I would like to thank Dr. Alex Chou, for maintaining all the equipments in ...... Hucknall A, Simnick AJ, Hill RT, Chilkoti A, Garcia A, Johannes MS, et al. Versatile.
iii LENGTH-SCALE EFFECTS ON THE DIFFERENTIAL ADHESION OF BACTERIA AND MAMMALIAN CELLS
ABSTRACT
The competition between microbial and mammalian cells to colonize a biomaterials surface begins immediately after insertion of a biomaterials implant or device in the human body. If mammalian cells win this competition, the result will be a healthy cellular layer covering the implant surface which is less vulnerable to bacterial virulence. However, biomaterials associated infection (BAI) can occur when bacteria win this competition and surface colonization occurs. The result can be significant patient morbidity and mortality. The biofilm mode of growth after certain amount of initial colonization can largely decrease the effectiveness of antibiotics. Minimizing the initial microbial adhesion is the first step to inhibit biofilm formation on and pathogenesis around a biomaterials implant. Nowadays biomaterials coatings are mainly developed to inhibit microbial adhesion, but often these only allow very little or no tissue integration. No biomaterials or functional coating exists that can resist microbial biofilm formation and support mammalian cell adhesion at the same time. This thesis aims to investigate to what extent spatially modified biomaterial surfaces, consisting of non‐adhesive hydrogel patterns, can favorably inhibit microbial adhesion, promote mammalian cell adhesion and spreading, and ultimately influence the race for the surface to reduce the chance of occurrence of biomaterials‐associated‐infection. We develop precise micro/nano-structured features using electron-beam patterning. We explore the properties of poly (ethylene glycol) (PEG) as negative electron beam resist for biointeractive applications, particularly as a means to spatially control protein adsorption and cell adhesion. The subsequent adhesion of bacteria and osteoblast-like cells are thus controlled with sub-micron precision in
iv flow-cell based on in vitro experimental models. We test length-scale modulated surfaces with different surface conditions (different adhesive area fraction, surface protein pre-treatment), and expose these to different microbial and mammalian cell strains (osteoblast-like cell and macrophages). A precise spatial controlled patterned surface could serve as a potential model to solve the general and fundamental problem of how to create surfaces that differentially interact with different cell types in biomaterials science. Moreover, when optimal dimensions of such patterns have been established with respect to the different biological processes occurring at a biomaterials interface in the human body, the concepts may become directly translatable to clinical practice.
Author: Yi Wang Advisor: Matthew Libera Date: April 22, 2013 Department: Chemical Engineering and Materials Science Degree: Doctor of Philosophy
v To my parents Xia Gan & De’an Wang
vi
Acknowledgements
My dissertation would not have been possible without all my supervisors. I want to specially thank my advisor, Prof. Matthew Libera for not only being my supervisor, but also encouraging and challenging me throughout my academic program. His creativity and persistence on research will always be inspiring to my professional career. I would also like to thank him for introducing me to various open resources especially the Center for Functional Nanomaterials in Brookhaven National Laboratory, NY, USA and the University Medical Center Groningen (UMCG), the Netherlands. Special gratitude is extended to my dissertation committee members: Professor Henk J. Busscher, Professor Xiaojun Yu and Professor Svetlana Sukhishvili. Your valuable time, constructive criticisms, and insights are faithfully appreciated. Ever since I got involved in the dual-PhD program between Stevens and UMCG, I owe my gratitude greatly to Prof. dr. Henk J. Busscher and Prof. dr. H.C. Van der Mei. Henk affected me with his passion both for research and for a fulfilling life. And I will always be grateful to Henny for her professional support on my thesis work. Thank you, Dr. Guruprakash Subbiahdoss, for sharing me with various experimental methods. And I cannot express my gratitude enough for all the great moments of sharing fun and frustrations together with all my colleagues and friends, both from Hoboken and Groningen. Thank you, Xianglin, Chris, Crystal, Marissa and Helen, thank you for always being there for me, listen to me and encourage me. I would like to thank Dr. Alex Chou, for maintaining all the equipments in the best way and also for being there patiently to guide me throughout. I enjoyed working with Xiaoguang, Yong, Emre and Yulong; they came to our lab of multiscale imaging later than I did, but I learnt a lot from them, both from their scientific way of thinking and also
vii from the smart way of playing board games. I always feel myself lucky and also I truly enjoyed working with such a professional team. It would not have been possible to write this doctoral thesis without the help and support of the kind people around me, to only some of whom it is possible to give particular mention here. Thank you Aaron, Hong, Qing, Joana, Xuening, Doctor Bu, Dr. Li Mei, Lei Song, Chongxia, Katya, Wenwen, for being my friends! Finally, and above all, I would like to thank my parents for giving me their unequivocal support, even merely by asking the date of my return. My small expression of thanks, as usual, does not suffice when facing this unconditional love. But I do know that the place where they are is my real and only home.
viii ABSTRACT ................................................................................................................................... iii List of Tables ............................................................................................................................... xiii List of Figures.............................................................................................................................xiiii Chapter 1 Introduction ................................................................................................................ 1 Mono-Functional to Multi-Functional Coatings: The Changing Paradigm to Control Biofilm Formation on Biomedical Implants (Book chapter accepted in "Biofilms in Bioengineering" Yi Wang, Matthew Libera, Henk J. Busscher, Henny C. van der Mei*) Chapter Summary ........................................................................................................................ 1 1.1 Introduction ............................................................................................................................ 1 1.2 Mono-Functional Coatings .................................................................................................... 5 1.3 Multi-Functional Coatings: A Shift in Paradigm ................................................................... 6 1.4 Selected Examples ................................................................................................................. 7 1.4.1 The Anti-adhesiveness of PEG-based Coatings .......................................................... .7 1.4.2 PEG-based Anti-adhesive Coating with an Antimicrobial Functionality .................... 8 1.4.3 PEG-based Anti-adhesive Coating with RGD-containing Peptides ............................ 8 1.4.4 Differentially Adhesive Surface ................................................................................... 9 1.5 Conclusion ........................................................................................................................... 10 References .................................................................................................................................. 11 Chapter 2 Poly(ethylene glycol) as a Biointeractive Electron-Beam Resist .......................... 19 (About to Submit to the Journal of Polymer Science: Polymer Physics Yi Wang, Emre Firlar, Xiaoguang Dai, Matthew Libera*) Chapter Summary ...................................................................................................................... 19 2.1 Introduction .......................................................................................................................... 20 2.2 Experimental and Modelling Procedures ............................................................................. 21 2.2.1 E-beam Patterning ...................................................................................................... 21 2.2.2 Fibronectin Adsorption .............................................................................................. 23 2.2.3 Morphological Analysis of PEG Microgel ................................................................ 24 2.2.4 Monte-Carlo Simulations ........................................................................................... 24 2.3 Results and Discussion ........................................................................................................ 26 2.3.1 Microgel-Patterned Surfaces ....................................................................................... 26 2.3.2 Monte Carlo Modelling of Electron-PEG Interactions ............................................... 30 2.3.3 Selecting Irradiation Conditions for Microgel Patterning ........................................... 40 2.4 Conclusion ........................................................................................................................... 44
ix References .................................................................................................................................. 46 Chapter 3 Length-Scale Mediated Differential Adhesion of Mammalian Cells and Microbes.................................................................................................................................... 50 (Advanced Functional Materials 2011; 21; 3916-23 Yi Wang, Guruprakash Subbiahdoss, Jan Swartjes, Henny C. van der Mei, Henk J. Busscher, and Matthew Libera*) Chapter Summary ...................................................................................................................... 50 3.1 Introduction .......................................................................................................................... 50 3.2 Experimental Section ........................................................................................................... 53 3.2.1 PEG Hydrogel Surface Patterning .............................................................................. 53 3.2.2 Experiments with Bacteria and Yeast ......................................................................... 54 3.2.3 Mammalian Cell Experiments .................................................................................... 55 3.2.4 AFM Force Measurement ........................................................................................... 57 3.3 Results and Discussion ........................................................................................................ 58 3.3.1 Surface-Patterned PEG Gels ....................................................................................... 58 3.3.2 Microbial Adhesion .................................................................................................... 60 3.3.3 Mammalian Cell Adhesion and Spreading ................................................................. 63 3.3.4 Length-Scale Mediated Differential Adhesion ........................................................... 67 References ................................................................................................................................. 69 Chapter 4 Effect of Adsorbed Fibronectin on the Differential Adhesion of Osteoblast-like Cells and Staphylococcus aureus with and without Fibronectin-Binding Proteins ............ 74 (Biofouling 2012; 28; 1011-1021 Yi. Wang, Guruprakash Subbiahdoss, Joop. de Vries, Matthew Libera *, Henny C. van der Mei, Henk J. Busscher) Chapter Summary ...................................................................................................................... 74 4.1 Introduction .......................................................................................................................... 74 4.2 Experimental Section ........................................................................................................... 77 4.2.1 PEG-Microgel Surface Patterning .............................................................................. 77 4.2.2 Fibronectin Pre-adsorption on PEG-Microgel-Patterns .............................................. 78 4.2.3 Bacterial Strains, Growth and Harvesting................................................................... 78 4.2.4 Atomic Force Microscopy .......................................................................................... 79 4.2.5 S.aureus Adhesion, Osteoblast-like Cells Adhesion ................................................... 80 4.4 Experimental Results ........................................................................................................... 82 4.4.1 Fn Adsorption on Patterned Surfaces.......................................................................... 82 4.4.2 S.aureus Adhesion Forces ........................................................................................... 82 4.4.3 S.aureus Adhesion to PEG-Microgel-Patterned Surfaces ........................................... 83
x 4.4.4 Mammalian Cell Adhesion and Spreading ................................................................. 83 4.5 Discussion and Conclusions ................................................................................................ 85 References ................................................................................................................................. 95 Chapter 5 Conditions of Lateral Surface Confinement that Favor Tissue-Cell Integration over Biofilm Growth .............................................................................................................. 101 (In preparation for Biomateirals Yi Wang, Joana F da Silva Domingues, Henny C. van der Mei, Henk J. Busscher, Matthew Libera*) Chapter Summary .................................................................................................................... 101 5.1 Introduction ........................................................................................................................ 102 5.2 Materials and Methods....................................................................................................... 105 5.2.1 Microgel Patterning .................................................................................................. 105 5.2.2 Fibronectin Adsorption ............................................................................................. 106 5.2.3 Bacterial Strain, Growth and Harvesting .................................................................. 107 5.2.4 Staphylococcal Adhesion and Growth ...................................................................... 107 5.2.5 Mammalian Cell Culature, Harvesting and Adhesion .............................................. 108 5.3 Results ................................................................................................................................ 110 5.3.1 Patterned Surfaces ..................................................................................................... 110 5.3.2 Mammalian Cell Interactions .................................................................................... 110 5.3.3 Bacterial Interactions ................................................................................................ 113 5.4 Discussion .......................................................................................................................... 118 5.5 Conditions that Promote Tissue-Cell Interactions and Inhibit Bacterial Colonization ...... 124 5.6 Conclusion ......................................................................................................................... 125 References ............................................................................................................................... 128 Chapter 6 In Vitro Interactions between Bacteria and Macrophages on PEG-Microgel Patterned Adhesive Patches in the Pathogenesis of Biomaterial-Associated Infection ... 134 (In preparation. Joana F. da Silva Domingues, Yi Wang, Matthew Libera, Guruprakash Subbiahdoss, Henny C. van der Mei, Henk J. Busscher.) 6.1 Introduction ........................................................................................................................ 134 6.2 Experimental Section ......................................................................................................... 136 6.2.1 PEG-Patterned Adhesive Patches ............................................................................. 136 6.2.2 Bacterial Strain, Growth and Harvesting .................................................................. 137 6.2.3 Macrophages Culturing and Harvesting ................................................................... 137 6.2.4 Staphylococcal Adhesion and Biofilm Growth on PEG-Patterned Adhesive Patches ................................................................................................................................................. 138
