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DESIGN AND MANUFACTURE OF BASE / FILAMENT

INTERCONNECTS AND FILAMENT PATTERNING OF

PIEZOELECTRIC COCHLEAR IMPLANT.

A Dissertation submitted to the

Division of Research and Advanced Studies

of the University of Cincinnati

in Partial Fulfillment of the Degree of

MASTER OF SCIENCE (MS)

in the Department of Chemical and Materials Engineering

of the College of Engineering

2004

by

Smita Mitra

B.S., Bhavans College, Mumbai, India, 1999

Committee Chair: Dr. Rodney Roseman

ABSTRACT

This dissertation describes the efforts undertaken towards the development of a

successful polymer electrode and interconnect based piezoelectric cochlear implant

hearing aid device. A thorough review of the cochlear environment and processes

resulted in the identification of the requirements of such a device, most critical of which

are high sensitivity, small dimensions and flexibility, impedance matching and

biocompatibility. Based on these requirements and initial experiments, proper electrode

material was identified. These devices are based on flexible polymer piezoelectric films,

used in the bending/flexure mode of deformation. The specific material chosen was

Polyvinylidene fluoride (PVDF), as PVDF is the most sensitive piezoelectric polymer

and high quality commercial films are readily available. Polypyrrole was identified as the

most appropriate electrode and interconnect material in working towards developing an

all-polymer system. The suitability of Polypyrrole (PPY) material for this application was

investigated. For the cochlear implant application, the devices with PPY electrodes and

interconnect showed much higher voltage responses than those from metal electroded

devices.

To measure the piezoelectric sensitivity of these devices to acoustic waves in-air

measurements were carried out. Devices measured display very high sensitivity (several

volts for conversational sounds at close range). Dependence of in-air sensitivity on device

dimensions and interconnect pattering was investigated. It was found that interconnects

connected in parallel showed a very high voltage response as compared to interconnect

connected in series. In work towards the development of interconnect and electrode for

cochlear implant it was seen that properties of PPY exceeded those of metal interconnect

and electrode material.

Conclusions were made followed by recommendations for future work.

ACKNOWLEDGEMENTS

I wish to express my profound gratitude to my advisor Prof. Rodney Roseman, for

his tireless guidance, patience and encouragement during the course of my Master’s

Thesis. I would also like to thank Prof. Relva Buchanan and Prof. Stephen J. Clarson, for

serving on the committee. I further want to acknowledge Prof. Ray Y. Lin for the use of

instruments and facilities in his lab. I would also like to acknowledge all other MSE

faculty, MSE departmental staff Douglas Bowling, Dale Weber, David Simpson my

guides over the years. Niloy Mukherjee, Srinivas Subramaniam, Brian Kendall, Arpit

Dwivedi, Ankur Kant, the University of Cincinnati, the UC College of Engineering, and

the people of Cincinnati.

Last, but not least, I wish to express my gratitude for my parents S.K Mitra, Jaya

Mitra, my husband Siddhartha Sengupta and my sister Soumita Mitra without whose

support and encouragement it would have been impossible to achieve this

accomplishment. I wish to acknowledge the National Science Foundation for supporting

this research work. Needless to say, if not for the help of the above individuals, this work

would have been very difficult indeed. A small paragraph can only do scant justice to the

enormity of the gratitude I feel for these people and many others whose names I have

omitted for brevity.

i

TABLE OF CONTENTS

Table Of Contents i

List Of Figures vi

List Of Tables xii

1 Background And Literature Review 1

1.1 Human Ear 2

1.1.1 The outer ear 3

1.1.2 The middle ear 3

1.1.3 The inner ear 3

1.1.4 Normal working of the human hearing mechanism 7

1.1.5 Sensorineural hearing loss 9

1.1.6 Working methodology of cochlear implants 11

1.1.7 Cochlear Implants 12

1.2 Piezoelectricity and PVDF2 16

1.2.1 Early developments 16

1.2.2 Definitions 16

1.2.3 Theory and basic equations 17

1.2.4 Piezoelectricity and Electrostriction 23

1.2.5 Piezoelectric Polymers 23

1.2.6 Structure of PVDF or PVF2 24

1.2.7 Morphology of PVDF 28

1.2.8 Polarization Process for PVDF 30

1.2.9 Transducer Applications of PVDF 33

1.2.10 Bending Mode Piezoelectricity in Polymers 35

1.2.11 Piezoelectric Cochlear Implant 37

ii

1.3 Device Requirements And Design 41

1.3.1 Device Requirements 41

1.3.2 Design 43

1.4 Polydimethylsiloxanes 47

1.4.1 Introduction 47

1.4.2 Health Effects 48

1.5 Conductive Polymers 49

1.5.1 Electrical Conductivity 49

1.5.2 Conductive Electrodes 50

1.5.3 Conductive Polymers 52

1.5.4 Electrical Properties 53

1.5.5 Doping 55

1.5.6 Applications Of Conductive Polymers 61

1.6 Polypyrrole 63

1.6.1 Introduction 63

1.6.2 Polymerization Mechanism for Polypyrrole 63

1.6.3 Monomer - Oxidant Ratio 65

1.6.4 Rate of The Polymerizing Reaction 70

1.6.5 Dopant 71

1.6.6 Anisotropic Properties In Polypyrrole 74

1.6.7 Double Layer Capacitance 78

1.6.8 Adhesion Properties 78

1.6.9 Self-Assembly Of Conducting Polymers 79

1.6.10 Methodology Of Self-Assembly 80

1.6.11 In-Situ Deposition Of Polypyrrole Films 81

1.6.12 Aging Effect In Polypyrrole Films 82

iii

1.6.13 Bio-Medical Applications 85

1.6.14 Medical Coating Characteristics 88

1.7 Four Point Probe 90

1.7.1 Working methodology 90

1.7.2 Correction Factor 93

1.8 Acoustic Studies 96

1.8.1 Basic Acoustic Phenomenon And Equation 96

1.8.2 Air Measurements: Semi-Quantitative Measurements 101

1.8.3 Underwater measurements: Quantitative 101

1.8.4 Comparison of Air Acoustics with Underwater Acoustics 102

1.8.5 Speed of Sound in Water 102

1.8.6 Acoustic Sensitivity Calculation 104

1.8.7 dB SPL Scale Sound Pressure Scale 105

1.8.8 dB (ref. 1 Vrms/1µPa) Transducer Sensitivity Scale 105

1.9 Electrode-Substrate Contacts 107

1.9.1 Drift of Carriers in Electric Field 107

1.9.2 Schottky Barriers 109

1.9.3 Rectifying Contacts 111

1.9.4 Ohmic Contacts 112

2 Experimental 116

2.1 Experimental Procedures 117

2.1.1 Materials and Substrates 117

2.1.2 Experimental Setup 118

2.1.3 Device Preparation 124

2.2 In-air Acoustics and Errors in Measurements & Voltage Response

Plots

126

iv

3 Results And Discussion 129

3.1 Polymer Substrate Characteristics 130

3.2 Four Probe Test 133

3.3 SEM Data Analysis 135

3.4 In Air Acoustics 139

3.4.1 Semi-Quantitative Acoustic Tests 139

3.4.2 Semi-Quantitative Acoustic Tests Using 28µm PVDF

Film

140

3.4.2.1 Voltage Response of Metal Electroded PVDF

Devices

140


Devices with

144

3.4.2.3 Varying Device Dimensions 153

3.4.2.4 Voltage Response of PPY Electroded PVDF

Devices

164

3.4.2.5 All Metal Device Vs All Polymer Device

Connected in Parallel

166

3.4.3 Semi-Quantitative Acoustic Tests Using 52µm

PVDF Film

166


Devices

166

3.4.3.2 Voltage Response of Polymer Electroded PVDF

Devices

174

3.4.3.3 Comparing metal electroded PVDF devices Vs

Polymer electroded PVDF devices

197

3.4.4 Film thickness 200

3.4.5 Dimensions and Interconnect thickness 201

v

4 Summary 202

5 Conclusion 204

6 Future Work 206

Bibliography 208

vi

LIST OF FIGURES

1.1.1 Schematic illustration of the ear [1 ] 2

1.1.2 (Upper) Cross-sectional view of the cochlea showing the three scalae and

the organ of Corti; (Lower) Close-up of the organ of Corti, showing hair

cells, stereocilia, basilar membrane [3].

5

1.1.3 Electrical ‘standing’ potential of different fluids within the cochlea with

respect to perilymph (0 mV). This is the potential difference between

perilymph fluid and other fluids in the inner ear. In the case of the

potential difference between perilymph and intracellular-fluid, this

potential is also called the resting membrane potential [1].

6

1.1.4 Schematic of human hearing 8

1.1.5 Simplified schematic of Human Ear 9

1.1.6 Schematic showing working mechanism of a Cochlear Implant 12

1.2.1 Schematic showing the piezoelectric effect 17

1.2.2 Equivalent circuit of a piezoelectric near resonance 22

1.2.3 Schematic summary of crystallization and inter-conversions of

polymorphic phases of PVF2[25].

25 1.2.4 Unit cells of (a) phase; (b) phase; and, (c) phase of PVF2. Only the

a-b Unit-cell plane is shown. Arrows indicate dipole directions normal to

molecular axis [4].

27

1.2.5 Spherulites of PVDF crystallized from crystallized from the melt at 1600

C. large spherulites are of the antipolar -phase; small ones belong to the

polar -phase [6].

28

1.2.6 Schematic representation of the structure of polymer spherulites [6] 29

1.2.7 Representation of the processes commonly employed to obtain

piezoelectrically active PVDF films [6].

30 1.2.8 Schematic depiction of the two most common chain conformations of

PVF2 crystals: (a) tg+tg_ (b) tttt. The arrows indicate the projections of

the CF2 dipole moments in the plane containing the carbon backbone [6].

31

vii

1.2.9 Typical Infrared absorption spectrum of PVDF film [29 35

1.3.1 Schematic of a device based on stereocilia. The vertical elements are

electrodedpolymer film. The base is a flexible bio-compatible polymer.

The device is shown placed in the cochlear scala tympani. The

piezoelectric elements face the incoming sound pressure.

46

1.5.1 Conductivity ranges for polymers (doped and undoped), inorganic

materials and Molecular crystals (from Cowie).

53

1.5.2 Insulator-Semiconductor-Metal (Band Diagram) 54

1.5.3 Typical properties of solitons, polarons and bipolarons [47]. 56

1.5.4 Polaron-Bipolaron (Band Diagram) [45] 57

1.5.5 Propagation of a polaron through a conjugated polymer chain by shifting

of double bonds [45].

58

1.5.6 Schematic showing Conductivity phenomenon in Polypyrrole [45] 59

1.5.7 Resonance forms (aromatic and quinoid) of PPY are not equivalent [45] 60

1.5.8 Technological Applications of Conductive Polymers [47] 62

1.6.1 Schematic showing chemical synthesis of Polypyrrole 64

1.6.2 Effect of Immersion time on normalized weight uptake and av. Sheet

resistance [70].

67

1.6.3 Effect of repeated PPY deposition [70]. 68

1.6.4 Conductivity and Activation energy of Polypyrrole along the Surface [81] 74

1.6.5 Conductivity and Activation energy of Polypyrrole in the thickness

direction [81]

74

1.6.6 In-plane Impedance of doped PANI film (a) Real and imaginary

impedance (b) Cole-Cole Plot [89]

76

1.6.7 Through-plane Impedance of doped PANI film (a) Real and imaginary


77

1.6.8 Effect of aging temperature (in air) on the conductivity of Glass/PPY (1-

step process) film at (1) 60°C; (2) 80°C; (3) 100°C; (4) 120°C; (1)

140°C; [106]

84

viii

1.6.9 Effect of aging temperature (in air) on the conductivity of Glass/PPY

(2-step process) film at (1) 60°C; (2) 80°C; (3) 100°C; (4) 120°C; (1)

140°C; [106]....

84

1.7.1. 2-point probe showing the contact resistance Rc the spreading

resistance Rsp and the semiconductor resistance Rs. [116]

90

1.7.2 A collinear Four point Probe. [116] 92

1.9.1 Drift of electrons and holes in a semiconductor bar [130] 108

1.9.2 A Schottky barrier formed by contacting a semiconductor with a metal

having a larger work function [133]

110

1.9.3 Effects of forward and reverse bias on the ideal metal-semiconductor

junction. (a) forward bias; (b) reverse bias; (c) typical I-V

characteristics [130]

111

1.9.4 Ohmic metal-semiconductor contacts: (a) ∅m < ∅S for an n-type

semiconductor, and (b) the equilibrium band diagram for the junction;

(c) ∅m > ∅S for a p-type semiconductor, and (d) the junction at

equilibrium [130 ]

114

1.9.5 Difference between the rectifying and Ohmic contacts [133] 115

1.9.6 I-V characteristics of rectifying and Ohmic contacts [133] 115

2.1 Process Schematic of Piezoelectric sensor device [124 120

2.2 Schematic showing in-situ Polymer deposition 123

2.3 Schematic showing Acoustic test in air 124

2.4 Picture of Water Tank and other instruments 125

2.5 Voltage response plot of a single element device 127

2.6 Voltage response plot of a 2 element device. 128

3.1.1 DSC of PDMS 132

3.3.1 Diluted silver paste on glass substrate 135

3.3.2 High purity silver paste on glass substrate 135

3.3.3 High Purity silver paste on Lecoset substrate – (at the junction) 136

3.3.4 High Purity silver paste on Lecoset substrate – (at the junction) 136

3.3.5 PPY on Lecoset substrate – (at the junction) 136

ix

3.3.6 PPY on Lecoset substrate – (Surface view) 136

3.3.7 PPY on Lecoset substrate 137

3.3.8 Globular morphology of PPY film on un-stretched PVDF film 138

3.4.1 Comparison of in-air test on std. Hydrophone and device prototype 139

3.4.2 (a)Multi element device connected in series and (b) Multi element

device connected in parallel

141

3.4.3 Voltage response of a single element device (a) Length of PVDF =

05mm and (b) Length of PVDF = 20mm

143

3.4.4 Voltage response of a 4 element device connected in parallel (a)

Length of PVDF = 20mm and (b) Length of PVDF = 05mm

146

3.4.5 Voltage response of a 4 element device connected in series (a) Length

of PVDF = 20mm and (b) Length of PVDF = 05mm

148

3.4.6 Voltage response of a 5 element device connected in parallel,

dimensions = 20 mm × 2 mm (a) Length of PVDF = 20mm and (b)

Length of PVDF = 05mm

151

3.4.7 Voltage response of a 5 element device connected in parallel,

dimensions = 20 mm × 1 mm (a) Length of PVDF = 20mm and (b)


152

3.4.8 Voltage response of a single element device (a) Length PPY

interconnect = 3 cms and (b) Length PPY interconnect = 1

155


Length PPY interconnect = 3 cms and (b) No interconnect

157

3.4.10 Voltage response of a 2 element device connected in series (a) Length

PPY interconnect = 3 cms and (b) No interconnect

158

3.4.11 4 element device connected in parallel with PPY electrodes and

interconnect.

159


Length PPY interconnect = 3 cms and (b) No interconnect

160

3.4.13 5 element device connected in parallel with PPY electrodes and

interconnect.

161

x

3.4.14 Voltage response of a 5 element device connected in parallel (a) Length


162

3.4.15 Voltage response of 8-element device connected in parallel [135] 168


dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b)

dimensions = 20 mm × 2 mm, Length of PVDF = 20mm

170




171




172

3.4.19 Voltage response of a single element device connected in parallel (a)



175




177




179

3.4.22 Prototype of a 4 element device connected in parallel (a)Design1 and

(b)Design2 3.4.23 Voltage response of a 4 element device connected

in parallel (Design 1)(a) dimensions = 20 mm × 1 mm, Length of

PVDF = 20mm and (b) dimensions = 20 mm × 2 mm, Length of

PVDF = 20mm

181

3.4.23 Voltage response of a 4 element device connected in parallel (Design

1)(a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b)


182

xi


2) (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b)


184




186




188

3.4.27 Prototype of a 6 element device connected in parallel (a) Design1 and

(b)Design2

189




190




192

3.4.30 Voltage Response Vs Length of interconnect 163

3.4.31 Future Experimental Setup 194

3.4.32 Voltage response of a single element device (a) Dimensions 2mm ×

0.5mm(b) Dimensions 10mm × 0.5mm

196

3.4.33 Comparison of metal VS polymer electroded device 198

xii

LIST OF TABLES

1.2.1 Physical Constants of some Piezoelectric Polymers 32

1.2.2 Acoustic Properties of Various materials [25] 34

1.2.3 Comparison of piezoelectric materials [29] 34

1.2.4 Bending Piezoelectric Constants [32] 37

1.8.1 Acoustic impedance of some common media [117] 100

1.8.2 Some typical sound pressure levels in the dB SPL scale [117] 101

1.8.3 Parameters entering formula for temperature dependence of sound speed

in pure water [118,119]

104

2.1 Voltage response of an all polymer 2 element device connected in parallel 126

2.2 Ten voltage response iterations of an all polymer 2 element device

connected in parallel

127 3.1.1 Solubility test of PDMS 130

3.2.1 Four probe test 133

3.4.1 Voltage response of metal electrode-PVDF devices Vs Number of

elements.

141

3.4.2 Voltage response of length metal electrode-PVDF devices Vs Number of

elements

145

3.4.3 (a) Lists the voltage response of a 5 element device 149

3.4.3 (b) Voltage response of 5 element metal electrode-PVDF devices Vs

changing dimensions

150

3.4.4 Voltage response of a single filament device 154

3.4.5 Voltage response of a 2 element device with PPY electrode and

interconnects

156

3.4.6 Voltage response of a 4 element device connected in parallel 159

3.4.7 Voltage response of a 5 element device connected in parallel 161

3.4.8 Comparison of Series and Parallel Connection Metal electroded PVDF

elements (20mm × 2mm) [135]

167

xiii

3.4.9 Comparison of Series and Parallel Connection PPY electroded PVDF


168

3.4.10 Comparison of Series and Parallel Connection PPY electroded PVDF


169

3.4.11 Voltage response of a metal and polymer single and multi element device 164

3.4.12 Voltage response of a multi-element device connected in parallel 169

3.4.13 Voltage response of an all polymer single element device connected in

parallel

174

3.4.14 Voltage response of an all polymer 2 element device connected in parallel 176

3.4.15 Voltage response of an all polymer 3 element device connected in parallel 178

3.4.16 Voltage response of an all polymer 4 element device connected in parallel

(Design 1)

181


(Design 2)

183


(Design 1)

185


(Design 2)

187


(Design 1)

189


(Design 2)

191

3.4.22 Voltage response of a single-element device Dimensions: 10mm × 0.5mm 195

3.4.23 Voltage response of a metal electroded device 197

3.4.23 (a) Voltage response of a polymer electroded device 198

3.4.24 Effect of peizo film thickness on voltage response (metal device) 200

3.4.25 Effect of peizo film thickness on voltage response (PPY device) 200

3.4.26 Effect of peizo film dimension on voltage response 201

1

Chapter 1

Background and Literature Review

2

1.1 Human Ear

The human ear and the brain are the hardware of the auditory mechanism.

Commonly, the ear is viewed as one of the conspicuous flap like appendages found on

the side of the head. Technically, the ear refers to the entire peripheral auditory system.

This involves everything from the readily visible outer ear to the intricate structures of

the inner ear located deep in the skull and the nerves connecting it to the brain [2]. A

schematic illustration of the ear is shown in Figure1.1.1 [1].

The peripheral auditory system consists of three major subdivisions referred to

classically as the external, middle and internal ear. In more common terms they are

known as the outer, middle and inner ear, respectively [2].

Figure 1.1.1: Schematic illustration of the ear [1]

3

1.1.1 The outer ear

The outer ear is comprised of those structures that gather air-borne vibrations and

funnel them down channels in the side of the head and is the essential first link in the

conduction of sound [1].

1.1.2 The middle ear

The middle ear in man is an acoustic and mechanical device, which exerts a

profound influence on the auditory stimuli arriving at the cochlea. In the most general

sense, the middle ear acts as a band-pass filter that dictates the shape of the minimum

audible pressure function and provides for an emphasis of those frequencies in the middle

range that are important for speech intelligibility (≈ 100 Hz - 4 kHz) [1]. The middle ear

has five major parts:

1. The cavity and mastoid air spaces;

2. The tympanic membrane (eardrum);

3. The auditory ossicles (bones), malleus, incus and stapes;

4. The middle ear muscles; and,

5. The Eustachian tube.

All of the structures and cavities comprising the middle ear influence the

transmission of sound energy from the air to the cochlea of the inner ear.

1.1.3 The inner ear

The human inner ear is divided into three portions, each having a specialized

sensory apparatus and a separate function. The oldest of the divisions lies within the

vestibule, or entryway, which contain sensory receptors. The semicircular canals, which

lie in orthogonal planes, comprise the second division of the inner ear and both provide

4

significant input to the central nervous system and share responsibility for our stability

while stationary or moving. The auditory portion of the inner ear, the cochlea, is a coiled

extension located on the other side of the vestibule. The cochlea (Figure 1.1.2) is

separated into three channels, or scalae, by the Reissner’s membrane and the basilar

membrane, both of which form the boundaries of the cochlear membranous labyrinth [3].

The membranous labyrinth that is suspended between the other two channels is the scala

media. Reissner’s membrane separates the scala media from the scala vestibuli and the

basilar membrane forms the boundary between the scala media and the scala tympani.

5

Figure 1.1.2: (Upper) Cross-sectional view of the cochlea showing the three scalae and

the organ of Corti; (Lower) Close-up of the organ of Corti, showing hair cells, stereocilia, basilar membrane [3].

The scala tympani and scala vestibuli are filled with perilymph fluid (~ 0.9 wt. %

NaCl). Fluid within the scala media is relatively low in sodium content and high in

potassium and is known as endolymph. The scala vestibuli communicates with the

vestibule at the basal end of the cochlea. The basal portion of the scala tympani

6

terminates at the round window. The apical end of the scala media almost, but not quite,

seals off the apical bony labyrinth, and the small opening that remains, the helicotrema,

allows for the communication between the scala vestibuli and the scala tympani. The

cochlea coils around a hollow core, the modiolus, through which the auditory nerve and

the cochlear blood supply pass [1, 3]. The aggregate of supporting cells, sensory cells,

membranous and other structures within the scala media comprises the auditory receptor

organ known as the organ of Corti. The organ of Corti is divided into “inner" and “outer"

regions. Receptor cells in these regions are known as inner and outer hair cells [3]. The

perilymph, endolymph and hair cells are not electrically neutral, and each has a ‘standing

potential’, quantified with respect to the perilymph whose potential is considered as zero.

Figure 1.3 shows the standing potential in each place in the cochlea [1].

Figure 1.1.3: Electrical ‘standing’ potential of different fluids within the cochlea with

respect to perilymph (0 mV). This is the potential difference between perilymph fluid and other fluids in the inner ear. In the case of the potential difference between perilymph and

intracellular-fluid, this potential is also called the resting membrane potential [1].

7

The hair cells project hair like structures called stereocilia that are arranged in

rows with graded heights, so that the tallest are on the side away from the modiolus.

