Conducting polymers for neural interfaces: Impact of ... - UNSWorks

Conducting polymers for neural interfaces: Impact of physico-chemical properties on biological performance

Rylie Green

A Thesis Submitted for the Degree of Doctor of Philosophy At the University of New South Wales, Faculty of Engineering, Graduate School of Biomedical Engineering

January 7, 2009

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Acknowledgements A number of people have been essential to the progress of my PhD research and completion of my dissertation, contributing their time, technical assistance, knowledge and support. Firstly, I am indebted to my supervisors Nigel Lovell and Laura Poole-Warren for taking a risk on a new project and supporting the development of a new laboratory with infrastructure to produce and characterise novel materials. Nigel’s advice and experience has been invaluable and I am exceedingly grateful for his support which allowed me the independence and resources to develop results. Laura’s willingness to join the project, contributing her expertise in the field provided direction to the project, instrumental to it’s progress and completion. I would like to thank both supervisors for their encouragement to gain experience in an overseas laboratory, opportunities to attend international conferences, and finally for their continued support and friendship throughout the research period, despite numerous extensions. For contributions to the everyday progression of my research Phil Preston must be mentioned and sincerely thanked. Phil could always be found when a piece of equipment needed servicing and was always willing to help. Specifically I’d like to thank Phil for designing and fabricating a number of “black box” devices and pieces of surprisingly effective equipment made from laboratory consumables (mostly pipette tips, alligator clips and perspex), that made the fabrication and testing of materials efficient and easy. The final dissertation would not be complete without significant effort from Ross Odell, for analysing, interpreting and producing statistics which would have taken me years to equal. Thank you Ross for your persistence and tolerance of my shrinking deadlines. I would to also like to thank the remaining staff and my other research colleagues at GSBmE, many of whom have offered advice, reviewed my work and helped out both in the laboratory and the office. The willingness of others to help both in GSBmE and across UNSW in general has made my 9 years of tertiary study a pleasure and resulted in many ongoing friendships. Specifically, I’d like to thank my lab companions, lunch buddies and desk neighbours at GSBmE, Pinki, Mai and Charles for their friendly support and ability to listen. Finally, I would like to thank my family, especially my parents for supporting me through my schooling and helping me into a position where I have had the opportunity to produce a PhD. Most of all, I would like to thank my husband for his belief in me, constant support, pursuit of understanding of my work (despite working in finance) and above all, his big bear hugs that make everything better. Oh, and thank you to Meatball for keeping my feet warm while writing and giving me an excuse to procrastinate and go for long walks!

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Originality Statement

’I hereby declare that this submission is my own work and to the best of my knowledge it contains no materials previously published or written by another person, or substantial proportions of material which have been accepted for the award of any other degree or diploma at UNSW or any other educational institution, except where due acknowledgment is made in the thesis. Any contribution made to the research by others, with whom I have worked at UNSW or elsewhere, is explicitly acknowledged in the thesis. I also declare that the intellectual content of the thesis is the product of my own work, except to the extent that the assistance from others in the project’s design and conception or in style, presentation and linguistic expression is acknowledged.’ Signed: Dated:

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Copyright Statement

‘I hereby grant the University of New South Wales or its agents the right to archive and to make available my thesis or dissertation in whole or part in the University libraries in forms of media, now or here after known, subject to the provisions of the Copyright Act 1928. I retain all proprietary rights, such as patent rights. I also retain the right to use in future works (such as articles or books) all or part of this thesis or dissertation. I also authorise Univeristy Microfilms to use the 350 word abstract of my thesis in Dissertation Abstract international. I have either used no substantial portions of copyright material in my thesis or I have obtained permission to use copyright material; where permission has not been granted I have applied/will apply for a partial restriction of the digital copy of my thesis or dissertation.’ Signed: Dated:

Authenticity Statement

’I certify that the Library deposit digital copy is a direct equivalent of the final officially approved version of my thesis. No emendation of content has occured and if there are any minor variations in formatting, they are the result of the conversion to digital format.’ Signed: Dated:

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Abstract

This research investigates the use of conducting polymer coatings on platinum (Pt) electrodes for use in neuroprostheses. Conducting polymers aim to provide an environment conducive to neurite outgrowth and attachment at the electrode sites, producing intimate contact between neural cells and stimulating electrodes. Conducting polymers were electropolymerised onto model Pt electrodes. Conventional polymers polypyrrole (PPy) and poly-3,4-ethylenedioxythiphene (PEDOT) doped with polystyrenesulfonate (PSS) and para-toluenesulfonate (pTS)were investigated. Improvement of material properties was assessed through the layering of polymers with multiwalled carbon nanotubes (MWNTs). The ability to incorporate cell attachment bioactivity into polymers was examined through the doping of PEDOT with anionic laminin peptides DCDPGYIGSR and DEDEDYFQRYLI. Finally, nerve growth factor (NGF), was entrapped in PEDOT during polymerisation and tested for neurite outgrowth bioactivity against the PC12 cell line. Each polymer modification was assessed for electrical performance over multiple reduction-oxidation cycles, conductivity and impedance spectroscopy, mechanical adherence and hardness, and biological response. Scanning electron microscopy was used to visualise film topography and x-ray photoelectron spectroscopy was employed to examine chemical constitution of the polymers. For application of electrode coatings to neural prostheses, optimal bioactive conducting polymer PEDOT/pTS/NGF was deposited on electrode arrays intended for implantation. PC12s were used to assess the bioactivity of NGF functionalised PEDOT when electrode size was micronised. Flexibility of the design was tested by tailoring PEDOT bioactivity for the cloned retinal ganglion cell, RGC-5, differentiated via staurasporine. It was established that PEDOT films had superior electrical and cell growth characteristics, but only PPy was able to benefit from incorporation of MWNTs. Bioactive polymers were produced through inclusion of both laminin peptides and NGF, but the optimum film constitution was found to be PEDOT doped with pTS with NGF entrapped during electrodeposition. Application of this polymer to an implant device was confirmed through positive neurite outgrowth on vision prosthesis electrode arrays. The design was shown to be flexible when tailored for RGC-5s, with differentiation occurring on both PEDOT/pTS and PEDOT/DEDEDYFQRYLI. Conducting polymers demonstrate the potential to improve electrode-cell interactions. Future work will focus on the effect of electrical stimulation and design of bioactive polymers with improved cell attachment properties.

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Contents Glossary of Abbreviations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . xvii

1 Introduction

1

1.1

Research Motive . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1

1.2

Project Overview, Hypotheses and Aims . . . . . . . . . . . . . . . . . . .

3

1.3

Summary and Thesis Structure . . . . . . . . . . . . . . . . . . . . . . . .

5

2 Background Information

8

2.1

Current Neuroprosthetic Designs and Applications . . . . . . . . . . . . .

8

2.2

Conducting Polymers for Neural Interfaces . . . . . . . . . . . . . . . . . .

11

2.3

Criteria for Assessment and Challenges to Viability . . . . . . . . . . . . .

17

2.3.1

Electrical Characterisation . . . . . . . . . . . . . . . . . . . . . .

17

2.3.2

Mechanical Behaviour . . . . . . . . . . . . . . . . . . . . . . . . .

21

2.3.3

Assessment and Impact of Chemical Composition . . . . . . . . . .

23

2.3.4

Biological Response . . . . . . . . . . . . . . . . . . . . . . . . . .

25

Summary and Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . .

30

2.4

iii

CONTENTS

iv

I

Design of Conducting Polymer Electrode Coatings

3 Synthesis and Characterisation of CPs

31

32

3.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

32

3.2

Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

36

3.2.1

Modelling of the Electrode Interface . . . . . . . . . . . . . . . . .

36

3.2.1.1

The Model Electrode . . . . . . . . . . . . . . . . . . . .

36

3.2.1.2

The Model Cell . . . . . . . . . . . . . . . . . . . . . . .

37

3.2.2

Conductive Polymer Constituents

. . . . . . . . . . . . . . . . . .

39

3.2.3

Electropolymerisation . . . . . . . . . . . . . . . . . . . . . . . . .

40

3.2.3.1

Scanning Electron Microscopy . . . . . . . . . . . . . . .

43

3.2.3.2

X-ray Photoelectron Spectroscopy . . . . . . . . . . . . .

44


44

3.2.4.1

Electrochemical Stability . . . . . . . . . . . . . . . . . .

45

3.2.4.2

Conductivity . . . . . . . . . . . . . . . . . . . . . . . . .

46

3.2.4.3

Impedance . . . . . . . . . . . . . . . . . . . . . . . . . .

48

3.2.4

3.2.5

3.2.6

Mechanical Characterisation

. . . . . . . . . . . . . . . . . . . . .

50

3.2.5.1

Film Hardness . . . . . . . . . . . . . . . . . . . . . . . .

50

3.2.5.2

Film Adhesion . . . . . . . . . . . . . . . . . . . . . . . .

51

In Vitro Neurite Outgrowth Assays

. . . . . . . . . . . . . . . . .

52

3.2.6.1

Neurite Length Measurement . . . . . . . . . . . . . . . .

54

3.2.6.2

Neurite Data Analysis and Statistical Evaluation . . . . .

54

CONTENTS 3.3

v

Experimental Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

55

3.3.1


60

3.3.2


64

3.3.3


. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

66

3.4

Summary of Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

70

3.5

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

71

3.6

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

75

4 Improving CPs through MWNT layering

77

4.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

77

4.2

Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

81

4.2.1

MWNT Conducting Polymer Composites . . . . . . . . . . . . . .

81

Experimental Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

82

4.3.1


87

4.3.2


91

4.3.3


4.3

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

95

4.4

Summary of Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

99

4.5

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

99

4.6

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 103

5 Cell Attachment Bioactivity

105

5.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 105

5.2

Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 109

CONTENTS

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5.3

5.2.1

Electropolymerisation . . . . . . . . . . . . . . . . . . . . . . . . . 109

5.2.2

Analytical Techniques . . . . . . . . . . . . . . . . . . . . . . . . . 110

5.2.3

In Vitro Cell Studies . . . . . . . . . . . . . . . . . . . . . . . . . . 111

Experimental Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 113 5.3.1

Electrical Characterisation . . . . . . . . . . . . . . . . . . . . . . 117

5.3.2


5.3.3


. . . . . . . . . . . . . . . . . . . . . 119 . . . . . . . . . . . . . . . . . 121

5.4

Summary of Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 126

5.5

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 126

5.6

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133

6 Neurotrophic Bioactivity

134

6.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134

6.2


6.3

6.4

6.2.1

Electropolymerisation . . . . . . . . . . . . . . . . . . . . . . . . . 138

6.2.2

Analytical Techniques . . . . . . . . . . . . . . . . . . . . . . . . . 139

6.2.3

In Vitro Cell Studies . . . . . . . . . . . . . . . . . . . . . . . . . . 139


Electrical Characterisation . . . . . . . . . . . . . . . . . . . . . . 144

6.3.2


6.3.3


. . . . . . . . . . . . . . . . . . . . . 148 . . . . . . . . . . . . . . . . . 152

Summary of Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156

CONTENTS

II

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6.5

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156

6.6

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 162

Application of CP Electrode Coatings to Neuroprostheses

7 Application of CP Coatings to Microelectrodes

165

166

7.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166

7.2


7.3

7.2.1

Fabrication of AVPG Microelectrodes . . . . . . . . . . . . . . . . 169

7.2.2

L929 Growth Inhibition Assay . . . . . . . . . . . . . . . . . . . . 171

7.2.3

Neurite Outgrowth on PEDOT/pTS/NGF Coated Microelectrodes 173


L929 Growth Inhibition . . . . . . . . . . . . . . . . . . . . . . . . 175

7.3.2

Neurite Outgrowth on PEDOT/pTS/NGF Coated Microelectrodes 175

7.4

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 181

7.5

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 184

8 Tailoring CP Bioactivity to an Alternate Cell Type

185

8.1

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185

8.2


8.3

8.2.0.1

Tailoring Polymer for RGC-5 Bioactivity . . . . . . . . . 190

8.2.0.2

RGC-5 Differentiation Assay . . . . . . . . . . . . . . . . 191

Experimental Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 194

CONTENTS

viii 8.4

Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 197

8.5

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200

9 Conclusions and Recommendations

201

9.1

Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201

9.2

Recommendations for Future Work . . . . . . . . . . . . . . . . . . . . . . 210

Appendices

215

A Effect of Dopant Concentration on Electrochemistry

216

A.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 216 A.2 Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 217 A.3 Results and Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 218 A.4 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 219

B Specifications for 12-Channel Galvanostat

221

C Effect of NGF concentration on Neurite Outgrowth

224

C.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 224 C.2 Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 224 C.3 Results and Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 226 C.4 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 228

D Statistical analysis of In Vitro Cell Studies

229

D.1 Design of Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 229

CONTENTS D.2 Preliminary Data Treatment

ix . . . . . . . . . . . . . . . . . . . . . . . . . 233

D.3 ANOVA for Basic Polymers . . . . . . . . . . . . . . . . . . . . . . . . . . 233 D.3.1 Cell Density . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 234 D.3.2 Neurite Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 234 D.4 ANOVA for MWNT Composite Polymers . . . . . . . . . . . . . . . . . . 236 D.4.1 Cell Density . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 236 D.4.2 Neurite Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 238 D.5 ANOVA for Laminin Peptide Doped PEDOT . . . . . . . . . . . . . . . . 239 D.5.1 Cell Density . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 239 D.5.2 Neurite Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 241 D.6 ANOVA for NGF loaded PEDOT . . . . . . . . . . . . . . . . . . . . . . . 242 D.6.1 Cell Density . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 243 D.6.2 Neurite Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 243 D.7 ANOVA for PEDOT/pTS/NGF coated AVPG electrode array . . . . . . 245

E Effect of Peptide Concentration on Electrochemistry

246

E.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 246 E.2 Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 246 E.3 Results and Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 248 E.4 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 249

F Effect of SS concentration on RGC-5 Differentiation

250

F.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 250

CONTENTS

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F.2 Materials and Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 251 F.3 Results and Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 253 F.4 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 254

Bibliography

254

List of Figures

2.1

AVPG epiretinal electrode array . . . . . . . . . . . . . . . . . . . . . . .

10

2.2

Structures of conducting polymers . . . . . . . . . . . . . . . . . . . . . .

13

2.3

Electrochemical behaviour of PPy . . . . . . . . . . . . . . . . . . . . . .

14

3.1

Custom manufactured well sandwich assembly . . . . . . . . . . . . . . . .

38

3.2

Sulfonate dopants: i. PSS and ii. pTS . . . . . . . . . . . . . . . . . . . .

40

3.3

Electropolymerisation set up . . . . . . . . . . . . . . . . . . . . . . . . .

41

3.4

Setup used for EIS . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

48

3.5

ASTM calibrated pencil hardness scale . . . . . . . . . . . . . . . . . . . .

50

3.6

SEMs of Basic PPy and PEDOT films . . . . . . . . . . . . . . . . . . . .

56

3.7

XPS of Basic PPy and PEDOT films . . . . . . . . . . . . . . . . . . . . .

57

3.8

Detailed spectra of sulfur in PEDOT . . . . . . . . . . . . . . . . . . . . .

59

3.9

Successive CV curves for PPy and PEDOT . . . . . . . . . . . . . . . . .

61

3.10 Loss of electroactivity for basic polymers . . . . . . . . . . . . . . . . . . .

62

3.11 Bode plot of impedance spectroscopy for basic polymers . . . . . . . . . .

65

3.12 Calibrated hardness for basic polymers . . . . . . . . . . . . . . . . . . . .

66

xi

LIST OF FIGURES

xii

3.13 Sample of x-cut images taken from basic polymers . . . . . . . . . . . . .

67

3.14 Basic film delamination by ASTM x-cut analysis . . . . . . . . . . . . . .

67

3.15 PC12 response to basic polymer substrates at 96 hr . . . . . . . . . . . . .

68

3.16 Sample images of PC12 neurite outgrowth on basic polymers . . . . . . .

69

4.1

SEMs of air-dried MWNTs . . . . . . . . . . . . . . . . . . . . . . . . . .

83

4.2

SEMs of layered composites of MWNTs and PPy . . . . . . . . . . . . . .

84

4.3

SEMs of layered composites of MWNTs and PEDOT . . . . . . . . . . . .

85

4.4

XPS of MWNT modified conducting polymer . . . . . . . . . . . . . . . .

86

4.5

Loss of electroactivity for MWNT composites . . . . . . . . . . . . . . . .

88

4.6

Example of a typical MWNT-PPy composite undergoing 400 cycles of CV

89

4.7

Bode plot of impedance spectroscopy for MWNT composite polymers . .

92

4.8

MWNT composite polymer hardness . . . . . . . . . . . . . . . . . . . . .

93

4.9

MWNT composite polymer delamination by x-cut test . . . . . . . . . . .

94

4.10 Sample of x-cut images taken from MWNT composite polymers . . . . . .

95

4.11 Sample images of PC12 neurite outgrowth on MWNT composite polymers

97

4.12 PC12 cell density on MWNT-polymer substrates . . . . . . . . . . . . . .

98

4.13 Neurite outgrowth of PC12s on MWNT-polymer composites . . . . . . . .

98

5.1

Electrode schematic with cell attachment bioactivity . . . . . . . . . . . . 105

5.2

SEMs of PEDOT doped with laminin peptides . . . . . . . . . . . . . . . 114

5.3

XPS of laminin doped PEDOT . . . . . . . . . . . . . . . . . . . . . . . . 116

5.4

Loss of electroactivity for PEDOT doped with laminin peptides . . . . . . 118

LIST OF FIGURES

xiii

5.5

Bode plot of impedance spectroscopy for laminin peptide doped PEDOT . 120

5.6

Laminin peptide doped polymer hardness . . . . . . . . . . . . . . . . . . 121

5.7

Laminin peptide doped polymer delamination by x-cut test . . . . . . . . 122

5.8

Sample of x-cut images taken from laminin peptide doped PEDOT . . . . 122

5.9

Sample images of PC12 neurite outgrowth on laminin peptide doped PEDOT124

5.10 PC12 cell density on laminin peptide doped PEDOT . . . . . . . . . . . . 125 5.11 Neurite outgrowth of PC12s on laminin peptide doped PEDOT . . . . . . 125 5.12 Redox behaviour of conducting polymers with large dopants . . . . . . . . 129

6.1

Bioactive electrode . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 135

6.2

SEMs of NGF modified PEDOT films . . . . . . . . . . . . . . . . . . . . 142

6.3

XPS of PEDOT with NGF incorporated . . . . . . . . . . . . . . . . . . . 143

6.4

Sample CV curve from PEDOT/DEDEDYFQRYLI/NGF . . . . . . . . . 146

6.5

Loss of electroactivity for PEDOT loaded with NGF . . . . . . . . . . . . 147

6.6

Bode plot of impedance spectroscopy for NGF loaded PEDOT . . . . . . 149

6.7

Effect of NGF loading on polymer hardness . . . . . . . . . . . . . . . . . 150

6.8

Delamination of NGF loaded polymers by x-cut test . . . . . . . . . . . . 151

6.9

Sample of x-cut images taken from NGF loaded polymers . . . . . . . . . 151

6.10 Sample images of PC12 neurite outgrowth on NGF loaded polymers . . . 153 6.11 PC12 cell density on peptide doped PEDOT + NGF . . . . . . . . . . . . 155 6.12 Neurite outgrowth of PC12s on laminin peptide doped PEDOT + NGF . 155

7.1

AVPG vision prosthesis schematic . . . . . . . . . . . . . . . . . . . . . . 167

LIST OF FIGURES

xiv 7.2

AVPG 98-electrode array for high resolution vision prostheses [1]. . . . . . 168

7.3

Well configuration for L929 growth inhibition assay . . . . . . . . . . . . . 171

7.4

L929 growth on electrode materials at 48 hr . . . . . . . . . . . . . . . . . 176

7.5

Normalised L929 growth inhibtion of electrode materials . . . . . . . . . . 177

7.6

PEDOT/pTS/NGF electrodeposited on AVPG electrode array. . . . . . . 178

7.7

PEDOT/pTS/NGF on AVPG electrode array with PC12s at 96 hr . . . . 179

7.8

Cell density and neurite outgrowth of PC12s at 96 hr on AVPG electrodes 180

8.1

Cell structure of the human retina . . . . . . . . . . . . . . . . . . . . . . 187

8.2

Structure of staurosporine (SS) molecule [2] . . . . . . . . . . . . . . . . . 191

8.3

Sample images of RGC-5 differentiation on SS tailored polymers . . . . . 195

8.4

RGC-5 cell density on SS tailored PEDOT . . . . . . . . . . . . . . . . . . 196

8.5

Neurite outgrowth of RGC-5s on PEDOT loaded with SS . . . . . . . . . 197

A.1 Electroactivity loss related to dopant concentration . . . . . . . . . . . . . 219

B.1 12-channel in-house manufactured galvanostat . . . . . . . . . . . . . . . . 223

C.1 NGF concentrations applied to PC12s . . . . . . . . . . . . . . . . . . . . 226 C.2 GFP-PC12 neurite outgrowth at various concentrations of NGF at 96 hr . 227

E.1 Electroactivity loss related to peptide dopant concentration . . . . . . . . 248

F.1 Concentrations of SS and glutamate for detection of RGC-5 differentiation 252 F.2 Optical density of NR following glutamate toxicity of RGC-5s . . . . . . . 253

List of Tables

2.1

Conductivity of common conducting polymers . . . . . . . . . . . . . . . .

18

2.2

Impedance magnitude at 1 kHz on 1250 μm2 electrodes . . . . . . . . . .

19

2.3

Toxicity limits of common polymer constituents . . . . . . . . . . . . . . .

27

3.1

Basic polymer solutions . . . . . . . . . . . . . . . . . . . . . . . . . . . .

42

3.2

Doping ratio of basic polymers calculated from XPS . . . . . . . . . . . .

60

3.3

Percentage of electroactivity loss for basic conducting polymers . . . . . .

63

3.4

Basic polymer conductivities . . . . . . . . . . . . . . . . . . . . . . . . .

63

3.5

Summary of results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

70

3.6

Impedance of basic conducting polymer films on Pt at 1 kHz. . . . . . . .

73

4.1

Doping ratios of MWNT layered polymers . . . . . . . . . . . . . . . . . .

86

4.2

Electroactivity losses for MWNT composites

. . . . . . . . . . . . . . . .

87

4.3

Conductivity of MWNT layered conducting polymer films . . . . . . . . .

90

4.4

Impedance of MWNT composite polymers at 1 kHz . . . . . . . . . . . .

91

4.5

Summary of results for MWNT layered conducting polymers . . . . . . .

99

xv

LIST OF TABLES

xvi 5.1

Laminin peptide domains and their associated bioactivity . . . . . . . . . 108

5.2

PEDOT/peptide electrolyte solutions . . . . . . . . . . . . . . . . . . . . . 110

5.3

XPS of laminin peptide doped PEDOT . . . . . . . . . . . . . . . . . . . 117

5.4

Electroactivity loss for PEDOT doped with laminin peptides . . . . . . . 117

5.5

Impedance of PEDOT doped with laminin peptides

5.6

Summary of results for laminin doped PEDOT . . . . . . . . . . . . . . . 126

6.1

PEDOT/peptide/NGF electrolyte solutions . . . . . . . . . . . . . . . . . 138

6.2

Elemental analysis of PEDOT loaded with NGF . . . . . . . . . . . . . . 144

6.3

Comparison of electroactivity losses for NGF loaded PEDOT . . . . . . . 146

6.4

Impedance of PEDOT loaded with NGF at 1 kHz . . . . . . . . . . . . . 148

6.5

Summary of results for NGF loaded PEDOT . . . . . . . . . . . . . . . . 156

8.1

Electrolyte solutions tailored for RGC-5 differentiation activity. . . . . . . 191

9.1

Summary of results for conducting polymers across assessed criteria . . . 202

. . . . . . . . . . . . 119

A.1 Basic polymer solutions with varied dopant concentration . . . . . . . . . 217

D.1 A two-way layout . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 230 D.2 A one-way layout . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 230 D.3 ANOVA table for the two-way layout . . . . . . . . . . . . . . . . . . . . . 231 D.4 ANOVA table for the one-way layout with pooled errors . . . . . . . . . . 232 D.5 ANOVA table for the randomised block design . . . . . . . . . . . . . . . 232 D.6 ANOVA table for basic mean cell number . . . . . . . . . . . . . . . . . . 234

LIST OF TABLES

xvii

D.7 Multiple comparison matrix for basic study mean cell density . . . . . . . 235 D.8 ANOVA table for basic neurite length . . . . . . . . . . . . . . . . . . . . 235 D.9 Multiple comparison matrix for basic study neurite length . . . . . . . . . 235 D.10 ANOVA table for mean cell density of MWNT layered polymers . . . . . 237 D.11 ANOVAs for the difference in mean cells density for MWNT composites . 237 D.12 Comparison matrix for cell densities in MWNT layering study . . . . . . . 237 D.13 ANOVA table for neurite length on MWNT layered polymers . . . . . . . 238 D.14 ANOVA of the difference in neurite length on MWNT-polymers . . . . . . 238 D.15 Comparison matrix for neurite lengths in MWNT layering study . . . . . 239 D.16 ANOVA table for cell density on peptide doped PEDOT . . . . . . . . . . 240 D.17 Comparison matrix for cell densities in peptide doping study . . . . . . . 241 D.18 Two-way ANOVA for neurite length on peptide doped PEDOT . . . . . . 241 D.19 One-way ANOVA for neurite length on peptide doped PEDOT . . . . . . 242 D.20 Comparison matrix for neurite length in peptide doped PEDOT study . . 242 D.21 ANOVA table for cell density on NGF loaded PEDOT . . . . . . . . . . . 243 D.22 ANOVA table for neurite length on NGF loaded PEDOT . . . . . . . . . 244 D.23 ANOVA table for the difference in neurite lengths on NGF loaded PEDOT 244 D.24 Comparison matrix for neurite length in NGF loading study . . . . . . . . 244 D.25 Analysis of variance for mean cell densities on AVPG electrodes . . . . . . 245 D.26 Analysis of variance for mean neurite lengths on AVPG electrodes . . . . 245

E.1 PEDOT/peptide electrolyte solutions . . . . . . . . . . . . . . . . . . . . . 247

LIST OF TABLES

xviii

Glossary of Abbreviations AFM AMD ANOVA ASIC AVPG BDNF BS C CNT CNTF CP CV DCDPGYIGSR

: : : : : : : : : : : : :

DEDEDYFQRYLI

:

DI DMEM DPBS EDOT EIS EPAni EtOH FBS FCS FTIR FTIR-MM GDNF GFP-PC12

: : : : : : : : : : : : :

GSBmE HA HS HSA HTD ITO LD50

: : : : : : :

MWNT N NGF

: : :

Atomic force microscopy Age-related macular degeneration Analysis of variance Application specifc integrated circuit Australian Vision Prosthesis Group Brain derived neurotrophic factor Benzene sulfonate Carbon Carbon nanotube Ciliary neutrophic factor Conducting polymer Cyclic voltammetry Synthetic laminin peptide comprised of amino acid sequence: Asp-Cys-Asp-Pro-Gly-Tyr-Ile-Gly-Ser-Arg Synthetic laminin peptide comprised of amino acid sequence: Asp-Glu-Asp-Glu-Asp-Tyr-Phe-Gln-Arg-Tyr-Leu-Ile Deionised Dulbecco’s modified Eagles medium Dulbecco’s phosphate buffered saline 3,4-ethylene dioxythiophene (monomer unit of PEDOT) Electrochemical impedance spectroscopy Emeraldine salt modified polyaniline Ethanol Fetal bovine serum Fetal calf serum Fourier transform infra red spectroscopy Fourier transform infrared microscopy with mapping Glial cell line derived neurotrophic factor Green fluorescent protein transfected PC12, visible by fluorescent reflection microscopy between 509 and 540 nm Graduate School of Biomedical Engineering, UNSW Hyaluronic acid Horse serum Human serum albumin Highest tolerated dose Indium tin oxide Lethal dose used to determine toxicity limits of chemical reagents Multi-walled carbon nanotubes Nitrogen Nerve growth factor

LIST OF TABLES

xix

NI NIH NO NR NSD NT-3 NT-4/5 O PAni PBS PC12 PDGF PDMS PEDOT PPy PSS Pt PTh pTS PU Redox RF RGC RGC-5 RP RPMI S SE SEM SGN SIDNE SLPF SNR SS TCP UNSW UNTHSC XPS YFQRYLI

: : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : : :

YIGSR

:

National Instruments Pty. Ltd. National Institute of Health, USA Nitric oxide Neutral red No significant difference Neurotrophin-3 Neurotrophin-4 or neurotrophin-5 Oxygen Polyaniline Phosphate buffered saline Cloned pheochromocytoma cell line Platelet derive growth factor Polydimethylsiloxane Poly-3,4-ethylenedioxythiphene Polypyrrole Polystyrenesulfonate Platinum Polythiophene Para-toluenesulfonate Polyurethane Reduction and oxidation Radio frequency Retinal ganglion cell Retinal ganglion cell type-5 clonal cell line Retinitis pigmentosa Roswell Park Memorial Institute medium Sulfur Standard error of the mean Scanning electron microscopy Spiral ganglion neurons Stimulus induced depression of neuronal excitability Silk-like polymer having fibronectin fragments Signal-to-noise ratio Staurosporine Tissue culture plastic University of New South Wales University of North Texas, Health and Sciences Centre X-ray photoelectron spectroscopy Laminin peptide active ligand comprised of amino acid sequence: Tyr-Phe-Gln-Arg-Tyr-Leu-Ile Laminin peptide active ligand comprised of amino acid sequence: Tyr-Ile-Gly-Ser-Arg

xx

LIST OF TABLES

Chapter 1

Introduction 1.1

Research Motive

Neuroprosthetic electrode configurations are designed to support functional restoration of a neural path, but many current designs are limited in their capacity for long-term stimulation and recording. In vivo, cortical cell responses to stimulation by microelectrodes degrade over a period of 100 days [3] and recordings made from cat sciatic nerve have been effective for a maximum of 7 months [4]. In the cochlear implant, the use of high stimulus rates intended to improve patient speech perception has been reported to significantly reduce neuronal excitability measured by electrically evoked auditory brainstem response [5].

The observed decrease in electrode performance can occur through two possible mechanisms. Firstly, in both recording and stimulating implants the impedance at the electrode interface can increase as a result of fibrous encapsulation associated with foreign body re1

2

CHAPTER 1. INTRODUCTION

actions, reducing the effective transfer of a signal [6]. Additionally, in stimulating implants a phenomenon described by McCreery and colleagues as stimulus induced depression of neuronal excitability (SIDNE), has been used to explain elevation of the neuronal firing threshold as time progresses [7]. The neuronal mechanisms associated with SIDNE are thought to be dependant on stimulus frequency, magnitude and selectivity [7, 8]. Studies conducted by Shepherd and colleagues using a cochlea implant stimulating auditory neurons shows concurring results, suggesting that the increase in threshold may be a result of an activity-induced depletion of neural energy resources required to maintain homeostasis [5, 9].

Electrodes are fabricated from a variety of conductive materials including gold, platinum, iridium oxide and glassy carbon. While all of these materials are considered safe, they are typically fabricated with smooth surfaces that are not conducive to tissue integration. In the case of metal electrodes, release of ions can also impact on tissue response. As a result the interface between a metal electrode and the tissue it stimulates/records is commonly associated with a significant fluid gap through which the electrical signal must be transduced. In stimulating applications, this distance from electrode to neural cells is critical in determining the current/charge threshold required for cell activation [10]. Similarly, the distance between the recording electrodes and neural tissue has a significant influence on the quality of signal obtained [11, 12]. A more desirable electrode material would support intimate contact between the tissue and electrode, minimising the attenuation of the signal within the extracellular fluid.

1.2. PROJECT OVERVIEW, HYPOTHESES AND AIMS

1.2

3

Project Overview, Hypotheses and Aims

This thesis addresses issues related to the long-term performance of stimulating electrodes through the use of surface modifications, specifically the coating of electrodes with conductive polymers. The research focuses on studying two polymers, polypyrrole (PPy) and poly(3,4-ethylene dioxythiophene) (PEDOT), doped with various chemical and biochemical dopants. It has been shown that this approach effectively stabilises the tissue-electrode interface, dramatically improves the signal-to-noise ratio (SNR) and enhances long-term performance of the neural implant [11, 13–15]. Additionally, conductive polymers loaded with neurotrophins have been used to successfully cause outgrowth of neurites from spiral ganglion neurons (SGNs) explanted from the cochlea [16]. While several research groups have investigated the use of conductive polymers in neural interfacing and nerve regeneration, there has been little published on the effect of these modifications on the mechanical performance and long-term electrochemical stability of the electrodes.

The ultimate objective of this project is to develop bioactive electrode interfaces with softer and more compliant properties than typical metal electrodes. The key to improving neuroprosthetic interfaces is in eliminating the distance between the electrode and the cell being stimulated. Bioactive properties such as cell attachment and neural cell outgrowth, incorporated within the electrode surface will produce an intimate and durable connection between the electrode and the surrounding tissue. As a consequence the current magnitude required for cell stimulation and spread of current to adjacent neurons would be reduced, producing a more effective charge transfer from the implant to the


4

biological system. Additionally, it is theorised that softer materials, such as conducting polymers, would reduce inflammation caused by strain mismatch at the tissue-electrode interface and subsequently minimise the foreign body reaction.

This research was motivated by three main hypotheses:

• Conducting polymers can provide a coating for typical neuroprosthetic metal electrodes which address the issues associated with poor cell interaction without compromising important physico-chemical or biological properties. • Doping conducting polymers with cell adhesion ligands can produce an interface where neural cells readily attach to the implant material, without compromising the electrical and mechanical stability of the electrode. • Incorporating neurotrophins will produce an interface with bioactivity to encourage neurite outgrowth from the surrounding cells across the electrode surface, further encouraging interactions between the cell and electrode.

This research explores the potential of bioactive conductive polymers in neural interface technology, encompassing the need to determine the effect of the various combinations of monomer, dopant and biomolecular inclusions on mechanical, electrical and biological properties of the electrode interface. The specific aims of this project were to:

1. Consistently prepare PPy and PEDOT conducting polymers using galvanostatic electrodeposition to produce electrode coatings.

1.3. SUMMARY AND THESIS STRUCTURE

5

2. Measure the influence of the various monomers, dopants and bioactive molecular incorporations on the physico-chemical properties of the polymer 3. Examine the relationship between polymer constitution (monomer, dopant and biomolecule inclusions) and in vitro cell response, using a cloned mammalian neural cell line. 4. Assess the bioactivity for both cell adherence and neurite outgrowth characteristics of the conductive polymer composite, and thus the ability of these films to produce an intimate connection at the neural interface.

1.3

Summary and Thesis Structure

To summarise, conducting polymers are materials that possess the ability to conduct electricity through their inherently unstable backbone. They require doping to stabilise the backbone in the passive state and retain conductivity. Current conducting polymer modifications of electrodes are not optimised for long-term performance in neuroprosthetic devices, such as the bionic eye. Additionally, the available body of literature does not provide crucial information on the performance of polymers across the range of criteria required for adequate assessment. The purpose of this project was to optimise the performance of conducting polymers across electrical, mechanical, chemical and biological properties. Commonly explored polymers were analysed across the specified criteria and then optimised through the incorporation of MWNTs and bioactive molecules. The remaining chapters of this thesis contain further review of the literature and are briefly

6


described below:

Chapter 2 reviews the literature on conducting polymers for neural interfaces, highlighting the current challenges in developing an effective long-term implant.

Chapter 3 explores the adaption of electropolymerisation of conducting polymers common to the literature and analysis across a range of criteria considered integral to the long-term performance of the material. A number of analytical techniques that are not reported in the literature are used to gain a thorough depiction of polymer performance.

Chapter 4 reviews the current limitations of common conducting polymers. Films are optimised through the incorporation of MWNTs and assessed using analytical techniques established in the previous chapter.

Chapter 5 reviews current technologies used to incorporate cell attachment bioactivity to conducting polymer films intended for neuroprosthetic devices. Laminin peptides are used to impart cell attachment bioactivity to conducting polymers and the effect of these molecules on film properties is explored. The efficacy of peptides is assessed through in vitro cell culture assays.

Chapter 6 presents the background information on neurotrophic bioactivity and reviews the current application of specific neurotrophins to implant devices. Both cell attachment molecules explored in the previous chapter and neurotrophins are introduced to conducting polymers. The effect of adding multiple bioactive molecules on polymer physico-

1.3. SUMMARY AND THESIS STRUCTURE

7

chemical properties is investigated. The efficacy of the neurotrophin is demonstrated through in vitro cell assays using a model neural cell line.

Chapter 7 demonstrates the application of conducting polymer coatings to microelectrodes more typical of neuroprostheses than the model electrode used in previous chapters. The optimal film constitution assessed in Part I, PEDOT/pTS/NGF, is applied to the vision prosthesis electrode array developed by the AVPG and assessed with in vitro cell assays.

Chapter 8 details the challenges associated with tailoring conducting polymer coatings for specific cells and tissues. The vision prosthesis application is again used and the target cell is the retinal ganglion cell. In this chapter the conducting polymer bioactivity is altered to be specific for differentiation of the RGC-5 clonal cell line.

Chapter 9 presents the conclusions and recommendations arising from these studies.