xi 6.2.5 Macrophages Interaction with Staphylococcal Biofilm on PEG-Patterns with Adhesive Patches ............................................................................................................... 138 6.3 Results and Discussion ...................................................................................................... 139 References ............................................................................................................................... 143 Chapter 7 Thesis Discussion .................................................................................................... 146 References ............................................................................................................................... 154 Vita.............................................................................................................................................. 156
xii List of Tables Table 1.1 Different mono-functional coatings that influence bacterial adhesion and biofilm formation together with some of their advantages and disadvantages. ........................................... 5 Table 2.1 Microgel diameters (FWTM) for various exposures...................................................... 41 Table 3.1 Maximum adhesion force of S. aureus on patterned glass surfaces with different intergel spacing ..................................................................................................................................... 60 Table 3.2 The initial deposition rates on unpatterned glass of the different microbial strains, together with Weibull scale (α) and shape (β) factors describing initial deposition rates on patterned surfaces as a function of inter-gel spacing. .............................................................. 63 Table 4.1 Initial S. aureus 8325-4 and S.aureus DU 5883 adhesion rates on silanized glass in the absence (-Fn) and presence (+Fn) of adsorbed Fn at a shear rate of 11 s-1 ................................. 85 Table 6.1 Adhesion and growth of S.aureus 8325-4 on different surfaces before and after exposure to J774A.1 macrophages .............................................................................................. 140
xiii
List of Figures Figure 1.1 Schematic presentation of a monolayer PEG coating: mushroom conformation (left)and brush conformation (right) ................................................................................................ 7 Figure 1.2 Schematic presentation of modified PEO polymer chains combining microbial repellency with cell-adhesive functinal groups................................................................................ 8 Figure 1.3 Scanning electron microscope image of a mammalian cell spreading on a PEGmicrogel patterned surface with an inter-gel spacing of 1.5µm. ..................................................... 9 Figure 2.1 Schematic description of the electron-beam lithography process................................ 25 Figure 2.2 SEM images of PEG microgels patterned on silanized Si surfaces ............................. 26 Figure 2.3 Different thickness of PEG homopolymer films can be generated by spin coating solutions of varying PEG concentration. ....................................................................................... 27 Figure 2.4 AFM images of PEG microgels pattenred on Si surfaces ............................................ 28 Figure 2.5 Confocal immunofluoresce images of surfaces patterned by PEG microgels ............. 29 Figure 2.6 The interaction volume of 30 keV electrons with a thin film of PEG on Si substrate .. 31 Figure 2.7 Monte Carlo calculations of the trajecories of 1000 electrons for different incident electron energies ............................................................................................................................ 32 Figure 2.8 Monte Carlo simulatins of deposited energy per unit volume ..................................... 33 Figure 2.9 Monte Carlo simulatin of fibronectin adsorption on patterned PEG surfaces ............ 38 Figure 3.1 Schematic presentation of cell and microbes adhering on different biomaterial surfaces .......................................................................................................................................... 52 Figure 3.2 Surface-patterned PEG microgels modulate the adhesion force experienced by an individual S. aureus bacterium. ..................................................................................................... 59 Figure 3.3 Microbial deposition rate decreases as inter-gel spacing approaches microbial dimensions. .................................................................................................................................... 61
xiv Figure 3.4 Osteoblast-like U-2 OS cells are able to adhere and spread on microgel-patterned surfaces with an inter-gel spacing of 1.0 μm or more. ................................................................ 64 Figure 3.5 SEM images of osteoblast-like cells adhere and spread on patterned surfaces .......... 64 Figure 3.6 Osteoblast-like cells adhere to and spread on surfaces with δ ≥ 1 µm........................ 66 Figure 4.1 Fn adsorption on PEG-microgel-patterned surfaces with inter-gel spacings of 0.5 μm and 2 μm......................................................................................................................................... 86 Figure 4.2 Adhesion forces between S. aureus 8325-4 and DU 5883 and microgel-patterned surfaces with different inter-gel spacings in the absence and presence of adsorbed Fn .............. 87 Figure 4.3 Number of S. aureus 8325-4 and S. aureus DU 5883 adhering to microgel-patterned surfaces with different inter-gel spacings in absence and presence of adsorbed Fn as a function of time................................................................................................................................................. 89 Figure 4.4 Phase-contrast and fluorescent micrographs of U-2 OS
cells adhering to and
spreading on microgel-pattern surfaces with different inter-gel spacings in the absence and presence of adsorbed Fn. ............................................................................................................... 90 Figure 4.5 (a)U-2 OS cells surface coverage after 1.5h seeding and after 48h of growth on microgel-patterned surfaces in the presence an absence of adsorbed Fn; (b) Fractional change in U-2 OS surface coverage from its initial value at 1.5 h to its value after 48 h of growth in the absence and presence of pre-adsorbed Fn. ................................................................................... 91 Figure 4.6 SEM micrograph of U-2 OS cells grow on PEG microgel patterned surface with 1μm intergel-spacing after 48 h of culture ............................................................................................ 93 Figure 5.1 The morphology of patterned surfaces with adhesive patches of 1 μm and 5 μm diameters at different adhesive area fractions ............................................................................. 109
xv Figure 5.2 U-2 OS cell surface coverage after 1.5 h on fibronectin adsorbed PEG-microgel patterns consisting of adhesive patches with different adhesive patch diameters and adhesive area fractions. ...................................................................................................................................... 110 Figure 5.3 U-2 OS cell adhesion after 1.5 h and after 48 h of growth on fibronectin adsorbed PEG-microgel patterns ................................................................................................................ 111 Figure 5.4 Representative time-resolved phase contrast imges show S.aureus deposition and growth .......................................................................................................................................... 113 Figure 5.5 The number of adhering S. aureus 8325-4 after 0.5 h (a) and 5 h (b)on fibronectin coated PEG-microgel patterns consisting of adhesive patches with different diameters of the adhesive areas and inter-patch distances .................................................................................... 115 Figure 5.6 Time-resolved phase –contrast images of S. aureus colonies developing on micropatterned surfaces............................................................................................................... 116 Figure 5.7 The number of bacteria within an individual colony increases with time, and the colony growth rate for micropatterned surfaces and unpatterned surface.................................. 118 Figure 5.8 SEM images of S.aureus biofilm and U-2 OS cell adhesion and spreading .............. 123 Figure 6.1 Number of adhering S. aureus 8325-4 on glass and different PEG-gel patterned surfaces ....................................................................................................................................... 139 Figure 6.2 Phase-contrast images of macrophage migration towards S. aureus and phagocytosis on PEG-patterned surfaces .......................................................................................................... 142 Figure 7.1 AFM images of PEG-microgel formed by focused-beam exposure with an electron dose of 10 fC and beam energy of 2 keV ...................................................................................... 148 Figure 7.2 Phase-contrast images of U-2 OS cell adhesion, spreading and growth at different time points in the presence of S. aureus 8325-4 on different PEG-microgel patterns in the presence of pre-adsorbed Fn. ...................................................................................................... 148
xvi Figure 7.3 Phase-contrast images of U-2 OS cell adhesion, spreading and growth after 48 h in the presence of S. epidermidis ATCC 35983 on PEG-microgel patterns with inter-gel spacings from 0.5 to 1.5 m. ....................................................................................................................... 151 Figure 7.4 SEM iamge of a DRG neurite grow on an e-beam patterned PEG gel arrays (A) and SEM imgae of U-2 OS cell and s.epidermidis co-culture on a patterned surface ....................... 153
1
Chapter 1
Introduction
Mono-Functional to Multi-Functional Coatings: the Changing Paradigm to Control Biofilm Formation on Biomedical Implants
Chapter Summary
Infection is the number one cause of failure of biomaterials implants and devices, despite decades of research into the development of anti-adhesive coatings. In this chapter we postulate that whereas anti-adhesive coatings may be of value for particular applications, such as urinary or intravenous catheters, contact lens cases, and voice prostheses, the paradigm underlying the ideal biomaterial for biofilm control on totally internal, permanent implants needs to change if we want to effectively reduce the occurrence of biomaterials-associated infections. Rather than focusing on the design of mono-functional coatings, multi-functional coatings need to be developed. These must include both anti-adhesive and anti-microbial functionalities, promote a proper immune response, and stimulate tissue integration while simultaneously reducing the threat of bacterial colonization of an implant surface.
1.1 Introduction
Biomaterials-associated infection (BAI) of medical implants and devices is a serious complication that often occurs after an otherwise technically successful surgical procedure or after the prolonged use of contact lenses, catheters, or other percutaneous devices. BAI of totally
Yi Wang, Matthew Libera, Henk J. Busscher, and Henny C. van der Mei*; Book Chapter accepted in “Biofilms in Bioengineering”.
2
internal implants nearly always results in the surgical removal and replacement of the implant, because antibiotics are not able to effectively kill bacteria in their biofilm mode of growth such as that typical on an implant surface. Similarly, the removal of an infected, temporary percutaneous device, such as a central intravenous catheter, poses a clinical dilemma. Removal often is synonymous with interruption of a desired therapy, and this can often have disastrous consequences in cases such as chemotherapies. On the other hand, not controlling the infection can also evolve into a life-threatening situation.
Over the past decades our knowledge of how bacteria adhere to surfaces and develop into biofilms has grown rapidly. Substratum surface properties like charge [1-3], hydrophobicity [4-6], and roughness [7-9] are now known to be major determinants of bacterial adhesion. Positively charged surfaces usually stimulate bacterial adhesion, because most bacterial strains carry a net negative cell surface charge [10]. However, at the same time, substrata carrying a positive surface charge through immobilization of, for example, quaternary ammonium compounds are known to become bactericidal upon contact [11, 12]. Acid-base interactions form the basis for the hydrophobicity of bacterial cell and substratum surfaces [13]. They also play a role in bacterial adhesion and, depending on the environmental conditions, can increase or decrease bacterial adhesion. Often hydrophobic surfaces attract equal numbers of adhering bacteria as do hydrophilic surfaces, but they appear free of biofilm after several hours or days under fluctuating environmental shear conditions [14]. Similarly, surface roughness has an effect on bacterial adhesion, but this depends on the scale of the roughness. Some studies have demonstrated that roughness on a scale much smaller than bacterial dimensions does not influence adhesion [15, 16], but other studies have shown that even nanometer-scale roughness may impact bacterial adhesion [16-18]. Interestingly, bacteria do not adhere preferentially on surface features such as
3
scratches, but they appear to have better opportunities to grow into a biofilm once they adhere to surface irregularities of bacterial or larger dimensions due to the fact that they are protected against environmental shear forces [19].
The body of literature on bacterial adhesion mechanisms has stimulated the development of anti-adhesive coatings for biomaterials. Whereas anti-adhesiveness may be a sufficient single functionality for a coating when applied on such devices as urinary or intravenous catheters, contact lens cases, and voice prostheses, an anti-adhesive functionality is not sufficient for totally internal, permanent implant coatings like vascular grafts, surgical meshes, and hip or knee prostheses. Tissue integration has been described to offer the best protection of such implants [20, 21]. However, many microbial strains known to cause BAI, adhere to the same serum- or plasmacoated surface receptor sites as do mammalian cells [22], and they can thus colonize unmodified implant surfaces. When a coating is applied to such a surface, however, the anti-adhesive functionality is typically not limited to microorganisms but also to mammalian cells. Hence, many anti-adhesive coatings, while successfully repelling bacteria also repel the desirable tissue cells.
Since anti-adhesiveness is never absolute and may involve at best a 2 or 3-log units reduction in the number of adherent bacteria, implant surfaces can still become contaminated during surgery and hospitalization or become colonized by blood-born organisms from infections elsewhere in the body when not fully tissue integrated. Furthermore, the natural defenses offered by macrophages are often frustrated in the presence of a biomaterial, and macrophages are thus less able to eradicate the infecting organisms. Consequently, in addition to anti-adhesive functionalities, antimicrobial functionalities are desirable attributes for an implant coating, as well
4
[23]. Antimicrobial functionalities can be introduced either by immobilizing antimicrobials on the surface yielding contact killing or by the application of antibiotic-release coatings.
Significantly, depending on the application, approximately 95% or more of all patients who receive biomaterial implants do not suffer adverse, long-term effects due to microbial interactions with the implant. However, the consequences when BAI does occur are severe. There can be substantial patient discomfort, complications that can lead to significant morbidity or mortality, significant stress on physicians and medical facilities, and a huge financial burden on the health care system. Clinically speaking, the major difference between revision patients receiving a secondary implant after BAI and primary implant patients receiving their first implant is that in revision surgery the tissue is compromised by infecting organisms [24]. Clearing the infection from surrounding tissue requires high local antibiotic concentrations, which can often be achieved by antibiotic-releasing beads, spacers, and sponges [25]. Thus, while antibiotic-release implant coatings can benefit primary orthopedic implants [26], such coatings are particular important for implants used in revision surgeries, and revision patients after BAI may require coatings with different functionalities than those used in a primary implant.
In this chapter we first briefly review past developments of biomaterials coatings and discuss their advantages and disadvantages. We argue that mono-functional coatings are applicable only for certain applications and the community must look beyond these for greater success in applications where an implant cannot be easily removed and where mammalian cells and microbial cells compete for surface colonization. Effective, infection-resisting permanent implants will require multi-functional surfaces that: (i) are not only anti-adhesive and antimicrobial but also stimulate tissue integration; (ii) do not hamper proper immune responses to
5
microbial presence, and (iii) in revision surgery after BAI, clear infecting organisms from the tissue surrounding the implant.
