| 014 | sami myllymaa | novel micro- and nano-technological … · 2010-11-10 · implant/tissue...

Publications of the University of Eastern FinlandDissertations in Forestry and Natural Sciences

Publications of the University of Eastern Finland

Dissertations in Forestry and Natural Sciences

isbn 978-�952-�61-�0216-�0

Sami Myllymaa

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

Recent advances in micro- and

nanotechnology offer a great oppor-

tunity to develop intelligent bioma-

terials and the next generation of

implantable devices for diagnostics,

therapeutics, and tissue engineering.

This dissertation is focusing on the

development of novel polymer-based

microelectrode arrays suitable for

use in intracranial electroencepha-

lographic recordings. Moreover,

the performances of novel thin film

materials and their surface modifica-

tions at micro- and nanoscales were

studied with physicochemical and

cellular experiments in order to de-

vise new solutions for further devel-

opment of biomedical microdevices.

dissertatio

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Sami MyllymaaNovel Micro- and

Nano-technological Approaches for Improving

the Performance of Implantable

Biomedical Devices

SAMI MYLLYMAA

Novel Micro- and Nano-

technological Approaches for

Improving the Performance

of Implantable Biomedical

Devices



Number 14

Academic Dissertation

To be presented by permission of the Faculty on Sciences and Forestry for public

examination in the Auditorium, Mediteknia Building at the University of Eastern

Finland, Kuopio, on Friday 19th November 2010, at 2 p.m.

Department of Physics and Mathematics

Kopijyvä Oy

Kuopio, 2010

Editors: Prof. Pertti Pasanen,

Prof. Tarja Lehto, Prof. Kai Peiponen

Distribution:

Eastern Finland University Library/Sales of Publications

P.O. Box 107, FI-80101 Joensuu, Finland

Tel: +358-50-3058396

http://www.uef.fi/kirjasto

ISBN 978-952-61-0216-0

ISSN 1798-5668

ISSNL 1798-5668

ISBN 978-952-61-0217-7 (PDF)

ISSN 1798-5676 (PDF)

Author’s address: Department of Physics and Mathematics

University of Eastern Finland

P.O. Box 1627

FI-70211 KUOPIO

FINLAND

E-mail: [email protected]

Supervisors: Prof. Reijo Lappalainen, Ph.D.



Prof. Juha Töyräs, Ph.D.



Reviewers: Prof. Timo Jämsä, Ph.D.

Department of Medical Technology

University of Oulu

Prof. Sami Franssila, Ph.D.

Department of Material Science and Engineering

Aalto University School of Science and Technology

Opponent: Prof. Jukka Lekkala, Ph.D.

Tampere University of Technology

Department of Biomedical Engineering

ABSTRACT

Biomaterials are used in a wide variety of in vivo applications,

ranging from joint and dental implants to neural prostheses. The

ultimate success or failure of implants mainly depends on the

biological interactions (molecular, cellular, tissue) at the

implant/tissue interface. Recent advances in micro- and

nanotechnology offer a great opportunity to develop intelligent

biomaterials and the next generation of implantable devices

may well be of achieving the desired tissue-implant interaction

and resolving various biomedical problems.

The main aim of this thesis work was to design, fabricate and

evaluate a novel flexible microelectrode array suitable for use in

sub- or epidural electroencephalographic recordings. Other aims

were to investigate the opportunities to improve the

electrochemical and biological properties of neural interfaces

using modern micro- and nanotechnology tools as well as to test

whether the micropatterning of thin films can be used to guide

the cellular response on biomaterial surface.

The developed microelectrode array was implemented on

polyimide with platinum which achieved both mechanical

flexibility and high quality electrochemical characteristics as

demonstrated via impedance spectroscopy. Somatosensory and

auditory evoked potentials were successfully recorded with

epidurally implanted array in rats with excellent signal stability

over two weeks. Subsequently, the signal levels declined, most

probably due to the thickening of dura and the growth of scar

tissue around the electrodes. It was hypothesized that one

obvious reason for this limited life-span was the poor

biocompatibility of photosensitive polyimide used as an

insulation material in these arrays. However, this possibility

was excluded by in vitro cytotoxicity studies according to ISO

10993-5 standard. Furthermore, ultra-short pulsed laser

deposition was demonstrated to be an effective method to

produce nanotextured platinum surfaces as well as ultrasmooth

insulators for further development of neural interfaces.

Experiments with osteoblast-like cells and mesenchymal

stem cells on micropatterned biomaterial surfaces indicated that

even partial coating of silicon with a biocompatible material is

an effective way to enhance the cytocompatibility of silicon-

based biomedical micro-electromechanical systems. Moreover, it

was demonstrated that not only the chemical composition of the

materials, but also the shape, edges (height) and size of the

features used for surface patterning have a remarkable effect on

cell guidance.

Overall, the results of the present thesis provide a solid basis

for the further development of neural interfaces as well as other

types of implantable devices.

National Library of Medicine Classification: QT 36, QT 37, WL 102,

WV 270

INSPEC Thesaurus: biomedical materials; biomedical electrodes; thin

films; microelectrodes; electroencephalography; bioelectric potentials;

polymer films; silicon; platinum; metals; adhesion; surface texture;

surface topography; surface energy; nanopatterning; nanotechnology;

microfabrication; vapour deposited coatings; pulsed laser deposition

Yleinen suomalainen asiasanasto: biomateriaalit; implantit; polymeerit;

pii; platina; pinnoitus; pinnoitteet; pintarakenteet; pintailmiöt; kuviot;

mikrotekniikka; nanotekniikka; mikroelektrodit; EEG

Preface

This study was carried out during the years 2005-2010 in the

Department of Physics and Mathematics, the Department of

Neurobiology and BioMater Centre, University of Eastern

Finland, Microsensor Laboratory, Savonia University of Applied

Sciences, Kuopio, and the Department of Medicine, Helsinki

University Central Hospital.

I wish to express my deepest gratitude to everyone who has

contributed to the studies included in this thesis. Especially, I

wish to mention the following persons. First of all I would like

to thank my principal supervisor, Professor Reijo Lappalainen

for giving me the opportunity to work in his research group and

providing great facilities and scientific guidance for research

work. I am also grateful to my second supervisor, Professor Juha

Töyräs for his professional guidance and never-ending

enthusiasm and optimism during this thesis work.

I want to thank my official reviewers Professor Timo Jämsä

and Professor Sami Franssila for their constructive criticism and

valuable suggestions. I am also grateful to Ewen MacDonald for

linguistic review of this thesis.

The studies included in this interdisciplinary thesis would

not have been possible without the excellent collaboration with

colleagues in the several research units. Professor Heikki Tanila

is gratefully acknowledged for his valuable ideas and guidance

related to neuroscience. Kaj Djupsund is gratefully thanked for

his patient testing of microelectrode array prototypes in the

animal model. Professor Yrjö T. Konttinen, Emilia Kaivosoja and

Vesa-Petteri Kouri are sincerely thanked for great co-operation

in studies aimed to clarify the interaction phenomena between

cells and engineered biomaterial surfaces. Professor Mikko

Lammi, Virpi Tiitu, Sanna Miettinen and Aila Seppänen are

gratefully acknowledged for their contributions in cytotoxicity

testing. Moreover, I would like to thank all other co-authors in

the original publications of this thesis including Professor Juha E.

Jääskeläinen, Irina Gureviciene and Tarvo Sillat.

I sincerely thank all of the members in the Biomaterial

Technology Research Group and BioMater Centre for providing

supportive working atmosphere. In particular, the contributions

of Hannu Korhonen, Markku Tiitta, and Juhani Hakala have

been valuable. Aimo Tiihonen is sincerely thanked for technical

assistance in electronics. The former and current staff of

Microsensor Laboratory, including Matti Sipilä, Pasi Kivinen,

Mikko Laasanen and Ari Halvari is greatly acknowledged for

providing excellent microfabrication facilities and giving their

contributions to my work. Picodeon Ltd. Oy is acknowledged

for providing nanotextured depositions.

I am greatly indebted to my parents Seija and Heimo for their

love and support during my life. I want to thank my parents-in-

law, Anneli and Jukka, for their help and encouragements. I am

also grateful to my friends and relatives for their support.

Finally, I owe my deepest thanks to my beloved wife Katja, for

her irreplaceable love, support, and understanding. As a

research colleague, your contribution to my research has been

invaluable. Words are not enough for me to describe how

important you are as a wife and as the mother of our two little

daughters. Dearest Lumia and Neelia, you are the sunshine of

my life and have helped me to remember that there are more

important and valuable things in life than work.

This study was financially supported by Finnish Funding

Agency for Technology and Innovation (TEKES), the National

Graduate School of Musculoskeletal Diseases and Biomaterials,

Otto A. Malm Foundation, Ulla Tuominen Foundation,

Foundation for Advanced Technology of Eastern Finland and

COST, European Cooperation in Science and Technology.

Kuopio, October 2010

Sami Myllymaa

ABBREVIATIONS AND SYMBOLS

µCP microcontact printing

AEP auditory evoked potential

AFM atomic force microscopy

Ag silver

Ag-AgCl silver-silver chloride

Al aluminium

Al2O3 alumina (aluminium oxide)

ALP alkaline phosphatase

ANOVA analysis of variance

AP action potential

Au gold

BHK-21 baby hamster kidney fibroblast

Bio-MEMS biomedical micro-electromechanical systems

Ca calcium

Cl chlorine

CN carbon nitride

CNS central nervous system

Cr chromium

CVD chemical vapour deposition

CZ Czochralski process for silicon crystallization

DBS deep brain stimulation

DC direct current

DLC diamond-like carbon

DMEM Dulbecco’s modified Eagle’s medium

DNA deoxyribonucleic acid

ECG electrocardiography

EEG electroencephalography

EIS electrical impedance spectroscopy

EMG electromyography

ESEM environmental scanning electron microscope

EP evoked potential

FCS fetal calf serum

FDA Food and Drug Administration

FZ float-zone process for silicon crystallization

HMDS hexamethyldisilazane

hMSC human mesenchymal stem cell

Ir iridium

ISO International Organization for Standardization

ITO indium tin oxide

K potassium

KOH potassium hydroxide

MEA microelectrode array

MSC mesenchymal stem cell

MEMS micro-electromechanical systems

MS mineral staining

MSCGM mesenchymal stem cell growth medium

MTS cell proliferation assay based on (3-(4,5-

dimethylthiazol-2-yl)-5-(3 carboxymethoxy-

phenyl)-2-(4-sulfophenyl)-2H-tetrazolium)-salt

Na sodium

OC osteocalcin

PGMEA propylene glycol methyl ether acetate

PBS phosphate buffered saline

PCB printed circuit board

PDMS polydimethylsiloxane

PE polyethylene

PI polyimide

PNS peripheral nervous system

PSPI photosensitive polyimide

Pt platinum

PVD physical vapour deposition

RIE reactive ion etching

SAM self assembled monolayer

SD standard deviation of the mean

SEM scanning electron microscopy

SEM standard error of the mean

SEP somatosensory evoked potential

SFE surface free energy

Si silicon

SS stainless steel

Ta tantalum

Ti titanium

TiN titanium nitride

USPLD ultra-short pulsed laser deposition

UV ultraviolet

VEP visual evoked potential

ZIF zero insertion force

WC Warburg (polarization) capacitance

f frequency

aR average surface roughness as arithmetic mean

deviation of surface

pvR peak-to-valley roughness

SR resistance of electrode solution

pvR Warburg (polarization) resistance

mV transmembrane potential

ppV peak-to-peak signal amplitude

Z impedance

S total surface free energy D

S dispersive component of total surface free energy P

S polar component of total surface free energy

zeta potential

LIST OF ORIGINAL PUBLICATIONS

This thesis is based on, but not limited to, results presented in

the following original publications, which are referred in the

text by their Roman numerals (I-VI):

I Myllymaa, S., Myllymaa, K., Korhonen, H., Gureviciene, I.,

Djupsund, K., Tanila, H. & Lappalainen, R. (2008) Development

of flexible microelectrode arrays for recording cortical surface

field potentials. In: Proceedings of the 30th Annual International

Conference of the IEEE engineering in Medicine and Biology Society,

Vancouver, British Columbia, Canada, 20-24 August, 2008, pp.

3200-3203.

II Myllymaa, S., Myllymaa, K., Korhonen, H., Töyräs, J.,

Jääskeläinen, J.E., Djupsund, K., Tanila, H. & Lappalainen, R.

(2009) Fabrication and testing of polyimide-based micro-

electrode arrays for cortical mapping of evoked potentials.

Biosensors and Bioelectronics 24, 3067-3072.

III Myllymaa, S., Myllymaa, K., Korhonen, H., Lammi, M.J., Tiitu,

V. & Lappalainen, R. (2010) Surface characterization and in vitro

biocompatibility assessment of photosensitive polyimide films.

Colloids and Surfaces B: Biointerfaces 76, 505-511.

IV Myllymaa, S., Myllymaa, K. & Lappalainen, R. (2009) Flexible

implantable thin film neural electrodes. In: Recent Advances in

Biomedical Engineering, In-Tech, Vukovar, Croatia, Editor:

Ganesh R. Naik, ISBN 978-953-307-004-9, Chapter 9, pp. 165-189.

V Myllymaa, S.*, Kaivosoja, E.*, Myllymaa, K., Sillat, T., Korhonen,

H., Lappalainen, R. & Konttinen, Y.T. (2010) Adhesion,

spreading and osteogenic differentiation of mesenchymal stem

cells cultured on micropatterned amorphous diamond, titanium,

tantalum and chromium coatings on silicon. Journal of Materials

Science: Materials in Medicine 21 (1), 329-341.

(* equal contribution)

VI Kaivosoja, E.*, Myllymaa, S.*, Kouri, V.-P., Myllymaa, K.,

Lappalainen, R. & Konttinen, Y.T. (2010) Enhancement of silicon

using micropatterned surfaces of thin films. European Cells and

Materials, 19, 147-157. (* equal contribution).

The publications are reprinted with the kind permission of the

copyright holders. This thesis also contains unpublished results

related to publications III and VI as well as unpublished data

focused on the electrochemical characterization of nanotextured

bioelectrode surfaces.

AUTHOR’S CONTRIBUTION

Publications I and II concern the development of flexible

microelectrode array suitable for recording cortical surface field

potentials. The original idea of this electrode system was

proposed by H. Tanila. The author of this thesis carried out most

of the development work related to the electrode design,

material selection and fabrication processes. Apart from

depositions of metal thin-films, which were performed by H.

Korhonen and R. Lappalainen, the author carried out all array

fabrication steps and characterization experiments with a

contribution from K. Myllymaa. In vivo testing of electrodes was

performed by K. Djupsund and I. Gureviciene. The author wrote

the articles, after receiving constructive comments from the co-

authors. R. Lappalainen supervised the work.

Publication III concerns the cytocompatibility of the novel

photosensitive polyimide used as an insulator material in our

sensor prototypes (papers I and II). The design and fabrication

of test samples were performed by the author. The author also

conducted the surface characterization together with H.

Korhonen and K. Myllymaa. The author planned cell

experiments together with V. Tiitu, who also performed the

experimental studies. The author was responsible for analyzing

the data, presenting the results and writing the article, while

receiving constructive comments from the co-authors. R.

Lappalainen, V. Tiitu and M.J. Lammi supervised the study.

The author was mainly responsible for writing publication IV

(book chapter), which is a literature review discussing the main

requirements and features of flexible thin film neural electrodes,

this being supplemented with personal results in the area of

sensor development.

Publications V and VI concern the interactions between cells

and biomaterial surfaces. R. Lappalainen and Y.T. Konttinen

originally presented the idea of studying the effect of surface

micropatterning on the cellular response. Apart from thin film

depositions, performed by H. Korhonen and R. Lappalainen, the

author carried out all sample design and sample fabrication

steps in collaboration with K. Myllymaa. The surface

characterization analysis was mainly performed by the author

with some contributions from H. Korhonen and K. Myllymaa.

Experimental work and data analysis on cellular studies was

planned and carried out by E. Kaivosoja, with contribution of

V.-P. Kouri in paper VI, and Y.T. Konttinen. The author and E.

Kaivosoja contributed equally to preparing results and writing

paper V, assisted by R. Lappalainen and Y. T. Konttinen. E.

Kaivosoja and Y.T. Konttinen chiefly wrote paper VI, assisted by

S. Myllymaa and R. Lappalainen. These studies were supervised

by R. Lappalainen and Y.T. Konttinen.

Contents

1 INTRODUCTION ........................................................................... 21

2 BIOMEDICAL MICRO-ELECTROMECHANICAL

SYSTEMS .......................................................................................... 25

2.1 Microfabrication techniques ........................................................... 26

2.1.1 Photolithography ...................................................................... 27

2.1.2 Soft lithography ........................................................................ 29

2.1.3 Thin film deposition ................................................................. 32

2.1.4 Lift-off processing ..................................................................... 34

2.1.5 Etching ..................................................................................... 35

2.2 Substrate materials ........................................................................... 37

2.2.1 Silicon ...................................................................................... 37

2.2.2 Glass ......................................................................................... 38

2.2.3 Polymers................................................................................... 38

2.3 Advantages of microfabrication .................................................... 40

2.4 Bio-MEMS applications ................................................................... 42

3 ORIGIN OF BIOELECTRIC SIGNALS ...................................... 45

3.1 Excitable nerve cell .......................................................................... 45

3.1.1 Cell membrane and ion channels ............................................. 47

3.1.2 Transmembrane potential and equilibrium potentials ............. 47

3.2 Synaptic potentials and action potentials ..................................... 49

3.3 Recording electrical activity of the brain ...................................... 51

4 BIOELECTRODES .......................................................................... 55

4.1 Implantable electrodes .................................................................... 56

4.2 Material requirements for implantable electrodes ...................... 59

4.2.1 Electrode materials ................................................................... 60

4.2.2 Substrate materials .................................................................. 63

5 BIOCOMPATIBILITY .................................................................... 67

5.1 Biocompatibility testing .................................................................. 70

5.2 Cell-biomaterial interactions .......................................................... 71

5.2.1 Effect of surface topography on cellular responses ................... 71

5.2.2 Effect of wettability properties on cell-biomaterial

interactions .............................................................................. 75

5.2.3 Effect of charge distribution on cell-biomaterial interactions .. 76

5.2.4 Effect of surface micropatterning on cellular responses ........... 78

6 AIMS OF THE PRESENT STUDY ............................................... 79

7 MATERIALS AND METHODS ................................................... 81

7.1 Micro- and nanofabrication of electrodes and other samples ... 82

7.1.1 Preparation of photomasks ....................................................... 82

7.1.2 Flexible polyimide-based microelectrode arrays ....................... 82

7.1.3 Micropatterned biomaterial surfaces on silicon ....................... 86

7.1.4 Tailored surfaces produced by ultra-short pulsed laser

deposition ................................................................................. 88

7.1.5 Spin-coated polyimide films for cytotoxicity testing ................ 89

7.2 Microscopic characterization of surfaces ...................................... 90

7.2.1 Scanning electron microscopy ................................................. 90

7.2.2 Atomic force microscopy .......................................................... 90

7.2.3 Contact angle measurements and determination of surface

free energies.............................................................................. 91

7.3 Electrochemical characterization of surfaces ............................... 92

7.3.1 Electrochemical impedance spectroscopy ................................. 92

7.3.2 Zeta potential measurements ................................................... 94

7.4 Recording of evoked potentials in rats ......................................... 95

7.5 Cell culture studies .......................................................................... 96

7.5.1 Cell lines ................................................................................... 96

7.5.2 MTS assay ................................................................................ 98

7.5.3 Scanning electron microscopy of cultured cells ....................... 98

7.5.4 Immunofluorescence and confocal laser scanning

microscopy ............................................................................... 99

7.6 Statistical analyses ......................................................................... 101

8 RESULTS ........................................................................................ 103

8.1 Evaluation of microelectrode arrays ........................................... 103

8.2 Evaluation of nanorough electrode surfaces .............................. 109

8.3 Cytocompatibility testing of insulator materials ....................... 111

8.4 Micropatterned biomaterial coatings .......................................... 115

8.4.1 Behaviour of human mesenchymal stem cells ........................ 116

8.4.2 Behaviour of osteoblast-like (SaOS-2) cells ............................ 120

9 DISCUSSION ................................................................................ 125

9.1 Microelectrode arrays for intracranial recordings .................... 125

9.2 Material science strategies to improve the performance of

microelectrode arrays .......................................................................... 128

9.3 Micropatterned biomaterial surfaces .......................................... 130

10 CONCLUSIONS .......................................................................... 133

11 REFERENCES .............................................................................. 135

APPENDIX: ORIGINAL PUBLICATIONS I-VI

21

1 Introduction

Developments in micro-electromechanical systems (MEMS)

technology have introduced a variety of breakthrough products

in the fields of microelectronics, telecommunications and the

automotive industry (Judy 2001, Madou 2002). In recent years,

the biomedical applications of micro-electromechanical systems

(bio-MEMS) have attracted growing interest in many

application areas such as diagnostics, therapeutics and tissue

engineering (Saliterman 2006, Grayson et al. 2004, Nuxoll &

Siegel 2009, Betancourt & Brannon-Peppas 2006, Bashir 2004,

Urban 2006). Bio-MEMS is a powerful technology capable of

ever-greater functionality and cost reduction in smaller

biomedical devices for improved diagnostics and treatment. The

merging of biology with micro/nanotechnology has been

postulated to trigger a scientific and technological revolution in

the future (Kotov et al. 2009). Although some applications such

as blood analysis cartridges, catheter pressure sensors and

cochlear implants have been already commercialized, the vast

majority of biomedical applications are still being investigated

or undergoing clinical trials (Saliterman 2006, Urban 2006).

Brain research can be considered as a one of the most

challenging scientific areas. The very first biological MEMS

devices, i.e. multi-sensor neural probes, were developed for

neuroscientists in order to facilitate studying of neuronal

activities at the tissue and cellular level (Urban et al. 2006). In

addition to pure scientific research, there has been a growing

interest in the clinical applications of stimulating and recording

neural electrodes and prostheses (DiLorenzo & Bronzino 2008).

These neural interfaces can benefit many patients with neural

disorders such as impaired hearing (cochlear implant) or

neuropathic pain (deep brain stimulators). Some extremely

challenging applications such as retina implants which may be

able to restore the vision or sieve and cuff electrodes potentially

1 - Introduction

22

evoking nerve regeneration in paralyzed patients are still under

intensive development (Stieglitz & Mayer 2006b, DiLorenzo &

Bronzino 2008).

To date, a wide variety of neural electrodes have been

utilized in neural interfaces starting with the early electrolyte-

filled micropipettes and subsequent metal electrodes to the

current emerging MEMS-based electrodes. The vast majority of

these MEMS devices have been fabricated on silicon (Si) even

though there is a huge mismatch between mechanical properties

of the neural tissue and Si causing many adverse tissue effects

(Polikov et al. 2005). Moreover, Si is not per se a biocompatible

material (Voskerician et al. 2003), and it is a poor substrate for

cell adhesion, even being slightly cytotoxic (Liu et al. 2007).

These weaknesses of Si could limit the integration of Si-based

devices into the human body. Recently, there has been a

growing interest toward developing polymer-based interfaces

that could be flexible enough to mimic biological tissue,

reducing mechanical damage and evoking less adverse tissue

reactions (Stieglitz & Mayer 2006b, Cheung 2007, HajjHassan et

al. 2008, DiLorenzo & Bronzino 2008).

Biomaterials are used in many in vivo applications, ranging

from joint and dental implants to neuroprosthetic devices.

Depending on particular specifications for each such application,

a set of different material characteristics, such as mechanical,

chemical and electrical properties, influence the performance of

biomaterial/prosthetic devices in a specific manner (Williams

2008). In all cases, however, the ultimate success or failure

depends on the biological interactions (molecular, cellular,

tissue) at the implant-tissue interface (Puleo & Nanci 1999,

Williams 2008, Navarro et al. 2008, Polikov et al. 2005).

Therefore, several approaches have been introduced to modify

surface properties, such as surface topography on the micro-

and nanoscales (Flemming et al. 1999, Martinez et al. 2009),

surface energy (van Kooten et al. 1992, Hallab et al. 2001, Lim et

al. 2008) and surface charge (Krajewski et al. 1996, Krajewski et

al. 1998, Cai et al. 2006), in order to achieve significant effects on

the functions of proteins and thus on cells. These enhanced

1 - Introduction

23

effects on cellular functions are crucial in subsequent

improvements in new tissue formation and integration of

implants in the surrounding tissues. Recently, much attention

has been paid to mesenchymal stem cells (MSCs) due to their

role in tissue repair and their potential in regenerative medicine

and tissue engineering in which these kinds of cells are grown

on biomaterial scaffold, which provides structural support and

substrate for cellular adhesion (Zippel et al. 2010, Pittenger et al.

1999, Stoddart et al. 2009). The ability to influence the adhesion,

distribution and behaviour of cells on biomaterial surfaces and

the knowledge of the nature of cell-substrate interaction has

therefore become increasingly important.

In this thesis, novel flexible polyimide (PI)-based

microelectrode arrays (MEA) suitable for sub- or epidural

electroencephalographic (EEG) recordings were developed. The

suitabilities of different MEMS materials as well as their surface

modifications for implantable applications were investigated

using electrochemical testings and cell experiments. Moreover,

the effect of surface micropatterning achieved using photo-

lithographic and thin film techniques on the behaviour of MCSs

and osteoblast-like cells was studied.

