSCGM MTS

Na OC PGMEA PBS PCB PDMS PE PI PNS PSPI Pt PVD RIE SAM SD SEM SEM SEP SFE Si SS Ta Ti

human mesenchymal stem cell iridium International Organization for Standardization indium tin oxide potassium potassium hydroxide microelectrode array mesenchymal stem cell micro-electromechanical systems mineral staining mesenchymal stem cell growth medium cell proliferation assay based on (3-(4,5dimethylthiazol-2-yl)-5-(3 carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium)-salt sodium osteocalcin propylene glycol methyl ether acetate phosphate buffered saline printed circuit board polydimethylsiloxane polyethylene polyimide peripheral nervous system photosensitive polyimide platinum physical vapour deposition reactive ion etching self assembled monolayer standard deviation of the mean scanning electron microscopy standard error of the mean somatosensory evoked potential surface free energy silicon stainless steel tantalum titanium

TiN USPLD UV VEP ZIF

titanium nitride ultra-short pulsed laser deposition ultraviolet visual evoked potential zero insertion force

CW

Warburg (polarization) capacitance

f Ra

frequency

R pv

peak-to-valley roughness

RS

resistance of electrode solution

R pv

Warburg (polarization) resistance

Vm

transmembrane potential

V pp

peak-to-peak signal amplitude

Z

impedance total surface free energy

S  SD  SP 

average surface roughness as arithmetic mean deviation of surface

dispersive component of total surface free energy polar component of total surface free energy zeta potential

LIST OF ORIGINAL PUBLICATIONS

This thesis is based on, but not limited to, results presented in the following original publications, which are referred in the text by their Roman numerals (I-VI): I

Myllymaa, S., Myllymaa, K., Korhonen, H., Gureviciene, I., Djupsund, K., Tanila, H. & Lappalainen, R. (2008) Development of flexible microelectrode arrays for recording cortical surface field potentials. In: Proceedings of the 30th Annual International Conference of the IEEE engineering in Medicine and Biology Society, Vancouver, British Columbia, Canada, 20-24 August, 2008, pp. 3200-3203.

II

Myllymaa, S., Myllymaa, K., Korhonen, H., Töyräs, J., Jääskeläinen, J.E., Djupsund, K., Tanila, H. & Lappalainen, R. (2009) Fabrication and testing of polyimide-based microelectrode arrays for cortical mapping of evoked potentials. Biosensors and Bioelectronics 24, 3067-3072.

III

Myllymaa, S., Myllymaa, K., Korhonen, H., Lammi, M.J., Tiitu, V. & Lappalainen, R. (2010) Surface characterization and in vitro biocompatibility assessment of photosensitive polyimide films. Colloids and Surfaces B: Biointerfaces 76, 505-511.

IV

Myllymaa, S., Myllymaa, K. & Lappalainen, R. (2009) Flexible implantable thin film neural electrodes. In: Recent Advances in Biomedical Engineering, In-Tech, Vukovar, Croatia, Editor: Ganesh R. Naik, ISBN 978-953-307-004-9, Chapter 9, pp. 165-189.

V

Myllymaa, S.*, Kaivosoja, E.*, Myllymaa, K., Sillat, T., Korhonen, H., Lappalainen, R. & Konttinen, Y.T. (2010) Adhesion, spreading and osteogenic differentiation of mesenchymal stem cells cultured on micropatterned amorphous diamond, titanium, tantalum and chromium coatings on silicon. Journal of Materials Science: Materials in Medicine 21 (1), 329-341. (* equal contribution)

VI

Kaivosoja, E.*, Myllymaa, S.*, Kouri, V.-P., Myllymaa, K., Lappalainen, R. & Konttinen, Y.T. (2010) Enhancement of silicon using micropatterned surfaces of thin films. European Cells and Materials, 19, 147-157. (* equal contribution).

The publications are reprinted with the kind permission of the copyright holders. This thesis also contains unpublished results related to publications III and VI as well as unpublished data focused on the electrochemical characterization of nanotextured bioelectrode surfaces.

AUTHOR’S CONTRIBUTION

Publications I and II concern the development of flexible microelectrode array suitable for recording cortical surface field potentials. The original idea of this electrode system was proposed by H. Tanila. The author of this thesis carried out most of the development work related to the electrode design, material selection and fabrication processes. Apart from depositions of metal thin-films, which were performed by H. Korhonen and R. Lappalainen, the author carried out all array fabrication steps and characterization experiments with a contribution from K. Myllymaa. In vivo testing of electrodes was performed by K. Djupsund and I. Gureviciene. The author wrote the articles, after receiving constructive comments from the coauthors. R. Lappalainen supervised the work. Publication III concerns the cytocompatibility of the novel photosensitive polyimide used as an insulator material in our sensor prototypes (papers I and II). The design and fabrication of test samples were performed by the author. The author also conducted the surface characterization together with H. Korhonen and K. Myllymaa. The author planned cell experiments together with V. Tiitu, who also performed the experimental studies. The author was responsible for analyzing the data, presenting the results and writing the article, while receiving constructive comments from the co-authors. R. Lappalainen, V. Tiitu and M.J. Lammi supervised the study. The author was mainly responsible for writing publication IV (book chapter), which is a literature review discussing the main requirements and features of flexible thin film neural electrodes, this being supplemented with personal results in the area of sensor development. Publications V and VI concern the interactions between cells and biomaterial surfaces. R. Lappalainen and Y.T. Konttinen originally presented the idea of studying the effect of surface micropatterning on the cellular response. Apart from thin film depositions, performed by H. Korhonen and R. Lappalainen, the author carried out all sample design and sample fabrication

steps in collaboration with K. Myllymaa. The surface characterization analysis was mainly performed by the author with some contributions from H. Korhonen and K. Myllymaa. Experimental work and data analysis on cellular studies was planned and carried out by E. Kaivosoja, with contribution of V.-P. Kouri in paper VI, and Y.T. Konttinen. The author and E. Kaivosoja contributed equally to preparing results and writing paper V, assisted by R. Lappalainen and Y. T. Konttinen. E. Kaivosoja and Y.T. Konttinen chiefly wrote paper VI, assisted by S. Myllymaa and R. Lappalainen. These studies were supervised by R. Lappalainen and Y.T. Konttinen.

Contents 1 INTRODUCTION ........................................................................... 21 2 BIOMEDICAL MICRO-ELECTROMECHANICAL SYSTEMS .......................................................................................... 25 2.1 Microfabrication techniques ........................................................... 26 2.1.1 Photolithography ...................................................................... 27 2.1.2 Soft lithography........................................................................ 29 2.1.3 Thin film deposition ................................................................. 32 2.1.4 Lift-off processing..................................................................... 34 2.1.5 Etching ..................................................................................... 35 2.2 Substrate materials........................................................................... 37 2.2.1 Silicon ...................................................................................... 37 2.2.2 Glass ......................................................................................... 38 2.2.3 Polymers................................................................................... 38 2.3 Advantages of microfabrication .................................................... 40 2.4 Bio-MEMS applications................................................................... 42

3 ORIGIN OF BIOELECTRIC SIGNALS ...................................... 45 3.1 Excitable nerve cell .......................................................................... 45 3.1.1 Cell membrane and ion channels ............................................. 47 3.1.2 Transmembrane potential and equilibrium potentials ............. 47 3.2 Synaptic potentials and action potentials ..................................... 49 3.3 Recording electrical activity of the brain ...................................... 51

4 BIOELECTRODES .......................................................................... 55 4.1 Implantable electrodes .................................................................... 56 4.2 Material requirements for implantable electrodes ...................... 59 4.2.1 Electrode materials ................................................................... 60 4.2.2 Substrate materials .................................................................. 63

5 BIOCOMPATIBILITY.................................................................... 67 5.1 Biocompatibility testing .................................................................. 70

5.2 Cell-biomaterial interactions .......................................................... 71 5.2.1 Effect of surface topography on cellular responses ................... 71 5.2.2 Effect of wettability properties on cell-biomaterial interactions .............................................................................. 75 5.2.3 Effect of charge distribution on cell-biomaterial interactions .. 76 5.2.4 Effect of surface micropatterning on cellular responses ........... 78

6 AIMS OF THE PRESENT STUDY ............................................... 79 7 MATERIALS AND METHODS ................................................... 81 7.1 Micro- and nanofabrication of electrodes and other samples ... 82 7.1.1 Preparation of photomasks ....................................................... 82 7.1.2 Flexible polyimide-based microelectrode arrays ....................... 82 7.1.3 Micropatterned biomaterial surfaces on silicon ....................... 86 7.1.4 Tailored surfaces produced by ultra-short pulsed laser deposition ................................................................................. 88 7.1.5 Spin-coated polyimide films for cytotoxicity testing................ 89 7.2 Microscopic characterization of surfaces ...................................... 90 7.2.1 Scanning electron microscopy ................................................. 90 7.2.2 Atomic force microscopy .......................................................... 90 7.2.3 Contact angle measurements and determination of surface free energies.............................................................................. 91 7.3 Electrochemical characterization of surfaces ............................... 92 7.3.1 Electrochemical impedance spectroscopy ................................. 92 7.3.2 Zeta potential measurements ................................................... 94 7.4 Recording of evoked potentials in rats ......................................... 95 7.5 Cell culture studies .......................................................................... 96 7.5.1 Cell lines................................................................................... 96 7.5.2 MTS assay................................................................................ 98 7.5.3 Scanning electron microscopy of cultured cells ....................... 98 7.5.4 Immunofluorescence and confocal laser scanning microscopy ............................................................................... 99 7.6 Statistical analyses ......................................................................... 101

8 RESULTS ........................................................................................ 103 8.1 Evaluation of microelectrode arrays ........................................... 103 8.2 Evaluation of nanorough electrode surfaces .............................. 109

8.3 Cytocompatibility testing of insulator materials ....................... 111 8.4 Micropatterned biomaterial coatings .......................................... 115 8.4.1 Behaviour of human mesenchymal stem cells ........................ 116 8.4.2 Behaviour of osteoblast-like (SaOS-2) cells ............................ 120

9 DISCUSSION ................................................................................ 125 9.1 Microelectrode arrays for intracranial recordings .................... 125 9.2 Material science strategies to improve the performance of microelectrode arrays .......................................................................... 128 9.3 Micropatterned biomaterial surfaces .......................................... 130

10 CONCLUSIONS.......................................................................... 133 11 REFERENCES .............................................................................. 135 APPENDIX: ORIGINAL PUBLICATIONS I-VI

1 Introduction Developments in micro-electromechanical systems (MEMS) technology have introduced a variety of breakthrough products in the fields of microelectronics, telecommunications and the automotive industry (Judy 2001, Madou 2002). In recent years, the biomedical applications of micro-electromechanical systems (bio-MEMS) have attracted growing interest in many application areas such as diagnostics, therapeutics and tissue engineering (Saliterman 2006, Grayson et al. 2004, Nuxoll & Siegel 2009, Betancourt & Brannon-Peppas 2006, Bashir 2004, Urban 2006). Bio-MEMS is a powerful technology capable of ever-greater functionality and cost reduction in smaller biomedical devices for improved diagnostics and treatment. The merging of biology with micro/nanotechnology has been postulated to trigger a scientific and technological revolution in the future (Kotov et al. 2009). Although some applications such as blood analysis cartridges, catheter pressure sensors and cochlear implants have been already commercialized, the vast majority of biomedical applications are still being investigated or undergoing clinical trials (Saliterman 2006, Urban 2006). Brain research can be considered as a one of the most challenging scientific areas. The very first biological MEMS devices, i.e. multi-sensor neural probes, were developed for neuroscientists in order to facilitate studying of neuronal activities at the tissue and cellular level (Urban et al. 2006). In addition to pure scientific research, there has been a growing interest in the clinical applications of stimulating and recording neural electrodes and prostheses (DiLorenzo & Bronzino 2008). These neural interfaces can benefit many patients with neural disorders such as impaired hearing (cochlear implant) or neuropathic pain (deep brain stimulators). Some extremely challenging applications such as retina implants which may be able to restore the vision or sieve and cuff electrodes potentially

21

1 - Introduction

evoking nerve regeneration in paralyzed patients are still under intensive development (Stieglitz & Mayer 2006b, DiLorenzo & Bronzino 2008). To date, a wide variety of neural electrodes have been utilized in neural interfaces starting with the early electrolytefilled micropipettes and subsequent metal electrodes to the current emerging MEMS-based electrodes. The vast majority of these MEMS devices have been fabricated on silicon (Si) even though there is a huge mismatch between mechanical properties of the neural tissue and Si causing many adverse tissue effects (Polikov et al. 2005). Moreover, Si is not per se a biocompatible material (Voskerician et al. 2003), and it is a poor substrate for cell adhesion, even being slightly cytotoxic (Liu et al. 2007). These weaknesses of Si could limit the integration of Si-based devices into the human body. Recently, there has been a growing interest toward developing polymer-based interfaces that could be flexible enough to mimic biological tissue, reducing mechanical damage and evoking less adverse tissue reactions (Stieglitz & Mayer 2006b, Cheung 2007, HajjHassan et al. 2008, DiLorenzo & Bronzino 2008). Biomaterials are used in many in vivo applications, ranging from joint and dental implants to neuroprosthetic devices. Depending on particular specifications for each such application, a set of different material characteristics, such as mechanical, chemical and electrical properties, influence the performance of biomaterial/prosthetic devices in a specific manner (Williams 2008). In all cases, however, the ultimate success or failure depends on the biological interactions (molecular, cellular, tissue) at the implant-tissue interface (Puleo & Nanci 1999, Williams 2008, Navarro et al. 2008, Polikov et al. 2005). Therefore, several approaches have been introduced to modify surface properties, such as surface topography on the microand nanoscales (Flemming et al. 1999, Martinez et al. 2009), surface energy (van Kooten et al. 1992, Hallab et al. 2001, Lim et al. 2008) and surface charge (Krajewski et al. 1996, Krajewski et al. 1998, Cai et al. 2006), in order to achieve significant effects on the functions of proteins and thus on cells. These enhanced

22

1 - Introduction

effects on cellular functions are crucial in subsequent improvements in new tissue formation and integration of implants in the surrounding tissues. Recently, much attention has been paid to mesenchymal stem cells (MSCs) due to their role in tissue repair and their potential in regenerative medicine and tissue engineering in which these kinds of cells are grown on biomaterial scaffold, which provides structural support and substrate for cellular adhesion (Zippel et al. 2010, Pittenger et al. 1999, Stoddart et al. 2009). The ability to influence the adhesion, distribution and behaviour of cells on biomaterial surfaces and the knowledge of the nature of cell-substrate interaction has therefore become increasingly important. In this thesis, novel flexible polyimide (PI)-based microelectrode arrays (MEA) suitable for sub- or epidural electroencephalographic (EEG) recordings were developed. The suitabilities of different MEMS materials as well as their surface modifications for implantable applications were investigated using electrochemical testings and cell experiments. Moreover, the effect of surface micropatterning achieved using photolithographic and thin film techniques on the behaviour of MCSs and osteoblast-like cells was studied.

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1 - Introduction

24

2 Biomedical microelectromechanical systems Micro-electromechanical systems refer to miniature devices that are fabricated using techniques originally developed and widely utilized in the microelectronics (integrated circuits) industry, and then modified, e.g. by adding mechanical components, such as beams, gears, and springs, for the creation of microstructures and microdevices such as sensors and actuators. MEMS, as well as its interchangeable acronyms, Micromachines (popular in Asia) and Microsystems (popular in Europe), refer to the miniature devices which have at least some of their dimensions in the micrometer range. According to the widest definition, MEMS comprises all devices and systems produced by micromachining other than integrated circuits or other conventional semiconductor devices (Judy 2001). At the beginning of MEMS era, in the 1970s and 1980s, the field was dominated by mechanical applications, but today most new applications are either communication/information-related or chemical and biological in nature (Madou 2002). MEMS technology has enabled low-cost, high-functionality microdevices in some widespread application areas, such as printer cartridges for ink jet printing and the accelerometers responsible for deployment of airbags in modern automobiles (Judy 2001). In recent years, the biological and biomedical applications of MEMS, which are commonly referred to as bio-MEMS, have gained increasing world-wide interest (Saliterman 2006, Nuxoll & Siegel 2009, Urban 2006). Bio-MEMS are generally defined as ‘‘devices or systems, constructed using techniques inspired from micro/nanoscale fabrication, that are used for processing, delivery, manipulation, analysis, or construction of biological and chemical entities’’ (Bashir 2004). At the present moment, bio-MEMS technology is a topic of intense research and

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2 – Biomedical micro-electromechanical systems

development activity with a wide variety of applications in the fields of cell and molecular biology, pharmaceutics and medicine so that it includes the areas of therapeutics, diagnostics and tissue engineering (Saliterman 2006, Grayson et al. 2004, Nuxoll & Siegel 2009, Betancourt & Brannon-Peppas 2006, Voldman et al. 1999, Bashir 2004).

2.1 MICROFABRICATION TECHNIQUES

A number of techniques are used to form bio-MEMS objects with dimensions ranging from the micrometer to millimeter scale. Some of these techniques have been adopted directly from the industry of integrated circuits whereas some others have been specifically developed for this novel purpose (Judy 2001, Voldman et al. 1999, Saliterman 2006, Li et al. 2003). The microfabrication process is typically a process flow that utilized these techniques in a sequential manner to produce the desired structure. MEMS devices can be built within the bulk of a substrate material in what is referred to as bulk micromachining, or if it is on the surface of the substrate it is known as surface micromachining (Madou 2002). However, a combination of bulk and surface micromachining is commonly used (Voldman et al. 1999). The most important microfabrication techniques in bioMEMS are photolithography, soft lithography, thin film deposition and etching (Saliterman 2006, Li et al. 2003). Photolithography is used to transfer the desired shape onto a material through the selective exposure of a photosensitive polymer (Madou 2002). Soft lithography is a set of different patterning techniques based on polymer block used as a stamp, mold or stencil for carrying out surface patterns and microstructures (Xia & Whitesides 1998, Li et al. 2003, Saliterman 2006). Thin film deposition which consists of numerous different techniques is used to form thin layers with thickness ranging from atomic layer to a few microns on the surface of a substrate. Etching is used to remove material selectively from the surface

26


of the microsystem by either chemical or physical processes (Madou 2002). 2.1.1 Photolithography Photolithography is a basic microfabrication technique widely employed to create the desired patterns onto a material. The photolithographic patterning is composed of a number of process steps (Fig. 1). photoresist

resist deposition and soft baking

substrate

UV exposure

photomask mask alignment (x, y, φ)

latent image

development

positive photoresist

negative photoresist

Figure 1: Photolithographic patterning process. A photomask with opaque regions in the desired pattern is used to selectively expose a photosensitive polymer (photoresist). Depending on the polarity of the resist used, it will become more soluble (positive-tone photoresist) or crosslinked (negative photoresist) after ultraviolet light illumination, thus generating the desired pattern after developing in liquid chemicals.

Firstly, a pattern is drawn using computer assisted design software and this is transferred onto a mask. The photomasks are typically transparent glass plate blanks covered with an opaque material (usually chromium, Cr) in the defined pattern (Madou 2002). The masks are usually prepared by a commercial mask manufacturer using electron beam or laser writing (Wu et

27


al. 2002). If the feature sizes and tolerances in the mask are relatively large, an option may be to utilize low-cost transparency films as a photomask (Deng et al. 1999, 2000, Wu et al. 2002, Gale et al. 2008). After producing the mask, a photosensitive polymer, i.e. photoresist, is spin-coated onto a substrate material, such as a Si wafer. The resulting photoresist thickness, t, is a function of spinner rotational speed, solution concentration, and molecular weight (determined by intrinsic viscosity) and can be estimated by (Madou 2002):

t

KC  



.

(1)

In this equation, K is the overall calibration constant, C is the polymer concentration (g/100 ml), η is the intrinsic viscosity and ω is the spinner rotational speed (rotations per minute, rpm). α, β and γ are exponential factors dependent on process. Typically, spinning speeds of 1500-8000 rpm are used to achieve resist thicknesses of about 0.5-2 µm (Madou 2002). The photoresist film is then prebaked at 75 to 100°C on a hotplate or in an oven to remove solvents and to promote adhesion of the resist layer to the substrate (Saliterman 2006, Madou 2002). In the following step of exposure, the photomask is placed on top of the photoresist-coated wafer in close proximity or even in contact with it and then ultraviolet (UV) light is used to illuminate the photoresist film through the photomask. With this procedure, the solubility difference between the exposed and unexposed regions is achieved, and depending on the polarity of the photoresist used, exposed or unexposed areas can then be removed by dissolving them in a developing solution. In the case of a positive-tone photoresist, the exposed regions break down and become more soluble in the developing solution, whereas in a negative-tone photoresist, the exposed areas become crosslinked and insoluble in the developing solution. The resulting photoresist pattern can now be used as a protective mask in following microfabrication processes such as in the etching step to prevent the covered substrate to be

28


removed and in thin film deposition to pattern metallization layer on the surface of substrate. After the followed process is finished, the photoresist can be removed, leaving the pattern design on the substrate. The removal of photoresist is usually performed by sonication in an organic solvent, e.g. acetone. (Madou 2002) The resolution of photolithography depends on the quality of mask and the wavelength of applied electromagnetic radiation (light). The typical wavelengths used are 436 nm, 365 nm, 248 nm and 193 nm (Madou 2002). The shorter the wavelength of light used, the higher resolution, i.e. smaller feature sizes are possible. Although photolithography is the main technology in the production of microscale features in MEMS, it has also a few short-comings. It requires clean-room facilities, and therefore it is not an inexpensive technology; it is poorly suited for patterning nonplanar surfaces and it is directly applicable only to a limited set of photosensitive polymers (Xia & Whitesides 1998). Furthermore, photolithography requires the use of strong solvents, meaning that it is poorly compatible with the patterning biological molecules used in many applications e.g. in biosensing, medical implants and the control of cell adhesion and growth (James et al. 1998, Li et al. 2003, St. John et al. 1997). 2.1.2 Soft lithography Soft lithography consists of a set of non-photolithographic patterning methods based on self-assembly and replica molding for carrying out micro- and nanofabrication (Saliterman 2006). It provides a low-cost, effective, and biocompatible strategy for the formation and manufacturing of surface patterns and structures with feature dimensions ranging from 30 nm to 100 µm (Xia & Whitesides 1998). The key element of soft lithography is an elastomeric block with patterned relief structures on its surface that subsequently utilized in different soft lithographic techniques such as in microcontact printing, stencil patterning and microfluidic patterning (Li et al. 2003). The commonest material used for the production of elastomeric block is polydimethylsiloxane (PDMS). This material has several

29


desirable properties such as durability, optical transparency, biocompatibility, flexibility, gas permeability and amenability to different surface modifications (Whitesides et al. 2001, Li et al. 2003). The major advantages of soft lithography compared to conventional photolithography are that numerous solid and liquid materials other than photoresists can be patterned, and it is suitable for curved or complicated surface geometrics and even 3D structures (Xia & Whitesides 1998, Gale et al. 2008). The soft lithography process starts with conventional photolithographic steps to create a PDMS block which will be used later as a mold or a stamp (Fig. 2a). A

B

PDMS

Alkanethiol ”ink”

photoresist

Au

Si Substrate

Pour PDMS over master Microcontact print PDMS Si

PDMS

Cure and peel off stamp PDMS

Substrate

Remove stamp SAM (2-3 nm) Substrate

Figure 2: (a) Fabrication of a polydimethylsiloxane (PDMS) stamp. PDMS is poured onto a silicon wafer coated with patterned photoresist. After the curing step, the stamp can be removed. (b) Microcontact printing with the PDMS stamp. The stamp is coated with alkenethiol ink, and then it is pressed on gold coated substrate. After application of gently pressure for a few seconds, a self-assembly monolayer (SAM) is formed on gold surface. Figure is inspired by Xia & Whitesides (1998).