Chapter 2

Background Information

[Portions of this chapter are reprinted from R.A. Green, N.H. Lovell, G.G. Wallace and L.A. Poole-Warren, Conducting polymers for neural interfaces: Challenges in developing an effective long-term implant with permissions from Biomaterials; 29 (2008) 3393 - 3399; DOI:10.1016. Accepted 28 April; Copyright 2008 Elsevier Ltd.]

2.1

Current Neuroprosthetic Designs and Applications

Bioelectrodes are critical elements in numerous medical devices and are typically used for biopotential recording or for neurostimulation. The cochlear implant and bionic eye are examples of stimulating implants. While these prostheses are used to electrically evoke responses in specific neural tissue through controlled stimulation paradigms, they can also perform recording functions that output information on the status of the neural-tissue interface. Current devices used for neural recording are typically more research-based and not in the commercial domain. However, with intensive research underway in the 8

2.1. CURRENT NEUROPROSTHETIC DESIGNS AND APPLICATIONS

9

areas of brain-computer interfacing and nerve interfacing for prosthetic control purposes, there will certainly be a large demand for neural recording devices in future years.

The cochlear implant is a commercially available implant that has restored hearing percepts to patients suffering from sensorineural deafness for several decades. The stimulating portion of the implant is a flexible electrode array that is threaded through the cochlea structure to directly interface with the spiral ganglion neurons (SGNs). The platinum electrodes used provide an effective neural stimulation platform. However, increased performance (lower power consumption and/or improved fidelity) could be achieved by using materials or electrode structures with increased electroactive surface area (and hence capacitance). Additionally an electrode structure capable of release of drugs and/or growth factors at appropriate times after implantation could be used to fight infection and/or to promote neurite outgrowth.

The bionic eye is a neurostimulation device that is currently being developed by a number of groups around the world [17, 18] including the University of New South Wales [19] to restore sight to patients blinded by degenerative diseases of the retina. In the latter case an implantable application specific integrated circuit (ASIC) is encapsulated in ceramic and with hermetic feedthroughs connects to an electrode array that is in close apposition to the epiretinal surface. In this manner electrical stimulation can excite retinal ganglion cells directly, thus bypassing the degenerative photoreceptors.

One of the major technological barriers in developing such a device is the construction

10

CHAPTER 2. BACKGROUND INFORMATION

of the high-density epiretinal electrode array. In this case, the planned approach is a construction of platinum foil within medical grade silicone rubber insulation [1]. Materials such as platinum and silicone have a proven track record in commercially-available implantable neurostimulation devices. The platinum foil of 18 μm thickness is micromachined using a Nd:YAG laser allowing feature sizes of down to 25 μm to be patterned as depicted in Figure 2.1. The laser is also used to expose the neural stimulation sites by way of appropriate sized openings (typically 200 μm) in the silicone insulation. This technology will allow a retinal neuroprosthesis with tens to hundreds of electrodes to be manufactured. Future prosthesis designs will no doubt strive for additional electrodes to improve the implantee’s visual perception. In order to construct high-density electrode arrays with thousands of electrodes, new technological approaches will be required. Conductive polymer coatings may offer one alternative to allow small electrodes to be constructed with appropriate characteristics including relatively low impedance and high charge storage densities.

Figure 2.1: Laser fabricated AVPG epiretinal electrode array; i. Prototype 19-electrode array, ii. Stacked electrodes form a potential high resolution 98-electrode array. Images courtesy of G. Suaning, AVPG, Sydney and J. Ordonez, IMTEK, Freiburg.

2.2. CONDUCTING POLYMERS FOR NEURAL INTERFACES

11

Important considerations for stimulating implants are the safe charge injection capacity of the electrode and the stimulation threshold of the surrounding neural tissue. Current designs for both stimulating systems outlined above use platinum electrodes, where typical charge storage density is 6 mC/cm2 [20] and safe charge injection capacity is 100 250 μC/cm2 above which hydrolysis of the platinum can result in the evolution of toxic chemical species [21].

From a neural recording viewpoint, a multichannel neural probe has been developed at the University of Michigan, primarily for investigations in the cortex [11, 22]. This gold coated electrode is designed for application as a diagnostic and surgical mapping tool. It is constructed as a single shank with an array of 8 active recording sites placed at varied distances around the tip to allow areas of the cortex to be analysed and mapped [22]. The challenge in improving recording electrodes focuses primarily on the signal to noise ratio (SNR) and thus the ability to create smaller electrodes that can faithfully record localised biopotentials. Studies have shown that the distance between the recording electrodes and cortical tissue has a significant influence on the SNR [11, 12]. An improved SNR could provide more accurate signals and the possibility of distinguishing signals that were previously unattainable, leading to better diagnostics and efficient surgical mapping [22].

2.2

Conducting Polymers for Neural Interfaces

Conducting polymer coatings aim to improve the electrode-tissue communication through providing a high surface area material more conducive to cell and tissue integration.

12


Charge transfer is improved through reduced impedance and greater selectivity for both recording and stimulating applications. Since polymers are typically softer materials, it is also hypothesised that inflammation is reduced due to reduction in strain mismatch between tissue and electrode surface. The outcome of this reduced inflammatory reaction is a decrease in thickness of the surrounding non-conductive fibrous tissue purported to cause signal degeneration. Conducting polymers polypyrrole (PPy), polythiophene (PTh) and its derivatives such as poly-3,4-ethylene dioxythiophene (PEDOT) are the most commonly investigated polymers, with numerous dopant types ranging from small salt ions to polymers or biospecific dopants such as peptides, proteins and neurotrophins. The chemical structures of reduced PPy and PEDOT are depicted in Figure 2.2. Polyaniline (PAni) is also shown, but is less commonly used in biomedical applications, despite recent evidence that the emeraldine salt modification (EPAni) is conductive and non inhibitive to cell growth and survival [23, 24].

Conductivity in these materials arises from the presence of conjugated double bonds along the backbone of an otherwise insulative structure. In conjugation, the bonds between the carbon atoms are alternatively single and double. Every bond in the backbone contains a localised ’sigma’ (s) bond which forms a strong chemical bond and every double bond also contains a less strongly localised ’pi’ (p) bond [25]. This alternate bond arrangement produces a low energy state and introduces a band width of 1.5eV, classifying these polymers as high energy gap semiconductors [26]. The polymer is transformed into a conductor by doping it with either an electron donator or an electron acceptor.


13

Figure 2.2: Structures of conducting polymers commonly studied for biomedical applications: i.polypyrrole, ii.polythiophene and iii.polyaniline.

Conducting polymers can store charge in two ways. In an oxidation process it could either lose an electron from the conduction or valence band or it could localise the charge over a small section of the polymer chain. Localising the charge causes a local distortion due to a change in geometry, decreasing the ionisation energy of the polymer chain and increasing its electron affinity. This method increases the energy of the polymer less than it would if the charge was delocalised, making it more able to accomodate newly formed charges, and hence takes place in preference of charge delocalisation [26]. A similar scenario exists for reductively doped polymers.

The main criteria for a conducting polymer dopant is that it must oxidise or reduce the polymer without lowering its stability [26]. In oxidative doping an electron is removed from the p-system of the backbone producing a free radical and a spinless positive charge.


14

The radical and cation are coupled to each other via local resonance and the combination is called a polaron. Upon further oxidation the free radical of the polaron is removed, creating a new spinless defect called a bipolaron [27]. Eventually, with continued doping, continuous bipolaron bands are formed along the conducting polymer backbone.

On application of a potential across a conducting polymer film, the main mechanism of conduction is the movement of charge carriers between localised sites or between bipolaron states. Ion movements in or out of the film are dependent on dopant charge and motility [16]. Figure 2.3 shows the ion movement of dopant A− that occurs in a polypyrrole film allowing the conduction of an applied electrical stimulus. Although bipolarons are known to be the main source of charge carriers, the precise mechanism of charge transport through conducting polymers is not fully understood [26]. It is important to consider that most conducting polymers are highly disordered, consisting of both amorphous and crystalline regions. Charge transport is known to occur not only along the polymer chains as shown in Figure 2.3, but also between chains and the complex boundaries established by the multiple phases.

Figure 2.3: Electrochemical behaviour of PPy doped with motile anion A.

Conducting polymers can be manufactured using a variety of methods that are either chemical or electrochemical in nature. Chemical polymerisation proceeds via many dif-


15

ferent routes, such as addition reactions stimulated by radicals, cations or anions; or condensation reactions [27]. These polymers can be tailored to include a great variety of modifications, including covalent attachments to the backbone. However chemical techniques result in films that often require post-fabrication doping to increase conductivity.

In electrodeposition the dopant is added to the electrolyte solution from which the polymer is synthesised. As electropolymerisation can be performed on the electrode intended for coating no further processing is required, making it a superior technique for neuroprosthetic applications. This technique is simple and reproducible, making it the most commonly used fabrication process since 1968 when Dall’Olio [28] and then Diaz [29] electrodeposited the first conducting polymers.

For oxidative polymerisation a dopant molecule is used with an overall anionic charge. Conversely, cationic dopants are used to produce polymers through reduction reactions. Oxidative or anodic electrodeposition is almost exclusively used for conductive polymer fabrication, however it is limited to systems in which the monomer can be oxidised in the presence of a potential to form reactive radical ion intermediates for polymerisation [27].

Much of the initial research into dopant types for conducting polymers applied as neural interfaces quantified the use of aromatic sulfonate variants para-toluenesulfonate (pTS), polystyrene sulfonate (PSS), and sodium benzenesulfonates (BS) to dope polymers [16, 30–34]. Other suitable dopants for oxidation polymerisation include buffer salts, I2 , BF4 , perchlorates and FeCl3 .

16


More recent research has focused on enhancing cell and tissue interactions with polymers through the use of biological signaling factors. These can be included either as an ionically charged dopant or a non-doping inclusion. A non-doping inclusion, unlike a biological dopant, does not have a significant overall charge and hence is not involved in balancing the charge across the polymer backbone. Biological dopants have included laminin peptide sequences, hyaluronic acid (HA), silk-like polymer having fibronectin fragments (SLPF) and polysaccharides [35–37]. Other inclusions such as neurotrophins, non-doping peptides, whole laminin, red blood cells and human serum albumin (HSA) have been incorporated by entrapment during polymerisation, post-fabrication adsorption or pre-fabrication covalent attachment to the monomer [35, 38]. These biological components can potentially enhance integration of excitable tissue with the electrode surface. However, each component of the polymer composite can also have a significant impact on polymer properties.

Studies of conducting polymers, focusing predominantly on PPy and PEDOT have provided insight into their capabilities as biomaterials and the issues associated with tailoring them to cell specific applications. While research shows that these materials are compelling candidates for electrode coatings, there are several challenges that need to be addressed prior to implantation including electrochemical stability, mechanical durability and long-term biological performance.

2.3. CRITERIA FOR ASSESSMENT AND CHALLENGES TO VIABILITY

2.3

17

Criteria for Assessment and Challenges to Long-Term Viability

During development of electrode coatings that are both conductive and bioactive, a number of properties need to be assessed. Typically research focuses on assessment of electrical characterisation, chemical composition and biological response, with few groups focusing on the impact of mechanical properties.

2.3.1

Electrical Characterisation

Four-point probe conductivity provides data on the resistance, (inversely conductance) of a material to the passage of direct current. Most metals provide negligible resistance to the passage of direct current, but conducting polymers have conductivity closer to semi-conductors. Probe conductivity is a good indication of the initial ability of a standalone polymer film to pass current. Table 2.3.1 demonstrates the range of conductivities produced from different dopants and deposition regimes. Highlighted for PPy/pTS is the conductivity data for three different monomer and dopant ratios, suggesting that higher dopant concentrations will yield polymers of higher conductivity. Electrical properties of more complex polymers such as those containing biological components are seldom presented in the literature. Four-point probe conductivity can only be performed accurately on films that can be removed from the electrode substrate. Strongly adherent and mechanically fragile films such as coatings with large peptide dopants or inclusions such as neurotrophins cannot be analysed using this method. Comparative data may be measured


18

on a substrate but the majority of the measured signal is known to travel through the underlying metal which has significantly higher conductivity. As a result other techniques are required to quantitatively assess conducting polymer electrical properties across the spectrum of possible constituents. Table 2.1: Conductivity of common conducting polymers Polymer Dopant Conductivity (S/cm) PPy 0.2M BS 0.3M 100 [39] PPy 0.2M pTS 0.3M 160 [39] PPy 0.1M pTS 0.1M 121 [40] PPy 0.2M pTS 0.05M 100 [41] PPy 0.1M PSS 0.05M 18 [42] PEDOT 0.1M pTS 0.05M 89 [42] a a PEDOT PSS 80 [43] PEDOT 0.01M LiClO4 0.1M 160 210 [44] a

Values not specified

Electrochemical impedance spectroscopy (EIS) provides information on the charge transfer characteristics of the film. Specifically it indicates how stimulus is transferred from the metal substrate through the coating and across the film-electrolyte interface. The impedance behaviour is characterised by fitting an equivalent electrical circuit to the data. The complex impedance is obtained as a function of the applied sinusoidal waveform frequency and the data are represented as a Bode or Nyquist plot. In general it is observed that polymers can reduce the impedance magnitude of an electrode by up to two orders of magnitude and Bode plots indicate that the cutoff frequency is at a significantly lower frequency for polymer coatings than for bare metal electrodes [20, 45]. Most commonly the data for 1 kHz is examined in detail as most neural cell communication occurs between 300 Hz - 1 kHz [30]. The electrolyte solutions used need to be physiologically


19

relevant and typically 0.9% saline and phosphate buffered saline (PBS) are used for neural electrode characterisation.

At 1 kHz conducting polymers have been reported to decrease electrode interface impedance from as high as 4 MΩ for bare gold microelectrodes to 120 kΩ for electrodeposited PPy/PSS coated gold. The use of neurotrophins incorporated during deposition creates a rougher film that decreases impedance further to around 15 kΩ [12]. Recent investigations into the use of hydrogel scaffolds through which conducting polymers are deposited have produced significant increases in surface area resulting in reported reduction in impedance from ≈1 MΩ on bare electrode to 7 kΩ for PPy/PSS coated electrodes [20]. Table 2.3.1 shows these results for a range of coatings and biomolecule incorporations, validating the use of polymer coatings for reducing impedance a range of polymers and and biomolecules. Table 2.2: Impedance magnitude at 1 kHz on 1250 μm2 electrodes Substrate Polymer coating Impedance magnitude @ 1kHz Gold 4 - 0.8 MΩ Gold PPy/PSS 120 kΩ Gold hydrogel + PPy/PSS 7 kΩ Gold PPy/PSS/NGF 15 kΩ Gold PPy/SLPF 360 kΩ Iridium 980 kΩ Iridium PEDOT/DCDPGYIGSR 130 kΩ

[12, 30] [46] [46] [12] [12] [47] [47]

Cyclic voltammetry (CV) is an electroanalytical technique used to assess the reduction and oxidation (redox) peaks of the polymer. In CV a potential is linearly scanned up to a switching potential and then reversed to its initial value. Many electrochemical studies favour this method during electrodeposition as it clearly shows the formation of conducting polymers and also indicates the potential range of charging and discharging.

20


A severe problem in the interpretation of conducting polymer redox states is the variety of possible shapes and forms of CV curves, even when the materials are prepared under similar conditions [48]. The setup used for CV should take into account material used for the counter electrode as well as the size (should have a surface area greater than 100 times the working electrode), the electrolyte solution constituents and the type of reference electrode used. For example, PPy/PSS alone has been reported as oxidising at -150, -250 and -340 mV [20, 31, 33] depending on these parameters.

Cyclic voltammetry is more gainfully employed in the assessment of electrochemical stability. Conducting polymers can be subjected to a continuous cycling of redox to determine the shifts in current carrying capacity as time progresses. Since the CV is performed at a specific rate, the area contained within the curve can be calculated by integrating the current at the film surface over time to give the charge carrying capacity of the film. Polymer coated electrodes have been shown to have charge carrying densities of up to 560 mC/cm2 compared to values of around 6 mC/cm2 for platinum [20]. The continued cycling of the potential difference across the film usually leads to a decrease in this calculated area, indicative of loss of charge carrying capacity over long-term stimulation.

The greatest challenge to the long-term viability of a conducting polymer electrode coating is the preservation of the conductivity or electrochemical stability. Losses of electroactivity are thought to occur predominantly through the loss of dopant molecules. PPy films doped with PSS have been shown to lose up to 95% of their conductivity when subjected to 16 hr of polarisation [33]. PEDOT films have been presented as being superior in main-


21

taining electrochemical activity with reports of only 30% loss of charge carrying capacity following 400 redox cycles [49] and an 11% reduction in conductivity following 16 hr of in vivo polarisation [33]. The superior electrochemical stability of PEDOT is attributed to the dioxyethylene bridging group, which blocks the possibility of coupling along the backbone, the defect that causes loss of stability in PPy conductors. Studies by Green et al. have shown that by creating a composite polymer layered with multi wall nanotubes (MWNTs), electrochemical stability can be improved by up to 50% in poorly performing film such as PPy/PSS when subjected to oxidation and reduction voltammetry for 400 successive cycles [49]. Alternately, layering of polymers with complementary properties could be a solution to providing electrochemical stability. Studies where PEDOT/pTS was layered over PPy/PSS showed that the composite film preserved 80% of its original electroactivity compared to 38% for PPy/PSS alone. This was only marginally lower than the PEDOT/pTS monolayer in which 89% electroactivity was preserved [42].

2.3.2

Mechanical Behaviour

Physico-mechanical properties need to be assessed to determine the surface roughness, polymer hardness, the potential for delamination and brittle fracture of coatings and the fatigue behaviour of the coatings. Scanning electron microscopy (SEM) and atomic force microscopy (AFM) are commonly performed to provide information on film roughness and topology. However, more quantitative studies performed on film structural mechanics are rare. Microtensile tests have yielded sparse data on the ultimate tensile stress of both PPy films and PEDOT films. Murray et al. [50] reported PPy/pTS as having an ultimate

22


tensile strength of 3.6 GPa and Lang and Dual [51] showed chemically produced, commercially available PEDOT/PSS (also known as Baytron) as having an ultimate tensile strength of 1 - 2.7 GPa and yield strength of 25 - 55 MPa. Several papers have reported electrochemically prepared conducting polymers as brittle and with limited ability to be handled [52, 53], however this is critically dependent on the dopant used [54, 55].

Nanoindentation was used in a study by Yang and Martin which showed a strong correlation between film softness and low impedance [45]. Nanoindentation shows promise as a technique that would provide preliminary information on in vivo responses to polymer implants as it is hypothesised that a softer interface would reduce inflammation caused by strain mismatch between the tissue and electrode. Consequently a reduction of inflammation might adequately reduce foreign body reactions and scar tissue formation responsible for stimulation failure in long term implants.

The durability of a coated electrode for the duration of an implant lifetime has not yet been studied. Mechanical studies are important for providing insight into possible delamination or mechanical erosion. A study by Tallman et al. into conducting polymers for corrosion control showed that pull-off tests conducted according to ASTM standard procedure D-4541-95 could provide valuable information on the adherence of PPy films to metal substrates [56]. Additional tests that have not been used in assessing conducting polymers but could provide insight into coating integrity are the ASTM adhesion and hardness assays. These assays are specifically designed to assess film adherence to a metallic substrate and film softness on a calibrated universal scale [57, 58].


23

Mechanical properties are integral to the performance of an implantable conducting polymer, but are rarely explored for biocompatible formulations. It is reported that conducting polymers are often brittle and the use of larger dopants can exacerbate this effect [37, 52, 53]. The molecular size of the dopant is an important aspect, because the polymer undergoes substantial volume changes during the oxidation and reduction if the dopant is motile. These dimensional changes can be extremely large and influence the mechanical strength of the polymer [59,60]. Data on the conductivity of conducting polymers with large biomolecule incorporations are primarily limited to information obtained through EIS. It is theorised that conductivity probes are not applied to these films as they are too brittle to be removed from the substrate to produce stand-alone films. Collier et al. reported that PPy films doped with biomolecule HA were more brittle than those doped with PSS and were difficult to handle [37]. Layering of polymers with complementary properties is also likely to provide a solution to improving mechanical durability. PPy/PSS and PPy/HA bilayer films were produced by Collier et al. to mitigate the poor mechanical and electrical performance of monolayer PPy/HA films, whilst preserving the superior biological interactions provided by inclusion of HA [37]. The limited information available on the mechanical performance of conducting polymers demonstrates that further research is required to produce a safe and durable long-term implant.

2.3.3

Assessment and Impact of Chemical Composition

Chemical composition is usually determined through x-ray photoelectron microscopy (XPS) or Fourier transform infra red spectroscopy (FTIR). Commonly these characteri-

24


sation techniques are used to confirm the presence of dopants or biological inclusions as they provide detailed data on the elemental composition of a material. Additionally, XPS and FTIR can be used to determine the doping ratio which is used to calculate both the theoretical thickness of the film as specified by Equation 2.1 adapted from Tallman et al. [56] and then the efficiency of polymerisation. The film thickness, d, is determined from the current density, j (A/cm2 ), deposition time t (s), the equivalent molecular weight MWeq (g/mol), Faradays constant F (C/mol) and the film density, ρ (g/cm3 ) as shown in Equation 2.1.

d=

j · t × M Weq F ·ρ

(2.1)

The equivalent molecular weight, MWeq is somewhat empirical, being determined from the dopant ratio by XPS or FTIR analysis. The dopant ratio is the proportion of dopant units to monomer units incorporated in the film. For electrodeposited conducting polymers this is usually between 0.1 and 0.5. Thus for each monomer unit the dopant only contributes 0.1 to 0.5 of its molecular weight. Knowing this the equivalent weight can be determined by the mass of combined polymer/dopant deposited per mole of electrons transferred. The theoretical thickness can be compared to the thickness measured by SEM or profilometer to determine the efficiency of an electrodeposition process. It has been suggested that conducting polymer films can be produced at almost 100% efficiency [56].

Chemical composition is important in determining the presence and amount of dopant


25

ions in the polymer. Dopant ions that are released during electrical cycling, mechanical erosion or through biological degradation can have a significant impact on the surrounding tissue. Chemical analysis techniques can not only determine the proportion of dopant and monomer in the film but also the presence of contaminants imparted to the films during processing and handling.

2.3.4

Biological Response

Biological performance testing of conducting polymers tends to focus on in vitro functional efficacy tests using neural-like cell lines or primary isolates from neural tissue. In these tests, model neural cells such as the rat pheochromocytoma (PC12) cloned cell line are often used. This cell line is robust and retains a number of neuronal characteristics, including the ability to sprout neurites in the presence of NGF. Length of neurite outgrowth and percentage of cells bearing neurites is used to indicate the relative functional efficacy of a given film. PC12 outgrowth assays have indicated that conventionally doped polymer films support cell attachment and growth. PC12s grown on PPy/PSS resulted in a median neurite length of 9.5 μm compared to tissue culture plastic controls where the median neurite length was 8.2 μm [32].

Toxicity testing tends to report favourable results suggesting that materials such as PPy and PEDOT have relatively good overall biological performance. Lack of toxicity is often inferred by successful growth of cells on the materials during efficacy tests. Fundamentally, the biological performance of conducting polymers is dependent on the chemical

26


characteristics of the material, its physical attributes and the presence of bioactive factors. Critical chemical characteristics include presence of unreacted monomers or small polymer chains, the motility and toxicity of dopant ions and the presence of process contaminants such as solvents. Although chemical analysis can determine the proportion of these components prior to implantation, their release from the polymer and subsequent impact on cells/tissue needs to be assessed using in vitro cell studies or implantation studies.

Toxicity data for representative monomers, dopants and typical process contaminants are shown in Table 2.3.4. The lethal dose (LD50) value is the amount of product that results in 50% of the animals exposed dying within a defined time period. While the routes of exposure are not the same as those of a neuroprosthetic, the LD50 values give an indication of the relative primary toxicity of a chemical. These data indicate that monomers tend to have higher toxicity than dopants however both components are slightly to moderately toxic. This clearly emphasises the need for removal of leachables and use of clean processing techniques, processes that apply to all biomaterials. While conductive polymer matrices may initially constrain dopants, if the dopant is motile it will be released from the film during oxidation. Often these ions will migrate away from the film and upon reduction an alternate ion from the surrounding environment such as Cl- will be taken back into the polymer. If in a constrained local environment such as the eye or cochlea, the risk to surrounding tissue from accumulation of potentially toxic molecules needs to be assessed.


27

Table 2.3: Toxicity limits of common polymer constituents [61] Toxicity Data: LD50 (mg/kg) Constituent Oral Subcutaneous Intraperitoneal Monomers Pyrrole 137 98 61 EDOT 615 894 Dopants BS 9378 pTS 1000 PSS 8000 15000 6000 LiClO4 1160 Solvent Acetonitrile 50 4480 1680

Other types of toxicity associated with conducting polymers such as mutagenicity and hypersensitivity have not been widely reported. A scan of the Toxnet Database (National Library of Medicine, National Institutes of Health, USA) had several records for mutagenicity tests of the monomers pyrrole and thiophene, all finding that the chemical is not mutagenic in bacterial point mutation assays [62]. Hypersensitivity of the monomers and dopants has not been widely reported.

Neuroprosthetics are intended for use in an active stimulating environment. Thus an important consideration in assessing cell and tissue responses is the effect of electrical activity. In vitro assays conducted by Schmidt et al. showed that PC-12 cells cultured on PPy films and subjected to an electrical stimulus through the film showed a significant increase in neurite lengths compared to the passive control [32]. Subsequent studies indicated that the longer neurites grown on PPy was due to increased protein adsorption from serum containing media mediated by the electrical stimulation [63]. However extended stimulation regimes indicated that this effect only extended over the first 24 hr of growth

28


and stimulations performed beyond this time frame showed no significant difference to passive controls.

As already discussed, the addition of bioactive factors to conducting polymers is a strategy for improving cell and tissue interaction that has had much attention. Although PC12 cells are the typical cell used in initial screening studies, polymer bioactivity evaluation has also been performed on SGN explants, glial cells and human neuroblasts [16, 64]. Richardson et al. demonstrated that PPy with NT-3 integrated during electrodeposition can promote neurite outgrowth from SGN explants [16]. Synthetic anionic laminin peptide DCPGYIGSR doped PEDOT was shown by Cui et al. to produce an interface with superior biological performance in comparison with polymeric anionic dopant PSS [30]. Other in vitro assays using rat glial cells and human neuroblastomas indicated that these cell types had greater affinity to coatings with biomolecule modifications [35].

Most in vitro studies published have been overwhelmingly positive in terms of the biological performance. However, in many studies polymer films tested are significantly different to films intended for implant use and there is little consideration given to the environment in which the electrode will be placed. The next stages in biological assessment are the use of application specific devices tested in their intended anatomical sites. Recording electrodes with polymer coatings have been assessed acutely in the guinea pig cerebellum [20] and rat motor cortex [47]. Initially SNR of neural recordings were improved by addition of a PEDOT coating, however over a 6 week period both the PEDOT coated and the bare electrodes registered large increases in impedance that mitigated the effect


29

of these improvements [47].

The addition of neurotrophic agents to polymers aims to produce an interface where the electrodes and cells are intimately contacting, providing a permanent attachment with low impedance for the implant lifetime. It is important to note that no studies to date have determined the ongoing performance of such an interface, with hypotheses suggesting that the contact may be lost as the neurotrophin is consumed. Studies conducted at the Intelligent Polymer Research Institute have used isotope tagged NT-3 to analyse release profiles of the neurotrophin NT-3 from PPy films. Both passive diffusion and the application of stimulation regimes were used to classify NT-3 release, and in all scenarios it was observed that an initial burst occurred within 24 hours followed by significantly lower levels of release for the remaining six days of the study [16, 64].

It is possible that the ongoing electrical stimulation of cells in the vicinity of the electrode may provide an environment conducive to neural cell maintenance but no long-term studies have been reported. Interfaces that contain cell adhesion peptides may provide a more permanent cell attachment but these may be subjected to denaturing by the body, rendering them inactive. Supporting this theory, in vivo studies on HA doped PPy revealed a higher level of vascularisation in rat subcutaneous implants than films doped with polymeric anion PSS. However, after 6 weeks in vivo there was no significant difference between the two film types. It was hypothesised that the initial vascularisation was associated with HA degradation products, that by 6 weeks were degraded or consumed [37]. It is clear that longer term in vivo studies with active electrical stimulation are required


30 in this area of research.

2.4

Summary and Conclusion

Conducting polymers have promise as platforms to enhance neural tissue integration at electrode interfaces. The major areas of development required for long-term performance include electrochemical stability, mechanical integrity and the maintenance of an intimate contact between the electrode surface and surrounding neural tissue. Incorporation of biomolecules appears to be highly beneficial in increasing cell interactions with electrodes. However, designing an optimised interface will require a trade-off between desired electrical, mechanical, chemical and biological properties. A better understanding of the impact of adding biological components on polymer properties is needed and understanding the changes polymers undergo in vivo will be integral in determining their role in improving performance of neural prosthetics.

Part I

Design of Conducting Polymer Electrode Coatings

31

Chapter 3

Synthesis and Characterisation of Conducting Polymers 3.1

Introduction

Conducting polymers have been studied extensively since the early 1980s [25]. There are currently over twenty five reported conducting polymer systems, all based on a conjugated carbon backbone [65, 66]. Most research on conducting polymers including polypyrrole (PPy), polythiophene derivative poly(3,4-ethylene dioxythiophene) (PEDOT) and polyaniline (PAni) has focused on the modification of biosensors and bioelectronics [6, 23, 24, 32, 67, 68]. The objectives of conducting polymer electrode coatings are to improve sensitivity, impart selectivity, to suppress interfering reactions and to hold and release biological factors [31]. While conducting polymers are unlikely to have equivalent conductivities to metal electrodes, they are hypothesised to provide a significantly increased interfacial area, resulting in a lower impedance than metal electrodes of the same size. Studies of these conducting polymers have provided insight into their characteristics 32

3.1. INTRODUCTION

33

and their potential capabilities as biomaterials.

An ideal conducting polymer for neuroprosthetic application will retain an acceptable level of charge transfer by having a low impedance and high surface area, be easily manufactured using accessible technologies and have evidence of biocompatibility in the processed form. The polymer coating must adhere strongly to the electrode substrate (ie. platinum), but should also be softer and more compliant than commonly utilised metals to reduce inflammation caused by strain mismatch at the tissue-electrode interface. Electrically and mechanically, the conducting polymer requires long-term stability.

PPy has been investigated for use in peripheral nerve regeneration scaffolds and bioactive electrode interfaces [11, 12, 16, 32]. It is considered a good candidate due to its relatively high conductivities, ease of processing and handling, and promising biological compatibility when assessed as a thin film in vitro [32]. PPy has been doped with a variety of reagents to produce a range of conductivities from 1 - 400 S/cm. PPy can be loaded with peptides for cell attachment [11, 34] and neurotrophins for growth promotion of neurites in targeted cell types [12, 16, 69].

While PPy has shown promise as a conducting polymer coating for neuroprosthetic applications, it has not been assessed across some significant criteria for biomaterials. Longterm electroactivity, mechanical properties and biological response to thick films have not been thoroughly investigated.

Polythiophenes can be fabricated using similar processes to polypyrrole, where the elec-

34

CHAPTER 3. SYNTHESIS AND CHARACTERISATION OF CPS

trode is immersed in a monomer solution. Polythiophene has the advantage of being stable in air and is not prone to the overoxidation reported in PPy [25, 70, 71]. PEDOT is the most commonly investigated polythiophene derivative. Studies have shown that PEDOT demonstrates much better electrochemical stability than PPy, due to the dioxyethylene bridging group, which blocks the possibility of α-β’ coupling [33], the cause of loss of stability in PPy. Cui et al. investigated PEDOT as an alternative to PPy on intracortical electrodes for neuron ingrowth and discovered that PEDOT retained greater electrochemical stability and decreased the impedance modulus by almost two orders of magnitude [30]. However, fabrication of the PEDOT coatings is relatively difficult due to the poor solubility of the EDOT monomer [30]. The use of water as a solvent can cause mislinkages and subsequent deterioration in polymer properties [72]. Consequently, the use of harsher solvents such as acetonitrile are required [6].

Several research groups are currently exploring PEDOT for use in biosensor and neuroprosthetic applications [6, 30, 33]. While research shows that PEDOT films have superior electrical properties to PPy films, very little literature is available on the long-term electrical properties. Scarce data is available on the mechanical performance of PEDOT regardless of its fabrication method or dopant type.

Conduction of charge through polymers is made possible by the introduction of doping ions to the structure during manufacture. The doping ions take the form of either an electron donator, if oxidative electrodeposition is used, or an electron acceptor for reductive electrodeposition. The dopant balances the charge across the unstable backbone while

3.1. INTRODUCTION

35

the polymer is in the passive form. On application of a potential across the film, biporonic conduction and the resulting ion partitioning causes the passage of charge through the film. This research will focus on the two dopants which have dominated research into conducting polymers for neural interfaces: PSS and pTS [16, 30–34].

The primary objective of this study was to gain a thorough understanding of conducting polymers previously described in literature across a variety of physico-chemical and biological properties. The profound effect of surface topography and mechanics on cell behaviour has been explored by tissue engineering researchers [73,74] and yet these properties in conducting polymer electrode research are not rigorously examined. Specifically, this study provided baseline data for controls and important new data on electrical and mechanical tests not commonly published, but integral to the long-term performance of an implant material. This chapter details the methodology used for synthesis of basic PPy and PEDOT electrode coatings doped with sulfonate ions PSS and pTS. The analytical techniques used to assess these polymers are also detailed and results collated to provide insight into the overall performance of these polymers.

The specific aims were to:

1. Prepare conductive polymer films described in the literature using common electropolymerisation parameters,

2. Assess the effect of polymer backbone chemistry and dopant type on the physical and chemical properties of the films. Specifically, surface morphology, chemical


36

composition, electrochemical stability, conductivity, impedance response, mechanical hardness and film adherence to the electrode will be evaluated, and

3. Examine the effect of polymer backbone chemistry and dopant type on mammalian cell interactions using a neural-like cell line.

3.2 3.2.1

Materials and Methods Modelling of the Electrode Interface

This research was conducted primarily in the context of neural interfaces for vision prosthesis applications. The electrodes intended for implantation are not suitable for bulk processing and high throughput assays. As such a model electrode system was designed for assessing films. Additionally a neural model cell was used for biological assays, allowing efficient assessment of a wide variety of film components in a consistent biological environment. Using this model it was possible to vary single components of the coating and assess the impact across electrical, mechanical and biological properties in a uniform, repeatable fashion.

3.2.1.1

The Model Electrode

Film characterisation was performed using a model electrode system that allowed assessment of a range of bulk properties using the same format for each assay. Properties were determined using a scaled up model such that results could be obtained easily without

3.2. MATERIALS AND METHODS

37

requiring micro-controlled probes or specialised equipment. Most neural implants use microelectrodes with a nominal diameter of 4 - 500μm. Assessment of properties at this scale is both difficult and expensive. For these purposes a model electrode system was developed to allow the characterisation of films using macro-manipulation and commonly used laboratory techniques.

The system developed consisted of three components depicted in Figure 3.1: a silicone well gasket, a substrate insert and sandwich assembly with lid. The silicone well gaskets were purchased from Greiner BioOne and consisted of an array of 6 x 2 wells of 5 mm diameter (Flexiperm Micro 12; Cat #. 90011436). The substrate used was either a Pt foil mounted on a glass microscope slide or a conductive indium tin oxide (ITO) coated microscope slide. A custom designed and manufactured polycarbonate sandwich assembly was then used to clamp the substrate and silicone gasket together, creating 12 discreet wells in which polymer films could be both fabricated and characterised.

3.2.1.2

The Model Cell

For biological characterisation a cloned neural cell line was used to determine the biological properties of the conductive polymer coating. The cell chosen was the PC12, transfected from transplantable rat adrenal pheochromocytoma. This line was developed at Harvard Medical School in 1976 as a useful model system for neurobiological and neurochemical studies [75]. The specific line used for this research was further transfected with green fluorescent proteins by Marinpharm GmbH (Cat #. PC-TG-PC-12). This

38


Figure 3.1: Custom manufactured well sandwich assembly


39

cell line is beneficial as cells can be viewed by fluorescent microscopy on opaque thick polymer films and platinum substrates, providing results that are a closer approximation to in vivo behaviour than literature results produced on microscope slides with thin transparent films. A neural clone was used in preference to a primary cell to allow high throughput assays with common literature control references. PC12s undergo differentiation in response to NGF supplemented low serum medium, which is quantifiable through the resultant neurite outgrowth.

3.2.2

Conductive Polymer Constituents

Conductive polymers manufactured through the use of electrodeposition require three components; the monomer, the dopant and a solvent. Initial studies were performed to characterise some commonly explored conductive polymers against reported literature values. These studies were determined as a baseline from which subsequent modifications could be compared. Pyrrole and 3,4-ethylene dioxythiophene (EDOT) monomers were doped with sulfonate variants PSS and pTS. Both monomers have been explored in literature and show promise in neural prosthetic applications. The dopants, shown in Figure 3.2 have also been characterised for use in conducting polymer coatings and are significantly different in size (PSS: MW=70,000-100,000 g/mol; pTS: MW=196 g/mol), enabling some comparison of the impact of dopant size on the film. Solvents were prepared to include the lowest proportion of non-aqueous components in an attempt to reduce possible contamination from harsh chemicals.