Towards the base of each hair cell, as shown in Figure 1.1.2, there are synaptic structures,

where the hair cells make contact with the afferent auditory nerve fibers. Large synaptic

terminals also exist on the outer hair cells. The central nervous system is able to influence

the state of the hair cells by means of this pathway. There are about 3000 inner hair cells,

arranged in a single row from base to apex of the cochlea. The outer hair cells, which

number 12,000 to 15,000, are arranged in 3 parallel rows [3]. The stereocilia are stiff,

rigid structures, composed of actin filaments, which run parallel to the long axis of the

stereocilium. Each of these actin filaments has certain sites along its length at which it

can bond to its neighbors. The stereocilia do not bend, but rather pivot at their base, until

the force becomes so great that they crack. The individual stereocilia are not separate

structures at the base of the hair cell, but are joined together by a large number of lateral

and upwards-running cross-links.

1.1.4 Normal working of the human hearing mechanism

The hearing system in mammals is an extraordinary example of nature’s

engineering. Sound is the mechanical vibration of gaseous, liquid or solid elastic medium

through which energy is transferred away from the source by progressive sound waves.

The human ear can respond to minute pressure variations in the air if they are in the

audible frequency range, roughly 20 Hz - 20 kHz (Figure 1.1.4).

8

Figure 1.1.4 Schematic of human hearing

The sensitive, sophisticated inner ear organ, Cochlea, performs the most important

function of converting the incoming mechanical energy into electrical signals in the

auditory nerve fibers terminating within it. Figure 1.1.5 shows a simplified diagram of the

human ear consisting of the outer, middle and inner ear.

Sound arrives at the outer ear in the form of a mechanical vibration, passes

through the outer ear canal experiencing the various amplifying mechanisms of the outer

ear, and impinges on the tympanic membrane (eardrum). The mechanical vibration is

transmitted from the tympanic membrane to the oval window of the cochlea by means of

the three middle ear ossicles. An increase in pressure is obtained as a result of the middle

ear amplification mechanisms. A mechanical traveling wave is introduced inside the

cochlear fluid as a result of the pressure differential between the oval window and the

round window. This wave travels with increasing amplitude and peaks at a certain

distance from the oval window, high frequencies peaking at short distance (first few

millimeters) and lower frequencies peaking at larger distances along the cochlea. The

upward force experienced by the basilar membrane due to the peaking wave causes

deflection of the basilar membrane in the direction of the Reissner’s membrane. This

movement results in the lateral movement of the cuticular plate (from which the

9

stereocilia arise) with respect to the tectorial membrane, thus causing shear displacement

of the stereocilia in the direction away from the modiolus. The displacement of the

stereocilia cause a change in the standing potential of the hair cell, resulting in generation

of action potentials in the synapse nerve fibers and sending a signal to the brain through

the eighth (auditory) nerve [4].

Figure 1.1.5 Simplified schematic of Human Ear

1.1.5 Sensorineural hearing loss

There are different kinds of hearing disorders in humans. Among all of them

sensorineural hearing disorder deals with hearing loss relating to the cochlea and is a

condition affecting a large percentage of severely and profoundly hearing impaired.

Sensorineural hearing loss results when the hair cell structures are unable to perform the

energy transduction task, often because they are either absent, destroyed or degraded

[7,8].

10

Numerous things can damage the complex and delicate water world of the

cochlea; however, the primary form of damage is the destruction of the tiny sensory hair

cells. When the cells die, contact with the hearing nerve fibers is broken and our

perception of particular frequencies is lost. Since the sensory hair cells of the auditory

system are not naturally replaced or repaired, the damage is permanent. Without a

limiting mechanism to prevent dangerous amounts of acoustic energy from being

transmitted to the inner ear, over-exposure to high intensity sound is the leading cause of

damage to the cochlea’s microscopic hair cells. Various infections can cause damage to

the nerves in the cochlea, such as spinal meningitis or syphilis. The most common viruses

that cause hearing loss are measles (rubella) and mumps. Viral infections in a pregnant

mother can also affect the fetus. Damage to the cochlea of an unborn child also can take

place when there is an Rh incompatibility with the mother or the mother has a fever.

Certain drugs have been identified as toxic to the cochlea, among them: quinine, large

doses of aspirin, certain drugs of the mycin family, aminoglycosides and certain chemo-

therapy agents [5].

In patients suffering from this disorder, the mechanical functions of the cochlea

are fully functional, with a large percentage of the nerve fibers also surviving

(progressive degeneration of the fibers could occur with age). Only the transduction link

between the mechanical and electrical energy is missing, rendering the patient profoundly

deaf. In these patients, traditional amplification with hearing aids or middle ear

reconstructive surgery cannot enhance hearing potential, and this type of deafness is

permanent and irreversible [7,8].

11

1.1.6 Cochlear Implants

Cochlear implants are a class of devices developed over the past ≈ 40 years in an

effort to counteract sensorineural hearing loss [9-11]. These devices, of which several

variations have been developed and are in use, are based on the direct stimulation of the

surviving nerve fibers in the cochlea by electrical charge. Sound is captured by an

external microphone worn by the patient, the signal(s) is subjected to conditioning

operations (filtering and compression) externally [9, 12], and the output is introduced to

electrode(s) implanted in the scala tympani of the cochlea. The charge on the electrodes

stimulates the nerve fibers, thus producing a sensation of hearing [9-11]. The implant is

surgically placed under the skin behind the ear.

An implant has four basic parts:

1. A microphone, which picks up sound from the environment.

2. A speech processor, which selects and arranges sounds, picked up by the

microphone.

3. A transmitter and receiver/stimulator, which receive signals from the speech

processor and convert them into electric impulses.

4. Electrodes, which collect the impulses from the stimulator and send them to

the brain.

An implant does not restore or create normal hearing. Instead, under the

appropriate conditions, it can give a deaf person a useful auditory understanding of the

environment and help him or her to understand speech. A cochlear implant is very

different from a hearing aid. Hearing aids amplify sound. Cochlear implants compensate

for damaged or non-working parts of the inner ear. When hearing is functioning

12

normally, complicated parts of the inner ear convert sound waves in the air into electrical

impulses. These impulses are then sent to the brain, where a hearing person recognizes

them as sound. A cochlear implant works in a similar manner. It electronically finds

useful sounds and then sends them to the brain. Hearing through an implant may sound

different from normal hearing, but it allows many people to communicate fully with oral

communication in person and over the phone.

1.1.7 Working methodology of cochlear implants

Figure 1.1.6 Schematic showing working mechanism of a Cochlear Implant

1. Sounds are picked up by a microphone and turned into an electrical signal.

2. This signal goes to the speech processor where it is "coded" (turned into a special

pattern of electrical pulses).

3. These pulses are sent to the coil and are then transmitted across the intact skin (by

radio waves) to the implant.

4. The implant sends a pattern of electrical pulses to the electrodes in the cochlea.

13

5. The auditory nerve picks up these tiny electrical pulses and sends them to the

brain.

6. The brain recognizes these signals as sound.

The cochlear implant is based on the idea that there are enough auditory nerve

fibers left for stimulation in the vicinity of the electrodes. Once the nerve fibers are

stimulated, they fire and propagate neural impulses to the brain. The brain interprets them

as sounds. The perceived loudness of the sound may depend on the number of nerve

fibers activated and their rates of firing. If a large number of nerve fibers are activated,

then the sound is perceived as loud. Likewise, if a small number of nerve fibers are

activated, then the sound is perceived as soft. The number of fibers activated is a function

of the amplitude of the stimulus current. Varying the amplitude of the stimulus current

can therefore control the loudness of the sound. The pitch on the other hand is related to

the place in the cochlea that is being stimulated. Low pitch sensations are elicited when

electrodes near the apex are stimulated, while high pitch sensations are elicited when

electrodes near the base are stimulated. In summary, the implant can effectively transmit

information to the brain about the loudness of the sound, which is a function of the

amplitude of the stimulus current, and the pitch, which is a function of the place in the

cochlea being stimulated. Two basic types of cochlear implants have been developed

when classified according to the nature of the signal processing electronics. Single

channel devices use only one channel for signal processing. The output signal may be

distributed in the vicinity of the auditory nerve (intra/extra-cochlear) by means of a single

electrode or by a multi-electrode array. The multi-channel devices break the input signal

from the microphone into several channels, each channel usually processing a specific

14

range of frequencies. This is done to enable place coding of frequencies in the cochlea.

Cochlear implants are considered as viable, effective and safe devices that have a high

success rate.

Over the past 40 - 45 years, work of implant investigators throughout the world

has shown that the implant provides, among other things, an awareness of environmental

sounds, improved speech reading, and possibly some minimal speech discrimination [13].

However, in the words of Dr. W .F. House [14], “The problem that plagues the implant’s

future is the same one that has plagued it for the past 20 years: namely that no one knows

how the implant works. We do not know that if it stimulates eighth nerve dendrites in the

basilar membrane or if it stimulates the spiral nucleus cells within the modiolus. We do

not know what quantity of neural tissue is necessary for a functioning cochlear implant,

or indeed if the number of viable eighth nerve fibers is related to the great variability of

response in our cochlear implant patients. We do not know how the electric currents

introduced into the inner ear flow through the structures of the cochlea, or how electron

field densities are built up and dissipated within the cochlea." These are some of the basic

problems that impede the further development of cochlear implants. On the more

practical side, there are a few other disadvantages of a cochlear implant:

1. They are bulky (consist an elaborate external and internal system).

2. They can be a source of discomfort to users and can even cause internal infections in

some.

3. Any malfunction in the internal electronic components necessitates surgery for

removal and repair/replacement.

4. They are costly.

15

The development of these devices has raised some fundamental questions about

the mechanism of their functioning and the mechanisms operating in the healthy cochlea.

The nature of the partially restored hearing experienced by the patient is significantly

different from normal hearing. The patient is generally able to at least be aware of

environmental sounds and a few patients are able to understand the verbal language due

to a combination of their re-learning skills, sound pattern recognition and high quality

device implant procedures [14, 15]. Moreover, single-electrode systems sometimes are as

effective as sophisticate multi-channel, multi-electrode cochlear implants. Reasons

behind this phenomenon are not well known and it is obvious that a better understanding

of the whole phenomenon would help better the engineering of the device implants [14].

16

1.2 Piezoelectricity and PVDF2

1.2.1 Definitions

Piezoelectric materials posses a unique property of generation of electric charge

when mechanical energy was applied to them. In other words they are transducers which

are capable of converting mechanical energy into electrical energy [16]. Cady [16]

defines piezoelectricity as “electric polarization produced by mechanical strain in

crystals belonging to certain classes, the polarization being proportional to strain and

changing sign with it”. The converse effect is also true; a piezoelectric material

experiences mechanical strain when an electric field is imposed across its electrodes.

1.2.2 Early developments

Piezoelectric effect was discovered by the Curie Brothers (Pierre & Jacques) in

1880. It was said that some crystals when compressed in particular directions show

positive and negative charges on certain portions of their surfaces, the charges being

proportional to pressure and disappearing when the pressure is withdrawn [16]. The term

Piezoelectricity (“piezo” in Greek means “to press”) was proposed by W. Hankel in

1881. The Curies observed this effect in crystals of Zinc Blende, Sodium Chlorate,

Boracite, Tourmaline, Quartz, Calamine, Topaz, Tartaric acid, cane sugar and Rochelle

salt. The inverse effect, which had been predicted theoretically by Lippmann, was also

verified [16-17].

17

Figure 1.2.1 Schematic showing the piezoelectric effect

The theory of Piezoelectricity is based on the thermodynamic principles stated by

Lord Kelvin, who suggested a molecular theory and produced a mechanical model of

piezoelectricity. The formulation was studied more by F. Pockels and P. Duhem and most

fully and rigorously by Woldemar Voigt in 1894. By combining the elements of

symmetry of elastic tensors and of electric vectors with the geometrical symmetry

elements of crystals, it was made clear in which of the 32 crystal classes the piezoelectric

effect may exist, and for each class it was shown which of the 18 possible piezoelectric

coefficients may have finite values. Max Born, who in his general theory of lattice

dynamics included a consideration of dielectric, pyroelectric, and piezoelectric effect

[18], performed the first rigorous treatment of the atomic theory of piezoelectricity. In

1920 he published, along with E. Bormann, the first theoretical calculation of the

piezoelectric constant in Zinc Blende.

1.2.3 Theory and basic equations

Piezoelectricity, as defined in a strict crystallographic sense, is a fundamental

property of crystal classes lacking a center of symmetry. When a non-centrosymmetric

18

unit cell is subjected to a uniform stress, a separation of the positive and negative charge

centers occurs within the unit cell, resulting in the formation of a dipole (or a change in

the pre-existing dipole moment in the case of a pyroelectric piezoelectric). Such local

events result in a change in the surface charge density of an electroded bulk sample,

which manifests itself as an open-circuit voltage or a closed-circuit current across the

electrodes. Mathematically, linear piezoelectricity can be quantified by means of four

different constants/coefficients, the most commonly used of which is the piezoelectric

stress coefficient (d), represented in tensor notation as:

Where D is dielectric displacement, is stress, e is strain and E is the electric

field. Stress and strain are rank two tensors, and dielectric displacement and electric field

are rank one tensors (vectors). Therefore, the piezoelectric constants are rank three

tensors, having 27 possible components [21]. However, dijk is symmetrical in j and k, thus

eliminating 9 of the 27 components, and leaving only 18 independent components. The j

and k subscripts are also replaced by a single digit, by adopting the following standards

[21]:

This enables a 2 subscript (instead of three) notation, and is called the matrix or

reduced notation. The piezoelectric compliance matrix of a crystal can then be written in

reduced notation as [21]:

19

By convention, the first character in the suffix of all piezoelectric parameters

represents the direction of the electrical parameter (E, D or P) and the second character

represents the direction of the mechanical parameter ( or e) [19-21]. The point group

symmetry of the crystal determines which of the 18 independent components in Equation

1.2.2 are non-zero.

Three more piezoelectric constants are defined. The piezoelectric stress constant

(g) is related to the d coefficient by the dielectric constant (r):

The piezoelectric strain constant (h) is defined as:

The negative signs arise in the case of the g and h coefficients because of

compressive stress, which is the stress more commonly used in piezoelectric

mathematical treatments/devices, and is thus, represented with a negative sign by

convention. The piezoelectric strain coefficient (e) is given as:

The denominator represents the change in strain. Although regrettable, from

conventional notation, the same symbol for strain and the piezoelectric strain coefficient

is used. The four piezoelectric constants are not independent of each other. If one is

known, the others can be calculated. A high d constant is required in applications where a

high stress response to an applied field is required, examples include oscillators and

20

ultrasonic generators, and in sensory applications where a high charge sensitivity is

required. A high g coefficient is the key parameter in sensory applications where high

voltage sensitivity is required, examples being microphones and igniters. The d and g

coefficients are the most important. They are also the easiest to measure. The

piezoelectric effect is thus, usually represented in the following manner:

The first term in the equation describes the contribution to dielectric displacement

(D) due to an applied field. The second term describes the contribution to D due to an

applied stress (piezoelectric effect).

Though the traditional crystallographic description of piezoelectricity implies that

the effect arises out of changes in dipole moments when a crystal is strained (primary

piezoelectricity), this is not the only manner in which a material can be piezoelectric. The

effect can also arise from dimensional changes when a material is stressed. This is called

the secondary piezoelectric effect, and is expected to play a greater role in materials

displaying relatively higher compressibility’s, e.g., polymers. In such materials, applied

stress causes a relatively large change in sample dimensions. As a result, the separation

distance of dipoles from the electrodes can change appreciably, thus resulting in a change

in the Charge State of the electrode [16]. An extension of this concept will lead to the

realization that any embedded/impurity charge present in a material will also have a

similar effect on the electrodes. This is an important fact, as it means that

embedded/impurity/space charges (usually present in polymers) can make a significant

contribution to the total piezoelectric activity. Another mechanism of piezoelectricity can

arise out of strain gradients, if present. This is called the second order piezoelectric effect.

21

The second order piezoelectric constant is a fourth rank tensor (first order constant is

third rank tensor), and is thus not represented by the piezoelectric compliance matrix of

Equation 1.2.3, and the four constants discussed above.

Coupling Coefficient

The piezoelectric coupling coefficient/factor (k) is a measure of the capability of a

piezoelectric to convert mechanical energy to electrical energy and vice-versa [19,22 and

23]. It is a convenient and direct measurement of the overall electromechanical effect. It

is defined as:

Since the conversion of mechanical energy into electrical energy and vice-versa is

always incomplete, k is always < 1 [23]. The coupling coefficient can be measured by

determining the frequency difference between the conditions of maximum impedance (fn)

and minimum impedance (fm) of an electroded sample. The effective coupling coefficient

is given by the following equation:

The exact coupling coefficient depends upon the geometry of the sample and the

mode of deformation. These are calculated using specific mathematical formulae [19], or

can be determined from pre-existing master curves of keff versus the specific k of interest

[24]. In the electrical equivalent circuit of Figure 1.2.2, two other conditions, known as

series resonance (fr) and parallel anti-resonance (fa), occur as the frequency of the applied

voltage is increased [23]. Series resonance occurs when the reactance of C1 becomes

22

equal and opposite to the reactance of L1. Anti-resonance is the condition when the total

impedance of the mechanical arm of the circuit becomes equal and opposite to the

impedance of Co and equal currents starts flowing in both the arms. These conditions

correspond to dE = 0 and dD = 0. In an ideal material, with no dissipation losses, fm = fr

and fn = fa. In reality, they vary from each other because of the finite amount of

dissipation (loss) present [23].

Figure 1.2.2: Equivalent circuit of a piezoelectric near resonance

Application

Piezoelectric materials are used in transducers, e.g., phonograph cartridges,

microphones, and strain gauges, which produce an electrical output from a mechanical

23

input, and in earphones and ultrasonic radiators, which produce a mechanical output from

an electrical input. Piezoelectric solids typically resonate within narrowly defined

frequency ranges; when suitably mounted they can be used in electric circuits as

components of highly selective filters or as frequency-control devices for very stable

oscillators.

1.2.4 Piezoelectricity and Electrostriction

Electrostriction is the mechanical deformation produced in a dielectric upon

application of an electric field. This phenomenon is essentially different from

piezoelectricity. The distinction lies in the fact that the deformation due to electrostriction

is proportional to the square of the applied electric field and is, therefore, independent of

the direction of the field. Piezoelectric strain on the other hand, is directly proportional to

the applied field and is, hence, directional (i.e., sign reverses with reversal of field) [16].

Electrostriction is an universal property exhibited by all dielectrics, whereas, non-

centrosymmetry of structure is an essential requirement for the existence of

piezoelectricity.

1.2.5 Piezoelectric Polymers

Introduction

Electroactive materials have been the primary source of actuation for Smart

Devices. Numerous applications have been reported, including ultrasonic and undersea

sonic techniques, robotics, vibration isolation and manipulation. The basic property of the

materials used for these applications is the strain response induced by an electric field,

and vice-versa. So far, most of the electroactive materials used for these applications are

electroactive ceramics. However, recently it has been found that polymers, such as

24

polyurethane and Polyvinylidene fluoride (PVDF) polymer, can exhibit very large

electrostrictive and piezoelectric effects [24]. The PVDF has been shown to exhibit

relatively large elastic modulus. The PVDF exhibits the highest elastic energy of all

electroactive materials, even considering ceramics and single crystal inorganic materials.

In certain environments where acoustic impedance matching is limited by size

constraints and conducting liquid medium, the advantage of polymers over ceramics is

their flexibility, low acoustic impedance, and high sensitivity. However, in order to

utilize these useful properties, the electrode material is an important issue, since the

conventionally used metal electrodes, such as Au, Ag and Al, have high acoustic

impedance and also impose mechanical clamping on the soft polymer which can

significantly reduce the electromechanical efficiency of the transducer [24].

In 1969 Kawai found out that the semi crystalline PVF2 or PVDF (Polyvinylidene

fluoride) polymer becomes strongly ferroelectric after having been subjected to the

effects of both mechanical stretching and the application of an electric field.

1.2.6 Structure of PVDF or PVF2

PVDF is a semi-crystalline material with a crystallinity of about 50%, of which

the crystalline form may exist in at least four polymorphs: -, -, - and polar - or -

phases [25]. Figure 1.2.3 shows the conversion of the various polymorphs:

25

Figure 1.2.3 Schematic summary of crystallization and inter-conversions of polymorphic

phases of PVF2[25].

26

- Phase

This is the most common form of PVDF and is normally obtained by

crystallization from the melt. The conformation of this phase is a slightly distorted

TG+TG- with a unit cell that is non-polar due to the anti parallel packing of the two

chains contained in the cell [25].

- Phase

This is the most important polymorph of PVDF because this phase is

predominantly responsible for the piezoelectric and pyroelectric properties of the

polymer. It has an all T configuration and a strong dipole moment normal to the chain

direction. It is formed by the mechanical deformation of the melt-crystallized - phase

films and the formation is aided by head-to-head (-CF2-CF2-) and tail to tail

(-CH2-CH2-) defects, which help to reduce intramolecular strain [25].

δ - Phase

This is a polar form and is produced by subjecting the -phase to a high electric

field, thus producing an inversion of the dipole moments in alternated chains. The unit-

cell dimensions and the configuration are the same as the -phase [25].

γ - Phase

This phase is formed by solution crystallization and readily transforms into -

phase on mechanical distortion. A number of arrangements have been suggested for this

polymorph, and the T3GT3G configuration is generally accepted. It is thus seen as an

intramolecular mix of both and forms [25].

27

Figure 1.2.4 Unit cells of (a) phase; (b) phase; and, (c) phase of PVF2. Only the a-b Unit-cell plane is shown. Arrows indicate dipole directions normal to molecular axis [4].

28

1.2.7 Morphology of PVDF

Polymer crystals are extremely small and when grown from the melt, are arranged

into essentially spherically symmetric poly-crystalline aggregates that have no net

polarization. These aggregates are called spherulites and result from nucleation of

primary crystals within the melt, followed by radial growth outward from these nuclei in

spherical envelopes [26].

Figure 1.2.5 Spherulites of PVDF crystallized from crystallized from the melt at 1600 C. large spherulites are of the antipolar -phase; small ones belong to the polar -phase [6].

29

Figure 1.2.6 Schematic representation of the structure of polymer spherulites [6]

This microstructure appears as radial fibers which are in fact stacks of very thin,

platelet-like crystals, called lamellae (about 10 nm thick and several micrometers in

lateral dimensions). These lamellae consist of macromolecular segments that are packed

crystallo-graphically, while the intervening amorphous regions contain chain segments in

disordered conformations. This two-phase structure of the solid state is typical of

crystallizable polymers. In PVDF, crystalline lamellae represent about 50% of the total

mass, the other half being amorphous [6].

30

1.2.8 Polarization Process for PVDF

The most common method for obtaining macroscopically polar films in PVDF

involves first mechanical extension and then electrical poling as described by the

following figure:

Figure 1.2.7 Representation of the processes commonly employed to obtain

piezoelectrically active PVDF films [6].

Mechanical drawing causes a breakdown of the original spherulitic structure into

an array of crystallites whole molecules are oriented in the direction of the force. At high

temperature (~150oC) deformations, the original TG+TG- chains get free to slide past

each other without altering their conformation, so that the resulting structure still remains

the - phase. However, a deformation at low temperatures (~ 90oC) results in a

molecularly oriented morphology belonging to - phase [6]. However, as seen from the

31

figure above, the dipole vectors are still not uniquely oriented but lie randomly in planes

normal to the molecular chains. Thus it is required to align these dipoles in the direction

of an externally applied field normal to the film.

Figure 1.2.8 Schematic depiction of the two most common chain conformations of PVF2 crystals: (a) tg+tg_ (b) tttt. The arrows indicate the projections of the CF2 dipole moments

in the plane containing the carbon backbone [6].