1.2 Mono-functional Coatings
Table 1 summarizes some of the many different mono-functional coatings and surfaces that have been designed over the past decades. There are clearly no coatings that can be expected to perform well in each and every application. The general concern with many of these coatings, especially those showing anti-adhesive functionality, is tissue integration.
Table 1.1 Different mono-functional coatings that influence bacterial adhesion and biofilm formation together with some of their advantages and disadvantages. Types of coatings Anti-adhesive
Basis Hydrophobic coatings
Hydrophilic coatings
Polymer brush coatings
Nano-patterned surfaces Tissue integrating
Fibronectin-based coatings
Immune-friendly
Polymer brush coatings Positively-charged surfaces
Antimicrobial
TiO2 photocatalytic surfaces
Advantages (+) and disadvantages (-) (+) reduced biofilm formation under fluctuating shear, depending on the strain involved (-) no tissue integration (+) reduced adhesion, depending on the strain involved (+) good mammalian cell interaction
Ref [27-29]
(+) very low bacterial adhesion (-) biofilms do develop ultimately, though weakly adhering (-) no tissue integration (+) reduced biofilm formation (+) easy to clean
[33-38]
(+) good adhesion and spreading of tissue cells (-) many bacterial strains use the same receptor-sites for their adhesion as tissue cells (+) allow macrophage mobility on a surface (-) no tissue integration (+) kill adhering organisms upon contact (-) layer of dead bacteria is left for subsequent adhesion (+) kill adhering bacteria after activation (+) can be re-activated depending on the application
[42-44]
[30-32]
[39-41]
[45-48] [49, 50]
[51-54]
6
Antibiotic-releasing
Antibiotic-loaded cements
Antibiotic-loaded cements for beads, spacers and sponges Antibiotic-releasing coatings
(+) high burst release of antibiotics in surrounding tissue (-) long-term, low concentration tail-release may induce antibiotic resistance (+) high burst release of antibiotics in surrounding tissue (+) low tail release and short stay in the human body (+) high burst release of antibiotics in surrounding tissue (-) no tail-release due to limited reservoir volume
[55]
[56]
[50, 57, 58]
1.3 Multi-functional Coatings: A Shift in Paradigm
Baier in 1982 [59] suggested that scientists and engineers should take lessons from nature when designing infection-resisting biomaterials coatings. Unfortunately, his approach was not widely adopted at that time, possibly because Baier defined the ideal biomaterials surface predominantly in terms of the so-called critical surface tension. This concept has since been recognized as too simple to cover all functionalities of the natural endothelium, designed by nature to withstand colonization by infecting organisms. Following Baier and taking lessons from nature, we postulate that the ideal biomaterials coating should possess multiple functionalities. The design of such multifunctional surfaces clearly requires a change in paradigm, abandoning the common concept of mono-functionality.
In the next section of this chapter, we briefly describe several examples of bi-functional coatings. Then, once the concept of multifunctional coatings is adopted, the number of possibilities of how to differentially control to surface interactions with mammalian cells and with infecting pathogens becomes huge.
7
1.4 Selected Examples 1.4.1 The Anti-adhesiveness of PEG-based Coatings Poly(ethylene glycol) (PEG)-based coatings are highly hydrated gel-like structures that constitute the most anti-adhesive surfaces hitherto known [60]. Often alternatively described as PEO for poly(ethylene oxide), PEG chains attached to a flat surface can exist in two different conformations, the so called “mushroom structure” at low chain grafting densities and the “brush structure” at higher chain grafting densities, where the chains are forced to stretch in the medium (see Figure 1.1). Both structures are expected to create a barrier between the microorganism and the surface thereby preventing adhesion.
Figure 1.1 Schematic presentation of a monolayer PEG coating: mushroom conformation (left) and brush conformation (right).
Since PEG-based coatings have superior anti-adhesive properties and, importantly, also allow for enhanced macrophage mobility, they are increasingly being used as a starting point for adding other functionalities. Among these are antimicrobial character and/or tissue-integrating factors. In the design of a multifunctional PEG-based coating, the difference in dimensions between mammalian cells (average range from about 10 to 50 µm depending upon their degree of spreading) and bacteria (on average having a diameter of around 1 µm) needs to be taken into account. Antimicrobial or tissue-integrating functionalities must be applied in surface densities that do not negate the anti-adhesiveness of a multifunctional coating against the smaller bacteria.
8
Figure 1.2 Schematic presentation of modified PEO polymer chains combining microbial repellency with cell-adhesive functional groups. The distance between the functional groups for cell attachment critically determines whether microbial repellency will be maintained or not.
1.4.2 PEG-based Anti-adhesive Coating with an Antimicrobial Functionality PEG-based coatings on silicone rubber strongly reduce bacterial adhesion, but, despite 2-log units reduction in initial adhesion, significant numbers of biofilms can nevertheless form [61]. Conjugation of lysozyme to PEG-molecules and coating of silicone rubber surfaces with 68/32% PEG/lysozyme has yielded a bi-functional polymer brush, possessing both anti-adhesive activity due to the polymer brush combined with the antibacterial activity of lysozyme [62].
1.4.3 PEG-based Anti-adhesive Coating with RGD-containing Peptides The anti-adhesiveness of PEG-based coatings extends to mammalian cells. In order to prepare bifunctional coatings which prevent microbial adhesion while supporting tissue cell growth, a biologically inert poly(L-lysine)-graft-poly(ethylene glycol) (PLL-g-PEG) copolymer has been equipped with the arginine-glycine-aspartic acid (RGD) peptide sequence [63]. This RGD peptide is known as one of the major recognition sites of integrin receptors through which cells connect to their extracellular matrix [64]. Reduced bacterial adhesion on bi-functional PLL-g-PEG/PEGRGD-modified surfaces has been demonstrated separately from their ability to support tissue cell
9
growth [65, 66]. Significantly, when the “race for the surface” between staphylococci and osteoblasts was studied on such bi-functional coatings in a co-culture experiment [67], mammalian cells appeared to be much more at an advantage than on common biomaterials surfaces.
1.4.4 Differentially Adhesive Surface PEG-based submicron-sized, non-adhesive microgels patterned on an otherwise cell-adhesive surface have recently been described as another example of a bi-functional, anti-adhesive and tissue-integrating coating. Such a structure creates a surfaces which is largely cell-adhesive but contains non-adhesive features, which, when spaced at distances comparable to bacterial dimensions, yield dramatically reduced staphylococcal deposition rates [68]. This has been attributed to the fact that staphylococci have relatively rigid cell walls that cannot easily conform to the substrate under conditions of modulated adhesiveness. In contrast, mammalian cells have flexible cell membranes and adhere to extracellular structures via transmembrane integrins. Mammalian cells can consequently adhere to surfaces with modulated cell adhesiveness by adhering to adhesive patches in between non-adhesive structures (Figure 1.3).
Figure 1.3 Scanning electron microscope image of a mammalian cell spreading on a PEGmicrogel patterned surface with an inter-gel spacing of 1.5 μm. Dotted structures are PEG-
10
microgels, while extensions from the cell circumference can be seen to grow over the microgels in search of the adhesive surface sites. The bar represents 1.5 μm.
1.5 Conclusion A great variety of coatings have been described in the literature and claim to be promising for the prevention of biomaterials-associated infection. Most are based on reducing bacterial adhesion and the subsequent formation of surface-attached biofilms. While such coatings are useful in certain applications, there has been a general lack of clinically favorable outcomes when using these coatings on internal biomedical devices such as implanted prostheses. Applications such as these, which are meant to be both totally internal and permanent, require more cleverly designed coatings possessing multiple functionalities. Importantly, this approach represents a significant paradigm change from one where coatings are designed to reduce infection by minimizing all biological interaction to one where interactions with desirable tissue cells are amplified while bacteria are simultaneously killed or blocked from the surface. Such an approach will require a far better understanding of how multiple cell types interact with synthetic surfaces together with a far greater level of creativity about how to design surfaces with differential cell-interactive properties.
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References 1.
Jones CR, Adams MR, Zhdan PA, Chamberlain AH. The role of surface physicochemical properties in determining the distribution of the autochthonous microflora in mineral water bottles. J Appl Microbiol 1999:86;917-27.
2.
Legeay G, Poncin-Epaillard F, Arciola CR. New surfaces with hydrophilic/hydrophobic characteristics in relation to (no)bioadhesion. Int J Artif Organs 2006;29:453-61.
3.
Gottenbos B, Van der Mei HC, Klatter F, Grijpma DW, Feijen J, Nieuwenhuis P, Busscher HJ. Positively charged biomaterials exert antimicrobial effects on Gram-negative bacilli in rats. Biomaterials 2003;24:2707-10.
4.
Andersen TE, Kingshott P, Palarasah Y, Benter M, Alei M, Kolmos HJ. A flow chamber assay for quantitative evaluation of bacterial surface colonization used to investigate the influence of temperature and surface hydrophilicity on the biofilm forming capacity of uropathogenic Escherichia coli. J Microbiol Methods 2010;81:135-40.
5.
Schwarz F, Sculean A, Wieland M, Horn N, Nuesry E, Bube C, Becker, J. Effects of hydrophilicity and microtopography of titanium implant surfaces on initial supragingival plaque biofilm formation. A pilot study. Mund Kiefer Gesichtschir 2007;11:333-8.
6.
Almaguer-Flores A, Olivares-Navarrete R, Wieland M, Ximenez-Fyvie LA, Schwartz Z, Boyan BD. Influence of topography and hydrophilicity on initial oral biofilm formation on microstructured titanium surfaces in vitro. Clin Oral Implants Res 2011;doi:101111/j.16000501.2011.02184.x.
7.
Wu Y, Zitelli JP, TenHuisen KS, Yu X, Libera MR. Differential response of Staphylococci and osteoblasts to varying titanium surface roughness. Biomaterials 2011;32:951-60.
8.
Tang L, Pillai S, Revsbech NP, Schramm A, Bischoff C, Meyer RL. Biofilm retention on surfaces with variable roughness and hydrophobicity. Biofouling 2011;27:111-21.
12
9.
Charman KM, Fernandez P, Loewy Z, Middleton AM. Attachment of Streptococcus oralis on acrylic substrates of varying roughness. Lett Appl Microbiol 2009;48:472-7.
10. Jucker BA, Harms H, Zehnder AJ. Adhesion of the positively charged bacterium Stenotrophomonas (Xanthomonas) maltophilia 70401 to glass and Teflon. J Bacteriol 1996;178:5472-9. 11. Tiller JC, Liao CJ, Lewis K, Klibanov AM. Designing surfaces that kill bacteria on contact. Proc Nat Acad Sci U S A; 2001;98:5981-5. 12. Majumdar P, Lee E, Patel N, Ward K, Stafslien SJ, Daniels J, Chisholm BJ, Boudjouk P, Callow ME, Callow JA, Thompson SE. Combinatorial materials research applied to the development of new surface coatings IX: an investigation of novel antifouling/foulingrelease coatings containing quaternary ammonium salt groups. Biofouling 2008;24(3):185200. 13. Van Oss CJ. Hydrophobicity of biosurfaces -- Origin, quantitative determination and interaction energies. Colloids Surf B: Biointerf 1995;5:91-110. 14. Hermannson M. The DLVO theory in microbial adhesion. Colloids Surf B:Biointerf 1999;14:105-19. 15. Li B, Logan BE. Bacterial adhesion to glass and metal-oxide surfaces. Colloids Surf B: Biointerf 2004;36:81-90. 16. Mitik-Dineva N, Wang J, Mocanasu RC, Stoddart PR, Crawford RJ, Ivanova EP. Impact of nano-topography on bacterial attachment. Biotechnol J 2008;3:536-44. 17. Mitik-Dineva N, Wang J, Truong VK, Stoddart PR, Malherbe F, Crawford RJ, Ivanova EP. Differences in colonisation of five marine bacteria on two types of glass surfaces. Biofouling 2009;25:621-31.