40


Figure 3.2: Sulfonate dopants: i. PSS and ii. pTS

3.2.3

Electropolymerisation

Electropolymerisation is commonly performed as a potentiostatic reaction using a threeelectrode system shown in Figure 3.3i. In this method the film is deposited on the working electrode with a reference electrode placed close to the working electrode, maintaining a constant potential across the film surface. The counter electrode supplies the electrons that are used to initiate the polymerisation reaction. However, when galvanostatic deposition is used the system can be reduced to having only two electrodes present in the electrolyte solution, as the reference is not required at the film surface to maintain a constant current. This method was preferable as the wells in which deposition occurred were relatively small, with a maximum capacity of approximately 200 μL. Additionally, this allowed for concurrent depositions of coatings across all 12 wells using a counter electrode array as depicted in Figure 3.3ii.


41

Figure 3.3: Electropolymerisation set up for: i. 3-electrode cell potentiostatic deposition; ii. 12-well array galvanostatic deposition. The silicone well gaskets described in Section 3.2.1.1 were used to produce coatings in a masked pattern on either a Pt foil or an ITO coated microscope slide. The Pt substrate was used for all assays that required the films in a composite format similar to their implant application, i.e. electroactive stability, impedance spectroscopy, film adherence and hardness and in vitro cell assays. Films were electrodeposited on ITO substrates for four-point probe conductivity measurements where they were required to be removed from the substrate to be assessed as stand-alone films. Scanning electron microscopy (SEM) and X-ray photoelectron spectroscopy (XPS) samples were also deposited on ITO as a disposable substrate was preferable to the Pt foil. Pt foil was obtained from Alloy and Gold, Sydney at 99.95% purity. All slides were 50 mm x 25 mm and 100 μm thick. Prior to each use the Pt was polished with 1200 Grit automotive grade abrasive paper to remove previous depositions and any surface scratches. The Pt was soaked in 1% Decon90 overnight, rinsed in deionised (DI) water and then autoclaved at 121 ◦ C for 20 mins in a waterbath. Foils were dried in a clean hood to minimise airborne contamination.

42


Wells were cleaned prior to each electrodeposition using a solution of 1% Decon90 in which they were soaked overnight, abraded of particulate contaminants with a testtube brush and then thoroughly rinsed and dried in a clean flow hood. Each silicone gasket was attached to either a Pt or ITO slide according to the application. The well system was clamped in the custom-built sandwich assembly described in section 3.2.1.1 to prevent leaking from adjacent well chambers.

Basic monomer solutions were made up in a glass vial as indicated in Table 3.1. Various other monomer-dopant ratios were explored, however initial testing to optimise the electrochemical stability revealed that 0.1M monomer (for pyrrole and EDOT) with 0.05M dopant (for pTS and PSS) was sufficient to produce a film with stable properties. It should be noted that molarity of the PSS dopant was calculated with respect to the styrene sulfonate monomer unit, or alternately, both dopant molarities were calculated with respect to their intrinsic net charge. Film stability studies are presented in Appendix A. Additionally, literature confirms that polymers produced from this monomer:dopant ratio have high efficiency of polymerisation and are not inhibitive to cell adherence and survival [31, 56].

Component Pyrrole EDOT PSS pTS Deionised (DI) Water Acetonitrile

Table 3.1: Basic polymer solutions Cat # PPy/PSS PPy/pTS PEDOT/PSS 131709 0.1M 0.1M 483028 0.1M 243051 0.05M 0.05M 402885 0.05M N/A 100% 100% 50% 271004 50%

PEDOT/pTS 0.1M 0.05M 50% 50%


43

All products were obtained from SigmaAldrich and pyrrole was vacuum distilled prior to each use to remove polymerised contaminants that reduce efficiency of electrodeposition. EDOT was found to have a longer shelf-life and was vacuum distilled on a monthly basis to remove polymerised components. Aliquots of 150μL of polymer solution were placed in each well with a Pt counter electrode. The working electrode was formed by the Pt or ITO substrate and electrically connected by an alligator clip. Polymers were electrodeposited using an in-house manufactured galvanostat (detailed in Appendix B) at 1mA/cm2 or 300μA per well for 10 minutes. Following deposition the deposition electrolyte was removed. Each film was washed with DI water four times. The final aliquot was left on the film for 24 hours in a 37 ◦ C incubator to leach any excess monomer, dopant or process contaminants.

The resultant film morphology and constitution was analysed through the use of SEM and XPS respectively.

3.2.3.1

Scanning Electron Microscopy

SEM was performed to obtain qualitative data on the film surface morphology. Following deposition on ITO microscope slides, films were air dried overnight and coated with an 8 - 12 nm layer of chromium using an EmiTech K575x high resolution coater. The samples were placed in a Hitachi S3400 SEM under vacuum with an accelerating potential of 20 kV. Images were captured at 1000 and 15000 times magnification.


44 3.2.3.2

X-ray Photoelectron Spectroscopy

XPS samples were produced on ITO slides, washed and then air dried overnight. The ITO slides were scored with a diamond cutter and then snapped to produce a single sample on a square of approximately 1cm2 . XPS analysis was carried out on a Kratos XSAM800 XPS and relevant peaks were analysed to determine presence and relative amount of monomer units and dopant 1 . From these values the doping ratio was calculated and used to determine the effectiveness and repeatability of the deposition. Typically doping ratios for conducting polymers are between 0.2 - 0.4. Values outside of this range indicate a poor deposition.

3.2.4


Electrical properties are integral to the performance of a neural interface. The addition of a polymer layer should not have a negative impact on the interface and in fact aims to improve the passage of charge between the Pt electrode and the neural tissue. Electrical properties are not determined simply by conductivity but rather a combination of the electrochemical stability, the conductivity and the electrochemical impedance. These three mechanisms were used to assess PPy and PEDOT films doped with PSS and pTS.

1

XPS was performed by Bill Bing Gong at the UNSW Surface Analytics Facility

3.2. MATERIALS AND METHODS 3.2.4.1

45

Electrochemical Stability

Electrochemical stability was determined through cyclic voltammetry (CV). CV is used in literature to determine the potential at which the conductive polymer will undergo oxidation and reduction. It uses a three-electrode cell with measurements performed in an electrolyte solution. For implant applications it is important that the electrolyte contains ions that are similar to those available in the cellular fluid such as saline solutions. The peaks seen during oxidation and reduction can vary significantly depending on the set up used in each laboratory. Additionally, CV can provide information about the charge carrying capacity of the film. Since the CV is performed at a specific rate, the area contained within the curve can be calculated by integrating the current across the film surface over time to give the charge carrying capacity of the film over the ramped potential. The continued cycling of the potential difference across the film usually leads to a decrease in this calculated area, indicative of loss of charge carrying capacity over longterm stimulation. The primary goal of CV in this research is to determine the relative electrochemical stability of the assessed polymers over a prolonged number of CV cycles.

PPy and PEDOT films doped with pTS and PSS as described in section 3.2.3 were deposited on Pt electrodes and analysed by CV to determine electrochemical stability. An eDAQ potentiostat and eCorder unit coupled with the supplied EChem software package (eDAQ Pty Ltd., Australia) was used to apply a cycling voltage from -800mV to 600mV. This range was chosen to allow the oxidation and reduction of the polymer within the limits of the water window, to prevent hydrolysis occurring at the electrodes. The scan

46


rate was set at 120mV/s for 400 continuous cycles and measurements were performed in 0.9% saline. The recordings were made with an isolated Ag/AgCl reference electrode and Pt counter electrode. The first stable curve was considered Cycle 1 and used to determine the original electroactivity of the film. The area contained within the oxidation-reduction curve was calculated to establish the current carrying capacity of the film. The area contained within the curve of successive cycles was calculated as a percentage of Cycle 1 to show the comparative loss of electroactivity over time. This process was repeated for three individually prepared samples and the mean curve for electroactivity loss was determined with the standard error of the mean (SE).

3.2.4.2

Conductivity

Polymer films were removed from the substrate and analysed using a Jandel four-point probe (Jandel Engineering Ltd., UK) with tungsten carbide tips having radii of 40 μm, tip spacing of 0.635 mm and set to a load of 60 g. A Jandel RM3 test unit was used as the current source and voltmeter. Basic PEDOT and PPy films were manually removed from the ITO glass substrate and air dried overnight. The stand-alone films were then subjected to currents ranging from 20 μA through to 2 mA with the linear four-point probe head. Current was applied in both directions and the average voltage output was taken. Values were adjusted to S/cm, using Equation 3.1.

1 Vav = × t × CF Conductivity I

(3.1)


47

Vav is the average voltage recorded using the four-point probe. I, is the applied current and t, the film thickness. CF is a correction factor based on the geometry of the sample and the probe, and their placement relative to each other. For all samples tested in this thesis, the linear probe was placed in the centre of the sample using micromanipulators supplied with the four-point probe platform. All samples were circular and on average 5 mm in diameter. The CF for samples in this set up was determined to be 3.973 as published in the United States National Bureau of Standards, 1964 [76]. For each material three samples were analysed, the mean and SE were calculated.

Polymer film thickness was measured using a variety of methods. Initially, the thickness of basic films was measured with a Dektak stylus profilometer. This was less than ideal as polymer debris was observed on the stylus following each pass, indicating that the stylus was scraping the surface and possibly removing thickness during measurements. Snapping an ITO slide to create a cross-section through the polymer and employing stage tilting during SEM was then used to gauge deposition thickness. Some measurements were obtained using this method and correlated well with profilometer results, however, snapping of the ITO through the film was unreliable and often caused film delamination from the slide. Finally, some measurements were obtained by optical profilometer. Due to the high cost of optical profilometers, samples were sent away for once-off trials with values obtained for a single sample of each film type from two companies. An average of these values was used as the thickness for conductivity calculations.

48 3.2.4.3

CHAPTER 3. SYNTHESIS AND CHARACTERISATION OF CPS Impedance

Electrochemical Impedance Spectroscopy (EIS) was used to determine the impedance characteristics of the system. This is different from conductivity in that AC signals can be used and film characteristics determined across a range of frequencies. This method relies on the use of an electrical model fit to the data to determine the resistive and capacitive elements of the system.

Figure 3.4: Setup used for EIS analysis. An AC function with no DC offset was created using the function generator. This signal was applied using the potentiostat and displayed on the oscilloscope. The generated waveform was passed through each sample using the potentiostat with a three-electrode setup. The working electrode and reference electrode maintained the required potential across the polymer surface and the counter electrode measured the resultant waveform. The signal recorded from each sample through the counter electrode was displayed on the oscilloscope. The amplitudes and difference in phase between the two waveform peaks were recorded manually.

The EIS system used is represented in Figure 3.4, consisting of a National Instruments


49

(NI) function generator controlling a house-made potentiostat with the input and output signals recorded on a NI oscilloscope. The signal was applied to the working electrode, comprised of the Pt substrate with polymer coating, in 0.9% NaCl electrolyte controlled by a Ag/AgCl reference and returned via a Pt counter electrode. The input signal amplitude (Vin ) was compared to the output signal amplitude (Vout ) and the time between each signal’s peak (Δt) was measured. The impedance magnitude and phase angle was calculated as per Equations 3.2 & 3.3 respectively. Where Rpotentiostat is the sense resistance of the potentiostat and f was the signal frequency. For all samples an AC sine wave with a 40mV amplitude was applied with 0V DC offset and frequencies from 10 Hz through to 100 kHz.

ImpedanceM agnitude =

P haseAngle =

Vin Vout /Rpotentiostat

Δt · 360◦ 1/f

(3.2)

(3.3)

Two data points were recorded at 10, 100, 500, 1000, 5000, 10000, 50000 and 100000 Hz, proceeding from low frequency to high frequency and then returning to low frequency. The values taken in each direction were averaged to give a single data point for each frequency. Values were presented on a Bode plot and compared to bare Pt foil. Three individually prepared films were sampled and the average curve plotted for both magnitude and phase with error presented as one standard deviation.

50

3.2.5



The use of softer more pliant materials in implant interfacing is aimed at reducing the strain mismatch between the electrode and tissue that can lead to the formation of scar tissue through inflammation and foreign body reactions. Few research groups assess the mechanical properties of conductive polymers, other than to indicate that they are softer than metals and hence have a greater compliance with neural tissue. During this research the film hardness was determined through the use of ASTM Hardness assays. The film was also assessed for adherence to the substrate through an ASTM x-cut adhesion test.

3.2.5.1

Film Hardness

A modified version of the ASTM standard test method for film hardness was performed on conductive polymer coated Pt substrates. The standard was developed as a procedure for rapid, inexpensive determination of the film hardness of an organic coating on a substrate in terms of pencil leads of known hardness [57]. The test was modified to account for small areas of films deposited on Pt foil. Calibrated Staedtler pencils meeting the scale of hardness illustrated in Figure 3.5 were prepared according to the standard.

Figure 3.5: ASTM calibrated pencil hardness scale

Approximately 5 to 6 mm of wood was removed from the point of each pencil to produce an undisturbed, unmarked, smooth cylinder of lead. The pencil was then rubbed at 90◦


51

to a piece of 400 grit abrasive paper until a flat, smooth and circular cross section free of nicks was obtained. Starting with the hardest lead, the pencil was held at a 45◦ angle and pushed away from the operator across the 5mm diameter film of polymer. Sufficient pressure was exerted to either cut or scratch the film or cause the pencil tip to crumble. The process was repeated for each pencil hardness and results were recorded on a LYNX magnifier with camera attachment (Vision Engineering Ltd., UK). The National Institute of Health (NIH) software Image J was used to assess results. The gouge hardness and scratch hardness were recorded as stipulated by the ASTM standard. The gouge hardness was scaled down from the ASTM guidelines to suit small area films and was defined as the hardest pencil that left the film uncut for 46% of the stroke length. The scratch hardness was the hardest pencil that did not rupture or scratch the film. Each test was conducted across six films (each sample was prepared from a separate batch) and standard error was given.

3.2.5.2

Film Adhesion

A modified ASTM tape adhesion test [58] was used to show the differences in mechanical stability. The Method B - X-cut Tape Test was performed using modified measurements to account for small areas of films deposited on Pt foil. The adhesion testing was carried out with low-medium adhesion Scotch 3M Blue masking tape for delicate surfaces (3M #2080). A 6H sharpened pencil (used in previous study Section 3.2.5.1 and known not to damage underlying Pt) was used to place two incisions in an X across the diameter of the film at 90 ◦ to each other exposing the substrate. A section of tape was placed over


52

each X and pressed to thoroughly adhere it to the damaged section. Tape was left for 5 minutes and then manually removed keeping the tape at a constant angle of 180 ◦ to the film surface as described by ASTM D3359 [58]. A LYNX eyepieceless stereo microscope (Vision Engineering Ltd., UK) with camera attachment was used to obtain images before and after tape removal. NIH software Image J was used to assess the films for loss of coating. A frequency histogram was produced based on the number of pixels present for each colour across the 255 grayscale tones from black to white. The number of pixels in the black region (tones 0 - 127) prior to the application of tape were compared to those following the test and a percentage of coating loss was recorded. The test was conducted across six films (each sample was prepared from a separate batch) and standard error reported.

3.2.6


For in vitro assays the silicone well gaskets were washed as above in section 3.2.3 and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. Wells and Pt slides were placed in a beaker of DI water, covered and autoclaved for 20 min at 121◦ C. All components were air dried in a clean laminar flow hood. The wells were attached to the Pt slides and the system was then clamped in the custom-built sandwich assembly disscussed in 3.2.1.1 to prevent leaking between adjacent well chambers.

Following deposition, films were washed carefully four times with DI water and placed in the incubator overnight covered with the last 150 μL of water to leach water soluble


53

contaminants. To disinfect the films 200 μL of 80% EtOH was placed in each well and left for 40 min. The EtOH was then removed and the films were washed twice with sterile Baxter’s water before being placed under UV for 2 hr.

Fluorescent PC12 cells were used for the characterisation of the basic polymer films and also for controls. This cell line discussed in section 3.2.1.2 is non-adherent and requires laminin or collagen to grow on most tissue culture surfaces. For basic polymer film characterisation, polymers were coated with 5 μg/mL laminin sourced from Engelbreth murine sarcoma (Sigma Aldrich, L2020) in Dulbecco’s phosphate buffered saline (DPBS). Wells were incubated for 12 hr at 37◦ C then washed three times with sterile DPBS prior to cell plating.

To plate cells, flasks of PC12s grown in Roswell Park Memorial Institute medium (RPMI) supplemented with 10% horse serum (HS) and 5% fetal calf serum (FCS), were pipetted into polypropylene centrifuge tubes and allowed to settle for 5 min. When a cell clump was visible the supernatant was removed, leaving approximately 0.5 mL of media. A 1 mL Eppendorf pipette tip was used to add 500 μL of fresh media to the cell suspension. Cells were aspirated gently using the 1 mL pipette tip to break up the clumps through 40 aspiration repetitions. Low serum RPMI with 1% horse serum was used to dilute the cell solution to 5 mL and a 0.5 mL sample was taken for counting. Cells were plated at 20,000 cells/cm2 and cultured for 96 hr in low serum medium supplemented with 50 ng/mL NGF (concentration was optimised thorough tissue culture plastic outgrowth assays found in Appendix C). All cell studies were carried out in an incubator at 37◦ C, 5% carbon dioxide


54

(CO2 ) and 100% humidity. Low serum media was renewed with NGF supplement at 48 hr by exchanging 2/3 or 100 μL of media in each well.

3.2.6.1

Neurite Length Measurement

PC12 cell experiments were analysed through fluorescent micrographs taken on a Carl Zeiss fluorescent microscope with camera attachment. The lengths of individual neurites were traced and calibrated length determined using NIH software Image J, with Neuron J plugin.

Neurites were measured when the length of the projection exceeded a single body length of the cell from which it extended. When neurites branched, one split was considered a continuation of the primary neurite and each alternate split of longer than a cell body in length was considered a new neurite.

3.2.6.2

Neurite Data Analysis and Statistical Evaluation

For each single experiment, each substrate type was run in three duplicate wells. For each well three fluorescent micrographs were taken by scanning the film such that each image was obtained from a different sector (third) of the well area. The cell density and neurite density supported by each substrate is presented as mean ± standard error.

Statistical significance was determined by an analysis of variance (ANOVA) and the Tukey simultaneous testing method 2 . A detailed description of the analysis design and calcu2

All advanced statistical analyses were performed by Dr Ross Odell, GSBmE, UNSW

3.3. EXPERIMENTAL RESULTS

55

lated values can be found in Appendix D. Cell numbers and measured neurite lengths were ranked and a pairwise analysis for each material type was conducted. A p-value was determined and a maximum value of 0.05 was used to indicate significance.

3.3

Experimental Results

All electrodeposited conducting polymers were opaque films. Films evolved from pyrrole were black and those formed from EDOT were dark blue. Polymers produced on ITO appeared slightly less cohesive than those produced on Pt, but removal of films from ITO was achieved manually during the washing step with the aid of a scalpel blade. Films fabricated on Pt did not lift during the washing stage.

SEM and XPS were used to confirm the constitution and morphology of the electrodeposited polymers. Sample SEMs of each film type are shown in Figure 3.6 at 1,000x and 15,000x magnification. Both PPy films were significantly smoother with less nodules than the correspondingly doped PEDOT films. Polymers doped with pTS appeared to be rougher, than those doped with PSS. At 1,000x magnification PPy/pTS had a nodular appearance, but at 15,000x magnification the nodules were revealed to be quite smooth. PPy/PSS was relatively featureless at 1,000x magnification, with sparse globular structures, protruding from the surface. At higher magnification it was observed that the film surface was comprised of nodular particles approximately 100 nm in diameter. Both PEDOT films were also comprised of agglomerated nodules, with pTS doped PEDOT having a rough topography with large crevices between nodular peaks.

56


Figure 3.6: SEMs at 1000x (top) and 15000x (bottom) magnification: i. PPy/pTS; ii. PPy/PSS; iii. PEDOT/pTS and iv. PEDOT/PSS.


57

XPS analysis was used to determine the elemental make up and relative quantities of the polymer components. A sample XPS plot of PPy/PSS and PEDOT/pTS is respresented in Figure 3.7. From these curves and associated data the doping ratios were calculated for each polymer. The doping ratio is used as an indication of the polymerisation efficiency, with typical values for conducting polymers ranging between 0.2 and 0.4, indicating that between 5 and 2.5 monomer units are stabilised by each dopant unit.

Figure 3.7: X-ray photoelectron spectra of i. PPy/PSS and ii. PEDOT/pTS depicting the relative elemental constitution of each polymer film. Unique elements N for PPy and S for PEDOT were used to calculate the doping ratio of each film type.

For pyrrole based polymers the size of the peaks seen in Figure 3.7i are proportional to the amount of each element present in the material. Since nitrogen (N) is unique to the pyrrole monomer and sulfur (S) is unique to the dopant (both PSS and pTS) these peaks alone are used to calculate the doping ratio. However, for EDOT based polymers both the monomer unit and dopant contain sulfur, with no element unique to either component. In this case the curve is analysed further to determine the type of sulfur, according to the

58


binding energies that comprise the sulfur peak.

A magnified section of the spectra for PEDOT/pTS is shown in Figure 3.8 with the important peaks marked. The uppermost curve that envelopes all other curves is the raw signal obtained during XPS. The curves below are fitted to the data and allow the user to obtain values that represent the amount of each type of sulfur present in the signal relative to the overall film constitution. Peak B corresponds to S(VI) with a binding energy of 168eV. This bond is seen in sulfur which is bound as SO3 and is present in the structure of sulfonate dopants, including PSS and pTS. At the lower binding energy of 164eV the sulfur bond corresponds to the -C-S-C- bonding observed in EDOT. Both Peak E and Peak F are identified as -C-S-C- bound species. Peak F corresponds to the sulfur component typically bound in the EDOT monomer, but additional -C-S-C- bonding is identified in Peak E. Peak E is considered a small variation in the species where additional hydrogen bonding or a break in the aromatic ring has resulted in a minor peak shift. This peak is still indicative of the bonding seen in EDOT and hence is still considered a component of the PEDOT film. These peaks are relative to the amount of each sulfur type seen on the polymer surface and are used to determine the doping ratio of PEDOT materials.

The area beneath the peak obtained during XPS is calculated as a percentage of the signal detected at the surface and from these values the doping ratio for each polymer was recorded in Table 3.2.

The polymer doping ratios reveal that all polymerisations by electrochemical deposition


59

Figure 3.8: Detailed spectral analysis of sulfur in PEDOT/pTS, used to ascertain the relative constitution of -C-S-C- bound sulfur in PEDOT (Peak E & F) and S(VI) bound sulfur in pTS (Peak B). The overlying curve is the raw data as obtained during XPS analysis. The curves fitted underneath are representative of the peaks that are required to create the raw data and are used to calculate the relative amounts of each element and bonding type.


60

Table 3.2: Doping ratio of basic polymers calculated from XPS analysis, (n=3). Polymer Doping Ratio SD PPy/PSS 0.32 +/- 0.12 PPy/pTS 1.18 +/- 0.07 PEDOT/PSS 0.40 +/- 0.18 PEDOT/pTS 0.23 +/- 0.06

were reproducible with low values of error seen across all materials. The doping ratio was within the acceptable range for all polymers except PPy/pTS, which had an unusually high ratio. These results could be an indication of poor polymerisation, however, it was noted that the film formed could be easily removed from the ITO substrate and handled as a stand-alone film. Additionally, the SD is low indicating a repeatable procedure. It is plausible that excess pTS was present on the surface of the film despite multiple washing steps and overnight leaching in an incubator.

3.3.1


CV revealed that continuous cycling of conducting polymers leads to a decrease in electroactivity for all polymer-dopant combinations. Typical curves showing CV progression for cycles 1, 100, 200, 300 and 400 are represented in Figure 3.9 for PPy/PSS and PEDOT/pTS. It was observed that the general shape of the curve was determined primarily by the monomer type, with PPy/pTS and PEDOT/PSS producing curves with similar shape to PPy/PSS and PEDOT/pTS respectively.

The data obtained from the CVs was used to produce Figure 3.10, depicting how losses in electroactivity progress. Each curve is normalised to the original electroactivity recorded


61

Figure 3.9: CV curves obtained over 400 cycles for i. PPy/PSS and ii. PEDOT/pTS depict the changing shape of the redox peaks as time progresses.

62


in Cycle 1. The majority of activity is lost within the first 100 cycles, with all curves showing a tendency to plateau across the subsequent 300 cycles.

Figure 3.10: Progression of conducting polymer losses of electrochemical activity over 400 cycles. Loss of electroactivity is plotted as a percentage of the original actvity measured at Cycle 1, normalising the polymers to a common origin, (n=3).

The final values for electroactivity loss are presented in Table 3.3, along with the final charge storage density of the films. Larger losses were observed for PPy films compared to PEDOT films. PPy/pTS performed poorly with an average of 68% loss of original activity. PEDOT/pTS showed the least loss of electroactivity with an average loss of less than 12%, corresponding to an 88% preservation. It is important to note that the charge storage capacity of PPy films is still relatively high compared to published values for Pt (6mC/cm2 [20]) despite high losses in electroactivity, due to the large original charge carrying capacity recorded for these materials.


63

Table 3.3: Percentage of electroactivity loss over 400 cycles of CV for basic conducting polymers, (n=3). Polymer Loss of Electroactivity SE Final Charge Storage (% of original activity) (%) Density (mC/cm2 ) a PPy/PSS 61.21 +/- 0.50 65.10 PPy/pTS 68.77 +/- 2.13 53.30 PEDOT/PSS 20.68 +/- 6.87 103.98 PEDOT/pTS 11.99 +/- 3.55 198.02 a

SE for electroactivity is also the SE for reported charge density as both values are derived from the area underneath the CV curve.

Conductivity was recorded with the 4-point probe and averaged across 6 individually prepared films. It was noted that PEDOT films were difficult to remove from the ITO substrate and were significantly more fragile than their PPy counterparts. PPy films appeared flexible and were removed without sustaining any damage. Conductivity measurements recorded in Table 3.4 support the CV data. The polymers performed similarly with PEDOT recording significantly higher values than PPy. Table 3.4: Conductivity of basic conducting polymer films, (n=6). Film Thickness (μm) Conductivity (S/cm) SE PPy/PSS 3.1 22.40 +/- 2.77 PPy/pTS 2.7 14.83 +/- 0.74 PEDOT/PSS 2.24 113.55 +/- 9.69 PEDOT/pTS 2.00 122.68 +/- 9.22

Finally, the impedance measurements were recorded and compared to bare Pt on Bode plots representing impedance magnitude and phase lag (Figure 3.11). It was shown that impedance magnitude of a metal electrode is significantly reduced by the addition of a polymer coating. All basic polymers showed a decrease in impedance at low frequency, when compared to Pt. Once again PEDOT showed superior performance, maintaining the

64


lowest impedance as frequency was increased. The phase lag plot indicates that the cutoff frequency (or corner frequency) is at a significantly lower frequency for polymer coatings than for bare Pt electrodes. The low frequency range is of importance as biological signals recorded from the body indicate they are produced between 300 and 1000 Hz [30]. In this region the polymers were shown to reduce impedance by up to 80% in both magnitude and phase lag. The Bode plot has very small error bars indicating good consistency in preparation of the films and repeatability of results.

3.3.2


An ASTM defined analysis was used to determine film hardness of the polymers deposited on the platinum substrate. The assay defines hardness of organic coatings on metallic substrates relative to calibrated pencil hardness. Two levels of hardness are defined gouge hardness (the hardest pencil that will remove less than 46% of the test stroke) and scratch hardness (the hardest pencil which removes none of the test stroke). It can be seen in Figure 3.12 that gouge hardness is one degree of hardness above scratch hardness for all pTS doped coatings and two degrees of hardness above scratch hardness for PSS doped coatings. Polymers doped with PSS were softer than those doped with pTS for both polymer types. PPy was in general a harder polymer than PEDOT.

ASTM x-cut adherence assays showed that higher surface areas of PPy were delaminated compared to PEDOT, however, it was noted that loss of polymer in the PEDOT films did not occur specifically at the x-cut but rather small areas across the film were removed, as


65

Figure 3.11: Bode plot of impedance spectroscopy for basic polymers. Impedance magnitude is plotted in the upper graph, with phase lag plotted below. Error bars are marked at one standard deviation. The black triangles represent the Pt control typical of a common neuroprosthetic electrode, (n=3).

66


Figure 3.12: Basic polymer hardness as measured by the ASTM hardness test for coatings. Gouge hardness is representative of a minor disruption to the polymer surface and scratch hardness depicts the level at which no polymer is removed from the Pt substrate, (n=6).

shown in Figure 3.13 sample images. Analysis of the images using NIH software package ImageJ, allowed the percentage of coating removed to be plotted as a percentage of the original coating coverage in Figure 3.14.

For both monomers it was observed that doping with pTS appeared to result in more adherent films. By looking at both mechanical assays, it can be stated that pTS dopants resulted in harder, more adherent films than PSS.

3.3.3


Cell culture assays revealed that PC12s could be grown on opaque, thick films (2-3μm) of all polymers with results varying both with polymer and dopant type. PEDOT films performed well when doped with both PSS and pTS, however, only PSS doped PPy


67

Figure 3.13: Sample of x-cut assay images following application of low-adhesion tape to disrupted portion of the film: i.PPy/PSS; ii.PPy/pTS; iii.PEDOT/PSS; and iv.PEDOT/pTS.

Figure 3.14: Film delamination shown as a percentage of the original film following ASTM x-cut analysis for basic polymers, (n=6).

68


showed significant neurite outgrowth. PPy/pTS had very little neurite outgrowth and a dramatically reduced cell number compared the nominal plating density. The majority of polymers showed improved cell and neurite densities compared to Pt, as represented in Figure 3.15. Sample micrographs in Figure 3.16 depict the typical cell response observed on each substrate.

Figure 3.15: PC12 neurite outgrowth on basic polymer substrates at 96hrs post-plating. Cell density and neurite length supported per cm2 are represented for each polymer type with Pt as the control substrate. Error bars represent the SE, (n=3); (∗ NSD); (∗∗ p < 0.05).

Statistical significance was determined through ANOVA and the Tukey test. Since significant differences were observed for the cell density data on nearly all materials (p < 0.05), those without significant differences (or no significant difference, NSD) were noted. For


69

Figure 3.16: Sample images of PC12 neurite outgrowth on basic polymer substrates at 96 hr post-plating: i. Bare Pt; ii. PPy/PSS; iii. PPy/pTS; iv. PEDOT/PSS and v. PEDOT/pTS.


70

cell densities, Pt was found to have NSD to PPy/PSS. However, a significantly greater cell density was marked for PEDOT compared to PPy/PSS. For neurite length supported per cm2 and cell density there was NSD between both PEDOT films. All other substrates were significantly different when compared in a pairwise analysis against each of the other substrates, with PPy/pTS performing significantly worse than all other substrates including Pt.

3.4

Summary of Results

Table 3.5 summarises the properties of each polymer type to aid in the comparison of materials. Where quantitative values are important, they are given. For other properties a qualitative interpretation of a range of values is used to indicate the polymer’s overall performance.

Analysis Surface Morphology Efficiency of Deposition Electrochemical Stability Conductivity (S/cm) Impedance (Ω at 1kHz) Hardness (Scratch) Adherence (% Loss) Neurite Outgrowth

Table 3.5: Summary of results PPy/PSS PPy/pTS PEDOT/PSS Rough Smooth Rough Good Poor Good Poor Poor Good 22.4 14.8 113.6 436.5 200.9 135.7 2H 4H B 51.11 4.52 20.43 Satisfactory Poor Good

PEDOT/pTS Very rough Good Good 122.7 138.7 3H 0.52 Good

3.5. DISCUSSION

3.5

71

Discussion

Polymers PPy and PEDOT are commonly explored in neural interfaces literature. In this chapter PPy and PEDOT were doped with conventional sulfonate dopants PSS and pTS, and analysed across a range of properties, including long-term electrical performance, mechanical properties and thick film cell responses not found in the literature. PPy films on the whole had inferior physico-chemical properties compared with PEDOT. Monomer type was the determining factor in formation of surface topography and electrical response, and dopant type appeared to dominate mechanical response. The resulting surface topography and mobility of dopant determined the in vitro cell response.

Electrical analyses revealed that polymers performed consistently in relation to each other across CV, conductivity and impedance measurements. PEDOT was superior to PPy in all electrical characterisations.

CV revealed that polymers can undergo significant losses to electroactivity when subjected to potentials that result in the polymer being continually cycled through oxidation and reduction reactions. The most significant portion is lost in the first 100 cycles with the loss plateauing through subsequent cycles. Yamato et al. have reported that coupling of the α-β’ bond across the PPy backbone leads to instability that results in dopant loss. The subsequent degradation of electroactivity is reflected in these results. PEDOT demonstrates much better electrochemical stability than PPy, due to the dioxyethylene bridging group, which blocks the possibility of α-β’ coupling [33]. Additionally, the results

72


of this study show that the larger dopant PSS (≈70kDa) is retained within the PPy matrix better than the smaller pTS (196Da). However, the reverse is true for PEDOT with pTS having only 12% loss compared to PSS with 20%. A possible explanation is that additional aromatic present on the PEDOT backbone creates a matrix in which PSS is not coupled as efficiently during elecrtodeposition as the significantly smaller pTS. This is supported by the charge density data which indicates that PEDOT/PSS has approximately half the charge carrying capacity of PEDOT/pTS and a lower conductivity.

Other electrical analyses show concurring data with PPy having a lower conductivity and higher impedance magnitude across the 100 - 1000 Hz range than PEDOT for both dopant types. The low frequency range is of importance as biological signals recorded from the body are most commonly observed between 300 and 1000 Hz [30]. These results coupled with the SEMs demonstrate the well defined phenomenon: that rougher materials with higher effective interfacial area have better electrical characteristics [30, 31, 45]. PEDOT films appear significantly rougher with a greater number of nodules than their PPy counterparts and as such higher conductivities and lower impedance magnitudes are observed. Literature reports that impedance magnitude of bare gold microelectrodes can be reduced by up to two orders of magnitude with conducting polymer coatings with an 85% reduction of magnitude seen at 1 kHz [20]. Similar results are seen in Table 3.6 for this research, where PEDOT/PSS is shown to reduce impedance magnitude at 1kHz by 80%.

Yang and Martin have reported that the lowest impedance films are also those that are the

3.5. DISCUSSION

73

Table 3.6: Impedance of basic conducting polymer films on Pt at 1 kHz. Electrode Coating Magnitude (Ω) Phase Lag (◦ ) Bare Pt 629.8 8.6 PPy/PSS 436.5 4.1 PPy/pTS 200.9 2.2 PEDOT/PSS 135.7 2.4 PEDOT/pTS 138.7 3.7

softest [45]. This is consistent with the ASTM hardness assays where PEDOT coatings were significantly softer than their PPy counterparts. Mechanical assays demonstrated that these properties are strongly dependent on the polymer dopants. Polymers doped with PSS were softer but also less adherent to the Pt substrate than the polymers doped with pTS. While a softer material is desirable to reduce strain mismatch at the tissue interface, the coating must also be durable and adherent.

The two parameters chosen to describe biological interaction with polymers were cell density and neurite outgrowth per cm2 . Although most PC12 neurite outgrowth assays report percentage of cells with outgrowth and average neurite length [12, 32, 77, 78], when cultured on conducting polymers these values did not present significant differences between materials. The clumping of cells commonly seen in PC12 cultures limits the percentage of cells capable of producing outgrowth and when stimulated by NGF the percentage of cells with outgrowth is most often between 30 and 40% [12, 78, 79]. Similarly, when neurites are measured as described in Section 3.2.6.1 the average neurite length is not meaningful unless coupled with the degree of branching or the total number of neurites [32, 78, 80]. In contrast, the cell density, indicative of the number of cells adhering to the substrate and the summation of neurite length across a given area can demonstrate the significant

74


differences between materials as demonstrated by Figure 3.15.

Cell culture assays demonstrated that all polymers can support cell growth, with only PPy/pTS demonstrating significant inhibition of cell adherence and growth. It is thought that pTS leaches from the PPy at concentrations that result in toxic death of cells. pTS has a LD50 much lower than that of PSS (1,000 mg/kg compared to 15,000 mg/kg subcutaneously delivered) [61], however, as demonstrated in the CV testing, pTS has greater mobility within the PPy matrix. In a media solution consisting of anionic buffer salts, it is hypothesised that the pTS could easily leach from the PPy matrix where other anionic molecules would replace them. PSS, being a much larger polymer dopant is less mobile and hence cannot escape the PPy matrix in such high concentrations. The XPS results also suggest that excess pTS remains on the PPy/pTS surface following deposition. If a similar constitution is present throughout the polymer, then a significantly higher proportion of pTS would be present in these cell culture samples and would also account for cell growth inhibition and death.

The lower neurite density seen on PPy/PSS compared to both PEDOT polymers could be attributed to the PEDOT surface topography. Literature has demonstrated that cells preferentially grow on rough surfaces [35]. Kim et al. have demonstrated that PPy has a lower neurite outgrowth percentage to PEDOT when identically doped [12]. These assays concur with this finding in two respects. It was demonstrated that rougher PEDOT substrates yielded significantly higher cell and neurite densities compared to PPy (p < 0.05); and polymers in general performed more favourably than bare, relatively smooth

3.6. CONCLUSIONS

75

Pt.