It has also been shown that the anisotropy disappears at high poling fields, that

dipoles are, in fact, reoriented during application of an electric field, and that other

typical phenomena accompanying ferroelectricity, hysteresis loops and Curie transitions

are also seen in PVDF [27].

Copolymers of vinylidene fluoride and trifluoroethylene also exhibit another

important aspect of ferroelectricity that has not been convincingly demonstrated in

PVDF. This is the Curie temperature at which a ferroelectric crystal undergoes reversibly

a solid-state transformation to a nonpolar (paraelectric) state. In these copolymers, the

32

Curie transitions found to involve primarily intramolecular changes of dipole directions

through introduction of g± bonds that alter the polar all-trans conformation to a somewhat

disordered arrangement of tg± and tt sequences.

Table 1.2.1: Physical Constants of some Piezoelectric Polymers Symbol Parameter PVDF Co polymer Units

t Thickness 9, 28, 52, 110 various µm (micron, 10-4)

d31 23 11

d33

Piezo strain constant

-33 -38

1012 (m/m)/(V/m)

or (C/m2)/(N/m2)

g31 236 162

g33

Piezo stress constant

-330 -542

10-3 (V/m) /(N/m2)

or (m/m)/ (C/m2)

k31 12% 20%

k1

Electromechanical

coupling factor 14% 25-29%

C Capacitance 380 for 28µm 68 for 110µm pF/cm2 @ 1kHz

Y Young’s Modulus 2-4 3.5 103 N/m2

1.5 2.3 V0 Speed of

sound

stretch

thickness 2.2 2.4

103 cm/s

ρ Pyroelectric coefficient 30 40 10-6 C/m2 °K

ε Permittivity 106-113 65-75 10-12 F/m

ε/ε0 Relative Permittivity 12-13 7-8

ρm Mass density 1.78 1.82 103 kg/m

ρe Volume resistivity >1012 >1011 Ohm meter

R

R

Surface Metallization

resistivity

tan δ Loss tangent 0.02 0.015 @1kHz

Yield strength 45-55 20-30 106 N/m2 (stretch axis)

Temperature range -40 to 80 -40 to 115 ... 145 °C

Water absorption <0.02 <0.02 %H2O

Maximum operating

voltage

750 (30) 750 (30) V/mil (V/µm), DC,@

25°C

Breakdown voltage 2000(80) 2000(80) V/mil (V/µm), DC,@

25°C

33

1.2.9 Transducer Applications of PVDF

Piezoelectric transducers are commonly used to measure the response in a

structure to impressed acoustic fields and other forces. By definition, an ideal sensor must

be non-intrusive and its output must be a close representation of the response of the

structure at that location. Thus the design of the sensor - its shape and size, material

properties relative to the structure, its location in the structure - all contribute to

sensitivity and fidelity. In addition, material damping in the plate and the transducer must

be taken into account. The sensor response involves computation of the voltage induced

in the transducer by an incident acoustic wave in the fluid or a force to which the plate

may be subjected to [28].

Microphones are transducers, which transform sound pressure waves into

electrical signals. There are several different types, but usually they convert pressure

variations into deflections of a diaphragm, resulting in a voltage whose rate of change

reflects the frequency of the sound, and whose amplitude corresponds to the amplitude of

the sound. A large number of potential applications have been suggested for PVDF. The

range of applications may be split conveniently into five types: pyroelectric sensors,

electromechanical devices, audio-frequency transducers, sonar hydrophones and

ultrasonic transducers.

PVDF with its low acoustic impedance (Table 1.2.1) and high g constant (about

15 times the value of some ceramics) makes an ideal transducer material for broad band

underwater receivers, i.e., hydrophones. In order to develop usable hydrophones, multiple

layers of PVDF can be laminated together using a thin layer of adhesive (~1 µm) to form

thicker stacks of material with electroded sheets and contacts where necessary. The

34

multilayer technique has enabled to produce hydrophones with a capacitance of 1500pF

and a sensitivity of 45dB ref µVPa-1 at 10 KHz, dropping off slightly to 36 dB at 100 Hz,

adequate to detect sea-state-zero noise levels [25]. These hydrophones show a very little

variation in sensitivity with temperature, storage or temperature cycling, and are a viable

alternative to ceramic devices, with increased frequency range.

Table 1.2.2: Acoustic Properties of Various materials [25]

Table 1.2.3: Comparison of piezoelectric materials [29]

35

This table provides a comparison of the piezoelectric properties of PVDF

polymer and two popular piezoelectric ceramic materials and gives an idea as to why

PVDF is suited for this research.

Figure 1.2.9 Typical Infrared absorption spectrum of PVDF film [29]

PVDF film’s typical optical transmission characteristics are shown in the Figure

1.2.9. The strong absorption of infrared energy at 7-20µm wavelengths makes piezo film

ideal for intrusion detection and energy management devices.

1.2.10 Bending Mode Piezoelectricity in Polymers

Bending mode piezoelectricity in polymer [30-32] offers very high sensitivity and

low mechanical input impedance (better sensitivity at lower frequencies) but has not been

studied in as much detail. The bending piezoelectric effect is described by the following

equation [31,32]:

Where D3 and E3 are the dielectric displacement and the electric field,

respectively, in the thickness direction, 3 the dielectric constant, R is the radius of

36

curvature for bending deformation, and β331 is defined as the bending piezoelectric

constant. If a cantilever film of thickness t, length l, width h and capacitance Co is bent

by x distance at the tip, and the voltage generated is V3, then β331 is given as [32]:

where both equations 1.2.8 and 1.2.9 are in CGS units. Typically, films of approximate

dimensions l = 1 cm; t = 50 - 100µm yield voltages in the mille volt range for bending

directions in the micrometer range [32, 33]. The bending piezoelectric constants of many

polymers, under different processing conditions, are presented in Table 1.2.4, reproduced

here from ref. [32]. It should be noted that, unlike the linear piezoelectric constants (d, g,

e and f), which are derived from crystallographic point group symmetry considerations

(resulting in a unique piezoelectric compliance matrix for each point group) [20], the

bending constant (β331) does not have a crystallographic origin. This is apparent from an

examination of the piezoelectric compliance matrix of the orthorhombic point group mm2

(symmetry of β phase PVDF) in Equation 1.2.2. According to the compliance matrix,

there is no electrical component in the 3' direction due a shear stress in the 31' plane ('5'

direction in reduced notation), i.e., a 331' (can be written as 35' in reduced notation)

electromechanical coupling does not exist for this point group symmetry. Thus, 331 must

simply be a bulk parameter; the material constant connecting the charge produced in the

3' direction (i.e., on film plane electrodes) in response to bending deformation.

37

Table 1.2.4: Bending Piezoelectric Constants [32] Polymer Mech.

Treatment Thickness Poling

Field (kV/cm)

Poling Temp. (°C)

331

Polytetrafluoroethylene undrawn

poll-drawn 290 110

430 700

225 200

2.1 1.6

Polyvinylideneflouride roll-drawn 125 400 80 19.1 Polyhlorotrifluroethylene undrawn 150 350 150 0.35

Polyphenyleneoxide undrawn 80 700 150 0.3-0.23 Teterafluoroethylene -ethylene-copolymer

undrawn 50 500 170 0.1

Tetrafluoroethylene -hexafluoropropylene

-copolymer

undrawn 25 930 150 0.1

Polysulphone undrawn 75 130 150 0.26-0.15 Polyethyleneterapthalate undrawn

poll-drawn 40 45

300 230 670

100 180 RT

0.2-0 0.1-0

0.10-0 Polycarbonate poll-drawn 35 590 120 0.18-0 Polyethylene undrawn 100 200

400 80 RT

0.1-0 0.1-0

Cellulose Triacetate undrawn stretched

70 100

340 0

150 0.54-0 0.41

1.2.11 Piezoelectric Cochlear Implant

Conceptually speaking then, piezoelectric materials are indeed the most suitable

for replacing “organs” for the nonexistent transduction mechanism in patients suffering

from profound Sensorineural hearing loss. A piezoelectric device, placed in the scala

tympani of the human ear will generate charge across its electrodes by virtue of the direct

piezoelectric effect resulting from the pressure waves. This charge, in turn would give

rise to ionic currents between the electrodes of the piezoelectric. If this current is similar

in magnitude as the charge/current used by the existing implants, it will fire the nerve

fibers in the organ of Corti, in much the same manner as the existing cochlear implants.

Consideration for use of piezoelectric devices also stems from the following facts:

38

1. Piezoelectric are used routinely in applications that involve detection of human

voice, musical notes, etc., demonstrating the ability to transduce complex acoustic

stimuli [21, 22, and 34-36].

2. Being natural transducers, piezoelectric function without the need for an external

power supply; presenting the possibility of obtaining an entirely “stand alone”

unit.

3. This device can provide a bipolar electrode arrangement [9, 11].

4. This device would be non-invasive to the cochlea once implanted, unlike present

cochlear implants.

Sound pressure levels in the cochlea have been measured to be a maximum of ~

120 dB SPL (ref. 20µPa) [37, 38]. It is also known that the existing cochlear implants use

the charges of the order of 1 - 100 nano-coulombs to stimulate the nerve fibers. This

suggests that a successful piezoelectric implant should be able to generate nano-coulombs

of charge at pressure of ~ 120 dB SPL.

Roseman, Willging, Buchanan, Wang and Dong [39, 40] performed initial

studies on the scope of implantable piezoelectric cochlear implants. In-vivo experiments

were carried out by implantation of ceramic piezoelectric transducers in the cochlea of

laboratory guinea pigs. The hearing capability of the guinea pigs was ablated by high

intensity of acoustic waves. Auditory brainstem response (ABR) test (tracing of neural

responses through the auditory pathway) were performed at various intervals on guinea

pigs with and without cochlear implants. After transducer implantation, the guinea pig.s

hearing ability had recovered 40-50 dB compared to the control guinea pigs. However,

the ABR response was not reproducible and also erratic, which may have been caused by

39

factors other than the piezoelectric charge generation. Though no formal conformation of

biocompatibility was carried out, ceramic samples implanted in guinea pig cochlea for

more than 6 months did not show any signs of erosion or fibrosis.

Mukherjee, Roseman and Willging [40, 41] carried out in-vitro experiments on

the feasibility of a ceramic implant in a mechanical model of the cochlea (test cell). The

most important findings were:

1) The specific acoustic impedance (Z) of water is ~ 1.5 x 106 kg/ (m2-sec) and that of

the most commonly used piezoelectric ceramic Lead Zirconate Titanate (PZT) is ~ 24

106 kg/ (m2-sec) [42]. As a result of this large difference in the acoustic impedance of

water and ceramic, the transduction efficiency of a ceramic piezoelectric device is quite

low and needs to be augmented by ‘matching’ layer(s) between ceramic and water.

Maximum impedance matching exists when the matching layer thickness is n/4 (where:

n = 1, 3, 5. and is the wavelength in the matching layer material) [43]. Using a

Neoprene coating as the Dielectric on the electrodes, the device sensitivity increases

monotonically with increasing dielectric layer thickness. This was attributed to an

improvement in impedance matching with increasing layer thickness. Though the devices

measured acquired significantly high sensitivities as the neoprene layer thickness was

increased, this method is not suited for a cochlear implant since the resulting device

would be too large to fit in the scala tympani.

2) Pressure - Distance test performed on these devices resulted in a finding that, within

the frequency range of test (100-600 Hz), the sound pressure level decays exponentially

with distance from the front diaphragm of the test cell. These results are related to the

events occurring in the cochlea.

40

3) The voltage output from the Shockwave tests in saline solutions of different

concentrations is independent of the solution salinity if the electrodes were coated with a

dielectric coating. This proved that the piezoelectric charge on electrodes in a saline

solution causes ionic polarization and electrolytic reactions to occur (hence some charge

is dissipated through the water and cannot be detected).This directly proved the capability

of a piezoelectric device to cause ionic polarization in perilymph fluid.

41

1.3 Device Requirements and Design

1.3.1 Device Requirements

The environment, conditions of use, and the properties required of a piezoelectric

cochlear implant are unique to this application. Therefore based on initial studies and

literature survey of micromechanics and acoustics the principal requirements for such an

application are discussed below:

Small Dimensions and Flexibility:

The scala tympani have a length of ≈ 15 mm from base to apex. The diameter of

this duct is ≈ 2 - 3 mm in humans [1]. These devices not only become difficult to

fabricate but also lowers the surface area and also limits the total charge produce. This

demands an increased sensitivity on the piezoelectric cochlear implant. In the transverse

cross-section of the duct the device cannot be larger than ≈ 2 mm × 2 mm. The device

should also not block the passage and prevent the flow of sound vibrations. The

structures inside the basilar membrane are extremely delicate. Therefore a stiff device

presents a danger of rupturing it. Ceramics have a very high stiffness. Therefore it posses

a potential problem if used. If the device is place length wise in the scala tympani, it can

be several centimeters long. The piezoelectric effect is sensitive to changes in sound

pressure. Therefore the device needs to be anchored in a way to prevent movement and

ensure repeatability over the years.

Specific Acoustic Impedance (Z) Matching:

The specific acoustic impedance mismatch between water/saline medium and

ceramic/metal is considerable. In this regard, polymers provide a distinct advantage over

42

these materials. Impedance (Z) of the polymer PVDF is 2 - 3 x 106 kg/ (m2 . sec) and is

comparable to that of the water, which has a Z value of ~1.5 x 106 kg/ (m2 . sec).

High Sensitivity:

Sensitivity is the most critical device requirement as the device needs to operate

under hydrostatic conditions. Under these conditions the piezoelectric stress (dh) and

voltage (gh) are very low. If a large degree of impedance mismatch also exists, the

effective sensitivity is further lowered. Therefore a considerably high figure of merit

(dhgh) is essential. Bending piezoelectricity in polymers could offer high sensitivities and

hence are the most likely candidates for this application.

Even Frequency Response:

The human ear is sensitive to sounds in the frequency range of ≈ 20 Hz - 20 kHz.

Although the threshold response of the healthy ear is not equal over this range, the

piezoelectric device must have a more or less flat response over this range of frequencies,

since it is also well known that the dynamic range of sensitivity of the ear to electrical

signals from scala tympani electrodes is far less than the dynamic range of sensitivity of

the healthy ear to sound pressures. To overcome this factor, a signal conditioning

operation called compression is performed on the signals used in the present cochlear

implants [9]. A piezoelectric device would not have this capability, and any resulting

limitation on device performance will have to be determined.

Biocompatibility:

The materials used for a cochlear implant device must be biocompatible. The

predominantly non-tissue environment of the scala tympani leads to the expectation that

biocompatibility issues will not be a limiting factor for such a device. However, it is

43

proposed that proven biocompatible, negligible swelling (in 0.9 wt. % saline), and

relatively chemically inert materials be selected for the piezoelectric and device

packaging materials proposed to be used for such an implant. This will largely prevent

complications at later stages of device development.

1.3.2 Design

Considering the requirements, it is noted that a device made of flexible polymer

like PVDF, used in the flexural/ bending mode would be best suited for the application.

The flexural/ bending mode is chosen because highest sensitivities are realized in this

mode. The voltage sensitivity (g - constant) of PVDF exceeds that of the ceramic

transducers but its charge sensitivity (d - constant) is lower [6]. However, polymer

devices could be made in large sizes thus improving upon the total charge obtained.

Electrode Design

The design of electrodes for cochlear prosthesis has been the focus of research

for over two decades. Some of the issues associated with electrode design are:

1. Electrode placement

2. Number of electrodes

3. spacing of contacts

4. Orientation of electrodes with respect to the excitable tissue, and

5. Electrode configuration.

Electrodes may be placed near the round window of the cochlea (extracochlear),

or in the scala tympani (intracochlear) or on the surface of the cochlear nucleus. Most

commonly, the electrodes are placed in the scala tympani because it brings the electrodes

in close proximity with auditory neurons, which lie along the length of the cochlea. This

44

electrode placement is preferred because it preserves the "place" mechanism used in a

normal cochlea for coding frequencies. That is, auditory neurons that are "tuned" for high

frequencies are stimulated whenever the electrodes near the base are stimulated, whereas

auditory neurons that are "tuned" for low frequencies are stimulated whenever the

electrodes near the apex are stimulated. In most cases, the electrode arrays can be inserted

in the scala tympani to depths of 22 - 30 mm within the cochlea.

The number of electrodes as well as the spacing between the electrodes affects

the place resolution for coding frequencies. In principle, larger the number of electrodes,

finer the place resolution for coding frequencies. Frequency coding is constrained,

however, by two factors, which are inherent in the design of cochlear prosthesis:

1. Number of surviving auditory neurons that can be stimulated at a particular

site in the cochlea, and

2. Spread of excitation associated with electrical stimulation. Unfortunately,

there is not much that can be done about the first problem, because it depends

on the etiology of deafness.

Ideally, we would like to have surviving auditory neurons lying along the length of the

cochlea. Such a neuron survival pattern would support a good frequency representation

through the use of multiple electrodes, each stimulating a different site in the cochlea.

A “toothbrush” type of design is proposed for the device, which is based on the scheme

and function of the stereocilia, which resides on the cuticular plate of the organ of corti

[3]. This design consists of several elements of the piezoelectric polymer film embedded

in a polymer substrate in the form of vertical cantilevers. The substrate is made up of

insulating material that can be shaped to fit into the scala tympani. The piezoelectric

45

elements, electroded with polymer thin films will face the incoming pressure waves

which will cause bending deformation, eliciting charges on the electrodes. The electrodes

of all the elements can be connected in a series electrical circuit (voltages add), by means

of interconnects, on the substrates. These interconnects can be metallic or polymeric. The

total voltage can be discharged by means of two noble metal electrodes. The phase

difference in the voltages between the elements is expected to be minimal as the

separation between the elements (≈ 1 - 2 mm) is several orders of magnitude smaller than

the wavelength of the sound waves (for example, = 1.5 m for 1 KHz underwater). If the

total voltage produced is applied to the perilymph through noble metal electrodes, the

electrical connections on the substrate and the electrodes on the piezoelectric elements

have to be insulated from the solution by a dielectric polymer barrier.

The ‘toothbrush’ design is flexible and open to further engineering. It could be

altered for research to produce a better device. The design can be extended depending

upon the requirements. For instance, piezoelectric elements can be varied in their heights

which will affect the frequency range of high sensitivity. The charges from different

element groups can be discharged separately at different places along the scala tympani;

or if high charge density is not required, the charge from each element can be allowed to

discharge through the saline independently (by refraining from coating the electrodes

with a dielectric polymer and connecting the elements electrically).

46

Figure 1.3.1 Schematic of a device based on stereocilia. The vertical elements are

electroded polymer film. The base is a flexible bio-compatible polymer. The device is shown placed in the cochlear scala tympani. The piezoelectric elements face the

incoming sound pressure.

47

1.4 Polydimethylsiloxanes

1.4.1 Introduction

Polydimethylsiloxanes (PDMS) are a type of polymeric organosilicon compounds

which are commonly referred to as "silicones". They are made by bonding organic groups

to silicon atoms in long (polymeric) silicon-oxygen chains. Usually, the organic groups

are methyl, longer alkyl, fluoroalkyl, phenyl and/or vinyl.

PDMS is used as a mold release agents; lubricants and greases; damping,

hydraulic, heat transfer, or dielectric fluids (transformer cooling fluids); in textile

softening and water repellant agents; as a softener in silicone rubber products; sealants

and caulking compounds; antifoam agents; in food processing; anticaking agent in

powdered food; in coatings, paints, inks, waxes and polishes; in personal care products

and cosmetics; and has many medical uses such as surgical tubing and implants and for

reducing gastrointestinal gas.

The versatility of silicones is largely attributed to the characteristics of the

siloxane backbone. Polydimetylsiloxanes exhibit relatively free rotation about the

siloxane bond and the spacing of the backbone allows free rotation of methyl group

attached to the silicon. The low barrier to rotation contributes excellent low temperature

properties of PDMS. In contrast to polymers with organic background PDMS remains

relatively constant over a wide temperature range [136].

48

1.4.2 Health Effects

Effects of Short-Term (Acute) Exposure

Inhalation:

Polydimethylsiloxanes (PDMS's) do not form vapours at room temperature. Based on

animal information, even very high exposures to aerosols are not expected to produce

health effects. Therefore, PDMS's are considered to be essentially non-toxic.

Skin Contact:

PDMS's are not expected to cause skin irritation, based on human and animal

information.

Effects of Long-Term (Chronic) Exposure

No harmful effects have been observed in long-term animal studies.

49

1.5 Conductive Polymers

1.5.1 Electrical Conductivity

Conductivity is defined by Ohm’s law:

Where, I is the current (in Amperes) through a resistor and V is the drop in potential (in

Volts) across it. The proportionality constant R is called the ‘resistance’, measured in

Ohms (Ω). The reciprocal of resistance (R-1) is called ‘conductance’. Ohm’s law is an

empirical law, related to irreversible thermodynamics (Ilya Priogogine, Noble Prize in

Chemistry, 1977), the flow I as a result of a gradient in potential leads to energy being

dissipated (RI2 Joules s-1).

It is known that not all materials obey Ohm’s law. Gas discharges, vacuum tubes,

semiconductors and what are termed one dimensional conductor (e.g. linear polyene

chain) generally all deviate from Ohm’s law.

In Ohmic material the resistance is proportional to the length l of the sample and

inversely proportional to the sample cross-section A:

Where, is the resistivity measured in Ω cm (in SI units). Its inverse = -1 is the

conductivity. The unit of conductance is the Siemens (S = Ω-1). The unit of conductivity

is S m-1.

Conductivity depends on the number density of charge carriers (number of

electrons n) and how fast they can move in the material (mobility µ):

50

Where, -e is the electron charge. In semiconductors and electrolyte solutions, positive

charge carriers (holes or cations) are also accounted for. Conductivity depends on

temperature, it generally increases with decreasing temperature for metallic materials

(some of which become superconductive below a certain critical temperature Tc), while it

generally decreases with lowered temperature for semiconductors and insulators.

In many materials, such as crystals, stretched polymers or liquid crystals,

macroscopic properties such as strength and optical and electrical properties generally

depend on direction. They are said to be anisotropic. Similarly, the material’s electrical

conductivity may depend on direction and be anisotropic. Anisotropy is also interesting

in other contexts of stretch aligned polymers: when the absorption of light is anisotropic

the materials acts as a polarizer. Also mechanical strength is anisotropic: aligned

polyacetylene fibers are known to be very strong along the orientation direction.

1.5.2 Conductive Electrodes

Conductive Electrodes are an integral part of every piezoelectric transducer. Their

most common form is a layer of metal intimately bonded to the substrate surface. Ideally,

the electrode should adhere strongly to the substrate, it should be very thin, should have

practically zero resistance and good chemical and physical durability. Also it should be

possible to solder leads to the electrodes so that the solder bond has good pull strength

and reliability. In practice, these ideals are not always obtainable [22].

Electrode adherence is highly critical. If there is any lack of intimate bonding, the

gap between the electrode and the high-K substrate acts as a series capacitance of low

value. The presence of an air gap lowers the effective capacitance of the transducer very

strongly. High voltage can cause heavy damage to the effectiveness of the transducer. In

51

extreme cases, the electrode can separate from the substrate, ruining its usefulness

completely.

The low elastic modulus and the ability to withstand high strain without failure

make the conducting polymer attractive for a wide range of acoustic applications based

on high-strain electroactive polymers [44]. The all-polymer composite films are flexible,

with strong coherent interfaces between the electrostrictive polymer layer and the

conductive polymer layer.