13
18. Emmerson IV RJ, Bergstrom TS, Liu Y, Soto ER, Brown CA, McGimpsey WG, Camesano TA. Microscale correlation between surface chemistry, texture, and the adhesive strength of Staphylococcus epidermidis. Langmuir 2006;22:11311-21. 19. Gristina AG. Biomaterial-centered infection: microbial adhesion versus tissue integration. Science 1987;237:1588-95. 20. Deacon JM, Pagliaro AJ, Zelicof SB, Horowitz HW. Prophylactic use of antibiotics for procedures after total joint replacement. J Bone Joint Surg Am 1996;78:1755-70. 21. Patti JM, Allen BL, McGavin MJ, Hook M. MSCRAMM-mediated adherence of microorganisms to host tissues. Annu Rev Microbiol 1994;48:585-617. 22. Boelens JJ, Dankert J, Murk JL, Weening JJ, Van der Poll T, Dingemans KP, Koole L, Laman JD, Zaat SA. Biomaterial-associated persistence of Staphylococcus epidermidis in pericatheter macrophages. J Infect Dis 2000;181:1337-49. 23. Broekhuizen CAN, Schultz MJ, Van der Wal AC, Boszhard L, De Boer L, VandenbrouckeGrauls CMJE, Zaat SAJ. Tissue around catheters is a niche for bacteria associated with medical device infection. Crit Care Med 2008;36:2395-402 . 24. Neut D, Van de Belt H, Van Horn JR, Van der Mei HC, Busscher HJ. Residual gentamicinrelease from antibiotic-loaded polymethylmethacrylate beads after 5 years of implantation. Biomaterials 2003;24:1829-31. 25. Engesaeter LB, Lie SA, Espehaug B, Furnes O, Vollset SE, Havelin LI. Antibiotic prophylaxis in total hip arthroplasty: effects of antibiotic prophylaxis systemically and in bone cement on the revision rate of 22,170 primary hip replacements followed 0-14 years in the Norwegian Arthroplasty Register. Acta Orthop Scand 2003;74:644-51. 26. Legeay G, Poncin-Epaillard F, Arciola CR. New surfaces with hydrophilic/hydrophobic characteristics in relation to (no)bioadhesion. Internat J Artif Organs 2006;29:453-61.
14
27. Harris LG, Mead L, Muller-Oberlander E, Richards RG. Bacteria and cell cytocompatibility studies on coated medical grade titanium surfaces. J Biomed Mater Res Part A 2006;78:50-8. 28. Karakecili AG, Gumusderelioglu M. Comparison of bacterial and tissue cell initial adhesion on hydrophilic/hydrophobic biomaterials. J Biomater Sci Polym Ed.2002;13:185-96. 29. Tenke P, Riedl CR, Jones GL, Williams GJ, Stickler D, Nagy E. Bacterial biofilm formation on urologic devices and heparin coating as preventive strategy. Int J Antimicrob Agents 2004;23:67-74. 30. An YH, Bradley J, Powers DL, Friedman RJ. The prevention of prosthetic infection using a cross-linked albumin coating in a rabbit model. J Bone Joint Surg Br 1997;79:816-9. 31. Wilson M, Harvey W. Prevention of bacterial adhesion to denture acrylic. J Dent 1989;17:166-70. 32. Pechey A, Elwood CN, Wignall GR, Dalsin JL, Lee BP, Vanjecek M, Welch I, Ko R, Razvi H, Cadieux PA.. Anti-adhesive coating and clearance of device associated uropathogenic Escherichia coli cystitis. J Urol 2009;182:1628-36. 33. Zoulalian V, Zurcher S, Tosatti S, Textor M, Monge S, Robin JJ. Self-assembly of poly(ethylene glycol)-poly(alkyl phosphonate) terpolymers on titanium oxide surfaces: synthesis, interface characterization, investigation of nonfouling properties, and long-term stability. Langmuir 2010;26:74-82. 34. Paulsson M, Kober M, Freij-Larsson C, Stollenwerk M, Wesslen B, Ljungh A. Adhesion of staphylococci to chemically modified and native polymers, and the influence of preadsorbed fibronectin, vitronectin and fibrinogen. Biomaterials 1993;14:845-53. 35. Sankar S, Rajalakshmi T. Application of poly ethylene glycol hydrogel to overcome latex urinary catheter related problems. Biofactors 2007;30:217-25.
15
36. Muller R, Ruhl S, Hiller KA, Schmalz G, Schweikl H. Adhesion of eukaryotic cells and Staphylococcus aureus to silicon model surfaces. J Biomed Mat Res Part A 2008;84:817-27. 37. Saldarriaga Fernandez IC, Van der Mei HC, Metzger S, Grainger DW, Engelsman AF, Nejadnik MR, Busscher HJ. In vitro and in vivo comparisons of staphylococcal biofilm formation on a cross-linked poly(ethylene glycol)-based polymer coating. Acta Biomater 2010;6:1119-24. 38. Hannig M, Kriener L, Hoth-Hannig W, Becker-Willinger C, Schmidt H. Influence of nanocomposite surface coating on biofilm formation in situ. J Nanosci Nanotechnol 2007;7:4642-8. 39. Diaz C, Schilardi PL, Salvarezza RC, De Mele MF. Nano/microscale order affects the early stages of biofilm formation on metal surfaces. Langmuir 2007;23:11206-10. 40. Battin TJ, Kammer FV, Weilhartner A, Ottofuelling S, Hofmann T. Nanostructured TiO2: transport behavior and effects on aquatic microbial communities under environmental conditions. Environ Sci Technol 2009;43:8098-104. 41. Tweden KS, Harasaki H, Jones M, Blevitt JM, Craig WS, Pierschbacher M, Helmus MN. Accelerated healing of cardiovascular textiles promoted by an RGD peptide. J Heart Valve Dis 1995;4:90-7. 42. Clubb FJ, Clapper DL, Deferrari DA, Hu SP, Seare WJ, Jr., Capek PP, Armstrong J, McGee MG, Billings LA, Fuqua JM, Parins SM. Surface texturing and coating of biomaterial implants: effects on tissue integration and fibrosis. Asaio J 1999;45:281-7. 43. Petrie TA, Raynor JE, Reyes CD, Burns KL, Collard DM, Garcia AJ. The effect of integrinspecific bioactive coatings on tissue healing and implant osseointegration. Biomaterials 2008;29:2849-57.
16
44. Hemmerlein JB, Trerotola SO, Kraus MA, Mendonca MS, Desmond LA. In vitro cytotoxicity of silver-impregnated collagen cuffs designed to decrease infection in tunneled catheters. Radiology 1997;204:363-7. 45. Atiyeh BS, Costagliola M, Hayek SN, Dibo SA. Effect of silver on burn wound infection control and healing: review of the literature. Burns 2007;33:139-48. 46. Widgerow AD. Nanocrystalline silver, gelatinases and the clinical implications. Burns 2010;36:965-74. 47. Bright KR, Gerba CP, Rusin PA. Rapid reduction of Staphylococcus aureus populations on stainless steel surfaces by zeolite ceramic coatings containing silver and zinc ions. J Hosp Infect 2002;52:307-9. 48. Cokeliler D, Goktas H, Tosun PD, Mutlu S. Infection free titanium alloys by stabile thiol based nanocoating. J Nanosci Nanotechnol 2010;10:2583-9. 49. Norowski PA, Bumgardner JD. Biomaterial and antibiotic strategies for peri-implantitis: a review. J Biomed Mater Res B Appl Biomater 2009;88:530-43. 50. Hochbaum AI, Aizenberg J. Bacteria pattern spontaneously on periodic nanostructure arrays. Nano Lett 2010;10:3717-21. 51. Kappell GM, Grover JP, Chrzanowski TH. Micro-scale surface-patterning influences biofilm formation. Electron J Biotechnol 2009;12:10-1. 52. Krsko P, McCann TE, Thach TT, Laabs TL, Geller HM, Libera MR. Length-scale mediated adhesion and directed growth of neural cells by surface-patterned poly(ethylene glycol) hydrogels. Biomaterials 2009;30:721-9. 53. Chung KK, Schumacher JF, Sampson EM, Burne RA, Antonelli PJ, Brennan AB. Impact of engineered surface microtopography on biofilm formation of Staphylococcus aureus. Biointerphases 2007;2:89-94.
17
54. Neut D, Dijkstra RJ, Thompson JI, Van der Mei HC, Busscher HJ. Antibacterial efficacy of a new gentamicin-coating for cementless prostheses compared to gentamicinloaded bone cement. J Orthop Res 2011: 29:1654-61 55. Diefenbeck M, Muckley T, Hofmann GO. Prophylaxis and treatment of implant-related infections by local application of antibiotics. Injury 2006;37(Suppl 2):S95-104. 56. Kwok CS, Horbett TA, Ratner BD. Design of infection-resistant antibiotic-releasing polymers. II. Controlled release of antibiotics through a plasma-deposited thin film barrier. J Control Rel 1999;62:301-11. 57. Davidoff SN, Call BP, Hogrebe PC, Grainger DW, Brooks AE. A robust method to coat allograft bone with a drug-releasing polymer shell - biomed 2010. Biomed Sci Instrum 2010;46:184-9. 58. Baier RE. Conditioning surfaces to suit the biomedical environment: recent progress. J Biomech Eng 1982;104:257-71. 59. Schulte VA, Diez M, Hu Y, Moller M, Lensen MC. Combined influence of substrate stiffness and surface topography on the antiadhesive properties of Acr-sP(EO-stat-PO) hydrogels. Biomacromol 2010;11:3375-83. 60. Nejadnik MR, Van der Mei HC, Norde W, Busscher HJ. Bacterial adhesion and growth on a polymer brush-coating. Biomaterials 2008;29:4117-21. 61. Muszanska AK, Busscher HJ, Herrmann A, Van der Mei HC, Norde W. Pluronic-lysozyme conjugates as anti-adhesive and antibacterial bifunctional polymers for surface coating. Biomaterials 2011;32:6333-41. 62. VandeVondele S, Voros J, Hubbell JA. RGD-grafted poly-L-lysine-graft-(polyethylene glycol) copolymers block non-specific protein adsorption while promoting cell adhesion. Biotechnol Bioeng 2003;82:784-90.
18
63. LeBaron RG, Athanasiou KA. Extracellular matrix cell adhesion peptides: functional applications in orthopedic materials. Tissue Eng 2000;6:85-103. 64. Maddikeri RR, Tosatti S, Schuler M, Chessari S, Textor M, Richards RG, Harris LG. Reduced medical infection related bacterial strains adhesion on bioactive RGD modified titanium surfaces: a first step toward cell selective surfaces. J Biomed Mat Res Part A 2008;84:425-35. 65. Shi Z, Neoh KG, Kang ET, Poh C, Wang W. Bacterial adhesion and osteoblast function on titanium with surface-grafted chitosan and immobilized RGD peptide. J Biomed Mater Res Part A 2008;86:865-72. 66. Subbiahdoss G, Kuijer R, Grijpma DW, Van der Mei HC, Busscher HJ. Microbial biofilm growth vs. tissue integration: "the race for the surface" experimentally studied. Acta Biomater 2009;5:1399-404. 67. Wang Y, Subbiahdoss G, Swartjes J, Van der Mei HC, Busscher HJ, Libera M. Length-scale mediated differential adhesion of mammalian cells and microbes. Adv Funct Mater 2011: 21:3916-23
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Chapter 2
Poly(ethylene glycol) as a Biointeractive Electron-Beam Resist
Chapter Summary Poly(ethylene glycol) (PEG) can serve as an electron-beam resist to modulate protein adsorption on surfaces. PEG preferentially crosslinks under e-beam irradiation to create microgels with controllable properties. Here atomic-force (AFM), scanning electron (SEM), and confocal microscopies are used to study discrete microgels formed from solvent-cast PEG thin films by focused electron beams with energies between 2 and 30 keV and point doses between 10 and 1000 fC. Consistent with experimental findings, Monte Carlo simulation of electron energy deposition identifies three structures within each microgel: a highly crosslinked core near the point of electron incidence; a lightly crosslinked near-corona surrounding the core; and a farcorona at the PEG-Si substrate created by backscattered electrons. The nature and relative sizes of these three regions and, hence, the microgel-protein interactions depend on the incident electron energy and dose. The far corona creates protein-repulsive surface that can extend hundreds of nanometers or more from the microgel core. The highly crosslinked core is largely shielded by the near corona. These findings can help guide the choice of irradiation conditions to most effectively modulate protein-surface interactions via PEG microgels patterned by electron beam lithography.
Yi Wang, Emre Firlar, Xiaoguang Dai, and Matthew Libera*; about to submit to Journal of Polymer Science: Polymer Physics.
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2.1 Introduction
The recognition that patterning at nano and micro length scales can have a profound effect on how synthetic surfaces interact with physiological systems has motivated a substantial interest in various surface-patterning technologies for biological applications.[1-5] Among these technologies is electron-beam (e-beam) lithography. It is attractive because of its ability to create features at high resolution in a user-specified pattern without the need to generate a mask.[6, 7] In one incarnation, it is employed in a manner similar to that primarily used by the semiconductordevice industry. Namely, a sacrificial polymer resist film is used as a medium to transfer a pattern to an underlying substrate, and this approach has been used to study such issues as how cellmaterial interactions can be influenced by surface topography patterned at sub-cellular length scales.[8] Increasingly, however, e-beam-patterning is being used to directly write structures that are either themselves biointeractive or can be chemically modified to confer some form of biospecific character to them.[5, 9-17]
The effects of incident electron energy and exposure dose, among other parameters, on the patterning of polymer resists used for microelectronic device fabrication are reasonably well understood. Energy imparted by incident electrons to a polymer film causes either e-beaminduced polymer chain scission and increased solubility (positive resist) or cross-linking and reduced solubility (negative resist). Typically, both chain-scission and crosslinking occur simultaneously during electron irradiation but at different rates. The overall nature of the resist is thus determined by the dominant of the two competing processes. There are materials where the dominant process depends on the incident electron energy and dose.