The modification of PPy to increase cell interactions has been explored by Li et al. who have reported the use of a PPy-hydrogel layered composite [81, 82]. The resulting surface modification was shown to support a higher density of PC12s and greater neurite outgrowth than TCP controls. Research by Li et al. suggests that the modification of PPy to enhance surface topography by increasing the surface area and nodularity, will result in a substrate more conducive to cell attachment and growth.

3.6

Conclusions

Preliminary investigations into the use of conventional conductive polymers for neuroprosthetic interfaces demonstrate that a wide variety of properties need to be analysed to prepare a material for an implant device. In characterising a range of basic polymers commonly found in literature, it was noted that an optimised interface will require a trade-off between desired electrical, mechanical, chemical and biological properties.

While PEDOT demonstrated superior performance in many assays, it can be further optimised to improve mechanical performance. The nodular surface topography coupled with the ability to retain dopants long-term make PEDOT the preferred choice for further modification. For PPy to be considered as an electrode coating it will require modifications that prevent the loss of electroactivity and increase the surface roughness to be more conducive to cell interactions.

76


The following chapter will focus on the modification of these basic films to increase longterm electroactivity and improve both impedance and cell interactions through altering the surface topography.

Chapter 4

Improving Polymer Properties through Multi-Walled Carbon Nanotube Layering

[Portions of this chapter are reprinted from R.A. Green, C.M. Williams, N.H. Lovell and L.A. Poole-Warren, Novel neural interface for implant electrodes: improving electroactivity of polypyrrole through MWNT incorporation with permissions from J Mater Sci Mater Med ; 19(4) 1625 – 1629; DOI:10.1007. Accepted 24 January; Copyright 2008 Springer]

4.1

Introduction

An alternate and emerging material for neural tissue engineering is the highly conductive carbon nanotube (CNT). CNTs are graphene sheets rolled into a cylinder with diameters ranging from 30 nm up to 500 nm [83] and lengths ranging from several hundred nanometres to micrometres. Multi-walled carbon nanotubes (MWNTs) are CNTs consist77

78

CHAPTER 4. IMPROVING CPS THROUGH MWNT LAYERING

ing of multiple graphene sheets arranged into concentric cylinders. MWNTs exhibit high mechanical strength (around 30 GPa) [84] and Young’s modulus (up to 4.15 TPa) [85]. They have a high surface area that is beneficial for low electrical impedance and exhibit substantial chemical and thermal stability. These properties have made MWNTs potentially a very attractive material for use in long-term implant devices, especially neural prostheses applications [86].

CNTs have commonly been investigated as an alternative to conducting polymers through dispersion into non-conducting polymer scaffolds, to produce a path through a normally insulative material [86]. CNTs are hypothesised to be a material that can incorporate a combination of high strength, low toxicity and ballistic electrical conductivity to neuroprosthetic materials. However, controlling their fabrication and incorporation into conventional biomaterials is a challenge that has yielded varied results. Their incorporation into polyurethanes (PUs) has shown that they are difficult to disperse evenly throughout a material and the resulting composite has far lower strength and conductivity than the nanotubes alone. Additionally the biological studies have been controversial, despite cytotoxicity studies showing that they are unlikely to be toxic, the nature of their small size indicates that they could perforate cell membranes and inadvertantly cause cell death.

CNTs have recently gained attention in biosensor research, where the incorporation of CNTs into conducting polymer films has resulted in sensors with high sensitivity and selectivity [87]. When incorporated into insulators such as PU, CNTs need to be dispersed throughout the polymer for the duration of the casting process which can take over 48 hr

4.1. INTRODUCTION

79

when a certified PU biopolymer is employed [88]. CNTs are significantly easier to handle when combined with conducting polymers to form a composite. They can be used as the polymer dopant [87], covalently tethered to the monomer [88] or simply entrapped during the polymerisation process [89, 90]. For most of these processes simple aqueous solutions of CNTs in transient suspension are used [87, 90] or alternately conducting polymers are deposited over CNTs grown on the substrate [89]. Research demonstrates that a material with the properties of both components can be obtained from the fabrication of CNT-conducting polymer composites. Additionally, the use of these composites in biosensor applications has shown that a sensor with superior electrochemical stability can be produced [89, 90].

In this study the incorporation of MWNTs into conducting polymer films for neural interfaces aims to improve the long-term electrical properties of PPy films, which were shown in Chapter 3 to have significantly lower conductivity and electrochemical stability than identically doped films of PEDOT. MWNTs also have the potential to impart increased interfacial area to an electrode, providing a lower impedance and an increased roughness conducive to cell attachment. The results of the previous study were not clear in determining whether mechanical/physical or chemical factors are dominant in cell adhesion and growth. Addition of MWNTs to the polymers will alter the physical properties while maintaining the same chemical composition at the cell interface, allowing the relationship between polymer and cell response to be further examined.

In this research composite films were fabricated in a bilayer, to address concerns over the

80


release of MWNTs that could perforate surrounding neural cells. MWNTs are of a size that would allow them to infiltrate single cells within the body and possibly accumulate, having a detrimental effect similar to that of asbestos. Mobile MWNTs in biomedical applications may be harmful through alternate mechanisms such as being undetected by the normal phagocytic defences allowing them access other tissues through the blood or modify protein structures, possibly altering their function or rendering them antigenic [91]. The bilayer configuration with a base layer of nanotubes and an overlying polymer layer, minimises the possibility of MWNT release into surrounding tissue. This approach also attempts to disperse the MWNTs evenly across the electrode surface to result in a composite with uniform properties across the film.

The specific aims of this study were to:

1. Prepare MWNT - conductive polymer composites in a bi-layer with nanotubes constrained by an overlying polymer layer,

2. Assess the effect of the MWNT layer on established conducting polymer physicochemical properties, and

3. Examine the effect of altering conducting polymer surface topography through MWNT layering on mammalian cell interactions using a neural-like cell line.


4.2

4.2.1

81

Materials and Methods

Multi-Walled Carbon Nanotube Conducting Polymer Composites

MWNTs with carboxylic acid functionality were obtained from Nanocyl (Cat. #3151). These nanotubes were used as they were able to be temporarily suspended in an aqueous solution without the use of harsh solvents. An aqueous solution (Millipore filtered DI water) of MWNTs was made up to 0.5 mg/mL and sonicated for 60 min to obtain a transient suspension. While suspended, a 150 μL aliquot, or 75μg of the sonicated MWNTs were placed in each 5mm diameter well and air dried in a laminar flow hood overnight. MWNTs dried slowly at room temperature to produce a layer that appeared confluent and uniform when viewed macroscopically. Conducting polymer solutions required careful pipetting to prevent disruption of the MWNT layer prior to electrodeposition. Where the MWNTs dried with uneven distribution or were disturbed during preparation for electrodeposition, the film was discarded. Solutions of PPy/PSS, PPy/pTS, PEDOT/PSS and PEDOT/pTS were produced and electrodeposited as described in Section 3.2.3.

Films were produced on both ITO and Pt, with analytical studies being performed as previously detailed in Section 3.2. All characterisation studies conducted on homogeneous polymers in the previous chapter were repeated for MWNT layered polymer composites. Cell outgrowth assays were conducted on MWNT layered polymers and their homogeneous counterparts were used as controls. Additionally, an uncoated MWNT layer dried on a Pt substrate with no additional processing was used to determine the possible impact of

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free MWNTs on cell growth. Results were recorded and where appropriate compared to values obtained in Chapter 3.

4.3


Overlying PPy and PEDOT films were evolved on the MWNTs, and subsequent washing demonstrated that the MWNTs were well constrained by the polymer matrix. Some free MWNTs were removed in the first washing cycle, however subsequent washes were not discoloured by their presence. Polymers produced on ITO were significantly less cohesive than those produced on Pt. Several attempts were required to produce samples on ITO as the MWNTs did not adhere well during drying. Samples were more easily produced on the Pt, which appeared to provide a surface more conducive to adherence of the MWNTs.

SEMs revealed that the MWNTs significantly changed the surface characterisitics of the polymer films. An SEM of the MWNTs prior to deposition of the overlying polymer is shown in Figure 4.1. This demonstrates that MWNTs are not confluent when airdried with the underlying Pt visible across a number of pores, ranging from 5 - 50 μm in diameter. At 15,000x magnification single MWNTs are discernable, and it is revealed that the MWNTs are clumped, with nodular aggregations across the electrode.

Following electrodeposition of the polymers, the SEMs reveal the extent of surface modification brought about by the underlying MWNTs. MWNT-PPy composites, depicted in Figure 4.2 demonstrate that the MWNTs greatly increase the surface roughness and


83

Figure 4.1: SEM of air-dried MWNTs at i.1000x and ii. 15000x magnification. presence of nodular features on the film surface. PPy/pTs is of particular interest, as the image at 15,000x magnification reveals fibrillar surface characteristics. PEDOT composites shown in Figure 4.3 also appear to have a greater degree of surface roughness, with PEDOT/PSS developing fibrillar extensions similar to PPy/pTS. While MWNTPEDOT/pTS, has a more nodular surface when viewed at 1000x magnification compared to homogeneous PEDOT/pTS, the surface characteristics at 15,000x magnification appear unaltered. It is noted that for all polymer composites there is no underlying Pt visible in the SEMs, indicating that during electrodeposition the pores previously observed in Figure 4.1 are filled by the polymer.

XPS analysis confirmed the presence of the polymer layer over the MWNTs. A bare MWNT layer was used as a control, producing the spectra in Figure 4.4(i), with approximately 84% pure carbon. Other elements present are due to the -COOH functionalisation of the commercially supplied MWNTs. Figure 4.4(ii) shows the surface of the

84


Figure 4.2: SEMs at 1000x (top) and 15000x (bottom) magnification: i. PPy/PSS; ii. MWNTs-PPy/PSS; iii. PPy/pTS and iv. MWNTs-PPy/pTS.


85

Figure 4.3: SEMs at 1000x (top) and 15000x (bottom) magnification: i. PEDOT/PSS; ii. MWNTs-PEDOT/PSS; iii. PEDOT/pTS and iv. MWNTs-PEDOT/pTS.

86


MWNTs-PPy/PSS composite film. This spectra, like all the polymer composites was very similar to the spectra obtained in Chapter 3 for the homogeneous polymer films. As a

Figure 4.4: X-ray photoelectron spectra of i. Bare MWNTs and ii. MWNT-PPy/PSS depicting the relative elemental constitution of each layer. No nitrogen or sulfur signal is seen in the MWNTs, which consist predominantly of carbon with a smaller oxygen peak due to the -COOH functionalisation. result, the corresponding doping ratios, given in Table 4.1 for the basic polymer films and their respective MWNT composites are not significantly different, with the exception of PPy/pTS. When produced as a composite film PPy/pTS had an improved doping ratio of 0.34, compared to 1.18 obtained for homogenous PPy/pTS. Table 4.1: Doping ratio of MWNT layered polymers compared to coatings, (n=3). Polymer Doping Ratio PPy/PSS 0.32 +/MWNTs-PPy/PSS 0.21 +/PPy/pTS 1.18 +/MWNTs-PPy/pTS 0.34 +/PEDOT/PSS 0.40 +/MWNTs-PEDOT/PSS 0.48 +/PEDOT/pTS 0.23 +/MWNTs-PEDOT/pTS 0.22 +/-

homogeneous polymer SD 0.12 0.07 0.07 0.13 0.18 0.04 0.06 0.02


4.3.1

87


Cyclic voltammetry curves had the same shape as reported in Section 3.3.1, with oxidation and reduction peaks occurring at the same monomer defined potentials. Electroactivity shown in Figure 4.5 degraded steadily across the first 200 cycles and tended to plateau in the subsequent 200 cycles. The peaks were shown to reduce in size as the cycles progressed, however, as seen in Table 4.2 the loss of electroactivity for most polymers was less than that seen for their homogeneous counterparts. Only PEDOT/pTS showed an increase in loss of electroactivity. Table 4.2: Comparison of electroactivity loss over 400 cycles of CV and final charge storage density for homogeneous polymers and MWNT composites, (n=3). Polymer Loss of Electroactivity SE Charge Storage (% of original activity) (%) Density (mC/cm( 2))a PPy/PSS 61.21 +/- 0.50 65.10 MWNT-PPy/PSS 32.54 +/- 3.82 108.48 PPy/pTS 68.77 +/- 2.13 53.30 MWNT-PPy/pTS 42.51 +/- 2.90 105.52 PEDOT/PSS 20.68 +/- 6.87 103.98 MWNT-PEDOT/PSS 10.23 +/- 3.14 131.74 PEDOT/pTS 11.99 +/- 3.55 198.02 MWNT-PEDOT/pTS 17.91 +/- 1.90 238.08 a

SE for electroactivity is also the SE for reported charge density as both values are derived from the area underneath the CV curve.

The most significant difference in electroactivity is seen in the PPy composites. The loss of electroactivity is shown to be markedly improved, but it is important to observe that the polymer still undergoes a significant reduction in redox activity. Figure 4.6, depicting 400 cycles of CV applied to a MWNT-PPy/PSS film, demonstrates the loss of redox activity observed. The anodic and cathodic peaks at Cycle 1 were pronounced and subsequent

88


Figure 4.5: Electrochemical activity loss for MWNT-polymer composites over 400 cycles of redox. Loss of electroactivity is plotted as a percentage of the original actvity measured at Cycle 1, normalising the polymers to a common baseline, (n=3).

cycles showed degradation. At the end of 400 cycles smaller broad peaks are observed, but in contrast to the homogeneous polymer, a high percentage of the area within the curve and hence electroactivity is preserved.

Conductivity was recorded from 4-point probe and averaged across 6 individually prepared films with standard error reported. Layering the films with MWNTs made it difficult to manually remove all films from the ITO, with PEDOT/PSS unable to be lifted without incurring substantial damage. PPy composites were the most flexible and were removed


89

Figure 4.6: An example of a typical MWNT-PPy/PSS composite undergoing 400 successive cycles of CV demonstrates the degradation of redox peaks over time.


90

without sustaining any damage. PEDOT/pTS was removed with the aid of a razor blade with minimal damage around the edges. Conductivity measurements recorded in Table 4.3 show the increase in conductivity that can be achieved across all polymers with MWNT layering. The greatest increase was recorded for PPy/pTS where conductivity was doubled and a statistically significant increase was observed for the MWNT modification (p < 0.05). Thickness measurements were made using optical profilometry on films prior to removal from the substrate. The MWNT layer was measured prior to coating with polymers and was shown to be approximately 13.5μm thick. Measurement was difficult due to the large differences in uniformity. The average thickness across a 5mm length was calculated for all composite materials. Table 4.3: Comparison of conductivity of homogeneous and MWNT layered conducting polymer films, (n=6). Film Thickness (μm) Conductivity (S/cm) SE PPy/PSS 3.1 22.40 +/- 2.77 PPy/PSS layered MWNTs 15.04 26.07 +/- 3.57 PPy/pTS 2.70 14.83 a +/- 0.74 a PPy/pTS layered MWNTs 14.91 31.46 +/- 2.48 PEDOT/PSS 2.24 113.55 +/- 9.69 b PEDOT/PSS layered MWNTs 15.80 PEDOT/pTS 2.00 122.68 +/- 9.22 PEDOT/pTS layered MWNTs 15.08 149.05 +/- 7.71 a

Statistically significant increase; p < 0.05. PEDOT/PSS films layered with MWNTs were unable to be removed from the ITO substrate and hence no conductivity reading was determined. b

Impedance spectroscopy results, plotted in Figure 4.7 demonstrated that MWNT polymer composites dramatically reduced the impedance magnitude and phase lag seen at the electrode surface compared to bare Pt. PSS doped polymer composites had very low magnitudes at the starting frequency of 10 Hz compared to other materials, however


91

the impedance magnitude of all polymers converged from 500 Hz onwards to plateau around 100 Ω. MWNT-PEDOT/pTS showed the lowest impedance at 1000 Hz, reducing the impedance magnitude recorded from Pt by 89% (1% below an order of magnitude). The phase lag was also significantly reduced across all substrates with the maximum lag recorded at 10 Hz being just below 30◦ or less than half the lag measured from Pt.

Compared to previously established measurements for homogeneous polymers the MWNT layered polymers had lower impedance magnitudes across the biological focal region of 300 - 1000 Hz. The phase lags at 1 kHz, recorded in Table 4.4, were not significantly different for any polymer composite. Table 4.4: Impedance of MWNT composite polymers compared to their homogeneous counterparts at 1 kHz, (n=3). Electrode Coating Magnitude (Ω) Phase (◦ ) Bare Pt 629.8 -8.6 PPy/PSS 436.5 -4.1 MWNTs - PPy/PSS 113.6 -4.3 PPy/pTS 200.9 -2.2 MWNTs - PPy/pTS 176.9 -3.12 PEDOT/PSS 135.7 -2.4 MWNTs-PEDOT/PSS 126.5 -3.1 PEDOT/pTS 138.7 -3.7 MWNTs-PEDOT/pTS 71.3 -3

4.3.2


ASTM defined tests for metal coatings were used to determine the effect of a MWNT layer on the mechanical properties of a conducting polymer. The results of the hardness test, represented in Figure 4.8 revealed that the MWNT inclusion had the greatest impact on

92


Figure 4.7: Bode plot of impedance spectroscopy for MWNT layered polymers. Impedance magnitude is plotted in the upper graph, with phase lag plotted below. Error bars are marked at one standard deviation, (n=3).


93

the hardest materials. The pTS doped polymers were significantly reduced in both gouge and scratch hardness. The PSS doped polymers that were found to be softer polymers in Section 3.3.2, did not appear to be affected by the MWNT inclusion.

Figure 4.8: MWNT composite polymer hardness compared to homogeneous polymers as measured by the ASTM hardness test for coatings. Gouge hardness is representative of a minor disruption to the polymer surface and scratch hardness depicts the level at which no polymer is removed from the Pt substrate. Error bars represent SE, (n=6).

The x-cut adherence test revealed that delamination was affected by the MWNTs. The graph in Figure 4.9 indicates that delamination problems seen with PPy/PSS are improved by the addition of the MWNT layer, however, looking at the image in Figure 4.10 it is observed that while delamination did not occur at the x-cut, there was significant loss of polymer across the film surface. The loss of small areas of polymer across the coating is also seen in MWNT-PPy/pTS and MWNT-PEDOT/PSS, but the percentage of coating

94


lost during the tape removal does not differ significantly from the amount removed from their homogeneous counterparts. It is important to note, that while the percentage of material delaminated from MWNT-PEDOT/pTS is low, it is still higher than observed for PEDOT/pTS, and hence it should be considered that the MWNT layer in this instance degrades its mechanical adherence.

Figure 4.9: Comparison of film delamination by ASTM x-cut analysis for MWNT composite polymers and their homogeneous controls. Error is represented as SE, (n=6).


95

Figure 4.10: Sample of x-cut assay images following application of low-adhesion tape to disrupted portion of the MWNT composite film: i.MWNT-PPy/PSS; ii.MWNTPPy/pTS; iii.MWNT-PEDOT/PSS; and iv.MWNT-PEDOT/pTS.

4.3.3


Due to their nanoscale sizing it is possible that mobile nanotubes could potentially threaten cell survival by perforating their membrane. The design used in this research aims to minimise the possible threat of cell perforation by constraining the MWNTs beneath a layer of conducting polymer. An uncoated MWNT layer was also assessed for control purposes and provided an insight into the possible effect free MWNTs may have on cells. Other control films included bare Pt and homogeneous polymers without MWNTs.

Sample images of MWNT layered polymers compared to bare Pt and an uncoated MWNT layer are shown in Figure 4.11. Results from similar images taken across three studies were

96


collated and graphed in Figure 4.12 and Figure 4.13 to quantify cell density and neurite length supported per cm2 respectively, following 96 hr of incubation on the material substrate. It is important to note that no negative impact resulted from the incorporation of MWNTs. In this study similar results are seen for the electrode material Pt, which is approved for use in commercial implant devices [92], as for the MWNT coated Pt.

Cell adhesion and neurite outgrowth was observed on all MWNT layered polymers. MWNT layered PPy/pTS films were far superior to homogeneous PPy/pTS, with significantly improved cell density and neurite outgrowth. It was noted that despite the significant improvement MWNTs imparted to this film, pictured in Figure 4.11(iv), cell growth on this film was still less than that observed for other MWNT modified polymers depicted in Figure 4.11 and the unmodified polymers, not pictured, but represented in Figures 4.12 and 4.13. PPy/PSS showed an increase in both cell density and neurite outgrowth, but NSD was seen for either property on PEDOT/PSS and PEDOT/pTS substrates.

It was observed from Figure 4.11 and Figure 4.12 that cells grown on PEDOT composites appeared to have neurite growth across the entire substrate and cell densities were close to the nominal 20,000 cells/cm2 plating density. These results suggest that neurite outgrowth is limited by other factors including available growth area, available NGF and proximity of adjacent cells, and hence it is propounded that cell response cannot be further improved within the limits of this study.


97

Figure 4.11: Sample images of PC12 neurite outgrowth on MWNT composite polymers at 96hrs post-plating: i. Bare Pt; ii. MWNTs; iii. MWNT-PPy/PSS; iv. MWNT-PPy/pTS; v. MWNT-PEDOT/PSS and vi. MWNT-PEDOT/pTS.

98


Figure 4.12: PC12 cell density on MWNT-polymer substrates at 96 hr post-plating. MWNT layered polymers are shown beside their homogeneous polymers. Error is represented as SE, (n=3); (∗ p < 0.05).

Figure 4.13: Neurite outgrowth of PC12s on MWNT-polymer composites at 96hrs postplating. Neurite outgrowth is represented as the total neurite length supported per cm2 . SE is given, (n=3); (∗ p < 0.05).

4.4. SUMMARY OF RESULTS

4.4

99

Summary of Results

Table 4.5 summarises the properties of each polymer type to aid in the comparison of materials. Where quantitative values are important, they are given. For other properties a qualitative interpretation of a range of values is used to indicate the polymer’s overall performance. Table 4.5: Summary of results for Analysis PPy/PSS Surface Morphology Rough Efficiency of Deposition Good Electrochemical Stability Satisfactory Conductivity (S/cm) 26.07 Impedance (Ω at 1kHz) 113.6 Hardness (Scratch) 2H Adherence (% Loss) 16.00 Neurite Outgrowth Good a

MWNT layered conducting polymers PPy/pTS PEDOT/PSS PEDOT/pTS Rough Very Rough Very rough Good Good Good Satisfactory Good Good a 31.46 149.05 176.9 126.5 71.3 2H B B 3.26 21.44 3.69 Satisfactory Excellent Excellent

No measurement obtained due to poor mechanical stability of stand-alone film.

4.5

Discussion

In the previous study PPy was shown to be inferior to PEDOT across a range of properties and most significantly in long-term electrochemical stability. PPy also had a poor cell response that was attributed to a comparatively featureless surface topography and mobility of dopant allowing toxic build-up and consequent cell death. It was hypothesised that MWNTs could be used to improve the electrical properties and physical morphology of conducting polymers, specifically PPy which performed poorly in these criteria. While MWNTs were applied to both PPy and PEDOT with both dopants, pTS and PSS, the

100


most noteable improvements were seen in PPy.

CV revealed that the long-term electroactivity of PPy could be significantly improved by the incorporation of MWNTs. Long-term electroactivity of the PPy/PSS films can be improved by up to 50% through the incorporation of MWNTs. Films that did not incorporate MWNTs were shown to lose 61% of activity over the course of 400 cycles. This concurs with results published by Yamato et al. who saw losses of 95% when PPy/PSS films were polarised for 16 hr [33]. When MWNTs were incorporated only 34% of electroactivity was lost. While the MWNTs cannot prevent the oxidation of the PPy backbone and loss of dopant described by Schlenoff and Hong [93], they can provide a conductive path that will not be subject to oxidative degradation, there by increasing the overall current carrying capacity of the composite material. The higher conductivity of all MWNT layered polymers suggests that the MWNT layer is dominant in carrying charge throughout the composite.

Mechanical assays showed that despite the high strength of CNTs, their presence in the conducting polymer matrices softened the overall composite. Although a softer polymer is desired to reduce strain mismatch at the tissue-electrode interface, this effect was not expected. The MWNT layering of PEDOT films also resulted in decreased stability of the stand-alone film. PEDOT/PSS was reduced to a powder when lifted from the ITO microscope slides. These results are most likely due to the disorder of the MWNTs when airdried on the Pt substrate and the resultant interaction of the MWNTs with the polymer matrices. The polymers doped with the larger PSS anion and and those formed from the

4.5. DISCUSSION

101

larger monomer, EDOT, were the most difficult to handle, indicating that these polymers interacted poorly with the rougher more nodular substrate. However, PPy/pTS was shown to form more efficiently on the MWNTs as indicated by its improved doping ratio of 0.21 compared to 1.18 for the film deposited directly on to Pt. It is hypothesised that the smallest monomer/dopant combination was able to use the increased surface area to polymerise more efficiently through increased access to available electrons. Additionally the smaller molecules could retain mobility around the protruding surface MWNTs, where the larger chains of PSS were more restricted, forcing a less efficient deposition at the MWNT-polymer interface. These results were not reflected in the doping ratios as XPS can only determine the chemical composition to a depth of nanometres [94].

Reduced adherence was also observed for MWNT-PEDOT composites. Since MWNTs are air dried onto the Pt substrate, adherence of the film is only formed where the polymer interacts with the underlying metal. By inspecting the SEM of the MWNTs it is clearly seen that there are less contact areas for the polymer to interact with the Pt. It is also important to note that this method of MWNT application results in a film with a much greater thickness than achieved with the polymers alone. The clumping of the MWNTs on drying results in areas where the film is 5 times thicker than an homogeneous coating, but the overlying polymer should be the same thickness as a result of identical electropolymerisation parameters. The resulting MWNT-polymer composite is predominantly composed of non-adhered MWNTs, explaining the reduced adherence seen in these films. It is hypothesised that a different technique of applying MWNTs, such as direct

102


growth on the electrodes, is required to produce a more uniform film with more equal volumes of each component. A more ordered composite would allow the benefits of each component to be more readily accessed.

Increased surface nodularity was observed for all MWNT layered polymers, but once again was most noticeable in the PPy modifications. The increased roughness was visually observed in the SEM and is also thought to be responsible for decreased impedance magnitudes. It is also thought that surface roughness was a contributing factor of the improved PC12 adherence and neurite outgrowth response seen in PPy composites. Significant improvement in cell density and neurite outgrowth was recorded for PPy/pTS when layered with MWNTs (p < 0.05). A less prominent increase in cell interactions was recorded for MWNT layered PPy/PSS. It is thought that PPy/pTS formed more efficiently on the MWNT layer with a reduced number of pTS ions per PPy monomer, as discussed above and indicated by the doping ratios. The reduced number of pTS ions is thought to have lowered their concentration to a level where pTS toxicity has less impact on cell growth. It is noted, however, that MWNT-PPy/pTS at best performs similarly to bare Pt, despite the increased surface roughness which has been shown to increase cell adherence and growth [12, 35].

An important feature of this design was the constraint of the MWNTs within the polymer matrix preventing accumulation or mobilisation of MWNTs within the surrounding cellular tissues. In this way the MWNT-polymer composite could maintain the superior electronic properties of the MWNTs without compromising the surrounding cellular en-

4.6. CONCLUSIONS

103

vironment. Despite this, neurite outgrowth assays demonstrated that PC12 growth and survival were unaffected by free MWNTs when cultured directly on top of them. NSD was seen between the Pt and MWNT substrates. These results indicate that incoporation of MWNTs into a biological system should not have a significant impact on the surrounding cellular environment.

MWNT modification of conductive polymers shows promise for neural interfaces with improved electrical properties, but it is hypothesised that greater benefit would be obtained from MWNT composites if the orientation and distribution of CNTs could be controlled. The growth or alignment of CNTs perpendicular to an electrode substrate would allow the surface topography and mechanical properties to be further optimised for neural tissue interaction.

4.6

Conclusions

Layering of PPy with MWNTs produces a film with far superior electrical properties than PPy alone. The addition of MWNTs significantly improves the electrochemical stability of PPy/PSS and PPy/pTS films for electrode coatings. Additionally, the inclusion of functionalised MWNTs into a biological system should not have a significant impact on cell growth. The incorporation of MWNTs into conductive polymer electrode coatings shows promise for neural interfaces with improved electrical properties. However, MWNTs showed no significant improvements when incorporated into PEDOT. Additionally, homogeneous PEDOT films still performed better across the range of analysed

104


properties than the optimised MWNT-PPy films. Although small improvements were seen in MWNT-PEDOT/pTS conductivity and impedance, the films were observed to be more fragile and less adherent. It is concluded that unmodified PEDOT/pTS is the most suitable polymer for modification of a conventional neural electrode.

The following chapters will focus on optimising PEDOT for cell interactions by incorporating biological signaling factors.

Chapter 5

Conducting Polymers with Cell Attachment Bioactivity 5.1

Introduction

A desirable neural interface would resemble the schematic in Figure 5.1, having the capacity to support tissue outgrowth and promote cell adherence through appropriate protein components. Extracellular matrix molecules are known to support cell attachment and growth when incorporated into conducting polymers or used as a coating.

Figure 5.1: Schematic of a conducting polymer electrode array with cell attachment bioactivity 105

106

CHAPTER 5. CELL ATTACHMENT BIOACTIVITY

When neuroprosthetic electrodes are placed in a cellular environment it is highly desirable to achieve a durable attachment of the neural tissue to the electrode for long-term stimulation and recording [95]. Molecules that have been implicated in these roles for cell regeneration include laminin, chondroitin sulfates and other proteoglycans; and hyaluronic acid (HA) [11, 15, 96].

Cui et al. has shown that adhesion of dorsal root ganglions and fibroblasts to cortical electrode surfaces is improved when the conducting polymer substrate is doped with a laminin peptide [11]. Additionally Collier et al. have demonstrated that PPy-HA composite biomaterials are promising candidates for tissue engineering and wound-healing applications that may benefit from both electrical stimulation and enhanced vascularisation [37]. While a number of cell attachment proteins have been incorporated into conducting polymers or onto the polymer surface, the effect of these molecules on polymer properties is not defined. In order to produce a long-term durable implant, the impact of large biomolecules on polymer structures needs to be explored.

Laminin, a basement membrane glycoprotein is known to provide a permissive substrate that binds to cell surface receptors and also can function to stimulate neurite extension [97, 98]. Laminin is composed of three polypeptide chains: A (≈400kDa), B1 (≈230kDa) and B2(≈220kDa) [99]. However, the use of the full multidomain protein or even a single domain bears some disadvantages. Proteins must be isolated from organisms and purified before use and consequently are a risk of infection and detrimental immune responses [100]. They are also subject to proteolytic degradation and thus may require refreshing

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to maintain cell attachment. This latter problem prevents them from being a useful tool when used for a permanent attachment in a long-term implant. Additionally, only specific sections of the laminin molecule contain the required receptors for cell adherence. Studies into the effects of proteolytically derived fragments of laminin have shown that partial functions of the large laminin molecule can be imitated by smaller fragmental components with specific functional binding [101].

Investigations by Huber et al. have found that synthetically prepared peptides have a greater stability than the native laminin protein molecule [95]. Huber and co-workers studied the effects of several laminin fragments including PDSGR, 18-mer SIKVAV, YFQRYLI and YIGSR. Results showed that synthetic peptides of both 18-mer SIKVAV and CDPGYIGSR mediate cell attachment and can enhance neurite extension in various animal models [11,35,95]. The laminin fragment CDPGYIGSR, has been used extensively by Cui et al. [35] and has been implanted on a coated probe into the guinea pig cortex. These positive results show that cell interaction of electrodes can be improved through use of adhesion peptides.

Many peptides of laminin can be synthetically manufactured and tailored for incorporation into conducting polymers. The peptide sequence is determined from the desired cell response. Table 5.1 shows a variety of laminin peptides reported as having specific cell response ligands. The addition of various amino acids to the end of a linear peptide chain can be used to control the overall ionic behaviour of the molecule. This ability to produce tailored laminin peptides as anions makes them ideal candidates for doping

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common conducting polymers. It is proposed that doping polymers with laminin peptides will enhance the biofunctionality of a neuroprosthetic implant. Table 5.1: Laminin peptide domains and their associated bioactivity Laminin Sequence Chain Domain Reported Bioactivity RGD Various Cell Adhesion [102] PDSGR β1 III Cell Adhesion [103] YIGSR β1 III Cell Adhesion [104] YFQRYLI α1 II Cell Adhesion and Neurite Outgrowth [105] SIKVAV α1 I Cell Adhesion and Neurite Outgrowth [106] RNIAEIIKDA γ1 I Neurite Outgrowth [107]

This study focuses on the use of laminin peptides to promote cell attachment to conducting polymers. Laminin peptides were chosen due to their ability to be tailored for both cell interactions and overall charge characteristics. Two peptide sequences derived from laminin were synthetically manufactured with the addition of amino acids to produce anionic peptides. CDPGYIGSR was chosen due to its positive results in cell attachment when previously explored by Cui et al [35]. In studies by Cui et al. synthetically manufactured DCDPGYIGSR was incorporated into PEDOT derivative PEDOT-MeOH films as a dopant, and showed good cell attachment properties, but was not characterised for electrochemical stability, mechanical hardness and adhesion to the electrode surface. An alternate anionically modified synthetic peptide DEDEDYFQRYLI, including the active sequence YFQRYLI will also be assessed in this chapter. The YFQRYLI ligand listed in Table 5.1 was identified by Tashiro et al. and reported to mediate cell attachment and promote neurite outgrowth in both PC12 cells and cerebellar microexplant cultures [105]. Tashiro et al. suggest that YFQRYLI is one of the active sites in laminin that regulate cell behaviour including neurite outgrowth, cell attachment and heparin binding. In this


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study it is hypothesised that the anionically modified synthetic peptide DEDEDYFQRYLI can be used as a dopant for PEDOT and providing increased cell adhesion and neurite outgrowth properties.


1. Produce conducting polymer coatings of PEDOT doped with synthetic anionic laminin peptides, DCDPGYIGSR and DEDEDYFQRYLI on model Pt electrodes,

2. Assess the effect of a large biomolecule dopant on established conducting polymer physico-chemical properties, and

3. Examine the effect of peptide doped polymers on mammalian cell interactions using a neural-like cell line. Additionally, the cell response to peptides containing specific bioactive ligands will be assessed.

5.2 5.2.1

Materials and Methods Electropolymerisation

Monomer and peptide dopoant solutions were made up in a polypropylene vial as indicated in Table 5.2. Due to the high cost of custom prepared synthetic peptides, initial studies into electrochemical stability were used to determine the concentration at which peptides were incorporated into the electrolyte solution, see Appendix E for CV data on 1 mg/mL compared to 5 mg/mL.


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Table 5.2: PEDOT/peptide electrolyte solutions Cat # Dopant type for PEDOT pTS DCDPGYIGSR DEDEDYFQRYLI EDOT 483028 0.1M 0.1M 0.1M pTS 402885 0.05M DCDPGYIGSR 12543000 5 mg/mL DEDEDYFQRYLI 1254300 5 mg/mL Deionised (DI) Water N/A 50% 50% 50% Acetonitrile 271004 50% 50% 50% Component

Custom peptides were obtained from Invitrogen. All other products were obtained from SigmaAldrich. Aliquots of 150μL of polymer solution were placed in each well with a Pt counter electrode. The working electrode was formed by the Pt substrate and electrically connected by an alligator clip. Polymers were electrodeposited using the inhouse manufactured galvanostat (detailed in Appendix B) at 1mA/cm2 or 300μA per well for 10 minutes. Following deposition the electrolyte was removed. Each film was washed with DI water four times. The final aliquot was left on the film for 24 hours in a 37 ◦ C incubator to leach any excess monomer, dopant or process contaminants. PEDOT/pTS was fabricated simultaneously and used as a control for all analyses.

5.2.2

Analytical Techniques

Physico-chemical analyses were performed according to methods detailed in Section 3.2, with the following adjustments:

• SEM and XPS studies were carried out on a small foil of Pt ≈ 25 x 25 mm. It was found during electrodeposition that samples containing biomolecules deposited


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poorly on ITO, but well on Pt. This piece of Pt was small enough to be placed in both analytical devices and enabled 4 samples to be assessed during each run. The Pt was cleaned between sample depositions using 1200 Grit sandpaper to remove prior samples, then washed in weak nitric acid and DI water.

• Conductivity studies using the four-point probe were not conducted. As stated above, films comprising biomolecules deposited poorly on ITO slides and were subsequently unable to be removed from the ITO substrate without significant damage. Films deposited and adhered well to Pt and were also unable to be removed from this substrate. Conductivity measurements performed on the Pt-polymer composite were biased by the underlying Pt layer. The four-point probe, designed for semiconductors was unable to operate with the Pt in place as the low voltage recordings were beyond its range.

5.2.3

In Vitro Cell Studies

Cell cultures were used to determine the activity of laminin peptide sequences YIGSR and YFQRYLI, used to dope PEDOT. The silicone well gaskets were washed as detailed in section 3.2.3 and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. Wells and Pt slides were placed in a beaker of DI water, covered and autoclaved for 20 min at 121◦ C. All components were air dried in a clean laminar flow hood. The wells were attached to the Pt slides and the system was then clamped in the custom-built sandwich assembly disscussed in 3.2.1.1 to prevent leaking between

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adjacent well chambers.

Following deposition in the presence of peptide doped electrolyte, the films were washed carefully four times with DI water and placed in the incubator overnight covered with the last 150 μL of water to leach water soluble contaminants. Sterilisation was performed by placing the wells without any overlying fluids under UV for 2 hr.