Polymers that exhibit a large strain response induced by electric fields have

attracted a great deal of attention in recent years. These polymers include Polyvinylidene

fluoride (PVDF) and its copolymer with trifluoroethylene, polyurethane, odd-numbered

nylons, etc. However, increased interest in using high strain electroactive polymeric

materials for electro-acoustic and electromechanical applications also raises the issue of

new electrode materials to meet new requirements and to provide better performance. For

instance, to achieve high acoustic transparency, very small acoustic impedance

mismatching between the electrode and electrostrictive polymers in required. Because of

a high elastic modulus compared with electrostrictive polymers, the commonly used

metal electrodes, such as Au, Al, Ni-Cu alloys, may impose mechanical clamping on the

polymer, which can reduce the electric field induced strain level and the efficiency of the

electromechanical transduction. Also, the electrostrictive polymer (PVDF) has a large

transverse strain, more than 3%. In general, at such a high strain level, thin metal

electrodes will crack and cause failure in the devices. Hence, a new electrode material

that can lower the clamping effect and withstand high strain is highly desirable. It is

believed that a conducting polymer electrode will meet these requirements. Due to

52

flexibility, low acoustic impedance, and elastic modulus of conductive polymer

electrodes, such all-polymer electroactive systems may improve the performance of

electromechanical polymer materials in acoustic and electromechanical applications [44].

1.5.3 Conductive Polymers

A key property of most polymers, which distinguishes them from metals, is their

inability to carry electricity. Whereas, the insulating properties of polymers are a

significant advantage for many applications of plastics.

During the past 25 years, however, a new class of organic polymers has been

devised with the remarkable ability to conduct electrical current. Part of a larger class of

materials called “synthetic metals”. Some of these conductive plastics are already under

development for practical applications.

A major obstacle to the rapid development of conductive polymers has been the

lack of understanding of how electrical conductivity works in these polymers. All

conductive polymers have one thing in common: They contain extended - conjugated

systems with single and double bonds alternating along the polymer chain. An

understanding of the relationship between the chemical structure of the repeating unit of

the polymer and its electrical properties would enable the electronic and mechanical

properties of these materials to be tailored at the molecular level [45].

53

Figure 1.5.1 Conductivity ranges for polymers (doped and undoped), inorganic materials

and Molecular crystals (from Cowie).

1.5.4 Electrical Properties

The electrical properties of any material are determined by its electronic structure.

Band theory explains the electronic structure of the materials. In the solid state, the

atomic orbitals of each atom overlap with the same orbitals of their neighboring atoms in

all directions to produce molecular orbitals similar to those in small molecules. In a solid,

the number of atomic orbitals is about 1022 per cc, and thus the number of molecular

orbitals would also be 1022. When this many orbitals are spaced together in a given range

of energies, they form what looks like continuous energy bands [45].

54

Figure 1.5.2 Insulator-Semiconductor-Metal (Band Diagram)

The energy spacing between the highest occupied and lowest unoccupied bands is

called the band gap. The highest occupied band is called the valence band, and the lowest

unoccupied band is the conduction band [45].

The electrical properties of conventional materials depend on how the bands are

filled. When the bands are filled or empty no conduction occurs. If the band gap is

narrow, at room temperature thermal excitation of electrons from the valence band to the

conduction band gives rise to conductivity. This is what occurs in classical

semiconductors. When the band gap is too wide, the thermal excitation at room

temperature is insufficient to excite electrons across the gap and the solid is an insulator.

The high conductivity of metals is due to partially occupied bands - a partially filled

conduction band, a partially empty valence band, or a zero band gap.

Conductive polymers are peculiar in that they conduct current without having a

partially empty or partially filled band. Their electrical conductivity cannot be explained

well by simple band theory. For example, the simple band theory cannot explain why the

55

charge carriers, usually electrons or holes, in polypyrrole (PPY), polyacetylene (PA) are

spinless.

When an electron is removed from the top of the valence band of a conjugated

polymer, such as PPY or PA, a vacancy (hole or radical cation) is created that does not

delocalize completely, as would be expected from classical band theory. Only partial

delocalization occurs, extending over several monomeric units and causing them to

deform structurally. The energy level associated with this radical cations represents a

destabilized bonding orbital and thus has a higher energy than the energies in the valence

band. In other words, its energy is in the band gap. This rise in energy is similar to the

rise in energy that takes place after an electron is removed from a filled bonding

molecular orbital [45].

1.5.5 Doping

Conductivity of polymer can be increased several-fold by doping it with oxidative

/ reductive substituents or by donor/ acceptor radicals. Shirakawa and Ikeda [46]

discovered that doping of PA with metallic regimes increases its conductivity by 9 - 13

orders of magnitude. Doping is accomplished by chemical methods of direct exposure of

the conjugated polymer to charge transfer agent (dopant) in the gas or solution phase, or

by electrochemical oxidation or reduction.

The doping is usually quantitative and the carrier concentration is directly

proportional to the dopant concentration. Dopant of conductive polymers involves

random dispersion or aggregation of structures of entangled chain and fibrils. Polymer

doping leads to the formation of conjugational defects, viz. solitons, polarons or bi-

polarons in the polymer chains. An x-ray diffraction study on iodine-doped polyacetylene

56

shows that the C-C bond length of the polyacetylene chain increases with donor doping

but decreases on acceptor doping [47].

The presence of localized electronic states of energies less than the band-gap

arising from changes in local bond order, including the formation of solitons, polarons

and bipolarons have led to the possibility of new types of charge conduction present in

the polymer systems.

Figure 1.5.3 Typical properties of solitons, polarons and bipolarons [47].

Doping agents or dopants are either strong reducing agents or strong oxidizing

agents. They may be neutral molecules and compounds or inorganic salts which can

easily form ions, organic dopants and polymeric dopants. The nature of dopants plays an

important role in the stability of conductive polymers. Stabilization of the conductive

polymers has been tried by two routes, viz. incorporation of antioxidants, such as

benzoquinone and hindered phenols or by using radical traps such as

azobisisobutylonitrile. Another method is ion implantation.

In solid-state physics, radical cation that is partially delocalized over some

polymer segment is called a Polaron. It stabilizes itself by polarizing the medium around

it, hence the name. Since it is really a radical cations, a polaron has a spin of ½.

57

If another electron now is removed from the already oxidized polymer containing

the polaron, two things can happen: This electron could come from either a different

segment of the polymer chain, thus creating another independent polaron, or from the

first polaron level (remove the unpaired electron) to create a special di-cation, which

solid-state physicists call a Bipolaron. Low doping levels give rise to polarons, whereas

higher doping levels produce bipolarons [45].

Figure 1.5.4 Polaron-Bipolaron (Band Diagram) [45]

The bipolaron also has structural deformation associated with it. The two positive

charges of the bipolaron are not independent, but act as a pair. Both polarons and

bipolarons are mobile and can move along the polymer chain by the rearrangement of

double and single bonds in the conjugated system that occurs in an electric field. If a

large number of bipolarons are formed, as a result of high doping, their energies can start

overlapping at the edges, which creates narrow bipolaron bands in the band gap.

58

Figure 1.5.5 Propagation of a polaron through a conjugated polymer chain by shifting of

double bonds [45].

In PPY, low doping concentrations create paramagnetic polarons, which, as the

degree of doping increases, convert to spinless bipolarons, which extend over about four

pyrrole rings. A polaron and bipolaron in polypyrrole are shown below:

59

Figure 1.5.6 Schematic showing Conductivity phenomenon in Polypyrrole [45]

The charged fragment made of four pyrrole rings can travel along the chain by the

rearrangement of double and single bonds.

In PA, which has a degenerate ground state (two equivalent resonance forms), the

bipolaron dissociates into two independent cations, which are spinless also and are called

Solitons. Solitons do not form in polymers with non-degenerate ground states, such as

60

polypyrrole, polythiophene and polyphenylene. These polymers are called non-

degenerate because their resonance forms are not identical if they are superimposed.

Figure 1.5.7 Resonance forms (aromatic and quinoid) of PPY are not equivalent [45]

Conduction by polarons and bipolarons is now thought to be the dominant

mechanism of charge transport in polymers with non-degenerate ground states. These

concepts also explain very well the optical absorption changes seen in these polymers

with doping.

Conductivity is the product of two important factors: the number of carriers,

electrons or holes and carrier mobility. The electrical conductivities of most conductive

polymers are in the same range as those of inorganic semiconductors. Semiconductors

have a very low number of carriers, typically 1016 to 1018 per cc, but very high

mobilities, typically 102 to 105 cm2 per volt-second, which results from a high degree of

crystallinity and purity and a lack of defects in these materials. However, most of the

conductive polymers are either amorphous or slightly crystalline at best.

61

On the other hand, the number of carriers in conductive polymers, typically 1021

to 1023 per cc, is four to five orders of magnitude higher than in inorganic

semiconductors. It is roughly equal to the number of cations per cc, and is therefore as

high as possible. But the mobility of conductive polymers is limited to only 10-4 to 10-5

cm2 per volt-second. Thus to achieve higher conductivities, higher mobilities are needed

which would result from a higher crystallinity, better orientation and a defect free

material [45].

1.5.6 Applications Of Conductive Polymers

Developments toward the synthesis of new and processable polymers as well as

discovering the broad range of physical phenomena and chemical flexibility has opened

opportunities for new technological applications [48-52]. The higher environmental

stability and modification of properties to suit a given end use and processability

achieved with the polymers derived from acetylene, pyrrole, thiophene, aniline and their

derivatives have emerged as materials to replace metals and semiconductors in the

electrical and electronics industry, as well offering themselves as the materials for

optoelectronic industry.

62

Figure 1.5.8 Technological Applications of Conductive Polymers [47]

63

1.6 Polypyrrole

1.6.1 Introduction

In recent years, intrinsic conducting polymers with conjugated double bonds have

attracted much attention as advanced materials. Among those conducting polymers,

Polypyrrole (PPY) is especially promising for commercial applications because of its

good environmental stability, facile synthesis, and higher conductivity than many other

conducting polymers. PPY can often be used as biosensors [53-54], gas sensors [55-56],

wires [57], micro actuators [58], anti-electrostatic coatings [59], solid electrolytic

capacitor [60,61], Electro-chromic windows and displays, and packaging, polymeric

batteries, electronic devices and functional membranes, etc. [62-64].

PPY coatings have excellent thermal stability and are good candidate for use in

carbon composites [65]. Furthermore, the electrochemical process parameters affecting

the properties of the PPY coatings have also been investigated [66]. PPY can easily be

prepared by either an oxidative chemical or electrochemical polymerization of pyrrole.

However, synthetically conductive PPY is insoluble and infusible, which restricts its

processing and applications in other fields.

In 1994 Biswas and Roy [67] studied the thermal, stable, morphological, and

conductive characteristics of the PPY prepared in aqueous medium. The results showed

that the PPY exhibits a spongy texture, the initial decomposition temperature at 180-

237°C and glass-transition temperature at 160-170°C.

1.6.2 Polymerization Mechanism For Polypyrrole

The chemical oxidation of pyrrole to produce an uncharacterized polymeric

‘pyrrole black’ has been known for many years.

64

The mechanism of polymerization is generally believed to involve the oxidation

of the monomer to yield a radical cation, so that the oxidation potential of the monomer

determines whether initiation can take place.

Figure 1.6.1 Schematic showing chemical synthesis of Polypyrrole

Provided there are no nucleophiles in the system capable of reacting with the

radical cations, they will couple to give a dimeric di-cation, which readily eliminates 2H+,

rearomatizing to five the pyrrole dimer. The oxidation potential of the dimer is slightly

lower than that of the monomer so the process of oxidative coupling of the units

continues as the polymer grows, becomes insoluble and precipitates on the electrode

surface where growth apparently continues to give a high molecular weight polymer.

The overall reaction of the chemical polymerization of pyrrole (PPy) in the

presence of an oxidizing agent, Ox, is usually presented as:

65

Where, Red is a reduced form of the oxidizing agent.

During chemical PPY synthesis the reaction conditions change with time. In

addition to the decreasing concentration of reactants and pH change, the oxidation

potential of the solution, E, expressed by the Nernst equation decreases due to the change

of the ratio Cox/CRed:

Where, Eθ is the standard solution potential. Large amount of oxidizing agent is

consumed for Pyrrole polymerization.

1.6.3 Monomer - Oxidant Ratio

The quality and composition of the polypyrrole films varies with the molar ratios

of the monomer to the oxidant of the reaction systems. Structural studies of PPY.s

through an ab-initio evaluation of bonding [68] and Monte Carlo growth approach to the

branch formation [69] suggests that the energy differences of different types of structures

mean that a great deal of branching is probable. The approximate probability function

generated in this manner is used in a statistical mechanical approach to estimate the

extent of bonding involving carbons as well as branching in PPY’s. The statistical

analysis of the growth process shows that the branching occurs even at very short chains.

The extent of branching does not depend on the chain length but it is a slowly varying

function of the temperature [69].

66

Monomer to oxidant ratio plays an important role in the quality of the film

developed using In-situ polymerization-deposition technique. Effects of the ratio have

been discussed rigorously by Saurin and Armes [70].

Most chemical synthesis of conducting polymers is carried out under non-

stoichiometric conditions, with the monomer usually present in excess relative to the

chemical oxidant. This method is believed to reduce the possibility of overoxidation of

the conducting polymer [70]. The stoichiometric oxidant-monomer mole ratios for the

polymerization of pyrrole by FeCl3 are 2.33:1. The Stoichiometric reaction conditions

minimize the rate of polymerization (which increases with both increasing oxidant and

monomer concentration [71]) for a given yield of conducting polymer. If the rate of

polymerization is too high, stable colloids are not formed and the conducting polymer is

instead obtained as an un-processable macroscopic precipitate.

Myers [72] pointed out that initial oxidant-monomer mole ratios, higher than

stoichiometric values, were actually desirable for the synthesis of polypyrrole in certain

non-aqueous solvents. Nicholau et al. [73] said that regardless of the solvent, such

synthesis should suppress the formation of soluble pyrrole oligomers and favor the

formation of insoluble high-molecular-weight polymer. Oligomer formation can be

suppressed by using higher initial oxidant-monomer mole ratios for the polymerization of

pyrrole [70]. Moreover, a sharp decrease in conductivity is also observed (< 10-5 Ω cm-1)

for the polypyrrole powders obtained from the two syntheses with stoichiometric and

higher than stoichiometric oxidant concentrations. Martin’s group [74] has reported that

surprisingly good-quality polypyrrole bulk powders can be prepared with oxidant-

monomer ratios as high as 50:1 provided mild oxidants such as FeCl3 are used.

67

Saurin and Armes have shown that increasing the immersion time in the chemical

deposition technique results in a monotonic increase in the weight uptake of polypyrrole.

On the other hand there is a relatively little decrease in the sheet resistance of

polypyrrole.

Figure 1.6.2 Effect of Immersion time on normalized weight uptake and av. Sheet

resistance [70].

Effects of increasing the number of immersions of the substrate in the

polymerizing solution have been discussed. It is found that initially, a compact film of

about 100-300 nm thickness is deposited on the substrate. But further polypyrrole

deposition results in a profound change to a less dense, globular conducting polymer

morphology. Figure 1.6.2 depicts that the weight uptake due to polypyrrole deposition is

almost linear, with an average weight up-take of approximately 1-mg g-1 per coating

treatment. Weight up-take is accompanied with a concomitant near-exponential decrease

in the sheet resistance from approximately 1000 Ω per square initially down to 84 Ω per

square. An increase in the deposition of polypyrrole initially yields a continuous thin film

68

of PPY formed over the substrate surface but subsequently the conducting polymer is

deposited with a more globular morphology [75].

Figure 1.6.3 Effect of repeated PPY deposition [70].

The study of surface morphology of electro-generated PPY films of various

thicknesses and with different dopant anions such as chloride, sulfate, perchlorate and

dodecylsulfate suggests that the characterization of surface roughness with one common

parameter (RMS, roughness factor or Ra) is not justified. In general two types of surface

globules with different heights are present on the film surface, independent of film

thickness and dopant nature [76]. Study of interaction energy anisotropy of the pyrrole

dimer and the influence of selected geometry variations on the interaction energy

components yields that any structure variation connected with the external energy is more

than compensated for by the repulsion energy [77].

PPY obtained under the co-existence of a sulfonic surfactant and a phenol

derivative with an electron-withdraw group has a high conductivity of about 40 S.cm-1

69

[78] and superior environmental stability. These favorable properties seem to be caused

by the surfactant anion being selectively incorporated into the PPY backbone as the

dopant and the electron-withdrawing substituent of the phenol derivative interacting with

pyrrole monomer so as to improve the regularity of the PPY backbone [79].

At the end of the synthesis of the polymer, PPY is generally in its p-doped

(oxidized) state. However, the polymer can be easily reduced to the neutral state in virtue

of a reversible doping - undoping chemical process which involves fraction of the charge

used for the total polymerization synthesis:

The reversibility and characteristics of this process makes PPY a very attractive

material for application as electrodes in rechargeable batteries. The kinetics of the PPY

doping and undoping reaction is controlled by the diffusion of the doping anions A-

through the polymer structure. Under these circumstances, the surface morphology and

the primary structure of the PPY film electrodes become a crucial factor in determining

the charging - discharging efficiency.

Properties of PPY film electrodes prepared in the presence of sulfonate and

chlorate ions have been studied by cyclovoltammetric curves (i.e. more developed

cathodic and anodic peaks). The PPY electrodes formed in the presence of sulfonate ions

show a higher charging - discharging efficiency than the other ones under the same

experimental conditions [80]. Superior behavior of PPY film electrodes can be ascribed

to the more developed surface and to the easier diffusion of dopant ions during the

doping - undoping process.

70

In Organic semiconductor, dopant density does not correspond to carrier density.

A possible reason for observing less radical (spin) density compared to dopant density is

that the dopants tend to aggregate among themselves and therefore the number of dopant

molecules that actually interact with the polymer chain can be grossly over counted. In

addition, radicals may decay via cross-linking or some other mechanisms [81].

1.6.4 Rate Of The Polymerizing Reaction

Some authors [71, 82] assume that the electron-transfer rate is the rate-

determining step for the polymerization. In addition they assume that this step proceeds

through an outer sphere, activated-complex mechanism. In that case, the polymerization

rate will depend on how effectively pyrrole and the solvated Fe3+ can approach each

other. Since pyrrole is a weak base, the approach is easier for Fe3+ ions which are

solvated by water than those which have one or two hydroxide ions in the salvation

sphere. Thus the polymerization rate increases with increasing H+ concentration as the

hydrolysis equilibria are shifted to the Fe3+ ions which are solvated by water [83].

For Fe(ClO4)3 in water the following equilibria have to be considered [83]:

71

Because of the low coordinating tendency of the perchlorate ion, only the last 3

equations are important; their equilibrium constants are 10-2.91, 10-3.05, and 10-3.26

respectively [84].

These equations can be used to find the concentrations of iron complexes and acid

for any given Fe (ClO4)3 and HClO4 concentration. Van den Schoor has shown in his

paper that the contribution of equilibrium equation (1.6.8) to the total acid concentration

is very small. There is no significant difference in (Fe(H2O)6)3+ concentration when

equation (1.6.8) is considered or not. The model calculation which has been supported by

the UV/VIS and pH measurements, show that only two Fe3+ equilibria, viz. equations

(1.4.6) and (1.4.7), determine the Fe3+ concentration in the polymer solution.

1.6.5 Dopant

Several kinds of soluble PPY have been synthesized and investigated because of

its poor processibility due to its insolubility and infusibility. Lee et al. [85] found that

PPY doped by p-dodecylbenzene (DBSA) with a conductivity of 2 S/cm dissolves in m-

cresol or chloroform in the presence of an extra amount of DBSA. Compared to the

substitution method to prepare soluble poly (3-substituted pyrrole) and poly (N-

substituted pyrrole), doping is simple for PPY which gives comparable conductivity to

poly (3-substituted pyrrole), but much higher than poly (N-substituted pyrrole) [85-86].

It has been reported that PPY prepared by in-situ doping polymerization in the

presence of -naphthalene sulfonic acid (NSA) can be dissolved in m-cresol and shows

high conductivity [86]. When using the in-situ polymerization method to prepare soluble

PPY, a sulfonic acid in the polymerization strongly affects the solubility, conductivity,

morphology and thermo-stability of PPY.

72

Shen and Wan [87] studied the effects of different sulfonic acids on in-situ

polymerized PPY. They suggested that sulfonic acid bearing a long alkyl group is

favorable for increasing the solubility of the doped PPY. Thus it can be postulated that, if

a chemical group in sulfonic acid has strong interaction with the solvent, e.g. hydrogen

bonding, the solubility of PPY doped with this acid may be enhanced.

It has been proposed that polarons of in-situ doped PPY are generated by adding

H+ to the -position of the pyrrole ring during the in-situ doping polymerization of

pyrrole [86]. Therefore, the absence of polarons in sulfonic acid doped PPY indicates

that, during the polymerization of pyrrole, these acids dope PPY poorly with protons.

Morphology of PPY prepared by ‘two step process’, in which PPY is firstly

prepared by polymerization of pyrrole by a similar method but in the absence of sulfonic

acid, and then doped with different sulfonic acids, is granular and independent of the kind

of sulfonic acid. This means that the morphology of PPY is induced by sulfonic acid

during the ins-situ doping polymerization of pyrrole. Quinone Sulfonic acid has

bipolarons as the charge carriers where as other sulfonic acids have both, polarons and

bipolarons as the charge carriers [87].

The films produced utilizing Anthraquinone-2-Sulfonic acid (AQSA) as the

dopant, demonstrate an extremely smooth morphology when compared to other dopants.

Because the typical doping levels for these materials is 0.30, and because of the high

molecular weight of aryl sulfonates, the dopant comprises as much as 50% of the

polymer film mass. The relative stability of the AQSA film, particularly during the initial

stages of conductivity loss, is due in part to the smooth surface which inhibits oxygen

penetration. The more porous morphology of the films produced using other dopants

73

contributes to the relative instability of their films. Since the rate determining degradation

mechanism involves the diffusion or reaction of oxygen, the conductivity loss in an inert

atmosphere will proceed by a slower, non oxygen-dependent process.

Anthraquinone-2-sulfonic acid doped polypyrrole films exhibit conductivity and

stability superior to all other dopants. This exceptional performance is due, at least in

part; to morphology factors which limit the diffusion of oxygen into the polymer film

[88].

74

1.6.6 Anisotropic Properties In Polypyrrole

The Electric conductivity of polypyrrole films is summarized in Figure (1.6.4) and

(1.6.5).

Figure 1.6.4 Conductivity and Activation energy of Polypyrrole along the Surface [81]

Figure 1.6.5 Conductivity and Activation energy of Polypyrrole in the thickness direction

[81]

Electric Conductivity measured in two different directions, i.e. the conductivity

along the surface direction and that in the thickness direction shows a significant

difference. The surface electric conductivity is higher by about 4 orders of magnitude that

bulk electric conductivity. Some possible explanations for this anisotropy in electrical

conductivity can be given as follows: One is that the dopants tend to remain on the

surface of the film more than they penetrate into the bulk because of diffusion barrier.

Therefore, the conductivity can be much higher along the surface than in the thickness

direction. The other is that the electric conductivity in the thickness direction is

75

underestimated because of the slight non-ohmic characteristics of the Substrate/PPY

contacts in the small voltage region. The conductivity along the surface could be

measured by four point probe method where as this method could not be used to measure

conductivity in the thickness direction. The non-ohmic characteristics may be caused by

the surface states on the polypyrrole surface [81].