Photoresist is spin-coated onto a substrate (e.g. Si wafer). The height of the resist layer (i.e. pattern relief) can be controlled by the viscosity of the solution and the spin speed (Madou 2002). EPON™ SU-8, originally developed by IBM, and recently marketed by MicroChem Corp. (Newton, MA, USA), is one the most popular photoresists being used in stamp fabrication (Li et al. 2003). This negative-tone epoxy-resist permits the fabrication

30


of tall structures, even more than 1 mm in height with a superior aspect ratio (Lorenz et al. 1997, 1998). The photoresist layer is illuminated with UV light through a photomask with the desired pattern, and then developed. In the case of SU-8, the resist areas exposed to UV light will remain in the development. Liquid PDMS precursor is then cast over the patterned substrate and cured at an elevated temperature. After cooling, the PDMS stamp can be peeled off from the master. A replica of the original template in a functional material can be generated by molding against the patterned PDMS block. (Li et al. 2003, Xia & Whitesides 1998, Whitesides et al. 2001) The microcontact printing (µCP), also known as microstamping, is based on the patterned transfer of the material (‚ink‛) from the surface of the PDMS stamp onto the receiving surface on a sub-micrometer scale (Fig. 2b) (Qin et al. 2010, Xia & Whitesides 1998, Kumar & Whitesides 1993). The material of interest to be transferred, e.g. alkanethiol, is applied onto the surface of the stamp with patterned surface relief. The dried stamp is then placed face down on the substrate, e.g. gold (Au)-coated glass, and gentle pressure is applied for a few seconds (Qin et al. 2010). The elastic PDMS material enables an excellent contact between the stamp and the substrate surface, and molecules that touch the surface are transferred from the stamp to the substrate, forming a self-assembly monolayer (SAM). The formed SAM layer can be used as a protective layer in subsequent microfabrication steps such as etching or deposition (Xia & Whitesides 1998). The chemicals used in the µCP to form SAMs typically have a chemical formula of Y(CH2)nX, where Y is the anchor and X is the headgroup (Saliterman 2006). A typical anchor group in alkanethiols is sulfide since this promotes the binding of thiol very tightly to Au. Typically, CH3 and COOH are used as the headgroups. The choice of headgroup has a great influence on the wettability properties, i.e. hydrophobicity/hydrophilicity, of the surface and subsequently to protein and cell binding. Furthermore, the headgroups can be chemically modified, e.g. an arginineglycine-aspartate peptide that enhances cell attachment (Hersel

31


et al. 2003, LeBaron & Athanasiou 2000, Shi et al. 2008, Chua et al. 2008). The µCP offers the advantage over traditional photolithographic patterning methods in that it requires no strong chemicals, making it suitable for patterning biologically active layers (Xia & Whitesides 1998, Whitesides et al. 2001, Saliterman 2006). In addition to alkanethiols, other peptides, proteins, polysaccharides and other molecules can also be stamped (Bernard et al. 1998, Branch et al. 1998, James et al. 1998, Li et al. 2003). Stencil patterning is the second soft lithography technique in which a membrane stencil is fabricated by casting PDMS to the top of a photoresist master, which creates holes with the shape of the master features. The PDMS stencil can be used as a mask for selective adsorption of cell-adhesive proteins to promote cellular patterning. (Folch et al. 2000, Li et al. 2003, Folch & Toner 2000) Micromoulding in capillaries, the third soft lithographic technique, employs a PDMS mold to build up microchannels against a substrate. These microchannels can be used to pattern fluid materials onto a substrate (Kim et al. 1995, 1996). A low viscosity prepolymer is applied at the open ends of the channels. The channels become filled with polymer due to capillary forces. After curing of the prepolymer, the PDMS mold is removed and the three-dimensional polymer microstructures are revealed. Subsequently, these microstructures can be used for selective delivery of different cell suspensions to specific locations of a substrate resulting in micropatterns of attached cells (Folch & Toner 1998, Tan & Desai 2003). 2.1.3 Thin film deposition The application of thin layers of materials is a common procedure in microfabrication. Thin films refer to thin material layers ranging from the thickness of atomic layers to several micrometers in thickness (Madou 2002). Thin films can play a structural or functional role in the device process (Smith 1995, Hsu 2008). All material classes, i.e. metals, ceramics, polymers, composites and biological compounds, can be deposited. They can be used either as sacrificial layers or mask layers that

32


selectively protect the substrate material during etching or they can be used as electrical insulators or conductors such as electrodes and transmission lines in sensors. Two common thin film materials used as insulators in Si-based MEMS are silicon dioxide (SiO2) and silicon nitride (Si3N4) (Voldman et al. 1999). Deposition techniques can be divided into physical vapour deposition (PVD) and chemical vapour deposition (CVD) methods (Madou 2002). In general, the PVD methods are based on the production of a condensable vapour by physical means and its deposition on a substrate as a thin film. Vacuum or low pressure gaseous environment is used in the PVD processing, in which deposition species are atoms or molecules or both. The most common PVD techniques are (1) evaporation in a vacuum, (2) sputter deposition, (3) arc-vapour deposition, (4) laser ablation, and (5) ion plating (Saliterman 2006). On the other hand, the CVD methods are based on chemical reactions which take place at a heated substrate surface to deposit a solid film. Gas phase reactions between chamber gases are not desirable because they often lead to poor adhesion, low density and high detect films. The main advantages of the CVD methods are the possibility to fill holes, cavities and other 3D structures, good adhesion between coating and substrate, and excellent thickness uniformity of coating. The major shortcomings are the toxicity of the by-products as well as a high processing temperature, being above 600°C which eliminates the use of CVD with heat sensitive materials like polymers (Hsu 2008, Saliterman 2006, Madou 2002). In addition, thin films can be produced by other techniques, such as spin coating, electrolytic deposition and electroless deposition. Spin-coating is typically used to deposit thin polymer films such as photoresists in photolithography. In the electroplating process, the substrate is immersed in an electrolyte solution. When a voltage is applied between a substrate (working electrode) and an inert counter electrode, such as platinum (Pt) in the liquid, chemical reduction-oxidation processes take place resulting in the formation of a layer of material on the surface of the substrate (Paunovic & Schlesinger

33


2006, Schlesinger & Paunovic 2000). An electroless plating process does not require any external electrical potential but the deposition spontaneously happens on any surface which forms a sufficiently high electrochemical potential with the electrolyte solution. Although electroless deposition (compared to the electroplating) is much easier to set-up and use, without the need for an electrical supply and its connections to electrodes, it is also more difficult to control with regards to film thickness and uniformity. Electroplating and electroless plating are typically used in MEMS to form conductive metal (e.g. platinum, gold, copper, nickel) or conductive polymer (e.g. polypyrrole, polyaniline) thin films (Madou 2002). 2.1.4 Lift-off processing Lift-off is a simple and easy method for patterning thin film metal layers. It is especially useful for patterning catalytic metals, such as Pt, which do not lend themselves well to direct wet etching (Madou 2002, Hsu 2008). The lift-off process starts with a photolithographic step, in which the photoresist layer is deposited and patterned as described in chapter 2.1.1. Thin film of desired material is then deposited all over the substrate, covering the photoresist and areas in which the photoresist has been cleared. Thereafter, the substrate is immersed in a solvent that dissolves the remaining soluble photoresist underneath the metal, and lifts off the metal. Only the metal which has been deposited directly on the substrate leaves and forms the final pattern on the substrate. In order to achieve clean results in liftoff processing, photolithographic and thin film deposition processes need to be optimized (Franssila 2004). Undercut photoresist patterns are beneficial and can be realized by reducing exposure dose and/or increasing the developing time of negative photoresist. The substrates should be kept at temperatures below 200 to 300°C during metal deposition to avoid hardening and flow of photoresist (Madou 2002).

34


2.1.5 Etching Etching techniques are used to create topographical patterns on a surface via the selective removal of material. These techniques can be roughly divided into wet and dry etching which are based on the utilization of liquid chemicals (etchants) or reactive ions/vapour-phase etchant, respectively (Madou 2002). The etching process is isotropic in its nature if it removes material in all directions equally, leading to mask undercutting and a rounded etch profile (Fig. 3a), or it is anisotropic if removal occurs with different etch rates in different directions in the material (Fig. 3b,c) (Voldman et al. 1999). In the simple wet etching technique, the sample is immersed in a container filled with a liquid solution (etchant). The sample is covered by a mask which leads to the selective removal of material (Madou 2002). The most critical task is to find a mask material that will not dissolve or at least is etched much more slowly than the sample material in question. A wide variety of etchants such as buffered hydrofluoric acid, Aqua regia, Piranha solution (i.e. mixture of sulfuric acid and hydrogen peroxide), potassium hydroxide (KOH), tetramethylammonium hydroxide and hydrochloric acid can be used to etch different MEMS materials (Williams & Muller 1996, Williams et al. 2003). Unfortunately, etchants are usually isotropic leading to etch profiles with large undercuts particularly when one is machining thick films. Additionally many wet etch chemicals are toxic and, thus they are poorly suited for the state-of-the-art bio-MEMS processes. Anisotropic wet etch is possible on some single crystal materials, such as Si in which anisotropic etchants (e.g. KOH) dissolve Si rapidly in the direction , but almost not at all in the direction of . Thus cavities with trapezoidal cross-sections with a characteristic 54.7° sidewall will be created on a [100]oriented wafer (Fig. 3c) (Madou 2002, Voldman et al. 1999).

35


A

B

C

etching mask 54.7

substrate

Figure 3: Different etching profiles resulted in (a) isotropic wet etching, (b) anisotropic dry etching, and (c) anisotropic wet etching of silicon using potassium hydroxide. Figure is modified from Voldman et al. (1999).

In order to achieve anisotropic etching with high resolution, different dry etching methods are used. Reactive ion etching (RIE) is one of the most popular dry etching methods in which the substrate is placed inside a reactor and several gases are introduced (Kovacs et al. 1998). The gas molecules can be made to disintegrate into ions using a radiofrequency (13.56 MHz) power source. Accelerated ions collide onto the surface of the material to be etched, react with the surface molecules forming another gaseous material. In addition to this chemical aspect, there is also a physical part present in the process. If the plasma particles, i.e. neutral radicals and ions, have enough energy, they can remove atoms out of the materials without any chemical reaction. This process can be quite complicated since there are many parameters to be adjusted. However, with optimal adjustment, it is possible to etch almost straight down without undercutting, which provides much higher resolution compared to wet etching. (Madou 2002) A special improvement of RIE technique is the deep RIE, based on the so called ‚Bosch process‛ (Laermer 1996) in which an etch depth of hundreds of micrometers can be achieved with vertical sidewalls. The deep RIE process is based on two different alternating gas compositions in the reactor. The first gas composition forms a polymer film on the surface of the substrate, and the second gas composition etches the substrate (Saliterman 2006). The polymer is immediately sputtered away by the second etching gas, but this happens only on the

36


horizontal surfaces and not the sidewalls protecting them from etching (Laermer 1996). Very high etching aspect ratios, even 50:1, can be achieved (Kovacs et al. 1998). Dry etching is a widely used method in the integrated circuit industry to achieve high resolution features, but in many cases it is not essential in the production of bio-MEMS.

2.2 SUBSTRATE MATERIALS

Traditionally, MEMS devices have been fabricated on microelectronics related materials, such as Si and glass. Recently, however, there has been a growing interest towards rubber and plastic substrate materials due to their suitable mechanical properties, enhanced biocompatibility, rapid prototyping and inexpensive manufacturing techniques available (Wilson et al. 2007). 2.2.1 Silicon Silicon is the most widely used material in microchips and MEMS devices since it is straightforward to grow oxide layers to form dielectrics on its surface and it has excellent semiconductor properties over a wide temperature range. Silicon wafers used in semiconductor/MEMS industry is produced using two alternative crystallization methods: Czochralski process (CZ) and float-zone (FZ) process (Pearce 1988). In the CZ method, a seed crystal is dipped into molten Si and pulled upwards with simultaneous rotation. By optimizing the rate of pulling and the speed of rotation as well as other relevant parameters such as temperature, it is possible to extract a large-diameter cylindrical single crystal ingot from the melt. Czochralski process is a more common and cheaper method compared to FZ, but the wafers produced using the CZ method contain slightly more impurities. Impurities in the CZ wafers originate from the quartz crucible, containing the Si melt that dissolves during the process. In the FZ method, a local melted zone produced by radiofrequency field is slowly passed along a polycrystalline rod. The seed

37


crystal is used at one end in order to launch the single crystalline growth. Impurities in the molten zone tend stay in the molten zone rather than being incorporated into the solidified region, hence allowing an extremely pure single crystal region being left after the molten zone has passed. The largest Si ingots, produced using CZ method, are 300-450 mm in diameter and 1 to 2 meters in length today. The diameters of the FZ ingots are typically less than 200 mm due to the surface tension limitations encountered during the crystallization process (Franssila 2004). Thin Si wafers with a typical thickness of 0.25-1.0 mm, are cut from these ingots and then polished to a very high smoothness. Silicon can be doped with impurity atoms such as boron or phosphorous, i.e. changing it into n-type or p-type Si, in order to enhance the electrical conductivity. (Madou 2002, Franssila 2004) In addition to the excellent semiconductor properties of Si, it has also superb mechanical properties, enabling the design of MEMS structures (Petersen 1982). Silicon micromachining techniques are established and widely available. The major drawbacks of Si for bio-MEMS applications include its limited biocompatibility, unfavourable mechanical properties (rigidity, fragility) as well as the relatively high material and processing cost (Cheung 2007), which makes it less attractive for use in disposable biomedical devices. 2.2.2 Glass In spite of the limited number of micromachining techniques available for glass substrates (compared to Si), this material is used in bio-MEMS applications due to some unique properties, most notably optical transparency. Glasses with a wide variety of compositions can be used. Fused silica (pure amorphous SiO2) and borosilicate (e.g. Pyrex®) are examples from commonly used glass materials (Voldman et al. 1999). 2.2.3 Polymers Polymers are very attractive for bio-MEMS applications due to their suitable mechanical properties, enhanced biocompatibility,

38


optical transparency, ability to modify their bulk and surface properties in order to improve their functionality, and the availability of rapid prototyping and mass production processes (e.g. injection molding, hot embossing) (Wilson et al. 2007). These issues enable the fabrication of low-cost disposable microdevices for clinical applications. Polydimethylsiloxane, polyimide, epoxy resins and parylene are some examples of widely utilized polymers in bio-MEMS (Seymour et al. 2009, Cheung 2007, Saliterman 2006). Polyimide has a long history of use in microelectronics e.g. as an encapsulation and stress buffer material on semiconductor chips as well as a substrate material in flexible printed boards/cables (Ghosh & Mittal 1996). It possesses numerous desirable properties such as excellent resistance to solvents, strong adhesion to metals and metal oxides and good dielectric properties (Wilson et al. 2007). Recently, PI has been introduced as a substrate/insulating material in numerous bio-MEMS applications, such as neural interfaces (Boppart et al. 1992, Rousche et al. 2001, Hollenberg et al. 2006, Molina-Luna et al. 2007, Takahashi et al. 2003, Cheung et al. 2007, Owens et al. 1995, Stieglitz 2001, Mercanzini et al. 2008, Patrick et al. 2008, Spence et al. 2007) due to its desirable mechanical and dielectric properties and good biocompatibility (Wilson et al. 2007, Stieglitz et al. 2000, Richardson et al. 1993). Polydimethylsiloxane, also familiar as a soft lithographic stamp material, is known to be biocompatible and it is approved for use as an implanted material in medical devices, i.e. approved by the Food and Drug Administration (FDA). It has been used as a biomaterial in catheters, pacemakers, and ear and nose implants (Visser et al. 1996). PDMS is a highly flexible and optically transparent material and it is permeable to gases. It is also amenable to different surface modifications, which can be used to tailor the biochemical functionality of the PDMS (McDonald & Whitesides 2002, Mata et al. 2005). PDMS can be conformed to submicron features to develop different microstructures (Xia & Whitesides 1998) such as microfluidic components (Ng et al. 2002, Folch et al. 1999) for bio-MEMS.

39


An epoxy-based negative photoresist SU-8 is one of the most widely used thick-film photoresists in MEMS (Lorenz et al. 1997, 1998, Wilson et al. 2007). It has been utilized as a structural material because it can be produced with a wide range of thicknesses from below 1 µm to even more than 1 mm (Lorenz et al. 1998). Recently, it has been employed in numerous bioMEMS applications, including analytical microfluidic systems (Tuomikoski 2007) and neural sensors (Hollenberg et al. 2006, Tijero et al. 2009, Altuna et al. 2010). Parylene is a thermoplastic polymer that can be vapour-deposited at room temperature to create pin-hole free, optically transparent barrier coatings that are stress-free, chemically and biologically inert, and minimally permeable to moisture (Saliterman 2006, Seymour et al. 2009).

2.3 ADVANTAGES OF MICROFABRICATION

Recent developments of MEMS and associated nanotechnology represent great opportunities in the design of miniature, smart, and low-cost biomedical devices that could revolutionize biomedical research and clinical practice (Urban 2006, DiLorenzo & Bronzino 2008, Nuxoll & Siegel 2009). Micro- and nanotechnology can be used either to improve the performance of an existing device or to enable development of an entirely new device. The bio-MEMS devices have many advantages over their macroscopic counterparts. The MEMS techniques enable development of small size devices that may be easier to use (e.g. portable and hand-held devices), or they can be truly innovative (e.g. implantable or even injectable devices) (Saliterman 2006, Grayson et al. 2004). The small size often saves on the costs of reagents, time and money (Voldman et al. 1999). For example, portable ‚point-of-care‛ hematological devices and test kits permit physicians to diagnose the patient’s condition more rapidly (Betancourt & Brannon-Peppas 2006). Diagnostic bioMEMS devices, also known as lab-on-a-chips or micro-total analysis systems, can be used to detect cells, microorganisms, viruses, proteins, deoxyribonucleic acid (DNA) and related

40


nucleic acid (Bashir 2004, Guber et al. 2004). In general, miniaturization offers several advantages over conventional analytical methods. Smaller sample volumes enable reducing assay cost and less waste disposal (Voldman et al. 1999). Due to the low sample volume, short diffusion distances and fast heating, less time is required for diagnostics (Saliterman 2006). Furthermore, miniaturized parallel operation allows highthroughput analysis, which plays an important role in genomic research and drug discovery (Bashir 2004). On the other hand, miniaturization of devices increases their surface area to volume ratio, leading often to situations where surface effects dominate volume effects. The larger surface area is essential in some applications e.g. to ensure heat removal avoiding adverse heat effects when high electric fields are used (Voldman et al. 1999). In addition, MEMS fabricated electrodes and sensor systems allow measurements and stimulations with higher spatial and temporal resolution than with conventional macroscale counterparts, and hence enable more precise biomedical research or clinical diagnostic and therapy (DiLorenzo & Bronzino 2008, Stieglitz & Mayer 2006a). Most MEMS fabrication processes can be performed simultaneously on many, even thousands of devices, similarly to the case of manufacturing of microchips on Si wafers. This kind of batch processing enables high volume manufacturing at low unit cost (Judy 2001). Processes are very reproducible, minimizing variations between objects that easily arise among individually constructed devices. Lastly, polymers and other cheap materials can be utilized to provide bio-MEMS devices, making feasible the manufacturing of disposable devices. Disposable devices are particularly crucial when handling blood and other biological fluids that may contain hazardous substances such as human immunodeficiency virus or hepatitis virus. Disposable devices also eliminate the risk of crosscontamination and subsequent analysis errors which are common in re-used devices (Wang & Soper 2007).

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2 – Biomedical micro-electromechanical systems 2.4 BIO-MEMS APPLICATIONS

During recent years, the bio-MEMS technology has been under intensive research and development activity. Numerous bioMEMS applications have been implemented in the fields of cell and molecular biology, pharmaceutics and medicine, including the areas of therapeutics, diagnostics and tissue engineering (Voldman et al. 1999, Bashir 2004, Cheung & Renaud 2006, Nuxoll & Siegel 2009, Saliterman 2006, Betancourt & BrannonPeppas 2006, Grayson et al. 2004). Although some applications, such as blood analysis cartridges, cochlear implants and deep brain stimulators (DBS) have been already commercialized, the vast majority of bio-MEMS objects are under research or undergoing clinical trials (Stieglitz & Mayer 2006a, DiLorenzo & Bronzino 2008). On the other hand, most of the commercially available systems are designed for in vitro diagnostics. The development cycle from the lab prototype to commercial manufacturing of implantable devices is long, perhaps even 1020 years, due to the extremely challenging environment inside the human body as well as strict requirements to pass the approval process (classified as FDA class III devices) (Ratner 1996). Examples of implantable bio-MEMS devices are listed in Table 1. A more detailed description of the electrodes used in neural prosthesis is given in Section 4 of the thesis.