As previously discussed in section 3.2.1.2 the PC12 is a non-adherent cell line which requires laminin or collagen to grow on most culture surfaces. For biological characterisation of laminin doped polymers two well plates containing three repeats of each polymer type were fabricated. The wells on one plate were coated using the adsorption method with 5 μg/mL whole laminin sourced from Engelbreth murine sarcoma (Sigma Aldrich, L2020) in sterile DPBS. The wells in the other plate had 150 μL of DPBS only placed over the polymers. The latter plate was used to assess the efficacy of the laminin peptides on the non-adherent cell line. The laminin coated plate was used as a control for the uncoated plate and was used to provide information on whether the peptides support additional cell growth when compared to similarly prepared PEDOT/pTS controls. Wells were incubated for 12 hr at 37◦ C then washed with DPBS prior to cell plating as detailed in Section 3.2.6. Cell image and statistical analyses were carried out as for previous studies outlined in Sections 3.2.6.1 and 3.2.6.2.


5.3

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All peptide results presented in this section were produced from either DEDEDYFQRYLI or DCDPGYIGSR. Both the full peptide structure and the active ligand, YFQRYLI and YIGSR, are used to refer to the polymer type.

Polymers doped with laminin peptides were produced galvanostatically on Pt foils. These depositions differed to prior studies as the electrolyte solution from which the polymer was evolved turned the same blue colour as the polymerised film, indicating that the polymer did not only form on the working electrode, but was dispersed throughout the electrolyte. The initial wash cycle removed excess material from the surface revealing the formation of a confluent polymer film. Subsequent washes showed evidence that polymer particulate was loosely adhered to the surface and were disrupted by the wash. The final wash cycle was free of blue particulate indicating that all loose matter had been removed.

SEMs were recorded at 1,000 x and 15,000 x magnification as for prior studies. The most noteable feature of DEDEDYFQRYLI doped PEDOT was the presence of two regions in the polymer morphology. Figure 5.2(ii) shows the presence of nodular outcrops emerging from a predominantly featureless surface. A very different morphology is depicted in Figure 5.2(iii) for DCDPGYIGSR doped PEDOT, with large globular features visible at 1,000 x magnification but a fairly uniform topography at 15,000 x magnification. Neither polymer has the degree of nodularity or uniformity of the PEDOT/pTS control.

XPS analysis confirmed the presence of the peptides within the polymer layer. Figure

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Figure 5.2: SEMs at 1000x (left) and 15000x (right) magnification: i. PEDOT/pTS; ii. PEDOT/DEDEDYFQRYLI; iii. PEDOT/DCDPGYIGSR.


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5.3 is an example of the three spectra which were used to analyse the elemental makeup of each polymer. The average of three scans from three individually prepared films was used to calculate the relative presence of the peptide in each polymer compared to the monomer presence. In previous studies this was referred to as the doping ratio, however, when doped with peptides, the number of available anions used in the polymerisation is variable and dependent on a number of factors. Effectively DEDEDYFQRYLI should have five available electrons and DCDPGYIGSR should have a single anionic charge available. However, this does not take into account the molecule polarity, folding or pH of the electrolyte. As such XPS was used to quantify the presence of peptide at the surface, but not was defined as the doping ratio since it is not possible to tell how many monomer units were charge balanced per dopant molecule.

Since the number of nitrogen (N) atoms can be calculated for each peptide from their molecular composition the XPS spectra was used to calculate the degree of peptide incorporation in the polymer relative to the EDOT monomer. The percentage of N from the peptides was compared to the S from PEDOT and is given in Table 5.3. The data obtained from XPS was not significantly different for any of the polymer combinations including the PEDOT/pTS control. The error indicates that these results are repeatable and that the peptides were incorporated into the PEDOT at a similar monomer to dopant molecule ratio as the pTS. This is somewhat unexpected as DEDEDYFQRYLI should have more available electrons than DCDPGYIGSR, however it is not unlikely that molecular size and structure has a significant role in the formation of these polymers.

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Figure 5.3: X-ray photoelectron spectra of i. PEDOT/pTS; ii. PEDOT/DEDEDYFQRYLI; iii.PEDOT/DCDPGYIGSR. Incorporation of peptides is determined from the XPS signal, the number of nitrogen (N) atoms in each laminin peptide is compared to the relative constitution of S from the EDOT monomer.


Table 5.3: XPS determination of peptide incorporation into PEDOT films pTS, (n=3). Polymer Monomer: Dopant Ratio PEDOT/pTS 0.23 +/PEDOT/DEDEDYFQRYLI 0.19 +/PEDOT/DCDPGYIGSR 0.20 +/-

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compared to SD 0.06 0.03 0.02


Cyclic voltammetry curves had the same shape as reported in Section 3.3.1 for PEDOT films, with oxidation and reduction peaks occurring at the same monomer defined potentials. Electroactivity shown in Figure 5.4 degraded steadily across the first 100 cycles and tended to plateau in the subsequent 300 cycles for PEDOT/pTS and PEDOT/CDCPGYIGSR. PEDOT/DEDEDYFQRYLI did not experience the same plateau and although the overall loss of activity is similar to that experienced by PEDOT/DCDPGYIGSR it is forseeable that electroactivity would degrade further if subjected to more cycles.

While the overall loss of electrical activity was not greatly reduced by the incorporation of laminin peptides, Table 5.4 also contains data on the polymer charge storage densities. Both peptide doped polymers have approximately one quarter of the capacity for charge storage when compared to the PEDOT/pTS control. Table 5.4: Comparison of electroactivity loss over 400 cycles of CV and final charge storage density for PEDOT doped with laminin peptides, (n=3). Electroactivity Final Charge Polymer Loss SE Density (% of original) (%) (mC/cm2 ) PEDOT/pTS 11.99 +/- 3.55 198.02 PEDOT/DEDEDYFQRYLI 18.77 +/- 4.28 53.31 PEDOT/DCDPGYIGSR 16.84 +/- 5.63 50.11

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Figure 5.4: Electrochemical activity loss for PEDOT doped with laminin peptides compared to conventional pTS dopant, over 400 cycles of redox. Loss of electroactivity is plotted as a percentage of the original actvity measured at Cycle 1, normalising the polymers to a common baseline, (n=3).

Impedance spectroscopy results, plotted in Figure 5.5 demonstrated that laminin doped PEDOT performed similarly to other polymers by reducing the impedance magnitude and phase lag seen at the electrode surface compared to bare Pt. Laminin doped PEDOT had very low magnitudes at the starting frequency of 10 Hz compared to other materials, however the impedance magnitude of all polymers converged from 500 Hz onwards to plateau around 130 Ω. PEDOT/pTS was not significantly different from either PEDOT/DEDEDYFQRYLI or PEDOT/DCDPGYIGSR at 1000 Hz, with all polymers reducing the impedance magnitude recorded from Pt by ≈ 80%. The phase lag was also significantly reduced across all substrates in the low frequency range with lag recorded at


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10 Hz for PEDOT/YFQRYLI being 12◦ or 17% of the lag measured from Pt.

Laminin peptide doped polymers had lower impedance magnitudes across the low frequency range of 10 - 100 Hz, but converged with PEDOT/pTS across the biological focal region of 300 - 1000 Hz. The phase lags at 1kHz, recorded in Table 5.5, were not significantly different for any polymer, but the impedance magnitude was still significantly better than for bare Pt. Table 5.5: Impedance of PEDOT doped with laminin peptides compared to chemical dopant pTS and bare Pt electrodes at 1 kHz, (n=3). Electrode Coating Magnitude Phase Bare Pt 629.8 -8.6 PEDOT/pTS 138.7 -3.7 PEDOT/YFQRYLI 126.5 -5.8 PEDOT/YIGSR 142.7 -4.3

5.3.2


Polymer hardness results graphed in Figure 5.6 demonstrated that peptide doping strongly affected the polymer hardness. DEDEDYFQRYLI doped PEDOT was the softest across both gouge and scratch hardness testing at 5 units softer than PEDOT/pTS for both measurements. PEDOT/DCDPGYIGSR was 1 unit harder than PEDOT/DEDEDYFQRYLI for both hardness measurements, however it was still 4 units softer than PEDOT/pTS. Error bars suggest that these results are repeatable, with little difference seen across the 6 samples.

As for prior analytical techniques, the peptide doped polymers shown in Figure 5.7 performed without significant difference. The dopant size appears to play a significant

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Figure 5.5: Bode plot of impedance spectroscopy for laminin peptide doped PEDOT electrode coatings. Impedance magnitude is plotted in the upper graph, with phase lag plotted below. Error bars are marked at one standard deviation, (n=3).


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Figure 5.6: Hardness of laminin doped PEDOT compared to pTS doped PEDOT as measured by the ASTM hardness test for coatings. Gouge hardness is representative of a minor disruption to the polymer surface and scratch hardness depicts the level at which no polymer is removed from the Pt substrate. Error bars represent SE, (n=6).

role in mechanical performance with the large DEDEDYFQRYLI (12-mer) and DCDPGYIGSR (10-mer) dopants having a significant impact on polymer adherence. Both PEDOT/DEDEDYFQRYLI and PEDOT/DCDPGYIGSR lost approximately 10% of their coating, compared to a negligible loss for PEDOT/pTS. As for previous studies the loss of coating from the Pt substrate shown by samples images in Figure 5.8 occurred across the entire film and was not concentrated at the x-cut disruption of the film.

5.3.3


Cell studies were performed to determine both the effect of the peptide dopants compared to conventional chemical dopants, and to determine the efficacy of the ligands contained

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Figure 5.7: Comparison of film delamination by ASTM x-cut analysis for laminin peptide doped PEDOT and pTS doped control. Error is represented as SE, (n=6).

Figure 5.8: Sample of x-cut assay images following application of low-adhesion tape to disrupted portion of the laminin peptide doped PEDOT: i.PEDOT/pTS (control); ii.PEDOT/DEDEDYFQRYLI; iii.PEDOT/DCDPGYIGSR.


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within the dopant. For this reason both uncoated and laminin coated polymers were used in the neurite outgrowth assays.

At 96 hr fluorescent micrographs were taken for cell analysis. Sample micrographs that were considered typical of the cell response on each substrate type are shown in Figure 5.9. The images indicate that despite the use of peptide dopants, the laminin coated version of each film is superior. The analysis of cell images are presented in Figure 5.10 and Figure 5.11 quantifying cell density and neurite length supported per cm2 , respectively.

The cell density data reveals an expected result in PEDOT/pTS, where the cell density with laminin coating is not significantly better than that seen for bare PEDOT/pTS. Additionally this density is not significantly worse than that seen for the laminin coated PEDOT/DEDEDYFQRYLI and is superior to all other uncoated substrates.

However, this performance is not replicated in the neurite length data. All substrates recorded significantly worse neurite lengths when the laminin coating was not placed over the polymer prior to cell plating. There was NSD seen in neurite length between the different polymer films when bare substrates were used. The addition of the laminin coating is clearly of benefit for all polymers, but most notably both peptide doped polymers were able to support a 400% increase in neurite length per unit area when coated with laminin. Compared to laminin coated PEDOT/pTS, the laminin coated and peptide doped polymers were able to support a 70% and 92% increase in neurite length for DEDEDYFQRYLI and DCDPGYIGSR doped PEDOT respectively.

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Figure 5.9: Sample images of PC12 neurite outgrowth on laminin doped polymers at 96 hr post-plating with bare polymer (left) and laminin coated (right): i. PEDOT/pTS; ii. PEDOT/DEDEDYFQRYLI; iii. PEDOT/DCDPGYIGSR.


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Figure 5.10: PC12 cell density on laminin peptide doped PEDOT at 96hrs post-plating. Polymers are shown both bare and with an additional whole laminin coating. Pt and PEDOT/pTS are used as controls. Error is represented as SE, (n=3); (∗ NSD).

Figure 5.11: Neurite outgrowth of PC12s on laminin peptide doped PEDOT at 96hrs post-plating. Neurite outgrowth is represented as the total neurite length supported per cm2 . SE is given, (n=3); (∗∗ p < 0.05).


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5.4

Summary of Results

Table 5.6 summarises the properties of each polymer type to aid in the comparison of materials. Where quantitative values are important, they are given. For other properties a qualitative interpretation of a range of values is used to indicate the polymer’s overall performance. Table 5.6: Summary of results for laminin doped PEDOT Analysis PEDOT/pTS PEDOT/YFQRYLI PEDOT/YIGSR Surface Morphology Very rough Smooth + rough phases Globular features Efficiency of Deposition Good Satisfactory Good Electrochemical Stability Good Satisfactory Satisfactory Impedance (Ω at 1kHz) 138.7 126.5 142.7 Hardness (Scratch) 3H 2B B Adherence (% Loss) 0.52 10.39 9.78 Neurite Outgrowth a Good Excellent Excellent Bioactivity b N/A c Limited Limited a

Laminin coated polymer With no laminin coating c Although this material was not intended to have bioactivity the surface topography promoted good cell attachment b

5.5

Discussion

PEDOT was doped with laminin peptides DEDEDYFQRYLI and DCDPGYIGSR through electropolymerisation using identical deposition parameters as for previous experiments. The resulting film was less uniform in macroscopic appearance to previous polymer compositions and physico-chemical properties were affected by the larger dopant. The hypothesised cell response was not clearly defined as increased interaction was only observed

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127

when the polymers were coated with the whole laminin molecule prior to cell plating.

Doping PEDOT with laminin peptides resulted in a difference in the formation of the polymer during electrodeposition. It was noted that the electrolyte solution was coloured a deep navy due to the presence of small polymerised particles of peptide doped PEDOT that were unable to be incorporated into the polymer film at the electrode. A description of PPy doped with HA, given by Collier et al. may provide insight into the mechanism responsible for this phenomenon. Collier indicated that the large size of the HA inhibited its incorporation as a dopant ion during PPy film synthesis, quantified by the time required to deposit a film of equal thickness to a PPy/PSS control [37]. Diffusion limitations in the HA solution were believed to give rise to inhomogeneous growth of the PPy/HA, resulting in films that had a less uniform appearance with nodular projections and were more brittle than PPy/PSS [37].

The observations made by Collier et al. for PPy/HA are mirrored in this study with both PEDOT/DEDEDYFQRYLI and PEDOT/DCDPGYIGSR having a less uniform appearance when viewed under SEM than the PEDOT/pTS control. Additionally, the larger DEDEDYFQRYLI peptide had a lower ratio of incorporation into the PEDOT, determined by XPS, than expected. The five available electrons on the anionic DEDED-tail suggest this peptide should be able to balance five times more EDOT units than DCDPGYIGSR, which only has one available electron. The results indicate that the larger DEDEDYFQRYLI molecule experienced significant diffusion limitations during electrodeposition, significantly reducing the efficiency of deposition. This may also be a contributing

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factor to the relatively featureless topography seen in the PEDOT/DEDEDYFQRYLI SEM. For the electrons available per molecule, the DCDPGYIGSR was incorporated at the same rate as the pTS, indicating this polymer was formed more easily and with possibly the same efficiency as the PEDOT/pTS control.

Electrical properties of peptide doped PEDOT were found to be consistent with those recorded for other large dopants such as PEDOT/PSS, characterised in Chapter 3. CV revealed a higher loss of electrochemical stability for both PEDOT/DEDEDYFQRYLI (19%) and PEDOT/DCDPGYIGSR (17%) compared to PEDOT doped with the small pTS anion (12%).

The manufacturer of the synthetic peptides reported that DED-

EDYFQRYLI had a molecular weight of 1.6 kDa and DCDPGYIGSR had a molecular weight of 1 kDa. When conducting polymers are reduced, small doping anions such as pTS migrate out of the film, upon oxidation they return to maintain the balance [108] as shown in Figure 5.12(i) for a PPy film. In contrast, large anionic dopants cannot leave the film as rapidly as required. In this instance the cations of the supporting electrolyte will move into the film to neutralise the negative charge of the dopant anions [108], represented in Figure 5.12(ii). For many systems, a mixture of both processes occur depending on the structure of the dopant, film thickness and the rate of redox induced [109]. In this study it is thought that during reduction the peptide dopants towards the surface of the electrolyte are able to migrate out of the polymer, but are quickly displaced on oxidation as they are unable to return due to lack of mobility. However, the bulk of the electroactivity is preserved as the majority of peptide dopants are restrained by the PEDOT matrix

5.5. DISCUSSION

129

and reduction occurs primarily through cation migration into the polymer (most likely Na+ in biological environments).

Figure 5.12: Oxidation and reduction behaviour of conducting polymers is dependent on dopant size. Small anion A− is able to migrate in and out of the PPy matrix to balance the charge across the backbone. Large anion B− is immobilised within the PPy matrix and hence relies on smaller cation X+ to migrate into the polymer from the surrounding electrolyte and maintain charge balance across the polymer.

Impedance of peptide doped films is not significantly different to that recorded from PEDOT/pTS, except in the low frequency range of 10 - 100 Hz. This trend concurs with studies by Stauffer and Cui, who doped PPy with peptides DCDPGYIGSR and RNIAEIIKDI and demonstrated that impedance was only significantly lower than that of the gold control in the 1 - 500 Hz range [110]. Earlier work by Cui et al. also recognised that PPy doped with an anionic modification of silk-like protein fragments (SLPF) had an impedance even higher than bare gold above 10 kHz, but in the lower frequency region from 10 to 100 Hz impedance was significantly lower than both gold and PPy/PSS [35]. Similar results are seen here for PEDOT with both peptide dopants performing with

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significantly lower impedance magnitudes and phase lags over the 10 - 100 Hz region, but converging with PEDOT/pTS above 500 Hz. The phase plot of impedance spectroscopy showed that the phase lag at 10 Hz was around 75◦ for Pt, which means the electrode was close to a pure capacitor. After coating with PEDOT the phase decreased to ≈ 30◦ for PEDOT/pTS and was further reduced to ≈ 20◦ for PEDOT/DEDEDYFQRYLI and ≈ 12◦ for PEDOT/DCDPGYIGSR. The reduction of phase lag is an indication of a more resistive and hence more desirable electrode for neuroprosthetic applications.

The dichotomy of mechanical performance is emphasised in this study with the benefit of softer polymers being outweighed by the lack of adherence and durability. Softer but less adherent films produced by peptide doping confirm that mechanical performance is primarily determined by dopant size. PEDOT/DEDEDYFQRYLI was both softer and less adherent than PEDOT/DCDPGYIGSR, confirming that the increase in molecular weight of the dopant had a noticeable effect on polymer properties. Collier et al. experienced similar problems when doping PPy with HA, describing the material produced as “brittle and difficult to handle” [37]. It is thought that the limitation of diffusion during electropolymerisation may contribute to the resulting mechanical properties.

Cell data obtained by in vitro assays using the PC12 clonal cell line revealed that cell adherence and neurite outgrowth are influenced by different mechanisms. When the laminin coating used in previous studies was not applied, the cell density on all substrates dropped dramatically with the exception of PEDOT/pTS. There was NSD between bare and laminin coated PEDOT/pTS (p > 0.05). However, it is noted that the higher density

5.5. DISCUSSION

131

of cells on PEDOT/pTS did not translate to a high outgrowth of neurites. NSD was observed in neurite length per cm2 between the different PEDOT substrates when no laminin coating was applied. There are a number of factors that might affect the performance of the peptide doped films in vitro, including peptide accessibility, concentration and possible degradation of ligands during electrodeposition. Coating the substrates with whole laminin chains provided insight into these facets of cell adherence and neurite outgrowth.

Surface analysis by XPS, used extensively in research to detect proteins [111–113], indicated that both peptides were present at the polymer surface and hence should be accessible to cells. While there was no evidence of PC12 cells accessing active peptides when plated directly on peptide doped PEDOT, there was a significant increase in both cell density and neurite outgrowth when these films were coated with an additional layer of whole laminin. The laminin coated, peptide doped PEDOT films had significantly longer neurite outgrowth supported per cm2 than both their uncoated counterparts and the laminin coated PEDOT/pTS (p < 0.05). The results suggest that peptides are accessible and active, but might not be present at a high enough concentration to promote attachment of the non-adherent PC12 cell line.

An alternate more adherent cell line may provide greater insight into the cell attachment properties of peptide dopants in conducting polymers. The addition of dopant peptides at a higher concentration is not recommended as the inhibition to polymer formation would be increased. It is also propounded that incorporation of peptides as dopants by electrodeposition is inefficient, requiring much higher concentrations of the costly synthetic anions

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in the electrolyte solution than can be incorporated into the polymer. Post-processing application of cell matrix proteins might provide a more efficient method of creating a conducting polymer with cell attachment properties.

Another polymer characteristic that requires consideration when evaluating the biological performance of peptide doped PEDOT is the film topography. By viewing the cell outgrowth results in conjunction with SEM imaging it was noted that PEDOT/pTS, the film with the highest cell density when not coated with whole laminin, provided a rough but uniform substrate which is a known factor in promoting cell adherence [12, 35]. The less uniform and relatively featureless substrate provided by the peptide doped PEDOT most likely contributed to the poor cell attachment properties recorded for these materials. Once laminin coated, all polymer substrates had sufficient cell adherence with NSD seen in the cell densities observed. It is therefore concluded that the peptides are present at the polymer surface, accessible and active but the film topography is unsuitable.

Despite the hypothesised dual bioactivity of the YFQRYLI ligand, PEDOT/DCGPGYIGSR recorded slightly higher neurite lengths than PEDOT/DEDEDYFQRYLI. This could indicate that the YFQRYLI ligand had no impact on neurite outgrowth bioactivity and only had adherence bioactivity, but may also indicate that the YIGSR ligand had a higher affinity for cell membrane attachment and provided support for neurite growth through cell attachment mechanisms. Although the exact mechanism through which YFQRYLI influenced neurite growth is not known, it can be concluded that both YFQRYLI and YIGSR were able to increase neurite outgrowth following cell adherence to the polymer

5.6. CONCLUSIONS

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substrate.

5.6

Conclusions

The incorporation of peptides into PEDOT has yielded varied results. The use of large synthetic peptides as anionic dopants inhibits the formation of the polymer during electrodeposition, resulting in inferior electrical properties to PEDOT doped with the significantly smaller pTS molecule. While the material produced is softer, it is also less durable. Cell studies revealed that without an additional laminin coating the cells were not able to access the benefits of the laminin peptides to promote cell adherence. However, peptide doped PEDOT did have an influence over neurite outgrowth with laminin coated films supporting significantly more neurite outgrowth than PEDOT/pTS and similar neurite outgrowth when no laminin was present despite having much lower cell numbers. The type of peptide used had little effect on cell interactions but the larger DEDEDYFQRYLI molecule was shown to interfere significantly with film formation and consequently this polymer had lower mechanical adherence than the DCDPGYIGSR doped PEDOT.

The use of a peptide dopant does retain the benefit of having a more natural material to interface neural tissue at an electrode surface. However, it is hypothesised that a much smaller peptide is required for optimal performance by reducing the impact of a bulky dopant on physico-chemical properties. Alternately, a different mechanism of peptide inclusion to conducting polymers, such as covalent tethering, could be used to present higher concentrations of peptides at the surface for improved cell interactions.

Chapter 6

Conducting Polymers with Neurotrophic Bioactivity 6.1

Introduction

The interface between conventional metal electrodes and the tissue they stimulate is commonly associated with a significant fluid gap through which the electrical signal must pass. The distance from an electrode to neural cells is critical in determining the longterm performance of a device including current threshold required for cell activation and the perceived signal resolution [10]. In the previous chapter it was shown that laminin peptides could be used to dope PEDOT creating a conducting polymer with cell attachment properties. An electrode material produced from conducting polymers loaded with neurotrophins has the potential to promote neural cell outgrowth from the tissue to the electrodes. The optimum design for a neural interface will have both cell attachment and neurite outgrowth properties, promoting the formation of the intimate connection depicted in Fig. 6.1, minimising the loss of signal to extracellular fluid and creating a 134

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long-term, robust interface.

Figure 6.1: Schematic of a desirable bioactive conducting polymer electrode array

Neurotrophins are a structurally related family of basic proteins of ≈13kDa, that have been shown to promote neuronal growth and survival [114]. Members of this family include nerve growth factor (NGF), brain derived neurotrophic factor (BDNF), glial cell line derived neurotrophic factor (GDNF), and neurotrophins 3 and 4/5 (NT-3 and NT4/5 respectively) [15]. Neurite extension is known to follow a gradient of growth factors [115–117].

The structural hallmark of all the neurotrophins is the characteristic arrangement of the disulfide bridges known as the cysteine knot, which has been found in other growth factors such as platelet derived growth factor (PDGF) [118]. There is a 95.8% homology between the rat and mouse forms, and a 85% homology between the human and mouse [118]. This study utilises NGF because of its reported ability to promote neurite outgrowth in the PC12 cell line [75] and prior evidence of its compatibility with conducting polymers in neuroprothesis applications [69]. Additionally, Carmignoto et al. have reported NGF

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CHAPTER 6. NEUROTROPHIC BIOACTIVITY

promoted survival of rat retinal ganglion cells following optic nerve axotomy [119].

Like other peptides and proteins, neurotrophins can be incorporated into polymers in a variety of ways including as a dopant, covalent tethering, adsorption and entrapment. However, these molecules are intended to be released where they can be taken into or placed in close proximity to the target cell and promote neurite outgrowth [120]. To provide cells access to the neurotrophins they are most commonly entrapped within the polymer matrix during electrodeposition and released during stimulation regimes consistent with the neuroprosthetic application [16, 64, 121].

Neurotrophins have been incorporated into polymers to tailor the application for specific tissues relating to the intended neuroprosthetic application. Studies conducted by the Bionic Ear Institute, Australia have shown that it is possible to resprout spiral ganglion cells with the release of neurotrophins and create axon growth towards the stimulating electrodes of the cochlear implant [122]. In vitro investigations were carried out on cultured spiral ganglion cells to determine the most appropriate neurotrophin. Neurotrophin 3 (NT-3) resulted in the greatest axon extension when released from a polypyrrole (PPy) electrode coating by cyclic voltammetry [122]. Acute in vivo testing of the NT-3 showed that it was possible to grow the spiral ganglion cells although no directional growth towards the stimulating implant was observed. Furthermore, the neurotrophin release was shown to prevent cell death and further regression of the remaining disease affected neurons during the course of NT-3 release [122]. Longer-term studies into NT-3 release have shown that once the neurotrophin source is depleted electrically stimulated cells can be

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maintained. Cells outside of the electrical field however will regress to the diseased state.

No research to date explores the affect of neurotrophins on conducting polymer physicochemical properties. Many studies have reported a presence of electrochemical activity directly following fabrication, but long-term electroactivity has not been reported. Additionally, the combination of both a laminin peptide for cell adherence and a neurotrophin for neurite extension has not been explored.

This study focuses on the incorporation of the neurotrophin NGF into laminin peptide doped PEDOT. The two anionically modified peptide sequences DEDEDYFQRYLI and DCDPGYIGSR characterised in Chapter 5 will be used in this study. Additionally, since PEDOT/pTS was found to have cell adherence properties it was also modified with the NGF molecule.


1. To entrap neurotrophin NGF in synthetic laminin peptide doped PEDOT and pTS doped PEDOT during electropolymerisation,

2. Assess the effect of incorporating two large biomolecular components on established conducting polymer physico-chemical properties, and

3. Examine the effect of incorporating neurotrophins into conducting polymer matrices on mammalian cell interactions using a neural-like cell line. Specifically, the bioactivity of the entrapped NGF will be assessed through neurite outgrowth.


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6.2 6.2.1

Materials and Methods Electropolymerisation

Monomer and peptide dopant solutions with neurotrophin additions were made up in a polypropylene vial as indicated in Table 6.1. Table 6.1: PEDOT/peptide/NGF electrolyte solutions Component Cat # Dopant type for PEDOT pTS DCDPGYIGSR DEDEDYFQRYLI EDOT 483028 0.1M 0.1M 0.1M pTS 402885 0.05M DCDPGYIGSR 12543000 5 mg/mL DEDEDYFQRYLI 1254300 5 mg/mL NGF 2.5S Grade I N240 1 μg/mL 1 μg/mL 1 μg/mL Deionised (DI) Water N/A 50% 50% 50% Acetonitrile 271004 50% 50% 50%

Custom peptides were obtained from Invitrogen and NGF sourced from mouse submaxillary glands was purchased from Alamone laboratories. All other products were obtained from SigmaAldrich. Aliquots of 150μL of polymer solution were placed in each well with a Pt counter electrode. The working electrode was formed by the Pt substrate and electrically connected by an alligator clip. Polymers were electrodeposited using the in-house manufactured galvanostat (detailed in Appendix ) at 1 mA/cm2 or 300μA per well for 10 minutes. Following deposition the electrolyte was carefully removed and each film was washed with 150 μL DI water four times. The final aliquot was left on the film for 24 hours in a 37 ◦ C incubator to leach any excess monomer, dopant or process contaminants. PEDOT films without NGF were fabricated for use as controls for analytical techniques.


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Analytical Techniques

Physico-chemical analyses were performed according to methods detailed in Section 5.2.2.

6.2.3

In Vitro Cell Studies

Cell assays were designed to determine the ability of NGF incorporated within the polymer matrix to diffuse and be recognised by the neural clone PC12, resulting in neurite extension. The silicone well gaskets were washed as detailed in section 3.2.3 and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. Wells and Pt slides were placed in a beaker of DI water, covered and autoclaved for 20 min at 121◦ C. All components were air dried in a clean laminar flow hood. The wells were attached to the Pt slides and the system was then clamped in the custom-built sandwich assembly disscussed in 3.2.1.1 to prevent leaking between adjacent well chambers.

Following deposition in the presence of peptide doped, NGF supplemented electrolyte, the films were washed carefully four times with DI water and placed in the incubator overnight covered with the last 150 μL of water to leach water soluble contaminants. Control films without the NGF supplement in the electrolyte were also fabricated with cell culture performed on both films simultaneously under identical conditions. Sterilisation was performed by placing the wells without any overlying fluids under UV for 2 hr.

As previously discussed in section 3.2.1.2 the PC12 is a non-adherent cell line which requires laminin or collagen to grow on most culture surfaces. The polymers were coated

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with 5 μg/mL laminin sourced from Engelbreth murine sarcoma (Sigma Aldrich, L2020) in sterile DPBS. Wells were incubated for 4 hr at 37◦ C then washed with DPBS prior to cell plating as detailed in Section 3.2.6. A reduced laminin adsorption time was used to prevent loss of NGF through diffusion mechanisms. The materials produced with NGF in the electrolyte were cultured for 96 hr in low serum (1% HS) RPMI without additional NGF. The control films fabricated without NGF were cultured as for previous studies in low serum medium supplemented with 50 ng/mL NGF. Cell image and statistical analyses were carried out as for previous studies outlined in Sections 3.2.6.1 and 3.2.6.2.

6.3


All peptide results presented in this section were produced from either DEDEDYFQRYLI or DCDPGYIGSR. Both the full peptide structure and the active ligand, YFQRYLI and YIGSR, are used to refer to the synthetically produced anionic molecules.

NGF modified polymers were produced galvanostatically on Pt foils. The presence of particulate matter similar to that observed for PEDOT/peptide polymerisations was noted during the electrodeposition, colouring the usually transparent electrolyte a deep opaque navy. This excess of polymer was removed during the washing cycles, however it was observed that the polymers containing both peptide and NGF were noticeably more fragile than those containing only one biomolecule. Despite extreme care during the washing cycles, on occasion shear forces resulted in film delamination or a major disruption of the sample. In these instances the sample was discarded.


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SEMs revealed that incorporation of NGF altered the surface topography of all polymer combinations. However, each polymer type was still able to be identified by comparing it to the images taken in Chapter 5 when NGF was not incorporated into the electrolyte, see Figure 5.2. PEDOT/pTS/NGF had a porous appearance and at 15,000 x magnification the surface was fibrillar in structure with pores of approximately 2 μm. Both PEDOT/DEDEDYFQRYLI/NGF and PEDOT/DCDPGYIGSR/NGF had a more nodular surface topography than PEDOT/pTS/NGF. Smaller protrusions were seen across the PEDOT/DEDEDYFQRYLI/NGF film with each nodule having a split at the apex, giving the impression of a burst bubble. Globular outcrop features similar to that seen in PEDOT/DEDEDYFQRYLI were also observed and are detailed at 15,000 x magnification in Figure 6.2(ii). PEDOT/DCDPGYIGSR/NGF closely resembled the surface topography of PEDOT/DCDPGYIGSR. The NGF incorporated film was more nodular than the peptide only film, but the presence of large toroidal features once again gave the impression of a burst bubble.

XPS analysis confirmed the presence of proteins within the polymers. Figure 6.3 is an example of the three spectra which were used to analyse the elemental makeup of each polymer. In previous studies this was referred to as the doping ratio, however, when doped with peptides and NGF the presence of amino acids in both molecules makes it impossible to quantify the contribution of each component. The XPS scan of PEDOT/pTS/NGF, Figure 6.3(i), shows a nitrogen peak previously not encountered in PEDOT/pTS spectra, and is consistent with the incorporation of NGF into the polymer.

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Figure 6.2: SEMs at 1000x (left) and 15000x (right) magnification: i. PEDOT/pTS/NGF; ii. PEDOT/DEDEDYFQRYLI/NGF; iii. PEDOT/DCDPGYIGSR/NGF.


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Figure 6.3: X-ray photoelectron spectra of i. PEDOT/pTS/NGF; ii. PEDOT/DEDEDYFQRYLI/NGF; iii.PEDOT/DCDPGYIGSR/NGF. Incorporation of peptides and NGF was determined from the XPS signal by comparison to the spectra without NGF.

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When NGF is incorporated into peptide doped PEDOT, the spectra does not produce a peak that is unique to each elemental contribution. Both sulfur and nitrogen, the elements which have been used in previous studies to compare the relative contribution of the dopant and monomer units, are found in NGF. In this scenario sulfur, S which is present in EDOT is also in the amino acid methionine, found in NGF. Additionally, nitrogen, N is present in all amino acids which are the building blocks of all proteins and are consequently abundant in both laminin peptides and NGF. However, by looking at the percentage contribution of all elements (oxygen, O; nitrogen, N; carbon, C; sulfur, S) to each polymer, detailed in Table 6.2, it can be observed that the NGF entrapped films consistently have more O, N and S, coupled with less C. This change in the elemental ratios is used to confirm the presence of a higher proportion of amino acid based molecules, as expected in the NGF loaded polymers. Table 6.2: Elemental analysis of PEDOT loaded with NGF compared to polymers without NGF, (n=3). Elemental Constitution Polymers O N C S (%) (%) (%) (% ) PEDOT/pTS 26.0 0.0 65.0 8.0 PEDOT/pTS/NGF 31.0 5.0 53.0 10.0 PEDOT/DEDEDYFQRYLI 29.0 8.0 58.0 4.0 PEDOT/DEDEDYFQRYLI/NGF 37.5 9.0 47.5 5.0 PEDOT/DCDPGYIGSR 27.5 9.0 57.5 5.0 PEDOT/DCDPGYIGSR/NGF 32.0 10.5 50.0 6.5

6.3.1


Cyclic voltammetry curves had the same shape as reported in Section 3.3.1 for PEDOT films, with oxidation and reduction peaks occurring at the same monomer defined poten-


145

tials, but with much smaller peak-to-peak amplitude than pTS doped PEDOT, depicted in Figure 6.4. The graph of electroactivity loss across 400 cycles of CV, Figure 6.5, shows an increase in electroactivity across the first 10 - 20 cycles for all polymers. This is consistent with an initial period of rearrangement and is commonly seen in polymers with a more complex structure. The subsequent cycles show a continual degradation in electrochemical behaviour. PEDOT/pTS/NGF starts to plateau after 200 cycles and PEDOT/DEDEDYFQRYLI plateaus after 300 cycles, but no noticeable flattening of the curve is observed for PEDOT/DCDPGYIGSR/NGF. The tendency for curves to flatten later in the cycling and in some cases not at all, is indicative of reduced electrochemical stability. Error bars represent SE and indicate that PEDOT/DEDEDYFQRYLI had substantial variation in electroactivity over the 400 cycles.

The overall loss of electrical activity and the final charge storage densities recorded in Table 6.3 imply that the incorporation of two biomolecules into the PEDOT matrix greatly impacts on the polymer electrochemical properties. It was observed however that the use of only one biomolecule, especially in the case of PEDOT/pTS/NGF, results in a more stable polymer which retains a higher charge storage density than those with two biomolecules.

The Bode plot of impedance, Figure 6.6, demonstrates that unlike electrochemical stability, the incorporation of two biomolecules does not have a negative impact on the impedance at the electrode surface. An overall decrease in impedance magnitude and phase lag is seen when NGF loaded films are compared to PEDOT/pTS and bare Pt

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Figure 6.4: Sample CV curve obtained from PEDOT/DEDEDYFQRYLI/NGF compared to PEDOT/pTS, shows reduction in peak-to=peak amplitude resulting from the incorporation of two large biomolecules into the PEDOT matrix.