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In bio-relevant applications, the established exposure principles and practical experience developed for semiconductor device applications can often be immediately translated. Much less is known, however, about how electrons interact with the water-soluble polymer precursors often used in bio-relevant patterning applications such as poly(ethylene glycol) [PEG],[18-20] poly(vinyl pyrrolidone) [PVP],[21, 22] poly (acrylic acid) [PAA],[23] or polyacrylimide.[24] PEG films, for example, can be e-beam processed to create surface-patterned hydrogels that resist nonspecific protein adsorption and cell adhesion.[12, 25] By using PEG precursors with chemically active groups, post-irradiation chemical functionalization can confer precise biospecific character to these gels.[10-12, 26] Since, however, biointeractive properties are determined by molecular-level processes, subtle changes in electron-beam processing parameters can have a substantial effect on how the resulting surfaces will ultimately interact with cells and proteins. Furthermore, because of electron-substrate scattering and the energy-dependent rate of electron energy deposition in a solid, the distribution of deposited energy at and around the point of electron incidence can be extremely nonuniform. Little work has been done to characterize these effects in PEG thin films and thus provide a basis from which to choose processing parameters optimized to create patterns appropriate for some particular experiment or application. Here we use a combination of experiments and Monte Carlo simulation to explore the effects of incident electron energy and dose on thin films of PEG solvent cast onto silicon substrates.
2.2 Experimental and Modelling Procedures 2.2.1 E-beam patterning The essential steps of the patterning processes are outlined schematically by figure 2.1. A thin film of poly(ethylene glycol) [PEG], approximately 100 nm thick, is spin cast onto a silicon substrate. A beam of focused electrons in a computer-controlled scanning electron microscope
22
(SEM) is focused onto the film to deliver a controllable dose of energetic electrons to specific positions on the specimen. After irradiation, the entire specimen is immersed in a good solvent for PEG. Unexposed regions are washed away leaving behind patterned microgels where both a critical crosslink density and sufficient binding to the substrate has been achieved.
Silicon substrates (10 mm x 10 mm; 0.5 mm thick) were cleaned and vinyl-silane modified using a procedure described previously.[27, 28] Briefly, Si were sonicated in ethanol (96%) for 5 min and then dried with flowing N2. The wafers were then exposed to Piranha etch solution (3:1 98% sulfuric acid and 30% H2O2) for 30 min, rinsed with deionized (DI) water, dried with flowing N2, and exposed to a low-pressure O2 plasma (~300 mTorr, 1.75 W) for 10 min. The wafers were then immersed in a 2% solution of [v/v] vinyl-methoxy siloxane homopolymer (viscosity: 8-12 cst, Gelest Inc., Morrisville, PA) in ethanol for 10 min. The substrates were rinsed with ethanol, dried, and baked at 110 °C for 2 hr. After cooling they were used immediately for spin coating. Thin films were cast from solutions of PEG (Mw = 6 kDa; Fluka) dissolved in tetrahydrofuran (THF) on Si substrates spinning for 5 min (4000 rpm). Unless otherwise noted, films were cast from 2 wt% PEG solutions. Samples were stored under vacuum (~10-3 Torr) until they were used for patterning.
Measurements of dry film thickness were made using a custom-built, single-wavelength, phase-modulated ellipsometer at 65° angle of incidence. The refractive indices for the native silicon oxide layer and dry PEG films on a silicon substrate were assumed to be 1.456 and 1.500, respectively. Three measurements were made from each specimen. Patterning was performed at room temperature using a Zeiss Auriga SEM with a Schottky fieldemission (FEG) source and a Nabity Nanometer Pattern Generation System. The patterning
23
experiments used a focused electron beam of known current which was allowed to dwell for controlled times at individual points on the specimen. The inter-pixel spacing, δ, between point exposures was varied between 0.1 and 5 µm, most typically being 3 µm. Incident electron energies of 2, 5, 10, or 30 keV were used. After electron exposure, the substrates were washed in a series of THF rinses to remove unexposed or insufficiently crosslinked PEG.
SEM imaging was done with same Zeiss Auriga SEM used to write the microgel patterns with fresh specimen fabricated only for the use of SEM imaging. And we also made patterned specimen specific for AFM (NanoInk, Nscriptor, Skokie, IL) measurements. We preceded scanning in contact mode with the research AFM software package (SPM Cockpit) provided by NanoInk. When making height measurements, gel heights were first determined in a dry state. A drop of purified water was put on top of the patterned area, and measurements of the gel height in the hydrated state were made after equilibrating the specimen in water for approximately 10 min. The atomic force microscopy (AFM) imaging force was minimized to limit the deformation of the gels by the AFM tip.
2.2.2 Fibronectin adsorption After washing with PBS buffer, patterned surfaces were exposed to human plasma fibronectin (Fn) purified protein (Merck Millipore, Temecula, CA) at 37 °C for 2 hrs with solution of 0.1 mg/mL Fn in PBS buffer. Samples were then rinsed (5 min) with PBS three times and exposed to a solution of mouse antihuman primary antibody (Monoclonal Anti-Fibronectin antibody produced in mouse, Sigma-Aldrich, St.Louis, MN) for 30 min at room temperature. After again rinsing in PBS, the samples were immersed in the FITC-conjugated goat anti-mouse secondary antibody (0.1 mg/mL in PBS) for 30 min at room temperature. They were rinsed with PBS and DI water,
24
and then dried under flowing N2 gas. Immunofluorescence confocal images were immediately taken with a Nikon E1000 upright microscope and C1 confocal imaging system using a 60x water-immersion lens with an NA of 1.2.
2.2.3 Morphological Analysis of PEG Microgel SEM (scanning electron microscope) imaging was done with same FEI Helios Dual-Beam FIBSEM used to write the microgel patterns with fresh specimen fabricated only for the use of SEM imaging. Quantitative measurements of the gel height using a NanoScope IIIa scanning probe microscope in contact mode (Digital Instruments, Veeco Metrology Group) employed Veeco Nanoprobe tips (model NP-20). When making these height measurements, specimens were mounted in a Digital Instruments liquid cell. Gel heights were first determined from microgels in the dry state. The liquid cell was then filled with deionized water, and measurements of the gel height in the hydrated state were made after equilibrating the specimen in water for approximately 10 min. The AFM imaging force was minimized to limit the deformation of the gels by the AFM tip. Lateral force mode was used in conjunction with topographical mode, and both topographical and lateral force images were generated simultaneously.
2.2.4 Monte-Carlo Simulations The interaction of electrons of various energies with thin PEG films on semi-infinate silicon substrates was simulated using a Monte Carlo model.[29, 30] Single-electron trajectories were followed, and decisions of scattering type - elastic or inelastic - scattering angle, and energy change due to electron-nuclei interactions were made using a random-number generator weighted by a screened Rutherford elastic cross section. Energy deposition into the PEG film was accounted for using the modified expression for the Bethe stopping power:[29, 31]
25
S
dE 78500Z 1.166( E KJ ) ln dx AE J
…[1]
where K =0.734Z0.037 and S, Z, A, E, and J are the stopping power, atomic number, atomic weight, instantaneous energy, and mean ionization potential, respectively. Individual electron trajectories were followed until the energy of the electron fell below 50 eV or until an electron reached the top surface of the specimen, (x, y, z = 0), where it escaped into the surrounding vacuum. In the former case, the remaining 50 eV of energy was ignored. Prior to this endpoint, the electron energy loss between scattering events due to the continuous slowing down of the electron in the specimen - given by the product of the instantaneous stopping power and the distance between successive scattering events – was accumulated in the 1 nm3 voxel at the end of the particular trajectory segment. Energy deposition in the Si substrate was ignored, since the semi-infinite Si is a good heat sink. This energy was assumed to be conducted away with no effect on the PEG film. The coordinates of each electron trajectory were recorded in one of the three Figure2.1 Schematic description of the electron-beam lithography process to form surface-patterned PEG microgels.
different arrays: (1) as a primary electron in the PEG film; (2) as a primary electron in the Si; and (3) as an electron backscattered from the Si substrate into the PEG film.
26
Figure 2.2 Secondary-electron SEM images of PEG microgels patterned in a 2x3 matrix on silanized Si surfaces using different incident electron energies and different point doses. The inter-gel spacing is 3 mm.
2.3 Results and Discussion 2.3.1 Microgel-Patterned Surfaces Electron-beam exposure and post-exposure rinsing (fig. 2.1) of PEG thin films leads to the formation of features patterned on the underlying silicon substrate. We have previously shown [14, 28, 32] that these features swell when immersed in water in a dose-dependent manner and thus correspond to microscopic hydrogels (microgels). Immunofluorescence imaging (see below and elsewhere [14, 25, 33]) demonstrates that these features resist protein adsorption consistent with hydrated PEG gels and brushes. Typical SEM images of these microgels are presented by figure 2.2. The spots of dark contrast correspond discrete microgels with center-to-center separations of 3 µm. In the cases shown in figure 2.2, point doses of 10 fC, 1000 fC, and 1000 fC at incident energies of 2, 5, 10, and 30 keV are all able to produce features on the silicon surface, though these images clearly show that the lateral size of each spot varies and does so very nonlinearly with increasing incident energy and dose. Furthermore, the fact that spots are present after the e-beam processing indicates that these exposure conditions are able to not only crosslink
27
the PEG itself but also graft the resulting microgel to the substrate below. Conditions of critical crosslinking and grafting to the underlying surface must both be met in order to form stable microgels patterned on a surface.
Figure 2.3 Different thickness of PEG homopolymer films can be generated by spin coating solutions of varying PEG concentration. The images show that 100 fC of 2 keV incident electrons are unable to both form microgels and bind these to the underlying substrate.
The ability of focused electron exposure to meet the requirement of microgel grafting to an underlying substrate is addressed by figure 2.3. This experiment exposed films of varying thickness to 100 fC point doses of electrons with different incident energies. The film thickness was varied from 50 nm to 250 nm by adjusting the concentration of PEG during spin casting from 1 wt% to 5 wt%. SEM imaging shows that 10 keV electrons are able to form stable microgels for film thicknesses of 50 nm and 250 nm (fig. 2.3 top), as well as for intermediate film thicknesses (data not shown). Thus, 10 keV electrons, as well as higher energy electrons, are able to traverse as much as 250 nm of PEG and induce radiation chemistry both within the film and between the film and the underlying substrate. In contrast, the SEM images show that the size of the
28
individual microgels formed from 2 keV electrons increases significantly as the film thickness increases indicating that the lateral range of the electron-polymer interactions depends strongly on film thickness for this incident energy. Furthermore, the fact that the array periodicity is imperfect in the case of the 200 nm film and that the spots are altogether missing in the 250 nm film indicates that, while 2 keV electrons may be able to critically crosslink PEG into microgels from films greater than about 200 nm, they are unable to deposit sufficient energy at the film/substrate interface to successfully graft the microgels to the underlying surface.
Figure 2.4 Phase-contrast AFM images of PEG microgels patterned on silanized Si surfaces with different incident electron energies and different point doses. The inter-gel spacing is 3 mm.
AFM phase images for the various doses and incident energies are presented in figure 2.4. The microgels spots are consistently smallest for the 10 fC dose, and they are unresolvable for the particular case of 10 fC of 30 keV electrons. The microgels exhibit nonlinear increases in diameter with increasing dose and incident electron energy. Note that the diameters of the microgels appear slightly larger in the AFM images than in the SEM images. We attribute this to
29
the fact that an AFM probe can more effectively detect very thin (e.g. monolayer) polymer films whereas the secondary electron signal detected in an SEM would be dominated by electron emission from the underlying substrate. Significantly, the AFM images indicate that there is structure within each spot not revealed by the SEM images (fig. 2.2). In the spot formed by 1000 fC irradiation by 2 keV electrons, for example, one can see a disk of dark contrast with a point of light contrast in its center. Similar structures can be resolved in the 5 keV and 10 keV images. While phase contrast in scanned probe microscopy is a relatively complex signal that depends both on material properties and on instrumental parameters, there is agreement that it is related to the local viscoelastic properties of the sample.[34, 35] Hence, these data indicate that the mechanical properties of the material in the center of the microgel have viscoelastic properties very different from that around its periphery.
Figure 2.5 Confocal immunofluorescence images of surfaces patterned by PEG microgels using different incident electron energies and different point doses and then exposed to human fibronectin followed by a primary and FITC-labeled secondary antibody. The inter-gel spacing is 3 µm.