1 - Introduction

24

25

2 Biomedical micro-

electromechanical systems

Micro-electromechanical systems refer to miniature devices that

are fabricated using techniques originally developed and widely

utilized in the microelectronics (integrated circuits) industry,

and then modified, e.g. by adding mechanical components, such

as beams, gears, and springs, for the creation of microstructures

and microdevices such as sensors and actuators. MEMS, as well

as its interchangeable acronyms, Micromachines (popular in

Asia) and Microsystems (popular in Europe), refer to the

miniature devices which have at least some of their dimensions

in the micrometer range. According to the widest definition,

MEMS comprises all devices and systems produced by

micromachining other than integrated circuits or other

conventional semiconductor devices (Judy 2001). At the

beginning of MEMS era, in the 1970s and 1980s, the field was

dominated by mechanical applications, but today most new

applications are either communication/information-related or

chemical and biological in nature (Madou 2002). MEMS

technology has enabled low-cost, high-functionality micro-

devices in some widespread application areas, such as printer

cartridges for ink jet printing and the accelerometers responsible

for deployment of airbags in modern automobiles (Judy 2001).

In recent years, the biological and biomedical applications of

MEMS, which are commonly referred to as bio-MEMS, have

gained increasing world-wide interest (Saliterman 2006, Nuxoll

& Siegel 2009, Urban 2006). Bio-MEMS are generally defined as

‘‘devices or systems, constructed using techniques inspired from

micro/nanoscale fabrication, that are used for processing,

delivery, manipulation, analysis, or construction of biological

and chemical entities’’ (Bashir 2004). At the present moment,

bio-MEMS technology is a topic of intense research and

2 – Biomedical micro-electromechanical systems

26

development activity with a wide variety of applications in the

fields of cell and molecular biology, pharmaceutics and

medicine so that it includes the areas of therapeutics,

diagnostics and tissue engineering (Saliterman 2006, Grayson et

al. 2004, Nuxoll & Siegel 2009, Betancourt & Brannon-Peppas

2006, Voldman et al. 1999, Bashir 2004).

2.1 MICROFABRICATION TECHNIQUES

A number of techniques are used to form bio-MEMS objects

with dimensions ranging from the micrometer to millimeter

scale. Some of these techniques have been adopted directly from

the industry of integrated circuits whereas some others have

been specifically developed for this novel purpose (Judy 2001,

Voldman et al. 1999, Saliterman 2006, Li et al. 2003). The

microfabrication process is typically a process flow that utilized

these techniques in a sequential manner to produce the desired

structure. MEMS devices can be built within the bulk of a

substrate material in what is referred to as bulk micromachining,

or if it is on the surface of the substrate it is known as surface

micromachining (Madou 2002). However, a combination of bulk

and surface micromachining is commonly used (Voldman et al.

1999). The most important microfabrication techniques in bio-

MEMS are photolithography, soft lithography, thin film

deposition and etching (Saliterman 2006, Li et al. 2003).

Photolithography is used to transfer the desired shape onto a

material through the selective exposure of a photosensitive

polymer (Madou 2002). Soft lithography is a set of different

patterning techniques based on polymer block used as a stamp,

mold or stencil for carrying out surface patterns and micro-

structures (Xia & Whitesides 1998, Li et al. 2003, Saliterman

2006). Thin film deposition which consists of numerous different

techniques is used to form thin layers with thickness ranging

from atomic layer to a few microns on the surface of a substrate.

Etching is used to remove material selectively from the surface


27

of the microsystem by either chemical or physical processes

(Madou 2002).

2.1.1 Photolithography

Photolithography is a basic microfabrication technique widely

employed to create the desired patterns onto a material. The

photolithographic patterning is composed of a number of

process steps (Fig. 1).

photoresist

substrate

photomask

UV exposure

resist deposition

and soft baking

latent image

mask alignment

(x, y, φ)

positive photoresist negative photoresist

development

Figure 1: Photolithographic patterning process. A photomask with opaque regions in

the desired pattern is used to selectively expose a photosensitive polymer (photoresist).

Depending on the polarity of the resist used, it will become more soluble (positive-tone

photoresist) or crosslinked (negative photoresist) after ultraviolet light illumination,

thus generating the desired pattern after developing in liquid chemicals.

Firstly, a pattern is drawn using computer assisted design

software and this is transferred onto a mask. The photomasks

are typically transparent glass plate blanks covered with an

opaque material (usually chromium, Cr) in the defined pattern

(Madou 2002). The masks are usually prepared by a commercial

mask manufacturer using electron beam or laser writing (Wu et


28

al. 2002). If the feature sizes and tolerances in the mask are

relatively large, an option may be to utilize low-cost

transparency films as a photomask (Deng et al. 1999, 2000, Wu

et al. 2002, Gale et al. 2008). After producing the mask, a

photosensitive polymer, i.e. photoresist, is spin-coated onto a

substrate material, such as a Si wafer. The resulting photoresist

thickness, t, is a function of spinner rotational speed, solution

concentration, and molecular weight (determined by intrinsic

viscosity) and can be estimated by (Madou 2002):

.

KCt (1)

In this equation, K is the overall calibration constant, C is the

polymer concentration (g/100 ml), η is the intrinsic viscosity and

ω is the spinner rotational speed (rotations per minute, rpm). α,

β and γ are exponential factors dependent on process.

Typically, spinning speeds of 1500-8000 rpm are used to

achieve resist thicknesses of about 0.5-2 µm (Madou 2002). The

photoresist film is then prebaked at 75 to 100°C on a hotplate or

in an oven to remove solvents and to promote adhesion of the

resist layer to the substrate (Saliterman 2006, Madou 2002). In

the following step of exposure, the photomask is placed on top

of the photoresist-coated wafer in close proximity or even in

contact with it and then ultraviolet (UV) light is used to

illuminate the photoresist film through the photomask. With

this procedure, the solubility difference between the exposed

and unexposed regions is achieved, and depending on the

polarity of the photoresist used, exposed or unexposed areas can

then be removed by dissolving them in a developing solution. In

the case of a positive-tone photoresist, the exposed regions

break down and become more soluble in the developing

solution, whereas in a negative-tone photoresist, the exposed

areas become crosslinked and insoluble in the developing

solution. The resulting photoresist pattern can now be used as a

protective mask in following microfabrication processes such as

in the etching step to prevent the covered substrate to be


29

removed and in thin film deposition to pattern metallization

layer on the surface of substrate. After the followed process is

finished, the photoresist can be removed, leaving the pattern

design on the substrate. The removal of photoresist is usually

performed by sonication in an organic solvent, e.g. acetone.

(Madou 2002)

The resolution of photolithography depends on the quality of

mask and the wavelength of applied electromagnetic radiation

(light). The typical wavelengths used are 436 nm, 365 nm, 248

nm and 193 nm (Madou 2002). The shorter the wavelength of

light used, the higher resolution, i.e. smaller feature sizes are

possible. Although photolithography is the main technology in

the production of microscale features in MEMS, it has also a few

short-comings. It requires clean-room facilities, and therefore it

is not an inexpensive technology; it is poorly suited for

patterning nonplanar surfaces and it is directly applicable only

to a limited set of photosensitive polymers (Xia & Whitesides

1998). Furthermore, photolithography requires the use of strong

solvents, meaning that it is poorly compatible with the

patterning biological molecules used in many applications e.g. in

biosensing, medical implants and the control of cell adhesion

and growth (James et al. 1998, Li et al. 2003, St. John et al. 1997).

2.1.2 Soft lithography

Soft lithography consists of a set of non-photolithographic

patterning methods based on self-assembly and replica molding

for carrying out micro- and nanofabrication (Saliterman 2006). It

provides a low-cost, effective, and biocompatible strategy for

the formation and manufacturing of surface patterns and

structures with feature dimensions ranging from 30 nm to 100

µm (Xia & Whitesides 1998). The key element of soft

lithography is an elastomeric block with patterned relief

structures on its surface that subsequently utilized in different

soft lithographic techniques such as in microcontact printing,

stencil patterning and microfluidic patterning (Li et al. 2003).

The commonest material used for the production of elastomeric

block is polydimethylsiloxane (PDMS). This material has several


30

desirable properties such as durability, optical transparency,

biocompatibility, flexibility, gas permeability and amenability to

different surface modifications (Whitesides et al. 2001, Li et al.

2003). The major advantages of soft lithography compared to

conventional photolithography are that numerous solid and

liquid materials other than photoresists can be patterned, and it

is suitable for curved or complicated surface geometrics and

even 3D structures (Xia & Whitesides 1998, Gale et al. 2008).

The soft lithography process starts with conventional

photolithographic steps to create a PDMS block which will be

used later as a mold or a stamp (Fig. 2a).

Si

PDMS

Si

photoresist

Pour PDMS over master

PDMS

Cure and peel off stamp

Alkanethiol ”ink”

PDMS

Substrate

Au

PDMS

Microcontact print

Substrate

Substrate

Remove stamp

SAM (2-3 nm)

A B

Figure 2: (a) Fabrication of a polydimethylsiloxane (PDMS) stamp. PDMS is poured

onto a silicon wafer coated with patterned photoresist. After the curing step, the stamp

can be removed. (b) Microcontact printing with the PDMS stamp. The stamp is

coated with alkenethiol ink, and then it is pressed on gold coated substrate. After

application of gently pressure for a few seconds, a self-assembly monolayer (SAM) is

formed on gold surface. Figure is inspired by Xia & Whitesides (1998).

Photoresist is spin-coated onto a substrate (e.g. Si wafer). The

height of the resist layer (i.e. pattern relief) can be controlled by

the viscosity of the solution and the spin speed (Madou 2002).

EPON™ SU-8, originally developed by IBM, and recently

marketed by MicroChem Corp. (Newton, MA, USA), is one the

most popular photoresists being used in stamp fabrication (Li et

al. 2003). This negative-tone epoxy-resist permits the fabrication


31

of tall structures, even more than 1 mm in height with a superior

aspect ratio (Lorenz et al. 1997, 1998). The photoresist layer is

illuminated with UV light through a photomask with the

desired pattern, and then developed. In the case of SU-8, the

resist areas exposed to UV light will remain in the development.

Liquid PDMS precursor is then cast over the patterned substrate

and cured at an elevated temperature. After cooling, the PDMS

stamp can be peeled off from the master. A replica of the

original template in a functional material can be generated by

molding against the patterned PDMS block. (Li et al. 2003, Xia &

Whitesides 1998, Whitesides et al. 2001)

The microcontact printing (µCP), also known as

microstamping, is based on the patterned transfer of the

material (‚ink‛) from the surface of the PDMS stamp onto the

receiving surface on a sub-micrometer scale (Fig. 2b) (Qin et al.

2010, Xia & Whitesides 1998, Kumar & Whitesides 1993). The

material of interest to be transferred, e.g. alkanethiol, is applied

onto the surface of the stamp with patterned surface relief. The

dried stamp is then placed face down on the substrate, e.g. gold

(Au)-coated glass, and gentle pressure is applied for a few

seconds (Qin et al. 2010). The elastic PDMS material enables an

excellent contact between the stamp and the substrate surface,

and molecules that touch the surface are transferred from the

stamp to the substrate, forming a self-assembly monolayer

(SAM). The formed SAM layer can be used as a protective layer

in subsequent microfabrication steps such as etching or

deposition (Xia & Whitesides 1998). The chemicals used in the

µCP to form SAMs typically have a chemical formula of

Y(CH2)nX, where Y is the anchor and X is the headgroup

(Saliterman 2006). A typical anchor group in alkanethiols is

sulfide since this promotes the binding of thiol very tightly to

Au. Typically, CH3 and COOH are used as the headgroups. The

choice of headgroup has a great influence on the wettability

properties, i.e. hydrophobicity/hydrophilicity, of the surface and

subsequently to protein and cell binding. Furthermore, the

headgroups can be chemically modified, e.g. an arginine-

glycine-aspartate peptide that enhances cell attachment (Hersel


32

et al. 2003, LeBaron & Athanasiou 2000, Shi et al. 2008, Chua et

al. 2008). The µCP offers the advantage over traditional photo-

lithographic patterning methods in that it requires no strong

chemicals, making it suitable for patterning biologically active

layers (Xia & Whitesides 1998, Whitesides et al. 2001, Saliterman

2006). In addition to alkanethiols, other peptides, proteins, poly-

saccharides and other molecules can also be stamped (Bernard

et al. 1998, Branch et al. 1998, James et al. 1998, Li et al. 2003).

Stencil patterning is the second soft lithography technique in

which a membrane stencil is fabricated by casting PDMS to the

top of a photoresist master, which creates holes with the shape

of the master features. The PDMS stencil can be used as a mask

for selective adsorption of cell-adhesive proteins to promote

cellular patterning. (Folch et al. 2000, Li et al. 2003, Folch &

Toner 2000)

Micromoulding in capillaries, the third soft lithographic

technique, employs a PDMS mold to build up microchannels

against a substrate. These microchannels can be used to pattern

fluid materials onto a substrate (Kim et al. 1995, 1996). A low

viscosity prepolymer is applied at the open ends of the channels.

The channels become filled with polymer due to capillary forces.

After curing of the prepolymer, the PDMS mold is removed and

the three-dimensional polymer microstructures are revealed.

Subsequently, these microstructures can be used for selective

delivery of different cell suspensions to specific locations of a

substrate resulting in micropatterns of attached cells (Folch &

Toner 1998, Tan & Desai 2003).

2.1.3 Thin film deposition

The application of thin layers of materials is a common

procedure in microfabrication. Thin films refer to thin material

layers ranging from the thickness of atomic layers to several

micrometers in thickness (Madou 2002). Thin films can play a

structural or functional role in the device process (Smith 1995,

Hsu 2008). All material classes, i.e. metals, ceramics, polymers,

composites and biological compounds, can be deposited. They

can be used either as sacrificial layers or mask layers that


33

selectively protect the substrate material during etching or they

can be used as electrical insulators or conductors such as

electrodes and transmission lines in sensors. Two common thin

film materials used as insulators in Si-based MEMS are silicon

dioxide (SiO2) and silicon nitride (Si3N4) (Voldman et al. 1999).

Deposition techniques can be divided into physical vapour

deposition (PVD) and chemical vapour deposition (CVD)

methods (Madou 2002). In general, the PVD methods are based

on the production of a condensable vapour by physical means

and its deposition on a substrate as a thin film. Vacuum or low

pressure gaseous environment is used in the PVD processing, in

which deposition species are atoms or molecules or both. The

most common PVD techniques are (1) evaporation in a vacuum,

(2) sputter deposition, (3) arc-vapour deposition, (4) laser

ablation, and (5) ion plating (Saliterman 2006). On the other

hand, the CVD methods are based on chemical reactions which

take place at a heated substrate surface to deposit a solid film.

Gas phase reactions between chamber gases are not desirable

because they often lead to poor adhesion, low density and high

detect films. The main advantages of the CVD methods are the

possibility to fill holes, cavities and other 3D structures, good

adhesion between coating and substrate, and excellent thickness

uniformity of coating. The major shortcomings are the toxicity of

the by-products as well as a high processing temperature, being

above 600°C which eliminates the use of CVD with heat

sensitive materials like polymers (Hsu 2008, Saliterman 2006,

Madou 2002).

In addition, thin films can be produced by other techniques,

such as spin coating, electrolytic deposition and electroless

deposition. Spin-coating is typically used to deposit thin

polymer films such as photoresists in photolithography. In the

electroplating process, the substrate is immersed in an

electrolyte solution. When a voltage is applied between a

substrate (working electrode) and an inert counter electrode,

such as platinum (Pt) in the liquid, chemical reduction-oxidation

processes take place resulting in the formation of a layer of

material on the surface of the substrate (Paunovic & Schlesinger


34

2006, Schlesinger & Paunovic 2000). An electroless plating

process does not require any external electrical potential but the

deposition spontaneously happens on any surface which forms

a sufficiently high electrochemical potential with the electrolyte

solution. Although electroless deposition (compared to the

electroplating) is much easier to set-up and use, without the

need for an electrical supply and its connections to electrodes, it

is also more difficult to control with regards to film thickness

and uniformity. Electroplating and electroless plating are

typically used in MEMS to form conductive metal (e.g. platinum,

gold, copper, nickel) or conductive polymer (e.g. polypyrrole,

polyaniline) thin films (Madou 2002).

2.1.4 Lift-off processing

Lift-off is a simple and easy method for patterning thin film

metal layers. It is especially useful for patterning catalytic metals,

such as Pt, which do not lend themselves well to direct wet

etching (Madou 2002, Hsu 2008). The lift-off process starts with

a photolithographic step, in which the photoresist layer is

deposited and patterned as described in chapter 2.1.1. Thin film

of desired material is then deposited all over the substrate,

covering the photoresist and areas in which the photoresist has

been cleared. Thereafter, the substrate is immersed in a solvent

that dissolves the remaining soluble photoresist underneath the

metal, and lifts off the metal. Only the metal which has been

deposited directly on the substrate leaves and forms the final

pattern on the substrate. In order to achieve clean results in lift-

off processing, photolithographic and thin film deposition

processes need to be optimized (Franssila 2004). Undercut

photoresist patterns are beneficial and can be realized by

reducing exposure dose and/or increasing the developing time

of negative photoresist. The substrates should be kept at

temperatures below 200 to 300°C during metal deposition to

avoid hardening and flow of photoresist (Madou 2002).


35

2.1.5 Etching

Etching techniques are used to create topographical patterns on

a surface via the selective removal of material. These techniques

can be roughly divided into wet and dry etching which are

based on the utilization of liquid chemicals (etchants) or reactive

ions/vapour-phase etchant, respectively (Madou 2002). The

etching process is isotropic in its nature if it removes material in

all directions equally, leading to mask undercutting and a

rounded etch profile (Fig. 3a), or it is anisotropic if removal

occurs with different etch rates in different directions in the

material (Fig. 3b,c) (Voldman et al. 1999). In the simple wet

etching technique, the sample is immersed in a container filled

with a liquid solution (etchant). The sample is covered by a

mask which leads to the selective removal of material (Madou

2002). The most critical task is to find a mask material that will

not dissolve or at least is etched much more slowly than the

sample material in question. A wide variety of etchants such as

buffered hydrofluoric acid, Aqua regia, Piranha solution (i.e.

mixture of sulfuric acid and hydrogen peroxide), potassium

hydroxide (KOH), tetramethylammonium hydroxide and

hydrochloric acid can be used to etch different MEMS materials

(Williams & Muller 1996, Williams et al. 2003). Unfortunately,

etchants are usually isotropic leading to etch profiles with large

undercuts particularly when one is machining thick films.

Additionally many wet etch chemicals are toxic and, thus they

are poorly suited for the state-of-the-art bio-MEMS processes.

Anisotropic wet etch is possible on some single crystal materials,

such as Si in which anisotropic etchants (e.g. KOH) dissolve Si

rapidly in the direction <100>, but almost not at all in the

direction of <111>. Thus cavities with trapezoidal cross-sections

with a characteristic 54.7° sidewall will be created on a [100]-

oriented wafer (Fig. 3c) (Madou 2002, Voldman et al. 1999).


36

etching mask

substrate

54.7

A B C

Figure 3: Different etching profiles resulted in (a) isotropic wet etching, (b)

anisotropic dry etching, and (c) anisotropic wet etching of silicon using potassium

hydroxide. Figure is modified from Voldman et al. (1999).

In order to achieve anisotropic etching with high resolution,

different dry etching methods are used. Reactive ion etching

(RIE) is one of the most popular dry etching methods in which

the substrate is placed inside a reactor and several gases are

introduced (Kovacs et al. 1998). The gas molecules can be made

to disintegrate into ions using a radiofrequency (13.56 MHz)

power source. Accelerated ions collide onto the surface of the

material to be etched, react with the surface molecules forming

another gaseous material. In addition to this chemical aspect,

there is also a physical part present in the process. If the plasma

particles, i.e. neutral radicals and ions, have enough energy, they

can remove atoms out of the materials without any chemical

reaction. This process can be quite complicated since there are

many parameters to be adjusted. However, with optimal

adjustment, it is possible to etch almost straight down without

undercutting, which provides much higher resolution compared

to wet etching. (Madou 2002)

A special improvement of RIE technique is the deep RIE,

based on the so called ‚Bosch process‛ (Laermer 1996) in which

an etch depth of hundreds of micrometers can be achieved with

vertical sidewalls. The deep RIE process is based on two

different alternating gas compositions in the reactor. The first

gas composition forms a polymer film on the surface of the

substrate, and the second gas composition etches the substrate

(Saliterman 2006). The polymer is immediately sputtered away

by the second etching gas, but this happens only on the


37

horizontal surfaces and not the sidewalls protecting them from

etching (Laermer 1996). Very high etching aspect ratios, even

50:1, can be achieved (Kovacs et al. 1998). Dry etching is a

widely used method in the integrated circuit industry to achieve

high resolution features, but in many cases it is not essential in

the production of bio-MEMS.

2.2 SUBSTRATE MATERIALS

Traditionally, MEMS devices have been fabricated on

microelectronics related materials, such as Si and glass. Recently,

however, there has been a growing interest towards rubber and

plastic substrate materials due to their suitable mechanical

properties, enhanced biocompatibility, rapid prototyping and

inexpensive manufacturing techniques available (Wilson et al.

2007).

2.2.1 Silicon

Silicon is the most widely used material in microchips and

MEMS devices since it is straightforward to grow oxide layers to

form dielectrics on its surface and it has excellent semiconductor

properties over a wide temperature range. Silicon wafers used

in semiconductor/MEMS industry is produced using two

alternative crystallization methods: Czochralski process (CZ)

and float-zone (FZ) process (Pearce 1988). In the CZ method, a

seed crystal is dipped into molten Si and pulled upwards with

simultaneous rotation. By optimizing the rate of pulling and the

speed of rotation as well as other relevant parameters such as

temperature, it is possible to extract a large-diameter cylindrical

single crystal ingot from the melt. Czochralski process is a more

common and cheaper method compared to FZ, but the wafers

produced using the CZ method contain slightly more impurities.

Impurities in the CZ wafers originate from the quartz crucible,

containing the Si melt that dissolves during the process. In the

FZ method, a local melted zone produced by radiofrequency

field is slowly passed along a polycrystalline rod. The seed


38

crystal is used at one end in order to launch the single

crystalline growth. Impurities in the molten zone tend stay in

the molten zone rather than being incorporated into the

solidified region, hence allowing an extremely pure single

crystal region being left after the molten zone has passed. The

largest Si ingots, produced using CZ method, are 300-450 mm in

diameter and 1 to 2 meters in length today. The diameters of the

FZ ingots are typically less than 200 mm due to the surface

tension limitations encountered during the crystallization

process (Franssila 2004). Thin Si wafers with a typical thickness

of 0.25-1.0 mm, are cut from these ingots and then polished to a

very high smoothness. Silicon can be doped with impurity

atoms such as boron or phosphorous, i.e. changing it into n-type

or p-type Si, in order to enhance the electrical conductivity.

(Madou 2002, Franssila 2004)

In addition to the excellent semiconductor properties of Si, it

has also superb mechanical properties, enabling the design of

MEMS structures (Petersen 1982). Silicon micromachining

techniques are established and widely available. The major

drawbacks of Si for bio-MEMS applications include its limited

biocompatibility, unfavourable mechanical properties (rigidity,

fragility) as well as the relatively high material and processing

cost (Cheung 2007), which makes it less attractive for use in

disposable biomedical devices.

2.2.2 Glass

In spite of the limited number of micromachining techniques

available for glass substrates (compared to Si), this material is

used in bio-MEMS applications due to some unique properties,

most notably optical transparency. Glasses with a wide variety

of compositions can be used. Fused silica (pure amorphous SiO2)

and borosilicate (e.g. Pyrex®) are examples from commonly

used glass materials (Voldman et al. 1999).

2.2.3 Polymers

Polymers are very attractive for bio-MEMS applications due to

their suitable mechanical properties, enhanced biocompatibility,


39

optical transparency, ability to modify their bulk and surface

properties in order to improve their functionality, and the

availability of rapid prototyping and mass production processes

(e.g. injection molding, hot embossing) (Wilson et al. 2007).

These issues enable the fabrication of low-cost disposable

microdevices for clinical applications. Polydimethylsiloxane,

polyimide, epoxy resins and parylene are some examples of

widely utilized polymers in bio-MEMS (Seymour et al. 2009,

Cheung 2007, Saliterman 2006).

Polyimide has a long history of use in microelectronics e.g. as

an encapsulation and stress buffer material on semiconductor

chips as well as a substrate material in flexible printed

boards/cables (Ghosh & Mittal 1996). It possesses numerous

desirable properties such as excellent resistance to solvents,

strong adhesion to metals and metal oxides and good dielectric

properties (Wilson et al. 2007). Recently, PI has been introduced

as a substrate/insulating material in numerous bio-MEMS

applications, such as neural interfaces (Boppart et al. 1992,

Rousche et al. 2001, Hollenberg et al. 2006, Molina-Luna et al.

2007, Takahashi et al. 2003, Cheung et al. 2007, Owens et al.