42

2 – Biomedical micro-electromechanical systems Table 1: Examples of implantable Bio-MEMS devices. Device

Brief description of its function

Manufacturer(s) General reference(s)

Cardiac pacemakers

Regulation of the heart beating by delivering electrical impulses to the electrodes in contact with heart muscles

Boston Scientific, USA; Medtronic, USA; St. Jude Medical, USA

Das et al. 2009

Cochlear implants

Electrical stimulation of auditory nerve through microelectrodes implanted in the inner ear to restore hearing

AllHear, USA; Advanced Bionics, USA; Cochlear, Australia; MEDEL, Austria; Neurelec, France

Loizou 1999, Eddington 2008, Wilson & Dorman 2008a,b

Deep-brain stimulators

Through intracortically implanted electrodes electrical stimulation is delivered to targeted regions of the brain to treat severe neurological disorders, such as Parkinson disease

Medtronic, USA

Kumar et al. 1997, Kern & Kumar 2007, Richardson 2008

Spinal cord stimulators

Electrical stimulation of spinal cord through electrodes implanted in epidural space in order to treat intractable pain and motor disorders

Advances Neuromodulation Systems, USA; Medtronic, USA

Waltz 1997, Mailis-Gagnon et al. 2004, Burton & Phan 2008

Vagus nerve stimulators

Electrical stimulation of vagus nerve in the neck via an implanted lead wire to treat certain types of intractable epilepsy and major depression

Cyberonics, USA

Ardesch et al. 2007, Boon et al. 2001

Drug delivery silicon chips

Controlled drug release from microfabricated Si reservoirs by applying electrical potential

Microchips, USA

Santini et al. 1999, 2000

Visual prosthesis

Electrical stimulation of neural tissue of blind patients to restore vision. Different approaches in which microelectrode arrays are implanted on or under the retina, around the optic nerve, and on or in the visual cortex are under development

Second Sight, USA; Retina Implant AG, Germany

Humayun et al. 2003, Stieglitz 2009, Greenberg 2008

43


44

3 Origin of bioelectric signals Bioelectric signals are generated as a result of electrochemical activity in certain type of cells, referred to as excitable cells which are the basic components of nervous, muscular and glandular tissue (Malmivuo & Plonsey 1995). An electrical potential difference, i.e. the transmembrane potential, exists between the internal and external environment of the cell (Hille 1992). When appropriately stimulated to cross a limiting threshold, the excitable cell will generate an action potential due to a flow of charged ions across the permeable cell membrane with the ions passing through ion channels. These action potentials are transferred from one cell to adjacent cells via their cellular extensions, i.e. axons and dendrites due to synaptic transmission (Tortora & Grabowski 2000). The activity of a group of excitable cells (i.e. population activity) generates an electric field that propagates through the biological tissue and this can be measured by field-potential measurements with surface electrodes attached to the surface of tissue, e.g. cerebellum cortex, or skin (Nunez & Srinivasan 2006, McAdams 2006). The field-potentials commonly monitored in a patient are those producing electrocardiograms (ECG) from the heart, EEG from the brain and electromyograms (EMG) from contraction of muscles.

3.1 EXCITABLE NERVE CELL

A nerve cell, also known as a neuron, is an excitable cell in the nervous system whose main function is to receive, handle and transmit cellular information via electrochemical phenomena. Neurons are the basic components of the brain, the spinal cord

45

3 – Origin of bioelectric signals

and the peripheral nerves. They can be divided into two categories: (1) principal cells that receive input from nearby or distant neurons and project to neurons in other nuclei; (2) interneurons that modulate the activity of principal cells in the same structure. Principal cells are usually excitatory and use glutamate as the neurotransmitter, whereas interneurons are usually inhibitory and use gamma-amino butyric acid as the neurotransmitter (Kandel et al. 2000). There are a huge variety of different nerve cells with sizes in the range of 600 – 70 000 µm3 (Schadé & Ford 1973). Most nerve cells consist of three main parts: a cell body and two kinds of cell extensions – dendrites and axons. The cell body (soma) is similar to that of all other cells containing the nucleus, mitochondria, endoplasmic reticulum, ribosomes, and other organelles (Tortora & Grabowski 2000). Most of the protein synthesis happens in this part of cell. Dendrites are typically short and highly-branched cellular extensions, through which the vast majority of signal input to the neuron occurs (afferent signals). Due to numerous branches (‚dendrite tree‛), a neuron may receive nerve impulses from thousands of other neurons (Nunez & Srinivasan 2006). The axon of a neuron is a single, cylindrical-shaped projection which carries nerve impulses away from the cell body toward the axon terminals passing the message on to other neurons. Although axons are very thin, typically about 1-20 µm in diameter, they can be very long (Malmivuo & Plonsey 1995). The length of axons, e.g. the human sciatic nerves, may be more than one meter long. The axons of many neurons are covered with an insulating layer called the myelin sheath, which is formed by either of two types of glial cells: Schwann cells or oligodendrocytes which insulate peripheral neurons or those of the central nervous system, respectively (Tortora & Grabowski 2000). The myelin sheath is not continuous but the gaps, known as nodes of Ranvier occur at regular intervals, dividing the sheath into short sections (Malmivuo & Plonsey 1995). This myelination structure enables much faster conduction (i.e. saltatory conduction) of nerve impulses so the impulse is transported faster than in nonmyelinated axons.

46


3.1.1 Cell membrane and ion channels The living cell is surrounded by the cell membrane which is an approximately 5 nm thick phospholipid bilayer where the hydrophobic tails of phospholipid molecules are pointing inside the membrane whereas the negatively polarized, hydrophilic heads of these molecules are directed outward of the membrane (Alberts et al. 1994). The main functions of the cell membrane are to form the boundaries for the intracellular components of the cell, maintaining cell homeostatis and regulating the incoming and outgoing substances (Tortora & Grabowski 2000). The cell membrane is permeable to only specific types of ions and molecules and thus the concept of selective permeability is used to describe this property of the membrane. The excitability of the membrane is based on embedded protein molecules, which form tiny ion channels and allow the flow of sodium (Na+), potassium (K+), calcium (Ca2+) and chloride (Cl-) ions through the membrane (Hille 1992). Ion channels can be divided into leakage channels and gated channels (Tortora & Grabowski 2000). The difference between them is that leakage channels are always open, whereas gated channels open and close in response to some kind of stimulus. The gated channels, on which the excitability of neural cells is based, can be divided into three groups according to the stimulus involved: (1) voltage-gated ion channels, (2) ligand-gated ion channels and (3) mechanically gated ion channels. The state (i.e. open or closed) of voltage-gated channels is adjusted by a change in the transmembrane potential. Ligand-gated ion channels open and close in response to specific chemical stimuli including neurotransmitters, hormones and ions, whereas mechanically gated ion channels are regulated by mechanical stimuli such as vibration, pressure or tissue stretching (Tortora & Grabowski 2000, Kandel et al. 2000). 3.1.2 Transmembrane potential and equilibrium potentials The selective permeability of the cell membrane maintains ion concentration difference between the inside and outside of the cell membrane (i.e. concentration gradient). For mammalian

47


muscle, the K+ concentration of the intracellular space is 155 mmol/l, while that within the extracellular space is only 4 mmol/l (Hille 1992). In addition, Na+, Ca2+ and Cl- ionic concentrations inside the cell differ greatly from the outside (Table 2). To balance the positive and negative charges inside the cell, there are also other negatively charged molecules, which are too large to permeate through the membrane. Table 2: Ion concentrations for mammalian muscle cells (Hille 1992). Ion

Intracellular concentration (mM/l)

Extracellular concentration (mM/l)

Na+

12

145

K+

155

4

Ca Cl-

2+

-4

10

1.5

4.2

123

This existing concentration gradient promotes the outflow of K ions and inflow of Na+, Ca2+ and Cl- ions in order to balance the ion concentrations. Potassium ions diffuse from inside to outside, making the interior of the cell more negative than the outside. Additionally, there is an active transportation of ions against a concentration gradient. These Na-K pumps transport two K+ ions into the cell for every three Na+ ions that the enzyme pumps out of the cell, and thus there is a net loss of positive charges within the cell (Hille 1992). The transmembrane potential, Vm, is defined as a potential difference between the internal and external environment of a cell. The resting membrane potential corresponds to the value of Vm when the cell is in the resting state in its natural, physiological environment. In neurons, the typical resting potential is around -70 mV (Tortora & Grabowski 2000). The membrane potential that just balances the ion concentration difference, i.e. equilibrium potential when no current flows across the membrane, can be determined for each specific ion by using the Nernst equation (derived by Water Hermann Nernst in 1888): +

48


V 

RT  Ci  ln . zF  Co 

(2)

Here R is the gas constant (8.314510 J/mol x K), T is absolute temperature, z is the valence of an ion, F is the Faraday’s constant and Ci and Co are ion concentrations inside and outside of the cell, respectively. It can be easily calculated by using the ion concentrations (Table 2) that the equilibrium potential for K+ ions at 37 °C is approximately -98 mV. The corresponding values for Na+ and Ca2+ ions are +67 mV and +128 mV, respectively. Since the equilibrium potential of K+ is quite close but not same as the value of the resting potential (typically -70 mV), at rest the cell membrane is much more permeable to K+ ions than to other ions. The Nernst equation can be extended to Golzman-HodgkinKatz equation, which produces the equilibrium potential when there are several types of ions in the intracellular and extracellular space, and when the membrane is permeable to all of them (Hille 1992). Then, the resting membrane potential can be estimated by:

VREST  

RT  PK ci , K  PNaci , Na  PCl co ,Cl  ln , F  PK co , K  PNaco , Na  PCl ci , Cl 

(3)

where PK, PNa and PCl are membrane permeabilities of K+ , Na+, Cl- ions, respectively.

3.2 SYNAPTIC POTENTIALS AND ACTION POTENTIALS

The transfer of information from one neuron to the next is based on junctions between an axon and the neighboring cell, called synapses. They can be divided into electrical and chemical synapses (Tortora & Grabowski 2000). An electrical synapse is a mechanical, electrically conductive link between two neurons which can transmit nerve impulses very quickly (Hormuzdi et al. 2004). Electrical synapses are common in the neural systems

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that require the fastest possible response, such as defensive reflexes. However, it is more common to operate with chemical synapses. The end of the axon has branching terminals through which neurotransmitter chemicals (e.g. dopamine, glutamate, serotonin) packaged into presynaptic vesicles, are released into a gap called the synaptic cleft in order to communicate with an adjacent neuron (Kandel et al. 2000). Most synapses are between axon and dendrite whereas also other combinations like axonto-axon and dendrite-to-dendrite exist (Tortora & Grabowski 2000). The width of synaptic cleft is only about 20-40 nm (Hormuzdi et al. 2004). Due to the small volume of the cleft, the neurotransmitter concentration can be raised and lowered rapidly. There are certain types of receptor molecules located on the cell membrane of the postsynaptic cell which after being occupied by neurotransmitter molecules, respond by opening nearby ion channels in the membrane, allowing charged ions to rush in or out. As a result, the local transmembrane potential of the cell is changed and thus is referred to as the post-synaptic potential. The result can be either excitatory (depolarizing) or inhibitory (hyperpolarizing) depending on the amount and type of neurotransmitter as well as the type of receptor molecules (Kandel et al. 2000). In an excitatory reaction, a net movement of positive ions into the cell changes the transmembrane potential making it less negative. Hyperpolarization is the exact opposite to depolarization, in which the change of Vm occurs in the negative direction. If the depolarizing response is great enough, the cell membrane is able to produce a characteristic bioelectric impulse called an action potential (Fig. 4) (Tortora & Grabowski 2000). At the beginning of depolarizing stimulus, the voltagegated Na+-selective ion channels become opened, allowing the flow of Na+ ions from outside to inside. The inside of the cell becomes more positive reaching a potential value of approximately +30 mV. Subsequently, more K+-selective ion channels will open allowing the flow of K+ ions from inside to outside. The flow of K+ ions returns the transmembrane potential to its resting value. The duration of the action potential is around 1 millisecond (Tortora & Grabowski 2000).

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Transmembrane potential (mV)

+30

0

Resting state Refractory period

-70

1

2

3

4

5

Time (ms)

Figure 4: The shape of an action potential. When a stimulus depolarizes the membrane beyond a threshold, an action potential is generated. At the end of depolarization phase, a membrane potential of 30 mV is reached. The flow of potassium ions returns the membrane potential to its resting value (-70 mV) during a short refractory period.

3.3 RECORDING ELECTRICAL ACTIVITY OF THE BRAIN

The human nervous system is a complex, highly organized network of billions of neurons and even more neuroglia (Tortora & Grabowski 2000). It is composed of two sub-systems: the central nervous system (CNS) and peripheral nervous system (PNS). Anatomically, the CNS contains the cerebrum, cerebellum, brainstem and spinal cord (Heimer 1995). The brain stem is a structure through which nerve fibers relay action potentials in both directions between spinal cord and higher brain centers. The cerebrum is divided almost equally into two halves, which both have subdivisions of the frontal, temporal, parietal and occipital lobes. The cerebral cortex, about 2-4 mm thick folded structure, forms the outer portion of the cerebrum (Tortora & Grabowski 2000). The surface area of the cerebral cortex varies from 1600 to 4000 cm2 and it contains perhaps as

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many as 10 billion neurons (Nunez & Srinivasan 2006). Those neurons are strongly interconnected. The surface of a large cortical neuron may be covered by as many as 105 synapses that transmit inputs from other neurons (Tortora & Grabowski 2000). Post-synaptic potentials in thousands of synchronously active pyramidal neurons produce an electric field propagating through the biological tissue and these can be measured on the scalp or on the surface of brain (subdural electrodes) or even inside the brain (probe electrodes) with an amplifier through appropriate electrodes (Nunez & Srinivasan 2006, Niedermayer & Lopes da Silva 2005). Scalp-EEG is a standard method in clinical neurophysiology to study the function of the brain. Both spontaneous activity and evoked potentials (EPs) can be measured. Spontaneous scalp-EEG is routinely utilized in many clinical cases, ranging from diagnosis of epilepsy to the analysis of sleep and the depth of anaesthesia. EPs are produced as a response to given sensory stimuli. The EPs can be classified according to the applied stimulus into three categories: somatosensory (SEP), visual (VEP) and auditory (AEP) evoked potentials (Niedermayer & Lopes da Silva 2005). The stimulus evokes electrical activity with a very short latency (delay) period over a limited area of the cerebral cortex. For example, VEPs are detected from the occipital region when a visual stimulus is presented, whereas AEPs are detected from the temporal region when an auditory stimulus is presented. Evoked potential recordings can be utilized to localize lesions, or monitor a sensory pathway during surgical procedures (Chiappa 1997). Stimulus is usually applied many times (even thousands), and then signal averaging is undertaken to achieve clean EP curves with a reduced effect of noise (Nunez & Srinivasan 2006). In clinical practice, scalp-EEG is typically measured using the standard international 10-20 electrode placement system in which 21 electrodes are attached on the scalp with exact intervals and potential differences between electrodes is recorded. The amplitudes of the scalp EEG varies from 10 to 100 µV (Niedermayer & Lopes da Silva 2005). The frequency band from 0.16 to 100 Hz is commonly adequate, but sometimes a

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much lower frequency response, down to direct current (DC) (Vanhatalo et al. 2003) or much higher (e.g. up to 3 kHz) is useful (Curio 2005). The scalp-EEG possesses good temporal resolution on a sub-millisecond scale, but the spatial resolution is poor. Spatial resolution can be improved by increasing the electrode numbers. Nowadays, multi-channel EEG caps with up to 256 electrodes are applied in experiments aimed at obtaining detailed information about brain sources (Väisänen 2008). Another choice to improve the source localization is to place the electrodes closer to the brain tissue. Both intracortical depth electrodes and subdural strip electrodes placed on the surface of cortex are used in intracranial recordings of local field potentials and occasionally spikes (Grill 2008). Compared to scalp-EEG, intracranial recordings are worthwhile in terms of their increased spatial resolution, signal amplitude level (0.5 – 4 mVpp vs. 10 – 100 µVpp), and reduced noise level, meaning that one can acquire more detailed maps directly of the brain (Lachaux et al. 2003, Niedermayer & Lopes da Silva 2005, Leuthardt et al. 2008). However, the use of penetrating electrode probes in humans is limited to the areas destined for surgical resection due to the risk of damaging brain tissue. Flexible subdural strip and grid electrodes are utilized as a less invasive alternative in some clinical cases, e.g. in mapping cerebellum cortex in patients undergoing epilepsy or brain tumor surgery. The signals recorded with subdural electrodes are considerably higher in amplitude because they are much closer to the source of the activity, and are separated from it by only a relatively high conductivity media (cerebrospinal fluid, brain parenchyma) (Nair et al. 2008, Leuthardt et al. 2008). The resistivity of the skull is estimated to be even 80 times higher than that of brain tissue (Rush & Driscoll 1968) that causes a decline in the signal amplitude and signal blurring in scalp-EEG recordings. Moreover, recordings made by subdural electrodes are almost free of artifacts (e.g. not affected by electrical signals originating from muscles) that are seen in scalp-EEG, and thus they yield a much higher signal-to-noise ratio (Nair et al. 2008).

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4 Bioelectrodes Biomedical electrodes are extensively utilized in numerous biomedical applications (McAdams 2006). Firstly, they are used to record bioelectrical events such as the ECG, EEG and EMG that intrinsically arise due to electrical activity of excitable cells. Secondly, bioelectrodes are utilized in applying electrical impulses to the body for treating neurological pain in many therapeutic methods such as in transcutaneous electrical nerve stimulation or deep brain stimulation. Deep brain stimulators consist of intracortically implanted electrodes through which a high frequency electrical stimulation is delivered to a targeted region of the brain. DBS is used in treating several neurological disorders, e.g. Parkinson’s disease (Kumar et al. 1997, Kern & Kumar 2007, Perlmutter & Mink 2006). Thirdly, they are used in alternating current impedance characterization of body tissues. These methods such as electrical impedance plethysmography and electrical impedance tomography (Cheney et al. 1999) do not monitor intrinsic biosignals emanating from the body, but inject small alternating currents/voltages and record the resultant voltages/currents. The electrical properties of the body can then be determined. Finally, bioelectrodes are utilized in electrotherapy by applying electrical potentials in order to facilitate the transdermal delivery of ionized molecules for therapeutic effect (iontophoresis). Moreover, recent activity on neuroengineering research has focused on implantable neural prosthetics to help restore the functions of the damaged nervous system (DiLorenzo & Bronzino 2006, Stieglitz & Mayer 2006b). The most successful examples of commercialized neural prosthesis are cardiac pacemakers and cochlear implants. Cochlear implant is a surgically implanted microelectronic device that utilizes an electrode array inserted into the scala tympani of the cochlea to stimulate the auditory nerve through

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the bone, enabling the deaf to hear sounds (Loizou 1999, Federspil & Plinkert 2004, Wilson & Dorman 2008a,b).

4.1 IMPLANTABLE ELECTRODES

The commercially available clinic subdural strip and grid electrodes (Fig. 5) typically consist of a number of conductive metal such as Pt, stainless steel (SS), and silver (Ag) disks, which are partially embedded in a thin sheet of biomedical grade silicone rubber (e.g., AD Tech Medical Instrument Corp., Racine, WI, USA; PMT Corp., Chanhassen, MN, USA). Typical diameter of disk contacts is about 2-3 mm, and the distance between adjacent electrodes around 10 mm.

Figure 5: Examples of clinical subdural strip and grid electrodes.

Despite major advances in microsystem technology and its great potential in the development of novel high-density MEAs, most human intracranial electrodes used to record brain activity are still handcrafted from Pt foils and silicone rubber sheet and therefore are quite bulky (see Fig. 5). Miniaturized neural microelectrodes are extensively developed approaches in neurophysiological (animal) studies aimed at understanding

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behaviour and function of the brain at the tissue and cellular level, but their human counterparts are still mainly under investigation or at best are undergoing clinical trials. Already over forty years ago it was discovered that the photolithographic, Si etching and thin-film deposition techniques originally developed for the production of integrated circuits would be also applicable for providing novel improved recording and stimulating electrodes for experiments with animal models (Wise 2005). Compared to the traditional approach of single microwires or microwire bundles (Nicolelis et al. 1997, Williams et al. 1999), microfabricated electrodes offered many advantages including precise control of the electrode sizes and separations between electrodes as well as a high degree of reproducibility in physical, chemical and electrical characteristics (Rutten 2002). Microfabrication is also a high yield, low-cost process once the design and processing steps have been optimized. Much of the pioneer work of Sibased microelectrodes has been done in three research centers in the USA, i.e. at the University of Stanford, University of Michigan, and University of Utah. Today Utah arrays (Campbell et al. 1991, Rousche & Normann 1998) and Michigan probes (Wise & Angell 1975, Najafi & Wise 1986, Vetter et al. 2004) are perhaps the best-known commercially available Si-based microelectrodes. A three dimensional Utah electrode array (Fig. 6a) consists of one hundred 1.5 mm long Si needles that project out from a 4 mm x 4 mm substrate (Campbell et al. 1991, Normann et al. 1999). The sharpened tip of each Si needle is coated with Pt. An advanced version of Utah array, i.e. Utah slanted electrode array, has also been developed (Fig. 6b). In this model, the length of Si needles can be increased from 0.5 to 1.5 mm permitting focal excitation of nerve fibers at different depths (Badi et al. 2003). Utah arrays have been used in studies of parallel information processing by the CNS at the visual and auditory cortex (Rousche & Normann 1998), and the control of muscle force and limb position by the PNS (Normann 2007). The typical Michigan probes are single-shank or multi-shank Si penetrating systems (Fig. 6c) where the electrodes are placed

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along the length of the shank allowing the measurement of neuronal activity at various depths of the brain tissue. Michigan probes have been successfully used in a number of neuroscience experiments throughout the CNS, PNS and even cardiac muscle (Vetter et al. 2004, Kipke et al. 2003, Wise 2005).

Figure 6: Silicon–based neural interfaces: (a) Utah array; (b) Slanted Utah array. These electrode arrays which both contain 100 microneedle-shaped electrodes are designed to be implanted in cerebral cortex or peripherally. (c) Michigan multi-shank probe contains active electronics integrated into the array. Reprinted by permission from Macmillan Publishers: Nature Clinical Practice Neurology (Normann 2007), copyright (2007).

Although rigid microwires and Si-based microelectrodes have long been used for recording of biosignals, the mechanical mismatch between the rigid electrode and soft biological tissue may evoke many adverse effects such as tissue damage, inflammatory reactions and scar formation (Polikov et al. 2005). Thus, there has been a growing interest in developing polymer-

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based (e.g. polyimide, benzocyclobutene, silicone rubbers and parylene-C) implants that could be sufficiently flexible to mimic biological tissue and to reduce adverse tissue reactions (Cheung 2007). The flexibility of polymer may decrease the mechanical mismatch between nervous tissue and the rigid implant reducing the risk of tissue damage and inflammation reactions. PI is one of the most widely studied polymers due to its suitable mechanical properties, biocompatibility and stability in wet microfabrication processes (Rubehn & Stieglitz 2010, Stieglitz et al. 2000). Several PI-based microelectrode constructions have been developed in the fields of neural engineering, including devices for in vitro recordings from brain slices (Boppart et al. 1992), cortical surface field potential recordings (Owens et al. 1995, Hollenberg et al. 2006, Takahashi et al. 2003, Hosp et al. 2008), intracortical multiunit neural activity recordings (Rousche et al. 2001, Cheung et al. 2007, Mercanzini et al. 2008) and for action potential recording in nerve and muscle tissues in vivo (González & Rodríguez 1997, Rodríguez et al. 2000, Spence et al. 2007). PI has also been applied for different types of sieve and cuff electrodes for interfacing with regenerating peripheral nerves (Stieglitz et al. 1997, 2000, Stieglitz 2001).