30
The fact that the e-beam patterned PEG microgels resist protein adsorption is illustrated by figure 2.5. This shows an array of immunofluorescence confocal images from patterned surfaces after fibronectin exposure (green). The dark contrast corresponds to the positions of the microgels. The fact that the Fn signal is weak or, in some cases, absent at these positions is consistent with the well-known antifouling properties of PEG and PEG hydrogels. Proteins tend to not bind to PEG both because of the energetically unfavourable need to displace relatively strong hydrogen bonds between water and the ether oxygen within the ethylene glycol unit and because of the decreased conformational entropy when proteins bind to a PEG segment in a gel. For point doses of 1 fC, none of the four incident electron energies produce microgels whose effect on Fn adsorption can be resolved using a 60x (NA = 1.2) water-immersion objective (data not shown). As the dose increases, points of dark contrast appear. Such points are visible for all four incident energies when the point dose is 1000 fC, though the diameters of the individual protein-repulsive spots vary substantially for different incident energies.
2.3.2 Monte Carlo Modelling of Electron-PEG Interactions When an energetic electron impinges from vacuum onto a solid substrate, the electron undergoes both elastic scattering and inelastic scattering. The nature and extent of these scattering processes depend on the initial energy of the incident electron and the physical characteristics of the substrate. Both the elastic and inelastic scattering cross sections scale nonlinearly with the average atomic number,
Z . For the case of a polymer thin film on a semi-infinite inorganic
substrate such as silicon, an electron thus traverses a region of low region of higher
Z before impinging on a next
Z where the scattering, as well as the consequences of that scattering, is very
different. Figure 2.6 presents the results of a Monte Carlo simulation describing the trajectories of 1000 electrons with an incident energy of 30 keV. This image corresponds to an x-z section
31
from a 3-D calculation. The electron trajectories in the Si substrate are represented by the blue colour. These dominate this image, and the interaction volume is approximately 8 µm in diameter. At higher magnification, the figure 2.6 inset shows the simulated trajectories of 1000 electrons with an incident energy of 2, 5, or 10 keV. The interaction volumes for these cases are substantially smaller. For the 2 keV case, the interaction volume is only ~200 nm in diameter and, unlike the 30 keV case, is confined largely to the PEG thin film.
Figure 2.6 The interaction volume of 30 keV electrons with a thin film (100 nm) of PEG on a silicon substrate. The inset shows the interaction volumes for 2, 5, and 10 keV incident electron energies.
Details of the trajectories as the incident electrons traverse the thin-film (100 nm) PEG resist into the semi-infinite Si substrate are presented for the four incident energies in figure 2.7. Primary electrons in the resist layer are coloured red. Primary electrons that pass through the resist into the substrate are coloured blue. Because of the higher atomic number in the substrate, the electrons there are, on average, scattered to higher angles. Some of these electrons are scattered multiple times and pass through the resist layer a second time. These backscattered electrons are
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coloured green. They can emerge from substrate many 10's or 100's of nm from the
Figure 2.7 Monte Carlo calculations of the trajectories of 1000 electrons for different incident electron energies. Red: primary electrons in the PEG film; Blue: primary electrons in the Si substrate; Green: backscattered electrons in the PEG film.
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initial point of electron incidence, and in e-beam lithography they are primarily responsible for the so-called proximity effect, where beam-induced resist exposure occurs in proximity to, but surrounding, the point of electron incidence. The proximity effect degrades the spatial resolution of patterning. Hence, understanding and controlling the proximity effect is an important element of precision electron-beam lithography.
Figure 2.8 Monte Carlo simulations of deposited energy per unit volume for 1000 fC point dose from electrons with incident energies of 2, 5, 10, and 30 keV.
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Chemical changes in the polymeric resist leading to chain scission or crosslinking occur because of the energy deposited by the electrons during inelastic scattering. Among the inelastic processes that can take place are: knock-on damage; core-loss, plasmon, or phonon excitation; bremstrahlung generation; and the production of both fast and slow secondary electrons, which themselves subsequently lose energy by the other processes. Each of these processes is characterized by a scattering cross section, which depends on, among other things, the instantaneous electron energy. The cross-section for knock-on processes, for example, becomes very small in most polymers at electron energies below about 30 keV,[36] and in graphene the threshold for knock-on damage of carbon is 80 keV.[37] Except for electron energies of about 1 – 1.5 keV, the cross section for carbon core-loss excitation is also relatively low. Most energy is thus deposited by the various low-loss excitations. While, in principle, Monte Carlo simulation of electron-solid interactions can track all of these individual excitations, the interactions are more commonly accounted for collectively by the total inelastic scattering cross section, which is embodied within the stopping power, S = dE/dx (equation [1]). The rate of energy deposition in the specimen increases by almost an order of magnitude as the electron energy decreases from 30 keV to about 2 keV and then decreases as the electron energy falls further.
Figure 2.8 maps the spatial distribution of energy deposited in a PEG thin film by 6.2M (1000 fC) electrons with each of the four different incident electron energies. These images correspond to x-z planes from a 3-D simulation where each pixel value represents the average of three 1 nm3 voxels sampled in the y direction. The colour codes denote different amounts of energy deposited per 1 nm3 voxel. These energies are particularly important when one considers the possible effects of the radiation the PEG thin film. One specific effect is crosslinking. We estimated a threshold energy density for crosslinking based on Gx values describing the number
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of crosslinking events per 100 eV of energy deposited in a sample. These range from about 0.5 to 5 and depend on the polymer and on the irradiation conditions.[38, 39] We assumed a Gx value of 2, and, together with the density of PEG (1.05 g/cm3) and the molecular weight (5 kDa), this value suggests a threshold energy density for crosslinking of 6 eV/nm3. Hence, the range of 5-8 eV (blue) in figure 2.8 indicates energy sufficient for approximately 1 PEG crosslinking event. 818 eV (purple) indicates approximately 3 PEG crosslinking events. The average degree of polymerization of 5 kDa PEG is 114, so 1-3 crosslinking events per molecule would lead to a structure that preserves substantial segmental flexibility, and the hydrophilic PEG would remain quite gel-like. Figure 2.8 manifests regions with these low energy densities on the periphery of the interaction volumes within the PEG, and these form a corona around each microgel.
In contrast to the lightly crosslinked corona, an energy density of 228 eV/nm3 would correspond to approximately 1 crosslinking event for every 3 monomer units. This structure would be far more rigid and not gel-like. Furthermore, at these higher doses, significant compositional changes are likely to occur, and one can argue that within the range of 18-228 ev/nm3 (yellow) the polymer will transition, both structurally and compositionally, away from being a PEG-like gel. The deposited energy close to the column along the direction of electron incidence can reach thousands of eV/nm3 (red in figure 2.8). In this region the exposed resist most likely bears little resemblance to the unexposed PEG. We refer to this region as the core. The core is surrounded by the corona.
While each irradiated spot has only one core, the corona can be subdivided into the near corona and the far corona. The near corona is a soft layer of gel that surrounds the chemically and structurally distinct core region. As suggested by the electron trajectories (fig. 2.7), the near
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corona is formed largely from energy deposited by primary electrons that have been scattered away from their initial incident direction (z). The far corona, on the other hand, has similar gellike properties but is localized at the substrate interface and can extend many tens of nm or more away from the core. The far corona is formed largely by backscattered electrons.
Importantly, the presence of a lightly crosslinked corona and a heavily crosslinked and chemically modified core is consistent with the phase-contrast AFM images (figure 2.4). These show a central spot with bright contrast whose mechanical properties are somewhat like those of the exposed silicon substrate. This corresponds to the core. Around the bright central feature is a circular area of material with dark contrast whose mechanical properties, as manifested by the phase-contrast AFM imaging, must be different from the core. This darker-contrast area corresponds to the corona, which we can expect to have both a lower viscosity and lower elastic modulus than the core.
Figure 2.8 shows that four different incident electron energies all produce a core, but the size and shape of the core varies substantially. With increasing accelerating energy the width of the core decreases consistent with decreasing scattering cross sections that preserve electron trajectories primarily in the forward (z) direction (fig. 2.7). In all four cases the core also extends through the entire PEG film and into the substrate below. This is important because the extensive radiation chemistry in the core would likely ensure strong binding to the substrate. The 2 keV case is particularly noteworthy since the core of heavy energy deposition does not penetrate far into the substrate. If, for example, the substrate thickness were increased from the 100 nm used in the present calculations to 200 nm, the simulation of figure 2.8 shows that there would be enough energy deposition into the PEG film to form a microgel but this microgel particle would
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not be bound to the substrate below. Hence, it would wash away during the development step. This finding is reasonably consistent with the experimental observation reported in figure 2.3, which shows that 2 keV electrons are able to bind microgels to the substrate for PEG film thicknesses of 200 nm and below but not for films whose thickness is 250 nm. The simulation of figure 2.8, together with the trajectory information of figure 2.7 inset, indicates that core energy deposition for the 10 keV incident energy extends deep into the substrate, and, consistent with the experimental results of figure 2.3, is able to bind microgels to the substrate surface even for PEG film thicknesses of 250 nm.
Figure 2.8 also shows that the corona regions vary between the four different accelerating energies. In the case of 2 keV incident electrons, for example, the near corona is relatively thick. However, there is essentially no far corona. In this case, the number of backscattered electrons is low and they have insufficient energy to travel far from the core. In contrast, higher-energy electrons can be backscattered from the substrate and can again provoke radiation chemistry on their second trip through the PEG. Furthermore, as illustrated by the electron trajectories presented in figure 2.6, these backscattered electrons can induce radiation chemistry many 10s or 100s of nm away from the core. In the case of 30 keV incident electrons, the distances can be as much as several microns. Much of the radiation chemistry induced by backscattered electrons will occur throughout the PEG film. However, much of this will be insufficiently extensive to graft the resulting gel to the substrate. Consequently, bits of gel that might be formed by this mechanism will be removed during the development step. However, crosslinking events that occur at or near the substrate can also involve grafting to the substrate. In this case, a layer of PEG gel or brush would remain on the surface after development. This layer would become increasing discontinuous with increasing distance from the core.
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Figure 2.9 Monte Carlo simulation of fibronectin adsorption on e-beam patterned PEG surfaces. Voxels within 5 nm of the substrate surface with deposited energies exceeding 6 eV/nm3 were assigned black (protein repulsive), and those with less energy were assigned green (protein adhesive). The intensity profiles plot the image intensity averaged over a 10 pixel width.
The near corona, the far corona, and the core all play a role in how the microgel interacts with protein. For example, since the core is so heavily modified by the electron beam, it no longer has the antifouling properties characteristic of high-swelling PEG gels or PEG brushes. We have shown previously [14] shown that the top of the core region can be exposed by patterning distinct point irradiations in close proximity (~ 200 nm spacings) to each other. Under high-dose irradiation, the core region swells little and is more adhesive to fibronectin than the surrounding silicon substrate. In the current involving surfaces patterned with discrete microgels spaced several microns apart, however, immunofluorscence imaging shows little evidence of a core
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region. In this case the core is surrounded by some amount of gel-like PEG that constitutes the near corona. It is this outer layer that interacts with proteins. The far corona will have similar antifouling character. Except perhaps for regions relatively close to the core when very high point doses are used, the far corona will be comprised primarily of discontinuous monolayer films grafted to the substrate surface. In a hard-materials application, such as the fabrication of a semiconductor device or microfluidic device, this far-corona material is typically of little concern since it can be rapidly removed by any number of physical or chemical etching procedures that would be used in a subsequent processing step. In a biological application, however, since only a monolayer of PEGylation is sufficient to modify the physicochemical properties of a surface, the far corona can have a substantial impact on the local protein and cell interactive properties of a surface.
To better visualize the effects of the corona, we modelled the adsorption of fibronectin around an e-beam patterned PEG microgel. We assumed that voxels within 5 nm of the substrate surface with deposited energies exceeding 6 eV/nm3 are PEGylated and protein resistant while those voxels with less energy present an unmodified silicon surface and are protein adhesive. The former were assigned to be black, and the latter were assigned to be green. The results are summarized by figure 2.9 for the range of incident electron energies and point doses. The general trends follow those observed experimentally by immunofluorescence confocal imagine (figure 2.5). For example, both the simulation and the experiment indicate that the 5 and 10 keV spots at the higher doses have the largest areas of protein resistance. The simulation provides more detailed information, however, since it does not suffer from the same resolution limits as the confocal microscope. The 30 keV microgel spots are clearly the finest. With increasing dose, however, these suffer from an increasingly apparent far corona, which manifests itself in the
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intensity profile as the deviations below the uniform background away from the core. For the 1000 fC 30 keV case, this far corona has a substantial impact on the Fn adsorption far from the core. On the other hand, the simulations show that the 2 keV microgel spots exhibit no far corona for any dose. In between these extremes, the simulations show that microgel spots formed using 5 or 10 keV incident electron energies will have a strongly protein repelling central region surrounded by a diffuse halo of decreasing antifouling character. The 5 and 10 keV cases also display the strongest spot size dependence on dose.