On a similar basis as that of conductivity, In-plane impedance of conductive

polymer thin films can be discussed. Generally, an impedance spectrum contains the

electrode response as well as the material response to an AC signal. Figure 1.6.4 shows

the frequency dependence of the real (Z*) and imaginary (Z*) parts of the in-plane

impedance (Z*) and Figure 1.6.5 shows the through-plane impedance of a polyaniline

film electrode.

76

Figure 1.6.6 In-plane Impedance of doped PANI film (a) Real and imaginary impedance

(b) Cole-Cole Plot [89]

The drop off of the real impedance and the peak of the imaginary impedance at

106 to 107 Hz shows there is a conductivity relaxation process occurring in this frequency

range. The good fit to a semicircle in the Cole-Cole plot means that this relaxation is

Debye-like and that there is a single relaxation time. The relaxation involves long range

conduction of charge carriers and is a direct result of the electrode material system being

measured.

77

Figure 1.6.7 Through-plane Impedance of doped PANI film (a) Real and imaginary


A similar conductivity relaxation is observed in the through plane impedance

case. However, compared with the in-plane impedance, the conductivity relaxation

occurs at a lower frequency. Also, the departure from a perfect semicircle can be

accounted for by a distribution of relaxation times [90] possibly due to inhomogeneous

dopant distribution through the film.

The frequency at which the conductivity relaxation occurs is strongly dependent

upon the doping level. As the doping level increases, the frequency at which conductivity

relaxation occurs, shifts to a higher frequency.

78

Similar conductivity anisotropy has been reported for polypyrrole film where the

in-plane conductivity is greater than the though-plane conductivity by four orders of

magnitude. Another reason responsible for the observed conductivity anisotropy is as

follows: the electric field for the through-plane measurement is 500 times the electric

field in the plane when the same oscillation voltage is used for in-plane and through-

plane measurements [89]. The relationship between the conductivity () and electric

field (E) is found to be ~ (ln E)-1. Other possible reason includes the difficulty of

dopants to fully penetrate the sample substrate.

1.6.7 Double Layer Capacitance

A study by Miller and Bockris [91], of PPY films synthesized by constant current

method on Pt electrodes yields that Polypyrrole shows a double layer capacitance of

37µF cm-2, which does not vary significantly over the potential range for which PPY is

highly conducting. The space charge capacity is constant but seems to decrease as the

potential is reduced. The surface state capacitance is also constant at the positive

potentials and increases as the potential is reduced. This indicates that as the potential is

reduced, approaching the redox potential of polypyrrole, the structure of the interphase

begins to change from a mixed metal-semiconductor situation to that of a semiconductor.

This corresponds with the theory that the charge carrier density of polypyrrole decreases

as the oxidized sites of the polymer are reduced, thereby increasing the semiconductor

character.

1.6.8 Adhesion Properties

The adhesion between a deposited polymer and a substrate is mainly determined

by the interfacial properties of the substrate. In many cases, the adhesion strength is

79

related to the surface free energy of the substrate [92]. Huang, Wang and Macdiarmid

have reported that the deposited thin films of polypyrrole adhere more strongly to the

hydrophilic, bare region on the substrate than to the hydrophobic region which can be

derivatized by using silicon on the substrate surface [93]. This property has been used to

fabricate both .positive. and .negative. patterns of polypyrrole using a procedure similar

to the adhesive tape test [94].

1.6.9 Self-Assembly Of Conducting Polymers

Ultra-thin organic films are currently gaining interest in many areas such as

integrated optics, biosensors, chemical sensors, friction reducing coating, surface

orientation layers and molecular electronics [43-44]. Most of the applications require

well-defined films composed of molecules with tailor-made properties in unique spatial

arrangements with respect to each other and to the substrates.

In-situ deposition, a recently emerged, extremely simple and yet very

sophisticated technique for fabricating multilayer thin films with precisely controlled

thickness and layer sequences, has many advantages over the other techniques for

polymer thin film deposition. This layer-by-layer molecule self-assembly process is

based on the spontaneous adsorption of polyions from dilute solutions on to surfaces that

carry a charge opposite to that of the depositing polymer. In essence, the excess charge of

a polyion adsorbed on to a substrate surface is used to attract a polyion of the opposite

charge on the surface. Multilayer thin films are fabricated by simply alternating the

dipping process in monomer solution and the dopant solutions. It has been shown that

this approach can also be used to manipulate electroactive polymers through the use of

conjugated polyions.

80

1.6.10 Methodology Of Self-Assembly

A well-cleaned substrate is immersed into a dilute aqueous solution of a cationic

polyelectrolyte for a period of time optimized for adsorption of a monolayer (~2nm

thick), then it is rinsed and dried [99]. The next step is the immersion of the

polyelectrolyte-monolayer-covered substrate into a dilute dispersion of negatively

charged nano-particles (or nano-platelets, or any other species with appropriate sizes and

charge distributions), also for a period of time optimized for adsorption of a mono-

particulate layer, followed by rinsing and drying. These operations complete the self-

assembly of a .sandwich unit. of a polyelectrolyte monolayer and a mono-particulate

layer of nano-particles onto the primed substrate.

The Self-Assembly is governed by a delicate balance between adsorption and

desorption equilibria. In the self-assembly of nano-particles, for example, the efficient

adsorption of one (and only one) mono-particulate layer of nano-particles onto the

oppositely charged substrate surface is the objective of immersion step. Desorption of

nano-particles forming a second layer and subsequent layers (and preventing the

desorption of the first added layer) is the purpose of the rinsing process. The optimization

of the self-assembly in terms of maximizing the adsorption of nano-particles from their

dispersions and minimizing their desorption upon rinsing requires the judicious selection

of stabilizer(s) and careful control of the kinetics of the process.

The most useful feature of self-assembly is that any oppositely charged species

(with the appropriate surface charge densities) can, in any order, be adsorbed layer-by-

layer. The oppositely charged species are held together by strong ionic bonds and form

long-lasting, uniform, and stable films that are often impervious to solvents. No special

81

film balance is required for the self-assembly; indeed, the method has been referred to as

.Molecular breaker epitaxy. [97]. Furthermore, self-assembly is economical (dilute

solutions and dispersions are used, and the materials can be recovered) and readily

amenable to scaling-up for the fabrication of large-are, defect-free devices on virtually

any shape and kind of surface [98].

The Molecular self-assembly process is driven by the attractions developed

between a positively charged, p-type-doped conducting polymer and a negatively charged

polyion. The positive charges of the conducting polymer are not created by the ionization

of permanently fixed side groups, but are rather in the form of partially delocalized defect

states that exist along the polymer backbone as a result of doping. Since the number of

these doping-induced defect states created along the backbone (polarons, bipolarons, etc.)

depends directly on the oxidation state of the conjugated polymer, they can be

systematically varied form none (neutral polymer, no multilayer assembly possible) to

about one for every three or four repeat units (highly doped polymer). The ability to

readily control the linear density of positive charges along the polymer backbone via

chemical doping provides an additional level of control over the polymer deposition

process [100]. This novel type of charge control is simply not possible with permanently

fixed ionic charges. It is interesting to know that such partially delocalized charge defect

states can be utilized to attract negatively charged polymers in a multilayer deposition

process.

1.6.11 In-Situ Deposition Of Polypyrrole Films

Among Conductive polymers, PPY is one of the extensively studied electronic

materials, and thus has received much attention because of various technological

82

applications. [101,102]. Though PPY receives considerable interest, processing it into an

ultra thin film is a challenging task because of its intractability, insolubility and

infusibility in most of the common organic solvents. PPY is usually synthesized by

electrochemical and chemical oxidative polymerization techniques. Its physical and

chemical properties are considerably dependent upon the dopants and polymerizing

conditions. Recent evidence shows that PPY can be self-assembled into multilayer thin

films with a suitable polyanion [103].

Using FeCl3 as the oxidant, there is a continuous formation of PPY because of

chemical oxidation of the monomer pyrrole. Due to the electrostatic interaction between

Fe+3 and polyelectrolytes, local perconcentration of both, pyrrole and Fe+3 are observed

which enhances the polymerization rate. The in-situ polymerized PPY molecules being

polycation are attracted towards polyanion chains via hydrophobic interaction resulting in

a higher local concentration. Polymerization can also be enhanced by the low

concentration of FeCl3 in the active solution [104].

The In-situ polymerization of the PPY film is a dynamic process where the

thickness of the PPY films initially increases linearly to the time, and then gradually

evidences to non-linear deposition after certain period of time. The remaining PPY that is

not deposited on the surface of the substrates precipitates gradually out of the active

solution [102].

1.6.12 Aging Effect In Polypyrrole Films

In any application, stability of the polymer films is a critical issue. Stability of

polypyrrole films has been studied and reported by many. Ageing problem is a major

drawback to its use in practical applications.

83

Understanding of the conductivity degradation with time in electronic conducting

polymers is very interesting. A unified dependence of the conductivity, , as a function of

the ageing time, ta, and of the temperature, T, is found to be:

Where, c (Ta) depends on the aging conditions. Equation (1.6.9) is only valid

under some conditions: the shorter the aging time, the lower the temperatures for which it

is valid. The variation of during aging at a given temperature, Ta, is thus a stretched

exponential for long aging times. After long aging times, the thermal variation of is

determined by a single parameter T0 that shows a linear dependence with ta [105]:

Thieblemont and Planche [106] studied the stability of chemically synthesized

polypyrrole films in detail. They adopted 2 methods to deposit the polypyrrole films on

various fabrics (woven glass, PVC and polyester). In the “one-step process” the substrate

was immersed in the polymerizing solution for certain time and removed. In the “two-

step process” the substrate was first dipped in the monomer solution followed by the

oxidant solution. The polypyrrole yield was more in the two-step process.

Ageing these PPY films in inert and ambient atmosphere was carried out at different

temperatures and surface resistivity was measured.

84

Figure 1.6.8 Effect of aging temperature (in air) on the conductivity of Glass/PPY (1-step

process) film at (1) 60°C; (2) 80°C; (3) 100°C; (4) 120°C; (1) 140°C; [106]

Figure 1.6.9 Effect of aging temperature (in air) on the conductivity of Glass/PPY (2-step

process) film at (1) 60°C; (2) 80°C; (3) 100°C; (4) 120°C; (1) 140°C; [106]

Thus, it was concluded that the one-step process leads to more stable products

than the two step one. On the other hand the nature of the substrate appears to have little

influence on the stability of the films.

The conductivity decrease of PPY doped with mineral anions in inert atmosphere

has been related to instability of these species with respect to the polymer chain

85

(reactions with polymer backbone). Thus it is assumed that the exclusive degradation

mechanism, underlying conductivity deterioration for doped PPY is an oxidation of the

polymer backbone by oxygen or water involving formation of carbonyl, ester and

hydroxyl groups [107-109].

The results of Truong et al. [110] seem to indicate that oxidation of the polymer

is a diffusion-controlled phenomenon. Assuming that the decrease of conductivity is

proportional to the amount of reactive species in the film, it is evident that the kinetics of

the conductivity decay is also controlled by diffusion and is temperature dependent,

described by the Arrhenius law in the range 60-140°C. This diffusion phenomenon

resembles that of solid state diffusion and could occur as a doping process involving an

exchange between oxygen and the doping species.

The electrical conductivity of PPY films on various substrates can be enhanced

by adding an extra sulfonate-doping agent in the polymerizing solution, owing to

preferential incorporation of this species in the polymer [106]. The stability in air of

these chemically synthesized films can be higher than those of films obtained by

electrochemical oxidation due to higher activation energy. Stability of the films depends

on the dopant species.

1.6.13 Bio-Medical Applications

Biomaterials are materials used for implants that replace or restore living tissues

and their functions. They can be made of metals, ceramics, polymers or composites. For

example, contact lenses are soft hydrogels made of polymers whilst bone substitutes are

hard bioceramic reinforced biopolymers. Applications are diverse and include artificial

86

assist devices (e.g. kidney dialysis membranes), tissue engineering, drug delivery and

gene therapy and artificial assist devices.

Inherently conductive polymer (or electronic polymer) technology is the most

innovative conductive polymer technology because it combines many of the desirable

properties of metals and plastics. Existing or potential advantages over metals or

inorganic semiconductors include higher electrical conductivity than other conductive

materials, plasticity and elasticity, low mass density, low heat conductivity, low

coefficient of expansion, resistance to chemicals and corrosion, anisotropic (axial)

molecular structure and conductivity, tunable optical properties (dependent on molecular

structure), electrochromism, and enhanced power storage and rechargeability (for

batteries). Applications of primary interest among developers include batteries and

capacitors, nonlinear optical materials (for use in telecommunications), photoconductors

for xerography (among several other optical applications), chip fabrication, sensors,

antistatic dissipation, display, anticorrosion, radio-frequency.interference shielding, and

Electromagnetic-interference shielding.

Conducting polymers are used to build actuators (muscles), transistors

(computations), wires (information transmission and energy delivery), super-capacitors

(energy storage) and strain gages (position sensing) thereby building most of the elements

needed for creating a bio-mimetic lifelike system [111].

Conductive polymers are potentially cheaper and can be bio-degradable, leading

to disposable electric motors. Polypyrrole based strain gages can undergo maximum

strains of several percent whereas silicon or metal film gages are limited to ~ 0.1 or 0.2%.

PPY gages are also used in applications that require bending or flexing of the gage itself.

87

In much the same way, the mammalian muscle spindles transduce position yet are as

flexible as the muscle fibers themselves. Artificial conducting polymer muscles exert

forces per cross-sectional area of up to 35 MPa, over 100 times the force per cross-

sectional area of skeletal muscle [111].

Polypyrrole based transistors doped with Chlorine ions are used in making

transistors, amplifiers and feedback loop circuits that are needed to create an intelligent

bio-mimetic system.

Conductive polymer super-capacitor could be used for energy storage for bio-

mimetic systems, either as a local source of energy within a muscle or as energy storage

for the bio-mimetic system as a whole.

Polymers have been processed to display permanent charges (electrets) or to

generate transient surface charges (piezoelectric materials) [112]. Studies using these

materials have demonstrated enhancement of nerve and bone cell growth in-vitro and in-

vivo. In contrast to the electrets and piezoelectric materials, Polypyrrole and

polythiophene generate electrical signals by electron transfer between different polymer

chains. In-vitro enhancement of nerve cell axonal extension has been reported using PPY

with application of either constant current or constant voltage. PPY has also been used as

a substrate to increase electronic interfacing between neurons and micro-machined

microelectrodes for potential applications in neural probes and prosthetic devices [113].

Electrically conducting polymers, unlike electrets and piezoelectric materials,

allow external control over the level and duration of stimulation, which is beneficial for

biomedical applications. Moreover, conductive polymers, in contrast to piezoelectric

materials, do not require extensive processing to render them electroactive. These

88

materials can also be modified by doping e.g. polypyrrole has been doped with biological

anions such as hyaluronan, which stimulates angiogenesis as it degrades, and adhesive

peptides, which enhance material-cell interactions. [112].

1.6.14 Medical Coating Characteristics

Medical device manufacturers have identified some primary material

characteristics to consider when selecting a medical coating. These include:

Biocompatibility: Coatings need to exhibit long-term compatibility and a non-reactive

relationship with body fluids and tissues. The coating should not undergo any chemical

interaction with the substrate to which it is being applied, nor should it produce any toxic

by-products or extracts that could be harmful to a patient or to the function of the item

being coated.

Coating Inertness: The product's coating must not contaminate the substrate with

outgassing or with by-products from catalysts, cure agents, solvents, or plasticizers.

Hydrophobic Characteristics: Hydrophobic.or hydrophilic.characteristics may be

important in certain medical applications.for example, a hydrophilic cardiovascular

catheter (slippery when wet) for ease of insertion versus a wet hydrophobic guidewire

that would be easy for the cardiologist to grip.

Cure Temperature: The cure temperature of a coating must be within the performance

range of the substrate.

Cure Forces: A coating's cure forces must not degrade or distort the underlying substrate.

Conformability: The coating must offer conformability to highly variable surface

geometries. It must provide effective isolation of all surfaces, including hidden areas,

crevices, etc., without bridging or pooling. The coating must be able to maintain

89

conformability at all magnitudes of substrate and surface feature sizes, from macro to

micro.

Finished Thickness: The finished thickness of a coating is important. The coating may

need to meet extremely tight dimensional tolerances, and therefore be quite thin, while at

the same time be able to provide uncompromised physical, chemical, or electrical

protection for the substrate with coverage that is free of voids and pinholes.

Mechanical Loading: A coating often needs to function dependably without significantly

altering the physical or mechanical properties of the substrate.

Resistance to Flaking: A coating needs to have considerable flaking resistance. It must be

sufficiently robust and adherent to prevent flaking from substrates or from itself.

Sterilizability: A coating must be capable of withstanding the effects of one or more

sterilization processes.

Polypyrrole, is one of a very few conducting polymers which when produced as a

thin film on appropriate substrates possesses all these characteristics and is a most suited

material for human implant devices.

90

1.7 Four Point Probe

1.7.1 Working methodology

The four-point probe technique is one of the most common methods for measuring the

semiconductor resistivity because two-point probe methods are more difficult to interpret.

Considering the two-point probe arrangement in Figure 1.7.1, each probe serves as a

current and as a voltage probe. The total resistance between the two probes is given by:

Where Rc is the contact resistance at each metal probe/semiconductor contact, Rsp is the

spreading resistance under each probe and Rs is the semiconductor resistance. The contact

resistance arises from the mechanical metal probe contacting the semiconductor. The

spreading resistance accounts for the resistance encountered by the current when it flows

from the small metal probe into the semiconductor. Neither Rc nor Rsp can be accurately

calculated so that Rs cannot be accurately extracted from the measured resistance.

Figure 1.7.1 2-point probe showing the contact resistance Rc the spreading resistance Rsp

and the semiconductor resistance Rs. [116]

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A solution to this dilemma is the use of four probes. Two probes carry current and

the other two probes are used for voltage sensing. The four-probe method was originally

proposed by Wenner in 1916 to measure the earth.s resistivity and the technique is

referred to in Geophysics as Wenner’s method. Valdes adopted it for semiconductor

wafer resistivity measurements in 1954. The probes are generally arranged in-line with

equal probe spacing, though other probe configurations are possible.

The use of four probes has an important advantage over two probes. Although the

two current-carrying probes still have contact and spreading resistance associated with

them, which is not true for the two voltage probes because the voltage is measured either

with a potentiometer which draws no current at all or with a high impedance voltmeter

which draws very little current. The two parasitic resistances Rc and Rsp are negligible in

either case because the voltage drops across them are negligibly small due to the very

small current that glows through them.

The potential ‘V’ at a distance ‘r’ from an electrode carrying a current ‘I’ in a

material of resistivity ‘ρ’, is given by [114]:

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Figure 1.7.2 A collinear Four point Probe. [116]

4 probes resting on a semi-infinite medium as in Figure 1.7.2, with current entering probe

1 and leaving probe 4, the voltage V becomes:

where r1 and r4 are the distances from probes 1 and 4, respectively. For probe spacing of

s1, s2, and s3, as shown in the Fig (1.2), the voltage at probe 2 is

and the voltage at probe 3 is

93

The total measured voltage V = V2 - V3 becomes

The semiconductor parameter of interest is the resistivity:

usually in the units of ohm-cm, with V measured in volts and I measured in amperes. For

most four-point probes the probe spacings are equal with s = s1 = s2 = s3 and the

Equation (1.7.7) reduces to:

Optimum probe spacing is on the order of 0.5 to 1.5 mm. The two most common probe

spacing is 0.635 mm (25mils) and 1.588mm.

1.7.2 Correction Factor

Semiconductor wafers are not semi-infinite in extent in either the lateral or the vertical

dimension. Equation (1.7.8) must be corrected for finite geometries. For an arbitrarily

shaped sample the resistivity is given by

where F is the correction factor, which depends on the sample geometry. F corrects for

edge effects, for thickness effects, and for probe placement effects, and it is usually a

product of several independent correction factors. For sample thicknesses greater than the

94

probe spacing, the simple, independent correction factors contained in F of Equation

(1.7.9) are no longer adequate due to interactions between thickness and edge effects.

Fortunately the sample thickness is generally smaller than the probe spacings, and the

correction factors can be independently calculated.

The following correction factors are for collinear or in-line probes with equal

probe spacing, s. The correction factor F is given as a product of three separate correction

factors:

F1 corrects for sample thickness, F2 corrects for lateral sample dimensions, and F3

corrects for placement of the probes relative to the sample edges. A parameter that must

be corrected for most practical measurement conditions is the sample thickness since

semiconductor wafers are not infinitely thick. Their thicknesses are usually on the order

of the probe spacing or less [115]. Introducing the correction factor for a non-conducting

bottom wafer surface boundary, where t is the wafer or layer thickness:

For a conducting bottom surface the correction factor becomes:

For thin samples, the correction factor becomes:

95

Equation (1.7.13) is valid for t ≤ s/2. For very thin samples for which F2 = F3 = 1, we

find from Equations (1.7.9), (1.7.10) and (1.7.13):

96

1.8 Acoustic Studies

1.8.1 Basic Acoustic Phenomenon And Equation

Sound is comprised of propagation of small mechanical perturbations in a

medium. The basic mathematical description of acoustic phenomena is obtained from

linearized fluid mechanics equations. There are three distinctly different approaches to

the study of acoustics: wave acoustics, ray acoustics and energy acoustics. The wave

acoustics approach is the most widely used and will be discussed here. (The ray and

energy acoustics approaches are used when the wave approach becomes too

complicated.) Acoustic waves are generally three dimensional (two dimensional waves

are seen on surfaces). As the wave moves away from its source, the curvature of the

wavefront decreases. At large distances from the source of sound therefore, the wavefront

may be regarded as a plane. This makes possible the description of what are called plane

waves (simplest mathematical formulation). Plane waves may be defined as waves that

have the same acoustic properties at any position on a plane perpendicular to the direction

of propagation of the wave [117].

If a disturbance in a thin element of fluid is considered, a mathematical

description can be written by making the assumptions that

1. the amount of fluid in the element is conserved,

2. the net longitudinal force is balanced by the inertia of the fluid in the element,

3. the process in the element is adiabatic, i.e., there is no flow of heat in or out of the

element, and

4. the undisturbed fluid is stationary (no flow) [117]

97

This equation is known as the one-dimensional equation of motion or the acoustic

wave equation. It relates the second rate of change in the pressure (p) with distance (x)

with the second rate of change of the pressure with time (t) through the square of the

speed of sound (c). The solution of Equation 1.8.1 takes the form [117].

where, f1 and f2 are arbitrary functions such as sine, cosine, exponential, log, etc. The first

term on the right in Equation 1.8.2 represents a wave traveling in the positive x-direction

and the second term represents a wave traveling in the negative x-direction (both

traveling with the speed c). Waves created by sound sources vibrating sinusoidally in

time (e.g., loudspeaker, piston) with angular frequency vary both in time and space in a

sinusoidal manner [117].

where p1 and p2 are amplitudes of waves traveling in the positive and negative x-

directions, k is the acoustic wave number, k = /c, and 1 and 2 are phase angles. This

equation means that:

1) At any point in space, x, the sound pressure is sinusoidal (simple harmonic) with

respect to time, and goes through one complete cycle when increases by 2. The

time required for a cycle is called the time period, T. Thus, T = 2; T = 2/;

98

2) At any instant t the sound pressure pattern is sinusoidal in space, and repeats itself

each time kx is increased by 2. Such a repetition is called a wavelength, .