1995, Stieglitz 2001, Mercanzini et al. 2008, Patrick et al. 2008,

Spence et al. 2007) due to its desirable mechanical and dielectric

properties and good biocompatibility (Wilson et al. 2007,

Stieglitz et al. 2000, Richardson et al. 1993).

Polydimethylsiloxane, also familiar as a soft lithographic

stamp material, is known to be biocompatible and it is approved

for use as an implanted material in medical devices, i.e.

approved by the Food and Drug Administration (FDA). It has

been used as a biomaterial in catheters, pacemakers, and ear and

nose implants (Visser et al. 1996). PDMS is a highly flexible and

optically transparent material and it is permeable to gases. It is

also amenable to different surface modifications, which can be

used to tailor the biochemical functionality of the PDMS

(McDonald & Whitesides 2002, Mata et al. 2005). PDMS can be

conformed to submicron features to develop different

microstructures (Xia & Whitesides 1998) such as microfluidic

components (Ng et al. 2002, Folch et al. 1999) for bio-MEMS.


40

An epoxy-based negative photoresist SU-8 is one of the most

widely used thick-film photoresists in MEMS (Lorenz et al. 1997,

1998, Wilson et al. 2007). It has been utilized as a structural

material because it can be produced with a wide range of

thicknesses from below 1 µm to even more than 1 mm (Lorenz

et al. 1998). Recently, it has been employed in numerous bio-

MEMS applications, including analytical microfluidic systems

(Tuomikoski 2007) and neural sensors (Hollenberg et al. 2006,

Tijero et al. 2009, Altuna et al. 2010). Parylene is a thermoplastic

polymer that can be vapour-deposited at room temperature to

create pin-hole free, optically transparent barrier coatings that

are stress-free, chemically and biologically inert, and minimally

permeable to moisture (Saliterman 2006, Seymour et al. 2009).

2.3 ADVANTAGES OF MICROFABRICATION

Recent developments of MEMS and associated nanotechnology

represent great opportunities in the design of miniature, smart,

and low-cost biomedical devices that could revolutionize

biomedical research and clinical practice (Urban 2006,

DiLorenzo & Bronzino 2008, Nuxoll & Siegel 2009). Micro- and

nanotechnology can be used either to improve the performance

of an existing device or to enable development of an entirely

new device. The bio-MEMS devices have many advantages over

their macroscopic counterparts. The MEMS techniques enable

development of small size devices that may be easier to use (e.g.

portable and hand-held devices), or they can be truly innovative

(e.g. implantable or even injectable devices) (Saliterman 2006,

Grayson et al. 2004). The small size often saves on the costs of

reagents, time and money (Voldman et al. 1999). For example,

portable ‚point-of-care‛ hematological devices and test kits

permit physicians to diagnose the patient’s condition more

rapidly (Betancourt & Brannon-Peppas 2006). Diagnostic bio-

MEMS devices, also known as lab-on-a-chips or micro-total

analysis systems, can be used to detect cells, microorganisms,

viruses, proteins, deoxyribonucleic acid (DNA) and related


41

nucleic acid (Bashir 2004, Guber et al. 2004). In general,

miniaturization offers several advantages over conventional

analytical methods. Smaller sample volumes enable reducing

assay cost and less waste disposal (Voldman et al. 1999). Due to

the low sample volume, short diffusion distances and fast

heating, less time is required for diagnostics (Saliterman 2006).

Furthermore, miniaturized parallel operation allows high-

throughput analysis, which plays an important role in genomic

research and drug discovery (Bashir 2004). On the other hand,

miniaturization of devices increases their surface area to volume

ratio, leading often to situations where surface effects dominate

volume effects. The larger surface area is essential in some

applications e.g. to ensure heat removal avoiding adverse heat

effects when high electric fields are used (Voldman et al. 1999).

In addition, MEMS fabricated electrodes and sensor systems

allow measurements and stimulations with higher spatial and

temporal resolution than with conventional macroscale counter-

parts, and hence enable more precise biomedical research or

clinical diagnostic and therapy (DiLorenzo & Bronzino 2008,

Stieglitz & Mayer 2006a).

Most MEMS fabrication processes can be performed

simultaneously on many, even thousands of devices, similarly

to the case of manufacturing of microchips on Si wafers. This

kind of batch processing enables high volume manufacturing at

low unit cost (Judy 2001). Processes are very reproducible,

minimizing variations between objects that easily arise among

individually constructed devices. Lastly, polymers and other

cheap materials can be utilized to provide bio-MEMS devices,

making feasible the manufacturing of disposable devices.

Disposable devices are particularly crucial when handling blood

and other biological fluids that may contain hazardous

substances such as human immunodeficiency virus or hepatitis

virus. Disposable devices also eliminate the risk of cross-

contamination and subsequent analysis errors which are

common in re-used devices (Wang & Soper 2007).


42

2.4 BIO-MEMS APPLICATIONS

During recent years, the bio-MEMS technology has been under

intensive research and development activity. Numerous bio-

MEMS applications have been implemented in the fields of cell

and molecular biology, pharmaceutics and medicine, including

the areas of therapeutics, diagnostics and tissue engineering

(Voldman et al. 1999, Bashir 2004, Cheung & Renaud 2006,

Nuxoll & Siegel 2009, Saliterman 2006, Betancourt & Brannon-

Peppas 2006, Grayson et al. 2004). Although some applications,

such as blood analysis cartridges, cochlear implants and deep

brain stimulators (DBS) have been already commercialized, the

vast majority of bio-MEMS objects are under research or

undergoing clinical trials (Stieglitz & Mayer 2006a, DiLorenzo &

Bronzino 2008). On the other hand, most of the commercially

available systems are designed for in vitro diagnostics. The

development cycle from the lab prototype to commercial

manufacturing of implantable devices is long, perhaps even 10-

20 years, due to the extremely challenging environment inside

the human body as well as strict requirements to pass the

approval process (classified as FDA class III devices) (Ratner

1996). Examples of implantable bio-MEMS devices are listed in

Table 1. A more detailed description of the electrodes used in

neural prosthesis is given in Section 4 of the thesis.


43

Table 1: Examples of implantable Bio-MEMS devices.

Device Brief description of its function Manufacturer(s) General

reference(s)

Cardiac

pacemakers

Regulation of the heart beating by delivering electrical impulses to

the electrodes in contact with

heart muscles

Boston Scientific, USA; Medtronic,

USA; St. Jude

Medical, USA

Das et al. 2009

Cochlear

implants

Electrical stimulation of auditory

nerve through microelectrodes implanted in the inner ear to

restore hearing

AllHear, USA;

Advanced Bionics, USA; Cochlear,

Australia; MED-

EL, Austria;

Neurelec, France

Loizou 1999,

Eddington 2008, Wilson & Dorman

2008a,b

Deep-brain

stimulators

Through intracortically implanted electrodes electrical stimulation is

delivered to targeted regions of

the brain to treat severe neurological disorders, such as

Parkinson disease

Medtronic, USA Kumar et al. 1997, Kern &

Kumar 2007,

Richardson 2008

Spinal cord

stimulators

Electrical stimulation of spinal cord through electrodes implanted in

epidural space in order to treat

intractable pain and motor

disorders

Advances Neuromodulation

Systems, USA;

Medtronic, USA

Waltz 1997, Mailis-Gagnon et

al. 2004, Burton

& Phan 2008

Vagus nerve

stimulators

Electrical stimulation of vagus

nerve in the neck via an implanted

lead wire to treat certain types of intractable epilepsy and major

depression

Cyberonics, USA Ardesch et al.

2007, Boon et al.

2001

Drug delivery

silicon chips

Controlled drug release from microfabricated Si reservoirs by

applying electrical potential

Microchips, USA Santini et al.

1999, 2000

Visual

prosthesis

Electrical stimulation of neural tissue of blind patients to restore

vision. Different approaches in

which microelectrode arrays are implanted on or under the retina,

around the optic nerve, and on or

in the visual cortex are under

development

Second Sight, USA; Retina

Implant AG,

Germany

Humayun et al. 2003, Stieglitz

2009, Greenberg

2008


44

45

3 Origin of bioelectric

signals

Bioelectric signals are generated as a result of electrochemical

activity in certain type of cells, referred to as excitable cells

which are the basic components of nervous, muscular and

glandular tissue (Malmivuo & Plonsey 1995). An electrical

potential difference, i.e. the transmembrane potential, exists

between the internal and external environment of the cell (Hille

1992). When appropriately stimulated to cross a limiting

threshold, the excitable cell will generate an action potential due

to a flow of charged ions across the permeable cell membrane

with the ions passing through ion channels. These action

potentials are transferred from one cell to adjacent cells via their

cellular extensions, i.e. axons and dendrites due to synaptic

transmission (Tortora & Grabowski 2000). The activity of a

group of excitable cells (i.e. population activity) generates an

electric field that propagates through the biological tissue and

this can be measured by field-potential measurements with

surface electrodes attached to the surface of tissue, e.g.

cerebellum cortex, or skin (Nunez & Srinivasan 2006, McAdams

2006). The field-potentials commonly monitored in a patient are

those producing electrocardiograms (ECG) from the heart, EEG

from the brain and electromyograms (EMG) from contraction of

muscles.

3.1 EXCITABLE NERVE CELL

A nerve cell, also known as a neuron, is an excitable cell in the

nervous system whose main function is to receive, handle and

transmit cellular information via electrochemical phenomena.

Neurons are the basic components of the brain, the spinal cord

3 – Origin of bioelectric signals

46

and the peripheral nerves. They can be divided into two

categories: (1) principal cells that receive input from nearby or

distant neurons and project to neurons in other nuclei; (2)

interneurons that modulate the activity of principal cells in the

same structure. Principal cells are usually excitatory and use

glutamate as the neurotransmitter, whereas interneurons are

usually inhibitory and use gamma-amino butyric acid as the

neurotransmitter (Kandel et al. 2000). There are a huge variety

of different nerve cells with sizes in the range of 600 – 70 000

µm3 (Schadé & Ford 1973). Most nerve cells consist of three

main parts: a cell body and two kinds of cell extensions –

dendrites and axons. The cell body (soma) is similar to that of all

other cells containing the nucleus, mitochondria, endoplasmic

reticulum, ribosomes, and other organelles (Tortora &

Grabowski 2000). Most of the protein synthesis happens in this

part of cell. Dendrites are typically short and highly-branched

cellular extensions, through which the vast majority of signal

input to the neuron occurs (afferent signals). Due to numerous

branches (‚dendrite tree‛), a neuron may receive nerve impulses

from thousands of other neurons (Nunez & Srinivasan 2006).

The axon of a neuron is a single, cylindrical-shaped projection

which carries nerve impulses away from the cell body toward

the axon terminals passing the message on to other neurons.

Although axons are very thin, typically about 1-20 µm in

diameter, they can be very long (Malmivuo & Plonsey 1995).

The length of axons, e.g. the human sciatic nerves, may be more

than one meter long. The axons of many neurons are covered

with an insulating layer called the myelin sheath, which is

formed by either of two types of glial cells: Schwann cells or

oligodendrocytes which insulate peripheral neurons or those of

the central nervous system, respectively (Tortora & Grabowski

2000). The myelin sheath is not continuous but the gaps, known

as nodes of Ranvier occur at regular intervals, dividing the

sheath into short sections (Malmivuo & Plonsey 1995). This

myelination structure enables much faster conduction (i.e.

saltatory conduction) of nerve impulses so the impulse is

transported faster than in nonmyelinated axons.


47

3.1.1 Cell membrane and ion channels

The living cell is surrounded by the cell membrane which is an

approximately 5 nm thick phospholipid bilayer where the

hydrophobic tails of phospholipid molecules are pointing inside

the membrane whereas the negatively polarized, hydrophilic

heads of these molecules are directed outward of the membrane

(Alberts et al. 1994). The main functions of the cell membrane

are to form the boundaries for the intracellular components of

the cell, maintaining cell homeostatis and regulating the

incoming and outgoing substances (Tortora & Grabowski 2000).

The cell membrane is permeable to only specific types of ions

and molecules and thus the concept of selective permeability is

used to describe this property of the membrane. The excitability

of the membrane is based on embedded protein molecules,

which form tiny ion channels and allow the flow of sodium

(Na+), potassium (K+), calcium (Ca2+) and chloride (Cl-) ions

through the membrane (Hille 1992). Ion channels can be divided

into leakage channels and gated channels (Tortora & Grabowski

2000). The difference between them is that leakage channels are

always open, whereas gated channels open and close in

response to some kind of stimulus. The gated channels, on

which the excitability of neural cells is based, can be divided

into three groups according to the stimulus involved: (1)

voltage-gated ion channels, (2) ligand-gated ion channels and (3)

mechanically gated ion channels. The state (i.e. open or closed)

of voltage-gated channels is adjusted by a change in the

transmembrane potential. Ligand-gated ion channels open and

close in response to specific chemical stimuli including

neurotransmitters, hormones and ions, whereas mechanically

gated ion channels are regulated by mechanical stimuli such as

vibration, pressure or tissue stretching (Tortora & Grabowski

2000, Kandel et al. 2000).

3.1.2 Transmembrane potential and equilibrium potentials

The selective permeability of the cell membrane maintains ion

concentration difference between the inside and outside of the

cell membrane (i.e. concentration gradient). For mammalian


48

muscle, the K+ concentration of the intracellular space is 155

mmol/l, while that within the extracellular space is only 4

mmol/l (Hille 1992). In addition, Na+, Ca2+ and Cl- ionic

concentrations inside the cell differ greatly from the outside

(Table 2). To balance the positive and negative charges inside

the cell, there are also other negatively charged molecules,

which are too large to permeate through the membrane.

Table 2: Ion concentrations for mammalian muscle cells (Hille 1992).

Ion Intracellular concentration (mM/l)

Extracellular concentration (mM/l)

Na+ 12 145

K+ 155 4

Ca2+ 10-4 1.5

Cl- 4.2 123

This existing concentration gradient promotes the outflow of

K+ ions and inflow of Na+, Ca2+ and Cl- ions in order to balance

the ion concentrations. Potassium ions diffuse from inside to

outside, making the interior of the cell more negative than the

outside. Additionally, there is an active transportation of ions

against a concentration gradient. These Na-K pumps transport

two K+ ions into the cell for every three Na+ ions that the enzyme

pumps out of the cell, and thus there is a net loss of positive

charges within the cell (Hille 1992). The transmembrane

potential, Vm, is defined as a potential difference between the

internal and external environment of a cell. The resting

membrane potential corresponds to the value of Vm when the

cell is in the resting state in its natural, physiological

environment. In neurons, the typical resting potential is around

-70 mV (Tortora & Grabowski 2000). The membrane potential

that just balances the ion concentration difference, i.e.

equilibrium potential when no current flows across the

membrane, can be determined for each specific ion by using the

Nernst equation (derived by Water Hermann Nernst in 1888):


49

.ln

o

i

C

C

zF

RTV (2)

Here R is the gas constant (8.314510 J/mol x K), T is absolute

temperature, z is the valence of an ion, F is the Faraday’s

constant and Ci and Co are ion concentrations inside and outside

of the cell, respectively.

It can be easily calculated by using the ion concentrations

(Table 2) that the equilibrium potential for K+ ions at 37 °C is

approximately -98 mV. The corresponding values for Na+ and

Ca2+ ions are +67 mV and +128 mV, respectively. Since the

equilibrium potential of K+ is quite close but not same as the

value of the resting potential (typically -70 mV), at rest the cell

membrane is much more permeable to K+ ions than to other ions.

The Nernst equation can be extended to Golzman-Hodgkin-

Katz equation, which produces the equilibrium potential when

there are several types of ions in the intracellular and

extracellular space, and when the membrane is permeable to all

of them (Hille 1992). Then, the resting membrane potential can

be estimated by:

,ln

,,,

,,,

CliClNaoNaKoK

CloClNaiNaKiK

RESTcPcPcP

cPcPcP

F

RTV

(3)

where PK, PNa and PCl are membrane permeabilities of K+ , Na+,

Cl- ions, respectively.

3.2 SYNAPTIC POTENTIALS AND ACTION POTENTIALS

The transfer of information from one neuron to the next is based

on junctions between an axon and the neighboring cell, called

synapses. They can be divided into electrical and chemical

synapses (Tortora & Grabowski 2000). An electrical synapse is a

mechanical, electrically conductive link between two neurons

which can transmit nerve impulses very quickly (Hormuzdi et

al. 2004). Electrical synapses are common in the neural systems


50

that require the fastest possible response, such as defensive

reflexes. However, it is more common to operate with chemical

synapses. The end of the axon has branching terminals through

which neurotransmitter chemicals (e.g. dopamine, glutamate,

serotonin) packaged into presynaptic vesicles, are released into

a gap called the synaptic cleft in order to communicate with an

adjacent neuron (Kandel et al. 2000). Most synapses are between

axon and dendrite whereas also other combinations like axon-

to-axon and dendrite-to-dendrite exist (Tortora & Grabowski

2000). The width of synaptic cleft is only about 20-40 nm

(Hormuzdi et al. 2004). Due to the small volume of the cleft, the

neurotransmitter concentration can be raised and lowered

rapidly. There are certain types of receptor molecules located on

the cell membrane of the postsynaptic cell which after being

occupied by neurotransmitter molecules, respond by opening

nearby ion channels in the membrane, allowing charged ions to

rush in or out. As a result, the local transmembrane potential of

the cell is changed and thus is referred to as the post-synaptic

potential. The result can be either excitatory (depolarizing) or

inhibitory (hyperpolarizing) depending on the amount and type

of neurotransmitter as well as the type of receptor molecules

(Kandel et al. 2000). In an excitatory reaction, a net movement of

positive ions into the cell changes the transmembrane potential

making it less negative. Hyperpolarization is the exact opposite

to depolarization, in which the change of Vm occurs in the

negative direction. If the depolarizing response is great enough,

the cell membrane is able to produce a characteristic bioelectric

impulse called an action potential (Fig. 4) (Tortora & Grabowski

2000). At the beginning of depolarizing stimulus, the voltage-

gated Na+-selective ion channels become opened, allowing the

flow of Na+ ions from outside to inside. The inside of the cell

becomes more positive reaching a potential value of

approximately +30 mV. Subsequently, more K+-selective ion

channels will open allowing the flow of K+ ions from inside to

outside. The flow of K+ ions returns the transmembrane

potential to its resting value. The duration of the action potential

is around 1 millisecond (Tortora & Grabowski 2000).


51

1 2 3 4 5 Time (ms)

Tra

nsm

em

bra

ne

po

ten

tia

l(m

V)

+30

0

-70

Resting stateRefractory

period

Figure 4: The shape of an action potential. When a stimulus depolarizes the membrane

beyond a threshold, an action potential is generated. At the end of depolarization

phase, a membrane potential of 30 mV is reached. The flow of potassium ions returns

the membrane potential to its resting value (-70 mV) during a short refractory period.

3.3 RECORDING ELECTRICAL ACTIVITY OF THE BRAIN

The human nervous system is a complex, highly organized

network of billions of neurons and even more neuroglia (Tortora

& Grabowski 2000). It is composed of two sub-systems: the

central nervous system (CNS) and peripheral nervous system

(PNS). Anatomically, the CNS contains the cerebrum,

cerebellum, brainstem and spinal cord (Heimer 1995). The brain

stem is a structure through which nerve fibers relay action

potentials in both directions between spinal cord and higher

brain centers. The cerebrum is divided almost equally into two

halves, which both have subdivisions of the frontal, temporal,

parietal and occipital lobes. The cerebral cortex, about 2-4 mm

thick folded structure, forms the outer portion of the cerebrum

(Tortora & Grabowski 2000). The surface area of the cerebral

cortex varies from 1600 to 4000 cm2 and it contains perhaps as


52

many as 10 billion neurons (Nunez & Srinivasan 2006). Those

neurons are strongly interconnected. The surface of a large

cortical neuron may be covered by as many as 105 synapses that

transmit inputs from other neurons (Tortora & Grabowski 2000).

Post-synaptic potentials in thousands of synchronously active

pyramidal neurons produce an electric field propagating

through the biological tissue and these can be measured on the

scalp or on the surface of brain (subdural electrodes) or even

inside the brain (probe electrodes) with an amplifier through

appropriate electrodes (Nunez & Srinivasan 2006, Niedermayer

& Lopes da Silva 2005). Scalp-EEG is a standard method in

clinical neurophysiology to study the function of the brain. Both

spontaneous activity and evoked potentials (EPs) can be

measured. Spontaneous scalp-EEG is routinely utilized in many

clinical cases, ranging from diagnosis of epilepsy to the analysis

of sleep and the depth of anaesthesia. EPs are produced as a

response to given sensory stimuli. The EPs can be classified

according to the applied stimulus into three categories:

somatosensory (SEP), visual (VEP) and auditory (AEP) evoked

potentials (Niedermayer & Lopes da Silva 2005). The stimulus

evokes electrical activity with a very short latency (delay) period

over a limited area of the cerebral cortex. For example, VEPs are

detected from the occipital region when a visual stimulus is

presented, whereas AEPs are detected from the temporal region

when an auditory stimulus is presented. Evoked potential

recordings can be utilized to localize lesions, or monitor a

sensory pathway during surgical procedures (Chiappa 1997).

Stimulus is usually applied many times (even thousands), and

then signal averaging is undertaken to achieve clean EP curves

with a reduced effect of noise (Nunez & Srinivasan 2006).

In clinical practice, scalp-EEG is typically measured using the

standard international 10-20 electrode placement system in

which 21 electrodes are attached on the scalp with exact

intervals and potential differences between electrodes is

recorded. The amplitudes of the scalp EEG varies from 10 to 100

µV (Niedermayer & Lopes da Silva 2005). The frequency band

from 0.16 to 100 Hz is commonly adequate, but sometimes a


53

much lower frequency response, down to direct current (DC)

(Vanhatalo et al. 2003) or much higher (e.g. up to 3 kHz) is

useful (Curio 2005). The scalp-EEG possesses good temporal

resolution on a sub-millisecond scale, but the spatial resolution

is poor. Spatial resolution can be improved by increasing the

electrode numbers. Nowadays, multi-channel EEG caps with up

to 256 electrodes are applied in experiments aimed at obtaining

detailed information about brain sources (Väisänen 2008).

Another choice to improve the source localization is to place the

electrodes closer to the brain tissue. Both intracortical depth

electrodes and subdural strip electrodes placed on the surface of

cortex are used in intracranial recordings of local field potentials

and occasionally spikes (Grill 2008). Compared to scalp-EEG,

intracranial recordings are worthwhile in terms of their

increased spatial resolution, signal amplitude level (0.5 – 4 mVpp

vs. 10 – 100 µVpp), and reduced noise level, meaning that one

can acquire more detailed maps directly of the brain (Lachaux et

al. 2003, Niedermayer & Lopes da Silva 2005, Leuthardt et al.

2008). However, the use of penetrating electrode probes in

humans is limited to the areas destined for surgical resection

due to the risk of damaging brain tissue. Flexible subdural strip

and grid electrodes are utilized as a less invasive alternative in

some clinical cases, e.g. in mapping cerebellum cortex in patients

undergoing epilepsy or brain tumor surgery. The signals

recorded with subdural electrodes are considerably higher in

amplitude because they are much closer to the source of the

activity, and are separated from it by only a relatively high

conductivity media (cerebrospinal fluid, brain parenchyma)

(Nair et al. 2008, Leuthardt et al. 2008). The resistivity of the

skull is estimated to be even 80 times higher than that of brain

tissue (Rush & Driscoll 1968) that causes a decline in the signal

amplitude and signal blurring in scalp-EEG recordings.

Moreover, recordings made by subdural electrodes are almost

free of artifacts (e.g. not affected by electrical signals originating

from muscles) that are seen in scalp-EEG, and thus they yield a

much higher signal-to-noise ratio (Nair et al. 2008).


54

55

4 Bioelectrodes

Biomedical electrodes are extensively utilized in numerous

biomedical applications (McAdams 2006). Firstly, they are used

to record bioelectrical events such as the ECG, EEG and EMG

that intrinsically arise due to electrical activity of excitable cells.

Secondly, bioelectrodes are utilized in applying electrical

impulses to the body for treating neurological pain in many

therapeutic methods such as in transcutaneous electrical nerve

stimulation or deep brain stimulation. Deep brain stimulators

consist of intracortically implanted electrodes through which a

high frequency electrical stimulation is delivered to a targeted

region of the brain. DBS is used in treating several neurological

disorders, e.g. Parkinson’s disease (Kumar et al. 1997, Kern &

Kumar 2007, Perlmutter & Mink 2006). Thirdly, they are used in

alternating current impedance characterization of body tissues.

These methods such as electrical impedance plethysmography

and electrical impedance tomography (Cheney et al. 1999) do

not monitor intrinsic biosignals emanating from the body, but

inject small alternating currents/voltages and record the

resultant voltages/currents. The electrical properties of the body

can then be determined. Finally, bioelectrodes are utilized in

electrotherapy by applying electrical potentials in order to

facilitate the transdermal delivery of ionized molecules for

therapeutic effect (iontophoresis). Moreover, recent activity on

neuroengineering research has focused on implantable neural

prosthetics to help restore the functions of the damaged nervous

system (DiLorenzo & Bronzino 2006, Stieglitz & Mayer 2006b).