Table 6.3: Comparison of electroactivity loss over 400 storage density for NGF loaded PEDOT, (n=3). Electroactivity Polymer Loss (% of original) PEDOT/pTS 11.99 PEDOT/pTS/NGF 12.13 PEDOT/DEDEDYFQRYLI 18.77 PEDOT/DEDEDYFQRYLI/NGF 50.76 PEDOT/DCDPGYIGSR 16.84 PEDOT/DCDPGYIGSR/NGF 60.64

cycles of CV and final charge

SE (%) +/- 3.55 +/- 4.27 +/- 4.28 +/- 11.42 +/- 5.63 +/- 8.59

Final Charge Density (mC/cm2 ) 198.02 117.81 53.31 12.49 50.11 18.74


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Figure 6.5: Electrochemical activity loss for PEDOT loaded with NGF for three dopant types, over 400 cycles of redox. Loss of electroactivity is plotted as a percentage of the original actvity measured at Cycle 1, normalising the polymers to a common baseline. Error bars represent SE, (n=3).


148

in the low frequency region. The inclusion of NGF in the PEDOT/pTS film results in impedance behaviour that mimics that of the peptide doped polymers. The phase lag recorded from 10 - 5000 Hz is consistently lower for PEDOT/pTS/NGF than other peptide doped, NGF loaded films.

Impedance magnitudes and phases were shown to converged for all polymers across the biological focal region of 300 - 1000 Hz. The phase lags at 1 kHz, recorded in Table 6.4, were not significantly different for any polymer, but the impedance magnitude was still significantly better than for bare Pt. PEDOT/pTS/NGF performed best with the lowest impedance magnitude and phase lag. Table 6.4: Impedance of PEDOT loaded with NGF compared to neurotrophin free controls at 1 kHz, (n=3). Electrode Coating Magnitude Phase Bare Pt 629.8 -8.6 PEDOT/pTS 138.7 -3.7 PEDOT/pTS/NGF 120.7 -1.4 PEDOT/DEDEDYFQRYLI 126.5 -5.8 PEDOT/DEDEDYFQRYLI/NGF 136.0 -5.0 PEDOT/DCDPGYIGSR 142.7 -4.3 PEDOT/DCDPGYIGSR/NGF 146.3 -3.1

6.3.2


ASTM hardness testing was performed on all NGF loaded films with results graphed in Figure 6.7 against the peptide doped PEDOT without NGF modification. Comparisons demonstrate that the addition of the NGF can reduce the hardness of conducting polymers. The only polymer where no difference was observed was PEDOT/DEDEDYFQRYLI, which was unchanged in the soft region. Both gouge and scratch hardness were reduced by


149

Figure 6.6: Bode plot of impedance spectroscopy for NGF loaded PEDOT electrode coatings. Impedance magnitude is plotted in the upper graph, with phase lag plotted below. Error bars are marked at one standard deviation, (n=3).

150


1 unit for PEDOT/DCDPGYIGSR by the addition of NGF, making it equal in softness to PEDOT/DEDEDYFQRYLI/NGF. PEDOT/pTS showed the most difference when NGF loaded, with gouge hardness reduced by 3 units and scratch hardness reduced by 2 units. Although the resulting film hardness was still higher than that recorded for peptide doped PEDOT with no NGF modification.

Figure 6.7: Effect of NGF loading on ASTM hardness of PEDOT based polymers. Gouge hardness is representative of a minor disruption to the polymer surface and scratch hardness depicts the level at which no polymer is removed from the Pt substrate. Error bars represent SE, (n=6).

As previously recognised, the dopant size appears to play a significant role in polymer mechanical performance. The incorporation of additional non-doping molecules, such as the large (26 kDa) NGF molecule appear to further decreases film adherence (Figure 6.8). Both PEDOT/DEDEDYFQRYLI/NGF and PEDOT/DCDPGYIGSR/NGF had a loss of approximately 30% of their coating, compared to ≈10% for their unmodified (without


151

Figure 6.8: Comparison of film delamination by ASTM x-cut analysis for PEDOT loaded with NGF. Error is represented as SE, (n=6).

Figure 6.9: Sample images of polymers following application of lowadhesion tape to x-cut portion of NGF loaded PEDOT: i.PEDOT/pTS/NGF; ii.PEDOT/DEDEDYFQRYLI/NGF; iii.PEDOT/DCDPGYIGSR/NGF.


152

NGF) counterparts. Most noteably, the inclusion of NGF in PEDOT/pTS resulted in a loss of adherence that was comparable to that of peptide doped PEDOT. Samples images in Figure 6.9 show polymer was lost from across the entire film for all materials and was not concentrated at the x-cut disruption of the film.

6.3.3


In vitro cultures of PC12 cells were grown under two conditions to assess the activity of NGF loaded in the various PEDOT matrices. Fluorescent micrographs,see Figure 6.10, taken at 96 hr show the degree of cell adherence and neurite outgrowth in the polymers loaded with NGF compared to those in which the media was supplemented with 50 ng/mL NGF. The media supplemented assays were performed on polymers that were not loaded with NGF and are used as growth controls. The control films were clearly superior in all instances.

Three fluorescent micrographs were taken for each well, with three replica wells used in each study. The results of three studies were averaged to produce the graphs of cell density, Figure 6.11 and neurite length, Figure 6.12, with error bars representing the SE across the multiple studies. It should be noted that the control films do not mirror those presented in the previous chapter due to reduced laminin coating periods used in this study to preserve diffusive NGF.

The cell densities reveal a distinct decrease in cell number when the NGF is incorporated into the polymer. PEDOT/pTS/NGF had the highest cell density when NGF


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Figure 6.10: Sample images of PC12 neurite outgrowth on PEDOT loaded with NGF (left) compared to polymers produced without NGF grown in NGF supplemented media (right): i. PEDOT/pTS/NGF; ii. PEDOT/pTS; iii. PEDOT/DEDEDYFQRYLI/NGF; iv. PEDOT/DEDEDYFQRYLI; v. PEDOT/DCDPGYIGSR/NGF; vi. PEDOT/DCDPGYIGSR.

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was not provided in the media, with 12,700 cells/cm2 compared to 16,900 cells/cm2 in the PEDOT/pTS control. This is equal to a 75% proportion of the control. PEDOT/DEDEDYFQRYLI/NGF supported only 53% of the PEDOT/DEDEDYFQRYLI control and PEDOT/DCDPGYIGSR/NGF supported 38.5% of the cell density recorded for its control film.

Similar results were seen for neurite length, where the NGF loaded films were unable to replicate the results seen for their control films in the presence of NGF supplemented media. Again, PEDOT/pTS/NGF came closest to replicating the in vitro cell behaviour of the control films. PEDOT/pTS/NGF was able to support an average of 17.5 cm of neurite extension compared to 23.5 cm of neurite extension measured on the PEDOT/pTS control.

The Pt control had negligible neurite extension despite a cell density of 4000 cells/cm2 when no NGF was present in the cell culture environment. This would suggest that some NGF was available in the peptide doped PEDOT films which supported some neurite outgrowth. PEDOT/DEDEDYFQRYLI/NGF supported approximately 7 cm of neurite extension per cm2 and PEDOT/DCDPGYIGSR/NGF supported approximately 9 cm of neurite extension per cm2 .


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Figure 6.11: PC12 cell density on laminin peptide doped PEDOT with NGF incorporation at 96hrs post-plating. Polymers are shown both with NGF incorporation and with NGF in media only. Pt and PEDOT/pTS are used as controls. Error is represented as SE, (n=3).

Figure 6.12: Neurite outgrowth of PC12s on laminin peptide doped PEDOT with NGF incorporation at 96hrs post-plating. Neurite outgrowth is represented as the total neurite length supported per cm2 . SE is given, (n=3); (∗ NSD).


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6.4

Summary of Results

Table 6.5 summarises the properties of each polymer type to aid in the comparison of materials. Where quantitative values are important, they are given. For other properties a qualitative interpretation of a range of values is used to indicate the polymer’s overall performance.

Table 6.5: Summary of results for NGF loaded PEDOT Analysis PEDOT/pTS PEDOT/YFQRYLI PEDOT/YIGSR Surface Morphology Very rough Smooth + rough phases Globular features Efficiency of Deposition Unquantifieda Unquantifieda Unquantifieda Electrochemical Stability Good Poor Poor Impedance (Ω at 1kHz) 120.7 136.0 146.3 Hardness (Scratch) H 2B 2B Adherence (% Loss) 9.61 30.07 32.39 Bioactivity of NGF Good Limited Limited a

NGF detected but unquantifiable by XPS

6.5

Discussion

NGF loaded PEDOT was produced by entrapping the NGF placed in the electrolyte solution during polymerisation. The combined effect of two large biomolecules, a peptide dopant and NGF, on polymer physico-chemical properties was found to be substantial. Where NGF was entrapped without severely degrading film properties, in the case of PEDOT/pTS/NGF, the cell response demonstrated that NGF biofunctionality was preserved and neurite extension was promoted.

6.5. DISCUSSION

157

NGF is known to have a net electropositive charge, but the molecule is very large and electropositivity is weak. In a report by Thompson et al. cyclic voltammetry was used to demonstrate that neurotrophins are not electroactive as they produce no detectable oxidation and reduction curve [64]. The basis of the neurotrophin incorporation is not fully understood but is thought to involve ionic interactions with dopant anions, hydrophobic interactions with either dopant molecules or monomer unit of the forming film, or physical entrapment of the protein molecules.

Electrodeposition of PEDOT coatings containing two biomolecules was difficult with evidence of diffusion limitations preventing the formation of polymer at the electrode surface. Films were fragile and easily damaged. SEM revealed the surface topography of the NGF loaded polymers was more porous with a higher density of surface features than their unmodified controls, presented in Chapter 5, Figure 5.2. XPS was used to confirm the presence of the NGF in the polymer films, but the inclusion of NGF and the peptide dopant was unable to be quantified using this analytical technique. Additional analysis of films using specific protein tagging assays are required to determine the availability of each biomolecule in the polymer matrix.

While impedance spectroscopy demonstrated that the NGF loaded polymers were more resistive than PEDOT/pTS, with reduced phase lag and impedance magnitude in the low frequency region, electrochemical stability was severely reduced by the inclusion of the NGF molecule. The large (≈ 26 kDa) non-doping molecule that is entrapped during the electropolymerisation reduces the efficiency of polymer formation and consequently

158


degrades the electrochemical activity of the polymer. The CV curves obtained from the peptide doped, NGF loaded materials are significantly smaller in peak-to-peak amplitude than unloaded peptide doped PEDOT. This is reflected in the final charge storage densities calculated for these polymers. PEDOT/DEDEDYFQRYLI/NGF contains the highest nominal molecular weight of biomolecules within the PEDOT, resulting in a charge storage density of 12 mC/cm2 , only two times greater than the published value of 6 mC/cm2 for Pt [20]. PEDOT/pTS/NGF with only one biomolecule inclusion has a significantly higher charge density (117.8 mC/cm2 ) and preservation of electrochemical activity (≈ 88%).

Where two biomolecules are included it is hypothesised that diffusion limitations dominate during the electrodeposition and the resulting PEDOT matrix is so poorly formed that even large dopants can migrate out of the polymer on reduction and are quickly displaced from the electrode surface resulting in poor electrochemical stability. Research indicates that lower charge densities are seen for polymers including peptides or neurotrophins [16, 35, 110] but longer-term analysis that requires repeated CV or polarisation cycles has not been reported.

Diffusion limitations effecting polymerisation is supported by the mechanical adherence test where both PEDOT/DEDEDYFQRYLI/NGF and PEDOT/DCDPGYIGSR/NGF are poorly formed and experience high (30%) loss of coating area from the relatively low force applied by low-tack tape. The loss of polymer is seen across the entire film and not concentrated at the x-cut disruption, confirming that the adherence properties of the polymer are uniformly poor across the electrode. Coating loss is significantly less for

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159

NGF loaded PEDOT/pTS at less than 10%, but the impact of NGF is still considerable compared to the PEDOT/pTS control loss of 0.5%. Reduced polymer adherence may be balanced by an increase in softness. PEDOT/pTS was one of the hardest basic polymers tested in Chapter 3 and the reduction in strain-mismatch provided by a softer interface would be beneficial.

The results of this study indicate that the presence of NGF as a non-doping molecule, severely disrupts the PEDOT matrix in the presence of peptide dopants. With a small pTS dopant the matrix is able to form efficiently while still constraining the NGF molecule with a relatively low impact on physico-chemical properties. However for peptide doped CPs where the dopant size has limited the formation of the PEDOT to a degree where lower mechanical and electrical performance is observed, the NGF molecule inhibits the PEDOT formation and has a substantial negative impact on these physical parameters. This phenomenon was not seen by Stauffer and Cui when they doped PPy with two peptides, but in this instance both molecules performed doping functions and the PPy matrix was not required to accommodate a higher volume of biomolecules [110]. Other groups who have incorporated large biomolecules into CPs have not performed mechanical testing and have only conducted limited electrical analysis [16, 37, 64]. Collier et al. reported that HA doped PPy was brittle and difficult to handle with a greatly reduced conductivity (four orders of magnitude lower) compared to PPy/PSS [37].

PC12 cell studies were able to confirm that NGF maintained activity when incorporated into PEDOT during electrodeposition. While the peptide doped, NGF loaded polymers

160


were not able to replicate the cell results recorded for NGF supplemented media, PEDOT/pTS/NGF had high cell densities and was able to produce neurite outgrowth with NSD to the media control. Low cell densities and neurite outgrowth was seen on both peptide doped substrates where NGF was supplied by the polymer compared to growth on substrates with NGF supplemented media. Despite low neurite outgrowth, these substrates were still superior to bare Pt, and a small degree of neurite outgrowth indicates some NGF is available and bioactive within these polymers.

Kim et al. have reported the incorporation of both collagen and NGF into PPy, and confirmed the activity of both biomolecules through PC12 neurite outgrowth assays [12]. Similar PEDOT studies were carried out, but collagen and NGF were not incorporated into the same film, presumably due to the poor film stability. PEDOT/NGF was shown to promote PC12 neurite outgrowth but only once coated with collagen prior to cell plating [12]. The mechanical and long-term electrochemical properties of these conducting polymers were not reported.

The difference in biological performance seen between PEDOT substrates can be explained by the formation of the polymer, reflected in the film physico-chemical properties. It is hypothesised that the accessibility of NGF in PEDOT/pTS/NGF was greater than that of peptide doped PEDOT. Where the PEDOT formed a durable, electrochemically stable film, as in the case of PEDOT/pTS/NGF, the polymer matrix limited the mobility and diffusion of NGF, maintaining neurite growth promotion for the 96 hr culture period. In comparison, less neurite outgrowth was seen for the peptide doped films and can be

6.5. DISCUSSION

161

explained by the polymer structure. If the polymers were not well formed as hypothesised, two film properties will contribute to poor cell results. Less NGF will be included in the films, reducing the amount available to the cells, and the little NGF incorporated in the films will quickly diffuse out, leaving a very low level of NGF to maintain the culture over the 96 hr growth period.

These studies suggest that more research is required to produce a stable conducting polymer with both cell attachment and neurite outgrowth bioactivity. Cell attachment mechanisms need to be controlled by either a much smaller peptide dopant or modification of the surface topography for optimal cell interactions. When large biomolecules interfere with conducting polymer formation, future modifications should look at alternate methods for delivering signaling factors. One method investigated by Li et al. for drug delivery applications employs a layered hydrogel-conducting polymer composite [81, 82]. The overlying hydrogel layer was reported to support higher cell densities than TCP and serves as a reservoir for high concentrations of biomolecules. Heparin was released from the overlying hydrogel through electrical stimulation and results suggest that this type of composite has potential to be applied in neuroprosthetic design. Research by Kim et al. has shown that PPy with an underlying hydrogel layer applied to cortical electrodes can be used to alter the conducting polymer surface topography [20, 46]. The resulting material had a lower impedance and improved cell attachment properties than PPy/PSS alone [20]. However, the characterisation of either of these hydrogel composites for long-term electrochemical stability, mechanical properties and the controlled release


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of neurotrophins to promote cell ingrowth has not yet been reported.

It is important to note that these cell studies were not done under electrical stimulation and the release of the neurotrophin component by stimulation regimes typically used in neuroprostheses may affect the biological performance of these conducting polymers.

6.6

Conclusions

The inclusion of both a laminin peptide dopant and non-doping neurotrophin in PEDOT has been demonstrated to severely degrade the performance of this polymer in mechanical, electrochemical and biological analyses. The NGF molecule was a dominant factor in the determination of polymer properties, with little difference seen between the two peptide doped variants. Most significantly the long-term electrochemical stability was greatly reduced when compared to the peptide doped PEDOT without NGF and charge storage capacity was significantly lower than values recorded for alternate polymer-dopant combinations. Mechanical adherence and hence durability deteriorated with the entrapment of NGF in peptide doped PEDOT. In vitro neither NGF loaded PEDOT/DEDEDYFQRYLI nor PEDOT/DCDPGYIGSR, was able to replicate the cell adherence and neurite outgrowth recorded for their respective controls.

PEDOT/pTS/NGF demonstrated superior performance to peptide doped PEDOT in all categories, including neurite outgrowth assays. Where some decrease was seen in longterm electrochemical stability and coating adherence it was offset by lower impedance

6.6. CONCLUSIONS

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and hardness values. The smaller dopant molecule provided a more cohesive coating than peptide doped PEDOT and the surface roughness was clearly compatible with in vitro cell attachment. Cell studies indicated that PEDOT/pTS entrapped NGF was accessible to the cells, with only a small reduction in neurite outgrowth compared to controls. Further research is required to quantify the inclusion of NGF into PEDOT/pTS during electrodeposition and its subsequent diffusion under cell culture conditions.

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Part II

Application of Conducting Polymer Electrode Coatings to Neuroprostheses

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Chapter 7

Application of Conducting Polymer Coatings to Microelectrodes 7.1

Introduction

All previous conducting polymer studies in this thesis have been performed on a model electrode system detailed in Section 3.2.1.1. The model system was efficient in assessing a variety of film physico-chemical and biological properties, but to advance the development of conducting polymers for electrode coatings they also require assessment on a micro-scale typical of a neuroprosthetic implant. The bioactivity of conducting polymers on the model electrode was most effective for PEDOT/pTS/NGF, but application to a significantly smaller electrode representative of an implant electrode needs to be characterised. There are a number of applications for which bioactive electrodes might improve efficacy and long-term viability of the implant. One such example is the vision prosthesis or bionic eye currently under development by the Australian Vision Prosthesis Group (AVPG) at 166

7.1. INTRODUCTION

167

UNSW.

The AVPG epiretinal vision prosthesis consists of three main components: a camera, a processor and a stimulating implant, as depicted in Figure 7.1. The camera mounted at eye level on sunglasses captures the image, which is then transmitted to a body worn external processor. The processor converts the image to a form that can be encoded and delivered to the implant by radio frequency (RF) telemetry. The signal is received by the implant that is surgically placed around the inner circumference of the eye to interface with the retinal tissue near the densely populated macular. The signal is decoded and stimulus delivered to neural cells via a microelectrode array.

Figure 7.1: AVPG vision prosthesis schematic

The epiretinal electrode array is an integral part of the bionic eye, allowing communication between implanted electronics and the retinal tissue. Through this communication, controlled electronic stimulation is delivered to neurons and vital impedance information can be recorded and relayed to the external processor. The current prototype uses a

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CHAPTER 7. APPLICATION OF CP COATINGS TO MICROELECTRODES

14-electrode array, but can be expanded by the stacking of numerous arrays to form the densely packed 98-electrode array pictured in Figure 7.2. The bondpads, interconnects and electrode sites are all fabricated from laser ablated flexible Pt foil embedded in a medical grade silicone [1].

Figure 7.2: AVPG 98-electrode array for high resolution vision prostheses [1].

To characterise the bioactivity of conducting polymers on microelectrodes the AVPG epiretinal electrode array was employed. Prior to assessing conducting polymer bioactivity, an L929 growth inhibition assay was performed to determine the possible toxic effects that could occur as a result of the electrode materials and their processing. Materials similar to those previously approved by the FDA for use in other medical applications such as the cochlear implant [92], non-cardiac pacemakers [123] and breast implants [124], were specifically chosen for electrode fabrication. However, the effect of specific material processing techniques required to manufacture the AVPG eletrode, including the application of solvents and laser ablation has not been established. The L929 cell line is commonly used in determining cytotoxicity of materials and hence is used in growth inhibition as-


169

says [125].

Following the assessment of electrode material toxicity, the best performing bioactive polymer, PEDOT/pTS/NGF was deposited on the microelectrode array and assessed through neurite outgrowth using the PC12 cell line. This assay was used to determine if bioactivity of conducting polymers is preserved when they are coated on a microelectrode array.

These studies aimed to:

1. Assess the impact of AVPG electrode array fabrication techniques on material properties through cell growth inhibtion assays, 2. Determine if cells will preferentially adhere to conducting polymers when deposited on the microelectrodes, and 3. Assess the bioactivity of polymer coatings present across a micro-sized exposed electrode area.

7.2 7.2.1

Materials and Methods Fabrication of AVPG Microelectrodes

Dummy electrodes were required for both the L929 growth inhibition and the PC12 neurite outgrowth assays. The microelectrode array was produced from materials used by the AVPG in the format of an epiretinal electrode [1, 126], adapted to the Flexiperm

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well configuration used in all previous cell studies. The electrode material composition is confined to that of 99.95% pure Pt foil and PDMS (MED-1000, NuSil, Carpinteria, CA, USA) diluted in a one-to-one ratio with a hydrocarbon solvent (n-Heptane, Ajax Chemicals Ltd, Sydney, NSW, Australia).

The electrode was produced on a glass microscope slide (standard, 75 mm x 25mm dimensions) at the AVPG laboratory, UNSW 1 . Initially, a layer of uncured, n-Heptane diluted PDMS was spin coated onto the slide. A 20 μm thickness was achieved using a 2000 rpm spin for 90 seconds at room temperature of 21◦ C. The layer was cured in a humid oven at 80◦ C for 20 min, allowing the n-Heptane to evaporate and ensuring layer stability for further processing. Removal from the oven was followed by immediate application of a layer of 99.95% pure Pt foil of 25 μm thickness (Surepure Chemetals, USA). The required electrode configuration is then patterned onto the Pt foil using a numerically-controlled mirror to direct a laser beam (Nd:YAG, Firescan DPL Genesis Marker, CAB GmbH, Karlsruhe, Germany). A second layer of PDMS/n-Heptane was spun on using identical process parameters such as to fully insulate the Pt. Finally, the overlying PDMS was removed by laser ablation to facilitate exposure of the metallic surface and provide access to the neural tissue in a manner representative of the electrode configuration.

1

Dummy microelectrode arrays were produced by Chris Dodds, GSBmE, UNSW


7.2.2

171

L929 Growth Inhibition Assay

Growth inhibition was assessed for the electrode array and also the component materials Pt and silicone. Positive control latex and negative control TCP were run in parallel. All materials assessed in this experiment were mounted on microscope slides and clamped into the well assembly described in Section 3.2.1.1, except the TCP negative control. Since TCP could not be mounted in the existing assembly, a 96-well sterile TCP plate with a cell growth area identical to the silicone wells was used for the negative control. Other materials were mounted on the microscope slide as shown in Figure 7.3.

Figure 7.3: Well configuration for assessing componental materials and the fabricated electrode toxicity through L929 growth inhibition

The electrode array wells were produced as detailed in Section 7.2.1. The silicone and Pt samples were mounted on the slide during fabrication of the electrode and the whole slide was cleaned by soaking in 1% Decon90 overnight. The samples were thoroughly rinsed in DI water and dried in a 60◦ C oven. The latex positive control sample was cut to size and placed in the appropriate location without additional adherence.

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The Flexiperm silicone well gaskets were washed as detailed in Section 3.2.3 and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. The wells were attached to the slides on which the electrode, component materials and latex were mounted and the system was then clamped in the custom-built sandwich assembly to prevent leaking between adjacent well chambers. Disinfection was performed by placing the wells under UV for 2 hr. The negative control TCP wells were supplied sterile and hence were not placed under UV.

L929s grown to confluency in 25 cm2 TCP flasks were detached by application of 2 mL trypsin. After 5 min incubation the trypsin was deactivated by the application of 5 mL DMEM containing 10% FBS. The supernatant was pipetted into a polypropylene centrifuge tubes and centrifuged at 1000 rpm for 5 min. The supernatant was dumped, leaving a visible cell clump. A 1 mL Eppendorf pipette tip was used to add 1 mL of fresh media to the cell suspension. Cells were aspirated gently using the 1 mL pipette tip to break up the clumps. DMEM with 10% FBS was used to dilute the cell solution to 5 mL and a 0.5 mL sample was taken for counting. Cells were plated on the various samples and three wells of the 96-well TCP plate at 10,000 cells/cm2 . Two plates for each material type were used with the first plate cultured for 4 hr and the second cultured for 48 hr in DMEM with 10% FBS at 37◦ C, 5% carbon dioxide (CO2 ) and 100% humidity.

At 4 hr and 48 hr the L929 cells were live stained using Calcein-AM (1 μg/mL in PBS incubated for 20 min) and Hoechst 3342 (Invitrogen, Australia) nuclear stain (0.5 μg/mL in PBS incubated for 10 min). After the stain uptake by living cells, wells were gently


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washed with 2 to 4 volumes of PBS to remove excess fluorescent dye. Cell experiments were analysed through fluorescent micrographs taken on a Carl Zeiss fluorescent microscope with camera attachment. Cell numbers were quantified by analysing the Hoescht nuclear stain as the Calcein AM revealed that adhered and proliferating cells were overlapping, making analysis difficult. By removing the green portion of the fluorescent image, NIH software Image J could be employed to count the viable nucleii in each image.

For each single experiment, each substrate type was run in three duplicate wells. For each well three fluorescent micrographs were taken by scanning the film such that each image was obtained from a different sector (third) of the well area. The cell growth was calculated by comparing the adhered cell number at 4 hr to the proliferated cell number at 48 hr. The cell proliferation for each material was normalised against the TCP negative control to give cell growth inhibition and is presented as mean ± standard error, (n=3).

7.2.3

PC12 Neurite Outgrowth on PEDOT/pTS/NGF Coated Microelectrodes

For the bioactivity assay the dummy electrode was cleaned by soaking in 1% Decon90 overnight, thoroughly rinsed in DI water and dried in a 60◦ C oven. The Flexiperm silicone well gaskets were washed as detailed in section 3.2.3 and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. The wells were attached to the slides on which the dummy microelectrodes were mounted and the system was then clamped in the custom-built sandwich assembly to prevent leaking between adjacent well chambers.

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The PEDOT/pTS/NGF was polymerised by electrodeposition using previously established parameters detailed in Section 6.2.1. The galvanostatic current was adjusted to the reduced area of exposed microelectrodes (approximately 300 electrodes of 200 μm diameter or a combined area of approximately 0.09 cm2 ) such that the current density was maintained at 1 mA/cm2 for 10 min. Following deposition the electrodes were washed carefully four times with DI water and placed in the incubator overnight covered with the last 150 μL of water to leach water soluble contaminants. Electrode arrays without polymer coating were fabricated for control purposes with cell culture performed on both substrates simultaneously under identical conditions. Sterilisation was performed by placing the wells without any overlying fluids under UV for 2 hr.

As previously discussed the PC12 is a non-adherent cell line which requires laminin or collagen to grow on most culture surfaces, however PEDOT/pTS was shown in Chapter 5 to promote good cell adherence without laminin coating. To validate PEDOT/pTS/NGF as an electrode material it was assessed without laminin coating. PC12s were directly plated on the electrodes as detailed in Section 3.2.6 and cultured for 96 hr in low serum (1% HS) RPMI without NGF supplements. Cell image and statistical analyses were carried out as for previous studies outlined in Sections 3.2.6.1 and 3.2.6.2.


7.3


7.3.1

L929 Growth Inhibition

175

Samples of the micrographs taken at 48 hr are shown in Figure 7.4. The latex sample was not shown as no cells survived on this known cytotoxic substrate. Conversely, it was observed that cells on the negative control TCP had spread and proliferated to a greater degree than the cells on any of the other substrates. Despite Pt and silicone being widely used in medical implant devices, the cell adherence and interaction with these substrates was poor.

The cell counts taken using ImageJ were normalised in Figure 7.5 to the TCP negative control. The results revealed that growth inhibition on the AVPG electrode substrate was not significantly different to the unprocessed electrode component materials, Pt and medical grade silicone. It can be concluded from these result that the fabrication techniques, employing both N-heptane thinning and laser ablation of the silicone, did not increase the growth inhibition of the fabricated electrode.

7.3.2

PC12 Neurite Outgrowth on PEDOT/pTS/NGF Coated Microelectrodes

PEDOT/pTS/NGF was deposited on the dummy electrode and LYNX reflective microscopy was used to confirm the presence of the polymer. The image in Figure 7.6 shows that the electrodes were well coated, appear uniform and discrete (without bridging of

176


Figure 7.4: Sample images of L929 growth on electrode materials at 48 hr: i. TCP negative control; ii. Pt; iii. MedNusil silicone rubber; iv. AVPG electrode array. (Latex, the positive control is not shown as this is a known cytotoxic material resulting in the death of the cell population.)


177

Figure 7.5: TCP normalised L929 growth inhibtion of electrode componental materials and the fabricated AVPG electrode. Error bars represent SE, (n=3). the polymer between adjacent electrodes).

It is important to note that the cells in this study were plated directly on the polymer coated electrode array without any laminin coating or NGF in the media. Previous studies indicated that PEDOT/pTS has surface topography ideal for cell attachment without cell attachment ligands. It has also been shown that NGF can be incorporated into the PEDOT/pTS matrix without substantially degrading the electrical and mechanical properties of the polymer. In previous cell studies PEDOT/pTS/NGF performed with similar results to PEDOT/pTS with NGF supplemented media. The current study extrapolates on this concept through polymer deposition on a microelectrode array used in prototyping the AVPG epiretinal electrode.

Fluorescent micrographs taken at 96 hr post-plating showed that PEDOT/pTS/NGF de-

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Figure 7.6: PEDOT/pTS/NGF electrodeposited on AVPG electrode array.

posited on Pt was successful in creating an environment conducive to both cell attachment and neurite outgrowth. In Figure 7.7 a sample cell image has been superimposed over a reflection micrograph of the electrode prior to plating and shows that cells preferentially attached to the polymer coated electrodes and sprouted neurites in response to NGF loaded in the PEDOT/pTS. Cell counts revealed that 74.6% of the viable PC12s were adhered to the PEDOT/pTS/NGF. Minimal adherence occurred on the silicone insulation despite making up more than 50% of the available surface area. Cells present on the control electrodes were all confined to the rough silicone edges around the border of the microelectrodes.

Values measured from the fluorescent micrographs were compared to those for the bare electrode with PC12s plated under identical conditions. It can be clearly seen that the PEDOT/pTS/NGF dramatically improves the cell response to the electrode array, plotted


179

Figure 7.7: PEDOT/pTS/NGF electrodeposited on AVPG electrode array with PC12s at 96 hr. Preferential adherence is seen on the electrode surface with good neurite outgrowth stimulated from the NGF loaded polymer. Note: the fluorescent cell image is superimposed over a reflection micrograph of the electrode prior to cell plating to highlight the location of cells relative to the polymer coated electrodes.

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in Figure 7.8. Both cell density and supported neurite length are significantly better when the exposed Pt is polymer coated (p < 0.05). This is particularly beneficial considering that the control electrode is fabricated from materials approved by the FDA for use in other neuroprosthetic devices [92,123] and in the preceeding assay demonstrated minimal cell growth inhibition against the L929 cell line.

Figure 7.8: Cell density and neurite outgrowth of PC12s on PEDOT/pTS/NGF coated AVPG electrode array compared to bare electrode. The bioactive polymer shows both cell attachment and neurite outgrowth functionality when placed in low serum media with no additional cell signaling factors. Error bars represent SE, (n=3); (∗ p < 0.05).

7.4. DISCUSSION

7.4

181

Discussion

Studies were carried out on electrode materials and the manufactured AVPG electrode array to determine possible cell growth inhibition arising from the fabrication process. It was shown that no significant cell inhibition occurred as a result of the processing, with inhibition rates for the processed electrode being similar to that of the component materials which are commonly used in commercially available implants [92, 123, 124]. The PC12 neurite outgrowth assay assessed the micronisation of electrodes comparing the bare electrode array to an array coated with bioactive polymer PEDOT/pTS/NGF. Preferential adherence of cells to the conducting polymer depositions was observed and bioactivity of the NGF was evident.

The cell growth inhibition assay determined the rate of L929 proliferation on homogeneous medical grade silicone, 99.95% pure Pt foil and the composite laser ablated Pt-silicone electrode array over a 48 hr period. All values were normalised to the TCP negative control where the rate of L929 proliferation was optimal. Proliferation rates on the silicone, Pt and the Pt-silicone composite were not significantly different, revealing that no additional growth inhibition and hence no toxic by-products are formed during the fabrication of the electrode with solvent N-heptane and laser ablation. Similar results were reported by Reuter et al. who assessed L929 proliferation on a cochlear implant developed at Cochlear Ltd., Sydney [127]. In this application two types of silicone and Pt were assessed. It was reported that cell growth occurred on all electrode materials but was lower on the silicone than on Pt [127]. By observing the cell images taken in

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this study (Figure 7.7), it can be seen that higher cell numbers adhere and flatten on the Pt substrate than the silicone, but neither substrate has the optimal properties of TCP, where the cells maintain a healthy flat morphology due to surface modifications which encourage cell interaction. While the fabrication process of the AVPG electrode does not impact on L929 cell interactions, it has become evident that material surface properties will play an integral role in electrode performance in vivo.

In respect to this thesis, subsequent cell experiments using the AVPG electrode array should not have been significantly impacted by fabrication by-products. The processed material should perform similarly to the Pt controls used throughout all other cell growth assays.

Throughout the course of this research neurite outgrowth assays have been performed on a model electrode with a 5 mm diameter. To assess the bioactivity of conducting polymers for use in neuroprosthetics the AVPG electrode array was used with 300 electrodes of 200 μm diameter exposed to the PC12s. The bioactive polymer PEDOT/pTS/NGF was assessed due to its superior performance across electrical, mechanical and biological properties. PC12 neurite outgrowth assays demonstrated that the PEDOT coated electrode array supported almost three times more cells than the bare electrode array, of which 75% were adhered to the polymer portion of the exposed composite. Consequently, the polymer coated electrode array was able to support a six fold increase in neurite length per unit area compared to the uncoated control.

7.4. DISCUSSION

183

Preferential adherence of cells to conducting polymers has been reported by a number research groups across a variety of cell types. Cui et al. demonstrated that glial cells preferentially bind to PPy incorporating fibronectin fragments and neuroblastomas favoured CDCPGYIGSR doped PPy [35]. Song et al. showed that micropatterns produced through covalent attachment of polylysine to PPy could be used to control neuron adhesion and extension [128]. This assay differs to studies reported in literature, in that a cell attachment ligand is not employed and the cell type is non-adherent, indicating that polymer surface topography is a dominant factor in controlling cell attachment. Additionally, most reports of preferential attachment of cells to conducting polymer substrates in the literature involve the use of cells that are more naturally adherent than the GFP-PC12 used in this assay.

It is worth recognising that the nominal cell density for this assay was 20,000 cells/cm2 and cell counts indicated an average density of 11,800 cells/cm2 on the PEDOT/pTS/NGF coated electrodes. Previous studies have shown that the polymers can support twice as many cells and up to five times greater neurite length when laminin coated. However, in this assay the preferential surface topography of the polymer minimises the attachment of cells on the insulating silicone and might be beneficial in reducing neuron spread and neurite outgrowth across more than one electrode. Cells with neurites spanning multiple electrodes could result in cross-talk between the electrodes, reducing the neuroprosthetic signal resolution.

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7.5


Conclusions

Processing of Pt and medical grade silicone through methods developed by the AVPG resulted in an electrode array without increased toxicity compared to the unprocessed controls. Coating this electrode with the optimal conducting polymer composite, PEDOT/pTS/NGF, created a substrate with significantly better cell interactions than that seen on the uncoated electrode. PC12 cells preferentially adhered to the polymer and minimal cells adhered to the surrounding silicone. The NGF loaded polymer can stimulate substantial neurite outgrowth. Future studies will examine the effect of common neuroprosthetic stimulation regimes on the release of NGF from PEDOT.

Chapter 8

Tailoring Conducting Polymer Bioactivity to an Alternate Cell Type, the RGC-5 8.1

Introduction

In previous studies the PC12 neural-like cell has been used to model the impact of conducting polymers on a biological environment. As detailed in Section 3.2.1.2 PC12 is a robust cell line that can be used for high-throughput assays with good repeatability of response and allows comparison with data in the literature as PC12s are widely used as a model strain. However, this cell is greatly removed from the primary cells that are intended to interface with neuroprosthetic implants. Having established that cells can be grown on conducting polymer coatings, it is desirable to test them against a more biologically appropriate cell. As for the previous study the vision prosthesis will provide the neuroprosthetic basis for application. 185

186CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE A commercially viable vision prosthesis will restore some form of visual perception to patients blinded by diseases of the retina, where a significant proportion of the optic nerve pathways are preserved. This encompasses blindness associated with retinitis pigmentosa (RP) and age-related macular degeneration (AMD), where cell death in the photoreceptor layer of the retina, causes a progressive loss of sight and subsequent, transneuronal degeneration of cells further along the visual path. AMD affects between 25 and 30 million people globally and is the leading cause of blindness in developed countries due to the increasing number of older people [129]. In these ageing populations the cost of AMD is estimated at AU$2.4 million per country [130] with the direct costs of AMD in Australia during 2004 totalling AU$19.4 million [131]. Similarly RP is the leading cause of inherited blindness among individuals between the ages of six and 60 and affects 1.5 million individuals worldwide [132].