Thus, k = 2; k = 2/. Two more important relationships emerge from the above:

where f is the frequency (= /2), and

Inside the medium sound propagation occurs by particle vibration in the direction

of the propagation of the wave. Such waves are known as longitudinal or compressional

waves. The wave moves at a constant speed c which depends upon the properties of the

medium and is independent of the frequency f. The particles of the medium oscillate back

and forth with a velocity, u, called the particle velocity. For a plane wave traveling in the

positive x-direction the following relationship holds [117].

where is the density of the medium. Similarly, for a plane wave traveling in the

negative x-direction

The quantity ‘c’ is called the acoustic impedance or the characteristic impedance

of the medium and is denoted by the symbol Z (units: kg/m2.sec). In practice, sound often

travels from one medium to another. As a result of acoustic impedance mismatch

99

between any two media, the sound transmission at the interface is not 100%; some

amount of sound is reflected, and some amount is transmitted. The percentage of sound

reflected depends upon the difference in acoustic impedance between the two media, and

is quantified by the reflectivity coefficient (R) [121].

This equation is valid for plane waves Z1 and Z2 are the acoustic impedances in

the two media. Table 1.4.1 lists the acoustic impedance of some common media [4, 122].

The sound intensity (I) is the rate at which the sound wave does work on a surface of unit

area in a direction perpendicular to the surface [117].

where p is the instantaneous pressure and p1 is the amplitude/maximum pressure of the

wave

Since the ranges of sound pressures usually encountered encompass over eight

orders of magnitude, logarithmic scales are used to quantify sound pressures. The units

used are usually decibels. Decibels are logarithmic of ratios of quantities which have the

same units. The most common decibel scale used in specifying sound pressures is the

Sound Pressure Level (SPL) scale, which uses a reference pressure of 20 µPascals

(approximate human threshold of hearing). Table 1.8.2 lists some typical sound pressure

levels [117]. As can be seen from this table, typical conversation produces pressures of

approximately 50 dB SPL. It must be remembered from the earlier discussions on human

ear, that the pressure levels present inside the cochlea, however, are far greater than what

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is incident on the outer ear (Table 1.8.2). The outer ear, middle ear and other structures of

the head amplify incident pressures ~ 20 - 30 dB SPL.

Table 1.8.1 Acoustic impedance of some common media [117]

101

Table 1.8.2 Some typical sound pressure levels in the dB SPL scale [117]

1.8.2 Air Measurements: Semi-Quantitative Measurements

Semi-quantitative acoustic measurements in air were performed to determine

primarily:

1. To check the integrity of all electrical connections

2. To obtain a semi-quantitative idea of the device sensitivity and parameters

affecting it.

1.8.3 Underwater measurements: Quantitative

The perilymph fluid environment of the scala tympani necessitates optimization

of the underwater (perilymph is a saline solution) sensitivity of a piezoelectric implant.

Moreover, as a result of the complicated geometry and structure of the cochlea, the

acoustics and micromechanics active therein are not well understood yet. Therefore, the

need for an underwater measurement in which the acoustics were relatively better

understood.

102

1.8.4 Comparison Of Air Acoustics With Underwater Acoustics

From Table 1.8.1, it is seen that the speed of sound in air is 344 m/s and the speed

of sound in water is 1500 m/s. The acoustic impedance of air is 0.000444 x 106 kg/m2.-

sec and that of water is 1.5 x 106 kg/m2-sec. Using this data and Equation 1.8.9,

comparisons can be made between the two media. For the same acoustic pressure (p) the

intensity of sound (rate of energy flow) in air and water can be compared [123].

This means that the same acoustic pressure corresponds to intensity in air, 3378 times

greater than the corresponding intensity in water. If, instead, the intensity is same in the

two media, the corresponding sound pressure levels can also be compared.

This means that if the intensity of sound is same in the two media, the pressure is about

58 times greater in water.

1.8.5 Speed Of Sound In Water

The speed of sound in any medium is governed by the density of the medium and

its adiabatic bulk modulus. From an acoustic perspective, the structure of liquids is

intermediate between gases and solids. A liquid is gas-like in the sense that it does not

offer resistance to shear stress and solid-like in the sense that its bulk modulus depends

103

upon intermolecular bonding forces and not on external factors like gravity, etc., as is the

case for gases [119]. Information on the structure of a liquid can be obtained from the

radial distribution function, obtained by x-ray and neutron diffraction. These type studies

have shown that liquids can contain regions of highly ordered structure; these regions are

called aggregates or clusters. In addition, there is an abundance of micro cavities, called

holes, of varying size and distribution, which are responsible for the fluidity of the liquid.

Some liquids do not show a tendency to form clusters; these are known as unassociated

liquids. Some other liquids, like water, show a substantial tendency to form clusters;

these are called associated liquids [119]. Water is considered to be a mixture of open

structures and closely-packed structures. As the temperature is increased towards the

boiling point, the loosely packed regions gain at the expense of the more tightly packed.

The competition between structural redistribution processes and thermal expansion is

responsible for such extraordinary properties of water as the density maximum at 4oC

[119]. Due to the complexities of structure, theoretical derivations of the speed of sound

in liquids have had very limited success. Instead, accurately measured empirical

relationships are used to describe the speed of sound. For water, the most prevalent such

expression is [118,119]

where c is speed of sound in meters/sec, Tc, a and the coefficients are given in Table 1.8.3

At 20oC this equation yields c = 1481.8 m/s. If plotted, it is found that the speed of sound

is maximum at 74oC, another effect of the extraordinary structure of water [124].

104

Table 1.8.3 Parameters entering formula for temperature dependence of sound speed in pure water [118,119]

1.8.6 Acoustic Sensitivity Calculation

The ranges of some quantities dealt with in engineering and medical fields vary

over several orders of magnitude. The decibel (dB) is a convenient logarithmic ratio

between two numerical values having the same units [120]. The decibel is defined as:

where x1 and x2 represent numerical values of variables such as Power, Energy, and

Intensity.

These variables are proportional to the square of variables such as Pressure,

Voltage and Current. Therefore, when numerical values of the latter, represented by y1

and y2, are used to determine the dB, the following equation is used [120]:

since, Decibels are relative quantities and do not provide information regarding

the absolute values of the variables involved, unless one of the values, by convention the

105

denominator, is a standard quantity which defines a particular scale [120]. In the

following, two-decibel scales, one commonly used to quantify sound pressures, the other

commonly used to quantify transducer sensitivity, are discussed.

1.8.7 dB SPL Scale Sound Pressure Scale

The dB Sound Pressure Level (SPL) scale, used to quantify sound pressures, is

based on a reference pressure of 20 µPascals (approximate human threshold of hearing).

For example, sound of intensity 50 dB SPL corresponds to a pressure of (X Pascals): 50 =

20 log (X/20.10-6). Therefore, X = 20.10-6. 102:5 = 6325 Pascals. Normal conversation

produces sound intensities of ~ 50 dB SPL at 1 meter distance [11]. Sound intensities

greater than ~ 85 dB SPL are considered harmful.

1.8.8 dB (ref. 1 Vrms/1µPa) Transducer Sensitivity Scale

The sensitivity of transducers is measured in volts/unit pressure. Since transducer

sensitivities can also vary over several orders of magnitude, decibel scales are used. The

most commonly used such scale specifies the sensitivity of a transducer in comparison to

an ideal transducer, which produces a voltage of 1 voltrms when a pressure of 1 µPascal is

incident on it. Such a transducer is thus highly sensitive, and most engineering

transducers (microphones, hydrophones) have sensitivities far inferior to it. Therefore, in

this scale, the sensitivities of most engineering transducers falls within the -150 to -250

dB range. In the following, the calculations used to convert the voltage measured from a

transducer to its dB sensitivity are presented. Suppose that a pressure of X Pascals is

incident on the transducer, and the transducer produces a voltage Vpeak to peak in response to

this pressure. The peak-peak voltage is first converted to the root mean square (rms)

value by the relation:

106

Thus, the transducer produces Vrms volts in response to X Pascals. The root mean square

voltage produced by the same transducer when the pressure is 1 µPascal is calculated by

the relation:

This simple proportionality equation assumes that the transducer is linear in this range of

pressures, and therefore must be used with caution. The sensitivity of the transducer can

now be calculated in the dB scale:

107

1.9 Electrode-Substrate Contacts

Semiconductor devices usually contain at least one junction between p-type and

n-type material. These p-n junctions are fundamental to the performance of functions

such as rectifications, amplification, switching, and other operations in electronic circuits.

The junction contacts have been discussed in this chapter.

1.9.1 Drift of Carriers in Electric Field

Knowledge of carrier concentrations in a conducting material is necessary for

understanding the current flow in the presence of electric or magnetic fields. The

collisions of the charge carries with the lattice and with the impurities are also important

factors to be considered. These processes affect the ease with which electrons and holes

can flow through the crystal, that is, their mobility within the material. These collisions

and scattering processes depend on temperature, which affects the thermal motion of the

lattice atoms and the velocity of the carriers [131].

The electron mobility µn, can be expressed as the average particle drift velocity

(vx) per unit electric field (x):

The units of mobility are (cm/sec)/ (V/cm) = cm2/V-sec. The minus sign in the definition

results in a positive value of mobility, since electrons drift opposite to the field.

The current density can be written in terms of mobility as:

This expression is based on the assumption that the current is carried primarily by

electrons. For hole conduction we can change n to p, -q to +q, and µn to µn, where µp =

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+<Vx>/x is the mobility of the holes. If both electrons and holes participate, we must

modify the equation 1.9.2 to [131]:

If we consider the semiconductor bar of Figure 1.9.1 and assume that it contains both

types of carriers, equation 1.9.3 gives the conductivity of the material. The resistance of

the bar is then given by:

Figure 1.9.1 Drift of electrons and holes in a semiconductor bar [130]

where is the resistivity (Ω-cm). The physical mechanism of carrier drift requires that

the holes in the bar move as a group in the direction of the electric field and that the

electrons move as a group in the opposite direction. Both the electron and the hole

components of the current are the in the direction of field, since conventional current is

positive in the direction of hole flow and opposite to the direction of electron flow. The

drift current described in equation 1.9.3 is constant throughout the bar. The contacts to

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this kind of semi-conducting bar are always ‘Ohmic’, which means that they are perfect

sources and sinks of both carrier types and have no special tendency to inject or collect

either electrons or holes [130].

1.9.2 Schottky Barriers

Many of the useful properties of a p-n junction can be achieved by simply

forming an appropriate metal-semiconductor contact. This approach has obviously been

attractive because of its simplicity of fabrication. Metal-semiconductor junctions are

particularly useful when high-speed rectification is required. The non-rectifying or the

Ohmic contacts are easy to fabricate too.

It may be stated here that the minimum energy required for an electron to escape

from the metal or the conductor into a vacuum is represented by q∅. The energy q∅ is

called the work function of the metal or the conducting material. It can also be defined as

the minimum energy required for an electron to be removed from the Fermi level to the

vacuum outside the material [132]. Whenever, negative charges are brought near the

conducting material surface, positive (image) charges are induced in the metal. When this

image force is combined with an applied electric field, the effective work function is

somewhat reduced. Such barrier lowering is called the Schottky effect. Although

Schottky effect is only a part of the explanation of metal-semiconductor effects,

rectifying contacts are generally referred to as Schottky barrier diodes.

Considering barriers in ideal metal-semiconductor junctions:

When a metal with work function q∅m is brought in contact with a semiconductor having

a work function q∅S, charge transfer occurs until the Fermi levels align at equilibrium, as

shown in figure 1.9.2.

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Figure 1.9.2 A Schottky barrier formed by contacting a semiconductor with a metal

having a larger work function [133]

When ∅m > ∅S, the semiconductor Fermi lever is initially higher than that of the

metal before contact is made. To align the two Fermi levels, the electrostatic potential of

the semiconductor must be raised (i.e., the electron energies must be lowered) relative to

that of the metal. In the n-type semiconductor (figure 1.9.2), a depletion region is formed

near the junction. The positive charge due to the uncompensated donor ions within the

depleted region matches the negative charge on the metal [130].

The equilibrium contact potential Vb which prevents further net electron

diffusion from the semiconductor conduction band into the metal, is the difference in

work function potential ∅m - ∅S. The potential barrier height ∅b for electron injection

from the metal into the semiconductor conduction band is ∅m - XS, where qXS (called the

electron affinity) is measured from vacuum level to the semiconductor conduction band

edge. The potential difference Vb can be decreased or increased by the application of

either forward or reverse-bias voltage.

111

1.9.3 Rectifying Contacts

When a forward-bias voltage V is applied to the Schottky barrier of figure 1.9.1,

the contact potential is reduced from Vb to Vb - V as shown in figure 1.9.3-a. As a result,

electron in the semiconductor conduction band can diffuse across the depletion region to

the metal [130]. This gives rise to a forward current (metal to semiconductor) through the

junction. Conversely, a reverse bias increases the barrier to Vb + Vr, and electron flow

from semiconductor to metal becomes negligible (figure 1.9.3-b). In either case flow of

electrons from the metal to the semiconductor is retarded by the barrier ∅m - XS.

Figure 1.9.3 Effects of forward and reverse bias on the ideal metal-semiconductor junction. (a) forward bias; (b) reverse bias; (c) typical I-V characteristics [130]

One important feature that can be predicted is that the saturation current should

depend upon the size of the barrier ∅b for electron injection from the metal into the

semiconductor. This barrier (which is ∅m - XS for the ideal case) is unaffected by the bias

112

voltage. With similar theory, it can be studied that in both the cases, i.e., metal-n-type

semiconductor and metal p-type semiconductor, the Schottky barrier diode is rectifying,

with easy current flow in the forward direction and little current in the reverse direction.

The forward current in each case is due to the injection of majority carriers from the

semiconductor into the metal. The absence of minority carrier injection and the

associated storage delay time is an important feature of Schottky barrier diodes.

1.9.4 Ohmic Contacts

In many cases it is desired to have an Ohmic contact between metal and

semiconductor. Ohmic contacts show linear I-V characteristics in both biasing directions.

These contacts have minimal resistance and no tendency to rectify signals.

Ideal metal-semiconductor contacts are Ohmic when the charge induced in the

semiconductor in aligning the Fermi levels is provided by majority carriers (figure 1.9.4).

For example, in the ∅m <∅S (n-type) case of figure 1.9.4-a, the Fermi levels are aligned

at equilibrium by transferring electrons from the metal to the semiconductor. This raises

the semiconductor electron energies (reduces the electrostatic potential) relative to the

meal at equilibrium (figure 1.9.4-b). In this case the barrier to electron flow between the

metal and the semiconductor is small and easily overcome by a small voltage [130].

Similarly, the case ∅m > ∅S (p-type) results in easy hole flow across the junction (figure

1.9.4-d).

113

Figure 1.9.4 Ohmic metal-semiconductor contacts: (a) ∅m < ∅S for an n-type

semiconductor, and (b) the equilibrium band diagram for the junction; (c) ∅m > ∅S for a p-type semiconductor, and (d) the junction at equilibrium [130 ]

Unlike the rectifying contacts discussed above, no depletion region occurs in the

semiconductor in these cases since the electrostatic potential difference required to align

the Fermi levels at equilibrium calls for accumulation of majority carriers in the

114

semiconductor. Figure 1.9.5 shows the difference between the rectifying and the Ohmic

contacts.

Figure 1.9.5 Difference between the rectifying and Ohmic contacts [133]

Figure 1.9.6 differentiates between the rectifying and Ohmic contacts with respect

to their I-V characteristics. It can be seen that because of negligible depletion region in

the case of Ohmic contacts electron flow through this region, we see a linear I-V curve.

Whereas in the case of Schottky barriers, electrons flow through (over) the barrier which

produces a non-linear I-V curve.

115

Figure 1.9.6 I-V characteristics of rectifying and Ohmic contacts [133]

116

Chapter 2

Experimental

117

2.1 Experimental Procedures

2.1.1 Materials and Substrates

Silanol-terminated PDMS (polydimethyl siloxane) of molecular weight ranging

from 20,000 to 30,000 (DMS-S31) and 80,000 to 130,000 (DMS- S45), the cross-linker

TEOS (tetraethoxysilane) and the catalyst STO (tin octate) were obtained from Gelest,

Inc.

Stretched, PVF2 polymer piezoelectric films with Nickel-Copper electrodes

(28µm, 52µm thick) and non- electroded films were obtained from Measurement

Specialties Inc., USA.

Pyrrole (98%), Anthraquninoe-2-sulfonic acid, sodium salt monohydrate (97%),

FeCl3.6H2O(98%), 5-Sulfosalicylic acid dehydrate (99%+), Octadecyltrichlorosilane

(OTS, 95%) were obtained from Aldrich.

Silver paste plus was obtained from SPI Inc. A UV curing gel was used as a mask

obtaind from Dymax. All chemicals were used as received. The final base used was

Lecoset. It is Leco’s newest rapid cold curing plastic.

118

2.1.2 Device Preparation

Polymer Base

The first aspect was to choose and optimize a suitable substrate material. The

principal requirements of the substrate material are:

1. Flexibility, as offered by elastomeric polymer;

2. Biocompatibility and negligible swelling with the perilymph and the structures of

the inner ear.

In addition, the substrate material needs to be amenable to easy fabrication and there also

needs to be a way to permanently embed the PVDF elements in the substrate relatively

easily.

Based on these requirements, the choice of the substrate material fell on the

inorganic elastomer poly-(dimethylsiloxane) (PDMS). Being an elastomer, PDMS

possesses rubber-like flexibility even after being cross-linked. The biocompatibility of

PDMS is already proven; the material is used in cosmetic implants. Additionally, the

material, which is a liquid (viscosity is a function of molecular weight), can be cross-

linked to form a solid relatively easily after being cast into the desired shape. The cross-

linking reaction can also be completed at room temperature within a suitable length of

time.

The amount of TEOS and STO used for x gms of PDMS was:

That is, TEOS and PDMS were added in stoichiometric amounts. The nominal amount of

catalyst added was 0.8 wt.% of this mix.

119

The actual amount of catalyst used was varied in an effort to obtain room

temperature cross linking, overnight. Many trials were performed with PDMS of various

molecular weights.

Overnight cross-linking time and best mechanical properties were obtained from

PDMS of molecular weight 20,000 - 30,000 (Gelest, Inc., product DMS-S31), and 2 wt.%

STO.

Lecoset Base

Lecoset is Leco’s newest rapid cold curing plastic. The liquid (resin monomer)

and powder (resin polymer) components are mixed easily in small batch sizes. It is then

poured in a mould with piezoelectric filaments and cured in air for 10 to 12 minutes.

After curing it forms an opaque base.

Device Preparation

The PVDF elements were cut into strips of approximate size of 20 mm x 2mm

and 20 mm x 1 mm. Strips were first embedded in a base of putty/modeling clay. After

the mold was filled with the liquid mixture, this arrangement was inverted into the mold

and the elements were thus suspended in the cast liquid and held stationary until the

cross-linking/gelling reaction was complete, whereupon the elements were anchored

securely in the now solid base. The gelling reaction was usually complete within 24 hrs

for the elastomeric substrate and about 2-3 hrs for the lecoset substrate. The putty support

was then removed from the elements, and the device was removed from the mold

carefully to prevent accidental tearing. No mold-releasing agents were required to

120

accomplish this task, as the gelling process results in shrinkage of the substrate enabling

relatively easy removal. The process flow chart is as shown in the figure below.

Figure 2.1 Process Schematic of Piezoelectric sensor devices [124]

The devices are made of a size somewhat larger than can be accommodated in the

cochlea, ≈ 2X of implantable size. As the principle thrust of this work has been to

establish feasibilities of the underlying science, it was thought that successful fabrication

of devices of this size was adequate to establish the feasibility of fabrication of a smaller

device suitable for implantation.

121

Some of the charge-discharging schemes demand that the PVDF elements be

connected in series or in parallel electrically. For this purpose two types of interconnects

were applied:

1. Silver paste was painted on the substrate and connected to the electrodes of the

PVF2 elements. The silver dries fully in approximately 24 - 48 hours at room

temperature.

2. Conducting polymer was deposited on the device masked with Dymax UV curing

gel.

High temperature curing for both the PDMS substrate as well as the silver is not

recommended, as PVDF begins to lose the piezoelectric property above 60oC. Important

observations from the silver painting process were:

1. The viscosity of the silver paste being applied needs to be closely controlled,

since a low viscosity paste easily migrates around the polymer film and short-

circuits the electrodes, and a high viscosity paste precludes thin conductive

coatings;

2. The solids loading of the silver paste affects the conductivity of the silver

connections. Highest conductivities (< 1 ohm) were easily obtained by using SPI,

Inc. high purity Silver Paste Plus (higher solids loading than conventional

microelectronic silver pastes);

3. The silver cannot be cured at high temperatures, the resulting interconnects are

brittle and develop cracks if the substrate is bent/stretched.

Deposition of conducting polymer on the elastomeric PDMS substrate was of a

very poor quality therefore Lecoset substrate was used.

122

Conducting Polymer Preparation

Monomer Solution

A solution was made containing 0.6-ml (0.009M) of pyrrole dissolved in 100 ml

of magnetically stirred deionized water.

Oxidant Solution

3.5g (0.02M) of ferric chloride was dissolved using magnetic stirring in 100 ml of

deionized water in a 250 ml beaker. The solution was yellow in color. After 5 minute of

stirring at room temperature, dopant acid, Anthraquninoe-2-sulfonic acid, sodium salt

monohydrate (0.003M) was added to the 100 ml of ferric chloride solution. This solution

was then magnetically stirred for 10 more minutes before adding 5.34 g of 5-

sulfosalicylic acid sodium salt (0.02M). The color of the solution changed from yellow to

purple.

Conducting Polymer Deposition

The prepared devices were suspended above the ferric chloride/dopant acid

mixture in 250 ml beaker using plastic clamps. A pyrrole solution was then poured

immediately into the beaker containing the ferric chloride and the acid. Device was

immersed completely in the polymerizing solution. To obtain conducting polymer films

of desirable quality and thickness, the device, was removed from the polymerizing

solution and rinsed with deionized water for ~10 seconds, several times during the

deposition process.

123

Figure 2.2 Schematic showing in-situ Polymer depositions

124

2.1.3 Experimental Setup

In-air Acoustic Measurements

Figure 2.3 Schematic showing Acoustic tests in air

These measurements were performed by speaking the word “hello" in the author’s

voice, into the device, from a distance of ≈ 3-4 inches. The device is connected to a

preamp. The response from the preamp is then feed to the Data Acquisition Unit. Here

the obtained response is measured and analyzed. Acoustic measurements in air were

performed to determine primarily:

1. The integrity of all electrical connections

2. To obtain the voltage responses of the prepared devices.

Underwater Acoustic Measurements

The perilymph fluid environment of the scala tympani necessitates optimization

of the underwater (perilymph is a saline solution) sensitivity of a piezoelectric implant.

Moreover, as a result of the complicated geometry and structure of the cochlea, the

acoustics and micromechanics active therein are not well understood yet. Therefore, the

need for an underwater measurement system in which the acoustics were relatively better

understood as in-air acoustic measurements were a semi-quantative method of predicting

the response of a device in the final environment. It is important to mention that all

underwater measurements were performed at this time, in tap water and not saline

125

solution. As tap water also contains dissolved salts, it is not expected that device response

in saline/Perilymph will be significantly different than the response measured in tap

water.

Figure 2.4 Picture of Water Tank and other instruments

Since this thesis deals with only the semi-quantative responses of the devices the

underwater acoustic measurement procedures have not been explained in detail.