4.2 MATERIAL REQUIREMENTS FOR IMPLANTABLE ELECTRODES

The materials used in implantable neural electrodes and prostheses must fulfill strict requirements in order to ensure optimal performance and longevity. According to Geddes and Roeder (2003), there are four criteria that must be considered when selecting a material for implantable use: (1) tissue response; (2) allergic response; (3) electrode-tissue impedance; and (4) radiographic visibility. An ideal implanted electrode and/or its insulation should be biocompatible without triggering any vigorous local or generalized host response or allergic reactions. The electrode-tissue impedance must be stable and low enough to enable high-resolution measurements. Moreover,

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the materials should be radiographically visible and be compatible with magnetic resonance imaging. It has been concluded in many studies that in spite of the extremely challenging processes used in fabrication of miniaturized electrodes, electrode failures due technical reasons, such as short-circuits and broken signal transmission lines, are rare, and the stable, long-term functioning of neural implant is mainly determined by the tissue response to chronically implanted electrodes (Polikov et al. 2005, Nicolelis et al. 2003, Rousche & Normann 1998, Williams et al. 1999). A general immune activation of the brain in response to the presence of a foreign body implant has been commonly believed to be the main reason for signal deterioration of chronically implanted electrodes (Nicolelis et al. 2003, Edell et al. 1992, Schmidt et al. 1993). The biocompatibility aspects of electrode materials will therefore be very crucial in the development of new generation chronic implantable devices.

4.2.1 Electrode materials Recording bioelectrodes are functional elements that convert the ionic current around the electrode site in the body into electron current in electronic circuitry. Electrodes can be roughly divided into polarizable and non-polarizable electrodes. Perfectly polarizable electrodes are those in which no actual charge crosses through the electrode-electrolytic interface, but changes the charge distribution within the solution near the electrode when a current is applied. This kind of electrode acts like a capacitor, and only the displacement current crosses the interface. In contrast, perfectly non-polarizable electrodes allow the current to pass freely through the electrode-electrolytic interface without changing the charge distribution in the electrolytic solution. However, neither of these ‚perfect‛ electrode types can be fabricated, but some practical electrodes can come close to possessing their features. (McAdams 2006, Neuman 1978)

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A silver-silver chloride (Ag-AgCl) electrode is an almost perfectly non-polarizable electrode. Due to many desirable properties, such as small and stable offset potential, low contact impedance and low intrinsic noise, this type of electrode has been found to be an excellent electrode material for surface (skin) biopotential recordings (McAdams 2006). Unfortunately, it AgAgCl cannot be used as an implantable electrode because of the rapid dissolution of Ag. Moreover, AgCl as well as pure Ag has a toxic effect on neural tissue (Yuen et al. 1987). Noble metals such as platinum, gold, iridium, rhodium and palladium are commonly used highly polarizable electrode materials in neural applications due to their excellent corrosion resistance and biocompatibility (Geddes & Roeder 2003, McAdams 2006). In the extensive study carried out by Stensaas and Stensaas (1978), twenty-seven different potential neural electrode materials were tested by implanting them into the cortices of rabbits. After 30 days, the histological examination revealed no adverse reaction in response to gold, platinum and tungsten. Unfortunately, noble metals usually tend to give rise to high electrode impedances and unstable potentials (McAdams 2006). Recently, different kinds of micro- and nanostructuring techniques have been investigated to increase the effective surface of noble metals and subsequently reduce the electrode-electrolyte interface impedance. Pt has been the most widely used material in implantable electrodes. Clinically, it has been utilized as an electrode material in various applications such as in cardiac pacemakers, cochlear implants and subdural strip and grid electrodes. Pt suffers from corrosion only slightly in both recording and stimulating use. Due to the mechanical softness of Pt, it is sometimes alloyed with iridium (Ir) to achieve electrodes for applications requiring high mechanical strength (McAdams 2006). Gold is also often used in recording electrodes, but it suffers from corrosion if used for stimulating purposes. On the contrary, Ir is mainly utilized as a stimulating electrode. The electrode properties of Ir can be further developed by electrochemical activation that leads to the adhesion of a porous and very stable oxide layer on the

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electrode surface. Activation provides much higher charge injection levels and also reduces the electrode contact impedance due to the porous microstructure (Cogan 2008, Blau et al. 1997). Titanium nitride (TiN) is another novel candidate suitable for use in stimulating electrodes, and it is already being utilized in cardiac pacemakers. The high charge delivery capacity of TiN is based on its high effective surface area due to the columnar structure of this material (Cogan 2008). Indium tin oxide (ITO) is a widely used electrode material in electrochemistry (Huang et al. 2009), but so far its use in neurophysiology is limited to in vitro MEAs for recording electrical signals from neurons. The biocompatibility of ITO is not completely understood although recent in vitro investigations on cellular responses (Bogner et al. 2006) and protein adsorption on ITO (Selvakumaran et al. 2008) have displayed promising results. The impedance of an electrode-electrolyte interface depends on the species of metal, the type of electrolyte used, the surface area of electrode, and the temperature (Geddes & Roeder 2001, 2003). Magnitude of impedance decreases with expanding surface area and increasing roughness. It also decreases with increasing frequency and increasing current density. Several different equivalent circuit models for the electrode-electrolyte have been suggested (reviewed by Geddes 1997). One of the most sophisticated models is presented in Fig. 7a. There are two identical electrodes in contact with electrolyte. Electrode polarization impedance consists of Warburg (polarization) resistance Rw and capacitance Cw. The Faradic resistance, Rf as a parallel with Warburg components, provides a route for direct current to pass through the interface. The resistance of the electrolyte solution is marked with Rs. Both Rw and the polarization reactance (Xw = 1/2πfCw) are frequency-dependent: they decrease as ( 1 / f ) with increasing frequency ( f ). At a sufficiently high frequency, the impedance between two electrodes in saline approaches asymptotically the value of Rs (Fig. 7b).

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B 1

Rw1 Cw1 Z

Rs Rw2

Cw2

2

Impedance |Z|(Ω)

A

Rf1+ Rf2 + Rs

Rs

Frequency (Hz) Rf2

Figure 7: (a) The equivalent circuit model for two identical bioelectrodes immersed in saline. (b) Typical electrode-electrolyte interface impedance curve measured between terminals 1 and 2 as a function of frequency. The curve approaches asymptotically the resistance value of the solution.

Due to the inverse relationship between electrode impedance and its surface area, it is advantageous to develop microelectrodes with high nanoscale surface topography since in this way it will be possible to achieve a high effective surface area without changing their geometrical dimensions (Kotov et al. 2009). Different techniques are used to increase the effective surface area including surface roughening, micro- and nanopatterning and chemical modification. In micro/nanopatterning techniques, different etching and machining techniques are utilized to form different micro/nanostructures, e.g. grooves and wells onto the surface of the substrate upon which the electrode material can then be deposited using traditional thin film techniques (Cheung 2002). Some examples of widely used surface modification techniques used to enhance nanoscale roughness includes Pt electroplating (Pt black) (de Haro et al. 2002) wet etching of Au and activation of Ir (Cogan 2008, Blau et al. 1997) to form a nanoporous iridium oxide layer as already discussed. 4.2.2 Substrate materials The surface area of substrate/encapsulation material is usually much larger compared to areas of the active recording sites. Therefore, biocompatibility aspects of substrate and insulator materials are very crucial in the development of neural

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electrodes and prosthesis which require long-term stability and functionality. These materials need to be biocompatible, biostable and possess good dielectric properties. An artificial material or implant can be considered to be biocompatible if (1) it does not evoke a toxic, allergenic or immunologic reaction, (2) it is not harmful to enzymes, cells and tissues, (3) it does not evoke thrombosis or tumors, and (4) it remains for a long term within the body without encapsulation or rejection (Heiduschka & Thanos 1998). On the other hand, biostability means that the implant material must not be susceptible to attack by biological fluids or any metabolic substances. Unfortunately, all implanted artificial materials in contact with neural tissue have an inconvenient tendency to induce significant glial scar tissue formation (Polikov et al. 2005). In response to implantation, glial cells such as astrocytes become activated and they transform into reactive phenotype which produces much larger size glial fibrillary acid protein filaments. Cell proliferation and migration capacity of astrocytes is also enhanced. The resultant scar tissue formation causes serious impairment of implant performance due to decreased local density of neurons and the formation of an encapsulation layer that increases electrode impedance and decreases the signal amplitudes. It has been noted that the astrocyte response to electrode implantation can be divided into two phases, the early response that occurs immediately after electrode implantation and the long-term chronic response. Within a few hours after implantation, the number of glial cells is significantly enhanced in the area surrounding the Si probes (Szarowski et al. 2003, Turner et al. 1999). The chronic response, starting approximately one week after implantation (Norton et al. 1992), consists of a compact glial scar tissue formation surrounding the electrode, which ultimately isolates the microelectrodes from neurons increasing the electrode contact impedance and causing signal amplitude deterioration (Edell et al. 1992, Nicolelis et al. 2003, Schmidt et al. 1993). Silicon with its common insulators, SiO2 and silicon nitride (Si3N4), have been the most widely used materials in neural interfaces as well as in the bio-MEMS technology overall. In

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spite of numerous desirable features of Si, there are at least two major weaknesses related to its use in neural applications. Firstly, there is a huge mismatch between the mechanical properties of the brain tissue and Si. The values of Young’s modulus for Si and brain tissue are 170 GPa and 3 kPa, respectively. The rigidity of Si may cause many adverse effects such as tissue damage, inflammation reactions and scar formation (Polikov et al. 2005). Secondly, the results from studies related to cytocompatibility of Si are somewhat contradictory (Kotzar et al. 2002). Some studies have shown that Si is a rather poor substrate for cell adhesion and even slightly cytotoxic (Liu et al. 2007, Madou & Tierney 1993), whereas some others have claimed that Si is almost as cytocompatible and immunogenic as tissue culture polystyrene (Ainslie et al. 2008). This uncertainty could limit the much-anticipated integration of Si-based neural implants into the human body. There is a growing interest in developing polymer-based implant materials that could be flexible enough to mimic biological tissue and to reduce mechanical damage, but not evoking any adverse tissue reactions. A wide variety of polymer materials, such as polyimide, benzocyclobutene, silicone rubbers, parylene-C and epoxy resins have been investigated as substrate and encapsulating materials in implantable neural interfaces (Cheung 2007, HajjHassan et al. 2008, Polikov et al. 2005). Polyimide possesses numerous desirable properties when used as a neural implant material. It has appropriate mechanical properties (Rubehn & Stieglitz 2010) and its biocompatibility has been demonstrated in many studies in vitro and in vivo (Stieglitz et al. 2000, Richardson et al. 1993, Seo et al. 2004, Lee et al. 2004, Yuen et al. 1987, Yuen & Agnew 1995). Stieglitz et al. (2000) evaluated the cytotoxicity of three commercial PI grades (Pyralin PI 2611, PI 2556, PI 2566; HD Microsystems GmbH, Bad Homburg, Germany) and reported excellent biocompatibility for PI 2611 and PI 2556 and good results for PI 2566. This last PI, since it is fluorinated, differs from the others with respect to its chemical structure (Kanno et al. 2002). In another study (Lee et al. 2004), it has been shown that fibroblasts attach, spread out

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and grow on the PI surfaces in a manner corresponding to the behaviour of control cells growing on the surface of polystyrene, i.e. the plastic that is used in cell culture well plates and flasks. Furthermore, Yuen et al. (1987) demonstrated that subdural implantations of two common insulators, parylene-C and PI (DuPont PI-2555), into the cerebral cortex of the cat evoked only minimal tissue reactions after 16 weeks. Tissue responses to PI and parylene-C were comparable to pure Pt (used as a negative control) whereas Ag-AgCl (positive control) triggered chronic inflammatory reactions (Yuen et al. 1987). In contrast, PI-Pt electrodes implanted into a rat sciatic nerve for 3 months have been shown to induce only mild scar tissue formation and local inflammation reactions, with these reactions being limited to a small area around the electrode (Lago et al. 2007). One shortcoming of PI is related to its permeability to environmental moisture and ions that could be crucial in terms of short circuiting and reducing the life-time of an implantable device. There are two common types of PIs on the market: nonphotosensitive PI and photosensitive polyimide (PSPI) (Ghosh & Mittal 1996). A special feature of PSPI, i.e. photopatternability, simplifies the fabrication process of the microdevice compared to conventional PI by eliminating the need for complex multilevel processes, i.e. oxygen RIE through a lithographically patterned metal aluminum (Al) mask (Wilson et al. 2007). PSPI is advantageous in electrode fabrication in which electrode sites and contact pads can be easily opened though an encapsulation layer while still insulating the transmission lines as well as in precision feature defining of the device’s outer shape. Although mechanical and electrical properties of PSPI are comparable to conventional PI, there will be a degree of uncertainty related to its biocompatibility due to the different chemical composition (e.g. existence of photoinitiators), imidization kinetics and solvent processes. So far very little data have been published addressing the biocompatibility of PSPI (Sun et al. 2009).

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5 Biocompatibility The main requirement for the successful clinical and experimental application of implantable devices constructed from artificial materials is that the surrounding tissues accept the implant, i.e. that the implant is biocompatible. Biocompatibility can be defined as ‚the ability of a material to perform with an appropriate host response in a specific situation‛ (Williams 1999). It is worth noting that the host response to specific individual materials could vary from one application site to another. An increasing number of applications require that the material should actively react in a desired manner with the tissue, or even degrade over time in the body, rather than remain constantly as an inert material. Biomaterials can be classified according to the chemical composition to metals, polymers, ceramics, composites and materials of biological origin (Navarro et al. 2008). Another classification is based on historical stages of biomaterial development. The first generation of implantable devices, developed during the years between 1940 and 1980, were constructed using materials that were chemically least reactive. It was thought that the best performance of implants could be achieved with a material that would be as passive as possible regarding adsorption of host proteins and cellular adhesion. Examples are inert metal materials that are very resistant to corrosion such as stainless steel, cobalt-chromium alloys, titanium alloys and the platinum group metals. With respect to the polymers, polytetrafluoroethylene, polymethylmethacrylate, polyethylene (PE) and silicones were widely used due to their high resistance to degradation (Williams 2008). The second generation biomaterials were specifically designed to be ‚bioactive‛, i.e. they could elicit a controlled host reaction and reaction in their physiological environment to achieve the desired result of the implantation. For example, certain

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compositions of bioactive glasses promote the desired cellular responses, like cellular adhesion, proliferation and differentiation into specific cell types, e.g. bone cells that will form new bone tissue and, thus, integrate the implant strongly into the natural peri-implant tissues. Recently, the third generation biomaterials have been and are being actively developed especially for in situ regeneration of damaged tissue and different tissue engineering applications. In these materials, living cells are used in combination with artificial materials (Navarro et al. 2008, Hench & Polak 2002, Ratner & Bryant 2004). There is a wide variety of chemical, biochemical, physiological, physical and other mechanisms which are involved in the contact between biomaterials and various tissues and thus will affect the biocompatibility (Williams 2008, Vadgama 2005). The concept of biocompatibility is not limited to bulk characteristics of the material used, but also includes physical and chemical surface properties as well as various decomposition products. In some cases, chemical composition of implant surface is of major importance, but in other cases some physical features such as stiffness and shape of the implant are the most important determinants of biocompatibility. The major material variables that may influence on the host response are summarized in Table 3. Table 3: Major material characteristics that affect the host response, modified from Williams (2008). Material variable The composition and structure of bulk material Crystallinity and crystallography Elastic constants Porosity Chemical composition of surface Surface roughness, texture Surface charge and electrical properties Surface wettability (hydrophilicity, hydrophobicity) Corrosion parameters, ion release profile and ion toxicity Degradation profile and toxicity of degradation products

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The interaction between an artificial material and a living system can be considered to consist of the local and systemic response of host tissue to the material as well as the response of the material to the host tissue (Williams 2008). After implantation of a biomaterial into the body, the host response involves a series of complex protein and cellular events. According to the ‚race for the surface‛ concept (Gristina 1987), there is in many occasions a critical competition between the microbial and host tissue cells for the implant surface. If this race is won by tissue cells, the implant surface is covered by tissue which provides resistance against bacterial colonization, and promotes tissue integration of the implant. On the other hand, if the race is won by microbes, the implant surface will be covered by bacteria and/or bacterial biofilm. In this case, the host cellular functions are hindered, resulting in adverse tissue reactions, and in the worst scenario, the implant has to be removed to protect the host. When a biomaterial is implanted into the living tissue or placed in cell culture, water molecules bind within nanoseconds to its surface. The extent and pattern of the interaction between the water molecules with the surface of implant depends on the surface properties of the material, such as surface free energy (SFE) (Roach et al. 2007, Kasemo 1998). Subsequently, within seconds to hours after implantation, the implant surface becomes covered by a monolayer of adsorbed serum or interstitial proteins. Since there are many hundreds of different proteins in blood and interstitial tissue fluid, there is a race for the surface between these molecules. Typically, the more abundant small proteins lead this race due to their rapid transport and attachment to the surface. Later, larger proteins with higher affinity towards the surface replace these small proteins (Roach et al. 2007). The protein layer formation process involves a very rapid initial phase, followed by a slower maturation phase as the protein coating of the implant approaches a steady-state phase. In the third stage, that occurs from as early as minutes after implantation, lasting up to several days, cells eventually adhere, spread out and differentiate on

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the protein molecules coating the biomaterial surface. These cellular processes are influenced by biological molecules such as extracellular matrix proteins, cell membrane proteins and cytoskeleton proteins as well as by the physicochemical properties of the surface such as its chemistry, topography, surface energy and electric charge (Puleo & Nanci 1999, Roach et al. 2007). Finally, the desired integration of implant will occur with the surrounding tissue. In biomedical implants, the final stage which can last from a few days (rapidly biodegradable devices) up to several decades (total hip replacement), can be seen as the continuum of the early implant stages. During this last stage, also adverse responses, such as clots, fibrous capsule formation, wear, corrosion, late hematogenic infections and device failure may occur.

5.1 BIOCOMPATIBILITY TESTING

Rigorous biocompatibility testing, both in vitro and in vivo, is needed when novel materials that are aimed for use in implantable medical devices, are developed. The International Organization for Standardization (ISO) has published the standard series ISO 10993 for biological evaluation of medical devices. This series of standards is currently composed of 18 parts. In vitro cytotoxicity tests (ISO 10993-5) are usually performed as the first stage acute toxicity and cytocompatibility testing for materials planned for use in medical devices. These tests are rapid to perform, sensitive, well standardized and cheap. Unnecessary use of animal models can be also avoided. In vitro tests are convenient in studies in which the adhesion, proliferation, differentiation and contact guidance of cells as well as the mineralization of matrix on biomaterial surface are investigated. These standardized and easily quantifiable tests (Pearce et al. 2007, Pizzoferrato et al. 1994) are also useful in studies aimed to clarify the individual effects of different surface

70


parameters such as SFE, surface charge and topography on the cellular responses. Various cell lines can be used for cytotoxicity testing; the cell line should be chosen to represent the cells that will be in contact with the biomaterial in its predicted in vivo use. Tests can be performed either via direct contact with a material and/or with extracts of the material. The direct contact test simulates the near-surface effects of the implanted biomaterial on tissues, whereas the extracts mimic the effect of the particles that are released from the surface. An incubation time ranging from 24 hours to 96 hours is recommended.

5.2 CELL-BIOMATERIAL INTERACTIONS

The clinical success of implants is largely based on adaptive and integrative interactions between the implant material and tissue cells. Although appropriate mechanical properties, durability and functionality of modern biomaterials and related devices are primarily derived from the bulk properties of materials, the biological responses to materials are affected by their physical and chemical surface properties. Thus, the interactions between cells and implants are regulated to a large extent by the surface properties of the implant material, although properties like elasticity can also affect host responses. The tendency of cells to adhere or not to adhere to the surface of implant is crucial for their performance, for example in total joint replacement prostheses or urinary stents (Tomlins et al. 2005). In order to improve the biological performance of implants, the effects of modified surface properties on cellular functions have been widely studied (Ratner & Bryant 2004, Williams 2008). Surface chemistry, roughness, texture, wettability and surface charge are among the characteristics that have an effect on the cellular response to a biomaterial. 5.2.1 Effect of surface topography on cellular responses Surface roughness is a measure of the random topography of a surface and it is quantified by the vertical variations of a real

71


surface from its ideal form. Amplitude parameters that are used in quantification of vertical deviations of the roughness profile from the midline can be grouped into two subclasses: averaging parameters and peak-to-valley parameters. The arithmetic mean deviation of the absolute value of the profile from a midline, Ra, is the most widely used parameter to quantify surface roughness. Ra is determined as: lr

1 Ra   z x  dx. lr 0

(4)

In this equation, z(x) is the absolute profile value from a midline and lr is the sampling length over which the surface profile has been measured (Tomlins et al. 2005). However, the Ra values are not always useful parameters for assessing biomaterial surfaces, because of the lack of detailed information about the geometry of the surface or the variations in peak heights or valley depths. Different peak-valley parameters have been developed to solve this problem including statistical distribution parameters called skewness and kurtosis. The simplest approach is to determine the highest and lowest points within the overall measuring length, and then the Rpv (peak-to-valley roughness) value can be estimated as the vertical distance between the farthest points. Both contact and non-contact techniques are available for measuring surface roughness. Contact methods are usually based on the use of stylus-based instruments in which a measurement needle is traversed over a test surface and the movement of the stylus arm can be detected e.g. using a piezoelectric crystal. However, these devices are not suitable for assessing the surface texture of soft materials, as they tend to damage the surface to be analyzed. Non-contact optical techniques such as laser profilometry and white light interferometry are especially useful for assessing the surface roughness of sensitive soft materials. Furthermore, different scanning probe microscopy techniques such as atomic force microscopy (AFM) and scanning tunneling microscopy can be

72


used for imaging surfaces with resolution at the atomic level. (Tomlins et al. 2005, Giessibl 2003, von Recum et al. 1996) The effect of surface roughness on cell functions has been studied in biomaterial surfaces which can be modified by several techniques such as sandblasting (Lampin et al. 1997, Iwaya et al. 2008, Orsini et al. 2000, Anselme & Bigerelle 2005), plasma treatment (Chu et al. 2002, Nebe et al. 2007, Saldana et al. 2006), mechanical polishing (Ponsonnet et al. 2003), acid etching (Iwaya et al. 2008, Orsini et al. 2000, Anselme & Bigerelle 2005) or laser treatment (Teixeira et al. 2007). It has been reported that surface roughness characteristics have an effect on the morphology, adhesion, proliferation and differentiation of osteoblast-like cells (Puckett et al. 2008, Kim et al. 2005, Khang et al. 2008, Das et al. 2009) as well as on fibroblasts (Könönen et al. 1992, Kunzler et al. 2007, Ponsonnet et al. 2003, Wirth et al. 2005). It is generally recognized that fibroblasts (connective tissue cells) prefer smooth surfaces whereas osteoblastic cells favor rougher surfaces (Kunzler et al. 2007, Ponsonnet et al. 2003, Wirth et al. 2005). Furthermore, enhanced differentiation and bone-specific gene expression of adhered human mesenchymal stem cells (hMSCs) have been demonstrated on nanorough surfaces (Ra < 50 nm) compared to rougher, machined surfaces (Ra ≈ 100 - 400 nm) (Mendonca et al. 2009, 2010, Dalby et al. 2006). However, the effects of surface roughness are also dependent on the substrate material being used (Hallab et al. 2001, Jinno et al. 1998). Moreover, it is important to note that a change in the surface roughness is also related to changes in the wettability properties of the surface. This is partly due to the fact that machining techniques commonly used to roughen or polish the surfaces often also introduce changes in the surface charges and tensions. Therefore, it may be difficult to conclude which mechanism is involved in modulating the cell behaviour at these instances (Khang et al. 2008, Das et al. 2009, Ponsonnet et al. 2003). Surface texture refers to a surface topography in which patterns of features are placed deliberately on the surface. Several microfabrication techniques, mainly originating from

73


the production of integrated circuits, have been recently utilized in the biomaterial research to study cellular reactions to a substrate having a three-dimensional structured surface topography in the micro- or nanoscale. The major part of published literature has focused on micro- and nanostructures produced through photolithography or electron beam lithography in conjunction with reactive ion etching processes on the surfaces of Si, quartz or Ti (Curtis & Wilkinsson 1997, Flemming et al. 1999, Martinez et al. 2009). More recently, polymer microfabrication techniques, such as hot embossing (Charest et al. 2004), injection molding (Myllymaa et al. 2009b) and soft lithography (Mata et al. 2002a,b) have been increasingly used. Different pattern feature types such as grooves/ridges, wells, pits and pillars with a wide variety of pattern dimensions have been examined (Martinez et al. 2009). Substrate micro- and nanotopography, independently of surface biochemistry, seems to have significant effects on cell orientation and adhesion, morphology, proliferation and differentiation. It has been demonstrated that cells seeded onto microgrooved surface aligned their shape and elongated in the direction of the grooves (Brunette et al. 1983, Chou et al. 1995, 1998, Britland et al. 1996, Curtis & Wilkinson 1997, Teixeira et al. 2003, Dalby et al. 2004, Ber et al. 2005, Mwenifumbo et al. 2007, Loesberg et al. 2007, den Braber et al. 1996, Clark et al. 1987, Charest et al. 2004, Hamilton & Brunette 2005; Hamilton et al. 2006). The cell membrane conformed to the widest grooves, but tended to bridge the narrowest grooves (Matsuzaka et al. 2003). Other geometrical features such as wells, pits and pillars with dimensions of less than 5 µm seem to lead to smaller, rounded cells with less organized cytoskeletons (Martinez et al. 2009, Andersson et al. 2003, Gallagher et al. 2002). The results from investigations related to the effects of nanoscale topographical structures on cell proliferation and differentiation are somewhat contradictory and highly dependent on the cell line and structures being investigated (Puckett et al. 2008, Andersson et al. 2003, Martinez et al. 2009).