Both RP and AMD are caused by photoreceptor losses from the outer layer of the retina, depicted in Figure 8.1. The loss of cell function at this level terminates the progression of a normal biological signal cascade throughout the remaining cellular layers and ultimately to the visual cortex. As a result the remaining neural cells including the horizontal, bipolar and amacrine cells of the intermediate cell layer and the retinal ganglion cells (RGCs) undergo losses through a mechanism known as transneuronal degeneration [133]. Retinal implants aim to minimise transneuronal cell death and artificially restore functionality to the intermediate cells and/or RGCs.

The feasibility of epiretinal and subretinal implants to restore sight has been investigated

8.1. INTRODUCTION

187

Figure 8.1: Cell structure of the human retina [134] by several groups throughout the world [135–139]. Clinical studies have shown that controlled electrical stimulation of retinal areas results in localised retinotopically correct perceptions in patients affected by RP and AMD [10, 137].

The target cells for stimulation, using an epiretinal vision prosthesis, are primarily the retinal ganglion cells (RGCs) [136, 138, 140–142]. RGCs are reduced in number by transneuronal degeneration subsequent to disease-related photoreceptor death associated with RP or AMD [133], however studies have shown that numbers of RGCs are still functional, although the exact percentage of RGCs retained varies with the type and the severity of the disease [133, 143–145]. Average preservation of RGCs specifically in the macular

188CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE region of exudative AMD patients is 52.7% and non exudative AMD patients is 57.2%, which does not differ significantly from age matched controls [133]. These results suggest that stimulation of RGCs in both RP and AMD patients is viable, however each individual will require assessment to locate the area of their retina that is most likely to respond to stimulation, the area most densely populated by RGCs.

To assess cell response to conducting polymers applied to vision prostheses it would be ideal to employ primary RGCs obtained from a retinal explant. Primary RGCs are most commonly harvested from rats and purified by immunopanning. This technique is most effective when performed on young (< 10 day old) rats or chicks when immunoreactivity of the protein marker used for sorting (Thy-1) is at its greatest [146–148]. This technique is time consuming, requires substantial resources and several animals have to be sacrificed to achieve sufficient cell yields [149, 150]. Additionally, in vitro primary RGC neurite outgrowth is most successful when a combination of neurotrophins and other factors, similar to those in developing mammalian eye, are available to the cells possibly including BDNF, CNTF, NT-3, bFGF and forskolin [151–154]. The cost of including all three biomolecules in a conducting polymer is considerable and as such a cloned RGC cell line will be employed. The use of a cloned RGC to replace primary culture has been suggested by a number of research groups [150, 155, 156], but the development of a stable line with characteristics similar to a primary RGC has only recently been reported [157].

To assess the validity of bioactive conducting polymers in a vision prosthesis application, the system was taken to the University of North Texas, Fort Worth Health Sciences Center

8.1. INTRODUCTION

189

(UNTHSC) where it was tailored and tested against the cloned RGC-5 cell line. The RGC-5 cell line was developed at UNTHSC to provide insights into vision neuroscience by allowing high throughput assays unable to be conducted on the more fragile and timeconsuming primary RGCs [157, 158]. Based on the expression pattern of various specific cell markers, Krishnamoorthy et al. developed the RGC-5 cell line and demonstrated its many characteristic phenotypes of primary RGCs [157]. The RGC-5s were shown to express Thy-1, Brn-3C, various neurotrophins and their receptors showed a dependence on trophic factors for survival [157, 159]. The more robust RGC-5 was considered an acceptable choice for assessing conducting polymers against a cell line closer to a primary RGC than the previously assessed neural model PC12.

In this assay both PEDOT/pTS and PEDOT/DEDEDYFQRYLI were assessed. Previous assays demonstrated little difference in physico-chemical properties of the polymers produced from the two different peptides, with the dominant factor being the large size of the dopant and the resultant surface topography. DEDEDYFQRYLI was chosen as it was initially hypothesised to have bioactivity specific to both cell attachment and neurite outgrowth and the adherent RGC-5 cell line may be more responsive to the ligand.

The aims of this study were to:

1. Determine if bioactive conducting polymers developed in previous studies can be tailored to an alternate neural cell type, and

2. Assess the tailored bioactive polymer coatings against a cell line closer to a primary

190CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE retinal ganglion cell.

8.2 8.2.0.1

Materials and Methods Tailoring Polymer for RGC-5 Bioactivity

PEDOT films were tailored for RGC-5 culture by replacing the differentiating agent for PC12s, NGF, with staurosporine (SS) pictured in Figure 8.2. SS is a 466.5 Da nonspecific kinase inhibitor more commonly implicated in cell apoptosis [2, 160]. Researchers at UNTHSC have shown that SS can induce the mitotically active RGC-5 cell line to stop proliferation, extend neurites, and express many of the electrophysiological and histochemical markers characteristic of primary RGCs [159]. RGC-5s are adherent cells that show some increased attachment behaviour when placed in contact with laminin. As such the DEDEDYFQRYLI laminin peptide doped PEDOT was assessed, along with PEDOT/pTS and bare Pt controls.

Electrodeposition of PEDOT films doped with pTS or DEDEDYFQRYLI and incorporating SS was carried out using the electrolyte solutions in Table 8.1. SS was applied to the electrolyte at a concentration determined from 24 hr differentiation assays on tissue culture plastic (TCP). These assays demonstrated amounts as low as 20 nM of SS were required to stimulate changes in cell morphology. For sufficient growth of PC12s from NGF loaded polymers the amount of differentiating agent placed in the electrolyte solution had to be 20 times greater than that placed directly in the media. The TCP results of the RGC-5 differentiation assay where SS was directly added to the media were scaled


191

Figure 8.2: Structure of staurosporine (SS) molecule [2] accordingly such that 400 nM of SS was placed in the electrolyte solution prior to polymerisation. Films were electrodeposited, washed and prepared for cell culture according to Section 3.2.3. Table 8.1: Electrolyte solutions tailored for RGC-5 differentiation activity. Component Cat # PEDOT/pTS/SS PEDOT/YFQRYLI/SS EDOT 483028 0.1M 0.1M pTS 402885 0.05M DEDEDYFQRYLI 1254300 5 mg/mL a Staurosporine 400 nM 400 nM Deionised (DI) Water N/A 50% 50% Acetonitrile 271004 50% 50% a

Supplied by N. Agarwal, School of Cell Biology and Genetics, UNTHSC

8.2.0.2

RGC-5 Differentiation Assay

For in vitro RGC-5 validation assays, the same cell culture set up was used as previously described for PC12s. The silicone well gaskets were washed and soaked in 80% ethanol (EtOH) overnight to eradicate contaminants from previous studies. Wells and Pt slides

192CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE were placed in a beaker of DI water, covered and autoclaved for 20 min at 121◦ C. All components were air dried in a cell culture laminar flow hood. The wells were attached to the Pt slides and the system was then clamped in the custom-built sandwich assembly disscussed in Section 3.2.1.1 to prevent leaking between adjacent well chambers.

Following galvanostatic deposition of tailored polymers at 1 mA/cm2 for 10 min, films were washed carefully four times with DI water and placed in the incubator overnight covered with the last 150 μL of water to leach water soluble contaminants. To disinfect the films 200 μL of 80% EtOH was placed in each well and left for 40 min. The EtOH was then removed and the films were washed twice with sterile Baxter’s water before being placed under UV for 2 hr.

RGC-5s grown to confluency in 25 cm2 TCP flasks were detached by application of 2 mL trypsin. After 5 min incubation the trypsin was deactivated by the application of 5 mL DMEM containing 10% FBS. The supernatant was pipetted into a polypropylene centrifuge tubes and centrifuged at 1000 rpm for 5 min. The supernatant was dumped, leaving a visible cell clump. A 1 mL Eppendorf pipette tip was used to add 1 mL of fresh media to the cell suspension. Cells were aspirated gently using the 1 mL pipette tip to break up the clumps through 40 aspiration repetitions. Low serum DMEM with 1% FBS was used to dilute the cell solution to 5 mL and a 0.5 mL sample was taken for counting. Cells were plated at 40,000 cells/cm2 and cultured for 96 hr in low serum medium. Polymers containing the SS differentiating agent were cultured in 1% FBS modified DMEM. Control polymers without SS in the electrolyte were cultured in the


193

presence of 1% FBS modified DMEM supplemented with 20 nM SS. All cell studies were carried out in an incubator at 37◦ C, 5% carbon dioxide (CO2 ) and 100% humidity. Low serum media was renewed (with SS supplement where applicable) at 48 hr by exchanging 2/3 or 100 μL of media in each well.

At 96 hr the RGC-5 cells were live stained using Calcein-AM (1 μg/mL in PBS incubated for 20 min). Cell experiments were analysed through fluorescent micrographs taken on a Olympus fluorescent microscope with camera attachment. The lengths of individual neurites were traced and calibrated length determined using NIH software Image J, with Neuron J plugin.

Neurites were measured when the length of the projection exceeded a single body length of the cell from which it extended. When neurites branched, one split was considered a continuation of the primary neurite and each alternate split of longer than a cell body in length was considered a new neurite.

For each single experiment, each substrate type was run in three duplicate wells. For each well three fluorescent micrographs were taken by scanning the film such that each image was obtained from a different sector (third) of the well area. The cell density and neurite density supported by each substrate is presented and standard error reported.

194CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE

8.3


The polymers were easily fabricated with the inclusion of the SS, which is a much smaller molecule than NGF used in the previous chapter (466Da c.f. 26 kDa). Although particulate matter was observed during the electrodeposition of PEDOT/DEDEDYFQRYLI/SS, it was attributed to diffusion limitations often seen in the presence of the large peptide, as discussed in Chapter 5. Cells were stained with Calcein AM at 96 hr to make them visible under fluorescent microscopy. Sample images typical of the cell response observed on each polymer are shown in Figure 8.3.

This experiment was repeated during the period of study at UNTHSC to produce comparisons for the tailored polymer across the different dopant types with a Pt control. The cell density plot, Figure 8.4 shows that RGC-5s respond quite differently to PC12s in the absence of differentiating conditions, proliferating until confluency. However, under the right conditions the Pt control shows no proliferation and some neurite sprouting. The cell densities on both PEDOT films indicate that SS loading of the polymers prevents cell proliferation.

Combined with the neurite length recorded per cm2 in Figure 8.5, these results suggest that SS is entrapped at sufficient concentration to stimulate RGC-5 differentiation. The slightly higher cell densities on the SS loaded polymers, while not statistically significant, suggest that differentiation may be delayed by incorporating the SS in the polymer matrix. When SS is supplied directly to the media it is instantly accessible to the cells and


195

Figure 8.3: Sample images of RGC-5 differentiation on PEDOT loaded with SS (left) compared to polymers produced without SS grown in SS supplemented media (right): i. Pt control; ii. Pt with SS in media; iii. PEDOT/pTS/SS; iv. PEDOT/pTS; v. PEDOT/DEDEDYFQRYLI/SS; vi. PEDOT/DEDEDYFQRYLI.

196CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE morphology changes are visible within 15 min of application. When the SS is loaded into the PEDOT it passively diffuses out of the polymer before cells can access the molecule, ceasing proliferation and promoting differentiation. The neurite length supported per cm2 is not significantly different between SS loaded polymers and the media supplemented controls. This would suggest that these polymers can provide biomolecules at a concentration similar to that of modified medium.

Figure 8.4: RGC-5 cell density on SS loaded PEDOT vs media controls at 96hrs postplating. Pt is used as a control and shows cell proliferation when SS is not present. Error is represented as SE, (n=2).

PEDOT/DEDEDYFQRYLI has higher cell densities and neurite lengths for both SS loaded PEDOT and the media supplmented control than PEDOT/pTS. It is hypothesised that the YFQRYLI ligand may have greater affinity to this more adherent cell type than the PC12 which required additional laminin to produce similar results in Chapter 5. The high cell density of PEDOT/DEDEDYFQRYLI/SS could be attributed to the inclusion

8.4. DISCUSSION

197

Figure 8.5: Neurite outgrowth of RGC-5s on PEDOT/pTS and PEDOT/DEDEDYFQRYLI with SS entrapment at 96 hr post-plating. Error bars represent SE, (n=2). of the larger DEDEDYFQRYLI molecule slowing the release of SS and hence allowing more time for proliferation.

8.4

Discussion

The tailoring of conducting polymer bioactivity can be performed through manipulation of the agent intended to promote neurite ingrowth to electrode sites. In this assay the previous model based on NGF stimulated PC12s was altered such that a neuronal cell closer in characteristics to the primary RGC could be differentiated. The differentiating agent SS was released from PEDOT through passive diffusion mechanisms to promote RGC-5 differentiation and neurite extension. Tailoring of the PEDOT was successful

198CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE with differentiation occurring on both PEDOT/pTS and PEDOT/DEDEDYFQRYLI.

The use of a cell type closer to that of a primary RGC allowed in vitro assessment of cell behaviour on laminin peptide doped PEDOT without the use of an additional laminin coating required for most substrates when cultured with non-adherent GFP-PC12s. The peptide investigated featured the YFQRYLI ligand, previously identified by Tashiro et al. as promoting both neural cell attachment and neurite outgrowth [105]. When the differentiating agent SS was delivered via the polymer matrix the DEDEDYFQRYLI doped PEDOT had a 26% higher cell density and supported 42% greater neurite length per cm2 than pTS doped PEDOT. The substantial increase in neurite length would support the theory that in this system the YFQRYLI ligand is active. However, the effect of having a large dopant molecule creating a tortuous path for the diffusion of SS out of the polymer must also be considered.

The tortuous path phenomenon has been extensively used in analytical techniques such as gel chromatography, but has also been applied in drug delivery devices including antibacterial urinary catheters [161–163]. A tortuous path is the longer path length traveled by a mobile molecule due to other impenetrable molecules blocking its direct route out of the polymer matrix [162]. In this application, it is proposed that a tortuous path is created by the immobile 1.6 kDa DEDEDYFQRYLI dopant molecule, slowing the diffusion of the 466 Da SS. The results indicate that SS diffusing from the PEDOT/DEDEDYFQRYLI/SS is most likely accessed by the cells at a later time period than the media SS control and is reflected in a 34% higher cell density indicative of RGC-5 proliferation prior to differenti-

8.4. DISCUSSION

199

ation. The cell density of PEDOT/DEDEDYFQRYLI/SS is 13,000 cells/cm2 higher than the nominal plating density, concurring with the hypothesis that some cell proliferation occurred following the cell plating. The only other sample where cell proliferation was observed was for Pt where SS was not applied to the system and proliferation continued for the 96 hr incubation period. This assay would suggest that bioactive conducting polymers need to be characterised not only for biofunctionality but also for the possible barrier effects on biomolecule delivery and the resulting impact on cell growth.

Tailoring polymers for specific target tissue is an important consideration in the design of bioactive conducting polymers for neural interfaces. Imparting biofunctionality requires the researcher to focus on a particular application such as cortical recording electrodes, the cochlear implant or peripheral nerve conduits [16, 27, 69, 110, 164, 165]. However, an optimal design will be applicable to all neuroprosthetic applications with incorporation of appropriate biomolecules. Cui et al. found that glial cells and neuroblastomas have affinities for different peptide ligands [35]. Glial cells preferentially bind to PPy incorporating SLPF and neuroblastomas favoured the PPy/CDPGYIGSR composite, demonstrating that specific cells within an application such as cortical recording, can also be targeted. In this assay, a conducting polymer coating designed for the neural model cell PC12 has been modified for the RGC-5 cell line with a view for applying the technology to vision prostheses. The in vitro assay has demonstrated that this design is flexible with PEDOT successfully tailored to promote RGC-5 differentiation. While the system requires further modification for primary RGC survival and in vivo application of a vision prosthesis, this

200CHAPTER 8. TAILORING CP BIOACTIVITY TO AN ALTERNATE CELL TYPE research indicates that alternate neurotrophins or cell signaling factors could be incorporated to produce an interface tailored to a different neuroprosthesis or even passive drug delivery devices.

8.5

Conclusions

Bioactive conducting polymers can be tailored to specific cell lines with biomolecule incorporations being chosen to stimulate responses in targeted neural tissue. This flexibility in polymer constitution demonstrates that these coatings can be applied to a wide range of neuroprosthetic implants. In this case the RGC-5 clone of the primary retinal ganglion cell was differentiated through release of SS incorporated into PEDOT/pTS and PEDOT/DEDEDYFQRYLI during electrodeposition. The differentiation assay confirmed that the SS loaded polymers performed with similar results to the controls where SS was supplied in the media. Additionally, it was shown that the peptide doped PEDOT was able to support higher cell densities with greater neurite length per unit area. This was possibly the result of the active YFQRYLI ligand encouraging cell attachment and neurite growth, but could also be due to the larger dopant slowing the diffusion of the SS into the media through barrier effects. Future studies on primary retinal ganglion cells will establish the true potential of tailoring conducting polymers for specific applications and neural cell types.

Chapter 9

Conclusions and Recommendations 9.1

Conclusions

Neural prostheses ideally have intimate contact between the excitable tissue and the electrode to maintain signal quality and activation of neural cells for the lifetime of the implant. Conducting polymers have been investigated in this role and can potentially enhance tissue/material contact by increasing the electrode surface area and roughness as well as allowing delivery of bioactive signals to neural cells. However, the impact of conventional dopants and biomolecules on long-term electroactivity, mechanical stability and their effect on biological interactions is largely unknown. These attributes are all critical for development of an interface with optimal properties for neural applications.

In the current study, conducting polymers were investigated across two monomer types, pyrrole and EDOT, with conventional sulfonate dopants, pTS and PSS, compared to 201

202

CHAPTER 9. CONCLUSIONS AND RECOMMENDATIONS

laminin peptide dopants, DEDEDYFQRYLI and DCDPGYIGSR. Additionally, the effect of non-doping components, MWNTs and NGF, were assessed. Table 9.1 summarises the results of this thesis showing the relative behavior of each of the materials studied. The properties of each polymer type were compared via a semi-quantitative rating system where 1 = poor; 2 = satisfactory; 3 = good; and 4 = excellent performance across the specific assays used to assess each property. It is clear that modification of conducting polymers for neural interfacing requires assessment across a range of physico-chemical properties and the design of an optimised interface requires a trade-off between these desired electrical, mechanical and biological characteristics. Table 9.1: Summary of results for conducting polymers across Conducting Polymer Performance Polymer Electrical Mechanical PPy/PSS Poor Poor PPy/pTS Poor Satisfactory PEDOT/PSS Good Poor PEDOT/pTS Good Satisfactory MWNT-PPy/PSS Satisfactory Good MWNT-PPy/pTS Satisfactory Good MWNT-PEDOT/PSS Good Poor MWNT-PEDOT/pTS Good Good PEDOT/DEDEDYFQRYLI Satisfactory Satisfactory PEDOT/DCDPGYIGSR Satisfactory Satisfactory PEDOT/pTS/NGF Good Good PEDOT/DEDEDYFQRYLI/NGF Poor Poor PEDOT/DCDPGYIGSR/NGF Poor Poor

assessed criteria Biological Satisfactory Poor Good Good Good Satisfactory Good Good Satisfactory Satisfactory Excellent Satisfactory Satisfactory

Total 4 4 7 8 8 7 7 9 6 6 10 4 4

The best performing material across all attributes was the PEDOT/pTS/NGF with a total score of 10 out of a maximum possible score of 12. This would indicate that biological modification of conducting polymers can be achieved without significant detrimental effects on electrical and mechanical properties, but further research will potentially produce

9.1. CONCLUSIONS

203

more ideal materials for implant application.

Three main hypotheses relating to polymer composition were examined. Firstly, it was proposed that conducting polymers could provide a coating for typical neuroprosthetic metal electrodes which address the issues associated with poor cell interaction without compromising important physico-chemical or biological properties. Next, doping conducting polymers with cell adhesion ligands was investigated in order to understand whether the resulting material would produce an interface where neural cells readily attach to the implant materials without compromising the electrical and mechanical stability of the electrode. Finally, it was propounded that incorporating neurotrophins could produce a bioactive interface to encourage neurite outgrowth from the surrounding cells across the electrode surface, further encouraging interactions between the cell and electrode. All polymers were subjected to a range of analytical techniques designed to characterise polymer electrical, mechanical and biological properties. The flexibility of the design was further explored through application of bioactive conducting polymers to vision prosthesis microelectrode arrays and tailoring polymer biofunctionality for the alternate neural cell type RGC-5.

Initial studies performed on basic conducting polymer combinations commonly seen in literature, PPy and PEDOT doped with pTS and PSS, revealed that PEDOT/pTS was superior to other polymers across the assessed criteria, as shown in Table 9.1. PEDOT/pTS had the highest electrochemical stability and conductivity, good adherence to the Pt substrate and superior cell density and neurite outgrowth when compared to PPy coatings.

204


PEDOT/PSS performed similarly in electrical and biological assays, but was significantly less adherent to the Pt. It was shown that mechanical performance is strongly associated with the polymer dopant size, but additionally this study revealed a dichotomy in conducting polymer mechanical performance not previously described in the literature. All polymers that were strongly adherent to the Pt substrate were substantially harder than the less adherent polymers, this phenomenon was found throughout all successive studies. As a result PEDOT/pTS was found to be harder than both PEDOT/PSS and PPy/PSS, but this was balanced by strong film adherence, integral to long-term performance in an implant. In vitro cell culture demonstrated that all polymers with the exception of PPy/pTS were able to increase cell interactions, determined through attachment and neurite outgrowth, compared to the bare Pt. The materials with the most nodular, rough appearance in SEMs, PEDOT/pTS and PEDOT/PSS, had the greatest degree of cell interaction.

The poor performance of PPy in the initial study was attributed to the mobility of the dopants in the polymer matrix and the relatively featureless surface topography. This result prompted an attempt to increase PPy conductivity and surface roughness through the layering of polymers with MWNTs. The MWNT layered polymer composites were noticeably rougher with most forming large nodules with fibrillar extensions. The impact of MWNT inclusions was in general positive for PPy, but had little effect on the PEDOT. MWNTs rescued PPy/pTS cell interactions and greatly improved PPy electrical properties, however, no result produced from these films was significantly better than those

9.1. CONCLUSIONS

205

established for PEDOT/pTS in previous studies. PEDOT mechanical properties were affected with PEDOT/PSS being extremely fragile and unable to be analysed for conductivity. This study concluded that although MWNTs show promise for neural interfaces with improved electrical properties, PEDOT/pTS was still the most suitable polymer for modification of a typical neural electrode.

The addition of cell attachment ligands in the form of anionic synthetic laminin peptides yielded varied results. Literature has reported that DCDPGYIGSR can be used to dope PPy but analysis of vital electrical and mechanical charactersitics have not been published. Synthetic peptides containing the YFQRYLI ligand have not been used to dope conducting polymers despite research that indicates it has both cell attachment and neurite outgrowth functionality. It was hypothesised that laminin peptides, DEDEDYFQRYLI and DCDPGYIGSR, could increase cell interactions at the electrode surface, but the impact on physico-chemical properties was not forseen. Diffusion limitations during electrodeposition of peptide doped PEDOT limited the efficiency of polymer formation resulting in reduced electrochemical stability and greatly decreased film adherence compared to PEDOT/pTS. The larger DEDEDYFQRYLI peptide was affected the most, and although impedance and softness was improved, without additional laminin coating the cell interactions were not significantly increased. The activity of ligands in the peptides were aimed at cell attachment (YFQRYLI and YIGSR) and neurite outgrowth (YFQRYLI). While these properties were not observed in polymers when laminin coating was not applied prior to cell plating, increases in cell attachment and neurite outgrowth were seen

206


in laminin coated peptide doped PEDOT compared to the PEDOT/pTS control.

An unexpected result from this study was that PEDOT/pTS had substantial cell adherence properties in the absence of a laminin coating. Cell adherence was attributed to the polymer’s nodular surface morphology. In conclusion, the doping of PEDOT with laminin peptides has the potential to increase cell interactions with neuroprostheses, but physicochemical properties are likely to be affected and are directly related to the size of the peptide dopant. A significant outcome is the finding that a dominant factor in conducting polymer performance in a biological environment is surface topography. Modification of surface topography may provide the key to producing a long-term implant conducive to neural tissue integration.

An optimal electrode coating will have the capacity to stimulate both cell attachment and neurite ingrowth to the electrode sites, providing an intimate connection between the implant and neural tissue. The combination of a laminin peptide doped polymer loaded with a neurotrophin has not been reported in the literature. Entrapment of neurotrophic factors in PEDOT was hypothesised to impart the neurite outgrowth component of bioactivity to the electrode coatings. Neurotrophin NGF (specific to PC12 differentiation) was incorporated by entrapment during electrodeposition into PEDOT doped with pTS, DEDEDYFQRYLI and DCDPGYIGSR. While the inclusion of NGF was shown to stimulate some neurite outgrowth, the PEDOT coatings with both a laminin peptide dopant and the non-doping NGF entrapment experienced severely degraded electrochemical and mechanical properties. The NGF molecule was a dominant factor in determining the properties

9.1. CONCLUSIONS

207

of peptide doped PEDOT, with little difference seen between the two peptide types. In vitro neither NGF loaded PEDOT/DEDEDYFQRYLI or PEDOT/DCDPGYIGSR, was able to replicate the cell adherence and neurite outgrowth recorded for their respective controls.

The incorporation of two large biomolecules clearly inhibited the formation of PEDOT, resulting in an inferior coating, however, PEDOT/pTS/NGF exhibited improved properties relative to the peptide doped PEDOT. Where some decrease was seen in long-term electrochemical stability and coating adherence it was offset by lower impedance and hardness values. The smaller dopant molecule provided a more cohesive coating and the surface roughness was again compatible with in vitro cell attachment. The NGF was entrapped in the PEDOT/pTS matrix with minimal evidence of interference in the polymer formation. PEDOT/pTS/NGF demonstrated good cell interaction, with only a small reduction in neurite outgrowth compared to the PEDOT/pTS controls where NGF was supplied via the media.

These studies present a relationship between physico-chemical properties of conducting polymers and their biological performance. Molecules that introduce diffusion limitations during electrodeposition inhibit the formation of a polymer and have a direct impact on polymer properties. For PEDOT a small dopant is desirable and the resulting polymer matrix is able to form and entrap a larger non-doping biofunctional molecule. In this arrangement PEDOT has a surface topography that is conducive to cell attachment without requiring additional coating with cell membrane proteins and the non-doping inclusion

208


can be tailored to stimulate cell ingrowth to the electrode surface. The trade-off between desired electrical, mechanical and biological properties indicates that PEDOT/pTS/NGF will provide a softer more compliant coating for metal electrodes and encourage cell interactions to produce a successful long-term neuroprosthetic interface.

In the second part of this thesis conducting polymers were applied to the AVPG vision prosthesis to assess the performance of these materials as part of a potential commercial device. PEDOT was assessed for performance on microelectrodes intended for implantation and tailored to stimulate differentiation in the biologically appropriate RGC-5 cell line.

Preliminary studies on the AVPG prototype electrode array assessed the possible toxicity imparted to materials during manufacture. Cell growth inhibition of the L929 cell line was used to determine the level of material toxicity arising from processing solvents and laser ablation. The processed electrode array was found to have no additional growth inhibition when compared to unprocessed Pt and medical grade silicone. Subsequently, the optimal conducting polymer, PEDOT/pTS/NGF, was deposited on the electrode array and a PC12 neurite outgrowth assay demonstrated that without any cell attachment proteins, cells preferentially adhere to the PEDOT/pTS/NGF substrate. Additionally, NGF entrapped in the conducting polymer matrix promoted neurite outgrowth from PC12s in the absence of neurotrophic loaded media.

Tailoring PEDOT for RGC-5 differentiation involved replacing the NGF, specific to PC12

9.1. CONCLUSIONS

209

differentiation, with RGC-5 differentiating agent SS. The SS modification was assessed on both PEDOT/pTS and PEDOT/DEDEDYFQRYLI. The study revealed that SS loaded polymers were able to replicate the degree of differentiation seen in the controls where SS was supplied in the media. DEDEDYFQRYLI doped PEDOT was able to support higher cell densities with greater neurite length per unit area than PEDOT/pTS. The apparent increase in the activity of YFQRYLI not seen in previous studies, was due to the combination of RGC-5s being more adherent than PC12s and the SS molecule being significantly smaller than NGF, resulting in a more cohesive polymer substrate. While this system requires further modification prior to use in a commercial vision prosthesis, this study indicates that cell signaling factors can be incorporated into PEDOT to produce an interface tailored for alternate neuroprosthetic applications.

In conclusion, the outcomes of this thesis are:

1. Conducting polymers can provide a coating for typical neuroprosthetic metal electrodes which increase cell interaction with minimal impact on important physicochemical and biological properties.

2. Doping conducting polymers with cell adhesion ligands can provide an interface where some cell types may experience increased attachment but the electrical and mechanical stability of the electrode may be reduced.

3. Incorporating neurotrophins by entrapment during electrodeposition, can produce an interface with bioactivity to encourage neurite outgrowth from the surrounding

210

CHAPTER 9. CONCLUSIONS AND RECOMMENDATIONS cells across the electrode surface, but the polymer constitution must be designed to form a cohesive film with ideal surface topography conducive to cell attachment.

9.2

Recommendations for Future Work

The impact of biological inclusions on polymer properties and their ongoing performance in neural prosthetics requires a greater understanding with future research aimed at controlling and optimising film characteristics for long-term performance. Future research should yield new approaches to processing conducting polymers with novel methods of incorporating biomolecules.

This thesis revealed that assessment of physico-chemical properties of conducting polymers modified for neural cell interactions is integral to producing a long-term implant material. Establishing polymer characteristics requires a number of analytical techniques to assess all aspects of chemical, electrical, mechanical and biological behaviour. Specifically assessment must include the analysis of electrochemical stability, mechanical performance and cell assays that are representative of the cell type and material quantity likely to be present around the implanted device. Future work will involve the integration of new techniques with existing analyses to more accurately predict in vivo material behaviour.

Alternate methods need to be designed for the quantification of polymer chemical constitution, specifically for analysing the incorporation of biological signaling factors within polymer composites. In the literature, quantification of neurotrophin release is performed

9.2. RECOMMENDATIONS FOR FUTURE WORK

211

through radio-labelling. This method is costly and radio-labelling can alter the structure of the biomolecule, potentially changing the way it interacts with the polymer matrix. More advanced techniques such as Fourier transform infrared microscopy with mapping (FTIR-MM) can provide in vitro information on the chemical elements present on a material surface and their bonding within that environment. The peak size is proportional to the available quantity, providing valuable information on the material constitution accessible to the surrounding cellular environment.

Electrochemical behaviour determined through cyclic voltammetry is a vital electrical analysis, providing information on reduction-oxidation potentials, potential charge storage density and long-term stability of these properties. Future studies should be coupled with intermittent stimulation of materials through in vivo implant specific stimulation regimes to give a more accurate indication of conducting polymer electrochemical stability across an implant lifetime.

Mechanical analyses in this thesis have clearly demonstrated that these properties vary significantly with polymer composition. The type of monomer, size of dopant and other non-doping inclusions greatly effect mechanical performance. In these studies ASTM tests for hardness and adherence revealed the variation in these film properties and their relationship to polymer composition, however, it was recognised that these tests did not provide enough information. The hardness test allowed comparisons to be drawn between polymers but did not give a definitive measure of hardness relative to the intended neural tissue application. The x-cut adherence test rarely showed delamination along the film

212


disruption, but rather poor adherence across the entire film. This would suggest that a different test without the x-cut would provide more information about the polymer adherence. A method that involves quantification of applied forces would allow comparisons to be made according to intended application, such as in vivo forces likely to be applied at the electrode interface. Nanoindentation is an analytical technique that will be used in the future to provide a more thorough picture of conducting polymer properties.

Finally, this research revealed that the surface topography of a polymer is integral to its in vitro performance. Polymer roughness is also implicated in the impedance behaviour of any conductor and polymers with nodular surfaces have been shown to have lower impedance magnitudes and phase lags in the low frequency region as a result of the increased surface area. However, analysis of soft, brittle conducting polymers is difficult using common techniques. Stylus profilometry is unacceptable as the stylus scrapes the polymer surface, removing material and clogging the stylus point. Atomic force microscopy (AFM) can be useful in imaging surface topography and measuring roughness of less nodular films, but the high degree of topographical variation observed in polymers with biomolecule modifications is beyond the capabilities of most AFM set ups. It is proposed that optical profilometer technology would be beneficial as it allows the user to build a three dimensional image of the polymer surface and can be used to determine both surface roughness and film thickness. Greater insight into the relationship between surface topography and cell interactions will be essential in designing electrodes for neuroprosthetics.

9.2. RECOMMENDATIONS FOR FUTURE WORK

213

While this research demonstrates that polymer topography is a major contributing factor in cell attachment, specific ligands have also been shown to increase attachment and neurite growth. An optimal electrode coating would present cell attachment ligands at the surface, incorporated in such a way that the connection was stable for the implant lifetime. Neurotrophic agents would be included in a manner that allows controlled release from the implant over a substantial period to stimulate and maintain cell ingrowth to the electrode sites. Animal studies to determine in vivo responses to these materials will also be an important part of ongoing research along with the effect of stimulation regimes commonly used in neuroprosthetic implants.

Continuing research will explore cell attachment properties of conducting polymers through both surface topography modification and peptide incorporation. Doping PEDOT with smaller specific peptides may produce a material with both optimal surface topography and increased biofunctionality. Alternately, the covalent tethering of peptides on the surface of conducting polymers may provide an environment conducive to long-term integration of neural tissue at the electrode interface.

Future work on cell ingrowth will look at incorporating signaling factors required for the in vivo application of conducting polymers, specifically to the vision prosthesis. The incorporation of more complex combinations of neurotrophic agents including BDNF, CNTF and forskolin, known to promote the survival of primary RGCs, will be explored. Alternate methods for delivering these signaling factors to cells including the use of hydrogel composites and biodegradable carrier components should be examined. It is important

214


to note that none of the cell studies in the thesis were done under electrical stimulation and that this may have an effect on the cell response. Although not within the scope of this thesis, these studies would be an important component of future work.

Bioactive conducting polymer coatings have the potential for improving the long-term performance of neuroprosthetic devices. Cell interactions can be encouraged through the coating of conventional metal electrodes with conducting polymers tailored for specific implant applications. The rigorous characterisation of potential conducting polymers across intgeral physico-chemical and biological properties is required for the development of a commercially viable implant component.

Appendices

215

Appendix A

Effect of Dopant Concentration on Electrochemical Stability A.1

Introduction

Electrochemical stability was identified in the literature as a major contributing factor to the long-term performance of conducting polymers for neuroprosthetics. As such this was the first analytical technique performed on all polymer types and a pilot study was used to determine the optimum doping concentration for the monomer electrolyte solutions. The doping pilot study was performed with PSS as the dopant and both pyrrole and EDOT monomers. PSS is a polymer but each monomeric unit contains a sulfonate compound capable of contributing an electron to charge conduction. As a result molarity of PSS solutions are calculated with respect to the monomeric unit and the optimal concentration determined in this study will adjusted for alternate dopants relative to their intrinsic net charge. 216

A.2. MATERIALS AND METHODS

A.2

217


Basic monomer solutions were made up in a glass vial as indicated in Table A.1. All products were obtained from SigmaAldrich and pyrrole was vacuum distilled prior to each use to remove polymerised contaminants that reduce efficiency of electrodeposition. EDOT was found to have a longer shelf-life and was vacuum distilled on a monthly basis to remove polymerised components. Aliquots of 150 μL of polymer solution were placed in each well with a Pt counter electrode. The working electrode was formed by the Pt or ITO substrate and electrically connected by an alligator clip. Polymers were electrodeposited using an in-house manufactured galvanostat (detailed in Appendix B) at 1 mA/cm2 or 300 μA per well for 10 minutes. Following deposition the deposition electrolyte was removed. Each film was washed with DI water four times. The final aliquot was left on the film for 24 hours in a 37 ◦ C incubator to leach any excess monomer, dopant or process contaminants.

Table A.1: Basic Component Pyrrole EDOT PSS Deionised (DI) Water Acetonitrile

polymer solutions with varied dopant concentration Cat # PPy/PSS PEDOT/PSS 131709 0.1M 483028 0.1M 243051 0.005M, 0.05M, 0.1M 0.005M, 0.05M, 0.1M N/A 100% 50% 271004 50%

PPy and PEDOT films doped at 0.005M, 0.05M and 0.1M were analysed by CV to determine electrochemical stability. An eDAQ potentiostat and eCorder unit coupled with the supplied EChem software package (eDAQ Pty Ltd., Australia) was used to apply

218APPENDIX A. EFFECT OF DOPANT CONCENTRATION ON ELECTROCHEMISTRY a cycling voltage from -800 mV to 600 mV. This range was chosen to allow the oxidation and reduction of the polymer within the limits of the water window, to prevent hydrolysis occurring at the electrodes. The scan rate was set at 120 mV/s for 400 continuous cycles and measurements were performed in 0.9% saline. The recordings were made with an isolated Ag/AgCl reference electrode and Pt counter electrode. The first stable curve was considered Cycle 1 and used to determine the original electroactivity of the film. The area contained within the oxidation-reduction curve was calculated to establish the current carrying capacity of the film. The area contained within the curve of successive cycles was calculated as a percentage of Cycle 1 to show the comparative loss of electroactivity over time. This process was repeated and the mean curve for electroactivity loss was calculated.