The most successful examples of commercialized neural

prosthesis are cardiac pacemakers and cochlear implants.

Cochlear implant is a surgically implanted microelectronic

device that utilizes an electrode array inserted into the scala

tympani of the cochlea to stimulate the auditory nerve through

4 – Bioelectrodes

56

the bone, enabling the deaf to hear sounds (Loizou 1999,

Federspil & Plinkert 2004, Wilson & Dorman 2008a,b).

4.1 IMPLANTABLE ELECTRODES

The commercially available clinic subdural strip and grid

electrodes (Fig. 5) typically consist of a number of conductive

metal such as Pt, stainless steel (SS), and silver (Ag) disks, which

are partially embedded in a thin sheet of biomedical grade

silicone rubber (e.g., AD Tech Medical Instrument Corp., Racine,

WI, USA; PMT Corp., Chanhassen, MN, USA). Typical diameter

of disk contacts is about 2-3 mm, and the distance between

adjacent electrodes around 10 mm.

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

Despite major advances in microsystem technology and its

great potential in the development of novel high-density MEAs,

most human intracranial electrodes used to record brain activity

are still handcrafted from Pt foils and silicone rubber sheet and

therefore are quite bulky (see Fig. 5). Miniaturized neural

microelectrodes are extensively developed approaches in

neurophysiological (animal) studies aimed at understanding

4 - Bioelectrodes

57

behaviour and function of the brain at the tissue and cellular

level, but their human counterparts are still mainly under

investigation or at best are undergoing clinical trials.

Already over forty years ago it was discovered that the

photolithographic, Si etching and thin-film deposition

techniques originally developed for the production of integrated

circuits would be also applicable for providing novel improved

recording and stimulating electrodes for experiments with

animal models (Wise 2005). Compared to the traditional

approach of single microwires or microwire bundles (Nicolelis

et al. 1997, Williams et al. 1999), microfabricated electrodes

offered many advantages including precise control of the

electrode sizes and separations between electrodes as well as a

high degree of reproducibility in physical, chemical and

electrical characteristics (Rutten 2002). Microfabrication is also a

high yield, low-cost process once the design and processing

steps have been optimized. Much of the pioneer work of Si-

based microelectrodes has been done in three research centers in

the USA, i.e. at the University of Stanford, University of

Michigan, and University of Utah. Today Utah arrays (Campbell

et al. 1991, Rousche & Normann 1998) and Michigan probes

(Wise & Angell 1975, Najafi & Wise 1986, Vetter et al. 2004) are

perhaps the best-known commercially available Si-based

microelectrodes. A three dimensional Utah electrode array (Fig.

6a) consists of one hundred 1.5 mm long Si needles that project

out from a 4 mm x 4 mm substrate (Campbell et al. 1991,

Normann et al. 1999). The sharpened tip of each Si needle is

coated with Pt. An advanced version of Utah array, i.e. Utah

slanted electrode array, has also been developed (Fig. 6b). In this

model, the length of Si needles can be increased from 0.5 to 1.5

mm permitting focal excitation of nerve fibers at different

depths (Badi et al. 2003). Utah arrays have been used in studies

of parallel information processing by the CNS at the visual and

auditory cortex (Rousche & Normann 1998), and the control of

muscle force and limb position by the PNS (Normann 2007). The

typical Michigan probes are single-shank or multi-shank Si

penetrating systems (Fig. 6c) where the electrodes are placed

4 – Bioelectrodes

58

along the length of the shank allowing the measurement of

neuronal activity at various depths of the brain tissue. Michigan

probes have been successfully used in a number of neuroscience

experiments throughout the CNS, PNS and even cardiac muscle

(Vetter et al. 2004, Kipke et al. 2003, Wise 2005).

Figure 6: Silicon–based neural interfaces: (a) Utah array; (b) Slanted Utah array.

These electrode arrays which both contain 100 microneedle-shaped electrodes are

designed to be implanted in cerebral cortex or peripherally. (c) Michigan multi-shank

probe contains active electronics integrated into the array. Reprinted by permission

from Macmillan Publishers: Nature Clinical Practice Neurology (Normann 2007),

copyright (2007).

Although rigid microwires and Si-based microelectrodes

have long been used for recording of biosignals, the mechanical

mismatch between the rigid electrode and soft biological tissue

may evoke many adverse effects such as tissue damage,

inflammatory reactions and scar formation (Polikov et al. 2005).

Thus, there has been a growing interest in developing polymer-

4 - Bioelectrodes

59

based (e.g. polyimide, benzocyclobutene, silicone rubbers and

parylene-C) implants that could be sufficiently flexible to mimic

biological tissue and to reduce adverse tissue reactions (Cheung

2007). The flexibility of polymer may decrease the mechanical

mismatch between nervous tissue and the rigid implant

reducing the risk of tissue damage and inflammation reactions.

PI is one of the most widely studied polymers due to its suitable

mechanical properties, biocompatibility and stability in wet

microfabrication processes (Rubehn & Stieglitz 2010, Stieglitz et

al. 2000). Several PI-based microelectrode constructions have

been developed in the fields of neural engineering, including

devices for in vitro recordings from brain slices (Boppart et al.

1992), cortical surface field potential recordings (Owens et al.

1995, Hollenberg et al. 2006, Takahashi et al. 2003, Hosp et al.

2008), intracortical multiunit neural activity recordings (Rousche

et al. 2001, Cheung et al. 2007, Mercanzini et al. 2008) and for

action potential recording in nerve and muscle tissues in vivo

(González & Rodríguez 1997, Rodríguez et al. 2000, Spence et al.

2007). PI has also been applied for different types of sieve and

cuff electrodes for interfacing with regenerating peripheral

nerves (Stieglitz et al. 1997, 2000, Stieglitz 2001).

4.2 MATERIAL REQUIREMENTS FOR IMPLANTABLE

ELECTRODES

The materials used in implantable neural electrodes and

prostheses must fulfill strict requirements in order to ensure

optimal performance and longevity. According to Geddes and

Roeder (2003), there are four criteria that must be considered

when selecting a material for implantable use: (1) tissue

response; (2) allergic response; (3) electrode-tissue impedance;

and (4) radiographic visibility. An ideal implanted electrode

and/or its insulation should be biocompatible without triggering

any vigorous local or generalized host response or allergic

reactions. The electrode-tissue impedance must be stable and

low enough to enable high-resolution measurements. Moreover,

4 – Bioelectrodes

60

the materials should be radiographically visible and be

compatible with magnetic resonance imaging.

It has been concluded in many studies that in spite of the

extremely challenging processes used in fabrication of

miniaturized electrodes, electrode failures due technical reasons,

such as short-circuits and broken signal transmission lines, are

rare, and the stable, long-term functioning of neural implant is

mainly determined by the tissue response to chronically

implanted electrodes (Polikov et al. 2005, Nicolelis et al. 2003,

Rousche & Normann 1998, Williams et al. 1999). A general

immune activation of the brain in response to the presence of a

foreign body implant has been commonly believed to be the

main reason for signal deterioration of chronically implanted

electrodes (Nicolelis et al. 2003, Edell et al. 1992, Schmidt et al.

1993). The biocompatibility aspects of electrode materials will

therefore be very crucial in the development of new generation

chronic implantable devices.

4.2.1 Electrode materials

Recording bioelectrodes are functional elements that convert the

ionic current around the electrode site in the body into electron

current in electronic circuitry. Electrodes can be roughly divided

into polarizable and non-polarizable electrodes. Perfectly

polarizable electrodes are those in which no actual charge

crosses through the electrode-electrolytic interface, but changes

the charge distribution within the solution near the electrode

when a current is applied. This kind of electrode acts like a

capacitor, and only the displacement current crosses the

interface. In contrast, perfectly non-polarizable electrodes allow

the current to pass freely through the electrode-electrolytic

interface without changing the charge distribution in the

electrolytic solution. However, neither of these ‚perfect‛

electrode types can be fabricated, but some practical electrodes

can come close to possessing their features. (McAdams 2006,

Neuman 1978)

4 - Bioelectrodes

61

A silver-silver chloride (Ag-AgCl) electrode is an almost

perfectly non-polarizable electrode. Due to many desirable

properties, such as small and stable offset potential, low contact

impedance and low intrinsic noise, this type of electrode has

been found to be an excellent electrode material for surface (skin)

biopotential recordings (McAdams 2006). Unfortunately, it Ag-

AgCl cannot be used as an implantable electrode because of the

rapid dissolution of Ag. Moreover, AgCl as well as pure Ag has

a toxic effect on neural tissue (Yuen et al. 1987).

Noble metals such as platinum, gold, iridium, rhodium and

palladium are commonly used highly polarizable electrode

materials in neural applications due to their excellent corrosion

resistance and biocompatibility (Geddes & Roeder 2003,

McAdams 2006). In the extensive study carried out by Stensaas

and Stensaas (1978), twenty-seven different potential neural

electrode materials were tested by implanting them into the

cortices of rabbits. After 30 days, the histological examination

revealed no adverse reaction in response to gold, platinum and

tungsten. Unfortunately, noble metals usually tend to give rise

to high electrode impedances and unstable potentials

(McAdams 2006). Recently, different kinds of micro- and

nanostructuring techniques have been investigated to increase

the effective surface of noble metals and subsequently reduce

the electrode-electrolyte interface impedance. Pt has been the

most widely used material in implantable electrodes. Clinically,

it has been utilized as an electrode material in various

applications such as in cardiac pacemakers, cochlear implants

and subdural strip and grid electrodes. Pt suffers from corrosion

only slightly in both recording and stimulating use. Due to the

mechanical softness of Pt, it is sometimes alloyed with iridium

(Ir) to achieve electrodes for applications requiring high

mechanical strength (McAdams 2006). Gold is also often used in

recording electrodes, but it suffers from corrosion if used for

stimulating purposes. On the contrary, Ir is mainly utilized as a

stimulating electrode. The electrode properties of Ir can be

further developed by electrochemical activation that leads to the

adhesion of a porous and very stable oxide layer on the

4 – Bioelectrodes

62

electrode surface. Activation provides much higher charge

injection levels and also reduces the electrode contact

impedance due to the porous microstructure (Cogan 2008, Blau

et al. 1997).

Titanium nitride (TiN) is another novel candidate suitable for

use in stimulating electrodes, and it is already being utilized in

cardiac pacemakers. The high charge delivery capacity of TiN is

based on its high effective surface area due to the columnar

structure of this material (Cogan 2008). Indium tin oxide (ITO) is

a widely used electrode material in electrochemistry (Huang et

al. 2009), but so far its use in neurophysiology is limited to in

vitro MEAs for recording electrical signals from neurons. The

biocompatibility of ITO is not completely understood although

recent in vitro investigations on cellular responses (Bogner et al.

2006) and protein adsorption on ITO (Selvakumaran et al. 2008)

have displayed promising results.

The impedance of an electrode-electrolyte interface depends

on the species of metal, the type of electrolyte used, the surface

area of electrode, and the temperature (Geddes & Roeder 2001,

2003). Magnitude of impedance decreases with expanding

surface area and increasing roughness. It also decreases with

increasing frequency and increasing current density. Several

different equivalent circuit models for the electrode-electrolyte

have been suggested (reviewed by Geddes 1997). One of the

most sophisticated models is presented in Fig. 7a. There are two

identical electrodes in contact with electrolyte. Electrode

polarization impedance consists of Warburg (polarization)

resistance Rw and capacitance Cw. The Faradic resistance, Rf as a

parallel with Warburg components, provides a route for direct

current to pass through the interface. The resistance of the

electrolyte solution is marked with Rs. Both Rw and the

polarization reactance (Xw = 1/2πfCw) are frequency-dependent:

they decrease as ( f/1 ) with increasing frequency ( f ). At a

sufficiently high frequency, the impedance between two

electrodes in saline approaches asymptotically the value of Rs

(Fig. 7b).

4 - Bioelectrodes

63

A B

Frequency (Hz)

Rs

Rw1

Rw2

Cw1

Cw2

Rf2

Rf1

Rf1+ Rf2 + Rs

Rs

1

2

Z

Imped

ance

|Z|(Ω

)

Figure 7: (a) The equivalent circuit model for two identical bioelectrodes immersed in

saline. (b) Typical electrode-electrolyte interface impedance curve measured between

terminals 1 and 2 as a function of frequency. The curve approaches asymptotically the

resistance value of the solution.

Due to the inverse relationship between electrode impedance

and its surface area, it is advantageous to develop

microelectrodes with high nanoscale surface topography since

in this way it will be possible to achieve a high effective surface

area without changing their geometrical dimensions (Kotov et al.

2009). Different techniques are used to increase the effective

surface area including surface roughening, micro- and nano-

patterning and chemical modification. In micro/nanopatterning

techniques, different etching and machining techniques are

utilized to form different micro/nanostructures, e.g. grooves and

wells onto the surface of the substrate upon which the electrode

material can then be deposited using traditional thin film

techniques (Cheung 2002). Some examples of widely used

surface modification techniques used to enhance nanoscale

roughness includes Pt electroplating (Pt black) (de Haro et al.

2002) wet etching of Au and activation of Ir (Cogan 2008, Blau et

al. 1997) to form a nanoporous iridium oxide layer as already

discussed.

4.2.2 Substrate materials

The surface area of substrate/encapsulation material is usually

much larger compared to areas of the active recording sites.

Therefore, biocompatibility aspects of substrate and insulator

materials are very crucial in the development of neural

4 – Bioelectrodes

64

electrodes and prosthesis which require long-term stability and

functionality. These materials need to be biocompatible,

biostable and possess good dielectric properties. An artificial

material or implant can be considered to be biocompatible if (1)

it does not evoke a toxic, allergenic or immunologic reaction, (2)

it is not harmful to enzymes, cells and tissues, (3) it does not

evoke thrombosis or tumors, and (4) it remains for a long term

within the body without encapsulation or rejection (Heiduschka

& Thanos 1998). On the other hand, biostability means that the

implant material must not be susceptible to attack by biological

fluids or any metabolic substances. Unfortunately, all implanted

artificial materials in contact with neural tissue have an

inconvenient tendency to induce significant glial scar tissue

formation (Polikov et al. 2005). In response to implantation, glial

cells such as astrocytes become activated and they transform

into reactive phenotype which produces much larger size glial

fibrillary acid protein filaments. Cell proliferation and migration

capacity of astrocytes is also enhanced. The resultant scar tissue

formation causes serious impairment of implant performance

due to decreased local density of neurons and the formation of

an encapsulation layer that increases electrode impedance and

decreases the signal amplitudes. It has been noted that the

astrocyte response to electrode implantation can be divided into

two phases, the early response that occurs immediately after

electrode implantation and the long-term chronic response.

Within a few hours after implantation, the number of glial cells

is significantly enhanced in the area surrounding the Si probes

(Szarowski et al. 2003, Turner et al. 1999). The chronic response,

starting approximately one week after implantation (Norton et

al. 1992), consists of a compact glial scar tissue formation

surrounding the electrode, which ultimately isolates the

microelectrodes from neurons increasing the electrode contact

impedance and causing signal amplitude deterioration (Edell et

al. 1992, Nicolelis et al. 2003, Schmidt et al. 1993).

Silicon with its common insulators, SiO2 and silicon nitride

(Si3N4), have been the most widely used materials in neural

interfaces as well as in the bio-MEMS technology overall. In

4 - Bioelectrodes

65

spite of numerous desirable features of Si, there are at least two

major weaknesses related to its use in neural applications.

Firstly, there is a huge mismatch between the mechanical

properties of the brain tissue and Si. The values of Young’s

modulus for Si and brain tissue are 170 GPa and 3 kPa,

respectively. The rigidity of Si may cause many adverse effects

such as tissue damage, inflammation reactions and scar

formation (Polikov et al. 2005). Secondly, the results from

studies related to cytocompatibility of Si are somewhat

contradictory (Kotzar et al. 2002). Some studies have shown that

Si is a rather poor substrate for cell adhesion and even slightly

cytotoxic (Liu et al. 2007, Madou & Tierney 1993), whereas some

others have claimed that Si is almost as cytocompatible and

immunogenic as tissue culture polystyrene (Ainslie et al. 2008).

This uncertainty could limit the much-anticipated integration of

Si-based neural implants into the human body.

There is a growing interest in developing polymer-based

implant materials that could be flexible enough to mimic

biological tissue and to reduce mechanical damage, but not

evoking any adverse tissue reactions. A wide variety of polymer

materials, such as polyimide, benzocyclobutene, silicone rubbers,

parylene-C and epoxy resins have been investigated as substrate

and encapsulating materials in implantable neural interfaces

(Cheung 2007, HajjHassan et al. 2008, Polikov et al. 2005).

Polyimide possesses numerous desirable properties when

used as a neural implant material. It has appropriate mechanical

properties (Rubehn & Stieglitz 2010) and its biocompatibility has

been demonstrated in many studies in vitro and in vivo (Stieglitz

et al. 2000, Richardson et al. 1993, Seo et al. 2004, Lee et al. 2004,

Yuen et al. 1987, Yuen & Agnew 1995). Stieglitz et al. (2000)

evaluated the cytotoxicity of three commercial PI grades

(Pyralin PI 2611, PI 2556, PI 2566; HD Microsystems GmbH, Bad

Homburg, Germany) and reported excellent biocompatibility for

PI 2611 and PI 2556 and good results for PI 2566. This last PI,

since it is fluorinated, differs from the others with respect to its

chemical structure (Kanno et al. 2002). In another study (Lee et

al. 2004), it has been shown that fibroblasts attach, spread out

4 – Bioelectrodes

66

and grow on the PI surfaces in a manner corresponding to the

behaviour of control cells growing on the surface of polystyrene,

i.e. the plastic that is used in cell culture well plates and flasks.

Furthermore, Yuen et al. (1987) demonstrated that subdural

implantations of two common insulators, parylene-C and PI

(DuPont PI-2555), into the cerebral cortex of the cat evoked only

minimal tissue reactions after 16 weeks. Tissue responses to PI

and parylene-C were comparable to pure Pt (used as a negative

control) whereas Ag-AgCl (positive control) triggered chronic

inflammatory reactions (Yuen et al. 1987). In contrast, PI-Pt

electrodes implanted into a rat sciatic nerve for 3 months have

been shown to induce only mild scar tissue formation and local

inflammation reactions, with these reactions being limited to a

small area around the electrode (Lago et al. 2007). One

shortcoming of PI is related to its permeability to environmental

moisture and ions that could be crucial in terms of short

circuiting and reducing the life-time of an implantable device.

There are two common types of PIs on the market: non-

photosensitive PI and photosensitive polyimide (PSPI) (Ghosh &

Mittal 1996). A special feature of PSPI, i.e. photopatternability,

simplifies the fabrication process of the microdevice compared

to conventional PI by eliminating the need for complex

multilevel processes, i.e. oxygen RIE through a lithographically

patterned metal aluminum (Al) mask (Wilson et al. 2007). PSPI

is advantageous in electrode fabrication in which electrode sites

and contact pads can be easily opened though an encapsulation

layer while still insulating the transmission lines as well as in

precision feature defining of the device’s outer shape. Although

mechanical and electrical properties of PSPI are comparable to

conventional PI, there will be a degree of uncertainty related to

its biocompatibility due to the different chemical composition

(e.g. existence of photoinitiators), imidization kinetics and

solvent processes. So far very little data have been published

addressing the biocompatibility of PSPI (Sun et al. 2009).

67

5 Biocompatibility

The main requirement for the successful clinical and

experimental application of implantable devices constructed

from artificial materials is that the surrounding tissues accept

the implant, i.e. that the implant is biocompatible.

Biocompatibility can be defined as ‚the ability of a material to

perform with an appropriate host response in a specific

situation‛ (Williams 1999). It is worth noting that the host

response to specific individual materials could vary from one

application site to another. An increasing number of

applications require that the material should actively react in a

desired manner with the tissue, or even degrade over time in the

body, rather than remain constantly as an inert material.

Biomaterials can be classified according to the chemical

composition to metals, polymers, ceramics, composites and

materials of biological origin (Navarro et al. 2008). Another

classification is based on historical stages of biomaterial

development. The first generation of implantable devices,

developed during the years between 1940 and 1980, were

constructed using materials that were chemically least reactive.

It was thought that the best performance of implants could be

achieved with a material that would be as passive as possible

regarding adsorption of host proteins and cellular adhesion.

Examples are inert metal materials that are very resistant to

corrosion such as stainless steel, cobalt-chromium alloys,

titanium alloys and the platinum group metals. With respect to

the polymers, polytetrafluoroethylene, polymethylmethacrylate,

polyethylene (PE) and silicones were widely used due to their

high resistance to degradation (Williams 2008). The second

generation biomaterials were specifically designed to be

‚bioactive‛, i.e. they could elicit a controlled host reaction and

reaction in their physiological environment to achieve the

desired result of the implantation. For example, certain

5 – Biocompatibility

68

compositions of bioactive glasses promote the desired cellular

responses, like cellular adhesion, proliferation and

differentiation into specific cell types, e.g. bone cells that will

form new bone tissue and, thus, integrate the implant strongly

into the natural peri-implant tissues. Recently, the third

generation biomaterials have been and are being actively

developed especially for in situ regeneration of damaged tissue

and different tissue engineering applications. In these materials,

living cells are used in combination with artificial materials

(Navarro et al. 2008, Hench & Polak 2002, Ratner & Bryant 2004).

There is a wide variety of chemical, biochemical,

physiological, physical and other mechanisms which are

involved in the contact between biomaterials and various tissues

and thus will affect the biocompatibility (Williams 2008,

Vadgama 2005). The concept of biocompatibility is not limited to

bulk characteristics of the material used, but also includes

physical and chemical surface properties as well as various

decomposition products. In some cases, chemical composition of

implant surface is of major importance, but in other cases some

physical features such as stiffness and shape of the implant are

the most important determinants of biocompatibility. The major

material variables that may influence on the host response are

summarized in Table 3.

Table 3: Major material characteristics that affect the host response, modified from

Williams (2008).

Material variable

The composition and structure of bulk material

Crystallinity and crystallography

Elastic constants

Porosity

Chemical composition of surface

Surface roughness, texture

Surface charge and electrical properties

Surface wettability (hydrophilicity, hydrophobicity)

Corrosion parameters, ion release profile and ion toxicity

Degradation profile and toxicity of degradation products


69

The interaction between an artificial material and a living

system can be considered to consist of the local and systemic

response of host tissue to the material as well as the response of

the material to the host tissue (Williams 2008). After

implantation of a biomaterial into the body, the host response

involves a series of complex protein and cellular events.

According to the ‚race for the surface‛ concept (Gristina 1987),

there is in many occasions a critical competition between the

microbial and host tissue cells for the implant surface. If this

race is won by tissue cells, the implant surface is covered by

tissue which provides resistance against bacterial colonization,

and promotes tissue integration of the implant. On the other

hand, if the race is won by microbes, the implant surface will be

covered by bacteria and/or bacterial biofilm. In this case, the

host cellular functions are hindered, resulting in adverse tissue

reactions, and in the worst scenario, the implant has to be

removed to protect the host.

When a biomaterial is implanted into the living tissue or

placed in cell culture, water molecules bind within nanoseconds

to its surface. The extent and pattern of the interaction between

the water molecules with the surface of implant depends on the

surface properties of the material, such as surface free energy

(SFE) (Roach et al. 2007, Kasemo 1998). Subsequently, within

seconds to hours after implantation, the implant surface

becomes covered by a monolayer of adsorbed serum or

interstitial proteins. Since there are many hundreds of different

proteins in blood and interstitial tissue fluid, there is a race for

the surface between these molecules. Typically, the more

abundant small proteins lead this race due to their rapid

transport and attachment to the surface. Later, larger proteins

with higher affinity towards the surface replace these small

proteins (Roach et al. 2007). The protein layer formation process

involves a very rapid initial phase, followed by a slower

maturation phase as the protein coating of the implant

approaches a steady-state phase. In the third stage, that occurs

from as early as minutes after implantation, lasting up to several

days, cells eventually adhere, spread out and differentiate on


70

the protein molecules coating the biomaterial surface. These

cellular processes are influenced by biological molecules such as

extracellular matrix proteins, cell membrane proteins and

cytoskeleton proteins as well as by the physicochemical

properties of the surface such as its chemistry, topography,

surface energy and electric charge (Puleo & Nanci 1999, Roach et

al. 2007). Finally, the desired integration of implant will occur

with the surrounding tissue. In biomedical implants, the final

stage which can last from a few days (rapidly biodegradable

devices) up to several decades (total hip replacement), can be

seen as the continuum of the early implant stages. During this

last stage, also adverse responses, such as clots, fibrous capsule

formation, wear, corrosion, late hematogenic infections and

device failure may occur.

5.1 BIOCOMPATIBILITY TESTING

Rigorous biocompatibility testing, both in vitro and in vivo, is

needed when novel materials that are aimed for use in

implantable medical devices, are developed. The International

Organization for Standardization (ISO) has published the

standard series ISO 10993 for biological evaluation of medical

devices. This series of standards is currently composed of 18

parts.

In vitro cytotoxicity tests (ISO 10993-5) are usually performed

as the first stage acute toxicity and cytocompatibility testing for

materials planned for use in medical devices. These tests are

rapid to perform, sensitive, well standardized and cheap.