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5.2.2 Effect of wettability properties on cell-biomaterial interactions Surface energy of materials is of special interest in the biomedical applications due to its obvious effect on cellbiomaterial interactions. Surface wetting properties have a significant effect on various biological events at the sub-cellular and cellular level including protein adsorption, cell attachment and spreading. Contact angle measurements are widely used in the assessment of wettability properties. In this technique, a small droplet (e.g. 15 µl) is placed on the surface and the angle between the drop (liquid-vapour interface) and the substrate (solid) is determined. Relatively more water wettable surfaces are called hydrophilic (high SFE), while less wettable surfaces are hydrophobic (low SFE) surfaces (Ratner 1996, Hasirci & Hasirci 2005). Although these are relative terms, it has been suggested that water contact angles smaller than 65° indicate hydrophilic surface, while those larger than 65° indicate hydrophobic surface (Vogler 1998). Different approaches, such as the Owens-Wendt (Owens & Wendt 1969) and van Oss (van Oss et al. 1988) models have been developed to estimate SFE of materials and their components on the basis of measured contact angles with polar and non-polar liquids whose surface tension values are known. Using the Owens-Wendt

model,

the

P polar  S and

dispersive  SD

components can be estimated:

(1  cos ) L  2(( SD LD )1/ 2  ( SP LP )1/ 2 ).

(5)

Here θ is the measured contact angle value and  L ,  LD and  LP stand for the liquid’s total SFE and its dispersive and polar components, respectively. The total SFE (  S ) is the sum of its dispersive and polar components:

 S   SD   SP .

(6)

75


The polar component of the SFE is related to hydrogen bonding, dipole-dipole and other site-specific interactions, whereas the dispersive component refers to the van der Waals and other non-site specific interactions between the solid surface and liquid applied on it. Increased wettability with a higher SFE has been shown generally to enhance the adhesion, and spreading of various cell types such as fibroblasts (Wei et al. 2007, Altankov et al. 1996, van Kooten et al. 1992), osteoblasts (Khang et al. 2008, Feng et al. 2003, Lim et al. 2008, Myllymaa et al. 2009b) and MSCs (Sawase et al. 2008). Moreover, through the SFE modifications, it is possible to influence the differentiation of MSCs into the desired phenotypes (Curran et al. 2005, 2006, Marletta et al. 2007, Calzado-Martin et al. 2010). It has been presented that the SFE is an even more important surface characteristic than surface roughness for cellular adhesion strength and proliferation and might be a useful parameter for modifying cell adhesion and cell colonization onto engineered tissue scaffolds (Hallab et al. 2001). The wettability of a solid material can be modified either by changing the surface roughness (Wang et al. 2006, Hao & Lawrence 2007) or the surface chemistry (Lim et al. 2008, Michiardi et al. 2007, Kennedy et al. 2006, Myllymaa et al. 2009b). In hydrophobic materials, surface roughening increases the hydrophobic nature of material, whereas in hydrophilic materials surface roughening enhances hydrophilicity. Surface chemistry modifications are usually performed by using different plasma treatments (Lim et al. 2008, Michiardi et al. 2007), proper surface coatings (Khang et al. 2008, Myllymaa et al. 2009b), SAMs (Kennedy et al. 2006, Folch & Toner 2000) or incorporating dopants into the bulk material (Myllymaa 2009a). 5.2.3 Effect of charge distribution on cell-biomaterial interactions The electric charge distribution in the vicinity of a biomaterial surface can be considered to be one of the main factors affecting the biological response of a biomaterial (Khorasani et al. 2006, Cai et al. 2006, Krajewski et al. 1998, Altankov et al. 2003). This

76


charge depends on several factors, such as the chemical composition of the material surface, the composition of the surrounding fluid and the environmental pH value (Krajewski et al. 1998). The charge and its distribution are closely related to protein adsorption, the strength of cellular adhesion and the shape of the cells growing on the surface of the biomaterial (Cheng et al. 2005, Bodhak et al. 2009, Thian et al. 2010). Therefore, the investigations of the relationship between electrical charge on a biomaterial surface and protein adsorption as well as subsequent cellular interactions are particularly useful ways to improve understanding of the mechanisms involved in the biological integration of materials with tissues. The interface between a solid surface and the surrounding electrolyte establishes a charge distribution which is different from the bulk phases of solid and liquid. In this electrochemical double layer model, at the vicinity of a solid-liquid interface, the charge carriers are fixed (a stationary layer), whereas they are mobile in the liquid phase (a mobile layer) at a greater distance The zeta potential (ζ) is termed as the potential decay between the solid surface and the bulk liquid phase at the shear plane (Lyklema 1995) (Fig. 8). Streaming current/streaming potential measurement is a convenient method for examining the existing interface charges of a solid biomaterial in an electrolyte solution (Roessler et al. 2002). In this method, an electrolyte is circulated through the measuring cell and forced under pressure to flow through a small gap formed by two sample surfaces. This fluid flow evokes a relative motion between the stationary and mobile layer and this leads to a charge separation which provides experimental access to the zeta potential. Higher positive values of the zeta potential at a fixed pH reflect a positive charge of the surface which would attract negatively charged entities such as anions or charged proteins while lower values of zeta potential at a fixed pH are evidence of a negative charge of the surface which tends to attract positively charged cations and particles (Cai et al. 2006). The relationship between the magnitude of zeta potential and cellular responses is not well understood (Altankov et al. 2003,

77


Hamdan et al. 2006), although in recent studies it has been demonstrated that negatively charged surfaces can enhance osteoblastic cell – biomaterial interactions (Bodhak et al. 2010, Thian et al. 2010). Instead, the effect of zeta potential on protein adsorption has been more extensively studied. The lower the absolute value of the zeta potential of the surface, the lower is the adsorption of serum proteins (Krajewski et al. 1996, 1998, ElGhannam et al. 2001). stationary layer

Electrokinetic potential

solid

-

mobile layer

+

+

-

+ + +

ζ

-

-

+

-

+ +

+

-

+

+

-

Distance Figure 8: The interface between a solid surface and a surrounding solution introduces an electrochemical double layer. The charge distribution is divided into a stationary and a mobile layer. These layers are separated by a plane of shear. The zeta potential (ζ ) is assigned to the potential decay at this shear plane.

5.2.4 Effect of surface micropatterning on cellular responses Surface micropatterning has been extensively employed for controlling the adhesion, proliferation, differentiation and contact guidance of cells (Folch & Toner 2000, Falconnet et al. 2006, Ito 1999). Photolithography and soft lithography are powerful techniques to produce surface patterns which differ from the background e.g. in terms of surface chemistry (Kane et al. 1999, Winkelmann et al. 2003, Scotchford et al. 2003, Levon et al. 2009, Myllymaa et al. 2009a), surface charge (Soekarno et al. 1993), SFE (Matsuda & Sugawara 1995), or the attachment of biological substances that can influence cellular function (McBeath et al. 2004, Chen et al. 1997, Blawas & Reichert 1998).

78

6 Aims of the present study In this study, micro- and nanofabrication methods have been utilized in the development of novel neural interfaces as well as intelligent implant materials capable of achieving the desired cellular responses on a biomaterial surface. It was hypothesized that these approaches would enhance the tissue response and functionality of the implants used in contact with soft (neural) or hard (bone) tissue. The main aims of this thesis were: 1. to design and fabricate novel thin film microelectrode arrays for stable and high-resolution intracranial recordings, 2. to study the suitability of flexible polymers (e.g. polyimide) as construction materials in bio-MEMS applications to replace materials based on silicon technology, 3. to study the opportunities to improve the electrochemical properties (stability, contact impedance) of bioelectrodes by using modern micro- and nanostructuring methods, 4. to investigate the cytocompatibility of different dielectric thin films produced by USPLD method and to clarify their suitability for neural interfaces, 5. to test whether the surface micropatterning could be used to enhance the cytocompatibility of cell unfriendly siliconbased implants and improve their biocompatibility during the initial phase of integration with tissue, and 6. to investigate the effect of surface micropatterning, produced using photolithographic and PVD techniques, on the behaviour (e.g. adhesion, spreading, and differentiation) of osteoblast-like cells and human mesenchymal stem cells.

79

6 – Aims of the present study

80

7 Materials and methods The present thesis consists of six studies, which are referred to by the Roman numerals (I-VI). In addition, some unpublished data are included. A summary of the materials and methods used in this thesis is presented in Table 4. Table 4: Overview of the materials and methods used in the present thesis. Study

Device/ sample

Characterization Cell line/ Testing methods Animal

I

Microelectrode array

Optical microscopy, EIS

Wistar rat

Acute cortical evoked potential recordings

II

Microelectrode array

Optical microscopy, SEM, AFM, EIS

Wistar rat

Acute and chronic cortical evoked potential recordings

III

PI films on Si,

AFM, Contact angle, SFE, Zeta potential

BHK-21

Qualitative SEM analysis of cultured cells, MTS assay, Fluorescence microscopy (live/dead)

Al203 on Si (*) C3N4 on Si (*) IV

Review article

V

Micropatterned Optical Ti, Ta, Cr and DLC microscopy, AFM, coatings on Si Contact angle, SFE, Zeta potential (*)

VI

Micropatterned Ti and DLC coatings on Si

(#)

Micropatterned Cr and Ta coatings on Si (*) Unpubl. data

nanorough USPLD Pt surfaces

hMSC

Quantitative and qualitative SEM analysis of cultured cells, Fluorescence microscopy of stained cells (focal adhesion, ALP, OC, MS)

SaOS-2

Quantitative and qualitative SEM analysis of cultured cells, Confocal laser scanning microscopy of stained cells (focal adhesion)

EIS, AFM

(*) Additional materials/methods: not included in original publications III, V or VI; (#) Sample surfaces were characterized in study V Abbreviations: EIS: electrical impedance spectroscopy; SEM: scanning electron microscopy; AFM: atomic force microscopy; PI: polyimide; Si: silicon; Al203: alumina; C3N4: carbon nitride; SFE: surface free energy; BHK-21: baby hamster kidney fibroblast; Ti: titanium; Ta: tantalum; Cr: chromium; DLC: diamond-like carbon; hMSC: human mesenchymal stem cell; ALP: alkaline phosphatase; OC: osteocalcin; MS: mineral staining; SaOS-2: sarcoma osteogenic cell; USPLD: ultra-short pulsed laser deposition; Pt: platinum

81

7 – Materials and methods 7.1 MICRO- AND NANOFABRICATION OF ELECTRODES AND OTHER SAMPLES

A wide variety of micro- and nanofabrication techniques as well as materials were used in this thesis work. A description of these fabrication processes is given in chapters of 7.1.1 – 7.1.5. 7.1.1 Preparation of photomasks In the studies I, II, V and VI, photolithographic methods were used in patterning thin deposition layers on PI or Si substrates. The photomasks used in these studies were designed by CleWin layout editor, version 4.0.2 (WieWeb software, Hengelo, The Netherlands) and fabricated using the laser scanning technique by Mikcell Oy (Ii, Finland) on 4-inches or 5-inches soda-lime glass plates with a structured Cr layer. The photomask used in studies I and II consists of two mask patterns both repeated five times enabling the ‚batch fabrication‛ at a prototyping scale. The first set of patterns was used in structuring the metallization layer and the second set in patterning the PI insulating layer to expose electrode sites. The photomasks used in studies V and VI consist of micropatterned areas ( 4 mm x 4 mm or 8 mm x 8 mm) containing regularly spaced microsquares (sides 5, 25, 75 or 125 µm) or microcircles (diameters 5, 25 or 125 µm). 7.1.2 Flexible polyimide-based microelectrode arrays A fabrication process for flexible MEAs that follows photolithographic and thin film procedures was developed and iterated to a fully functional solution during a two year trial period such that appropriate sensor layouts and materials with their thicknesses for base, metallization and insulating layers were optimized. The most relevant variations are presented in Table 5. The developed arrays consisted of either 8 or 16 roundshaped microelectrodes with diameters of 200 µm or 100 µm, respectively, in an area of about 2 mm x 2 mm at the end of a PI ribbon. The layouts of these electrodes are presented in Fig. 9.

82

7 – Materials and methods Table 5: Developed microelectrode array prototypes. Type

Channels Substrate material and thickness

Electrode material, its diam.

Insulation layer, its thickness

Reference

A

8

Polyimide PI-2525a: 3 layers, total 30 µm

Ti/Pt, 200 µm

PI-2771a, 3µm

Study I

B

8

Kapton HN filmb 25 µm + PI-2525 5 µm

Ti/Pt, 200 µm

PI-2771, 3µm

Study II

C

8

Kapton HN film 25 µm + PI-2525 5 µm

Ti/Au, 200 µm

PI-2771, 3µm

Myllymaa et al. 2008

D

16


Ti/Pt, 100 µm

PI-2771, 3µm


E

16


Ti/Pt, 100 µm

SU-8c, 3 µm


Abbreviations: Ti: titanium; Pt: platinum; Au: gold a HD Microsystems GmbH, Bad Homburg, Germany; c Microchem Corp., Newton, MA, USA

A

b

Goodfellow Ltd., Cambridge, UK;

Electrode sites

B

C

32 mm

1 mm

Connector pads

Figure 9: (a) The layout of the 16-channel microelectrode array. The length of arrays is designed to be about 32 mm. There are connector pads at one end of the device allowing connection of the array to miniature connector and further to recording instrumentation. The magnified schema of the 16-channel (b) and 8-channel (c) array viewed at the recording ends. The sizes of circular-shaped microelectrodes are 100 µm (b) and 200 µm (c).

83

7 – Materials and methods

Similar processes were used in studies I and II except for the construction of a base layer and the curing parameters for insulating PI layer. In study I, a base layer was constructed from three spin-coated PI (PI-2525, HD Microsystems) layers, i.e. the thickness of the base was about 30 µm. In study II, a 25 µm thick Kapton HN film (Goodfellow Ltd., Cambridge, UK) was utilized together with a spin-coated single PI-2525 layer, resulting in a base with total thickness of 30 µm. The process flow is demonstrated in Fig. 10. A

E

UV

UV

B F

C

electrode opening

connector pad

G D

glass slide

polyimide, d ~ 30 µm

Ti/Pt thin film, d ~ 200 nm

photoresist

photosensitive polyimide, d ~ 3 µm

Figure 10: The fabrication process flow of polyimide-based microelectrode arrays. The process is based on photolithographic and magnetron sputter deposition techniques. Electrode sites, transmission lines and connector pads are sandwiched between two layers of polyimide. See details in the text.

First, one of the above mentioned practices was used to form a 30 µm thick PI layer onto 2‛ x 3‛ microscope glass slides (Logitech Ltd., Glasgow, Scotland, UK) which were used to ensure a rigid support throughout the process (Fig. 10a). The PI layer was cured in an oven according to the manufacturer’s recommendations. After cooling, 20 % hexamethyldisilazane

84


(HMDS) (Riedel-de-Haen AG, Seelze, Germany) in xylene was spun upon the PI base layer to enhance adhesion between the PI and the photoresist. A negative photoresist (ma-N 1420, Micro Resist Technology GmbH, Berlin, Germany) was spun (5000 rpm, 30 s), pre-baked on a hot plate (100°C, 120 s), exposed with i-line (365 nm) UV light (Karl Suss MA45, Suss Microtec Inc., Waterbury Center, VT, USA) for 30 s, developed in a ma-D 553s developer (Micro Resist Technology), rinsed with deionized water and post-baked in an oven (100°C, 30 min) (Fig. 10b). Thin films of Ti and Pt were deposited at thicknesses of 20 nm and 200 nm, respectively, by using magnetron sputtering (Stiletto Serie ST20, AJA International Inc., North Scituate, MA, USA) and high purity (99.9 % or better) Ti and Pt targets (Goodfellow Metals Ltd., Huntingdon, UK) (Fig. 10c). Ti serves as an adhesion promoter between the PI and the Pt. The metallization layer was patterned via a lift-off procedure in resist remover, mr-Rem 660 (Micro Resist Technology) forming electrodes, transmission lines and connector pads on the PI base (Fig. 10d). Photosensitive polyimide PI-2771 (HD Microsystems) was used as an insulation layer. PI-2771 was spun (5000 rpm, 60 s) and pre-baked on the hot plate (Fig. 10e). After cooling, it was exposed to UV light (12 s) and developed in cyclopentanone (Sigma-Aldrich Corp., St. Louis, MO, USA) and rinsed with propylene glycol methyl ether acetate (PGMEA, Sigma-Aldrich). The PI-2771 layer was cured in an oven in which the temperature was gradually raised either to 200 °C after which it was maintained for 30 minutes (study I) or to 300 °C after which it was maintained for 60 minutes (study II). PI-2771 formed about a 3 µm thick insulation layer that was present on all positions other than the areas of the electrode sites and connector pads (Fig. 10f). Finally, the arrays were detached from the glass slide (Fig. 10g) and cut into their final shapes using scissors and a knife. Since only the edge regions were tightly attached to glass slides (detailed descriptions of these protocols are given in original papers I and II), it was possible to achieve detachment easily with by immersing in deionized water.

85


Thin film connector pads were designed to fit into a 16channel 0.5 mm pitch zero-insertion-force (ZIF) connector (JST Ltd., Halesworth, UK). This ZIF connector was soldered with lead-free solder to a thin printed circuit board (PCB) adapter (Kytkentälevy Oy, Helsinki, Finland) containing also a 2 x 8 channel surface mount microsocket (CLM-serie, Samtec Inc., New Albany, IN, USA) through which the array was connected to the recording instrumentation. Casco strong epoxy-resin (Akzo Nobel Coating Oy, Vantaa, Finland) was used as an encapsulant at the interface between the array and the ZIF connector and the signal tracks in the PCB in order to avoid short-circuiting when implanted onto the surface of rat cortex. 7.1.3 Micropatterned biomaterial surfaces on silicon In studies V and VI, micropatterned titanium (Ti), tantalum (Ta), chromium (Cr) and diamond-like carbon (DLC) coatings were fabricated using photolithography and PVD methods. Si wafers were used as substrates. Prior to application of the photoresist, the Si wafers were dry-baked and coated with adhesion promoter (20 % HDMS in xylene). Then, the chosen photoresist, i.e. ma-N 1420 (Micro Resist Technology) in study V or SU-8 (MicroChem) in study VI was applied, patterned and cured using the optimized process parameters described in the original papers (V, VI). Two PVD methods were used to carry out thin film depositions. Magnetron sputtering was used to deposit Ta, Ti and Cr coatings and a filtered pulsed plasma arc discharge method was used to produce DLC coatings onto the surface of the patterned wafers. The process parameters were optimized (details are described in the original articles V and VI) to achieve well-adhering, smooth coatings with a thickness of about 200 nm. The micropatterns were revealed via a lift-off procedure by immersing the wafers in a resist remover (acetone or mr-Rem 660) ultrasonic bath. The biomaterial coating deposited on the top of the photoresist was removed together with the dissolved resist, and the final microfeatures were formed. The Si wafers were cut into individual samples (10 mm x 10 mm) with a custom-made device using a diamond knife.

86


The size of each final micropatterned sample used in the cellular studies was 10 mm x 10 mm. In study VI, three different sample sets were utilized. In the first sample set, each chip contained four 4 mm x 4 mm sample areas. Three of these areas consisted of circles with diameters of 5, 25 or 125 µm or squares with the length of the sides being 5, 25 or 125 µm composed of either DLC or Ti on a Si background with the fourth sample area containing Ti or DLC as a homogeneous surface (Fig. 11a,b). Regardless of the size of the patterns, the circles and squares covered 30.6 % and 25 % of the total sample surface area, respectively. In the second set of samples, the pattern and background were reserved (‚inverse samples‛) so that the patterns were composed of uncoated Si squares or circles on a DLC or Ti background. In the third set, each sample contained a micropatterned area of 8 mm x 8 mm at the center of the sample. Micropatterning in these areas consisted of 75 µm squares with 100 µm spacing between adjacent squares in two orthogonal directions (Fig. 11c). Thus, the (Ti, Ta, Cr or DLC) patterns occupied 18.4 % of the total sample surface area. This last sample type was also used in study V.

Figure 11: Sample types used in studies V-VI to clarify cellular behaviour on micropatterned surfaces: (a) circular patterning, diameters of circles correspond to 125 µm, 25 µm and 5 µm plus plain reference area (bottom right), (b) rectangular patterning, sides of squares correspond to 125 µm, 25 µm and 5 µm plus plain reference area (bottom right) and (c) rectangular patterning, sides of squares are 75 µm.