126

2.2 In-air Acoustics and Errors in Measurements & Voltage Response Plots

The acoustic measurements taken from a device were an average of ten iterations.

Shown below are the voltage responses of a 52µm all polymer device with 2 elements

connected in parallel. Since the response in the table is an average therefore when

repeated will not give the same result.

Table 2.1 Voltage response of an all polymer 2 element device connected in parallel

Parallel Voltage (Vp volts) No. of

elements

Length of

Peizo elements

(mm)

2mm

(Element Width)

1mm

(Element Width)

10 mm 5.4 4.0 2

05 mm 2.1 3.0

Table 2.2 shows us the voltage response iterations of a 2 element all polymer

device. As seen from the table the response of the same device is not necessarily

repeatable. But is ± 0.5 from the average value noted. This is primarily because the in-air

acoustic test requires the author to perform the “Hello test”. Each of these tests had

variations in it causing the response to vary accordingly. Also the test was performed

from the distance of about 3-4 inches. A change in this distance also led to some of the

variations mentioned.

This type of error mentioned is completely eliminated while performing

underwater acoustic. For underwater tests tones are used which reduces the errors causes

because of the “Hello test”. The distance between the source of sound and device is also

kept constant mechanically eliminating any human errors. These factors help get an

accurate response as compare to in-air acoustics.

127

Table 2.2 Ten voltage response iterations of an all polymer 2 element device connected in

parallel

Parallel Voltage (Vp volts) No. of

elements

Iteration

Number

Length of

Peizo

elements

(mm)

2mm

(Element Width)

1mm

(Element Width)

Iteration 01 5.3 3.6 Iteration 02 5.1 4.5 Iteration 03 4.9 3.9 Iteration 04 5.0 3.7 Iteration 05 5.4 4.4 Iteration 06 5.7 4 Iteration 07 5.9 3.5 Iteration 08 5.8 4.4 Iteration 09 5.6 4 Iteration 10 5.5 4

2

Average

10 mm

5.4 4.0

Figure 2.5 Voltage response plot of a single element device

128

Figure 2.6 Voltage response plot of a 2 element device

Factors causing error in voltage response also causes the voltage response spectra

to differ for each acoustic measurement for each device. There is also a percentage of

human error in these plots. There is a time lag between the Hello test and the actual

measurement of the plot which causes the spectra to start from any arbitrary value and

not zero.

The factor noise is does not effect semi-quantative in-air acoustic measurements.

This is because we get in-air acoustic measurements from the hello test. The interference

of noise is practically zero for this test. But it is an important factor for underwater

acoustics as electrical interference is the main source of noise in this case.

129

Chapter 3

Results and Discussion

130

3.1 Polymer Substrate Characteristics

Solubility Test

PDMS (Silanol Terminated PolyDimethyl Siloxane) was the initial substrate

choice for cochlear implants. Its long term effect on the weight of the base material in

0.9% NaCl solution was seen for a period of over 1 year. 0.9% NaCl was chosen as the

environment because scala tympani and scala vestibuli in the inner ear where our actual

cochlear implant will be placed is filled with perilymph fluid which is approximately a

0.9% NaCl solution.

Table 3.1.1 Solubility test of PDMS

No. of readings Date Weight of PDMS

(gms)

Initial Reading 4th april’02 0.093

1 1st july’02 0.093

2 1st october’02 0.088

3 1st january’03 0.086

4 1st april’03 0.085

5 1st july’03 0.084

From the table above it was seen that the PDMS substrate has an extremely low

solubility in 0.9% NaCl solution. The substrate also shows no swelling. From this

experiment it can be concluded that PDMS being physically stable in the chosen

environment is a good initial choice as a substrate material. Therefore the device once

implanted in the inner ear will not disintegrate in a couple of years. But the final in-air

acoustic tests were not done with PDMS as it wasn’t the ideal substrate for deposition of

131

silver and polypyrrole. This is because both the materials had a very weak adherence to

PDMS making it a bad choice of substrate for the final all polymer device. Therefore the

physical properties of our final substrate material can be summarized as follows.

1. Desktop gelling since PVDF losses its piezoelectric properties if heated.

2. Gelling time should not exceed more than 24 hours.

3. Zero to low solubility.

4. No swelling in 0.9% saline medium.

5. It should be biocompatible.

6. Should be a good substrate material for polymer deposition as interconnect

material.

It is critical for our cochlear substrate to fulfill all the above properties. This is

because the final device should be easily manufactures and also not disintegrate once

implanted.

DSC (differential scanning calorimetry)

From the DSC measurements, the real-time signals, plot, and method progress of

our sample (PDMS) are analyzed. Differential scanning calorimetry is a technique used

to study the reaction of the material when heated. In this case the material used was

PDMS. This technique was used to study the thermal transitions of the material. The

normal range for human body temperature is 97 to 100 degrees Fahrenheit or 36.1 to

37.8 degrees Celsius. This is the working temperature range of the cochlear implants.

From figure 3.1.1 no thermal transitions were observed in the internal temperature range

132

of a human body. This shows that PDMS is thermally stable when implanted in the inner

ear.

Figure 3.1.1 DSC of PDMS

From the Solubility tests and DSC plots, the physical and thermal stability of

PDMS when used as a substrate material for cochlear implants was seen.

133

3.2 Four Probe Test

A four probe test was performed to find out and compare the resistance and

resistivity of silver paste and conducting polymer (polypyrrole) as an interconnect

between filaments and as an electrode on the filaments.

Table 3.2.1 Four probe test

Conducting

Polymer on

PDMS

(1.5 cms wide)

Conducting

Polymer on

PDMS

(2mm wide)

High purity

silver paste

(2mm wide)

Diluted High

purity silver

paste

(2mm wide)

4 probe result

(volts)

0.83 0.98 0.08 0.85

Resistance

(ohms)

375.99 443.94 40.31 385.05

Thickness of

film (meters)

6 × 10-8 6 × 10-8 25 × 10-8 29 × 10-8

Resistivity

(ohmmeters)

0.23 × 10-4 0.27 × 10-4 0.10 × 10-4 1.12 × 10-4

From the table 3.2.1 it is shown that high purity silver paste offers the least

amount of resistance. But when dilute the resistivity value decreases by a whole order of

magnitude. Diluted silver paste had to be used as interconnects for the experiment as in

its pure form it would not to stick to the substrate material. But when high purity silver

paste is used in its diluted form its resistivity value is lower than that of the conducting

polymer (0.27 × 10-4-m). As seen from the table, vast differences in the thickness of

the two types of interconnect materials used is observed. The deposited conducting

polymer has a thickness of 6µm as compared to 29µm for the high purity silver paste. For

134

a thin film resistivity is directly proportional to thickness of the film. Therefore,

resistivity offered by the conducting polymer now becomes comparable to high purity

silver paste as an interconnect material.

The resistivity of conducting polymer as an interconnect can be further reduced,

thereby increasing charge flow and conductivity by decreasing the thickness of the

conducting polymer. Polypyrrole (conducting polymer) is deposited by in-situ technique.

It undergoes polymerization and simultaneous adsorption of PPY on the substrate

surface. Precipitates of PPY tend to coagulate on the substrate surface while the

polymerization is active. When the deposition process is over, there is a significant

amount of residue of un-adhered PPY precipitate left in the reaction solution. The

deposited precipitate has some loosely adhered flakes on the surface. The basic film is

observed to be made up of such flakes. The excess flakes can be removed through the

scotch tape. This test was initially performed to investigate film adherence [135]. It was

found that this reduces the over all thickness of PPY without changing the resistance of

the PPY film, and results in a further decrease in resistivity while simultaneously

increasing film integrity.

Thus, it was found that PPY can be substituted in place of silver paste as an

interconnect and electrode material without losing electrical behavior.Also, filament

flexibility and acoustic impedance matching is enhanced with the conducting polymer

electrode, as will be discussed subsequently.

135

3.3 SEM Data Analyses

SEM (scanning electron microscopy) was used to observe the morphology of the

deposited silver paste and PPY. It was also performed to compare the integrity of both of

these type of materials used as interconnects.

Figure 3.3.1 Diluted silver paste on glass substrate

Figure 3.3.2 High purity silver paste on glass substrate

Figures 3.3.1 & 3.3.2 above are SEM micrographs of silver paste as interconnect

material on glass substrate. As seen from the figure below, diluted silver paste gives a

more consistent thickness and flexibility as compared to the high purity silver

interconnect which tends to be very brittle with a very inconsistent thickness. But as seen

from the 4 probe test in section 3.2 the conductivity of the high purity silver interconnects

far exceeds that of the diluted interconnect.

136

Figure 3.3.3 High Purity silver paste on

Lecoset substrate – (at the junction) Figure 3.3.4 High Purity silver paste on

Lecoset substrate – (at the junction)

Figure 3.3.3 & Figure 3.3.4 are SEM micrographs of high purity silver paste as

interconnect material at the junction with a PVF2 element embedded in a hard resin

substrate. High purity silver paste when dried is very brittle and stiff. As seen from the

SEM images below any slight movement of the single element device is likely to

separates it from the element and the substrate.

Figure 3.3.5 PPY on Lecoset substrate – (at

the junction)

Figure 3.3.6 PPY on Lecoset substrate –

(Surface view)

137

Figure 3.3.7 PPY on Lecoset substrate

Figure 3.3.5 to Figure 3.3.7 are SEM micrographs of conducting polymer PPY

interconnects between single element device embedded in a hard resin substrate. As

indicated in the figures above, PPY as an interconnect on our device has better physical

properties then silver interconnects as listed below:

1. Thickness of interconnect is less than 5 µm (very thin compared to silver paste)

2. The polymer film is very flexible and takes the shape of the substrate that it is

deposited on.

138

Figure 3.3.8 Globular morphology of PPY film on un-stretched PVDF film [128]

Polypyrrole films also tend to show a globular morphology. The average globule

size deduced from SEM micrographs were approximately ½ µm. Globular morphology

was obtained on PVDF substrate surfaces as depicted in the figure 3.3.8. The deposition

time was 90 minutes using the in-situ deposition technique [128].

139

3.4 In Air Acoustics

3.4.1 Semi-Quantitative Acoustic Tests

Acoustic measurements in air were performed to determine quality and integrity

of the electrical connections and to obtain information on sensitivity of the device and

parameters affecting its response.

Figure 3.4.1 shows a comparison of the signal output from a standard hydrophone

(PCMH-2) and that from a typical PVDF device (28µm, height 20mm and width ~ 2mm),

which generates much higher voltages (> 5 volts), compared to the hydrophone [124].

The frequency response of this device is also comparable to that of the high quality

standard hydrophone (ceramic stack). Such high sensitivities of PVDF bending devices in

air are indicative of their suitability for a cochlear implant [124].

Figure 3.4.1 Comparison of in-air test on std. Hydrophone and device prototype [124]

140

3.4.2 Semi-Quantitative Acoustic Tests Using 28µµµµm PVDF Film

Single and multi element devices were made of 28µm PVDF films using various

combinations of electrode interconnect material. The voltage response obtained from the

“Hello” test of each response was compared. PVDF films obtained from MSI Inc. with

Ni-Cu electrodes were used to make the metal electrode PVDF devices. High purity

silver paste was used as an interconnect material in metal electrode-PVDF devices.

Polymer electroded PVDF devices were made from non-electroded PVDF films and both

interconnects and electrodes were made by 90 minute in-situ deposition of polypyrrole.

3.4.2.1 Voltage Response of Metal Electroded PVDF Devices

Table 3.4.1 lists the voltage response of single and multi-element devices

obtained from “Hello” test performed on devices made out of 28µm metal electrode

PVDF devices. The dimensions of the devices were 20mm × 2mm with high purity silver

paste interconnects. The distance between the electrodes was kept constant at 1cm. The

multi-element devices were connected in series and parallel.

141

(a)

(b)

Figure 3.4.2 (a) Multi element device connected in series and (b) Multi element device

connected in parallel

As seen from figure 3.4.2 (a) the device is connected in series in such a way that

there is only one path for the flow of charge. In figure 3.4.2 (b), the devices are shown to

be connected in parallel.

Table 3.4.1 Voltage response of metal electrode-PVDF devices Vs Number of elements Voltage response (± 0.5 volts variation)

Number of Elements Series connection

(Vs volts)

Parallel connection

(Vp volts)

1 2.4

2 4.5 3.6

3 4.4 4.8

4 3.9 5.5

5 3.2 ~6.0

6 1.6 ~7.0

142

As seen from the table above, the response of devices connected in series

and parallel are completely different. As the number of elements increases there is a

considerable amount of voltage scaling upwards in the device where the elements are

connected in parallel. In a series device the voltage is seen to scale down, with an

increase in number of elements. It is shown that for a 6 element device Vs is 1.6 volts

whereas Vp is approximately 7.5.

The PVDF elements with any electrode material is thus shown to be similar to a

capacitor system. When deformed, there is a generation (accumulation) of positive and

negative charges on the surface. Therefore, in the cochlear environment the surface with

positive charge would attract the negative ionic species (Cl- ions) and the negatively

charged surface would attract the positive ionic species (Na+ ions). Each of these

elements work like capacitors, thus when in parallel they scale up the potential because of

high charge accumulation and when in series, scale down because of charge

neutralization. Charge neutralization in these devices has to be understood at a detailed

level.

As explained in section 2.3 the voltage responses taken were an average of ten

iterations. Therefore when repeated again there will a probability of ± 0.5 variation in the

voltage result.

143

Figure 3.4.3 Voltage response of a 4 element device connected in parallel

Figures 3.4.3 shows the voltage response of a four filament metal electroded

PVDF device connected in parallel. As seen from the figures, the device connected in

parallel gives a high voltage response as compared to the device connected in series. The

voltage response obtained in table 3.4.1 is an average of 10 results obtained from the

“Hello” test. The results obtained can be duplicated within the range of ± 0.5 of the

obtained value. The voltage response spectrum of the devices depends on a lot of factors.

This causes each response spectra to be different. Some of the factors affecting the

response are as follows:

1. The devices are very sensitive. Any external noise has an influence on the voltage

spectra.

2. The “Hello” test though done by the same person has variations in it. This also causes

the voltage spectra to be different each time.

144

3.4.2.2 Voltage Response of Metal Electroded PVDF Devices with Varying Device

Dimensions

Table 3.4.2 lists the voltage response of metal electroded PVDF devices Vs

number of elements obtained from “Hello” test performed on devices made out of 28µm

Ni-Cu electroded PVDF elements of dimensions 20mm × 2mm with high purity silver

paste interconnects. The variable factor in this case was the element length. The distance

between the elements was kept constant at 1cm. The multi-element devices were

connected both in series and parallel.

145

Table 3.4.2 Voltage response of length metal electrode-PVDF devices Vs Number of elements

Voltage Response(± 0.5 volts variation) Number of Elements

Length of PVDF element

(mm)

Series connection (Vs volts)

Element Width 2mm

Parallel connection (Vp volts)

Element Width 2mm

20 2.8 15 3.3 10 2.8

1

05 1.4 20 4.7 3.5 15 3.9 3.5 10 1.0 3.0

2

05 0.9 1.4 20 4.3 5.0 15 2.0 4.7 10 0.4 4.6

3

05 0.2 1.2 20 4.0 5.3 15 2.7 4.2 10 2.0 2.4

4

05 0.1 1.2

From table 3.4.2 it is shown that as the length of the PVDF elements in the single

and multi-element devices decreases, the voltage response also decreases. Comparing the

voltage response of series and parallel device it was found that the device connected in

parallel has a greater voltage response. Even for a device with dimensions 20mm ×

05mm the voltage response obtained by the parallel connection always exceeds the

voltages response obtained from the device connected in series.

146

(a)

(b)

Figure 3.4.3 Voltage response of a single element device (a) Length of PVDF = 05mm and (b) Length of PVDF = 20mm

147

(a)

(b)

Figure 3.4.4 Voltage response of a 4 element device connected in parallel (a) Length of PVDF = 20mm and (b) Length of PVDF = 05mm

148

(a)

(b)

Figure 3.4.5 Voltage response of a 4 element device connected in series (a) Length of PVDF = 20mm and (b) Length of PVDF = 05mm

149

As explained in section 2.3 the discrepancies in the voltage response spectra is

because of the different human error which cannot be eliminates for in-air acoustics.

Tables 3.4.3 (a) and (b) shows the voltage response of a 5 element metal

electroded PVDF device. In the table 3.4.3 (a) length of the element and type of

interconnect connection is the varying factor. In table 3.4.3 (b) voltage response is taken

for a device connected in parallel and varying the dimensions of the device.

Table 3.4.3(a) lists the voltage response of a 5 element device

Voltage Response(± 0.5 volts variation) Number of

Elements

(parallel)

Length of PVDF

element (mm)

Element Width

2mm

Series connection

(Vs volts)

Parallel connection

(Vp volts)

20 4.9 5.8

15 3.5 3.2

10 2.2 2.2

5

05 0.7 1.0

For table 3.4.3 (b) the results were obtained from “Hello” test performed on

devices made out of 28µm Ni-Cu electroded PVDF elements of dimensions 20mm ×

2mm and 20mm × 1mm with high purity silver paste interconnects. The distance between

the electrodes was kept constant at 1cm. Since the voltage response of the series voltage

decreases with increase in the number of elements, the 5 element devices were connected

in parallel and tested.

150

Table 3.4.3(b) Voltage response of 5 element metal electrode-PVDF devices Vs

changing dimensions

Dimension of element(± 0.5 volts variation)

Element width 2 mm Element width 1 mm

Number of Elements

(Parallel Connection)

Length of

PVDF element

(mm) Vp(p-p) Volts Vp(p-p) Volts

20 3.6 2.6

15 3.1 1.2

10 2.6 1.0

5

05 1.0 0.5

As seen from both the table above the voltage response from the 5 element device

in parallel has a better response that the device connected in series. Changing the

dimensions from 20 mm × 2 mm to 20 mm × 1 mm also decreases the response.

Moreover length of the element in a PVDF device also plays an important part. It was

seen that a decrease in the element height made the element stiffer causing a lower

voltage response.

151

(a)

(b)

Figure 3.4.6 Voltage response of a 5 element device connected in parallel, dimensions = 20 mm × 2 mm (a) Length of PVDF = 20mm and (b) Length of PVDF = 05mm

152

(a)

(b)

Figure 3.4.7 Voltage response of a 5 element device connected in parallel, dimensions = 20 mm × 1 mm (a) Length of PVDF = 20mm and (b) Length of PVDF = 05mm

153

From sections 3.4.2.1 and 3.4.2.2 it can be concluded that the voltage response of

a metal electroded PVDF device depends on the parameters mentioned below:

1. Type of interconnect. This is by far the most important factor affecting the voltage

response of the device. The response of a device connected in parallel exceeds that of a

device in series.

2. Length of element in the device. As the element length increases it increases the

sensitivity of the device. But the important thing to note is that for a device with element

of height 05mm does give us a good response which can be increased by increasing the

number of elements.

3. Dimension of the elements in the device. A change in dimensions also affects the

working of the device.

3.4.2.3 Voltage Response of PPY Electroded PVDF Devices

Table 3.4.4 lists the voltage response of a single element device obtained from

“Hello” test performed on devices made out of 28µm PVDF elements of dimensions

20mm × 2mm with polypyrrole electrode and interconnect. PPY was deposited in the all

polymer devices using in-situ deposition technique for a period of 90 minutes. Voltage

response is obtained by changing the length of the conducting polymer interconnects.

154

Table 3.4.4 Voltage response of a single filament device

Voltage response (± 0.5 volts variation)

Length of interconnect

(cms)

Series connection

(Vs volts)

Parallel connection

(Vp volts)

3.0 2.8

2.5 3.6

2.0 3.4

1.5 3.4

1.0 3.6

155

(a)

(b)

Figure 3.4.8 Voltage response of a single element device (a) Length PPY interconnect = 3 cms and (b) Length PPY interconnect = 1 cms

156

It is seen from table 3.4.4 that as the length of the interconnect decreases the

voltage response of the device increases. The difference in the voltage spectrum as

explained in section 3.4.2.1 is because of the highly sensitive nature of the device which

picks up even the slightest noise variations in the background. Since it is not possible to

eliminate all noise sources from the experimental environment the variations in the

spectrum will exist.

Table 3.4.5 lists the voltage response of a 2 element device obtained from “Hello”

test performed on devices made out of 28µm PVDF elements of dimensions 20mm ×

2mm with polypyrrole electrode and interconnect. Voltage response is obtained by

changing the length of the conducting polymer interconnects. The interconnects were

patterned in series and parallel.

Table 3.4.5 Voltage response of a 2 element device with PPY electrode and interconnects

Voltage response(± 0.5 volts variation)


(cms)

Series connection

(Vs volts)

Parallel connection

(Vp volts)

3.5 1.4 2.9

3.0 1.4 2.8

2.5 1.5 3.2

2.0 1.3 3.4

1.5 2.5 3.5

1.0 2.4 3.4

0.5 2.5 3.5

No interconnect 2.9 3.6

157

From the table above it was seen that the response from a parallel connection was

always higher than the device connected in series. Also as we decrease the interconnect

length it causes an increase in response from the device.

(a)

(b)

Figure 3.4.9 Voltage response of a 2 element device connected in parallel (a) Length PPY interconnect = 3 cms and (b) No interconnect

158

(a)

(b)

Figure 3.4.10 Voltage response of a 2 element device connected in series (a) Length PPY interconnect = 3 cms and (b) No interconnect

159





patterned in parallel.

Figure 3.4.11 4 element device connected in parallel with PPY electrodes and

interconnect

Table 3.4.6 Voltage response of a 4 element device connected in parallel


(cms)

Voltage response (± 0.5 volts variation)

Parallel connection

(Vp volts)

3.5 2.6

3.0 2.7

2.5 2.9

2.0 2.9

1.5 3.4

1.0 3.7

0.5 3.7

No interconnect 4.0

160

(a)

(b)

Figure 3.4.12 Voltage response of a 4 element device connected in parallel (a) Length PPY interconnect = 3 cms and (b) No interconnect

161





patterned in parallel.

Figure 3.4.13 5 element device connected in parallel with PPY electrodes and

interconnect

Table 3.4.7 Voltage response of a 5 element device connected in parallel


(cms)

Voltage response(± 0.5 volts variation)

Parallel connection

(Vp volts)

3.0 2.9

2.5 3.6

2.0 3.9

1.5 4.5

1.0 4.7

No interconnect 4.8

162

(a)

(b)

Figure 3.4.14 Voltage response of a 5 element device connected in parallel (a) Length


163

As seen from tables 3.4.4 to 3.4.7 for a single and multi element device the factors

affecting the voltages response can be summaries as below.

1. Length of the interconnect. To get optimum voltage response from the device the

interconnect should be very small.

2. Type of interconnect. Device connected in parallel gives the best response. It is

therefore necessary to study the factors effecting a parallel connection in detail.

Voltage response Vs Length of Interconnect

0

1

2

3

4

5

6

3.5 3 2.5 2 1.5 1 0.5 No

Length of interconnect (cms)

Vol

tage

Res

pons

e (V

olts

1 element

2 element

4 element

5 element

Figure 3.4.30 Voltage Response Vs Length of interconnect

The above graph shows the voltage response of multi element devices as a

function of interconnect length. It can be seen that a decrease in interconnect length

increases its voltage response. This is a very encouraging result because as final device

has to be very small the length of the interconnect will also be very small.

164

3.4.2.4 All Metal Device Vs All Polymer Device Connected in Parallel

Table 3.4.11 compares the voltage response of an all polymer and metal single

and multi-element 28µm PVDF device.