2.3.3 Selecting Irradiation Conditions for Microgel Patterning One of the important distinctions of an exposed resist to be used directly for a biological application is the role that very thin layers of exposed resist can play on the resulting properties. A monolayer of PEG is sufficient, for example, to convert a surface from protein adhesive to nonadhesive. In contrast, a monolayer of resist is usually inconsequential when the lithography step is followed by an etching step, since a monolayer of polymer can rapidly be removed to expose underlying substrate while 100's of nms of adjacent resist will survive the etching and transfer the pattern. The role of thin layers in the case of e-beam-patterned PEG microgels is clear in at least two places: the core and the far corona. Despite the fact that the core is physically and chemically very different from unexposed PEG, it can be shielded from a protein-containing solution by the gel-like PEG in the adjacent near corona. We have shown previously that the top of the protein-adhesive core can be shielded when the e-beam patterning conditions or the choice of PEG precursor molecular weight is such that the size of the core is less than the radius of gyration (Rg) of the PEG segments in the adjacent loosely crosslinked near corona. For the purposes of patterning bio-active function onto a surface, the role of thin PEG layers is perhaps even more important in the far corona. Depending on the combination of incident electron energy
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and radiative dose, a sufficient number of backscattered electrons can be generated which graft PEG to the substrate 10's or 100's of nanometers from the core. When the average spacing between such grafted PEG molecules is comparable to the characteristic size of approaching proteins and biomolecules, the modified surface will resist adhesion. Hence, the proximity effect can have an even bigger impact in a bio-interactive application than in a semiconductor device fabrication application.
The apparent size of individual microgels depends not only on the incident electron energy and electron dose but also on the measurement method. Table I summarizes microgel diameters determined from phase-contrast AFM imaging, immunofluorescence confocal imaging, and Monte Carlo simulation. These were determined by drawing a line profile of intensity across a microgel, establishing an average background by measuring an average signal intensity away from the microgel, and then determining the full-width at tenth maximum (FWTM) of the microgel intensity below the average background. Each of the values from the AFM and confocal imaging represent the average and standard deviation of at least nine different measurements. Those by AFM imaging are the most accurate. The spatial resolution and
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sensitivity are high enough to detect microgels for all of the irradiation conditions except that of 30 keV 10 fC exposure. Furthermore, the phase shifting can be influenced by monolayers of polymer and, unlike secondary electron imaging in an SEM (microgel diameter measurements not tabulated), is sensitive to the edges of the far corona. In contrast, the diameters measured by confocal imaging are both less precise and consistently larger than those measured by AFM. There are at least two reasons for the discrepancies. One is that the resolution of the confocal microscope is less than that of the AFM. The smallest detectable feature by this approach, for example, corresponds to the 2 keV 10 fC microgel with a diameter of 0.7 µm whereas corresponding AFM measurement gives a FWTM of 0.37 µm. The difference is in part due to the blurring effect of the point spread function of the optical microscope, which, for the imaging conditions used (FITC excitation and emission wavelengths of 490 and 525 nm, respectively; 60x water-immersion lens with NA = 1.2), can be modelled as a Gaussian with a FWTM of approximately 400 nm.[40] Deconvoluting this point-spread function from the fluorescence images would bring better agreement with the AFM measurements. The fluorescence images also suffer from the fact that they have substantially less signal and more noise than the AFM images, particularly at the edges of the far corona that define microgel size and particularly for the cases such as the 2 keV 10 fC irradiation where the protein-resistant spot is weakly detectable relative to the background (figure 2.5).
The spot-size predictions of the Monte Carlo simulations follow the same trends as the experimental measurements though there are several important discrepancies. Perhaps most critical among the assumptions of the model is the energy-dependent stopping power, which is least well understood at lower electron energy levels. The fact that the simulations predict a maximum penetration depth for 2 keV electrons of just over 100 nm (figure 2.7) while the
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experiments indicate this value is closer to 200 nm maximum depth is one indication of uncertainty in the stopping-power model. Furthermore, there is no precise measure of a G value for crosslinking in PEG. Assuming a value lower than 2, for example, would indicate crosslinking events farther from the core and, hence, larger protein-resistant spot sizes. Nevertheless, the Monte Carlo model correctly predicts the trend for the 5 and 10 keV irradiations that the spot size will increase with increasing dose, because the proximity effect manifests itself at farther distances from the core with increasing dose. The model also predicts that 30 keV irradiation produces small spots, including ones that may be undetectable by AFM or confocal imaging (e.g. 10 fC 30 keV) but which could nevertheless influence protein-material and cellmaterial interactions at nanolength scales. Note, however, that the 30 keV predictions of spot size do not reflect the proximity effects very effectively. These can produce a background of proteinresistant PEGylation both uniformly and at very large distances from the core because of the relatively long trajectories of 30 keV electrons. Finally, the simulations reinforce the experimental measurements that 2 keV irradiation leads to relatively small spot sizes. The simulations do not, however, exhibit the same increasing spot size with increasing dose as seen experimentally. Within the current assumptions, the 2 keV electrons can traverse about 100 nms of PEG, and this is about how far they go not only in the forward (z) direction but also in the lateral directions (x and y). Hence, in the model, few electrons reach the substrate, and there is essentially no dose-dependent proximity effect. The discrepancies may again be in part due to the uncertainties in stopping power for electrons with energies of 2 keV and below. An overestimate of the rate of energy loss would underestimate the maximum electron trajectory. If, as suggested by figure 2.3, 2 keV electrons can travel through the PEG film and into the Si substrate, then a fraction of these would be backscattered and lead to a far corona whose size would increase from zero with increasing incident dose.
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The incident electron energies used in e-beam lithography range from about 1 keV to about 100 keV. The highest energies (e.g. 100 keV) are used to improve resolution, since most of the exposure is by primary electrons as they traverse the resist. Most common, perhaps, are incident energies of 10-30 keV, in part because affordable systems derived from scanning-electron microscopes, able to also perform imaging functions, are commercially available and have been relatively accessible to a broad community for many years. Lower energies (e.g. 1-2 keV) are of particular interest because of their ability to work with electrically-nonconductive substrates such as glass. E-beam lithography using higher energy electrons (30 keV, 100 keV) with nonconductive substrates can be a challenge because the charging of the underlying substrate creates electrostatic fields that deflect incident electrons and substantially degrade pattern integrity. One solution is to use an electrically conductive layer above the resist layer, but such an approach in a bio-relevant application would compromise the requisite biointeractive properties of the resist itself. The fact that low electron energies (~200 eV – 2 keV) are useful for SEM imaging of uncoated and electrically nonconductive specimens, because of dynamic charge neutralization, is well established. The present experiments indicate that a combination of 2 keV incident electrons with a resist thickness on the order of 100 nm is sufficient to not only generate surface-patterned microgels but ones whose size and properties are relatively insensitive to incident dose.
2.4 Conclusion We studied the effects of incident electron energy, exposure dose, and other parameters on the patterning of PEG resist used for bio-relevant applications such as protein adsorption. We found that low beam energy (2 keV) and lower exposure point dose (< 100 fC) were more suitable for
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bio-relevant patterning applications. With most of the energy deposited within the PEG resist, low beam energy helped to decrease the proximity effect (backscattering) dramatically while at the same time maximally preserved the bio-interactivities provided by low-crosslinking density.
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References 1.
Anselme K, Davidson P, Popa AM, Giazzon M, Liley M, Ploux L. The interaction of cells and bacteria with surfaces structured at the nanometre scale. Acta Biomater.2010;6(10):382446.
2.
Falconnet D, Csucs G, Michelle Grandin H, Textor M. Surface engineering approaches to micropattern surfaces for cell-based assays. Biomaterials. 2006;27(16):3044-63.
3.
Kane RS, Takayama S, Ostuni E, Ingber DE, Whitesides GM. Patterning proteins and cells using soft lithography. Biomaterials. 1999;20(23-24):2363-76.
4.
Khademhosseini A, Langer R, Borenstein J, Vacanti JP. Microscale technologies for tissue engineering and biology. Proc Natl Acad Sci USA. 2006;103(8):2480-7.
5.
Kolodziej CM, Maynard HD. Electron-beam lithography for patterning biomolecules at the micron and nanometer scale. Chem. of Mater. 2012;24(5):774-80.
6. Menon R, Patel A, Gil D, Smith HI. Maskless lithography. Materials Today. 2005;8(2):26-33. 7.
Rai-Choudary P. Handbook of Microlithography, Micromaching, and Microfabrication: Volume I. Bellingham: SPIE; 1997.
8.
Wilkinson CDW, Dalby M, Curtis ASG. Making structures for cell engineering. Eur. Cells Mater. 2004;8:21-6.
9.
Ballav N, Thomas H, Winkler T, Terfort A, Zharnikov M. Making protein patterns by writing in a protein-repelling matrix. Angewandte Chemie 2009;48(32):5833-6.
10. Christman KL, Schopf E, Broyer RM, Li RC, Chen Y, Maynard HD. Positioning multiple proteins at the nanoscale with electron beam cross-linked functional polymers. J Am Chem Soc. 2009 Jan 21;131(2):521-7. 11. Dai X, Yang W, Firlar E, Marras SAE, Libera M. Surface-patterned microgel-tethered molecular beacons. Soft Matter. 2012.
47
12. Hong Y, Krsko P, Libera M. Protein surface patterning using nanoscale PEG hydrogels. Langmuir. 2004 Dec 7;20(25):11123-6. 13. Kolodziej CM, Chang C-W, Maynard HD. Glutathione S-transferase as a general and reversible tag for surface immobilization of proteins. J. of Mater. Chem. 2011;21(5):1457-61. 14. Krsko P, Mansfield M, Sukhishvili S, Clancy R, Libera M. Electron-Beam Patterned Poly(Ethylene Glycol) Microhydrogels. Langmuir. 2003;19:5618-25. 15. Pesen D, Haviland DB. Modulation of cell adhesion complexes by surface protein patterns. ACS Appl Mater Interfaces. 2009;1(3):543-8. 16. Senaratne W, Sengupta P, Harnett C, Craighead H, Baird B, Ober CK. Molecular templates for bio-specific recognition by low-energy electron beam lithography. Nanobiotechnology. 2005;1(1):23-33. 17. Zhang GJ, Tanii T, Zako T, Hosaka T, Miyuke T, Kanari Y, et al. Nanoscale patterning of protein using electron beam lithography of organosilane self-assembled monolayers. Small. 2005;1(8-9):833-7. 18. Krsko P, Libera M. Biointeractive hydrogels. Mater Today. 2005;8(12):36-44. 19. Krsko P, Kaplan JB, Libera M. Spatially controlled bacterial adhesion using surfacepatterned poly(ethylene glycol) hydrogels. Acta Biomater. 2009;5(2):589-96. 20. Lussi JW, Tang C, Kuenzi P-A, Staufer U, Csucs G, Vörös J, et al. Selective molecular assembly patterning at the nanoscale: a novel platform for producing protein patterns by electron-beam lithography on SiO2/indium tin oxide-coated glass substrates Nanotechnology. 2005;16. 21. Burkert S, Schmidt T, Gohs U, Mönch I, Arndt K-F. Patterning of thin poly(N-vinyl pyrrolidone) films on silicon substrates by electron beam lithography. J Appl Polym Sci. 2007;106(1):534-9.
48
22. Lee BM, Kang DW, Jung CH, Choi JH, Hwang IT, Hong SK, et al. Patterning of polymer nanocomposite resists containing metal nanoparticles by electron beam lithography. J Nanosci Nanotechnol. 2011 Aug;11(8):7390-3. 23. Saaem I, Tian J. e-Beam Nanopatterned Photo-Responsive Bacteriorhodopsin-Containing Hydrogels. Adv Mater. 2007;19(23):4268-71. 24. Tirumala VR, Divan R, Ocola LE, Mancini DC. Direct-write e-beam patterning of stimuliresponsive hydrogel nanostructures. J Vac Sci Technol B. 2005;23(6):3124-8. 25. Krsko P, McCann TE, Thach TT, Laabs TL, Geller HM, Libera MR. Length-scale mediated adhesion and directed growth of neural cells by surface-patterned poly(ethylene glycol) hydrogels. Biomaterials. 2009 Feb;30(5):721-9. 26. Broyer RM, Schopf E, Kolodziej CM, Chen Y, Maynard HD. Dual Click reactions to micropattern proteins. Soft Matter. 2011;7(21):9972-7. 27. Krsko P, Sukhishvili S, Mansfield M, Clancy R, Libera M. Electron-Beam Surface-Patterned Poly(ethylene glycol) Microhydrogels. Langmuir. 2003 2003/07/01;19(14):5618-25. 28. Krsko P, Saaem I, Clancy R, Geller H, Soteropoulos P, Libera M. E-beam-patterned hydrogels to control nanoscale surface bioactivity. Proc SPIE. 2005;6002. 29. Joy DC. Monte Carlo Modeling for Electron Microscopy and Microanalysis. New York: Oxford University Express; 1995. 30. Shimizu R, Ze-Jun D. Monte Carlo modelling of electron-solid interactions. Rep Prog Phys. 1992:487-531. 31. Joy DC, Luo S. An empirical stopping power relationship for low-energy electrons. Scanning. 1989;11(4).