Unnecessary use of animal models can be also avoided. In vitro

tests are convenient in studies in which the adhesion,

proliferation, differentiation and contact guidance of cells as

well as the mineralization of matrix on biomaterial surface are

investigated. These standardized and easily quantifiable tests

(Pearce et al. 2007, Pizzoferrato et al. 1994) are also useful in

studies aimed to clarify the individual effects of different surface


71

parameters such as SFE, surface charge and topography on the

cellular responses.

Various cell lines can be used for cytotoxicity testing; the cell

line should be chosen to represent the cells that will be in

contact with the biomaterial in its predicted in vivo use. Tests

can be performed either via direct contact with a material and/or

with extracts of the material. The direct contact test simulates

the near-surface effects of the implanted biomaterial on tissues,

whereas the extracts mimic the effect of the particles that are

released from the surface. An incubation time ranging from 24

hours to 96 hours is recommended.

5.2 CELL-BIOMATERIAL INTERACTIONS

The clinical success of implants is largely based on adaptive and

integrative interactions between the implant material and tissue

cells. Although appropriate mechanical properties, durability

and functionality of modern biomaterials and related devices

are primarily derived from the bulk properties of materials, the

biological responses to materials are affected by their physical

and chemical surface properties. Thus, the interactions between

cells and implants are regulated to a large extent by the surface

properties of the implant material, although properties like

elasticity can also affect host responses. The tendency of cells to

adhere or not to adhere to the surface of implant is crucial for

their performance, for example in total joint replacement

prostheses or urinary stents (Tomlins et al. 2005). In order to

improve the biological performance of implants, the effects of

modified surface properties on cellular functions have been

widely studied (Ratner & Bryant 2004, Williams 2008). Surface

chemistry, roughness, texture, wettability and surface charge are

among the characteristics that have an effect on the cellular

response to a biomaterial.

5.2.1 Effect of surface topography on cellular responses

Surface roughness is a measure of the random topography of a

surface and it is quantified by the vertical variations of a real


72

surface from its ideal form. Amplitude parameters that are used

in quantification of vertical deviations of the roughness profile

from the midline can be grouped into two subclasses: averaging

parameters and peak-to-valley parameters.

The arithmetic mean deviation of the absolute value of the

profile from a midline, Ra, is the most widely used parameter to

quantify surface roughness. Ra is determined as:

lr

a dxxzlr

R0

.1

(4)

In this equation, z(x) is the absolute profile value from a midline

and lr is the sampling length over which the surface profile has

been measured (Tomlins et al. 2005).

However, the Ra values are not always useful parameters for

assessing biomaterial surfaces, because of the lack of detailed

information about the geometry of the surface or the variations

in peak heights or valley depths. Different peak-valley

parameters have been developed to solve this problem

including statistical distribution parameters called skewness

and kurtosis. The simplest approach is to determine the highest

and lowest points within the overall measuring length, and then

the Rpv (peak-to-valley roughness) value can be estimated as the

vertical distance between the farthest points.

Both contact and non-contact techniques are available for

measuring surface roughness. Contact methods are usually

based on the use of stylus-based instruments in which a

measurement needle is traversed over a test surface and the

movement of the stylus arm can be detected e.g. using a

piezoelectric crystal. However, these devices are not suitable for

assessing the surface texture of soft materials, as they tend to

damage the surface to be analyzed. Non-contact optical

techniques such as laser profilometry and white light

interferometry are especially useful for assessing the surface

roughness of sensitive soft materials. Furthermore, different

scanning probe microscopy techniques such as atomic force

microscopy (AFM) and scanning tunneling microscopy can be


73

used for imaging surfaces with resolution at the atomic level.

(Tomlins et al. 2005, Giessibl 2003, von Recum et al. 1996)

The effect of surface roughness on cell functions has been

studied in biomaterial surfaces which can be modified by

several techniques such as sandblasting (Lampin et al. 1997,

Iwaya et al. 2008, Orsini et al. 2000, Anselme & Bigerelle 2005),

plasma treatment (Chu et al. 2002, Nebe et al. 2007, Saldana et al.

2006), mechanical polishing (Ponsonnet et al. 2003), acid etching

(Iwaya et al. 2008, Orsini et al. 2000, Anselme & Bigerelle 2005)

or laser treatment (Teixeira et al. 2007). It has been reported that

surface roughness characteristics have an effect on the

morphology, adhesion, proliferation and differentiation of

osteoblast-like cells (Puckett et al. 2008, Kim et al. 2005, Khang et

al. 2008, Das et al. 2009) as well as on fibroblasts (Könönen et al.

1992, Kunzler et al. 2007, Ponsonnet et al. 2003, Wirth et al. 2005).

It is generally recognized that fibroblasts (connective tissue cells)

prefer smooth surfaces whereas osteoblastic cells favor rougher

surfaces (Kunzler et al. 2007, Ponsonnet et al. 2003, Wirth et al.

2005). Furthermore, enhanced differentiation and bone-specific

gene expression of adhered human mesenchymal stem cells

(hMSCs) have been demonstrated on nanorough surfaces (Ra <

50 nm) compared to rougher, machined surfaces (Ra ≈ 100 - 400

nm) (Mendonca et al. 2009, 2010, Dalby et al. 2006). However,

the effects of surface roughness are also dependent on the

substrate material being used (Hallab et al. 2001, Jinno et al.

1998). Moreover, it is important to note that a change in the

surface roughness is also related to changes in the wettability

properties of the surface. This is partly due to the fact that

machining techniques commonly used to roughen or polish the

surfaces often also introduce changes in the surface charges and

tensions. Therefore, it may be difficult to conclude which

mechanism is involved in modulating the cell behaviour at these

instances (Khang et al. 2008, Das et al. 2009, Ponsonnet et al.

2003).

Surface texture refers to a surface topography in which

patterns of features are placed deliberately on the surface.

Several microfabrication techniques, mainly originating from


74

the production of integrated circuits, have been recently utilized

in the biomaterial research to study cellular reactions to a

substrate having a three-dimensional structured surface

topography in the micro- or nanoscale. The major part of

published literature has focused on micro- and nanostructures

produced through photolithography or electron beam

lithography in conjunction with reactive ion etching processes

on the surfaces of Si, quartz or Ti (Curtis & Wilkinsson 1997,

Flemming et al. 1999, Martinez et al. 2009). More recently,

polymer microfabrication techniques, such as hot embossing

(Charest et al. 2004), injection molding (Myllymaa et al. 2009b)

and soft lithography (Mata et al. 2002a,b) have been increasingly

used. Different pattern feature types such as grooves/ridges,

wells, pits and pillars with a wide variety of pattern dimensions

have been examined (Martinez et al. 2009). Substrate micro- and

nanotopography, independently of surface biochemistry, seems

to have significant effects on cell orientation and adhesion,

morphology, proliferation and differentiation. It has been

demonstrated that cells seeded onto microgrooved surface

aligned their shape and elongated in the direction of the grooves

(Brunette et al. 1983, Chou et al. 1995, 1998, Britland et al. 1996,

Curtis & Wilkinson 1997, Teixeira et al. 2003, Dalby et al. 2004,

Ber et al. 2005, Mwenifumbo et al. 2007, Loesberg et al. 2007, den

Braber et al. 1996, Clark et al. 1987, Charest et al. 2004, Hamilton

& Brunette 2005; Hamilton et al. 2006). The cell membrane

conformed to the widest grooves, but tended to bridge the

narrowest grooves (Matsuzaka et al. 2003). Other geometrical

features such as wells, pits and pillars with dimensions of less

than 5 µm seem to lead to smaller, rounded cells with less

organized cytoskeletons (Martinez et al. 2009, Andersson et al.

2003, Gallagher et al. 2002). The results from investigations

related to the effects of nanoscale topographical structures on

cell proliferation and differentiation are somewhat contradictory

and highly dependent on the cell line and structures being

investigated (Puckett et al. 2008, Andersson et al. 2003, Martinez

et al. 2009).


75

5.2.2 Effect of wettability properties on cell-biomaterial

interactions

Surface energy of materials is of special interest in the

biomedical applications due to its obvious effect on cell-

biomaterial interactions. Surface wetting properties have a

significant effect on various biological events at the sub-cellular

and cellular level including protein adsorption, cell attachment

and spreading. Contact angle measurements are widely used in

the assessment of wettability properties. In this technique, a

small droplet (e.g. 15 µl) is placed on the surface and the angle

between the drop (liquid-vapour interface) and the substrate

(solid) is determined. Relatively more water wettable surfaces

are called hydrophilic (high SFE), while less wettable surfaces

are hydrophobic (low SFE) surfaces (Ratner 1996, Hasirci &

Hasirci 2005). Although these are relative terms, it has been

suggested that water contact angles smaller than 65° indicate

hydrophilic surface, while those larger than 65° indicate

hydrophobic surface (Vogler 1998).

Different approaches, such as the Owens-Wendt (Owens &

Wendt 1969) and van Oss (van Oss et al. 1988) models have been

developed to estimate SFE of materials and their components on

the basis of measured contact angles with polar and non-polar

liquids whose surface tension values are known. Using the

Owens-Wendt model, the polar PS and dispersive D

S

components can be estimated:

).)()((2)cos1( 2/12/1 P

L

P

S

D

L

D

SL

(5)

Here θ is the measured contact angle value and L , DL and P

L stand for the liquid’s total SFE and its dispersive and

polar components, respectively. The total SFE ( S ) is the sum of

its dispersive and polar components:

.P

S

D

SS

(6)


76

The polar component of the SFE is related to hydrogen

bonding, dipole-dipole and other site-specific interactions,

whereas the dispersive component refers to the van der Waals

and other non-site specific interactions between the solid surface

and liquid applied on it.

Increased wettability with a higher SFE has been shown

generally to enhance the adhesion, and spreading of various cell

types such as fibroblasts (Wei et al. 2007, Altankov et al. 1996,

van Kooten et al. 1992), osteoblasts (Khang et al. 2008, Feng et al.

2003, Lim et al. 2008, Myllymaa et al. 2009b) and MSCs (Sawase

et al. 2008). Moreover, through the SFE modifications, it is

possible to influence the differentiation of MSCs into the desired

phenotypes (Curran et al. 2005, 2006, Marletta et al. 2007,

Calzado-Martin et al. 2010). It has been presented that the SFE

is an even more important surface characteristic than surface

roughness for cellular adhesion strength and proliferation and

might be a useful parameter for modifying cell adhesion and cell

colonization onto engineered tissue scaffolds (Hallab et al. 2001).

The wettability of a solid material can be modified either by

changing the surface roughness (Wang et al. 2006, Hao &

Lawrence 2007) or the surface chemistry (Lim et al. 2008,

Michiardi et al. 2007, Kennedy et al. 2006, Myllymaa et al. 2009b).

In hydrophobic materials, surface roughening increases the

hydrophobic nature of material, whereas in hydrophilic

materials surface roughening enhances hydrophilicity. Surface

chemistry modifications are usually performed by using

different plasma treatments (Lim et al. 2008, Michiardi et al.

2007), proper surface coatings (Khang et al. 2008, Myllymaa et al.

2009b), SAMs (Kennedy et al. 2006, Folch & Toner 2000) or

incorporating dopants into the bulk material (Myllymaa 2009a).

5.2.3 Effect of charge distribution on cell-biomaterial

interactions

The electric charge distribution in the vicinity of a biomaterial

surface can be considered to be one of the main factors affecting

the biological response of a biomaterial (Khorasani et al. 2006,

Cai et al. 2006, Krajewski et al. 1998, Altankov et al. 2003). This


77

charge depends on several factors, such as the chemical

composition of the material surface, the composition of the

surrounding fluid and the environmental pH value (Krajewski

et al. 1998). The charge and its distribution are closely related to

protein adsorption, the strength of cellular adhesion and the

shape of the cells growing on the surface of the biomaterial

(Cheng et al. 2005, Bodhak et al. 2009, Thian et al. 2010).

Therefore, the investigations of the relationship between

electrical charge on a biomaterial surface and protein adsorption

as well as subsequent cellular interactions are particularly useful

ways to improve understanding of the mechanisms involved in

the biological integration of materials with tissues.

The interface between a solid surface and the surrounding

electrolyte establishes a charge distribution which is different

from the bulk phases of solid and liquid. In this electrochemical

double layer model, at the vicinity of a solid-liquid interface, the

charge carriers are fixed (a stationary layer), whereas they are

mobile in the liquid phase (a mobile layer) at a greater distance

The zeta potential (ζ) is termed as the potential decay between

the solid surface and the bulk liquid phase at the shear plane

(Lyklema 1995) (Fig. 8). Streaming current/streaming potential

measurement is a convenient method for examining the existing

interface charges of a solid biomaterial in an electrolyte solution

(Roessler et al. 2002). In this method, an electrolyte is circulated

through the measuring cell and forced under pressure to flow

through a small gap formed by two sample surfaces. This fluid

flow evokes a relative motion between the stationary and mobile

layer and this leads to a charge separation which provides

experimental access to the zeta potential.

Higher positive values of the zeta potential at a fixed pH

reflect a positive charge of the surface which would attract

negatively charged entities such as anions or charged proteins

while lower values of zeta potential at a fixed pH are evidence

of a negative charge of the surface which tends to attract

positively charged cations and particles (Cai et al. 2006). The

relationship between the magnitude of zeta potential and

cellular responses is not well understood (Altankov et al. 2003,


78

Hamdan et al. 2006), although in recent studies it has been

demonstrated that negatively charged surfaces can enhance

osteoblastic cell – biomaterial interactions (Bodhak et al. 2010,

Thian et al. 2010). Instead, the effect of zeta potential on protein

adsorption has been more extensively studied. The lower the

absolute value of the zeta potential of the surface, the lower is

the adsorption of serum proteins (Krajewski et al. 1996, 1998, El-

Ghannam et al. 2001).

stationary layer mobile layer

-

-

-

-

-

-

-

-

+

+

+

+

+

+

+

+

-

-

-

+

Distance

Ele

ctro

kine

tic p

oten

tial

ζ

+ -

+

solid

Figure 8: The interface between a solid surface and a surrounding solution introduces

an electrochemical double layer. The charge distribution is divided into a stationary

and a mobile layer. These layers are separated by a plane of shear. The zeta potential (ζ )

is assigned to the potential decay at this shear plane.

5.2.4 Effect of surface micropatterning on cellular responses

Surface micropatterning has been extensively employed for

controlling the adhesion, proliferation, differentiation and

contact guidance of cells (Folch & Toner 2000, Falconnet et al.

2006, Ito 1999). Photolithography and soft lithography are

powerful techniques to produce surface patterns which differ

from the background e.g. in terms of surface chemistry (Kane et

al. 1999, Winkelmann et al. 2003, Scotchford et al. 2003, Levon et

al. 2009, Myllymaa et al. 2009a), surface charge (Soekarno et al.

1993), SFE (Matsuda & Sugawara 1995), or the attachment of

biological substances that can influence cellular function

(McBeath et al. 2004, Chen et al. 1997, Blawas & Reichert 1998).

79

6 Aims of the present study

In this study, micro- and nanofabrication methods have been

utilized in the development of novel neural interfaces as well as

intelligent implant materials capable of achieving the desired

cellular responses on a biomaterial surface. It was hypothesized

that these approaches would enhance the tissue response and

functionality of the implants used in contact with soft (neural)

or hard (bone) tissue. The main aims of this thesis were:

1. to design and fabricate novel thin film microelectrode arrays

for stable and high-resolution intracranial recordings,

2. to study the suitability of flexible polymers (e.g. polyimide)

as construction materials in bio-MEMS applications to

replace materials based on silicon technology,

3. to study the opportunities to improve the electrochemical

properties (stability, contact impedance) of bioelectrodes by

using modern micro- and nanostructuring methods,

4. to investigate the cytocompatibility of different dielectric

thin films produced by USPLD method and to clarify their

suitability for neural interfaces,

5. to test whether the surface micropatterning could be used to

enhance the cytocompatibility of cell unfriendly silicon-

based implants and improve their biocompatibility during

the initial phase of integration with tissue, and

6. to investigate the effect of surface micropatterning,

produced using photolithographic and PVD techniques, on

the behaviour (e.g. adhesion, spreading, and differentiation)

of osteoblast-like cells and human mesenchymal stem cells.

6 – Aims of the present study

80

81

7 Materials and methods

The present thesis consists of six studies, which are referred to

by the Roman numerals (I-VI). In addition, some unpublished

data are included. A summary of the materials and methods

used in this thesis is presented in Table 4.

Table 4: Overview of the materials and methods used in the present thesis.

Study Device/ sample Characterization Cell line/

Animal

Testing methods

I Microelectrode array

Optical microscopy, EIS

Wistar rat Acute cortical evoked potential recordings

II Microelectrode array

Optical microscopy, SEM,

AFM, EIS

Wistar rat Acute and chronic cortical evoked potential

recordings

III PI films on Si,

Al203 on Si (*)

C3N4 on Si (*)

AFM, Contact angle, SFE, Zeta

potential

BHK-21

Qualitative SEM analysis of cultured cells, MTS

assay, Fluorescence

microscopy (live/dead)

IV Review article

V Micropatterned Ti, Ta, Cr and DLC

coatings on Si

Optical microscopy, AFM,

Contact angle,

SFE, Zeta

potential (*)

hMSC Quantitative and qualitative SEM analysis

of cultured cells,

Fluorescence

microscopy of stained cells (focal adhesion,

ALP, OC, MS)

VI Micropatterned Ti and DLC coatings

on Si

Micropatterned Cr and Ta coatings

on Si (*)

(#) SaOS-2 Quantitative and qualitative SEM analysis

of cultured cells,

Confocal laser scanning

microscopy of stained

cells (focal adhesion)

Unpubl. data

nanorough USPLD Pt surfaces

EIS, AFM

(*) Additional materials/methods: not included in original publications III, V or VI;

(#) Sample surfaces were characterized in study V

Abbreviations: EIS: electrical impedance spectroscopy; SEM: scanning electron

microscopy; AFM: atomic force microscopy; PI: polyimide; Si: silicon; Al203: alumina;

C3N4: carbon nitride; SFE: surface free energy; BHK-21: baby hamster kidney fibroblast;

Ti: titanium; Ta: tantalum; Cr: chromium; DLC: diamond-like carbon; hMSC: human

mesenchymal stem cell; ALP: alkaline phosphatase; OC: osteocalcin; MS: mineral

staining; SaOS-2: sarcoma osteogenic cell; USPLD: ultra-short pulsed laser deposition;

Pt: platinum

7 – Materials and methods

82

7.1 MICRO- AND NANOFABRICATION OF ELECTRODES AND

OTHER SAMPLES

A wide variety of micro- and nanofabrication techniques as well

as materials were used in this thesis work. A description of

these fabrication processes is given in chapters of 7.1.1 – 7.1.5.

7.1.1 Preparation of photomasks

In the studies I, II, V and VI, photolithographic methods were

used in patterning thin deposition layers on PI or Si substrates.

The photomasks used in these studies were designed by CleWin

layout editor, version 4.0.2 (WieWeb software, Hengelo, The

Netherlands) and fabricated using the laser scanning technique

by Mikcell Oy (Ii, Finland) on 4-inches or 5-inches soda-lime

glass plates with a structured Cr layer. The photomask used in

studies I and II consists of two mask patterns both repeated five

times enabling the ‚batch fabrication‛ at a prototyping scale.

The first set of patterns was used in structuring the metallization

layer and the second set in patterning the PI insulating layer to

expose electrode sites. The photomasks used in studies V and VI

consist of micropatterned areas ( 4 mm x 4 mm or 8 mm x 8 mm)

containing regularly spaced microsquares (sides 5, 25, 75 or 125

µm) or microcircles (diameters 5, 25 or 125 µm).

7.1.2 Flexible polyimide-based microelectrode arrays

A fabrication process for flexible MEAs that follows photo-

lithographic and thin film procedures was developed and

iterated to a fully functional solution during a two year trial

period such that appropriate sensor layouts and materials with

their thicknesses for base, metallization and insulating layers

were optimized. The most relevant variations are presented in

Table 5. The developed arrays consisted of either 8 or 16 round-

shaped microelectrodes with diameters of 200 µm or 100 µm,

respectively, in an area of about 2 mm x 2 mm at the end of a PI

ribbon. The layouts of these electrodes are presented in Fig. 9.


83

Table 5: Developed microelectrode array prototypes.

Type Channels Substrate material

and thickness

Electrode material,

its diam.

Insulation layer, its

thickness

Reference

A 8 Polyimide PI-2525a: 3 layers, total 30 µm

Ti/Pt, 200 µm

PI-2771a, 3µm

Study I

B 8 Kapton HN filmb 25 µm + PI-2525 5 µm

Ti/Pt, 200 µm

PI-2771, 3µm

Study II

C 8 Kapton HN film 25 µm + PI-2525 5 µm

Ti/Au, 200 µm

PI-2771, 3µm

Myllymaa et al. 2008

D 16 Kapton HN film 25 µm + PI-2525 5 µm

Ti/Pt, 100 µm

PI-2771, 3µm

Myllymaa et al. 2008

E 16 Kapton HN film 25 µm

+ PI-2525 5 µm

Ti/Pt,

100 µm

SU-8c,

3 µm

Myllymaa

et al. 2008

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

A

B C

32 m

m

1 mm

Electrode sites

Connector

pads

Figure 9: (a) The layout of the 16-channel microelectrode array. The length of arrays is

designed to be about 32 mm. There are connector pads at one end of the device allowing

connection of the array to miniature connector and further to recording

instrumentation. The magnified schema of the 16-channel (b) and 8-channel (c) array

viewed at the recording ends. The sizes of circular-shaped microelectrodes are 100 µm

(b) and 200 µm (c).


84

Similar processes were used in studies I and II except for the

construction of a base layer and the curing parameters for

insulating PI layer. In study I, a base layer was constructed from

three spin-coated PI (PI-2525, HD Microsystems) layers, i.e. the

thickness of the base was about 30 µm. In study II, a 25 µm thick

Kapton HN film (Goodfellow Ltd., Cambridge, UK) was utilized

together with a spin-coated single PI-2525 layer, resulting in a

base with total thickness of 30 µm. The process flow is

demonstrated in Fig. 10.

polyimide, d ~ 30 µm photoresist

Ti/Pt thin film, d ~ 200 nm

UVUV

glass slide

A

B

C

D

connector padelectrode opening

E

F

G

photosensitive polyimide, d ~ 3 µm

Figure 10: The fabrication process flow of polyimide-based microelectrode arrays. The

process is based on photolithographic and magnetron sputter deposition techniques.

Electrode sites, transmission lines and connector pads are sandwiched between two

layers of polyimide. See details in the text.

First, one of the above mentioned practices was used to form

a 30 µm thick PI layer onto 2‛ x 3‛ microscope glass slides

(Logitech Ltd., Glasgow, Scotland, UK) which were used to

ensure a rigid support throughout the process (Fig. 10a). The PI

layer was cured in an oven according to the manufacturer’s

recommendations. After cooling, 20 % hexamethyldisilazane


85

(HMDS) (Riedel-de-Haen AG, Seelze, Germany) in xylene was

spun upon the PI base layer to enhance adhesion between the PI

and the photoresist. A negative photoresist (ma-N 1420, Micro

Resist Technology GmbH, Berlin, Germany) was spun (5000

rpm, 30 s), pre-baked on a hot plate (100°C, 120 s), exposed with

i-line (365 nm) UV light (Karl Suss MA45, Suss Microtec Inc.,

Waterbury Center, VT, USA) for 30 s, developed in a ma-D 553s

developer (Micro Resist Technology), rinsed with deionized

water and post-baked in an oven (100°C, 30 min) (Fig. 10b). Thin

films of Ti and Pt were deposited at thicknesses of 20 nm and

200 nm, respectively, by using magnetron sputtering (Stiletto

Serie ST20, AJA International Inc., North Scituate, MA, USA)

and high purity (99.9 % or better) Ti and Pt targets (Goodfellow

Metals Ltd., Huntingdon, UK) (Fig. 10c). Ti serves as an

adhesion promoter between the PI and the Pt. The metallization

layer was patterned via a lift-off procedure in resist remover,

mr-Rem 660 (Micro Resist Technology) forming electrodes,

transmission lines and connector pads on the PI base (Fig. 10d).

Photosensitive polyimide PI-2771 (HD Microsystems) was used

as an insulation layer. PI-2771 was spun (5000 rpm, 60 s) and

pre-baked on the hot plate (Fig. 10e). After cooling, it was

exposed to UV light (12 s) and developed in cyclopentanone

(Sigma-Aldrich Corp., St. Louis, MO, USA) and rinsed with

propylene glycol methyl ether acetate (PGMEA, Sigma-Aldrich).

The PI-2771 layer was cured in an oven in which the

temperature was gradually raised either to 200 °C after which it

was maintained for 30 minutes (study I) or to 300 °C after which

it was maintained for 60 minutes (study II). PI-2771 formed

about a 3 µm thick insulation layer that was present on all

positions other than the areas of the electrode sites and

connector pads (Fig. 10f). Finally, the arrays were detached

from the glass slide (Fig. 10g) and cut into their final shapes

using scissors and a knife. Since only the edge regions were

tightly attached to glass slides (detailed descriptions of these

protocols are given in original papers I and II), it was possible to

achieve detachment easily with by immersing in deionized

water.