Table 3.4.11 Voltage response of a metal and polymer single and multi element device

Number of elements

(Parallel Connection)

Voltage response of an all

polymer device

Vp volts

(± 0.5 volts variation)

Voltage response of a metal

device

Vp volts


1 2.8 2.4

2 3.6 3.6

3 4.8 3.7

4 5.5 3.7

5 ~6.0 4.8

From the table above it is show that voltage response of the all polymer compared

to the metal device is higher. Theoretically devices made out of polymer (polypyrrole)

electroded filaments should show a higher voltage response than the metal (Ni-Cu)

electroded filament device. This is because of less stiffness rendered to the filaments by

polymer electrodes than by the metallic electrodes. Polymer electrodes are highly flexible

and do not restrict the vibrations in the piezoelectric PVDF films. Practically patterning

of the interconnects plays an important role in the output voltage response. The

interconnect pattern which obtains optimum response from a metal device might not

produce the optimum response in polymer devices.

165

Figure 3.4.13 shows the interconnect pattern used. In this pattern since the

elements are spread over a longer area the intensity of sound wave interacting with each

of the elements maybe different.

166

3.4.3 Semi-Quantitative Acoustic Tests Using 52µµµµm PVDF Film

Single and multi element devices were made of 52µm PVDF films with different

electrode materials and different interconnect material. The voltage response obtained

from the “Hello” test of each response was compared. PVDF films obtained from MSI

Inc. with Ni-Cu electrodes were used to make the metal electrode PVDF devices. High

purity silver paste was used as an interconnect materials in metal electrode-PVDF

devices. Polymer electroded PVDF devices were made from non-electroded PVDF films

and interconnects and electrodes were made by 90 minute in-situ deposition of

polypyrrole.

3.4.3.1 Voltage Response of Metal Electroded PVDF Devices

Multi element devices were made and tested. Table 3.4.8 lists the voltage

obtained from “Hello Test” performed on devices made out of 52µm Ni-Cu electroded

PVDF elements of dimensions 20mm x 2mm.

167

Table 3.4.8 Comparison of Series and Parallel Connection Metal electroded PVDF

elements (20mm x 2mm) [135]

No. of elements

Series Voltage


(Vs Volts)

Parallel Voltage


(Vp Volts)

1 3.0 - 3.5

2 4 - 4.5 3.5 - 4.0

3 3.0 - 3.5 4.5 - 5.0

4 3.0 - 3.5 >5.0 (≈ 5.6)

5 ~3.0 > 6.0

6 ~2.5 > 7.0

7 ~0.5 >> 7.0

8 ~0.5 >> >7.0(≈ 8.5)

12 ~0.45 >> >>8.0(≈ 13)

Table 3.4.9 Comparison of Series and Parallel Connection PPY electroded PVDF

elements (20mm ×××× 2mm) [135]

No. of elements

Series Voltage


(Vs Volts)

Parallel Voltage


(Vp Volts)

4 ~3.0 >7.0(≈ 7.3)

8 ~2.0 >>7.5(≈ 9)

12 ~1.0 >>>7.5(≈ 15)

168

Table 3.4.10 Comparison of Series and Parallel Connection PPY electroded PVDF

elements (10mm ×××× 1mm) [135]

No. of elements

Series Voltage


(Vs Volts)

Parallel Voltage


(Vp Volts)

4 ~3.0 >> 8.0

8 ~2.6 >> 8.0

12 ~2.5 > >9.0(≈ 10)

Figure 3.4.15 Voltage response of 8-element device connected in parallel [135]

Table 3.4.8 and 3.4.9 list the voltages obtained from long (20mm x 2mm) and

shorter (10mm x 1mm) elements with polypyrrole (PPY) electrodes from the “Hello

Test”. Because of the limitation of the software in use, ez-analyst, used for data

acquisition in the acoustic tests, it was not possible to predict the accurate voltage

response, if it increased above +5 volts. Figure 3.4.15 shows the response of an eight-

element device (dimensions: 10mm x 1mm), connected electrically in parallel. The

voltage response listed for this device in the above table is >>8.0 volts. As can been in

the plot, the curve is cutoff at 5 volts value because of the software limitations. General

169

prediction from this graph makes understand that total voltage out of this device would

definitely be greater than 5 volts.

Table 3.4.12 lists the voltage response of a multi-element device obtained from

“Hello” test performed on devices made out of 52µm Ni-Cu electroded PVDF elements

of dimensions 20mm × 2mm and 20mm × 1mm with silver paste interconnect. Voltage

response is obtained by changing the length of the PVDF elements. The interconnects

were patterned in parallel.

The voltage response in Table 3.4.12 does not follow the rule that was noticed in

the earlier sections. This is because of the device response is very sensitive to

interconnect patterning. This is studied in detail in the following sections.

Table 3.4.12 Voltage response of a multi-element device connected in parallel

Parallel Voltage (Vp volts) (± 0.5 volts variation) No. of elements (Parallel

Connection)

Length of Peizo elements

(mm)

Element width 2 mm

Element width 1 mm

20 mm 7.2 7.3

15 mm 6.5 6.2

10 mm 2.4 2.1

2

05 mm 1.5 1.0

20 mm 6.0 6.0

15 mm 4.0 6.1

10 mm 3.8 5.2

3

05 mm 0.06 1.1

20 mm 6.4 6.2

15 mm 5.1 4.8

10 mm 3.6 3.9

4

05 mm 2.0 2.6

170

(a)

(b)

Figure 3.4.16 Voltage response of a 2 element device connected in parallel (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2

mm, Length of PVDF = 20mm

171

(a)

(b)

Figure 3.4.17 Voltage response of a 3 element device connected in parallel (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm ×

2 mm, Length of PVDF = 20mm

172

(a)

(b)

Figure 3.4.18 Voltage response of a 4 element device connected in parallel (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2


173

From tables 3.4.8 to 3.4.12 the following important characteristics can be noted in

the devices.

1. Thickness of PVDF. Comparing the results obtained from devices made of 28µm and

52µm PVDF we see that a 52µµµµm PVDF device gives a higher voltage response.

This is because as the thickness increases flexibility increases giving a higher

response. Decreasing the thickness of the PVDF films results in curling and distortion

of the elements. This tends to drastically reduce the voltage response to acoustic

waves.

2. Type of interconnect. A parallel connection always gives a high voltage response

as compared to a series type. But factors such as element spacing affect the

response of the device. Type of spacing of elements giving optimum voltage response

has to be understood in greater detail.

3. Length of element in device. An increase in the length of the element causes the

voltage response to decrease. This is because the elements in the device are less

flexible. This decreases the sensitivity of the device. But the important thing to note is

that though the sensitivity decreases it doesn’t become zero for a device with element

height 05mm which is very small. This is a very encouraging result because as final

device has to be very small the height of the elements in the device will also be very

small.

4. Dimensions of elements and interconnect thickness. Decreasing the dimension of the

device also decreases the sensitivity of the device. Since the main goal is

manufacturing a device which is small enough to be implanted in human cochlea.

Though the voltage response decreases with decreasing dimensions, increasing the

174

number of elements in the device in turn increases its response. This will help obtain

the desired voltage from a device with small dimensions.

3.4.3.2 Voltage Response of Polymer Electroded PVDF Devices

Table 3.4.13 to 3.4.21 lists the voltage response of a multi-element device

obtained from “Hello” test performed on all polymer devices. The devices were made out

of 52µm PPY electroded PVDF elements of dimensions 20mm × 2mm and 20mm × 1mm

with PPY interconnect. Voltage response is obtained by changing the length of the PVDF

elements. The interconnects were patterned in parallel.

Table 3.4.13 Voltage response of an all polymer single element device connected in

parallel

Parallel Voltage (Vp volts) (± 0.5 volts variation) No. of

elements

Length of

Peizo elements

(mm)

Element width 2mm Element width 1mm

20 mm >>8.1 (≈10.1) >>8.3(≈10.3)

15 mm 6.2 >8.0

10 mm 2.8 5.7

1

05 mm 2.6 3.3

175

(a)

(b)

Figure 3.4.19 Voltage response of a single element device connected in parallel (a)

dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2


176

As discussed in section 3.4.2, because of the limitation of the software in use it

was not possible to predict the accurate voltage response, if it increased above +5 volts.

Figure 3.4.19 shows the response of an single-element device (dimensions: 20mm ×

2mm, 20mm × 1mm), connected electrically in parallel. The voltage response listed for

this device in the above table is >>8.3 and >>8.1 volts. General prediction from this

graph makes understand that total voltage out of this device would definitely be greater

than 5 volts.

Table 3.4.14 Voltage response of an all polymer 2 element device connected in parallel


elements

Length of

Peizo elements

(mm)

Element Width 2mm Element Width 1mm

20 mm >>8.0 (≈11) >>8.4 (≈11.4)

15 mm >7.3 (≈9) >>7.5 (≈9.5)

10 mm 5.4 4.0

2

05 mm 2.1 3.0

177

(a)

(b)

Figure 3.4.20 Voltage response of a 2 element device connected in parallel (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2 mm,


178



elements

Length of

Peizo elements

(mm)


20 mm >>8.2 (≈11.2) >>8.5 (≈11.5)

15 mm >7.0 (≈9) >7.3 (≈9.7)

10 mm 6.1 6.8

3

05 mm 3.5 2.6

179

(a)

(b)

Figure 3.4.21 Voltage response of a 3 element device connected in parallel (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2 mm,


180

From table 3.4.13 to 3.4.15 shows voltage response of single and multi-element

device connected in parallel and also changing the dimensions of the element. It is found

that decreasing the length of the peizo elements the voltage response also decreases. But

comparing the response of an all metal device table 3.4.12 to all polymer device, the

response obtained by the all polymer device is a lot higher. Also short “element” all

polymer devices have a higher voltage response than the all metal devices.

This is because of less stiffness rendered to the filaments by polymer electrodes

than by the metallic electrodes. Polymer electrodes are highly flexible and do not restrict

the vibrations in the piezoelectric PVDF films. It is known and can be visually observed

in the case of metal electroded piezo filaments, that the less flexibility of the device is

mainly due to the stiff metallic coatings. This result too adds up towards the final device

design, as the final device would require highly flexible filaments with high sensitivity

movements on interaction with acoustic wave fronts.

Table 3.4.16 to 3.4.21 lists the voltage response of a multi-element device

obtained from “Hello” test performed on devices made out of 52µm PPY electroded

PVDF elements of dimensions 20mm × 2mm and 20mm × 1mm with PPY interconnect.

Voltage response is obtained by changing the length of the PVDF elements. The

interconnects were patterned in parallel.

181

(a)

(b)

Figure 3.4.22 Prototype of a 4 element device connected in parallel (a) Design 1 and (b) Design 2

From the figure 3.4.22 two different types of parallel connection can be seen. In

design 1 the elements are connected in parallel in a single row whereas in design 2 the

elements are spaced over a broader area.

Table 3.4.16 Voltage response of an all polymer 4 element device connected in

parallel (Design 1)


elements

Length of

Peizo elements

(mm)


20 mm > 4.4 (≈5.4) > 5.3 (≈6.4)

15 mm 3.4 3.5

10 mm 3.2 3.2

4

05 mm 0.5 0.9

182

Design 1

(a)

(b)

Figure 3.4.23 Voltage response of a 4 element device connected in parallel (Design 1)(a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2


183


(Design 2)


elements

Length of

Peizo elements

(mm)


20 mm >>>9.2 (≈14.5) >>>8.4 (≈14)

15 mm >>8.2 (≈12.2) >>>8.2 (≈13)

10 mm 5.0 6.4

4

05 mm 2.1 4.3

184

Design 2

(a)

(b)

Figure 3.4.24 Voltage response of a 4 element device connected in parallel (Design 2) (a) dimensions = 20 mm × 1 mm, Length of PVDF = 20mm and (b) dimensions = 20 mm × 2


185


(Design 1)


elements

Length of

Peizo elements

(mm)


20 mm 4.7 5.3

15 mm 3.7 4.5

10 mm 3.6 4.0

5

05 mm 0.7 1.2

186

Design 1

(a)

(b)



187


(Design 2)


elements

Length of

Peizo elements

(mm)

Element Width 2mm Element width 1mm

20 mm >>>8.0 (≈16) >>>8.7(≈16.5)

15 mm >>>7.8(≈14) >>>6.8(≈15)

10 mm 3.6 4.1

5

05 mm 3.2 3.1

188

Design 2

(a)

(b)



189

(a)

(b)

Figure 3.4.27Prototype of a 6 element device connected in parallel (a) Design 1 and (b) Design 2


(Design 1)


elements

Length of

Peizo elements

(mm)


20 mm > 5.5 (≈6.5) > 6.2 (≈7.2)

15 mm 5.3 5.7

10 mm 3.9 4.7

6

05 mm 1.1 1.4

190

Design 1

(a)

(b)



191


(Design 2)


elements

Length of

Peizo elements

(mm)


20 mm >>>9.3 (≈20) >>>8.5 (≈17)

15 mm >>>8.9 (≈16) >>>8.2 (≈15)

10 mm >>7.7 (≈12) 5.0

6

05 mm 3.3 3.4

192

Design 2

(a)

(b)



193

Tables 3.4.16 to 3.4.21 shows the voltage response of multi-element device

changing the length of the interconnects and the changing dimension. Figures 3.4.22 (a)

and (b) shows the interconnect patterns used. The voltage response obtained by “Design

2” is significantly higher than by “Design 1”. This can be explained as below.

In “Design 1” the elements were placed behind one another in a long row.

Therefore the acoustic waves hitting each element of the device is vastly unequal. It is

suspected that the elements at the beginning of the row being closer to the acoustic source

may vibrate or feel a greater acoustic force than the elements in the back where the

strength of the acoustic wave might be significantly less. Therefore the collective voltage

response of the device decreases.

In “Design 2” the element were scattered over a broader area. This caused the

acoustic waves to hit the elements in the device in a much more even fashion as

compared to design 1. This causes the collective voltage response of a device made from

design 2 to have a higher voltage response as all the elements of the device vibrates with

fairly equal strength.

Decreasing the width of the elements increases their flexibility. Therefore they do

not restrict the vibrations in the piezoelectric PVDF films. This result adds up towards the

final device design, as the final device would require highly flexible filaments with high

sensitivity movements on interaction with acoustic wave fronts.

194

Figure 3.4.31 Future Experimental Setup

Separation distances and actual design will need further investigation. This can be

done by studying the acoustic response of each element of the device. This can be done

by connecting each element separately to the current ez-analyst software which in its

present state is capable of capturing voltage response from 16 separate elements of a

device.

Table 3.4.22 lists the voltage response of a single-element device obtained from

“Hello” test performed on devices made out of 52µm PPY electroded PVDF elements of

dimensions 10mm × 0.5mm with PPY interconnect. Voltage response is obtained by

changing the length of the PVDF elements.

195

Table 3.4.22 Voltage response of a single-element device. Dimensions: 10mm × 0.5mm

No. of elements Length of Peizo elements

(mm)

Voltage Response (volts)


10 mm 3.5

9 mm 3.3

8 mm 2.2

7 mm 2.2

6 mm 2.0

5 mm 1.7

4 mm 1.4

3 mm 1.3

2 mm 1.4

1

1 mm 1.4

196

(a)

(b)

Figure 3.4.32 Voltage response of a single element device (a) Dimensions 2mm × 0.5mm (b) Dimensions 10mm × 0.5mm

10mm × 0.5mm PPY device is a very small and thin device. This is the smallest

device that can be made manually. The voltage response of the single element device

10mm × 0.5mm as seen in the table above gives a fairly good result of 3.5 volts. Reducing

197

the dimension to 01mm × 0.5 mm still gives a voltage response of 1.4 volts. This is very

encouraging as is shows that making devices for actual cochlear implants is not only

possible practically but also theoretically.

3.4.3.3 Comparing metal electroded PVDF devices Vs Polymer electroded PVDF

devices

From table 3.4.8 and tables 3.2.12 to 3.4.21 voltage responses from all metal to all

polymer devices can be compared.

Table 3.4.23 Voltage response of a metal electroded device

No. of elements Parallel Voltage (Vp Volts) (± 0.5 volts variation)

20mm × 2 mm

1 3.0-3.5

2 3.5 - 4.0

3 4.5 - 5.0

4 >5.0 (≈ 5.5)

5 6

6 7

198

Table 3.4.23 (a) Voltage response of a polymer electroded device


elements

Length of Peizo

elements Element width 2mm Element width 1mm

1 20 mm >>8.1 (≈10.1) >>8.3(≈10.3)

2 20 mm >>8.0 (≈11) >>8.4 (≈11.4)

3 20 mm >>8.2 (≈11.2) >>8.5 (≈11.5)

4 20 mm >>>9.2 (≈14.5) >>>8.4 (≈14)

5 20 mm >>>8.0 (≈16) >>>8.7(≈16.5)

6 20 mm >>>9.3 (≈20) >>>8.5 (≈17)

Table 3.4.23 and 3.4.23(a) shows voltage response of metal and all polymer

device connected in parallel. Figure 3.4.28 and 3.4.29 shows the response of a multi

element device. General prediction from the graph makes us understand that total voltage

out of this device would definitely be greater than 5 volts.

Figure 3.4.33 Comparison of metal VS polymer electroded device

199

From the above graph the following can be summarized.

1. An all polymer device has a higher voltage response than an all metal device.

2. Decreasing the dimension of the device does not affect the sensitivity of the device by

a large value.

The study of behavior of the device and interconnects show that an increasing number

of elements also produces higher voltage response. But with higher number of elements

problems such as wave front interaction on each element of the device. It is also required

to understand and model the vibrational analysis of these elements in a fluid medium of

different densities and volume. Also recommended is a study of phase differences

produced (if any) as a result of spatial separations.

As has been found from the literature, cochlear implants need a total charge density

of about 10 nC/mm2 and an optimum voltage for depolarizing the eighth nerve, although

there are no fixed values of charge and voltage found yet, it is recommended that such a

device needs to be fabricated and tested in guinea pigs. It is with these tests, we would be

able to optimize the total charge and voltage figures required. Research needs to be done

to understand the combination of both charge and voltage for the depolarization of the

auditory nerve and obtain the optimum figure of merit.

200

3.4.4 Film thickness

Hello Test was performed on devices, made out of polypyrrole electroded PVDF

films of 28 and 52µm thicknesses. These 5-element devices, connected in parallel, had

element dimensions of 20mm x 2mm each. The variable used in this experiment was the

PVDF film thickness.

Table 3.4.24 Effect of peizo film thickness on voltage response (metal device)

Device specifications Voltage response(± 0.5 volts variation)

Vp volts

Device A: 28µm all metal device 4.9

Device B: 52µm all metal device 6.4

Table 3.4.25 Effect of peizo film thickness on voltage response (PPY device)

Device specifications Voltage response(± 0.5 volts variation)

Vp volts

Device A: 28µm all PPY device 5.0

Device B: 52µm all PPY device >>8.0

From the tables 3.4.24 and 3.4.25 we see that thicker the PVDF film higher is the

voltage response for both metal and polymer devices. The devices made from 52µm

vibrated more in response to the sound waves as compared to 28µm device. This cause

the net voltage response to be higher for the thicker PVDF elements.

201

3.4.5 Dimensions and Interconnect thickness

Hello Test was performed on devices, made out of polypyrrole electroded PVDF

films of 52µm thicknesses. This single element device had decreasing with a factor of 2.

The variable used in this experiment was the device dimensions.

Table 3.4.26 Effect of peizo film dimension on voltage response

Device specifications Voltage Response(± 0.5 volts variation)

V(peak - peak) volts

Device A: 10mm x 2mm 3.0

Device B: 05mm x 1mm 2.9

Device C: 2.5mm x 0.5mm 1.4

From the table above we see that as the device dimensions get smaller the voltage

response also decreases, but even the smallest dimension gives a good response. This is

an encouraging result as the final device would be very small.

202

Chapter 4

SUMMARY

Polymer Electrode and Interconnect

• High flexibility, better adhesion to substrate material, high electrical conductivity and

ease of fabrication (in-situ polymerization deposition technique) makes PPY a very

good candidate replacing Ni-Cu as electrode and high purity silver paste as

interconnect material in cochlear implant device application.

• Devices made out of polymer (polypyrrole) electroded filaments show a higher

voltage response than the metal (Ni-Cu) electroded filament device. Polymer

electrodes are highly flexible and do not restrict the vibrations in the piezoelectric

PVDF films.

• Elements with polymer electrodes, connected in parallel are better than those with

metal electrodes, connected in parallel.

• Devices with thicker PVDF elements show an increase in voltage response.

Device Design

• Results shows that there is an increase in voltage response when the elements are

connected in parallel whereas the voltage response decreases when connected in

series.

• Interconnect patterning plays a very important role. Results show that when elements

are connected over a broader area as compared to that in a long row gives a higher

voltage response.

203

• Multi-element device with elements placed closer to each other show higher response

than those with farther separations. Spatial distances between the elements, plays an

important role in the device design.

204

Chapter 5

CONCLUSION

• High electrical conductivity, flexibility, ease of fabrication (in-situ polymerization

deposition technique), stability in the application frequency range and temperature,

and strong adhesive properties make polypyrrole (PPY) a strong candidate for being

the electrode for the piezoelectric cochlear implant device application. With a better

acoustic response obtained from devices having PPY electrode and interconnect than

with the Ni-Cu electroded devices and silver paste interconnect materials, it can be

concluded that PPY has the ability to replace metal electrodes from the piezoelectric

cochlear implant application.

• Multi-element device was fabricated and tested for in-air acoustic sensitivity. The

results of these tests have been the basis of the conclusion that parallel connections

work better than the series connections for the elements. Higher voltages are obtained

as we increase the number of elements connected electrically in parallel than in series.

• Acoustic test results have shown that shorter element devices show better response

than the longer element devices. Increasing the number of elements, so as to improve

the charge densities could increase the effective surface area of the device. Higher

voltages are obtained from these devices because of the materials piezoelectric nature,

whereas higher charge densities can only be obtained by increasing the effective

surface area.

• Spatial separations of the elements and patterning of interconnects in the device

plays an important role in determining the effective voltage response from these

205

multi-element devices. These factors should be taken care of while designing the final

device.

206

Chapter 6

FUTURE WORK

• Device Design & Modeling

From the results obtained the toothbrush device shows high sensitivity in-air acoustic.

Current studies are focused on the feasibility of an all polymer device i.e., polymer

electrodes and interconnects. Further studies need to be done on electrode and

interconnect patterning. Other design issues, like the width dependence of piezo elements

on the acoustic sensitivity, a system to discharge the voltage into the cochlea fluid using

a non-wire system, critical number of elements to be connected in parallel for desired

sensitivities, electrical circuit combinations for desired voltage response need to be

investigated. A computer program which incorporates the different results from the above

studies needs to be written which can predict voltage response and charge density.

Miniaturization of the device to implantable size is the next task to test them in real-time

environment in guinea pigs. Nano-fabrication techniques are recommended for such

miniaturized devices.

• In-air & Underwater Testing

Further improvisation is required in the design of external and underwater test stations.

We need to understand the device design and acoustic wavefront progresses and

interaction with the device and its elements, and also understand the cancellation or

addition of voltage response taking place, if any.

The test station should be able to

1. Get voltage response from each element of a multi element device

2. Get voltage response from a combination of elements of a multi element device

207

3. Get voltage response of the device as a whole

These tests would be able to give a feedback about a future direction for the research to

develop the final, fully implantable device.

208

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