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32. Wang Y, Subbiahdoss G, Swartjes J, Van Der Mei HC, Busscher HJ, Libera M. Length-scale mediated differential adhesion of mammalian cells and microbes. Adv. Funct. Mater. 2011;21(20):3916-23. 33. Wang Y, Subbiahdoss G, de Vries J, Libera M, van der Mei HC, Busscher HJ. Effect of adsorbed fibronectin on the differential adhesion of osteoblast-like cells and Staphylococcus aureus with and without fibronectin-binding proteins. Biofouling. 2012;28(9):1011-21. 34. Passeri D, Rossi M, Tamburri E, Terranova ML. Mechanical characterization of polymeric thin films by atomic force microscopy based techniques. Anal. Bioanal. Chem. 2013;405(5):1463-78. 35. Wang Y, Song R, Li Y, Shen J. Understanding tapping-mode atomic force microscopy data on the surface of soft block copolymers. Surface Science. 2003;530(3):136-48. 36. Egerton RF. Mechanisms of radiation damage in beam-sensitive specimens, for TEM accelerating voltages between 10 and 300 kV. Microsc Res Techniq. 2012;75(11):1550-6. 37. Meyer JC, Eder F, Kurasch S, Skakalova V, Kotakoski J, Park HJ, et al. Accurate measurement of electron beam induced displacement cross sections for single-layer graphene. Phys Rev Lett. 2012;108(19). 38. Chapiro A. Radiation Chemistry of Polymeric Systems. New York: John Wiley & Sons; 1962. 39. Woods RJ, Pikaev AK. Applied Radiation Chemistry. New York: John Wiley & Sons; 1994. 40. Zhang B, Zerubia J, Olivo-Marin JC. Gaussian approximations of fluorescence microscope point-spread function models. Applied Optics. 2007;46(10):1819-29.
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Chapter 3
Length-Scale Mediated Differential Adhesion of Mammalian Cells
and Microbes
Chapter Summary Surfaces of implantable biomedical devices are increasingly engineered to promote their interactions with tissue. However, surfaces that stimulate desirable mammalian cell adhesion, spreading, and proliferation also enable microbial colonization. The biomaterials-associated infection that can result is now a critical clinical problem. We have identified an important mechanism to create a surface that can simultaneously promote healing while reducing the probability of infection. We created surfaces with submicron-sized, non-adhesive microgels patterned on an otherwise cell-adhesive surface. Quantitative force measurements between a staphylococcus and a patterned surface show that the adhesion strength decreases significantly at inter-gel spacings comparable to bacterial dimensions.Time-resolved flow-chamber measurements show that the microbial deposition rate dramatically decreases at these same spacings. Importantly, the adhesion and spreading of osteoblast-like cells is preserved despite the sub-cellular non-adhesive surface features. Since such length-scale-mediated differential interactions do not rely on antibiotics, this mechanism can be particularly significant in mitigating biomaterials-associated infection by antibiotic-resistant bacteria such as MRSA. into mapped or unreachable spaces.
3.1 Introduction Restoration of human function using implantable biomedical devices and prostheses is indispensable to modern medicine, and the surfaces of modern biomaterials are now highly Yi Wang, Guruprakash Subbiahdoss, Jan Swartjes, Henny C. van der Mei, Henk J. Busscher, and Matthew Libera*; Advanced Functional Materials 2011; 21; 3916-23
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engineered to regulate their interactions with physiological systems. However, many of the same surface properties that influence mammalian cell interactions also enable bacterial adhesion (Figure 3.1A). The subsequent biomaterials-associated-infection that can occur is now recognized as a major clinical problem. When bacteria win the race with mammalian cells to colonize an implant surface [1-4], they can develop into biofilms where they are both extremely resistant to antibiotics and able to evade the host immune system [5-9]. While antibiotics can mitigate the short-term symptoms of systemic infection, they are usually unable to resolve the localized biomaterials-associated infection. In such cases, the implant is removed, the infection is resolved over periods of week to months, and a revision surgery, with a higher probability of re-infection, is performed in a site with less native tissue [10-12]. Biomaterials-associated infection is a concern with all biomedical devices that contact tissue and is usually assumed as a given in cases involving, for example, long-term percutaneous structures or serious trauma involving large and contaminated wounds. The general problem is being increasingly exacerbated by the growing preponderance of antibiotic-resistant bacteria and the concomitant decline of new antibiotic development [13, 14]. Surface modification to mitigate bacterial colonization has largely followed one of two routes. The first incorporates antimicrobials by various drug-delivery mechanisms [15-17]. Such an approach, however, must determine a priori the appropriate antimicrobial and often delivers it when not needed, thus promoting antibiotic resistance. Consequently there is a growing focus on alternatives involving, for example, metal ions, cationic peptides, and quorum-sensing targets. A second route concentrates on the modification of the surface itself, much of which has focused on antifouling coatings. Among these, surfaces that incorporate poly(ethylene glycol) [PEG] have been extensively studied because of their ability to resist both protein adsorption and cell adhesion [18, 19]. These have served as a model for the development of other highly hydrophilic
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non-adhesive and multi-functional surfaces [20-23]. Such surfaces protect against microbial colonization, but they also resist mammalian cell adhesion (Figure 3.1B). Thus, while they can mitigate biomaterials-associated infection, they simultaneously compromise healing.
Figure 3.1 (A) Both mammalian cells and microbes can adhere to a fully cell-adhesive biomaterial surface; (B) Neither mammalian cells nor microbes can adhere to a fully nonadhesive biomaterial coating; (C) A surface patterned with submicron-sized, non-adhesive features on an otherwise cell-adhesive surface can enable mammalian-cell adhesion but reduce microbial adhesion due to their smaller size.
A fundamental problem in biomaterials science centers on how to create surfaces that differentially interact with different cell types. In the case of biomaterials-associated infection, this problem sharpens to the question of how to create a surface that differentially promotes interactions with desirable mammalian cells while simultaneously inhibiting microbial colonization [24, 25]. Significantly, there are important physico-chemical differences between mammalian cells and bacteria around which differentially interactive surfaces can be designed. For example, mammalian cells, such as osteoblasts, are typically 10-100 µm in diameter, have
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flexible cell membranes that can conform to a substratum, and adhere to surfaces through multiple, integrin-mediated submicron-sized focal contacts [26]. In contrast, microbes have welldefined shapes and are a few micrometers or less in size. In particular, the staphylococci most often implicated in biomaterials-associated infection – Staphylococcus epidermidis and Staphylococcus aureus - are spherically shaped with diameters of about 1 µm and have relatively rigid cell walls. Combining these differences with emerging concepts of surface patterning and compartmentalization [27, 28], we hypothesize that a surface whose cell adhesiveness is laterally modulated at microscopic length scales will enable mammalian cell adhesion while reducing microbial adhesion (Figure 3.1C). Using a combination of electron-beam surface patterning, quantitative staphylococcal-surface adhesive force measurement, and in situ characterization of microbe/osteoblast surface colonization, we show this hypothesis to be true: laterally modulated adhesiveness significantly reduces microbial colonization, in the absence of antibiotics, when the spacing between non-adhesive features is comparable to microbial dimensions (1-2 m), while osteoblast-like cells can nevertheless adhere to and spread over these surfaces. Such a differentially adhesive surface is one that can promote healing while simultaneously reducing the risk of infection.
3.2 Experimental Section 3.2.1 PEG Hydrogel Surface Patterning Patterned glass slides were prepared using established procedures [31, 45]. Slides were sonicated in ethanol for 5 min and dried with N2 gas. Piranha etch (3:1 98% sulfuric acid and 30% H2O2) was used for a second cleaning. After a water rinse and N2 dry, the slides were exposed to an O2 plasma (300 mTorr, 1.75 W) for 10 min and silanized with 2% [v/v] vinyl-methoxy siloxane (Gelest) in ethanol for 10 min, rinsed with ethanol and baked at 110°C (2 h). A solution of 2 wt%
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PEG (6 kDa; Fluka) in tetrahydrofuran was used to make thin films by spin casting on these slides. The film thickness after solvent evaporation was ~60 nm.
E-beam patterning used an FEI Helios SEM controlled by a Nanometer Pattern Generation System (Nabity). A typical point dose (dose per single microgel) was 10 fC (2 keV) to locally crosslink the PEG. The crosslinking of PEG under such e-beam irradiation has been attributed to a free-radical polymerization mechanism after ionization of C-H bonds [27, 45]. After exposure, the slides were washed in de-ionized water (30 min) to remove unexposed PEG. The resulting surface consisted of surface-bound microgels separated from each other by silanized glass. Patterns were made as 200 μm × 200 μm microgel arrays. Within a given array, was fixed. Three identical copies with a particular were created on each substratum. Arrays sampling inter-gel spacings from 0.5 - 8 µm were patterned on the same substratum. Each array was separated from an adjacent array by 100 µm. After patterning, substrata were stored under vacuum (50 mTorr). Prior to adhesion experiments, the substrates were placed in phosphate buffered saline (PBS) for at least 30 min.
3.2.2 Experiments with Bacteria and Yeast Bacterial inoculum Two different staphylococcal strains were used: S. aureus ATCC 12600 and S. epidermidis ATCC 35983. For each flow-chamber experiment a colony from an agar plate was inoculated into 10 mL of tryptone soya broth (TSB), cultured for 24 h at 37 ºC, and used to inoculate a second culture grown for 17 h in 200 mL TSB. Bacteria were harvested by centrifugation (6500 g) for 5 min at 10 ºC and washed twice in sterile phosphate-buffered saline (PBS; 10 mM potassium phosphate and 0.15 M NaCl, pH 7). Bacterial aggregates were broken by mild sonication on ice
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for 3 × 10 s at 30 W (Wibra Cell model 375, Sonics and Materials Inc., Danbury, Connecticut, USA) and then resuspended to a concentration of 3 × 106 bacteria/mL in sterile PBS.
Yeast cell inoculum Candida albicans GB1/2 yeast cells were used. A colony from a brain heart infusion (BHI) agar plate was inoculated in 10 mL BHI, cultured for 24 h at 37 oC, and used to inoculate a second culture grown for 17 h in 200 mL BHI. Yeast cells were harvested by centrifugation (5000 g) for 5 min at 10 oC and washed twice in sterile PBS. The harvested cells were resuspended in 10 mL sterile PBS and diluted to 3 x 106 cells/mL in PBS.
Time-resolved microbial adhesion Microbial adhesion was studied by incorporating patterned slides into a parallel-plate flow chamber [4]. Microbial deposition was monitored by digital phase-contrast microscopy (Olympus BH-2; 40x). After removing air bubbles by flowing PBS, bacteria or yeast cells in PBS were perfused through the chamber (11 s-1 shear rate) for 3 h at room temperature. Images were taken from each patterned array, and from the unpatterned silanized glass (control), at 1 min intervals. From these, the number of adhering bacteria or yeast on each array was determined. Since yeast tended to sediment, the yeast inoculum reservoir was gently stirred.
3.2.3 Mammalian Cell Experiments Mammalian cell culturing and harvesting U-2 OS osteosarcoma cells were cultured in low-glucose Dulbecco’s Modified Eagles Medium (DMEM) supplemented with 10% fetal calf serum (FBS) and 0.2 mM of ascorbic acid-2phosphate. U-2 OS cells were maintained in T75 culture flasks at 37 oC in humidified 5% CO2
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and harvested at 90% confluency using trypsine/ethylenediamine–tetraacetic acid. The harvested cells were diluted to 6 x 105 cells/mL.
Time-resolved mammalian cell adhesion U-2 OS adhesion and spreading on patterned substrata were studied by in situ imaging in the parallel-plate flow chamber and by ex situ immunofluorescence/SEM imaging. The flow chamber was kept at 37 °C. Once fully filled and bubble free, a U-2 OS cell suspension in modified medium (DMEM + 10% FBS and 2% TSB [4]) was introduced. After filling the chamber, flow was stopped (1.5 h) to allow for U-2 OS adhesion. Modified culture medium supplemented with 2% HEPES buffer was then flowed at 0.14 s-1 shear rate for 48 h. Phase-contrast images were taken from each patterned array and from the silanized-glass control at 1.5 h to determine the number of adhering cells per unit area and total surface coverage of spread U-2 OS cells. Statistical analysis of variance used Tukey’s HSD post hoc test and a p-value of