87


7.1.4 Tailored surfaces produced by ultra-short pulsed laser deposition Ultra-short pulsed laser deposition (USPLD) technique was utilized in the creation of novel coatings for applications of bioMEMS. Firstly, ultrasmooth alumina (Al2O3) and carbon nitride (C3N4) dielectric films were deposited on Si wafers to clarify their cytocompatibility properties using similar fibroblast experiments as in study III (unpublished findings). Secondly, Pt depositions with varied nanotextured surface topography were carried out on SS, Si and Kapton substrates aiming to find out new solutions to improve the electrochemical properties of bioelectrodes. Neural electrodes need to be encapsulated in a material that provides excellent barrier properties and shows acceptable biocompatibility. In this context, Al2O3 and C3N4 were considered as interesting coating candidates. Biomaterials coated with Al2O3 have been used for a long time in medical applications, particularly in fields of dentistry and orthopedics (Szeitzer et al. 2006). The biocompatibility of Al2O3 has been previously demonstrated as a bulk material (Heimke et al. 1998, Takami et al. 1997) as well as an atomic layer deposited film (Finch et al. 2008). The dielectric properties of pulsed laser deposited Al2O3 are also superior (Katiyar et al. 2005) and USPLD deposited Al2O3 provides good insulating properties even in liquid environment (R. Lappalainen, personal communication 2009). Carbon nitride is a promising novel coating material for biomedical applications. It has superb chemical inertness and its tribological and dielectric properties are also favourable (Cui & Li 2000). The abbreviation ‚CN‛ is commonly used to describe all carbon nitride coatings with different C/N ratios. The bio- and hemocompatible nature of CN has been also demonstrated (Du et al. 1998, Cui & Li 2000, Cui et al. 2005). However, in neural applications, the testing of Al2O3 and CN films has been rare. The USPLD method utilizes intense laser pulse (typically in ps time scale) to erode a target and deposit the eroded material in the form of high speed ionized plasma onto a substrate. Just

88


prior to the depositions, the sample surfaces were gently cleaned using Ar+ ion sputtering (SAM-7KV, Minsk, Belarus). A new type of mode-locked fiber laser (Corelase Oy, Tampere, Finland) and the ColdabTM deposition technology developed by Picodeon Ltd. Oy (Helsinki, Finland) were used for USPLD. A high purity sintered target of C3N4 and Al2O3 was used for deposition. In both cases, optimized laser parameters (Amberla et al. 2006), such as pulse energy, pulse repetition rate and scanning length were used. Maximum average power was 20 W at 4 MHz, which results in 5 µJ pulse energy. The pulse length was 15-20 ps. The laser beam was focused into a 15-25 micron spot giving the maximum fluences about 4 J/cm2 at 5 µJ. Both Al2O3 and C3N4 films were deposited at thicknesses of about 200 nm on p-type Si wafers. The USPLD technique was also used to produce nanorough Pt depositions on SS, Si and Kapton substrates. Laser deposition parameters were modified to obtain thin (100-300 nm) Pt films with a different nanotextured surface topography (unpublished data). Due to the high energy of Pt plasma, good adhesion was achieved without using any special intermediate layers or special techniques. The main deposition parameter to control the surface topography was the Ar gas atmosphere with a pressure varied in the range 1-10-6 mbar. 7.1.5 Spin-coated polyimide films for cytotoxicity testing In order to evaluate the cytocompatibility of different PI grades, non-photosensitive PI-2525 (HD Microsystems) and photosensitive PI-2771 (HD Microsystems) were deposited on Si wafers by using a spin-coating technique. Prior to application of the PI precursor, 2-inches, p-type Si wafers (Si-Mat, Landsberg am Lech, Germany) were baked for 20 min at 200 °C and coated with VM-651 (HD Microsystems) to remove moisture and to enhance the adhesion, respectively. Then, the PI-2525 and PI-2771 precursors were spin-coated (5000 rpm, 60 s) under low acceleration onto Si wafers. Subsequently, the PI-2771 deposits were pre-baked (100 °C, 90 s), exposed to 365nm UV light (Karl Suss MA-45), developed in cyclopentanone (Sigma–

89


Aldrich) and rinsed with PGMEA (Sigma–Aldrich). Half of the PI-2771 coated wafers and all of the PI-2525 coated wafers were polymerized by curing them in an oven in which the temperature was gradually raised to 350 °C over a period of 60 min and then maintained for 60 min at that temperature. Another half of the PI-2771 coated wafers was cured at a lower temperature (200 °C). Therefore, three different PI groups were investigated: (1) non-photosensitive PI-2525, cured at 350°C; (2) photosensitive PI-2771, cured at 200 °C and (3) photosensitive PI-2771, cured at 350 °C.

7.2 MICROSCOPIC CHARACTERIZATION OF SURFACES

Different microscopic methods, including optical microscopy, scanning electron microscopy (SEM), atomic force microscopy (AFM) and contact angle measurements were used in this thesis in quality control monitoring and in the characterization of the produced samples and devices. Brief descriptions of these characterization techniques are given in the following sections of 7.2.1– 7.2.3. 7.2.1 Scanning electron microscopy Scanning electron microscopy was used for examining the quality of photolithographically patterned surfaces (studies V and VI), spin-coated PI films (study III) and the MEAs (study II). Prior to imaging with a Philips XL30 ESEM-TMP (Fei Company, Eindhoven, The Netherlands), the samples were usually coated with a thin layer of Au using a Sputter Coater E 5100 (Polaron Equipment Ltd., Hertfordshire, UK). 7.2.2 Atomic force microscopy Atomic force microscopy was used in this thesis (studies II, III, and V) to analyze the surface topography (roughness) of biomaterial samples and to determine the thickness of thin film depositions. The characterization was performed using a PSIA XE-100 (Park Systems Corp., Suwon, Korea) AFM system at

90


ambient temperature and humidity. Al-coated Si cantilevers (Acta-10, ST Instruments B.V., LE Groot-Ammers, The Netherlands) were used in a non-contact mode to probe the surface across an area of 2 x 2 µm. The scanning rate of 0.25 Hz was used. The thicknesses of thin film depositions as well as the average surface roughness (Ra) and peak-to-valley roughness (Rpv) values of sample surfaces were determined using the analysis software (XIA) of the AFM instrument. 7.2.3 Contact angle measurements and determination of surface free energies The contact angle measurements were performed to clarify wettability properties of material surfaces (studies III and V). Prior to the measurements, the samples were ultrasonicated in ethanol and deionized water and allowed to dry. Contact angles were measured using the sessile drop method at constant laboratory conditions (22 °C and in 45% relative humidity) with the aid of a custom made apparatus that consisted of a stereo microscope SZ-PT (Olympus Corp., Tokyo, Japan) equipped with a digital camera (Camedia C-3030-ZOOM, Olympus). In this method, a small droplet (15 µl) was pipetted onto the surface under study, and after 5 seconds, the droplet was photographed. Gnu Image Manipulation program (GIMP, www.gimp.org) was used to determine the contact angle from the shape geometry of the droplet. One polar liquid, i.e. water and one non-polar liquid, i.e. diiodomethane was used to access the total SFE as well as its polar and dispersive components. The Owens-Wendt model, described in chapter 5.2.2, was used to determine SFEs.

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7 – Materials and methods 7.3 ELECTROCHEMICAL CHARACTERIZATION OF SURFACES

The electrochemical properties of electrodes and other biomaterial surfaces were investigated by electrochemical impedance spectroscopy (EIS) and zeta potential measurements. Descriptions of methods used are given in the following sections. 7.3.1 Electrochemical impedance spectroscopy Electrochemical impedance spectroscopy was used (studies I and II) to characterize the electrochemical properties of the fabricated microelectrodes, i.e. to assess the recording capabilities of the electrode for making neural measurements. The measurements were performed with an LCR meter (3531 Z HiTester, HIOKI E.E. Corp., Nagano, Japan) (study I) or with a Solartron 1260 impedance gain/phase analyzer (Solartron Analytical, Farnborough, UK) (study II). In studies I and II, the MEA was immersed in physiological saline solution (0.9% NaCl) at room temperature and a small sinusoidal perturbation signal (100 mV without any DC offset) was applied between the individual microelectrode and the Pt counter electrode having a much larger surface area (Fig. 12a). The induced current and its phase were recorded. Frequency bands of 50 Hz - 1 MHz and 1 Hz-100 kHz were used in the studies I and II, respectively. Electrode-electrolyte interface impedance spectra of SS electrodes with and without a nanotextured Pt layer were also measured (unpublished data). The EIS measurements were implemented in a custom-made miniature measurement cell (Fig. 12b), in which a working electrode and a counter (Pt) electrode were clamped against the opposing walls of the cell. The holes with a diameter of 8 mm at the both walls exposed both the working and the counter electrodes with equal areas of 50 mm2. The measurements were carried out at frequencies between 1 Hz and 1 MHz by applying a sinusoidal voltage of 100 mV without any DC offset using the Solartron 1260 impedance analyzer.

92


A

microelectrode

Pt counter electrode

B

Pt counter electrode Sample under test saline solution

Figure 12: Set-ups used to measure the electrode-electrolyte interface impedances. Electrical impedance spectroscopy was carried out with a two-electrode configuration, in which a platinum sheet was used as a counter electrode. Working electrode, i.e. microelectrode (a) or nanotextured platinum coating on stainless steel (b) was immersed in saline (0.9 % NaCl) solution or exposed to contact with it through the hole (Ø: 8 mm) drilled at the wall of polytetrafluoroethylene-based pool, respectively.

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7.3.2 Zeta potential measurements In study III, the streaming current measurement technique was used to examine existing interface charges of the studied solid biomaterials in a liquid. The streaming current measurements were performed with an electrokinetic analyzer (SurPASS, Anton Paar GmbH, Graz, Austria) equipped with an adjustable gap cell. In each measurement, two identical test samples, i.e. pieces of Si wafer coated with biocompatible thin film were fixed on the sample holders with a cross section of 20 mm x 10 mm by using a double-sided adhesive tape. The sample holders were then inserted in the adjustable gap cell and a gap of approximately 100 µm was adjusted between the surfaces of the samples during the measurement. An electrolyte was circulated through the adjustable gap cell and forced under pressure to flow through a narrow gap. This evoked a comparative movement of charges in the electrochemical double layer which could be detected by two Ag-AgCl electrodes placed at both ends of the gap channels. The measurements were performed using 1 mM KCl as an electrolyte solution at a fixed pH of 7.4 ± 0.2. The pH of KCl was adjusted with 0.1 M HCl and 0.1 M NaOH. An electrolyte was circulated through the adjustable gap cell and forced under pressure to flow through a narrow gap. This evoked a comparative movement of charges in the electrochemical double layer which could be detected by two Ag-AgCl electrodes placed at both ends of the gap channels. The measurements were performed using 1 mM KCl as an electrolyte solution at a fixed pH of 7.4 ± 0.2. The pH of KCl was adjusted with 0.1 M HCl and 0.1 M NaOH. The zeta potential (ζ) was obtained from streaming current measurements according to the Helmholtz-Smoluchowski equation (Lyklema 1995):

 

dI  L   . dp    0 A

(7)

Here dI/dp is the slope of the streaming current versus the differential pressure,  is the electrolyte viscosity,  0 is the vacuum permittivity,  is the dielectric constant of the

94


electrolyte and L is the length of the streaming channel, and A is the cross-section of the streaming channel (the rectangular gap between the planar samples).

7.4 RECORDING OF EVOKED POTENTIALS IN RATS

The suitability of the MEAs for cortical mapping of evoked potentials was investigated in acute and chronic recordings in Wistar rats in the A.I.V. Institute of the University of Eastern Finland (studies I and II). All experiments were conducted in accordance with the Council of Europe guidelines and approved by the Institutional Animal Care and Use committee and the State Provincial Office of Eastern Finland. Prior to surgery, Wistar rats were anesthetized with urethane (1.2-1.5 g/kg) for acute recordings and with medetomidine/ ketamine (0.5 mg/kg and 75 mg/kg, respectively) for implanting the chronic electrodes, and placed in a stereotaxic apparatus. A microelectrode array was placed on the dura over the pre- and postparietal cortices between lambda and bregma (Fig. 13).

Figure 13: Depiction of the implanted electrode location in Wistar rat skull. Microelectrode array was inserted on the dura over the pre- and postparietal cortices between lambda and bregma. A stainless steel screw used as a reference/ground electrode (R) was placed 2-3 mm in front of bregma.

A stainless steel screw inserted 2-3 mm in front of bregma was used as a reference electrode connected to a ground potential. In the chronic recordings, electrodes were anchored onto the rat’s skull with screws and bone cement. After surgery

95


under medetomidine/ketamine anesthesia, the rat was aroused by giving the antagonist, atimepazole, and rats were allowed to recover for a minimum of 7 days before the first experiments. In the chronic trials, the implanted rats were either allowed to move freely or were gently immobilized by holding it securely in a towel. Both auditory and electrical current stimuli were used to generate EPs. The auditory stimuli consisted of 25 sets of paired pulses (pulse duration: 10 ms) with pulse intervals of 500 ms, followed by a 9.5 s quiet interval. The square wave current stimuli consisted of 25 sets of paired pulses (pulse duration: 1 ms) with pulse intervals of 500 ms, followed by a 9.5 s quiet interval, delivered to the left front paw of the rat. Auditory and somatosensory evoked field potentials were amplified with a preamplifier (Neuralynx Inc., Bozeman, MT, USA) and with a main amplifier (Grass Instruments, West Warwick, RI, USA) and led to the data acquisition PC running SciWorks (DataWave Technologies, Loveland, CO, USA). In study II, in vivo electrode impedance measurements were performed using a multi-task tester (Fredrich Hare Inc., Brunswick, ME, USA) in order to clarify the long-term performance of the MEAs.

7.5 CELL CULTURE STUDIES

Studies III, V and VI concentrated on investigating cellular responses on the surface of the biomaterial samples. The following chapters (7.5.1 – 7.5.4) describe the methodology used in these studies. 7.5.1 Cell lines In study V, hMSCs (PoieticsTM, Lonza Group Ltd., Basel, Switzerland) were cultured in Lonza Mesenchymal Stem Cell Growth Medium (MSCGM) at 37 °C in 5% CO2 in air. The cell monolayer was washed twice with 140 mM phosphate buffered saline (PBS, pH 7.4) and the cells were removed by applying 0.25% trypsin in PBS-EDTA (ethylenediamine-tetraacetic acid)

96


solution for 5 min at room temperature. The suspension was then transferred to a Falcon tube and trypsin was removed by centrifugation. The cells were resuspended in culture medium, seeded onto the biomaterial surfaces at a density of 0.52 x 104 cells/cm2 and cultured for 7.5 h or 5 days. In the osteogenic induction experiments, the hMSC cells were seeded at a density of 0.31 x 104 cells/cm2. Half of the samples were induced 24 h after seeding by replacing MSCGM with Osteogenesis Induction Medium (Lonza). Induced hMSCs were supplied every 3–4 days with fresh medium. Noninduced hMSCs (control samples) were fed with MSCGM using the same schedule. In study III, baby hamster kidney fibroblasts (BHK 21, clone 13, HPA Culture Collections, Salisbury, UK) were cultured in Dulbecco’s Modified Eagle’s Medium (DMEM, EuroClone S.p.A., Pero, Italy) supplemented with 10% fetal calf serum (FCS, PAA Laboratories GmbH, Linz, Austria), 50 µg/ml ascorbic acid (Sigma-Aldrich), 2mM l-glutamine (PAA Laboratories), 20 IU/ml penicillin (EuroClone), and 200 µg/ml streptomycin sulfate (EuroClone). The cells were seeded onto the surface of the samples (7mm x 7mm) by adding a medium which contained 5.0 × 104 cells/ml. The cells were cultured for 24 h at 37 °C in 5% CO2 in air. In study VI, human primary osteogenic sarcoma SaOS-2 (ECACC 890500205) cells were cultured in McCoy’s 5A culture medium containing GlutaMAXTM (Gibco BRL/Life Technologies Inc., Gaithersburg, MD, USA) and supplemented with 10% v/v FCS, 100 IU/ml of penicillin and 100 µg/ml streptomycin. Cells were seeded onto the biomaterial surfaces (10 mm x 10 mm) at a density of 15-25 x 103 cells/cm2, and cultured for 48 h or 120 h at 37 °C in 5% CO2 in air. Tissue culture-treated polystyrene 12-well (studies V and VI) and 24-well cell culture plates (study III) supplied from Corning Inc. (NY, USA) and Greiner Bio-One GmbH (Frickenhausen, Germany), respectively, were used in all experiments.

97


7.5.2 MTS assay The MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3 carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium)-based proliferation assay (CellTiter 96® Aqueous One Solution Reagent, Promega Corp., Madison, WI, USA) was used as a colorimetric method to assess the cytotoxic effects of different PI grades (study III) and also other potential material candidates for neural sensors (additional data). The MTS assay, a modification from the more conventional MTT (3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltetrasodium bromide) assay, is based on the reduction of MTS salt in the presence of phenazine methosulfate to a watersoluble formazan product in actively functioning mitochondria. Since it is water-soluble, no solubilization step is required with MTS as is the essential with the MTT assay. The quantity of formazan formed was determined by measuring the absorbance at 490 nm, which is directly proportional to the number of living cells in the culture. After 24-h cultivation of the BHK-21 cells, the samples were transferred into unused wells, and a fresh DMEM medium plus MTS reagent (200 µl) was added. The following incubation period lasted 1 h, after which the media were removed, the wells were washed with PBS and the absorbances were measured at 490 nm with an ELISA reader (Molecular Devices, Inc., Sunnyvale, CA, USA). The relative numbers of viable cells were determined by normalizing the optical density values to the highest measured values. 7.5.3 Scanning electron microscopy of cultured cells Scanning electron microscopy was used to analyze the cellular behaviour on biomaterial surfaces. In studies V and VI, after incubation period, samples were washed twice with PBS and fixed overnight in 2.5 % glutaraldehyde (Sigma-Aldrich) at +4°C. After this, the samples were washed with PBS and dehydrated in a graded ethanol series (from 50 % solution to absolute ethanol). The dehydration was completed using a Bal-Tec CPD 030 Critical point drying unit (Bal-Tec AG, Balzers, Liechtenstein). Samples were mounted on SEM stubs, coated with a Pt layer with an Agar sputter device (Agar Scientific Ltd.,

98


Stansted, UK) and examined using a Zeiss DSM 962 SEM (Carl Zeiss GmbH, Oberkochen, Germany) at an accelerating voltage ranging from 8 kV to 10 kV. In study III, the cells were fixed with 2.5% (w/v) glutaraldehyde (Sigma-Aldrich) in sodium cacodylate buffer (pH 7.4) and dehydrated with ethanol gradient and HMDS (Sigma-Aldrich). These samples were examined with an environmental SEM (Philips XL30 ESEM-TMP, Fei Company) at an accelerating voltage of 8–15 kV after being coated with a thin layer of Au using a Sputter Coater E 5100 (Polaron Equipment). 7.5.4 Immunofluorescence and confocal laser scanning microscopy After either a 7.5 or 120-hour incubation period with the hMSCs (study V) or after a 48-hour incubation period with the SaOS-2 cells (study VI), the specimens were triple stained with fluorescence dyes to reveal cytoskeletal (actin) and focal adhesion (vinculin) proteins and cell nuclei. A detailed description of the fixation, staining and viewing of the cells is given in original articles V and VI. Briefly, the cells grown on Si pieces were washed with PBS solution, fixed with paraformaldehyde solution, rinsed in PBS and permeabilized in Triton X-100 solution. After immersing in normal goat serum, the cells were incubated with the primary antibodies (antivinculin) and washed. Cells were then incubated simultaneously with fluorochrome-conjugated stains containing Alexa fluor 488 (Molecular Probes Inc., Eugene, OR, USA) and Alexa fluor 568 (Invitrogen Corp., Carlsbad, CA, USA) to detect focal adhesion (vinculin) and actin cytoskeleton. Cell nuclei were stained with DAPI (study V) or with TO-PRO-3 (study VI) probes and then the samples were washed again, and mounted on objective slides. Stained cells were viewed in a fluorescence microscope (Olympus AX70) and in more detail using a Leica TCS SP2 confocal laser scanning microscope (Leica Microsystems GmbH, Mannheim, Germany) equipped with appropriate filter sets.

99


In study V, the development of the osteoblastic phenotype was followed using alkaline phosphatase (ALP), osteocalcin (OC) and mineral staining (MS) as differentiation markers at 14, 17 and 21 days, respectively. A detailed description of the staining of hMSCs is given in original article V. After staining, the samples were fixed onto objective slides and coverslipped. The cells were observed using an Olympus AX70 fluorescence microscope in ALP and OC assays and a Leitz Diaplan microscope (Leica Microsystems GmbH, Wetzlar, Germany) in a mineralization assay. In study III, specimens were stained using a combination of 2 fluorescent probes, i.e. a cell-impermeable DNA-binding dye, propidium iodide (Sigma-Aldrich), and a cell-permeable fluorescein diacetate (Fluka Chemie GmbH, Buchs, Switzerland) in PBS to assess the viability of the BHK-21 fibroblasts cultivated for 24 hours on the samples. Detailed information on staining procedure can be found in original article III. Fluorescein diacetate is a non-fluorescent molecule that penetrates through the cell membranes. Inside the cell, fluorescein diacetate is hydrolyzed by non-specific esterases resulting in fluorescein release. Fluorescein emits green fluorescence when irradiated by laser light. Propidium iodide is a common fluorescent molecule used to stain DNA and to identify dead cells in a cell population. Since propidium iodide is membrane impermeable, it only stains dead cells which have leaking membranes. Propidium iodide emits a red fluorescence when excited. Since only viable cells are able to produce fluorescein, these two stains can be used to separate the living and dead cells in a cell populations. Cells are either stained with propidium iodide or fluorescein diacetate, but not with both. The samples were viewed in a confocal scanner (PerkinElmer Life Sciences, Wallac-LSR, Oxford, UK) on a Nikon Eclipse TE300 microscope (Nikon Corp., Tokyo, Japan), using the wavelengths 488/10 nm (excitation) and 525/50 nm (emission) for the fluorescein, and 568/10 nm (excitation) and 607/45 nm (emission) for the propidium iodide. Finally, the numbers of dead (red-staining) cells (nuclei/mm2) and the surface area of live (green-staining) cells were counted

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using Image J software 1.37c (National Institute of Health, Bethesda, MD) (Abramoff et al. 2004).

7.6 STATISTICAL ANALYSES

The numerical results are expressed as mean ± standard deviation of the mean (SD) or as mean ± standard error of the mean (SEM). In studies III, V and VI, statistical analyses were conducted with the SPSS software (versions 14.0 and 16.0, SPSS Inc., Chicago, IL, USA). One-way analysis of variance (ANOVA) followed by Tukey Post-Hoc Tests were applied to determine the statistical significance of the differences observed between material groups. Values of p < 0.05 were considered statistically significant.