86

Thin film connector pads were designed to fit into a 16-

channel 0.5 mm pitch zero-insertion-force (ZIF) connector (JST

Ltd., Halesworth, UK). This ZIF connector was soldered with

lead-free solder to a thin printed circuit board (PCB) adapter

(Kytkentälevy Oy, Helsinki, Finland) containing also a 2 x 8

channel surface mount microsocket (CLM-serie, Samtec Inc.,

New Albany, IN, USA) through which the array was connected

to the recording instrumentation. Casco strong epoxy-resin

(Akzo Nobel Coating Oy, Vantaa, Finland) was used as an

encapsulant at the interface between the array and the ZIF

connector and the signal tracks in the PCB in order to avoid

short-circuiting when implanted onto the surface of rat cortex.

7.1.3 Micropatterned biomaterial surfaces on silicon

In studies V and VI, micropatterned titanium (Ti), tantalum (Ta),

chromium (Cr) and diamond-like carbon (DLC) coatings were

fabricated using photolithography and PVD methods. Si wafers

were used as substrates. Prior to application of the photoresist,

the Si wafers were dry-baked and coated with adhesion

promoter (20 % HDMS in xylene). Then, the chosen photoresist,

i.e. ma-N 1420 (Micro Resist Technology) in study V or SU-8

(MicroChem) in study VI was applied, patterned and cured

using the optimized process parameters described in the

original papers (V, VI). Two PVD methods were used to carry

out thin film depositions. Magnetron sputtering was used to

deposit Ta, Ti and Cr coatings and a filtered pulsed plasma arc

discharge method was used to produce DLC coatings onto the

surface of the patterned wafers. The process parameters were

optimized (details are described in the original articles V and VI)

to achieve well-adhering, smooth coatings with a thickness of

about 200 nm. The micropatterns were revealed via a lift-off

procedure by immersing the wafers in a resist remover (acetone

or mr-Rem 660) ultrasonic bath. The biomaterial coating

deposited on the top of the photoresist was removed together

with the dissolved resist, and the final microfeatures were

formed. The Si wafers were cut into individual samples (10 mm

x 10 mm) with a custom-made device using a diamond knife.


87

The size of each final micropatterned sample used in the

cellular studies was 10 mm x 10 mm. In study VI, three different

sample sets were utilized. In the first sample set, each chip

contained four 4 mm x 4 mm sample areas. Three of these areas

consisted of circles with diameters of 5, 25 or 125 µm or squares

with the length of the sides being 5, 25 or 125 µm composed of

either DLC or Ti on a Si background with the fourth sample area

containing Ti or DLC as a homogeneous surface (Fig. 11a,b).

Regardless of the size of the patterns, the circles and squares

covered 30.6 % and 25 % of the total sample surface area,

respectively. In the second set of samples, the pattern and

background were reserved (‚inverse samples‛) so that the

patterns were composed of uncoated Si squares or circles on a

DLC or Ti background. In the third set, each sample contained a

micropatterned area of 8 mm x 8 mm at the center of the sample.

Micropatterning in these areas consisted of 75 µm squares with

100 µm spacing between adjacent squares in two orthogonal

directions (Fig. 11c). Thus, the (Ti, Ta, Cr or DLC) patterns

occupied 18.4 % of the total sample surface area. This last

sample type was also used in study V.

Figure 11: Sample types used in studies V-VI to clarify cellular behaviour on

micropatterned surfaces: (a) circular patterning, diameters of circles correspond to 125

µm, 25 µm and 5 µm plus plain reference area (bottom right), (b) rectangular

patterning, sides of squares correspond to 125 µm, 25 µm and 5 µm plus plain

reference area (bottom right) and (c) rectangular patterning, sides of squares are 75

µm.


88

7.1.4 Tailored surfaces produced by ultra-short pulsed laser

deposition

Ultra-short pulsed laser deposition (USPLD) technique was

utilized in the creation of novel coatings for applications of bio-

MEMS. Firstly, ultrasmooth alumina (Al2O3) and carbon nitride

(C3N4) dielectric films were deposited on Si wafers to clarify

their cytocompatibility properties using similar fibroblast

experiments as in study III (unpublished findings). Secondly, Pt

depositions with varied nanotextured surface topography were

carried out on SS, Si and Kapton substrates aiming to find out

new solutions to improve the electrochemical properties of

bioelectrodes.

Neural electrodes need to be encapsulated in a material that

provides excellent barrier properties and shows acceptable

biocompatibility. In this context, Al2O3 and C3N4 were

considered as interesting coating candidates. Biomaterials

coated with Al2O3 have been used for a long time in medical

applications, particularly in fields of dentistry and orthopedics

(Szeitzer et al. 2006). The biocompatibility of Al2O3 has been

previously demonstrated as a bulk material (Heimke et al. 1998,

Takami et al. 1997) as well as an atomic layer deposited film

(Finch et al. 2008). The dielectric properties of pulsed laser

deposited Al2O3 are also superior (Katiyar et al. 2005) and

USPLD deposited Al2O3 provides good insulating properties

even in liquid environment (R. Lappalainen, personal

communication 2009). Carbon nitride is a promising novel

coating material for biomedical applications. It has superb

chemical inertness and its tribological and dielectric properties

are also favourable (Cui & Li 2000). The abbreviation ‚CN‛ is

commonly used to describe all carbon nitride coatings with

different C/N ratios. The bio- and hemocompatible nature of CN

has been also demonstrated (Du et al. 1998, Cui & Li 2000, Cui et

al. 2005). However, in neural applications, the testing of Al2O3

and CN films has been rare.

The USPLD method utilizes intense laser pulse (typically in

ps time scale) to erode a target and deposit the eroded material

in the form of high speed ionized plasma onto a substrate. Just


89

prior to the depositions, the sample surfaces were gently

cleaned using Ar+ ion sputtering (SAM-7KV, Minsk, Belarus). A

new type of mode-locked fiber laser (Corelase Oy, Tampere,

Finland) and the ColdabTM deposition technology developed by

Picodeon Ltd. Oy (Helsinki, Finland) were used for USPLD. A

high purity sintered target of C3N4 and Al2O3 was used for

deposition. In both cases, optimized laser parameters (Amberla

et al. 2006), such as pulse energy, pulse repetition rate and

scanning length were used. Maximum average power was 20 W

at 4 MHz, which results in 5 µJ pulse energy. The pulse length

was 15-20 ps. The laser beam was focused into a 15-25 micron

spot giving the maximum fluences about 4 J/cm2 at 5 µJ. Both

Al2O3 and C3N4 films were deposited at thicknesses of about 200

nm on p-type Si wafers.

The USPLD technique was also used to produce nanorough

Pt depositions on SS, Si and Kapton substrates. Laser deposition

parameters were modified to obtain thin (100-300 nm) Pt films

with a different nanotextured surface topography (unpublished

data). Due to the high energy of Pt plasma, good adhesion was

achieved without using any special intermediate layers or

special techniques. The main deposition parameter to control

the surface topography was the Ar gas atmosphere with a

pressure varied in the range 1-10-6 mbar.

7.1.5 Spin-coated polyimide films for cytotoxicity testing

In order to evaluate the cytocompatibility of different PI grades,

non-photosensitive PI-2525 (HD Microsystems) and photo-

sensitive PI-2771 (HD Microsystems) were deposited on Si

wafers by using a spin-coating technique. Prior to application of

the PI precursor, 2-inches, p-type <100> Si wafers (Si-Mat,

Landsberg am Lech, Germany) were baked for 20 min at 200 °C

and coated with VM-651 (HD Microsystems) to remove

moisture and to enhance the adhesion, respectively. Then, the

PI-2525 and PI-2771 precursors were spin-coated (5000 rpm, 60 s)

under low acceleration onto Si wafers. Subsequently, the PI-2771

deposits were pre-baked (100 °C, 90 s), exposed to 365nm UV

light (Karl Suss MA-45), developed in cyclopentanone (Sigma–


90

Aldrich) and rinsed with PGMEA (Sigma–Aldrich). Half of the

PI-2771 coated wafers and all of the PI-2525 coated wafers were

polymerized by curing them in an oven in which the

temperature was gradually raised to 350 °C over a period of 60

min and then maintained for 60 min at that temperature.

Another half of the PI-2771 coated wafers was cured at a lower

temperature (200 °C). Therefore, three different PI groups were

investigated: (1) non-photosensitive PI-2525, cured at 350°C; (2)

photosensitive PI-2771, cured at 200 °C and (3) photosensitive

PI-2771, cured at 350 °C.

7.2 MICROSCOPIC CHARACTERIZATION OF SURFACES

Different microscopic methods, including optical microscopy,

scanning electron microscopy (SEM), atomic force microscopy

(AFM) and contact angle measurements were used in this thesis

in quality control monitoring and in the characterization of the

produced samples and devices. Brief descriptions of these

characterization techniques are given in the following sections

of 7.2.1– 7.2.3.

7.2.1 Scanning electron microscopy

Scanning electron microscopy was used for examining the

quality of photolithographically patterned surfaces (studies V

and VI), spin-coated PI films (study III) and the MEAs (study II).

Prior to imaging with a Philips XL30 ESEM-TMP (Fei Company,

Eindhoven, The Netherlands), the samples were usually coated

with a thin layer of Au using a Sputter Coater E 5100 (Polaron

Equipment Ltd., Hertfordshire, UK).

7.2.2 Atomic force microscopy

Atomic force microscopy was used in this thesis (studies II, III,

and V) to analyze the surface topography (roughness) of

biomaterial samples and to determine the thickness of thin film

depositions. The characterization was performed using a PSIA

XE-100 (Park Systems Corp., Suwon, Korea) AFM system at


91

ambient temperature and humidity. Al-coated Si cantilevers

(Acta-10, ST Instruments B.V., LE Groot-Ammers, The

Netherlands) were used in a non-contact mode to probe the

surface across an area of 2 x 2 µm. The scanning rate of 0.25 Hz

was used. The thicknesses of thin film depositions as well as the

average surface roughness (Ra) and peak-to-valley roughness

(Rpv) values of sample surfaces were determined using the

analysis software (XIA) of the AFM instrument.

7.2.3 Contact angle measurements and determination of

surface free energies

The contact angle measurements were performed to clarify

wettability properties of material surfaces (studies III and V).

Prior to the measurements, the samples were ultrasonicated in

ethanol and deionized water and allowed to dry. Contact angles

were measured using the sessile drop method at constant

laboratory conditions (22 °C and in 45% relative humidity) with

the aid of a custom made apparatus that consisted of a stereo

microscope SZ-PT (Olympus Corp., Tokyo, Japan) equipped

with a digital camera (Camedia C-3030-ZOOM, Olympus). In

this method, a small droplet (15 µl) was pipetted onto the

surface under study, and after 5 seconds, the droplet was

photographed. Gnu Image Manipulation program (GIMP,

www.gimp.org) was used to determine the contact angle from

the shape geometry of the droplet. One polar liquid, i.e. water

and one non-polar liquid, i.e. diiodomethane was used to access

the total SFE as well as its polar and dispersive components. The

Owens-Wendt model, described in chapter 5.2.2, was used to

determine SFEs.

http://www.gimp.org/


92

7.3 ELECTROCHEMICAL CHARACTERIZATION OF SURFACES

The electrochemical properties of electrodes and other

biomaterial surfaces were investigated by electrochemical

impedance spectroscopy (EIS) and zeta potential measurements.

Descriptions of methods used are given in the following sections.

7.3.1 Electrochemical impedance spectroscopy

Electrochemical impedance spectroscopy was used (studies I

and II) to characterize the electrochemical properties of the

fabricated microelectrodes, i.e. to assess the recording

capabilities of the electrode for making neural measurements.

The measurements were performed with an LCR meter (3531 Z

HiTester, HIOKI E.E. Corp., Nagano, Japan) (study I) or with a

Solartron 1260 impedance gain/phase analyzer (Solartron

Analytical, Farnborough, UK) (study II). In studies I and II, the

MEA was immersed in physiological saline solution (0.9% NaCl)

at room temperature and a small sinusoidal perturbation signal

(100 mV without any DC offset) was applied between the

individual microelectrode and the Pt counter electrode having a

much larger surface area (Fig. 12a). The induced current and its

phase were recorded. Frequency bands of 50 Hz - 1 MHz and 1

Hz-100 kHz were used in the studies I and II, respectively.

Electrode-electrolyte interface impedance spectra of SS

electrodes with and without a nanotextured Pt layer were also

measured (unpublished data). The EIS measurements were

implemented in a custom-made miniature measurement cell

(Fig. 12b), in which a working electrode and a counter (Pt)

electrode were clamped against the opposing walls of the cell.

The holes with a diameter of 8 mm at the both walls exposed

both the working and the counter electrodes with equal areas of

50 mm2. The measurements were carried out at frequencies

between 1 Hz and 1 MHz by applying a sinusoidal voltage of

100 mV without any DC offset using the Solartron 1260

impedance analyzer.


93

Pt counter electrode

microelectrode

Pt counter

electrode

Sample under test

saline

solution

A

B

Figure 12: Set-ups used to measure the electrode-electrolyte interface impedances.

Electrical impedance spectroscopy was carried out with a two-electrode configuration,

in which a platinum sheet was used as a counter electrode. Working electrode, i.e.

microelectrode (a) or nanotextured platinum coating on stainless steel (b) was

immersed in saline (0.9 % NaCl) solution or exposed to contact with it through the

hole (Ø: 8 mm) drilled at the wall of polytetrafluoroethylene-based pool, respectively.


94

7.3.2 Zeta potential measurements

In study III, the streaming current measurement technique was

used to examine existing interface charges of the studied solid

biomaterials in a liquid. The streaming current measurements

were performed with an electrokinetic analyzer (SurPASS,

Anton Paar GmbH, Graz, Austria) equipped with an adjustable

gap cell. In each measurement, two identical test samples, i.e.

pieces of Si wafer coated with biocompatible thin film were

fixed on the sample holders with a cross section of 20 mm x 10

mm by using a double-sided adhesive tape. The sample holders

were then inserted in the adjustable gap cell and a gap of

approximately 100 µm was adjusted between the surfaces of the

samples during the measurement. An electrolyte was circulated

through the adjustable gap cell and forced under pressure to

flow through a narrow gap. This evoked a comparative

movement of charges in the electrochemical double layer which

could be detected by two Ag-AgCl electrodes placed at both

ends of the gap channels. The measurements were performed

using 1 mM KCl as an electrolyte solution at a fixed pH of 7.4 ±

0.2. The pH of KCl was adjusted with 0.1 M HCl and 0.1 M

NaOH. An electrolyte was circulated through the adjustable gap

cell and forced under pressure to flow through a narrow gap.

This evoked a comparative movement of charges in the

electrochemical double layer which could be detected by two

Ag-AgCl electrodes placed at both ends of the gap channels.

The measurements were performed using 1 mM KCl as an

electrolyte solution at a fixed pH of 7.4 ± 0.2. The pH of KCl was

adjusted with 0.1 M HCl and 0.1 M NaOH. The zeta potential (ζ)

was obtained from streaming current measurements according

to the Helmholtz-Smoluchowski equation (Lyklema 1995):

.0 A

L

dp

dI

(7)

Here dI/dp is the slope of the streaming current versus the

differential pressure, is the electrolyte viscosity, 0 is the

vacuum permittivity, is the dielectric constant of the


95

electrolyte and L is the length of the streaming channel, and A is

the cross-section of the streaming channel (the rectangular gap

between the planar samples).

7.4 RECORDING OF EVOKED POTENTIALS IN RATS

The suitability of the MEAs for cortical mapping of evoked

potentials was investigated in acute and chronic recordings in

Wistar rats in the A.I.V. Institute of the University of Eastern

Finland (studies I and II). All experiments were conducted in

accordance with the Council of Europe guidelines and approved

by the Institutional Animal Care and Use committee and the

State Provincial Office of Eastern Finland.

Prior to surgery, Wistar rats were anesthetized with urethane

(1.2-1.5 g/kg) for acute recordings and with medetomidine/

ketamine (0.5 mg/kg and 75 mg/kg, respectively) for implanting

the chronic electrodes, and placed in a stereotaxic apparatus. A

microelectrode array was placed on the dura over the pre- and

postparietal cortices between lambda and bregma (Fig. 13).

Figure 13: Depiction of the implanted electrode location in Wistar rat skull.

Microelectrode array was inserted on the dura over the pre- and postparietal cortices

between lambda and bregma. A stainless steel screw used as a reference/ground

electrode (R) was placed 2-3 mm in front of bregma.

A stainless steel screw inserted 2-3 mm in front of bregma

was used as a reference electrode connected to a ground

potential. In the chronic recordings, electrodes were anchored

onto the rat’s skull with screws and bone cement. After surgery


96

under medetomidine/ketamine anesthesia, the rat was aroused

by giving the antagonist, atimepazole, and rats were allowed to

recover for a minimum of 7 days before the first experiments. In

the chronic trials, the implanted rats were either allowed to

move freely or were gently immobilized by holding it securely

in a towel.

Both auditory and electrical current stimuli were used to

generate EPs. The auditory stimuli consisted of 25 sets of paired

pulses (pulse duration: 10 ms) with pulse intervals of 500 ms,

followed by a 9.5 s quiet interval. The square wave current

stimuli consisted of 25 sets of paired pulses (pulse duration: 1

ms) with pulse intervals of 500 ms, followed by a 9.5 s quiet

interval, delivered to the left front paw of the rat. Auditory and

somatosensory evoked field potentials were amplified with a

preamplifier (Neuralynx Inc., Bozeman, MT, USA) and with a

main amplifier (Grass Instruments, West Warwick, RI, USA)

and led to the data acquisition PC running SciWorks

(DataWave Technologies, Loveland, CO, USA). In study II, in

vivo electrode impedance measurements were performed using

a multi-task tester (Fredrich Hare Inc., Brunswick, ME, USA) in

order to clarify the long-term performance of the MEAs.

7.5 CELL CULTURE STUDIES

Studies III, V and VI concentrated on investigating cellular

responses on the surface of the biomaterial samples. The

following chapters (7.5.1 – 7.5.4) describe the methodology used

in these studies.

7.5.1 Cell lines

In study V, hMSCs (PoieticsTM, Lonza Group Ltd., Basel,

Switzerland) were cultured in Lonza Mesenchymal Stem Cell

Growth Medium (MSCGM) at 37 °C in 5% CO2 in air. The cell

monolayer was washed twice with 140 mM phosphate buffered

saline (PBS, pH 7.4) and the cells were removed by applying

0.25% trypsin in PBS-EDTA (ethylenediamine-tetraacetic acid)


97

solution for 5 min at room temperature. The suspension was

then transferred to a Falcon tube and trypsin was removed by

centrifugation. The cells were resuspended in culture medium,

seeded onto the biomaterial surfaces at a density of 0.52 x 104

cells/cm2 and cultured for 7.5 h or 5 days. In the osteogenic

induction experiments, the hMSC cells were seeded at a density

of 0.31 x 104 cells/cm2. Half of the samples were induced 24 h

after seeding by replacing MSCGM with Osteogenesis Induction

Medium (Lonza). Induced hMSCs were supplied every 3–4 days

with fresh medium. Noninduced hMSCs (control samples) were

fed with MSCGM using the same schedule.

In study III, baby hamster kidney fibroblasts (BHK 21, clone

13, HPA Culture Collections, Salisbury, UK) were cultured in

Dulbecco’s Modified Eagle’s Medium (DMEM, EuroClone

S.p.A., Pero, Italy) supplemented with 10% fetal calf serum (FCS,

PAA Laboratories GmbH, Linz, Austria), 50 µg/ml ascorbic acid

(Sigma-Aldrich), 2mM l-glutamine (PAA Laboratories), 20

IU/ml penicillin (EuroClone), and 200 µg/ml streptomycin

sulfate (EuroClone). The cells were seeded onto the surface of

the samples (7mm x 7mm) by adding a medium which

contained 5.0 × 104 cells/ml. The cells were cultured for 24 h at

37 °C in 5% CO2 in air.

In study VI, human primary osteogenic sarcoma SaOS-2

(ECACC 890500205) cells were cultured in McCoy’s 5A culture

medium containing GlutaMAXTM (Gibco BRL/Life Technologies

Inc., Gaithersburg, MD, USA) and supplemented with 10% v/v

FCS, 100 IU/ml of penicillin and 100 µg/ml streptomycin. Cells

were seeded onto the biomaterial surfaces (10 mm x 10 mm) at a

density of 15-25 x 103 cells/cm2, and cultured for 48 h or 120 h at

37 °C in 5% CO2 in air.

Tissue culture-treated polystyrene 12-well (studies V and VI)

and 24-well cell culture plates (study III) supplied from Corning

Inc. (NY, USA) and Greiner Bio-One GmbH (Frickenhausen,

Germany), respectively, were used in all experiments.


98

7.5.2 MTS assay

The MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3 carboxymethoxy-

phenyl)-2-(4-sulfophenyl)-2H-tetrazolium)-based proliferation

assay (CellTiter 96® Aqueous One Solution Reagent, Promega

Corp., Madison, WI, USA) was used as a colorimetric method to

assess the cytotoxic effects of different PI grades (study III) and

also other potential material candidates for neural sensors

(additional data). The MTS assay, a modification from the more

conventional MTT (3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyl-

tetrasodium bromide) assay, is based on the reduction of MTS

salt in the presence of phenazine methosulfate to a water-

soluble formazan product in actively functioning mitochondria.

Since it is water-soluble, no solubilization step is required with

MTS as is the essential with the MTT assay. The quantity of

formazan formed was determined by measuring the absorbance

at 490 nm, which is directly proportional to the number of living

cells in the culture. After 24-h cultivation of the BHK-21 cells,

the samples were transferred into unused wells, and a fresh

DMEM medium plus MTS reagent (200 µl) was added. The

following incubation period lasted 1 h, after which the media

were removed, the wells were washed with PBS and the

absorbances were measured at 490 nm with an ELISA reader

(Molecular Devices, Inc., Sunnyvale, CA, USA). The relative

numbers of viable cells were determined by normalizing the

optical density values to the highest measured values.

7.5.3 Scanning electron microscopy of cultured cells

Scanning electron microscopy was used to analyze the cellular

behaviour on biomaterial surfaces. In studies V and VI, after

incubation period, samples were washed twice with PBS and

fixed overnight in 2.5 % glutaraldehyde (Sigma-Aldrich) at +4°C.

After this, the samples were washed with PBS and dehydrated

in a graded ethanol series (from 50 % solution to absolute

ethanol). The dehydration was completed using a Bal-Tec CPD

030 Critical point drying unit (Bal-Tec AG, Balzers,

Liechtenstein). Samples were mounted on SEM stubs, coated

with a Pt layer with an Agar sputter device (Agar Scientific Ltd.,


99

Stansted, UK) and examined using a Zeiss DSM 962 SEM (Carl

Zeiss GmbH, Oberkochen, Germany) at an accelerating voltage

ranging from 8 kV to 10 kV.

In study III, the cells were fixed with 2.5% (w/v)

glutaraldehyde (Sigma-Aldrich) in sodium cacodylate buffer

(pH 7.4) and dehydrated with ethanol gradient and HMDS

(Sigma-Aldrich). These samples were examined with an

environmental SEM (Philips XL30 ESEM-TMP, Fei Company) at

an accelerating voltage of 8–15 kV after being coated with a thin

layer of Au using a Sputter Coater E 5100 (Polaron Equipment).

7.5.4 Immunofluorescence and confocal laser scanning

microscopy

After either a 7.5 or 120-hour incubation period with the hMSCs

(study V) or after a 48-hour incubation period with the SaOS-2

cells (study VI), the specimens were triple stained with

fluorescence dyes to reveal cytoskeletal (actin) and focal

adhesion (vinculin) proteins and cell nuclei. A detailed

description of the fixation, staining and viewing of the cells is

given in original articles V and VI. Briefly, the cells grown on Si

pieces were washed with PBS solution, fixed with

paraformaldehyde solution, rinsed in PBS and permeabilized in

Triton X-100 solution. After immersing in normal goat serum,

the cells were incubated with the primary antibodies (anti-

vinculin) and washed. Cells were then incubated

simultaneously with fluorochrome-conjugated stains containing

Alexa fluor 488 (Molecular Probes Inc., Eugene, OR, USA) and

Alexa fluor 568 (Invitrogen Corp., Carlsbad, CA, USA) to detect

focal adhesion (vinculin) and actin cytoskeleton. Cell nuclei

were stained with DAPI (study V) or with TO-PRO-3 (study VI)

probes and then the samples were washed again, and mounted

on objective slides. Stained cells were viewed in a fluorescence

microscope (Olympus AX70) and in more detail using a Leica

TCS SP2 confocal laser scanning microscope (Leica

Microsystems GmbH, Mannheim, Germany) equipped with

appropriate filter sets.


100

In study V, the development of the osteoblastic phenotype

was followed using alkaline phosphatase (ALP), osteocalcin (OC)

and mineral staining (MS) as differentiation markers at 14, 17

and 21 days, respectively. A detailed description of the staining

of hMSCs is given in original article V. After staining, the

samples were fixed onto objective slides and coverslipped. The

cells were observed using an Olympus AX70 fluorescence

microscope in ALP and OC assays and a Leitz Diaplan

microscope (Leica Microsystems GmbH, Wetzlar, Germany) in a

mineralization assay.