A.3

Results and Discussion

PSS doped polymers were all subjected to 400 cycles of CV and the resulting curve is presented in Figure A.1 with values normalised to the first curve for each material. Doping both PPy and PEDOT with 0.005M PSS resulted in higher losses of electroactivity than the other dopant concentrations. There was no significant difference between 0.1M and 0.05M PSS for either polymer.

It is thought that the 0.005M PSS is an insufficient concentration of dopant for effective polymerisation from the electrolyte solution, resulting in a material with poor electrochemical stability. Optimal electrochemical performance appears to occur with 0.05M

A.4. CONCLUSIONS

219

Figure A.1: Electroactivity loss related to dopant concentration for PSS in PPy and PEDOT PSS with only a small increase observed when the higher concentration of 0.1M is used. 0.05M PSS provides sufficient dopant to create an electrochemically balanced polymer. Since the amount of dopant included in a polymer is determined from the backbone chemistry and one dopant molecule balances approximately three monomer units, the provision of excess dopant, as in the case of 0.1M PSS, is not beneficial.

A.4

Conclusions

A minimal concentration of dopant is desired to reduce the quantity of synthetic materials placed in a biological environment. As such the optimal concentration of 0.05M PSS will be used in fabrication of conducting polymers for neural interfaces. This concentration will also be used for other dopants with respect to their net charge. Alternate dopant

220APPENDIX A. EFFECT OF DOPANT CONCENTRATION ON ELECTROCHEMISTRY pTS, has an identical net charge to a monomer unit of PSS, due the presence of the sulfonate compound. Both conventional dopants will be supplied at 0.05M in the electrolyte solution.

Appendix B

Specifications for In-House Manufactured 12-Channel Galvanostat

This appendix describes the operation of a simple 12-channel galvanostat/biphasic stimulator. This device was purpose built by Phil Preston at GSBmE to be used in an electropolymerisation system, where polymer is coated into platinum in a small well system.

The device does two jobs:

1. Acts as a galvanostat to provide electrodeposition at a constant current

2. Provides biphasic currents to release neurotrophins from conducting polymers

The printed circuit board consists of 4 sections: 221

222

APPENDIX B. SPECIFICATIONS FOR 12-CHANNEL GALVANOSTAT • A simple microcontroller (AVR mega168) that provides control • 3 banks of 4-channel current sources

It is powered by an AC plug-pack. This is design decision was driven by the need to take the device to the United States, where it was relatively simple to replace the 240 V/50 Hz plug pack with a US style 120 V/60 Hz plug pack. The device runs from a 12V AC plug pack, rated at 500 mA.

The device shown in Figure B.1 features three LEDS: Green is power; red and yellow are under software control. It provides current through a 12-wire ribbon cable and a single black alligator clip returns current. This clip goes to the ITO slide or platinum substrate. The ribbon cable connects to a set of platinum probe cards that are custom manufactured to fit the Flexiperm 12-well silicone gasket. The current control knob on the front allows the operator to set the current manually, and a software read-out is visible on the screen. Initially this read-out was calibrated with the use of an oscilloscope.

A serial communications link with the device from a desktop computer allows the user to set both the constant current and duration of the electrodeposition. Simple commands instruct the device to begin polymerisation and when the set duration is reached the current is terminated.

The printed circuit board was designed using Eagle-CAD. The separate sections were then combined into a single layout and etched in-house using the toner transfer method.

223

Figure B.1: 12-channel in-house manufactured galvanostat for concurrent electrodeposition of conducting polymers on 12 electrodes in custom sandwich assembly.

Appendix C

Effect of NGF concentration on PC12 Neurite Outgrowth C.1

Introduction

The concentration of NGF required to promote neurite outgrowth from GFP transfected PC12s was determined through a titration assay on TCP. Although subsequent cultures were performed in smaller silicone wells adhered to novel substrates, the TCP provides an optimal environment in which no other material interaction affects cell response.

C.2


Sterile 12-well TCP plates (Cat #. 353043, BD Falcon, USA) were coated with 5 μg/mL laminin sourced from Engelbreth murine sarcoma (Sigma Aldrich, L2020) in Dulbecco’s phosphate buffered saline (DPBS). Wells were incubated for 2 hr at 37◦ C then stored overnight at 4◦ C. On the morning before plating the coated wells were washed three 224

C.2. MATERIALS AND METHODS

225

times with sterile DPBS and air dried in a PCII laminar flow hood.

To plate cells, flasks of GFP transfected PC12s (Marinpharm GmbH Cat #. PC-TGPC-12) grown in Roswell Park Memorial Institute medium (RPMI) supplemented with 10% horse serum (HS) and 5% fetal calf serum (FCS), were pipetted into polypropylene centrifuge tubes and allowed to settle for 5 min. When a cell clump was visible the supernatant was removed, leaving approximately 0.5 mL of media. A 1 mL Eppendorf pipette tip was used to add 500 μL of fresh media to the cell suspension. Cells were aspirated gently using the 1 mL pipette tip to break up the clumps through 40 aspiration repetitions. Low serum RPMI with 1% horse serum was used to dilute the cell solution to 5 mL and a 0.5 mL sample was taken for counting. Cells were plated at 20,000 cells/cm2 and cultured for 96 hr in low serum medium supplemented with NGF (NGF2.5 Grade 1 from Alomone Laboratories) at the various concentrations depicted in Figure C.1. All cell studies were carried out in an incubator at 37◦ C, 5% carbon dioxide (CO2 ) and 100% humidity. Low serum media was renewed with NGF supplement at 48 hr by exchanging 2/3 or 1.3 mL of media in each well.

PC12s on TCP were analysed using phase contrast microscopy with images captured on a Carl Zeiss camera attachment. The lengths of individual neurites were traced and calibrated length determined using NIH software Image J, with Neuron J plugin. Neurites were measured when the length of the projection exceeded a single body length of the cell from which it extended. When neurites branched, one split was considered a continuation of the primary neurite and each alternate split of longer than a cell body in length was

226APPENDIX C. EFFECT OF NGF CONCENTRATION ON NEURITE OUTGROWTH

Figure C.1: Concentrations of NGF in a 12-well plate, used to determine optimum concentration for GFP-PC12 neurite outgrowth

considered a new neurite.

For each experiment, each NGF concentration was run in three duplicate wells. For each well three micrographs were taken by scanning the film such that each image was obtained from a different sector (third) of the well area. The cell density and neurite density supported by each substrate is presented for two repetitions as mean ± standard error.

C.3


The GFP transfected PC12s were analysed at 96 hr and shown to grow in response to an increasing concentration of NGF, graphed in Figure C.2. No growth was seen on negative controls where no NGF was in the media. Cell density was significantly better and neurite outgrowth was recorded at 10 ng/mL NGF. Further increases in both cell

C.3. RESULTS AND DISCUSSION

227

density and neurite outgrowth were seen at 50 ng/mL. Cell density at 100 ng/mL NGF was no different to that recorded for 50 ng/mL and while neurite length supported per cm2 increased, it was not significantly better.

Figure C.2: GFP-PC12 cell density and neurite outgrowth at various concentrations of NGF on TCP at 96 hr, (n=2).

These values correlate with experiments previously performed at GSBmE on non-fluorescent PC12s grown under similar conditions. A minimal concentration of NGF will be used to promote optimal neurite outgrowth in all PC12 assays performed as part of this thesis. As such 50 ng/mL will be applied as the media supplement for neurite outgrowth assays on novel substrates.

228APPENDIX C. EFFECT OF NGF CONCENTRATION ON NEURITE OUTGROWTH

C.4

Conclusions

TCP assays performed on GFP tranfected PC12s indicate that they differentiate in response to NGF in the same manner as their non-transfected predecessors. The optimal concentration for NGF was found to be 50 ng/mL of media.

Appendix D

Statistical analysis of In Vitro Cell Studies

Prepared by Ross Odell, PhD, Lecturer at the Graduate School of Biomedical Engineering, UNSW.

D.1

Design of Analysis

The data were obtained in several experiments investigating the effect of material (M) with 4 or 5 levels, which generally comprised a monomer and dopant, crossed with another factor (A) at 2 levels, such as coating with laminin or not, on the number of cells and the total length of neurites. Measurements were made on 3 images in each of 3 wells for each combination of the experimental factors and 3 replicates (S) were conducted of each experiment. The design is illustrated in Table D.1.

This can also be arranged as 3 replicates of a 1D array of 8 cells as in Table D.2. 229

230

APPENDIX D. STATISTICAL ANALYSIS OF IN VITRO CELL STUDIES

Table D.1: A two-way Replication 1 m1 m2 m3 a1 9 9 9 a2 9 9 9

a1m1 9

a1m1 9

a1m1 9

a1m2 9

a1m2 9

a1m2 9

layout m4 9 9

a1 a2

Replication 2 m1 m2 m3 9 9 9 9 9 9

m4 9 9

a1 a2

Replication 3 m1 m2 m3 9 9 9 9 9 9

m4 9 9

Table D.2: A one-way layout Replication 1 a1m3 a1m4 a2m1 a2m2 a2m3 9 9 9 9 9

a2m4 9

a1m3 9

Replication 2 a1m4 a2m1 9 9

a2m2 9

a2m3 9

a2m4 9

a1m3 9

Replication 3 a1m4 a2m1 9 9

a2m2 9

a2m3 9

a2m4 9

D.1. DESIGN OF ANALYSIS

231

When the experiment is replicated, there will be variation between the observations for each cell of the design. The variation will arise from the measurement error associated with the 9 observations within a given cell in one replicate and an experimental error arising from variations in experimental conditions between replicates.

The analysis of variance table for the two-way layout will have the form illustrated in Table D.3 [166]. The within-cell mean square (MSE) will be referred to as the measurement error and the interactions of A, M and AM with replicate will be referred to as experimental errors. The F-ratio for testing the effect of a factor has, in the denominator, the mean square for the interaction of that factor with replicate and therefore has a relatively small number of degrees of freedom. As a result, the power of the test is reduced.

Table D.3: Analysis of variance table for the two-way layout. df Mean Square F-ratio A 1 MSSA MSA /MSA×Rep M 3 MSM MSM /MSM ×Rep AM 3 MSAM MSAM /MSAM ×Rep Rep 2 MSRep A × Rep 2 MSA×Rep M × Rep 6 MSM ×Rep AM by Rep 6 MSAM ×Rep Within cell 8∗24 = 192 MSE Total 215

If the three mean squares for the interactions with replicate are similar, they might be viewed as three estimates of the same experimental error and they might therefore be pooled as in Table D.4 in order to increase the denominator degrees of freedom and the power.

232


Table D.4: Analysis of variance table for the one-way layout with pooled errors. df Mean Square F-ratio Between cells 23 Treatment 7 MST MST / MST ×R Rep 2 MSR Treat × Rep 14 MST ×R Within cell 8∗24 = 192 MSE Total 215

An alternative method of analysis, which can be used to compare the two levels of factor A, is to calculate the difference between the mean results for level A1 and A2 for each material and replicate and to analyse these as a randomised block experiment as in Table D.5. In this analysis, the M effect corresponds to the AM interaction in Table 3 and the M Rep interaction is the error term for testing the effect of M. Confidence intervals for the true difference between A1 and A2 can be calculated as described by Zar [167]. The limits of a 95% confidence interval are calculated from Equation D.1.

s x ± t0.05,v √ n

(D.1)

Where x is the mean difference, s is the error mean square from the ANOVA table, v is the error degrees of freedom and n is the number of replicates.

Table D.5: Analysis of variance table for the randomised block design. df Mean Square F-ratio M 3 MSM MSM / MSM ×R Rep 2 MSR M × Rep 6 MSM ×R Total 11

D.2. PRELIMINARY DATA TREATMENT

D.2

233

Preliminary Data Treatment

The data were analysed in summarized form. The mean cell number and total neurite length were calculated for each set of conditions and further analysis was performed on these. Since there were the same number of observations, n = 9, for each case, the analysis was exactly equivalent to that illustrated in Table D.3.

The number of cells and, especially, the total neurite length, varied over a decade or more, often an indication that a variance-stabilising transformation might be needed. Examination of the residual indicated that analysis of untransformed cell numbers was satisfactory but that for neurite length, the power transformation z = y0.25 gave a superior distribution of residuals [168].

Analyses were performed with Minitab 15 (Minitab Inc. 2007) using the GLM command and the Tukey option for multiple comparisons, when appropriate.

D.3

ANOVA for Basic Polymers

This study examined two monomers (PPy and PEDOT) and two dopants (PSS and pTS) along with a platinum control. These were expressed as a single factor with 5 levels.

1. Platinum

2. PPy/PSS

234


3. PPy/pTS 4. PEDOT/PSS 5. PEDOT/pTS

D.3.1

Cell Density

The analysis of variance performed on the summarized data, Table D.6, indicates highly significant differences among materials (MM), p < 0.001. Table D.6: Analysis of variance table for mean cell number, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P MM 4 17140.9 17178.0 4294.5 37.66 0.000 RR 2 456.1 456.1 228.1 2.00 0.206 Error 7 798.3 798.3 114.0 Total 13 18395.3 S = 10.6791 R-Sq = 95.66% R-Sq(adj) = 91.94%

Table D.7 shows the mean differences in the number of cells between materials (row material - column material) with the p-value in brackets below (based on the Tukey multiple comparison procedure). For example, the difference between materials 2 (PPy/PSS) and 4 (PEDOT/PSS) was 44.9 and was highly significant (p = 0.008).

D.3.2

Neurite Length

The neurite length data were analysed after transformation. Table D.8 indicates that there were highly significant differences among the materials (M), p < 0.0001.

D.3. ANOVA FOR BASIC POLYMERS

235

Table D.7: Multiple comparison matrix for mean cell density (based on the Tukey multiple comparison procedure) 1 2 3 4 2 5.1 -0.983 3 -44.7 -49.9 -0.017 -0.005 4 50 44.9 94.7 -0.009 -0.008 0 5 42.5 37.3 87.2 -7.5 -0.022 -0.021 0 -0.901 Table D.8: Analysis of variance table for neurite length (transformed), using adjusted SS for tests Source DF Seq SS Adj SS Adj MS F P MM 4 47.4503 46.6635 11.6659 39.44 0.000 RR 2 2.6983 2.6983 1.3492 4.56 0.054 Error 7 2.0707 2.0707 0.2958 Total 13 52.2194 S = 0.543891 R-Sq = 96.03% R-Sq(adj) = 92.64% The matrix of comparisons among materials, based on the Tukey multiple comparison procedure (Table D.9) indicates that the difference between materials 2 (PPy/PSS) and 4 (PEDOT/PSS) was not significant, p = 0.288. Table D.9: Multiple comparison matrix for basic neurite length 1 2 3 4 2 1.075 -0.311 3 -2.674 -3.75 -0.007 0 4 2.042 0.967 4.717 -0.029 -0.288 0 5 2.121 1.046 4.796 0.079 -0.024 -0.232 0 -1

236

D.4


ANOVA for MWNT Composite Polymers

In the MWNT study there were 5 materials (M), both as homogeneous polymers and as composites with nanotubes (MWNTs). The materials were

1. Platinum 2. PPy/PSS 3. PPy/pTS 4. PEDOT/PSS 5. PEDOT/pTS

The main purpose of the study was to compare the materials with and without the MWNTs (factor C = “MWNT” vs. C = “Basic”).

D.4.1

Cell Density

The cell number data were analysed without transformation. The ANOVA table (Table D.10 indicates a significant MM × CC interaction; in other words, the difference between “MWNTs” and “Basic” was not the same for all materials. The difference between homogeneous polymer (CC = “MWNTs”) and the corresponding composite (CC = “Basic”) was analysed as a randomized block design (Table D.11) in order to compute confidence intervals for the difference. Confidence intervals for the difference are given in Table

D.4. ANOVA FOR MWNT COMPOSITE POLYMERS

237

Table D.10: Analysis of variance table for mean cell density of MWNT layered polymers, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P RR 2 45.8 45.8 22.9 0.09 0.912 MM 4 23683.5 23683.5 5920.9 23.99 0.000 CC 1 804.0 804.0 804.0 3.26 0.088 MM∗CC 4 3039.8 3039.8 759.9 3.08 0.043 Error 18 4442.6 4442.6 246.8 Total 29 32015.7 S = 15.7103 R-Sq = 86.12% R-Sq(adj) = 77.64%

D.12. The confidence interval for material 3 (PPy/pTS) does not include zero, indicating a significant difference, p < 0.05.

Table D.11: Analysis of variance table for the difference in mean cells densitied between MWNT composite and basic polymers, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P M 4 6079.6 6079.6 1519.9 7.36 0.009 R 2 869.7 869.7 434.8 2.11 0.184 Error 8 1651.2 1651.2 206.4 Total 14 8600.4 S = 14.3665 R-Sq = 80.80% R-Sq(adj) = 66.40%

Table D.12: Comparison matrix for cell densities in MWNT layering study; including the difference between coated and uncoated, the standard error of the difference, and lower and upper limits of a 95% confidence interval for the difference. Cell Density 95% Confidence Limits M Basic MWNTs diff se(diff) Lower Upper 1 78.9 74.5 -4.4 8.29 -23.5 14.7 2 78.9 96.4 17.5 8.29 -1.6 36.6 3∗ 21.7 65.2 43.5 8.29 24.4 62.6 4 127.4 112.1 -15.3 8.29 -34.4 3.8 5 111.6 121.9 10.3 8.29 -8.8 29.4

238

D.4.2


Neurite Length

The analysis of variance on the transformed neurite length data is presented in Table D.13. It indicates a highly significant material × MWNT interaction (MM∗CC). For this reason a blanket statement that MWNTs improve or degrade performance is not justified; it depends on the material.

Table D.13: Analysis of variance table using adjusted SS for tests. Source DF Seq SS RR 2 2.2621 MM 4 51.7983 CC 1 2.6516 MM∗CC 4 9.0887 Error 18 3.8601 Total 29 69.6609

for neurite length on MWNT layered polymers, Adj SS 2.2621 51.7983 2.6516 9.0887 3.8601

Adj MS 1.1311 12.9496 2.6516 2.2722 0.2145

F 5.27 60.38 12.36 10.60

P 0.016 0.000 0.002 0.000

The analysis of the difference between materials with and without MWNTs are presented in Table D.14 and the calculated confidence intervals in Table D.15. Only the confidence interval for material 3 (PPy/pTS ) does not include zero, indicating a significant positive effect of MWNTs for this material, p < 0.05.

Table D.14: Analysis of variance of the difference in neurite length on MWNT layered polymers, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P M 4 18.1774 18.1774 4.5443 14.68 0.001 R 2 0.1484 0.1484 0.0742 0.24 0.792 Error 8 2.4766 2.4766 0.3096 Total 14 20.8023 S = 0.556391 R-Sq = 88.09% R-Sq(adj) = 79.17%

D.5. ANOVA FOR LAMININ PEPTIDE DOPED PEDOT

239

Table D.15: Comparison matrix for neurite lengths (transformed) for materials with and without MWNTs; including the difference between coated and uncoated, the standard error of the difference, and lower and upper limits of a 95% confidence interval for the difference. Neurite Lengths 95% Confidence Limits M Basic MWNTs diff se(diff) Lower Upper 1 5.03 4.46 -0.57 0.32 -1.31 0.17 2 5.65 6.22 0.57 0.32 -0.17 1.31 3* 2.19 4.84 2.65 0.32 1.91 3.39 4 6.88 6.79 -0.09 0.32 -0.83 0.65 5 6.78 7.19 0.41 0.32 -0.33 1.15

D.5

ANOVA for Laminin Peptide Doped PEDOT

In this study there were 4 materials which were tested with and without a laminin coating.

1. Platinum

2. PEDOT/pTS

3. PEDOT/DEDEDYFQRYLI

4. PEDOT/DCDPGYIGSR

D.5.1

Cell Density

The ANOVA results in Table D.16 indicate that the material × coating interaction (MM∗LL) was highly significant (p = 0.003). That is, the magnitude of the effect of coating with laminin was different for different materials. The difference between coated and uncoated was smaller for material 2 (PEDOT/pTS) than for the other materials.

240


Table D.16: Analysis of variance table for cell density on peptide adjusted SS for tests Source DF Seq SS Adj SS Adj MS F RR 2 380.3 380.3 190.1 0.14 MM 3 7295.2 7295.2 2431.7 2.23 LL 1 8271.7 8271.7 8271.7 22.27 MM∗LL 3 2080.0 2080.0 693.3 10.94 RR∗MM 6 6533.0 6533.0 1088.8 17.18 RR∗LL 2 743.0 743.0 371.5 5.86 Error 6 380.2 380.2 63.4 Total 23 25683.4 S = 7.96073 R-Sq = 98.52% R-Sq(adj) = 94.32%

doped PEDOT, using P 0.875 x 0.185 0.042 0.008 0.002 0.039

Further investigation of the difference between coated and uncoated materials is complicated by the fact that the three estimates of the experimental error, RR∗MM, RR∗LL and Error are quite different; the differences among materials (MM) varied much more between replicates (RR) than did the differences between the coated and uncoated (LL). Pooling these error terms would give an inflated error with respect to comparisons between coated/uncoated.

Instead, the difference in cell numbers between coated and uncoated materials was calculated for each study and these were then analysed as a randomized block experiment [167]. The table below (Table D.17 gives the mean difference, the standard error of the mean difference, and upper and lower 95% confidence limits for the difference. For material 2 the 95% confidence interval does not include zero, indicating that the difference between coated and uncoated was not significant at the 5% level for this material.

D.5. ANOVA FOR LAMININ PEPTIDE DOPED PEDOT

241

Table D.17: Comparison matrix for cell densities in peptide doping study. Cell Density 95% Confidence Limits M Uncoated Coated Diff. SE(Diff) Lower Upper 1 37.7 69.6 32 6.5 16.1 47.9 2 94.7 103.6 8.9 6.5 -7 24.8 3 45.3 98.4 53.1 6.5 37.2 69 4 62.5 117.1 54.6 6.5 38.7 70.5

D.5.2

Neurite Length

The neurite lengths were transformed before analysis and the data for the platinum control were excluded before analysis. In Table D.18, the experimental errors have been pooled. There is an suggestion, not quite significant, of a material × coating interaction (p = 0.076) and a clear effect of laminin coating (LL). Table D.18: Analysis of variance table for neurite length on peptide doped PEDOT (transformed) using two-way method and adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P RR 2 2.9264 2.9264 1.4632 9.97 0.004 MM 2 1.1936 1.1936 0.5968 4.07 0.051 LL 1 22.6845 22.6845 22.6845 154.59 0.000 MM∗LL 2 0.9897 0.9897 0.4948 3.37 0.076 Error 10 1.4674 1.4674 0.1467 Total 17 29.2615 S = 0.383061 R-Sq = 94.99% R-Sq(adj) = 91.48%

In order to compare the 3 laminin-coated materials, the data were re-analysed in the oneway layout of Table D.2 with resultes recorded in Table D.19. The factor BB, with levels 1 to 8, corresponds to the numbering from left to right of the cells in Table D.2. Thus the 3 coated materials correspond to BB = 6,7,8. The multiple comparison matrix in Table D.20 indicates that there was a marginally significant difference (p = 0.04) between laminin-coated PEDOT/pTS and laminin-coated PEDOT/DCDPGYIGSR (BB=6 and

242


BB=8). Table D.19: Analysis of variance table for neurite length on peptide doped PEDOT (transformed) using one-way ANOVA and adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P RR 2 2.9264 2.9264 1.4632 9.97 0.004 BB 5 24.8677 24.8677 4.9735 33.89 0.000 Error 10 1.4674 1.4674 0.1467 Total 17 29.2615 S = 0.383061 R-Sq = 94.99% R-Sq(adj) = 91.48% Table D.20: Multiple comparison matrix for neurite length in peptide doped PEDOT study(based on the Tukey multiple comparison procedure). Gives the difference (row number - column number) and p-value for pairwise comparisons. 2 3 4 6 7 3 0.071 -1 4 0.038 -0.033 -1 -1 6 1.597 1.526 1.559 -0.005 -0.006 -0.005 7 2.519 2.448 2.481 0.922 0 0 0 -0.11 8 2.729 2.658 2.691 1.132 0.21 0 0 0 -0.04 -0.982

D.6

ANOVA for NGF loaded PEDOT

In the NGF study there were four materials (M).

1. Platinum 2. PEDOT/pTS 3. PEDOT/DEDEDYFQRYLI 4. PEDOT/DCDPGYIGSR

D.6. ANOVA FOR NGF LOADED PEDOT

243

The were tested under two conditions, with NGF incorporated into the material or with NGF added to the media. The main interest was to determine whether adding NGF to the polymer (factor N = “Polymer”) improved performance to equal the performance with NGF in the media (factor N = “Media”).

D.6.1

Cell Density

The analysis of variance of cell numbers(Table D.21) indicates that there was no significant MM∗NN interaction. In other words, there is no strong indication that the effect of adding NGF to the polymer was different for the different polymers. For this reason, testing for the effect of NGF with individual polymers is not warranted. The conclusions from this analysis are that adding NGF to the media as opposed to the polymer improved performance (p < 0.001). Table D.21: Analysis of adjusted SS for tests Source RR MM NN MM∗NN Error Total

D.6.2

variance table for cell density on NGF loaded polymers, using DF 2 3 1 3 14 23

Seq SS 1044.1 8690.3 9713.7 1527.0 3649.7 24624.8

Adj SS 1044.1 8690.3 9713.7 1527.0 3649.7

Adj MS 522.1 2896.8 9713.7 509.0 260.7

F 2.00 11.11 37.26 1.95

P 0.172 0.001 0.000 0.168

Neurite Length

The analysis of variance for the transformed lengths is shown in Table D.22. The MM∗NN interaction is highly significant.

244


Table D.22: Analysis of variance table for neurite length on NGF loaded PEDOT (transformed), using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P SS 2 1.5167 1.5167 0.7584 5.02 0.023 MM 3 46.4272 46.4272 15.4757 102.39 0.000 NN 1 29.2298 29.2298 29.2298 193.40 0.000 MM*NN 3 8.7729 8.7729 2.9243 19.35 0.000 Error 14 2.1160 2.1160 0.1511 Total 23 88.0626 S = 0.388767 R-Sq = 97.60% R-Sq(adj) = 96.05%

Analysis of variance for the difference in neurite length between NN = “Polymer” and NN = “Media” is shown in Table D.23 and the resulting 95% confidence intervals are shown in Table D.24. The confidence interval for PEDOT/pTS includes zero, indicating that for this polymer the difference is not significant.

Table D.23: Analysis of variance table for the difference in neurite length on NGF loaded PEDOT, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P M 3 17.5458 17.5458 5.8486 24.53 0.001 S 2 1.0494 1.0494 0.5247 2.20 0.192 Error 6 1.4307 1.4307 0.2385 Total 11 20.0259 S = 0.488320 R-Sq = 92.86% R-Sq(adj) = 86.90%

Table D.24: Mean transformed neurite lengths and 95% confidence limits for the difference between NGF in the polymer and NGF in the media. Neurite Lengths 95% Confidence Limits M Basic MWNTs diff se(diff) Lower Upper 1 3.87 2.93 -0.94 0.28 -1.6 -0.3 2 5.01 4.51 -0.5 0.28 -1.2 0.2 3 4.7 3.77 -0.93 0.28 -1.6 -0.2 4 5.1 3.6 -1.5 0.28 -2.2 -0.8

D.7. ANOVA FOR PEDOT/PTS/NGF COATED AVPG ELECTRODE ARRAY 245

D.7

ANOVA for PEDOT/pTS/NGF coated AVPG electrode array

This was a single factor study comparing a bare electrode with an electrode coated with PEDOT/pTS/NGF. There were 3 replicates, each with 3 wells and 3 images/well.

The analysis of variance for cells and neurite lengths are in Tables D.25 and D.26, respectively. In both cases, the effect of coating the electrode (the factor EE) was significant despite the small number of replicates. Table D.25: Analysis of variance for mean cell densities on AVPG electrodes, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P RR 2 164.6 164.6 82.3 2.56 0.281 EE 1 3246.0 3246.0 3246.0 101.12 0.010 Error 2 64.2 64.2 32.1 Total 5 3474.8 S = 5.66576 R-Sq = 98.15% R-Sq(adj) = 95.38% Table D.26: Analysis of variance for mean neurite lengths on AVPG electrodes, using adjusted SS for tests. Source DF Seq SS Adj SS Adj MS F P RR 2 0.0124 0.0124 0.0062 0.03 0.971 EE 1 29.5950 29.5950 29.5950 141.07 0.007 Error 2 0.4196 0.4196 0.2098 Total 5 30.0269 S = 0.458030 R-Sq = 98.60% R-Sq(adj) = 96.51%

Appendix E

Effect of Peptide Concentration on Electrochemical Stability E.1

Introduction

Electrochemical stability was identified in the literature as a major contributing factor to the long-term performance of conducting polymers for neuroprosthetics. As such this was the first analytical technique performed on all polymer types and a pilot study was used to determine the optimum doping concentration for laminin peptides in the monomer elecrtrolyte solutions. Two concentrations were assessed with the higher being obtained from studies by Cui et al. [35]

E.2


Basic monomer solutions were made up in a glass vial as indicated in Table E.1. Aliquots of 150 μL of polymer solution were placed in each well with a Pt counter electrode. The 246

E.2. MATERIALS AND METHODS

247

working electrode was formed by the Pt or ITO substrate and electrically connected by an alligator clip. Polymers were electrodeposited using an in-house manufactured galvanostat (detailed in Appendix B) at 1 mA/cm2 or 300 μA per well for 10 min. Following deposition the deposition electrolyte was removed. Each film was washed with DI water four times. The final aliquot was left on the film for 24 hr in a 37 ◦ C incubator to leach any excess monomer, dopant or process contaminants. Table E.1: PEDOT/peptide electrolyte solutions Cat # Dopant type for PEDOT DCDPGYIGSR DEDEDYFQRYLI EDOT 483028 0.1M 0.1M DCDPGYIGSR 12543000 5 mg/mL and 1 mg/mL DEDEDYFQRYLI 1254300 5 mg/mL and 1 mg/m Deionised (DI) Water N/A 50% 50% Acetonitrile 271004 50% 50% Component

PEDOT films doped with DEDEDYFQRYLI and DCDPGYIGSR at both 1 mg/mL and 5 mg/mL were analysed by CV to determine electrochemical stability. An eDAQ potentiostat and eCorder unit coupled with the supplied EChem software package (eDAQ Pty Ltd., Australia) was used to apply a cycling voltage from -800 mV to 600 mV. This range was chosen to allow the oxidation and reduction of the polymer within the limits of the water window, to prevent hydrolysis occurring at the electrodes. The scan rate was set at 120 mV/s for 400 continuous cycles and measurements were performed in 0.9% saline. The recordings were made with an isolated Ag/AgCl reference electrode and Pt counter electrode. The first stable curve was considered Cycle 1 and used to determine the original electroactivity of the film. The area contained within the oxidation-reduction curve was calculated to establish the current carrying capacity of the film. The area contained

248APPENDIX E. EFFECT OF PEPTIDE CONCENTRATION ON ELECTROCHEMISTRY within the curve of successive cycles was calculated as a percentage of Cycle 1 to show the comparative loss of electroactivity over time. This process was repeated for three individually prepared samples and the mean curve for electroactivity loss was determined with standard error.

E.3


The resultant CV curves of the peptide doped polymers made it clear that the higher concentration of laminin peptide was required in the electrolyte solution to preserve electrochemical stability. The curves obtained from analysing the CVs are presented in Figure E.1.

Figure E.1: Electroactivity loss related to peptide dopant concentration for both DEDEDYFQRYLI and DCDPGYIGSR, (n=3).

E.4. CONCLUSION

249

It is thought that at 1 mg/mL insufficient dopant is available in the electrolyte solution. Diffusion limitations resulting from the large biomolecules are further affected by the low number of available anions. As such the resultant polymer has limited electroactivity which degrades quickly. In the electrolytes with higher concentrations of peptides, the dopant molecules are abundant enough to produce a polymer with a relatively stable backbone, despite the limited mobility of the anions.

E.4

Conclusion

Laminin peptides will be included in electrolyte solutions at a concentration of 5 mg/mL. Higher concentrations are not desired as they may result in electrodeposition diffusion limitations that prevent the formation of the PEDOT.

Appendix F

Effect of SS concentration on RGC-5 Differentiation F.1

Introduction

RGC-5 differentiation in response to staurosporine (SS) was characterised through glutamate toxicity assays. Glutamate acts as a normal neurotransmitter in the retina, but its high levels are neurotoxic when used on retinal neuronal cells, resulting in apoptosis of retinal ganglion cells (RGCs). Glutamate stimulates the production of large quantities of nitric oxide (NO) which triggers cell death by apoptosis [158]. Differentiation of the RGC-5 by SS renders them susceptible to toxic death when glutamate is present in high concentrations.

Viable cell numbers following glutamate toxicity can be quantified through neutral red (NR) assays. The NR cytotoxicity assay procedure is a cell survival/viability chemosensitivity assay, based on the ability of viable cells to incorporate and bind NR dye. NR 250

F.2. MATERIALS AND METHODS

251

readily penetrates cell membranes by non-ionic diffusion, accumulating intracellularly. Alterations of the cell surface or lysosomal membrane result in a decreased uptake and binding of NR. It is thus possible to distinguish between viable or dead cells and hence determine the concentration of SS required to promote RGC-5 differentiation.

F.2


RGC-5s grown to confluency in 25 cm2 TCP flasks were detached by application of 2 mL trypsin. After 5 min incubation the trypsin was deactivated by the application of 5 mL DMEM containing 10% FBS. The supernatant was pipetted into a polypropylene centrifuge tubes and centrifuged at 1000 rpm for 5 min. The supernatant was dumped, leaving a visible cell clump. A 1 mL Eppendorf pipette tip was used to add 1 mL of fresh media to the cell suspension. Cells were aspirated gently using the 1 mL pipette tip to break up the clumps through 40 aspiration repetitions. DMEM with 10% FBS was used to dilute the cell solution to 5 mL and a 0.5 mL sample was taken for counting. Cells were plated in a 24-well TCP plate at 10,000 cells/cm2 and cultured for 24 hr in DMEM with 10% FBS at 37◦ C, 5% carbon dioxide (CO2 ) and 100% humidity.

At 24 hr post-plating the cell media was replaced with low serum (1% FBS) DMEM supplemented with SS. SS was applied to wells at 0 nM, 20 nM, and 50 nM as depicted in Figure F.1. An earlier pilot study revealed that higher concentrations of SS resulted in some cell death over the incubation period. Cells were incubated for a further 72 hr and then treated with glutamate at various concentrations to determine the highest tolerated

252APPENDIX F. EFFECT OF SS CONCENTRATION ON RGC-5 DIFFERENTIATION dose (HTD). Glutamate was applied to wells at 0 μM, 200 μM, 500 μM and 1 mM, as shown in Figure F.1, in serum free DMEM and incubated for a further 24 hr.

Figure F.1: Assay concentrations of SS and glutamate in a 24-well plate, used to determine optimum concentration of SS for RGC-5 differentiation

To conduct the NR assay, media was washed from the cells with HEPES buffer (125 mM NaCl, 5 mM KCl, 1.8 mM CaCl2 , 2 mM MgCl2 , 0.5 mM NaH2 PO4 , 5 mM NaHCO3 , 10 mM D-glucose, and 10 mM HEPES [pH 7.2]). NR was added to a final concentration of 0.033% in HEPES buffer and 0.5 mL aliquots were incubated with cells for 2 hr at room temperature. After the NR uptake by living cells, wells were gently washed with 2 to 4 volumes of HEPES buffer to remove excess dye. Cells were allowed to air dry for 20 min and then treated with ice-cold solubilisation buffer (1% acetic acid/50% ethanol; 500 μL) to extract the dye taken up by the cells. Optical densities were measured 20 min later when NR was extratced and visible in the buffer. Samples were placed in a plate reader

F.3. RESULTS AND DISCUSSION

253

and values measured at 570 nm.

F.3


Optical densities for the SS concentrations were plotted against the glutamate concentrations in Figure F.2. All samples without SS treatment were not affected by glutamate revealing that these RGC-5s were not differentiated. Cell survival in samples with 20 nM and 50 nM SS were affected by glutamate, even at the lowest concentration of 200 μM. Increased cell death was shown to occur at higher glutamate concentrations. Both con-

Figure F.2: Optical densities of NR dye, used to quantify cell viability following glutamate toxicity of differentiated RGC-5s at varying concentrations of SS. Error bars represent SEM, (n=2).

254APPENDIX F. EFFECT OF SS CONCENTRATION ON RGC-5 DIFFERENTIATION centrations of SS were able to promote differentiation of the RGC-5s, with no significant difference in cell susceptibility to glutamate toxicity.

Controls where SS was added but glutamate insults were not performed (0 μM glutamate), had lower optical densities than the undifferentiated (no SS) RGC-5s, indicating that viable cell numbers were lower. This was thought to be due to the reduction in cell proliferation brought on by the SS. Microscopic observations confirmed this hypothesis, as the RGC-5s in the wells without SS treatment appeared confluent at 72 hr post-plating.

F.4

Conclusions

The RGC-5 differentiation assay performed using glutamate toxicity, revealed that 20 nM SS was sufficient to promote cell differentiation, detected at even the lowest concentration of glutamate. It is possible that lower concentrations could be used, but working aliquots would be difficult to manage in the smaller well environment employed for polymer studies.

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