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102

8 Results 8.1 EVALUATION OF MICROELECTRODE ARRAYS

A number of MEA prototypes were developed during the 2-year trial period by optimizing the substrate material, substrate thickness, electrode material, sensor layout and insulation material (Table 6). It was demonstrated that the magnetron sputter deposited and photolithographically patterned Pt thin films between two layers of PI represented a very powerful and reproducible way to produce miniaturized neural interfaces (Fig. 14). The ZIF type connector allowed the easy and reliable connection of a thin film MEA to recording instrumentation (Fig. 14b). The use of the ZIF connector eliminated the need for soldering, which had been observed to be very challenging to perform on the thin Pt pads produced by the sputtering technique on PI. Furthermore, the use of PSPI (PI-2771) simplified the electrode fabrication considerably because this material could be patterned directly in the same way as photoresists without using complex multilevel processes (i.e. lithography, etching mask formation, etching mask dissolution) which were unavoidable steps when conventional PIs were used. It was possible to pattern the PI-2771 layer with good alignment to the metallization layer (Fig. 14f,g). Moreover, one could pattern the PI-2771 layer with an almost vertical pattern profile and without residuals (Fig. 14h). The electrode type B, originally presented in study II, was found to be the most suitable for recording cortical surface field potentials in rats. The major shortcomings of the other prototypes compared to type B are summarized in Table 6.

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8 – Results

A

B

connector board

transmission lines

recording sites

C

5 mm

D

5 mm

E

1 mm

F

G

200 µm

H

1 mm

50 µm

Figure 14: The developed polyimide-based microelectrode arrays (a). The PI ribbons are connected via a connector board consisting of a ZIF- type connector (JST Ltd.) and a surface mount microsocket (Samtec Inc.) to the recording instrumentation (b). The developed array is highly miniaturized as shown in comparison to a clinical subdural grid electrode (c), and extremely flexible (d). A SEM-image of the 8-channel array viewed at the end of the recording sites before deposition of the insulation layer (PI2771) (e). Patterning of this polyimide insulator layer achieved a good alignment with the metallization layer (f,g). The 3 µm thick polyimide layer had an almost vertical pattern profile and was virtually free of residuals (h).

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8 - Results Table 6: The summary of developed microelectrode array prototypes. The main disadvantages of each array type compared to type B are presented. Type Channel number

Substrate

Electrode Insulator diameter (µm)

A

8

PI-2525a (3 layers)

Ti/Pt, 200

PI-2771a

Troublesome fabrication

B

8

Kapton HNb + PI-2525

Ti/Pt, 200

PI-2771

–

C

8

Kapton HN + PI-2525

Ti/Au, 200

PI-2771

Higher electrode impedances

D

16

Kapton HN + PI-2525

Ti/Pt, 100

PI-2771

Short-circuiting

E

16

Kapton HN + PI-2525

Ti/Pt, 100

SU-8c

Short-circuiting and cracking

Abbreviations: Ti: titanium; Pt: platinum; Au: gold a HD Microsystems GmbH, Bad Homburg, Germany; c Microchem Corp., Newton, MA, USA

b

Weaknesses

Goodfellow Ltd., Cambridge, UK;

Due to the viscosity and chemical properties of PI-2525, the formation of a 30 µm thick substrate layer required three spin coating cycles (type A). This was somewhat complicated compared to type B (study II), in which the same thickness could be readily achieved by using Kapton HN film together with one spin-coated PI-2525 layer. Gold electrodes (type C) have significantly higher electrode-electrolyte interface impedance levels compared to Pt electrodes especially at low frequencies below 1 kHz which are the most relevant when biological signals need to be recorded (Fig. 15). The microelectrode surface area in high-density arrays (types D and E, electrode diameter 100 µm) was only 25 % of the area of the larger electrode (types A-C, electrode diameter 200 µm) causing significantly higher electrode impedances.

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8 – Results

Au Pt

107

|Z| (Ω)

106 95.5 k

105 25.5 k

104 103

frequency (Hz) 10

Phase angle (°)

-20

102

103

104

105

Au Pt

-40 -60 -74.1

-80

-81.0

Figure 15: Impedance spectroscopy of platinum (Pt) and gold (Au) microelectrodes (diameter 200µm), measured in 0.9 % saline solution at room temperature. The impedance magnitude and phase are presented with error bars corresponding to standard deviations (SD). In some instances, SD values were often smaller than the symbol size and thus not shown. The absolute values of the impedance magnitude and the phase are also given at 1 kHz.

Designs D and E were also susceptible to short-circuiting by cerebrospinal fluid when the array was implanted onto the surface of the rat cortex. Moreover, epoxy-resist SU-8 (type C) was observed to be totally unsuitable for present electrode application. While PI- insulated arrays were very robust and they could be bent even into sharp angles without any cracks, the bending/implantation did create many cracks and even delaminations of the SU-8 layers. The developed electrodes were tested in acute and chronic evoked potential recordings in Wistar rats. The flexibility of PI was particularly suitable for following the curved cortex surface,

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8 - Results

thus ensuring good contact between electrode site and the dura mater. This enabled reliable acquisition of cortical signals. During the acute recording sessions, the MEA type B was capable of capturing stable recordings as observed in the form of the standard response parameters, such as latencies, onsets and decays of the main components in the potential traces (Fig. 16a). The recorded signal traces were biologically meaningful and reproducible, this being verified by comparing the responses at the beginning of the experiments with those obtained at the end. In chronic recordings, the electrode was capable of yielding signals comparable with the presented acute signals for approximately two weeks after the array implantation (Fig. 16b). However, after 16 days, the responses decayed. This decline in signal amplitudes was concurrent with the observed jump in the electrode impedances which were systematically measured in vivo at three-day intervals (Fig. 16c). The physical appearance of the dura around the electrode occasionally changed during chronic implantation. It became thicker and encapsulated the electrode. However, in microscopic examination, the surface of the brain appeared to be undamaged with the electrode location being slightly depressed. Although no macroscopic scar formation was seen in the examination of the brain after the array was removed, this kind of jump in the in vivo impedances was compatible with a thickening of the dura, a common finding in chronic experiments, and microscopic growth of non-conductive fibrous tissue around the electrodes, shrinking the freely exposed surface area of the recording electrode.

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8 – Results

Figure 16: (a) An acute recording of somatosensory evoked potentials (SEP) in the rat parietal cortex. Profiles of averaged SEP traces using the 8-channel microelectrode array are shown on the right with their correspondence to the recording sites which are numbered in schema in left side. The position of reference/ground electrode is marked as “R”. The current stimuli consisted of 25 sets of paired pulses with pulse intervals of 500 ms, followed by a 9.5 s quiet interval, delivered to the left front paw of the rat. (b) A chronic recording of auditory evoked potentials in the rat parietal cortex. The stimuli consisted of 25 sets of paired pulses at intervals of 500 ms, followed by a 9.5 s quiet interval. (c) Average electrode impedance (mean ± SD, n = 8) measured in vivo.

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8 - Results 8.2 EVALUATION OF NANOROUGH ELECTRODE SURFACES

The USPLD technique was used to prepare the Pt coatings on SS, PI and Si with a desired nanostructured surface topography. The adhesion of coatings deposited upon all the tested substrates was good without any signs of delamination. The USPLD process parameters were adjusted to achieve different surface topographies. Fig. 17 shows an example where the feature size is about 30 nm. The nanostructured electrode surface clearly increased the effective surface area compared to a smooth Pt surface typically achieved using sputtering.

Figure 17: AFM image of nanostructured Pt-electrodes produced using the USPLD technique. The average roughness of the surface is 18 ± 4 nm (mean ± SD)

Electrochemical impedance spectroscopy was applied in order to study the electrode-electrolyte impedance behaviour of different Pt-coated SS electrodes (unpublished data). A total of six sample areas (50 mm2) of each plate were examined using the measurement set-up described in chapter 7.3.1. The mean impedance magnitude and phase data for Pt coating, whose surface topography was demonstrated in Fig. 17, are shown in Fig. 18. Pure SS electrode and sputter-coated Pt on SS are

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8 – Results

presented as reference materials. The results clearly reveal that the presence of the nanostructures on an electrode surface had increased the surface area and decreased impedance. The impedance magnitudes (mean ± SD) at 1 Hz were 7.0 kΩ ± 0.5 kΩ, 17.8 kΩ ± 1.1 kΩ and 21.7 kΩ ± 1.0 kΩ for nanostructured USPLD-Pt, sputtered-Pt and pure SS, respectively. The differences in impedance values between the materials were statistically significant (p < 0.001) at low frequencies. Due to the relative large geometrical surface area, all coatings reached the same impedance magnitude level of approximately 180-200 Ω by 1 kHz, which corresponds to the resistance value of electrolyte solution. All surfaces are highly polarizable, but nanostructuring decreases the phase difference between current and voltage making electrodes less polarizable. 105 SS.dat Sputtered Pt.dat USPLD-Pt.dat

0

104

-40 103

Phase angle (°)

|Z| (ohm)

-20

-60

102 100

101

102

103

104

-80 105

Frequency (Hz) Figure 18: Impedance magnitude and phase of a nanostructered platinum-coating (USPLD-Pt) and sputter-coated Pt-coating deposited on stainless steel (SS). Results for pure SS electrode are presented for reference. A clear drop in electrode impedance can be seen for USPLD-Pt, especially at low frequencies. Note the logarithmic scale.

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8 - Results 8.3 CYTOCOMPATIBILITY TESTING OF INSULATOR MATERIALS

Prior fibroblastic BHK-21 cell experiments, spin-coated PIs (PI2525 films and PI-2771 films cured at 200 °C and at 350 °C) and USPLD coatings (C3N4 and Al2O3) were characterized in terms of their roughness, wettability and zeta potential (Table 7). The surface roughness values obtained were all very low, i.e. the Ra values were below 1 nm. However, the Ra value for the PI-2525 was significantly higher than for other coatings (p < 0.001). The contact angles for water on both PI-2771 film types were significantly higher than for the PI-2525 (p < 0.001). The same tendency was also noticed when contact angles were measured for diiodomethane (results are shown in the original article III, Table 1) highlighting the more hydrophobic nature of the PI2771 with significantly lower polar and dispersive surface energy values (p < 0.001). Somewhat surprisingly, the effect of the chosen curing protocol (200 °C vs. 350 °C) for the PI-2771 on roughness or on wettability was not statistically significant. The contact angles for water on C3N4 were in the same range as on PI-2525 film, evidence of a slightly hydrophobic thin film coating. In contrast, the Al2O3 coating was very similar to PI2771 film with respect to its wettability properties. All of the surfaces were negatively charged, and the zeta potential values measured at a fixed pH of 7.4 were in a narrow range of -54.2 mV – -66.2 mV. The PI-2771-200 film followed by the PI-2525 film was the most negatively charged and their zeta potential values differed significantly from the all other test materials (p < 0.001).

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8 – Results Table 7: Average surface roughness (Ra), water contact angle and zeta potential values of tested polyimide and ceramic coatings. All of the materials were extremely smooth although the roughness of PI-2525 was significantly higher than that of other materials. The PI-2525 and carbon nitride (C3N4) were more hydrophilic compared to alumina (Al2O3) and both PI-2771 types. Zeta potential values were all in the same range by demonstrating the negative surface charge of these coatings. Material

Ra (nm)

Contact angle (°)

Zeta potential (mV)

PI-2525

0.89 ± 0.13 *

69.5 ± 1.8 ##

-59.9 ± 0.3 *

PI-2771-200

0.43 ± 0.15

95.1 ± 1.3

-66.2 ± 1.6 *

PI-2771-350

0.39 ± 0.04

93.2 ± 0.8

-54.2 ± 0.1

C3N4

0.24 ± 0.05 #

70.8 ± 0.9 ##

-55.1 ± 1.2

Al2O3

0.18 ± 0.03 #

94.0 ± 1.3

-56.6 ± 3.4

* p < 0.001, as compared to all other materials # p < 0.001, as compared to all polyimides ## p < 0.001, as compared to all other materials, except of C3N4/PI-2525 The values are mean ± SD. One-way ANOVA was used to determine the statistical differences between materials.

In this study, the cytotoxicity of the insulator candidate materials was examined with in vitro tests conducted along to the guidelines of the international standard ISO-10993-5. Latex rubber and PE were used as positive and negative control material, respectively. The MTS assay defined the number of viable BHK-21 cells attached on the surfaces 24 h after cellseeding. The results were normalized relative to the highest value, i.e., the value for the PE control (Fig. 19). The relative cell numbers in the different polyimide groups and C3N4 were at a level of 62-70 % of the PE controls. The number of cells attached to the Al2O3 surface was lower, i.e. 50 %, but still significantly higher than attained with the latex rubber control. Fluorescence microscopy studies with live/dead staining probes revealed that all test materials were highly cytocompatible, in contrast to latex rubber, on which viable cells were rarely found, instead dead cells had most probably detached and were floating in the medium (illustrated in the original article III). Results from the quantitative analysis included the determination of the density of dead cells

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(nuclei/mm2) and the surface area of live cells (µm2) is presented in the Table 8. The numbers of dead cells were very low on each surface, i.e., mean values ranged from 13 cells/mm2 to 48 cells/mm2. Even though the dead cell density was also minimal on the surface of latex rubber, the cytotoxic effect could be clearly seen when the surface area of viable cells was examined. The total surface area of viable cells on latex rubber was significantly lower (p < 0.001) than that of the other test materials, being only around 2% of the corresponding values obtained for other surfaces. The values for C3N4 and Al2O3 were slightly but not significantly higher than that for PE control or PI materials. Table 8: The density of the dead cells (nuclei/mm2) and the surface area of viable cells (µm2) on polyethylene (PE, negative control), different polyimide surfaces, alumina (Al2O3), carbon nitride (C3N4) and latex rubber (positive control) after a 24-hour cultivation period evaluated using a combination of two fluorescent probes and Image J software. Material

Density of dead cells (1/mm2)

Surface area of viable cells (µm2)

PE (n = 10)

13.9 ± 5.4

70400 ± 2900

PI-2525 (n = 13)

13.5 ± 5.5

71900 ± 8500

PI-2771-200 (n = 8)

18.8 ± 8.2

67300 ± 13200

PI-2771-350 (n = 8)

30.8 ± 8.8

63000 ± 5800

C3N4 (n = 13)

48.0 ± 10.3

83900 ± 11600

Al2O3 (n = 15)

47.4 ± 5.8

82600 ± 12300

Latex rubber (n = 8)

24.9 ± 8.6

1400 ± 320*

The values are the mean ± SEM. *p < 0.001 (One-way ANOVA).

SEM imaging was used to monitor cell morphology, adhesion and spreading on the test surfaces. The BHK-21 cells which had been cultivated for 24 h on each PI surface, Al2O3 and C3N4 exhibited a normal fibroblastic morphology, adhered well with filamentous extensions, and had started to form clusters (Fig. 20). Thus, all insulator materials were considered not to be toxic and equally good for allowing the adhesion and spreading of cultured fibroblasts.

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Relative cell number (%)

A

120 100

*

80 60 40

*

20 0 PE PE

Cell coverage (%)

B

20

PI-2525 PI2771-200 PI-2771-200PI2771-350 PI-2771-350 Alumina Al2O3 PI2525

CCN 3N4

Latex

15

10

5

*

0 PE

PI-2525 PI2771-200 PI-2771-200PI2771-350 PI-2771-350 alumina Al2O3 PI-2525

C CN 3N4

Latex Latex

Figure 19: (a) Relative number of BHK-21 cells on polyethylene (PE, negative control), conventional polyimide (PI-2525), two differently cured photosensitive polyimides (PI2771-200 and PI-2771-350), alumina (Al2O3), carbon nitride (C3N4) and latex rubber (positive control). The data was obtained by MTS assay and normalized to the highest optical density value. (b) Surface area covered by BHK-21 cells at 24 h on the same test materials. The data was acquired from confocal laser scanning microscope images of the samples with live/dead-stained cells. All polyimide and ceramic coatings can be considered to be non-cytotoxic and they allowed adhesion and spreading of fibroblasts. The error bars indicate standard errors of the means. *p < 0.001 (one-way ANOVA).

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Figure 20: SEM images of BHK-21 cells cultured for 24 h on PI-2771-200 at two different magnifications. The cells exhibited a normal fibroblastic morphology, adhered well with filamentous extensions, and had even started to form clusters. Cells cultured on the other test materials, i.e. PI-2525, PI-2771-350, C3N4 and Al2O3, appeared to be very similar and thus, are not shown.

8.4 MICROPATTERNED BIOMATERIAL COATINGS

The interactions of the hMSCs and SaOS-2 cells on micropatterned biomaterial (Ti, Ta, Cr, DLC) coatings were investigated in studies V and VI. The roughness and surface energy data of test materials were originally published in paper V, but these values are also valid for the samples used in study VI. A summary of these results, supplemented with the zeta potential values achieved for smooth (non-patterned) surfaces is given in Table 9. All of the materials used in these studies were extremely smooth, with Ra values of less than 2 nm leading to the hypothesis that the observed differences in cell behaviour were not due to the differences in surface topography but were mediated by different chemical compositions and physical patterns of material studied. The only relevant topographical cues of the samples were attributable to the thickness of coating (~ 200 nm) that forms steps between micropatterns and the background. All coating materials also possessed very similar wettability properties, i.e. their SFE values were in a range between 47.3 - 52.1 mJ/m2. The Si wafer used as background material in these experiments had a significantly higher SFE (68.5 mJ/m2) compared to all other coating materials. The negative zeta potential values obtained from streaming current

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measurements at a fixed pH of 7.4 disclosed that the surfaces of all coating materials were negatively charged. The differences in zeta potential values between the surfaces were statistically significant (p < 0.001 for all). Table 9: Average surface roughness (Ra), water contact angle, surface free energy and zeta potential values for the surfaces of different test materials. All surfaces were extremely smooth (i.e. mirror finish) and their contact angle/surface energy values were in a narrow range (except for background Si), although some statistically significant differences existed. Material

Ra (nm) *

Contact angle (°)

Surface energy (mJ/m2)

Zeta potential (mV) *

DLC

0.3 ± 0.1

67.4 ± 2.0

49.6 ± 1.2

-70.2 ± 1.4

Ti

1.8 ± 0.2

67.6 ± 2.0

49.7 ± 1.1

-53.9 ± 0.3

Ta

1.5 ± 0.3

64.1 ± 1.6#

52.1 ± 1.1#

-49.6 ± 1.1

Cr

0.8 ± 0.1

71.2 ± 1.1#

47.3 ± 0.7#

-39.6 ± 2.2

Si

0.2 ± 0.1

32.4 ± 1.1#

68.5 ± 0.8#

-43.5 ± 1.3

*p < 0.005 between all materials #p < 0.05, as compared to all other materials The values are mean ± SD. One-way ANOVA was used to determine the statistical differences between materials.

8.4.1 Behaviour of human mesenchymal stem cells An incubation time of 7.5 hours was used to investigate the early-stage adhesion, contact guidance and morphology of hMSC cells on the micro-patterned DLC, Ti, Ta and Cr thin films deposited on Si (study V). Based on observations with the SEM, hMSCs seemed to prefer the biomaterial patterns over a Si background. This kind of guiding effect seemed to be evident with all materials, although only Ti and DLC surfaces are presented here (Fig. 21a,b). Quantitative analysis showed that on Ta, Cr and DLC, the density of hMSCs was 3.0-3.5 times higher than on the background (p < 0.0005 for all, Fig. 22). One exception was Ti, on which the density of hMSCs on the patterns did not differ from the density of the cells grown on the Si background (p = 0.14). The analysis of the surface area of patterns and Si background covered by the MSCs at 7.5 h gave

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results consistent with the results from the cell density assessment (Fig. 23). From this viewpoint, the guiding effect was also clear on all the materials (p < 0.05) except for Ti (p = 0.46).

Figure 21: hMSC cells after 7.5 h culture (a,b) and 5 days culture (c,d) on patterned titanium (a,c) and diamond-like carbon (b,d) surfaces. The cells seemed to prefer the patterns over background and some cells were aligned along the edges of the patterns (a,b). At 5 days, cells were much larger and they had grown to form a subconfluent monolayer. The side length of the biomaterial square is 75 µm.

One interesting observation was made about the size of cells. The mean diameter of the cell lying on patterns (68 µm) was significantly lower (p < 0.05) than that of cells growing on a Si background (102 µm). Moreover, it was observed that numerous cells were aligned along the edges of the patterns. This phenomenon was confirmed in the immunofluorescence triple staining, which revealed that focal vinculin adhesions and actin cytoskeleton were outgoing from the pattern edges i.e. cells were believed to be taking on the geometric square shapes.

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Background (Si) Patterns Average

140

Cell density (cells/mm2)

120 100 80 60 40 20 DLC

Cr

Ta

Ti

Figure 22: Density of hMSCs at 7.5 hours. On lithographically patterned diamondlike carbon (DLC), chromium (Cr) and tantalum (Ta) samples, the cells significantly preferred the biomaterial patterns over a silicon (Si) background. The cell density on titanium (Ti) patterns was marginally lower (# p ≈ 0.06, one-way ANOVA) than that on Cr and Ta patterns. The error bars indicate standard deviations of the means. 100 Background (Si)

Cell coverage (%)

80

Patterns

*

*

*

*

60

#

40

+ 20

0 DLC (7.5 h)

DLC (5 d)

Cr (7.5 h)

Cr (5 d)

Ta (7.5 h)

Ta (5 d)

Ti (7.5 h)

Ti (5 d)

Figure 23: Coverage of hMSCs at 7.5 hours and 5 days. A clear guiding effect seen at 7.5 hours was nearly lost by 5 days. * p < 0.05, as compared to DLC and Cr, # p < 0.05, as compared to Cr and Ta, + p < 0.05, as compared to DLC. The error bars indicate standard deviations of the means. One-way ANOVA was utilized to determine the statistical differences.