In study III, specimens were stained using a combination of 2

fluorescent probes, i.e. a cell-impermeable DNA-binding dye,

propidium iodide (Sigma-Aldrich), and a cell-permeable

fluorescein diacetate (Fluka Chemie GmbH, Buchs, Switzerland)

in PBS to assess the viability of the BHK-21 fibroblasts cultivated

for 24 hours on the samples. Detailed information on staining

procedure can be found in original article III. Fluorescein

diacetate is a non-fluorescent molecule that penetrates through

the cell membranes. Inside the cell, fluorescein diacetate is

hydrolyzed by non-specific esterases resulting in fluorescein

release. Fluorescein emits green fluorescence when irradiated by

laser light. Propidium iodide is a common fluorescent molecule

used to stain DNA and to identify dead cells in a cell population.

Since propidium iodide is membrane impermeable, it only

stains dead cells which have leaking membranes. Propidium

iodide emits a red fluorescence when excited. Since only viable

cells are able to produce fluorescein, these two stains can be

used to separate the living and dead cells in a cell populations.

Cells are either stained with propidium iodide or fluorescein

diacetate, but not with both. The samples were viewed in a

confocal scanner (PerkinElmer Life Sciences, Wallac-LSR,

Oxford, UK) on a Nikon Eclipse TE300 microscope (Nikon Corp.,

Tokyo, Japan), using the wavelengths 488/10 nm (excitation) and

525/50 nm (emission) for the fluorescein, and 568/10 nm

(excitation) and 607/45 nm (emission) for the propidium iodide.

Finally, the numbers of dead (red-staining) cells (nuclei/mm2)

and the surface area of live (green-staining) cells were counted


101

using Image J software 1.37c (National Institute of Health,

Bethesda, MD) (Abramoff et al. 2004).

7.6 STATISTICAL ANALYSES

The numerical results are expressed as mean ± standard

deviation of the mean (SD) or as mean ± standard error of the

mean (SEM). In studies III, V and VI, statistical analyses were

conducted with the SPSS software (versions 14.0 and 16.0, SPSS

Inc., Chicago, IL, USA). One-way analysis of variance (ANOVA)

followed by Tukey Post-Hoc Tests were applied to determine

the statistical significance of the differences observed between

material groups. Values of p < 0.05 were considered statistically

significant.

http://www.citeulike.org/user/christofseiler/author/Abramoff:MD


102

103

8 Results

8.1 EVALUATION OF MICROELECTRODE ARRAYS

A number of MEA prototypes were developed during the 2-year

trial period by optimizing the substrate material, substrate

thickness, electrode material, sensor layout and insulation

material (Table 6). It was demonstrated that the magnetron

sputter deposited and photolithographically patterned Pt thin

films between two layers of PI represented a very powerful and

reproducible way to produce miniaturized neural interfaces (Fig.

14). The ZIF type connector allowed the easy and reliable

connection of a thin film MEA to recording instrumentation (Fig.

14b). The use of the ZIF connector eliminated the need for

soldering, which had been observed to be very challenging to

perform on the thin Pt pads produced by the sputtering

technique on PI. Furthermore, the use of PSPI (PI-2771)

simplified the electrode fabrication considerably because this

material could be patterned directly in the same way as

photoresists without using complex multilevel processes (i.e.

lithography, etching mask formation, etching mask dissolution)

which were unavoidable steps when conventional PIs were used.

It was possible to pattern the PI-2771 layer with good alignment

to the metallization layer (Fig. 14f,g). Moreover, one could

pattern the PI-2771 layer with an almost vertical pattern profile

and without residuals (Fig. 14h).

The electrode type B, originally presented in study II, was

found to be the most suitable for recording cortical surface field

potentials in rats. The major shortcomings of the other

prototypes compared to type B are summarized in Table 6.

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104

A B

C D

E F

G H

recording

sites

connector board

transmission

lines 5 mm

5 mm

1 mm 1 mm

200 µm 50 µm

Figure 14: The developed polyimide-based microelectrode arrays (a). The PI ribbons

are connected via a connector board consisting of a ZIF- type connector (JST Ltd.) and

a surface mount microsocket (Samtec Inc.) to the recording instrumentation (b). The

developed array is highly miniaturized as shown in comparison to a clinical subdural

grid electrode (c), and extremely flexible (d). A SEM-image of the 8-channel array

viewed at the end of the recording sites before deposition of the insulation layer (PI-

2771) (e). Patterning of this polyimide insulator layer achieved a good alignment with

the metallization layer (f,g). The 3 µm thick polyimide layer had an almost vertical

pattern profile and was virtually free of residuals (h).

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105

Table 6: The summary of developed microelectrode array prototypes. The main

disadvantages of each array type compared to type B are presented.

Type Channel

number

Substrate Electrode

diameter (µm)

Insulator Weaknesses

A 8 PI-2525a (3 layers)

Ti/Pt, 200 PI-2771a Troublesome fabrication

B 8 Kapton HNb

+ PI-2525 Ti/Pt, 200 PI-2771 –

C 8 Kapton HN + PI-2525

Ti/Au, 200 PI-2771 Higher electrode impedances

D 16 Kapton HN

+ PI-2525

Ti/Pt, 100 PI-2771 Short-circuiting

E 16 Kapton HN

+ PI-2525

Ti/Pt, 100 SU-8c Short-circuiting

and cracking

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

Due to the viscosity and chemical properties of PI-2525, the

formation of a 30 µm thick substrate layer required three spin

coating cycles (type A). This was somewhat complicated

compared to type B (study II), in which the same thickness

could be readily achieved by using Kapton HN film together

with one spin-coated PI-2525 layer. Gold electrodes (type C)

have significantly higher electrode-electrolyte interface

impedance levels compared to Pt electrodes especially at low

frequencies below 1 kHz which are the most relevant when

biological signals need to be recorded (Fig. 15). The

microelectrode surface area in high-density arrays (types D and

E, electrode diameter 100 µm) was only 25 % of the area of the

larger electrode (types A-C, electrode diameter 200 µm) causing

significantly higher electrode impedances.

8 – Results

106

Au

Pt

10 102 103 104 105

frequency (Hz)

|Z| (

Ω)

Pha

sean

gle

(°)

107

106

105

104

103

-20

-40

-60

-80-81.0

-74.1

95.5 k

25.5 k

Au

Pt

Figure 15: Impedance spectroscopy of platinum (Pt) and gold (Au) microelectrodes

(diameter 200µm), measured in 0.9 % saline solution at room temperature. The

impedance magnitude and phase are presented with error bars corresponding to

standard deviations (SD). In some instances, SD values were often smaller than the

symbol size and thus not shown. The absolute values of the impedance magnitude and

the phase are also given at 1 kHz.

Designs D and E were also susceptible to short-circuiting by

cerebrospinal fluid when the array was implanted onto the

surface of the rat cortex. Moreover, epoxy-resist SU-8 (type C)

was observed to be totally unsuitable for present electrode

application. While PI- insulated arrays were very robust and

they could be bent even into sharp angles without any cracks,

the bending/implantation did create many cracks and even

delaminations of the SU-8 layers.

The developed electrodes were tested in acute and chronic

evoked potential recordings in Wistar rats. The flexibility of PI

was particularly suitable for following the curved cortex surface,

8 - Results

107

thus ensuring good contact between electrode site and the dura

mater. This enabled reliable acquisition of cortical signals.

During the acute recording sessions, the MEA type B was

capable of capturing stable recordings as observed in the form

of the standard response parameters, such as latencies, onsets

and decays of the main components in the potential traces (Fig.

16a). The recorded signal traces were biologically meaningful

and reproducible, this being verified by comparing the

responses at the beginning of the experiments with those

obtained at the end. In chronic recordings, the electrode was

capable of yielding signals comparable with the presented acute

signals for approximately two weeks after the array

implantation (Fig. 16b). However, after 16 days, the responses

decayed. This decline in signal amplitudes was concurrent with

the observed jump in the electrode impedances which were

systematically measured in vivo at three-day intervals (Fig. 16c).

The physical appearance of the dura around the electrode

occasionally changed during chronic implantation. It became

thicker and encapsulated the electrode. However, in microscopic

examination, the surface of the brain appeared to be undamaged

with the electrode location being slightly depressed. Although

no macroscopic scar formation was seen in the examination of

the brain after the array was removed, this kind of jump in the

in vivo impedances was compatible with a thickening of the

dura, a common finding in chronic experiments, and

microscopic growth of non-conductive fibrous tissue around the

electrodes, shrinking the freely exposed surface area of the

recording electrode.

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108

Figure 16: (a) An acute recording of somatosensory evoked potentials (SEP) in the rat

parietal cortex. Profiles of averaged SEP traces using the 8-channel microelectrode

array are shown on the right with their correspondence to the recording sites which are

numbered in schema in left side. The position of reference/ground electrode is marked

as “R”. The current stimuli consisted of 25 sets of paired pulses with pulse intervals of

500 ms, followed by a 9.5 s quiet interval, delivered to the left front paw of the rat. (b)

A chronic recording of auditory evoked potentials in the rat parietal cortex. The stimuli

consisted of 25 sets of paired pulses at intervals of 500 ms, followed by a 9.5 s quiet

interval. (c) Average electrode impedance (mean ± SD, n = 8) measured in vivo.

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109

8.2 EVALUATION OF NANOROUGH ELECTRODE SURFACES

The USPLD technique was used to prepare the Pt coatings on

SS, PI and Si with a desired nanostructured surface topography.

The adhesion of coatings deposited upon all the tested

substrates was good without any signs of delamination. The

USPLD process parameters were adjusted to achieve different

surface topographies. Fig. 17 shows an example where the

feature size is about 30 nm. The nanostructured electrode

surface clearly increased the effective surface area compared to a

smooth Pt surface typically achieved using sputtering.

Figure 17: AFM image of nanostructured Pt-electrodes produced using the USPLD

technique. The average roughness of the surface is 18 ± 4 nm (mean ± SD)

Electrochemical impedance spectroscopy was applied in

order to study the electrode-electrolyte impedance behaviour of

different Pt-coated SS electrodes (unpublished data). A total of

six sample areas (50 mm2) of each plate were examined using

the measurement set-up described in chapter 7.3.1. The mean

impedance magnitude and phase data for Pt coating, whose

surface topography was demonstrated in Fig. 17, are shown in

Fig. 18. Pure SS electrode and sputter-coated Pt on SS are

8 – Results

110

presented as reference materials. The results clearly reveal that

the presence of the nanostructures on an electrode surface had

increased the surface area and decreased impedance. The

impedance magnitudes (mean ± SD) at 1 Hz were 7.0 kΩ ± 0.5

kΩ, 17.8 kΩ ± 1.1 kΩ and 21.7 kΩ ± 1.0 kΩ for nanostructured

USPLD-Pt, sputtered-Pt and pure SS, respectively. The

differences in impedance values between the materials were

statistically significant (p < 0.001) at low frequencies. Due to the

relative large geometrical surface area, all coatings reached the

same impedance magnitude level of approximately 180-200 Ω

by 1 kHz, which corresponds to the resistance value of

electrolyte solution. All surfaces are highly polarizable, but

nanostructuring decreases the phase difference between current

and voltage making electrodes less polarizable.

100 101 102 103 104 105102

103

104

105

Frequency (Hz)

|Z| (o

hm

)

SS.datSputtered Pt.datUSPLD-Pt.dat

-80

-60

-40

-20

0

Ph

ase

an

gle

(°)

Figure 18: Impedance magnitude and phase of a nanostructered platinum-coating

(USPLD-Pt) and sputter-coated Pt-coating deposited on stainless steel (SS). Results

for pure SS electrode are presented for reference. A clear drop in electrode impedance

can be seen for USPLD-Pt, especially at low frequencies. Note the logarithmic scale.

8 - Results

111

8.3 CYTOCOMPATIBILITY TESTING OF INSULATOR

MATERIALS

Prior fibroblastic BHK-21 cell experiments, spin-coated PIs (PI-

2525 films and PI-2771 films cured at 200 °C and at 350 °C) and

USPLD coatings (C3N4 and Al2O3) were characterized in terms of

their roughness, wettability and zeta potential (Table 7). The

surface roughness values obtained were all very low, i.e. the Ra

values were below 1 nm. However, the Ra value for the PI-2525

was significantly higher than for other coatings (p < 0.001). The

contact angles for water on both PI-2771 film types were

significantly higher than for the PI-2525 (p < 0.001). The same

tendency was also noticed when contact angles were measured

for diiodomethane (results are shown in the original article III,

Table 1) highlighting the more hydrophobic nature of the PI-

2771 with significantly lower polar and dispersive surface

energy values (p < 0.001). Somewhat surprisingly, the effect of

the chosen curing protocol (200 °C vs. 350 °C) for the PI-2771 on

roughness or on wettability was not statistically significant. The

contact angles for water on C3N4 were in the same range as on

PI-2525 film, evidence of a slightly hydrophobic thin film

coating. In contrast, the Al2O3 coating was very similar to PI-

2771 film with respect to its wettability properties. All of the

surfaces were negatively charged, and the zeta potential values

measured at a fixed pH of 7.4 were in a narrow range of -54.2

mV – -66.2 mV. The PI-2771-200 film followed by the PI-2525

film was the most negatively charged and their zeta potential

values differed significantly from the all other test materials (p <

0.001).

8 – Results

112

Table 7: Average surface roughness (Ra), water contact angle and zeta potential

values of tested polyimide and ceramic coatings. All of the materials were extremely

smooth although the roughness of PI-2525 was significantly higher than that of other

materials. The PI-2525 and carbon nitride (C3N4) were more hydrophilic compared to

alumina (Al2O3) and both PI-2771 types. Zeta potential values were all in the same

range by demonstrating the negative surface charge of these coatings.

Material Ra (nm) Contact angle (°) Zeta potential (mV)

PI-2525 0.89 ± 0.13 * 69.5 ± 1.8 ## -59.9 ± 0.3 *

PI-2771-200 0.43 ± 0.15 95.1 ± 1.3 -66.2 ± 1.6 *

PI-2771-350 0.39 ± 0.04 93.2 ± 0.8 -54.2 ± 0.1

C3N4 0.24 ± 0.05 # 70.8 ± 0.9 ## -55.1 ± 1.2

Al2O3 0.18 ± 0.03 # 94.0 ± 1.3 -56.6 ± 3.4

* p < 0.001, as compared to all other materials # p < 0.001, as compared to all polyimides ## p < 0.001, as compared to all other materials, except of C3N4/PI-2525

The values are mean ± SD. One-way ANOVA was used to determine the statistical differences between materials.

In this study, the cytotoxicity of the insulator candidate

materials was examined with in vitro tests conducted along to

the guidelines of the international standard ISO-10993-5. Latex

rubber and PE were used as positive and negative control

material, respectively. The MTS assay defined the number of

viable BHK-21 cells attached on the surfaces 24 h after cell-

seeding. The results were normalized relative to the highest

value, i.e., the value for the PE control (Fig. 19). The relative cell

numbers in the different polyimide groups and C3N4 were at a

level of 62-70 % of the PE controls. The number of cells attached

to the Al2O3 surface was lower, i.e. 50 %, but still significantly

higher than attained with the latex rubber control.

Fluorescence microscopy studies with live/dead staining

probes revealed that all test materials were highly

cytocompatible, in contrast to latex rubber, on which viable cells

were rarely found, instead dead cells had most probably

detached and were floating in the medium (illustrated in the

original article III). Results from the quantitative analysis

included the determination of the density of dead cells

8 - Results

113

(nuclei/mm2) and the surface area of live cells (µm2) is presented

in the Table 8. The numbers of dead cells were very low on each

surface, i.e., mean values ranged from 13 cells/mm2 to 48

cells/mm2. Even though the dead cell density was also minimal

on the surface of latex rubber, the cytotoxic effect could be

clearly seen when the surface area of viable cells was examined.

The total surface area of viable cells on latex rubber was

significantly lower (p < 0.001) than that of the other test

materials, being only around 2% of the corresponding values

obtained for other surfaces. The values for C3N4 and Al2O3 were

slightly but not significantly higher than that for PE control or PI

materials.

Table 8: The density of the dead cells (nuclei/mm2) and the surface area of viable cells

(µm2) on polyethylene (PE, negative control), different polyimide surfaces, alumina

(Al2O3), carbon nitride (C3N4) and latex rubber (positive control) after a 24-hour

cultivation period evaluated using a combination of two fluorescent probes and Image J

software.

Material Density of dead cells

(1/mm2)

Surface area of viable

cells (µm2)

PE (n = 10) 13.9 ± 5.4 70400 ± 2900

PI-2525 (n = 13) 13.5 ± 5.5 71900 ± 8500

PI-2771-200 (n = 8) 18.8 ± 8.2 67300 ± 13200

PI-2771-350 (n = 8) 30.8 ± 8.8 63000 ± 5800

C3N4 (n = 13) 48.0 ± 10.3 83900 ± 11600

Al2O3 (n = 15) 47.4 ± 5.8 82600 ± 12300

Latex rubber (n = 8) 24.9 ± 8.6 1400 ± 320*

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

SEM imaging was used to monitor cell morphology,

adhesion and spreading on the test surfaces. The BHK-21 cells

which had been cultivated for 24 h on each PI surface, Al2O3 and

C3N4 exhibited a normal fibroblastic morphology, adhered well

with filamentous extensions, and had started to form clusters

(Fig. 20). Thus, all insulator materials were considered not to be

toxic and equally good for allowing the adhesion and spreading

of cultured fibroblasts.

8 – Results

114

0

20

40

60

80

100

120

PE PI2525 PI2771-200 PI2771-350 Alumina CN Latex

Re

lative

ce

ll n

um

be

r (%

)*

*

0

5

10

15

20

PE PI-2525 PI2771-200 PI2771-350 alumina CN Latex

Ce

ll co

ve

rag

e (

%)

*

PE PI-2525 PI-2771-200 PI-2771-350 Al2O3 C3N4 Latex

PE PI-2525 PI-2771-200 PI-2771-350 Al2O3 C3N4 Latex

A

B

Figure 19: (a) Relative number of BHK-21 cells on polyethylene (PE, negative control),

conventional polyimide (PI-2525), two differently cured photosensitive polyimides (PI-

2771-200 and PI-2771-350), alumina (Al2O3), carbon nitride (C3N4) and latex rubber

(positive control). The data was obtained by MTS assay and normalized to the highest

optical density value. (b) Surface area covered by BHK-21 cells at 24 h on the same test

materials. The data was acquired from confocal laser scanning microscope images of

the samples with live/dead-stained cells. All polyimide and ceramic coatings can be

considered to be non-cytotoxic and they allowed adhesion and spreading of fibroblasts.

The error bars indicate standard errors of the means. *p < 0.001 (one-way ANOVA).

8 - Results

115

Figure 20: SEM images of BHK-21 cells cultured for 24 h on PI-2771-200 at two

different magnifications. The cells exhibited a normal fibroblastic morphology, adhered

well with filamentous extensions, and had even started to form clusters. Cells cultured

on the other test materials, i.e. PI-2525, PI-2771-350, C3N4 and Al2O3, appeared to be

very similar and thus, are not shown.

8.4 MICROPATTERNED BIOMATERIAL COATINGS

The interactions of the hMSCs and SaOS-2 cells on

micropatterned biomaterial (Ti, Ta, Cr, DLC) coatings were

investigated in studies V and VI. The roughness and surface

energy data of test materials were originally published in paper

V, but these values are also valid for the samples used in study

VI. A summary of these results, supplemented with the zeta

potential values achieved for smooth (non-patterned) surfaces is

given in Table 9. All of the materials used in these studies were

extremely smooth, with Ra values of less than 2 nm leading to

the hypothesis that the observed differences in cell behaviour

were not due to the differences in surface topography but were

mediated by different chemical compositions and physical

patterns of material studied. The only relevant topographical

cues of the samples were attributable to the thickness of coating

(~ 200 nm) that forms steps between micropatterns and the

background. All coating materials also possessed very similar

wettability properties, i.e. their SFE values were in a range

between 47.3 - 52.1 mJ/m2. The Si wafer used as background

material in these experiments had a significantly higher SFE

(68.5 mJ/m2) compared to all other coating materials. The

negative zeta potential values obtained from streaming current

8 – Results

116

measurements at a fixed pH of 7.4 disclosed that the surfaces of

all coating materials were negatively charged. The differences in

zeta potential values between the surfaces were statistically

significant (p < 0.001 for all).

Table 9: Average surface roughness (Ra), water contact angle, surface free energy and

zeta potential values for the surfaces of different test materials. All surfaces were

extremely smooth (i.e. mirror finish) and their contact angle/surface energy values

were in a narrow range (except for background Si), although some statistically

significant differences existed.

Material Ra (nm) *

Contact angle (°)

Surface energy (mJ/m2)

Zeta potential (mV) *

DLC 0.3 ± 0.1 67.4 ± 2.0 49.6 ± 1.2 -70.2 ± 1.4

Ti 1.8 ± 0.2 67.6 ± 2.0 49.7 ± 1.1 -53.9 ± 0.3

Ta 1.5 ± 0.3 64.1 ± 1.6# 52.1 ± 1.1# -49.6 ± 1.1

Cr 0.8 ± 0.1 71.2 ± 1.1# 47.3 ± 0.7# -39.6 ± 2.2

Si 0.2 ± 0.1 32.4 ± 1.1# 68.5 ± 0.8# -43.5 ± 1.3

*p < 0.005 between all materials #p < 0.05, as compared to all other materials

The values are mean ± SD. One-way ANOVA was used to determine the statistical differences between materials.

8.4.1 Behaviour of human mesenchymal stem cells

An incubation time of 7.5 hours was used to investigate the

early-stage adhesion, contact guidance and morphology of

hMSC cells on the micro-patterned DLC, Ti, Ta and Cr thin films

deposited on Si (study V). Based on observations with the SEM,

hMSCs seemed to prefer the biomaterial patterns over a Si

background. This kind of guiding effect seemed to be evident

with all materials, although only Ti and DLC surfaces are

presented here (Fig. 21a,b). Quantitative analysis showed that

on Ta, Cr and DLC, the density of hMSCs was 3.0-3.5 times

higher than on the background (p < 0.0005 for all, Fig. 22). One

exception was Ti, on which the density of hMSCs on the

patterns did not differ from the density of the cells grown on the

Si background (p = 0.14). The analysis of the surface area of

patterns and Si background covered by the MSCs at 7.5 h gave

8 - Results

117

results consistent with the results from the cell density

assessment (Fig. 23). From this viewpoint, the guiding effect was

also clear on all the materials (p < 0.05) except for Ti (p = 0.46).

Figure 21: hMSC cells after 7.5 h culture (a,b) and 5 days culture (c,d) on patterned

titanium (a,c) and diamond-like carbon (b,d) surfaces. The cells seemed to prefer the

patterns over background and some cells were aligned along the edges of the patterns

(a,b). At 5 days, cells were much larger and they had grown to form a subconfluent

monolayer. The side length of the biomaterial square is 75 µm.

One interesting observation was made about the size of cells.

The mean diameter of the cell lying on patterns (68 µm) was

significantly lower (p < 0.05) than that of cells growing on a Si

background (102 µm). Moreover, it was observed that numerous

cells were aligned along the edges of the patterns. This

phenomenon was confirmed in the immunofluorescence triple

staining, which revealed that focal vinculin adhesions and actin

cytoskeleton were outgoing from the pattern edges i.e. cells were

believed to be taking on the geometric square shapes.

8 – Results

118

140

120

100

80

60

40

20

Cel

lden

sity

(cel

ls/m

m2 )

DLC Cr Ta Ti

Background (Si)

Patterns

Average

Figure 22: Density of hMSCs at 7.5 hours. On lithographically patterned diamond-

like carbon (DLC), chromium (Cr) and tantalum (Ta) samples, the cells significantly

preferred the biomaterial patterns over a silicon (Si) background. The cell density on

titanium (Ti) patterns was marginally lower (# p ≈ 0.06, one-way ANOVA) than that

on Cr and Ta patterns. The error bars indicate standard deviations of the means.

0

20

40

60

80

100

DLC (7.5 h)

DLC (5 d)

Cr (7.5 h)

Cr (5 d)

Ta (7.5 h)

Ta (5 d)

Ti (7.5 h)

Ti (5 d)

Cell

covera

ge (

%)

Background (Si)

Patterns*

**

*

+

#

Figure 23: Coverage of hMSCs at 7.5 hours and 5 days. A clear guiding effect seen at

7.5 hours was nearly lost by 5 days. * p < 0.05, as compared to DLC and Cr, # p < 0.05,

as compared to Cr and Ta, + p < 0.05, as compared to DLC. The error bars indicate

standard deviations of the means. One-way ANOVA was utilized to determine the

statistical differences.

8 - Results

119

The spreading of the hMSCs was investigated after 5 days of

incubation (Fig. 21c,d). At this stage, the individual cells had

grown considerably achieving a size which greatly exceeded

that of a singular square pattern (75 µm x 75 µm), which meant

that it was no longer possible to calculate the cell densities.

Surface coverage analysis indicated that over 65 % of the sample

surface (of each material) was covered by cells at 5 days (Fig. 23).

The guiding effect was virtually lost since the cell density was

only 1-1.9 times higher on patterns than that on Si background.