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The spreading of the hMSCs was investigated after 5 days of incubation (Fig. 21c,d). At this stage, the individual cells had grown considerably achieving a size which greatly exceeded that of a singular square pattern (75 µm x 75 µm), which meant that it was no longer possible to calculate the cell densities. Surface coverage analysis indicated that over 65 % of the sample surface (of each material) was covered by cells at 5 days (Fig. 23). The guiding effect was virtually lost since the cell density was only 1-1.9 times higher on patterns than that on Si background. The only statistically significant difference was observed between the DLC patterns and the Si background. Somewhat surprisingly, the cell density on the very same Si background varied statistically significantly, depending on the material used in microfabricated patterns. At 7.5 h, the hMSC-covered surface of the background Si was significantly lower on Ti than on DLC patterns (p = 0.015), but by the 5th culture day this situation had reversed (p = 0.006) (Fig. 23). This indicates that apart from direct material-cell interactions, also probably indirect cell-cell interactions were present. The cells on material islands had remote interactions with cells growing on Si background. After 5 days, vinculin containing focal adhesion were already well developed all over the sample surface without any preferential localization on the square patterns as compared to the background. Osteogenic differentiation of hMSCs was investigated by using ALP, OC and MS as the differentiation markers. The results showed that the patterned surfaces, except for DLC, allowed induced osteogenic differentiation although this was less effective than on plain samples. It seemed to be so that cells growing on the micropatterned surfaces did not have enough space to organize themselves into multicellular bone tissue-type structures. Moreover, the rectangular shape which the hMSC cells took due to microsquare shaped islands might not be ideal with respect to osteogenesis. The DLC coating did not support osteogenesis, and for that reason an inert DLC seems to be a potential coating material for applications which come into

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contact with bone but which need to be removed later, e.g. various fracture fixation devices, such as plates and screws. 8.4.2 Behaviour of osteoblast-like (SaOS-2) cells The SEM examination of SaOS-2 cells cultured on micropatterned Ti and DLC coatings (on Si) for 48-hours revealed a clear cellular preference for the biomaterial patterns over the Si background. The cells preferentially adhered and spread well on the large-sized (125 µm) circular (Fig. 24a) or rectangular (Fig. 24b) Ti patterns, which facilitated their occupation by several cells. On the reverse-patterned samples containing Si features (circles or squares) on a Ti background (Fig. 24c), the inverse phenomenon was observed, with almost all of the cells now attached to the Ti background, but they seemed to avoid spreading out into the Si circles. The cytocompatibilityenhancing and cell-guiding effects of DLC were also clear, though much weaker than those encountered with Ti, as demonstrated in samples containing DLC circles on Si (Fig. 24d) and in the reverse-patterned samples containing Si squares on a DLC background (Fig. 24e). The influences of the pattern shape and its size on adherence and morphology of SaOS-2 cells were studied using rectangular and circular patterns of three different sizes (diameters of circles and sides of squares: 5, 25 and 125 µm). Large-sized patterns facilitated the adhesion of several cells onto one island and the boundaries of the outermost cells followed the shape of the edges of the patterns to which they had attached (Fig. 25a,b). The medium pattern size (25 µm) was designed so that singular cells could inhabit individual patterns. This pattern size seemed to restrict the growth of cells although they no longer adapted so strictly to the geometrical shape of the patterns (Fig. 25c). However, neighboring patterns were so far away that cells could not really join together to form clusters. Particularly on circularly patterned surfaces, cells occasionally covered two or more islands, leading to longitudinal or star-like morphologies (Fig. 25d). On the small-sized patterns, cell bodies covered several of the small patterns (Fig. 25e), but their thin and slender

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filopodia seemed to prefer the more cell-friendly material, and they seemed to try to avoid bare Si areas (Fig. 25f).

Figure 24: SaOS-2 cells cultured for 48 h on surfaces containing large-sized (125 µm) features. There is a clear cellular adhesion preference for the Ti circles (a) or squares (b) over background Si, and in the case of reverse-patterned samples for the Ti background over the Si circles (c). The cell-guiding effect of DLC was also clear even though weaker than that of Ti as demonstrated in samples containing DLC circles on Si (d) and in reverse-patterned samples containing Si squares on a DLC background (e). Scale bar is 100 µm.

Immunofluorescence triple staining disclosed that the actin cytoskeleton was well organized in SaOS-2 cells growing on the biomaterial islands, but poorly organized in the cells lying on Si.

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Furthermore, it was observed that the formation of vinculin containing focal adhesions was much enhanced in cells adhering onto the biomaterial patterns compared to cells lying on the Si backgrounds.

Figure 25: SEM images of SaOS-2 cells cultured for 48 h on micropatterned Ti (a-d) and DLC (e,f) surfaces. Large-sized (125 µm) squared (a) and circular (b) features facilitated the adhesion of several cells on one Ti island with the cells aligning themselves along the edges of the cell-friendly material. The cells adhering to mediumsized Ti islands no longer conformed strictly to the geometrical shape of the patterns (c) but particularly on circularly patterned surfaces, star-like cellular morphologies appeared (d). On small-sized inverse DLC samples, the cell bodies non-selectively covered large micro-patterned areas (e), but their filopodia clearly showed a preference for DLC trying to avoid bare Si circles (f).

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Quantitative analysis of the SaOS-2 cell density, coverage and average cell size was performed with samples containing 75 µm squared patterns of Ti, Ta, Cr and DLC on the Si substrate. The relative cell density at 48h was 2.9-3.7 times higher on all four coating materials than that achieved on the Si background (p < 0.001). The relative coverage was 2.0-3.0-fold higher for Ti, Ta, Cr and DLC compared to the corresponding value of Si (Fig. 26). A comparison between the four patterning materials indicated that the proportions of surface area covered by cells were higher for Ti and Ta than for DLC and Cr (p < 0.05). Further, the Si background was better covered with the Ti- and Cr-patterned samples than the background of the DLC and Ta samples (p < 0.001).

Cell coverage (%)

80

60

40

20

DLC

Cr

Ta

Ti

Figure 26: Coverage of SaOS-2 cells was always higher on biomaterial patterns than on silicon (Si) background. Titanium (Ti) and tantalum (Ta) were better covered than diamond-like carbon (DLC) and chromium (Cr) as designated with # (p < 0.05) and the Si background on Ti- and Cr patterned samples was better covered than the background on DLC or Ta samples (##, p < 0.05). The error bars indicate standard deviations of the means. One-way ANOVA was applied to determine the statistical significance of the differences observed between material groups.

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The average sizes of the cells cultured for 48-hours on each sample type are given in Table 10. The cells growing on the biomaterial patterns (each material) were significantly smaller (13-41 %) than the cells growing on Si. A comparison between different biomaterials disclosed that the cells growing on Ti were larger than the cells on Cr (p < 0.005), and the cells on Ta patterns were larger than the cells on DLC or Cr patterns (p < 0.05). Further, the cells growing on Ti-patterned samples were also significantly larger on the Si background than on backgrounds containing patterns fabricated from other coating materials (p < 0.01). This was somewhat surprising since the background material and the preparation techniques used in sample fabrication were the same for all samples. Table 10: The size of SaOS-2 cells on micropatterned samples. The SaOS-2 cells were significantly larger on background silicon compared to the cells growing on patterns. In addition, the cells were significantly larger on the Si background of Ti-patterned samples than on the backgrounds of other materials. Material

Cell size on Si background (µm2)

Cell size on patterns (µm2)

DLC

1282 ± 60

962 ± 45

Cr

1404 ± 102

954 ± 25

Ta

1322 ± 86

1151 ± 65#

Ti

1786 ± 77*

1059 ± 20**

* p < 0.01, as compared to all other backgrounds # p < 0.05, as compared to DLC and Cr ** p < 0.005, as compared to Cr The values are mean ± SD. One-way ANOVA was used to determine the statistical differences.

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9 Discussion 9.1 MICROELECTRODE ARRAYS FOR INTRACRANIAL RECORDINGS

In studies I, II and IV, the design, fabrication and experimental use of novel flexible polymer-based microelectrode arrays was demonstrated. Successful acute and chronic intracranial multichannel recordings performed on the surface of the rat cortex clearly revealed the potential of these arrays. In particular, the PI-Pt-PI sandwich type was proved to be capable of acquiring stable signals with a high signal-to-noise ratio in evoked potential recordings. The PI-based arrays demonstrated excellent flexibility and mechanical strength during handling and implantation onto the surface of rat cortex. The flexibility also allowed the array to follow the movements of the rat brains by keeping the recording points in close contact with the target area, thus promoting reliable signal acquisition. Although SU-8 epoxy-resist has been recently utilized as an insulating material in implantable neural electrodes (Hollenberg et al. 2006, Yeager et al. 2008), implantation (bending) of SU-8 insulated array developed here caused several cracks in the SU-8 layer. Sputtering and evaporation are traditional techniques used in producing thin film microelectrodes. The results obtained from EIS with our magnetron sputtered Pt and Au microelectrodes were consistent with the previously reported values (Cheung et al. 2007, Metz et al. 2004, Stieglitz et al. 2000, Takahashi et al. 2003). The Pt electrodes displayed significantly lower electrodeelectrolyte impedance magnitudes compared to the Au electrodes, particularly at biologically relevant frequencies, i.e. below 1 kHz. With the help of AFM imaging, the surface of sputtered Pt or Au was observed to be extremely smooth (Ra < 1 nm) i.e. its effective surface area was nearly the same as its geometrical dimensions. The impedance levels for the Pt

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electrodes used in the final array prototypes (particularly for electrodes with diameter of 200 µm) were adequate, but if one needs to develop smaller-sized neural electrodes for highresolution application, it would be advantageous to increase the effective surface area of microelectrodes. The fabrication process developed here enables the fabrication of miniaturized arrays with high channel densities, which have benefits over microwire bundles in terms of precise control of the electrode sizes and separations between electrodes. The lithography-based process also allows implementation of several other neural interfaces with different layouts, material selections or target areas either for recording or stimulation purposes. Furthermore, these reproducible devices can be mass produced in batches at low cost. As concluded in paper IV, the technology developed in this thesis work also offers several attractive applications for clinical use. The high density arrays could possibly be utilized to achieve more accurate mapping the brain e.g. in patients suffering from epilepsy, Parkinson’s disease or brain tumors, enabling more precise planning of the neurosurgery. The flexibility also eases the insertion of the array without the need for removal of large skull areas, in contrast to the commercially available, somewhat bulky, designs of subdural strip and grid electrodes. In addition, it is obvious that prosthetic systems used in electrical stimulation of neural tissue in order to treat neuropathic pain, replace lost sensory or motor function will also benefit from further development of this kind of technology. As a result of study IV, a comprehensive summary of miniaturized flexible microelectrode arrays was obtained. Some of these MEAs have been developed from the point of view of human applications (Kitzmiller et al. 2006, 2007, Molina-Luna et al. 2007, Takahashi et al. 2003, Hollenberg et al. 2006). The technique of small-field cortical surface recordings, introduced by Kitzmiller et al. (2006) was based on lithographically patterned Pt electrodes on the PDMS sheet. Although being highly miniaturized compared to present clinical electrodes, this design did not take full advantage of all that microfabrication

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can offer but used fragile wire bonding to create the electrical wiring. It was realized that the transmission lines could be formed in the same deposition step as the electrodes. Thin film transmission lines firmly attached to the substrate surface were demonstrated to represent a robust way to establish the signal acquisition. By using somewhat similar technology as applied in the present thesis, Molina-Luna et al. (2007) introduced epidurally implantable thin-film MEA for cortical stimulation mapping. Their approach was based on 72 round TiN contacts (diameters 100 µm) deposited on the surface of PI foil which was utilized for mapping the motor cortex of the rat. In addition, a Kapton-based MEA consisting of 8 x 8 Au electrodes has been developed for epidural recording of surface field potential from the barrel cortex of rats (Hollenberg et al. 2006). However, the array developed by Hollenberg et al. is thicker (75 µm) and thus stiffer than the system developed here and it is also unsuitable for chronic recordings due to the bulky circuitboard that the rat would have to carry on its head. A flexible PIbased MEA with a bending structure introduced by Takahashi et al. (2003) was successfully tested in epidural mapping of AEP recordings. However, that system was not designed for chronic implantation. On the contrary, the connector board with the two miniconnectors used in our arrays was shown to be small and light enough to allow chronic recordings. The acute EP recordings obtained with the present MEAs were comparable in their signal-to-noise ratios with previously reported recordings with wire electrodes made with platinum (Hayton et al. 1999), tungsten (Ureshi et al. 2004) or silver (Kalliomäki et al. 1998, Jones & Barth 1999). They were also comparable to previous recordings obtained with MEAs (Stett et al. 2003, Hollenberg et al. 2006, Takahashi et al. 2003). The quality of the recorded signal at the beginning of chronic EP experiments was similar to those obtained in acute experiments for about two weeks. Subsequently, the signal amplitude declined due to the increase in electrode impedance resulting, most probably from the thickening of the dura and the growth of electrically insulating fibrous tissue around the electrodes.

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9 – Discussion 9.2 MATERIAL SCIENCE STRATEGIES TO IMPROVE THE PERFORMANCE OF MICROELECTRODE ARRAYS

Due to the limited long-term performance of the developed arrays, the confirmation about biocompatibility of materials used in those arrays was sought in study III. Moreover, different material science strategies were investigated in attempts to devise a solution to making the functional life span as long as possible (unpublished findings). Unfortunately, when implanted in neural tissue, all artificial materials have an inconvenient tendency to induce the formation of scar tissue, which tends to encapsulate the electrode surface, increases electrode impedances and lowers signal amplitudes. In present arrays, the greatest doubt with this vigorous scar tissue formation and reduced biocompatibility was addressed to PSPI (PI-2771, HD Microsystems) used as an insulation material. However, the results from study III, performed according to the ISO 10993-5 standard, revealed that PSPI is almost as non-cytotoxic as conventional PI and PE (used as a negative control material). The PSPI films did not adversely affect the BHK-21 fibroblast cell function, and hence it seems to be an appropriate material for incorporation into implantable biomedical devices. However, studies were performed only with one cell line and with a moderately short culture period. Therefore, further in vitro and in vivo studies will be needed to clarify its long-term effects. Platinum electroplating has been shown to represent an effective method to produce sponge-like Pt surfaces with much higher roughness and increased effective surface areas (de Haro, et al. 2002). Although this technique is a powerful way to decrease electrode impedance (Gonzales & Rodriguez 1997), and to prolong the effective life-span of microelectrode (Chen et al. 2009), it has also some serious disadvantages. Electrodeposited Pt adheres weakly to the electrode surface, resulting in unstable electrode impedance. Further, electrolytes commonly used in Pt electroplating contain lead, which is known to be a highly toxic heavy metal evoking a degree of uncertainty about the

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biocompatibility of electrodeposited Pt surfaces (Schuettler et al. 2005). In this study, the USPLD technique (ColdabTM ) was used to deposit Pt thin films with a nanotextured surface. Compared to the sputter-coated smooth Pt surface, a significant reduction in electrode-electrolyte impedances was achieved. This may be attributable to the larger effective surface area. The USPLD technique also conferred several other advantages in electrode fabrication. Firstly, the USPLD films displayed high adhesion to different substrates such as PI, SS and Si. Secondly, the samples containing heat-sensitive materials, such as photoresists, could be kept normally very close to room temperature, thus avoiding hardening or flow during deposition. This meant that it was possible to achieve both high-quality films and easy lift-off. Thirdly, it was possible to vary deposition parameters easily to obtain Pt films ranging from smooth (Ra < 1 nm) to relatively rough (Ra = 100 nm) meaning that optimal results could be obtained for each purpose. Fourthly, the biocompatibility of Pt films is guaranteed because high purity solid source materials are directly transferred from the target to the sample. It is desirable that the electrode surface should also support attachment and growth of neurons. It has been reported that thin Pt films are non-toxic and can support the attachment and the growth of cortical rat neurons (Thanawala et al. 2007). It has been also shown that electrode surface morphology has a major impact on neurocompatibility (Turner et al. 2000). Enhanced functions of neurons have been demonstrated on biomaterials structured at the nanoscale (Bayliss et al. 1999, Raffa et al. 2007, Kelly et al. 2008). Furthermore, it has been reported that astrocyte adhesion declines whereas the extension of neurons increases on nanoporous Si surfaces (Moxon et al. 2004). There is decreased adhesion as well as reduced proliferation and poorer functional activation of astrocytes on carbon nanofiber materials (diameter 60 nm) compared to conventional carbon fibres (diameters 125-200 nm) (McKenzie et al. 2004). These results provide clear evidence that materials structured at the nanoscale may have promising neural applications, because they increase the effective surface area and interact positively with neurons,

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while they simultaneously reduce the functions of astrocytes, leading to decreased glial scar formation (Kotov et al. 2009). The vast majority of the tissue contact in the neural interfaces occurs under the substrate/encapsulation material. Thus, this material has to fulfill high demands in order to allow the stable, long-term function of the implant. An optimal insulating material should be biocompatible, biostable and still possess appropriate dielectric properties. Although PI was successfully used in the present arrays as well as in several other neural interfaces developed by other research groups (Cheung et al. 2007, Stieglitz et al. 2000, Takahashi et al. 2003, Rousche et al. 2001), some degree of uncertainty is related to the relatively large water uptake, which is typically in the range of 0.5 – 4% (Stieglitz et al. 2000). In this context, one important advantage of the USPLD method compared to several other methods is that it is possible to create dielectric films that do not contain microparticles or droplets, which would cause pinholes in the films. This is due to the fact that the laser ablation pulses are so short that the energy penetrates into only a very shallow (tens of nanometers) surface layer, which is effectively converted to plasma and the target surface remains smooth without any flows or deteriorations. These coatings do not seem to contain any microdefects which could lead to short circuiting, especially when they come into contact with body fluids. Alumina and carbon nitride coatings deposited by the USPLD technique were demonstrated to be cell-friendly materials at least to fibroblastic (BHK-21) cells. The ultrasmoothness of these depositions might provide an additional benefit by avoiding the activation of glial cells and the formation of scar tissue.

9.3 MICROPATTERNED BIOMATERIAL SURFACES

In studies V and VI, photolithographic micropatterning of biomaterial surfaces was used for the production of model surfaces with accurately defined features. These kinds of model surfaces were demonstrated to be powerful tools for

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investigating the effects of different surface on the adhesion, spreading and morphology of cells and cellular communities. Different pattern sizes (5-125 µm), shapes (rectangular or circular) and surface chemistries (Ta, Ti, Cr, DLC) were utilized in the creation of engineered surfaces on silicon wafers. The optimized photolithography process with correct baking, exposure and development parameters allowed the formation of precisely defined photoresist features with an undercut edge profile. The shape of profile was observed to be extremely important for straightforward lift-off after deposition. The surface roughness, a parameter widely known to modulate protein adsorption and cellular responses, could be kept almost constant with all coatings. The AFM characterization revealed that variation in roughness between sample types was below 2 nm and the only relevant topographical cues of the samples were the steps between the patterns and background. By precise adjustment of the deposition time with two different types of PVD methods, i.e. magnetron sputtering for metals and filtered pulsed plasma arc discharge for DLC, it was possible to ensure that the coating thickness variations were kept as low as possible (200 ± 20 nm). This indicates that the possible clear-cut differences observed between different materials and patterns could not have been due to differences in surface topography and instead suggests that the differences seen and discussed below originated from the different chemical compositions and physical patterns of the materials being studied. Silicon is the most widely used substrate material in microchips and MEMS devices. However, biocompatibility properties of Si are often unsuitable for use in bio-MEMS devices. In studies V and VI, the cytocompatibility enhancement of Si was achieved using its partial coating with more biocompatible materials such as Ti or DLC. This suggests that even partial coating of Si-based bio-MEMS devices can be used to enhance their cytocompatibility and to facilitate their integration with the host during the critical initial phase of integration. The rapid integration of host cells also reduces the risk of implant-related infections according to the concept of the

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‚race for the surface‛ (Gristina 1987). In cellular experiments with both SaOS-2 and hMSC cells, a clear cellular preference for the biomaterial patterns over background Si was observed. Both cell types initially tended to adopt the geometrical shape of the patterns, resulting in ‚engineered‛ square or circle shaped cellular communities. The results from SaOS-2 cell experiments with different pattern sizes disclosed that the size and spacing of the patterns provide further options to guide the cells. The size of the SaOS-2 cells growing on the same Si background varied significantly, depending on the type of biomaterial used for patterning. This is a very interesting observation when one takes into account the fact that all surfaces were produced using the same fabrication methods on the same background Si. Another interesting observation was made in the focal adhesion staining experiments. Preferential localization of vinculin containing focal adhesions in SaOS-2 cells was clearly observed on the biomaterial patterns, but hMSCs did not show any such preference for the patterns. One could speculate that the patterns used in the hMSC experiments (75 µm squares) were too small to allow proper organization of the actin cytoskeleton and subsequent preferential forming of vinculin-containing adhesion plaques. Moreover, it has been demonstrated that the pattern size has an effect on the differentiation of hMSCs (McBeath et al. 2004) and in that case, relatively large islands are needed to ensure osteogenesis. The patterning techniques developed in this thesis provide an effective way to enhance the biocompatibility of Si-based microdevices. To allow proper functionality of these bio-MEMS devices, only the areas which are not under the integrated electronic components may be coated with biocompatible thin films. These patterning techniques also provide an effective means to guide the attachment, size and the shape of the cells as well as to regulate the differentiation of MSCs into the desired phenotypes. It is hoped that these aspects will stimulate the design of even more advanced biomaterials and constructs for tissue engineering and regenerative medicine.

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10 Conclusions In this work, novel flexible epidural microelectrode arrays were successfully developed and tested to permit recordings of acute and chronic evoked potentials in an animal model. In addition, novel materials and surface textures at a micro- and nanoscales were investigated in vitro aiming to improve the biological and electrochemical characteristics of neural interfaces. The knowledge about cell-biomaterial interactions obtained in this thesis work can be used in the further development of neural interfaces as well as in other implantable biomedical devices. On the basis of the present series of studies, the following detailed conclusions can be drawn: 1. The developed flexible microelectrode arrays possess adequate mechanical and electrochemical characteristics and they are capable of acquiring signals with high signal-tonoise ratios in evoked potential recordings when implanted onto the surface of the rat cortex. 2. The polyimide-platinum-polyimide sandwich structure, in which the flexibility of PI and the superior electrochemical properties of Pt were combined, was observed to be the most suitable realization of the array. A novel finding was the confirmation of the biocompatibility of the photosensitive polyimide used in the developed arrays. 3. Nanostructures on an electrode surface can be used to increase the effective surface area and to lower electrode impedance as observed with nanotextured Pt surfaces created with an ultra-short pulsed laser deposition technique. 4. Thin film dielectrics, alumina and carbon nitride produced by the USPLD method were proved to be highly cytocompatible in in vitro studies with fibroblasts. However,

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further in vitro and in vivo studies will be needed to clarify their long-term effects. 5. Micropatterning with biomaterials can be used to enhance the cytocompatibility of silicon-based bio-MEMS devices. Not only the chemical composition of the materials, but also the shape, edges (height) and size of the features used for surface patterning have a remarkable effect on cell guidance. 6. The morphology of cells cultured on micropatterned surfaces seemed to be aligned along the edges of the features e.g. forming round- or square-shaped cell morphologies having an interesting effect of the differentiation of stem cells into certain phenotypes. These results may be particularly useful in applications involving regenerative medicine and tissue engineering where cells are grown on biomaterial scaffolds.

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161

Sami Myllymaa Novel Micro- and Nano-technological Approaches for Improving the Performance of Implantable Biomedical Devices

Recent advances in micro- and nanotechnology offer a great opportunity to develop intelligent biomaterials and the next generation of implantable devices for diagnostics, therapeutics, and tissue engineering. This dissertation is focusing on the development of novel polymer-based microelectrode arrays suitable for use in intracranial electroencephalographic recordings. Moreover, the performances of novel thin film materials and their surface modifications at micro- and nanoscales were studied with physicochemical and cellular experiments in order to devise new solutions for further development of biomedical microdevices.

Publications of the University of Eastern Finland Dissertations in Forestry and Natural Sciences isbn 978-952-61-0216-0