The only statistically significant difference was observed

between the DLC patterns and the Si background. Somewhat

surprisingly, the cell density on the very same Si background

varied statistically significantly, depending on the material used

in microfabricated patterns. At 7.5 h, the hMSC-covered surface

of the background Si was significantly lower on Ti than on DLC

patterns (p = 0.015), but by the 5th culture day this situation had

reversed (p = 0.006) (Fig. 23). This indicates that apart from

direct material-cell interactions, also probably indirect cell-cell

interactions were present. The cells on material islands had

remote interactions with cells growing on Si background. After

5 days, vinculin containing focal adhesion were already well

developed all over the sample surface without any preferential

localization on the square patterns as compared to the

background.

Osteogenic differentiation of hMSCs was investigated by

using ALP, OC and MS as the differentiation markers. The

results showed that the patterned surfaces, except for DLC,

allowed induced osteogenic differentiation although this was

less effective than on plain samples. It seemed to be so that cells

growing on the micropatterned surfaces did not have enough

space to organize themselves into multicellular bone tissue-type

structures. Moreover, the rectangular shape which the hMSC

cells took due to microsquare shaped islands might not be ideal

with respect to osteogenesis. The DLC coating did not support

osteogenesis, and for that reason an inert DLC seems to be a

potential coating material for applications which come into

8 – Results

120

contact with bone but which need to be removed later, e.g.

various fracture fixation devices, such as plates and screws.

8.4.2 Behaviour of osteoblast-like (SaOS-2) cells

The SEM examination of SaOS-2 cells cultured on micro-

patterned Ti and DLC coatings (on Si) for 48-hours revealed a

clear cellular preference for the biomaterial patterns over the Si

background. The cells preferentially adhered and spread well on

the large-sized (125 µm) circular (Fig. 24a) or rectangular (Fig.

24b) Ti patterns, which facilitated their occupation by several

cells. On the reverse-patterned samples containing Si features

(circles or squares) on a Ti background (Fig. 24c), the inverse

phenomenon was observed, with almost all of the cells now

attached to the Ti background, but they seemed to avoid

spreading out into the Si circles. The cytocompatibility-

enhancing and cell-guiding effects of DLC were also clear,

though much weaker than those encountered with Ti, as

demonstrated in samples containing DLC circles on Si (Fig. 24d)

and in the reverse-patterned samples containing Si squares on a

DLC background (Fig. 24e).

The influences of the pattern shape and its size on adherence

and morphology of SaOS-2 cells were studied using rectangular

and circular patterns of three different sizes (diameters of circles

and sides of squares: 5, 25 and 125 µm). Large-sized patterns

facilitated the adhesion of several cells onto one island and the

boundaries of the outermost cells followed the shape of the

edges of the patterns to which they had attached (Fig. 25a,b).

The medium pattern size (25 µm) was designed so that singular

cells could inhabit individual patterns. This pattern size seemed

to restrict the growth of cells although they no longer adapted

so strictly to the geometrical shape of the patterns (Fig. 25c).

However, neighboring patterns were so far away that cells

could not really join together to form clusters. Particularly on

circularly patterned surfaces, cells occasionally covered two or

more islands, leading to longitudinal or star-like morphologies

(Fig. 25d). On the small-sized patterns, cell bodies covered

several of the small patterns (Fig. 25e), but their thin and slender

8 - Results

121

filopodia seemed to prefer the more cell-friendly material, and

they seemed to try to avoid bare Si areas (Fig. 25f).

Figure 24: SaOS-2 cells cultured for 48 h on surfaces containing large-sized (125 µm)

features. There is a clear cellular adhesion preference for the Ti circles (a) or squares (b)

over background Si, and in the case of reverse-patterned samples for the Ti background

over the Si circles (c). The cell-guiding effect of DLC was also clear even though

weaker than that of Ti as demonstrated in samples containing DLC circles on Si (d)

and in reverse-patterned samples containing Si squares on a DLC background (e).

Scale bar is 100 µm.

Immunofluorescence triple staining disclosed that the actin

cytoskeleton was well organized in SaOS-2 cells growing on the

biomaterial islands, but poorly organized in the cells lying on Si.

8 – Results

122

Furthermore, it was observed that the formation of vinculin

containing focal adhesions was much enhanced in cells adhering

onto the biomaterial patterns compared to cells lying on the Si

backgrounds.

Figure 25: SEM images of SaOS-2 cells cultured for 48 h on micropatterned Ti (a-d)

and DLC (e,f) surfaces. Large-sized (125 µm) squared (a) and circular (b) features

facilitated the adhesion of several cells on one Ti island with the cells aligning

themselves along the edges of the cell-friendly material. The cells adhering to medium-

sized Ti islands no longer conformed strictly to the geometrical shape of the patterns (c)

but particularly on circularly patterned surfaces, star-like cellular morphologies

appeared (d). On small-sized inverse DLC samples, the cell bodies non-selectively

covered large micro-patterned areas (e), but their filopodia clearly showed a preference

for DLC trying to avoid bare Si circles (f).

8 - Results

123

Quantitative analysis of the SaOS-2 cell density, coverage and

average cell size was performed with samples containing 75 µm

squared patterns of Ti, Ta, Cr and DLC on the Si substrate. The

relative cell density at 48h was 2.9-3.7 times higher on all four

coating materials than that achieved on the Si background (p <

0.001). The relative coverage was 2.0-3.0-fold higher for Ti, Ta,

Cr and DLC compared to the corresponding value of Si (Fig. 26).

A comparison between the four patterning materials indicated

that the proportions of surface area covered by cells were higher

for Ti and Ta than for DLC and Cr (p < 0.05). Further, the Si

background was better covered with the Ti- and Cr-patterned

samples than the background of the DLC and Ta samples (p <

0.001).

DLC Cr Ta Ti

Cel

lcov

erag

e(%

)

80

60

40

20

Figure 26: Coverage of SaOS-2 cells was always higher on biomaterial patterns than

on silicon (Si) background. Titanium (Ti) and tantalum (Ta) were better covered than

diamond-like carbon (DLC) and chromium (Cr) as designated with # (p < 0.05) and

the Si background on Ti- and Cr patterned samples was better covered than the

background on DLC or Ta samples (##, p < 0.05). The error bars indicate standard

deviations of the means. One-way ANOVA was applied to determine the statistical

significance of the differences observed between material groups.

8 – Results

124

The average sizes of the cells cultured for 48-hours on each

sample type are given in Table 10. The cells growing on the

biomaterial patterns (each material) were significantly smaller

(13-41 %) than the cells growing on Si. A comparison between

different biomaterials disclosed that the cells growing on Ti

were larger than the cells on Cr (p < 0.005), and the cells on Ta

patterns were larger than the cells on DLC or Cr patterns (p <

0.05). Further, the cells growing on Ti-patterned samples were

also significantly larger on the Si background than on

backgrounds containing patterns fabricated from other coating

materials (p < 0.01). This was somewhat surprising since the

background material and the preparation techniques used in

sample fabrication were the same for all samples.

Table 10: The size of SaOS-2 cells on micropatterned samples. The SaOS-2 cells were

significantly larger on background silicon compared to the cells growing on patterns.

In addition, the cells were significantly larger on the Si background of Ti-patterned

samples than on the backgrounds of other materials.

Material Cell size on Si background (µm2)

Cell size on patterns (µm2)

DLC 1282 ± 60 962 ± 45

Cr 1404 ± 102 954 ± 25

Ta 1322 ± 86 1151 ± 65#

Ti 1786 ± 77* 1059 ± 20**

* p < 0.01, as compared to all other backgrounds

# p < 0.05, as compared to DLC and Cr

** p < 0.005, as compared to Cr

The values are mean ± SD. One-way ANOVA was used to determine the statistical

differences.

125

9 Discussion

9.1 MICROELECTRODE ARRAYS FOR INTRACRANIAL

RECORDINGS

In studies I, II and IV, the design, fabrication and experimental

use of novel flexible polymer-based microelectrode arrays was

demonstrated. Successful acute and chronic intracranial multi-

channel recordings performed on the surface of the rat cortex

clearly revealed the potential of these arrays. In particular, the

PI-Pt-PI sandwich type was proved to be capable of acquiring

stable signals with a high signal-to-noise ratio in evoked

potential recordings. The PI-based arrays demonstrated

excellent flexibility and mechanical strength during handling

and implantation onto the surface of rat cortex. The flexibility

also allowed the array to follow the movements of the rat brains

by keeping the recording points in close contact with the target

area, thus promoting reliable signal acquisition. Although SU-8

epoxy-resist has been recently utilized as an insulating material

in implantable neural electrodes (Hollenberg et al. 2006, Yeager

et al. 2008), implantation (bending) of SU-8 insulated array

developed here caused several cracks in the SU-8 layer.

Sputtering and evaporation are traditional techniques used in

producing thin film microelectrodes. The results obtained from

EIS with our magnetron sputtered Pt and Au microelectrodes

were consistent with the previously reported values (Cheung et

al. 2007, Metz et al. 2004, Stieglitz et al. 2000, Takahashi et al.

2003). The Pt electrodes displayed significantly lower electrode-

electrolyte impedance magnitudes compared to the Au

electrodes, particularly at biologically relevant frequencies, i.e.

below 1 kHz. With the help of AFM imaging, the surface of

sputtered Pt or Au was observed to be extremely smooth (Ra < 1

nm) i.e. its effective surface area was nearly the same as its

geometrical dimensions. The impedance levels for the Pt

9 – Discussion

126

electrodes used in the final array prototypes (particularly for

electrodes with diameter of 200 µm) were adequate, but if one

needs to develop smaller-sized neural electrodes for high-

resolution application, it would be advantageous to increase the

effective surface area of microelectrodes.

The fabrication process developed here enables the

fabrication of miniaturized arrays with high channel densities,

which have benefits over microwire bundles in terms of precise

control of the electrode sizes and separations between

electrodes. The lithography-based process also allows

implementation of several other neural interfaces with different

layouts, material selections or target areas either for recording or

stimulation purposes. Furthermore, these reproducible devices

can be mass produced in batches at low cost. As concluded in

paper IV, the technology developed in this thesis work also

offers several attractive applications for clinical use. The high

density arrays could possibly be utilized to achieve more

accurate mapping the brain e.g. in patients suffering from

epilepsy, Parkinson’s disease or brain tumors, enabling more

precise planning of the neurosurgery. The flexibility also eases

the insertion of the array without the need for removal of large

skull areas, in contrast to the commercially available, somewhat

bulky, designs of subdural strip and grid electrodes. In addition,

it is obvious that prosthetic systems used in electrical

stimulation of neural tissue in order to treat neuropathic pain,

replace lost sensory or motor function will also benefit from

further development of this kind of technology.

As a result of study IV, a comprehensive summary of

miniaturized flexible microelectrode arrays was obtained. Some

of these MEAs have been developed from the point of view of

human applications (Kitzmiller et al. 2006, 2007, Molina-Luna et

al. 2007, Takahashi et al. 2003, Hollenberg et al. 2006). The

technique of small-field cortical surface recordings, introduced

by Kitzmiller et al. (2006) was based on lithographically

patterned Pt electrodes on the PDMS sheet. Although being

highly miniaturized compared to present clinical electrodes, this

design did not take full advantage of all that microfabrication

9 - Discussion

127

can offer but used fragile wire bonding to create the electrical

wiring. It was realized that the transmission lines could be

formed in the same deposition step as the electrodes. Thin film

transmission lines firmly attached to the substrate surface were

demonstrated to represent a robust way to establish the signal

acquisition. By using somewhat similar technology as applied in

the present thesis, Molina-Luna et al. (2007) introduced

epidurally implantable thin-film MEA for cortical stimulation

mapping. Their approach was based on 72 round TiN contacts

(diameters 100 µm) deposited on the surface of PI foil which

was utilized for mapping the motor cortex of the rat. In

addition, a Kapton-based MEA consisting of 8 x 8 Au electrodes

has been developed for epidural recording of surface field

potential from the barrel cortex of rats (Hollenberg et al. 2006).

However, the array developed by Hollenberg et al. is thicker (75

µm) and thus stiffer than the system developed here and it is

also unsuitable for chronic recordings due to the bulky circuit-

board that the rat would have to carry on its head. A flexible PI-

based MEA with a bending structure introduced by Takahashi

et al. (2003) was successfully tested in epidural mapping of AEP

recordings. However, that system was not designed for chronic

implantation. On the contrary, the connector board with the two

miniconnectors used in our arrays was shown to be small and

light enough to allow chronic recordings.

The acute EP recordings obtained with the present MEAs

were comparable in their signal-to-noise ratios with previously

reported recordings with wire electrodes made with platinum

(Hayton et al. 1999), tungsten (Ureshi et al. 2004) or silver

(Kalliomäki et al. 1998, Jones & Barth 1999). They were also

comparable to previous recordings obtained with MEAs (Stett et

al. 2003, Hollenberg et al. 2006, Takahashi et al. 2003). The

quality of the recorded signal at the beginning of chronic EP

experiments was similar to those obtained in acute experiments

for about two weeks. Subsequently, the signal amplitude

declined due to the increase in electrode impedance resulting,

most probably from the thickening of the dura and the growth

of electrically insulating fibrous tissue around the electrodes.

9 – Discussion

128

9.2 MATERIAL SCIENCE STRATEGIES TO IMPROVE THE

PERFORMANCE OF MICROELECTRODE ARRAYS

Due to the limited long-term performance of the developed

arrays, the confirmation about biocompatibility of materials

used in those arrays was sought in study III. Moreover, different

material science strategies were investigated in attempts to

devise a solution to making the functional life span as long as

possible (unpublished findings).

Unfortunately, when implanted in neural tissue, all artificial

materials have an inconvenient tendency to induce the

formation of scar tissue, which tends to encapsulate the

electrode surface, increases electrode impedances and lowers

signal amplitudes. In present arrays, the greatest doubt with this

vigorous scar tissue formation and reduced biocompatibility

was addressed to PSPI (PI-2771, HD Microsystems) used as an

insulation material. However, the results from study III,

performed according to the ISO 10993-5 standard, revealed that

PSPI is almost as non-cytotoxic as conventional PI and PE (used

as a negative control material). The PSPI films did not adversely

affect the BHK-21 fibroblast cell function, and hence it seems to

be an appropriate material for incorporation into implantable

biomedical devices. However, studies were performed only

with one cell line and with a moderately short culture period.

Therefore, further in vitro and in vivo studies will be needed to

clarify its long-term effects.

Platinum electroplating has been shown to represent an

effective method to produce sponge-like Pt surfaces with much

higher roughness and increased effective surface areas (de Haro,

et al. 2002). Although this technique is a powerful way to

decrease electrode impedance (Gonzales & Rodriguez 1997), and

to prolong the effective life-span of microelectrode (Chen et al.

2009), it has also some serious disadvantages. Electrodeposited

Pt adheres weakly to the electrode surface, resulting in unstable

electrode impedance. Further, electrolytes commonly used in Pt

electroplating contain lead, which is known to be a highly toxic

heavy metal evoking a degree of uncertainty about the

9 - Discussion

129

biocompatibility of electrodeposited Pt surfaces (Schuettler et al.

2005). In this study, the USPLD technique (ColdabTM ) was used

to deposit Pt thin films with a nanotextured surface. Compared

to the sputter-coated smooth Pt surface, a significant reduction

in electrode-electrolyte impedances was achieved. This may be

attributable to the larger effective surface area. The USPLD

technique also conferred several other advantages in electrode

fabrication. Firstly, the USPLD films displayed high adhesion to

different substrates such as PI, SS and Si. Secondly, the samples

containing heat-sensitive materials, such as photoresists, could

be kept normally very close to room temperature, thus avoiding

hardening or flow during deposition. This meant that it was

possible to achieve both high-quality films and easy lift-off.

Thirdly, it was possible to vary deposition parameters easily to

obtain Pt films ranging from smooth (Ra < 1 nm) to relatively

rough (Ra = 100 nm) meaning that optimal results could be

obtained for each purpose. Fourthly, the biocompatibility of Pt

films is guaranteed because high purity solid source materials

are directly transferred from the target to the sample. It is

desirable that the electrode surface should also support

attachment and growth of neurons. It has been reported that

thin Pt films are non-toxic and can support the attachment and

the growth of cortical rat neurons (Thanawala et al. 2007). It has

been also shown that electrode surface morphology has a major

impact on neurocompatibility (Turner et al. 2000). Enhanced

functions of neurons have been demonstrated on biomaterials

structured at the nanoscale (Bayliss et al. 1999, Raffa et al. 2007,

Kelly et al. 2008). Furthermore, it has been reported that

astrocyte adhesion declines whereas the extension of neurons

increases on nanoporous Si surfaces (Moxon et al. 2004). There is

decreased adhesion as well as reduced proliferation and poorer

functional activation of astrocytes on carbon nanofiber materials

(diameter 60 nm) compared to conventional carbon fibres

(diameters 125-200 nm) (McKenzie et al. 2004). These results

provide clear evidence that materials structured at the nanoscale

may have promising neural applications, because they increase

the effective surface area and interact positively with neurons,

9 – Discussion

130

while they simultaneously reduce the functions of astrocytes,

leading to decreased glial scar formation (Kotov et al. 2009).

The vast majority of the tissue contact in the neural interfaces

occurs under the substrate/encapsulation material. Thus, this

material has to fulfill high demands in order to allow the stable,

long-term function of the implant. An optimal insulating

material should be biocompatible, biostable and still possess

appropriate dielectric properties. Although PI was successfully

used in the present arrays as well as in several other neural

interfaces developed by other research groups (Cheung et al.

2007, Stieglitz et al. 2000, Takahashi et al. 2003, Rousche et al.

2001), some degree of uncertainty is related to the relatively

large water uptake, which is typically in the range of 0.5 – 4%

(Stieglitz et al. 2000). In this context, one important advantage

of the USPLD method compared to several other methods is that

it is possible to create dielectric films that do not contain

microparticles or droplets, which would cause pinholes in the

films. This is due to the fact that the laser ablation pulses are so

short that the energy penetrates into only a very shallow (tens of

nanometers) surface layer, which is effectively converted to

plasma and the target surface remains smooth without any

flows or deteriorations. These coatings do not seem to contain

any microdefects which could lead to short circuiting, especially

when they come into contact with body fluids. Alumina and

carbon nitride coatings deposited by the USPLD technique were

demonstrated to be cell-friendly materials at least to fibroblastic

(BHK-21) cells. The ultrasmoothness of these depositions might

provide an additional benefit by avoiding the activation of glial

cells and the formation of scar tissue.

9.3 MICROPATTERNED BIOMATERIAL SURFACES

In studies V and VI, photolithographic micropatterning of

biomaterial surfaces was used for the production of model

surfaces with accurately defined features. These kinds of model

surfaces were demonstrated to be powerful tools for

9 - Discussion

131

investigating the effects of different surface on the adhesion,

spreading and morphology of cells and cellular communities.

Different pattern sizes (5-125 µm), shapes (rectangular or

circular) and surface chemistries (Ta, Ti, Cr, DLC) were utilized

in the creation of engineered surfaces on silicon wafers. The

optimized photolithography process with correct baking,

exposure and development parameters allowed the formation of

precisely defined photoresist features with an undercut edge

profile. The shape of profile was observed to be extremely

important for straightforward lift-off after deposition. The

surface roughness, a parameter widely known to modulate

protein adsorption and cellular responses, could be kept almost

constant with all coatings. The AFM characterization revealed

that variation in roughness between sample types was below 2

nm and the only relevant topographical cues of the samples

were the steps between the patterns and background. By precise

adjustment of the deposition time with two different types of

PVD methods, i.e. magnetron sputtering for metals and filtered

pulsed plasma arc discharge for DLC, it was possible to ensure

that the coating thickness variations were kept as low as

possible (200 ± 20 nm). This indicates that the possible clear-cut

differences observed between different materials and patterns

could not have been due to differences in surface topography

and instead suggests that the differences seen and discussed

below originated from the different chemical compositions and

physical patterns of the materials being studied.

Silicon is the most widely used substrate material in

microchips and MEMS devices. However, biocompatibility

properties of Si are often unsuitable for use in bio-MEMS

devices. In studies V and VI, the cytocompatibility enhancement

of Si was achieved using its partial coating with more

biocompatible materials such as Ti or DLC. This suggests that

even partial coating of Si-based bio-MEMS devices can be used

to enhance their cytocompatibility and to facilitate their

integration with the host during the critical initial phase of

integration. The rapid integration of host cells also reduces the

risk of implant-related infections according to the concept of the

9 – Discussion

132

‚race for the surface‛ (Gristina 1987). In cellular experiments

with both SaOS-2 and hMSC cells, a clear cellular preference for

the biomaterial patterns over background Si was observed. Both

cell types initially tended to adopt the geometrical shape of the

patterns, resulting in ‚engineered‛ square or circle shaped

cellular communities. The results from SaOS-2 cell experiments

with different pattern sizes disclosed that the size and spacing of

the patterns provide further options to guide the cells. The size

of the SaOS-2 cells growing on the same Si background varied

significantly, depending on the type of biomaterial used for

patterning. This is a very interesting observation when one takes

into account the fact that all surfaces were produced using the

same fabrication methods on the same background Si. Another

interesting observation was made in the focal adhesion staining

experiments. Preferential localization of vinculin containing

focal adhesions in SaOS-2 cells was clearly observed on the

biomaterial patterns, but hMSCs did not show any such

preference for the patterns. One could speculate that the

patterns used in the hMSC experiments (75 µm squares) were

too small to allow proper organization of the actin cytoskeleton

and subsequent preferential forming of vinculin-containing

adhesion plaques. Moreover, it has been demonstrated that the

pattern size has an effect on the differentiation of hMSCs

(McBeath et al. 2004) and in that case, relatively large islands are

needed to ensure osteogenesis.

The patterning techniques developed in this thesis provide

an effective way to enhance the biocompatibility of Si-based

microdevices. To allow proper functionality of these bio-MEMS

devices, only the areas which are not under the integrated

electronic components may be coated with biocompatible thin

films. These patterning techniques also provide an effective

means to guide the attachment, size and the shape of the cells as

well as to regulate the differentiation of MSCs into the desired

phenotypes. It is hoped that these aspects will stimulate the

design of even more advanced biomaterials and constructs for

tissue engineering and regenerative medicine.

133

10 Conclusions

In this work, novel flexible epidural microelectrode arrays were

successfully developed and tested to permit recordings of acute

and chronic evoked potentials in an animal model. In addition,

novel materials and surface textures at a micro- and nanoscales

were investigated in vitro aiming to improve the biological and

electrochemical characteristics of neural interfaces. The

knowledge about cell-biomaterial interactions obtained in this

thesis work can be used in the further development of neural

interfaces as well as in other implantable biomedical devices.

On the basis of the present series of studies, the following

detailed conclusions can be drawn:

1. The developed flexible microelectrode arrays possess

adequate mechanical and electrochemical characteristics and

they are capable of acquiring signals with high signal-to-

noise ratios in evoked potential recordings when implanted

onto the surface of the rat cortex.

2. The polyimide-platinum-polyimide sandwich structure, in

which the flexibility of PI and the superior electrochemical

properties of Pt were combined, was observed to be the most

suitable realization of the array. A novel finding was the

confirmation of the biocompatibility of the photosensitive

polyimide used in the developed arrays.

3. Nanostructures on an electrode surface can be used to

increase the effective surface area and to lower electrode

impedance as observed with nanotextured Pt surfaces

created with an ultra-short pulsed laser deposition technique.

4. Thin film dielectrics, alumina and carbon nitride produced

by the USPLD method were proved to be highly cyto-

compatible in in vitro studies with fibroblasts. However,

10 – Conclusions

134

further in vitro and in vivo studies will be needed to clarify

their long-term effects.

5. Micropatterning with biomaterials can be used to enhance

the cytocompatibility of silicon-based bio-MEMS devices.

Not only the chemical composition of the materials, but also

the shape, edges (height) and size of the features used for

surface patterning have a remarkable effect on cell guidance.

6. The morphology of cells cultured on micropatterned

surfaces seemed to be aligned along the edges of the features

e.g. forming round- or square-shaped cell morphologies

having an interesting effect of the differentiation of stem

cells into certain phenotypes. These results may be

particularly useful in applications involving regenerative

medicine and tissue engineering where cells are grown on

biomaterial scaffolds.

135

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Publications of the University of Eastern FinlandDissertations in Forestry and Natural Sciences



isbn 978-�952-�61-�0216-�0

Sami Myllymaa

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

Recent advances in micro- and

nanotechnology offer a great oppor-

tunity to develop intelligent bioma-

terials and the next generation of

implantable devices for diagnostics,

therapeutics, and tissue engineering.

This dissertation is focusing on the

development of novel polymer-based

microelectrode arrays suitable for

use in intracranial electroencepha-

lographic recordings. Moreover,

the performances of novel thin film

materials and their surface modifica-

tions at micro- and nanoscales were

studied with physicochemical and

cellular experiments in order to de-

vise new solutions for further devel-

opment of biomedical microdevices.

dissertatio

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ovel M

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Sami MyllymaaNovel Micro- and

Nano-technological Approaches for Improving

the Performance of Implantable

Biomedical Devices

| 014 | sami myllymaa | novel micro- and nano-technological … · 2010-11-10 · implant/tissue...

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