ELECTROCHEMICAL CHARACTERIZATION OF TISSUE ENGINEERED ELECTRONIC NERVE INTERFACES

By

SRUTHI NATT

A THESIS PRESENTED TO THE GRADUATE SCHOOL

OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF

MASTER OF SCIENCE

UNIVERSITY OF FLORIDA

2017

© 2017 Sruthi Natt

To Aravind Krishna For his advice, love and faith

4

ACKNOWLEDGMENTS

Foremost, I would like to thank Dr. Jack Judy for being an enthusiastic advisor

and providing me with wonderful opportunities and invaluable lessons. I’ve been

inspired by his sense of optimism, integrity and willingness to collaborate. His

coursework in MEMS provided me with a solid foundation that was required to

understand a significant portion of this research.

I thank my committee members, Dr. Erin Patrick and Dr. Yong-Kyu Yoon for

their guidance and support. Dr. Patrick provided valuable feedback that gave a better

direction to my work. Dr. Yoon was always willing to brainstorm any research ideas.

I would also like to thank the talented group of collaborators who were

instrumental in electrode design and fabrication presented in this work. Special thanks

to the post-doc of our lab, Cary Kuliasha, for his willingness to help me fix any problems

with my electrochemistry experiments. Trying to set-up the experiment on my own

would have been frustrating. I would also like to extend many thanks to the previous

post-doc, Vidhi Desai, for encouraging to complete my thesis in a timely manner. I am

grateful to Jose Alcantara, a fellow biomedical engineering graduate student, for

assisting me with Matlab coding.

I would like to thank all my friends here in Gainesville for giving me many good

memories that I will cherish forever.

I would also like to express my deepest gratitude to my cousin, Praveen Natt, for

his steadfast support and motivation throughout my student life. Finally, I would like to

thank my parents, Sriram Natt and Sai Lakshmi, and my brother, Sudarshan Natt

Srinivas, who have been my biggest fans and a major source of moral support.

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TABLE OF CONTENTS page

ACKNOWLEDGMENTS .................................................................................................. 4

LIST OF TABLES ............................................................................................................ 7

LIST OF FIGURES .......................................................................................................... 8

LIST OF ABBREVIATIONS ........................................................................................... 10

ABSTRACT ................................................................................................................... 11

CHAPTER

1 INTRODUCTION .................................................................................................... 13

2 LITERATURE REVIEW .......................................................................................... 15

Background on the Electrode-Electrolyte Interface ................................................. 15 Capacitance Models ......................................................................................... 15

Helmholtz Model ........................................................................................ 16 Gouy-Chapman Model ............................................................................... 17 Stern Model ................................................................................................ 18

Constant-Phase Element.................................................................................. 19 Series Resistance ................................................................................................... 19

Charge-Transfer Resistance ................................................................................... 20 Randles Model ........................................................................................................ 21

Variations of Randles Model ............................................................................. 22 Polarizable Electrode Behavior ........................................................................ 22

Low-Frequency Behavior.................................................................................. 22 High-Frequency Behavior ................................................................................. 23

3 EXPERIMENTAL METHODS ................................................................................. 26

TEENI Design and Layout ...................................................................................... 26

TEENI Fabrication ............................................................................................ 27 PCB-TEENI integration ........................................................................................... 28 Experimental Set-up ............................................................................................... 30

Electrochemical Cell ......................................................................................... 30 Instrumentation ................................................................................................. 31 Data Acquisition and Analysis .......................................................................... 32

Electrochemical Impedance Spectroscopy ............................................................. 33

4 RESULTS AND DISCUSSION ............................................................................... 37

EIS of TEENI with 3 Threads .................................................................................. 37

6

EIS of TEENI with 4 threads ................................................................................... 38

ACA Bonded TEENI-PCB Assembly ...................................................................... 38

Equivalent Circuit Modeling .................................................................................... 39

5 CONCLUSION AND FUTURE WORK .................................................................... 50

EIS Data and Modeling ........................................................................................... 50 Measurement Techniques ...................................................................................... 51 In-vivo Measurements............................................................................................. 52

Final Words ............................................................................................................. 54

LIST OF REFERENCES ............................................................................................... 56

BIOGRAPHICAL SKETCH ............................................................................................ 59

7

LIST OF TABLES

Table page 4-1 Impedance magnitude and phase values of all electrodes at 1KHz, from

largest electrode to smallest ............................................................................... 48

4-2 Impedance magnitude and phase values of all electrodes at 1KHz ................... 48

4-3 Literature values of various parameters to calculate the double-layer capacitance and solution resistance ................................................................... 48

4-4 Theoretical double-layer capacitance, series resistance and fitted frequency- dependent exponent values for different electrode sizes. ................................... 49

8

LIST OF FIGURES

Figure page 2-1 Potential distribution across double layer for various models. A- Helmholtz

theory, B-Guoy Chapman theory, C- Stern theory .............................................. 23

2-2 Randles circuit model ......................................................................................... 24

2-3 Equivalent circuit to represent ideal polarizable electrode behavior ................... 24

2-4 Equivalent circuit to represent low-frequency electrode behavior ....................... 24

2-5 Equivalent circuit to represent high-frequency electrode behavior ..................... 25

3-1 TEENI design. .................................................................................................... 34

3-2 Implant regions of TEENI thread-sets with 3 threads or 4 threads. .................... 34

3-3 TEENI microfabrication process flow. ................................................................. 35

3-4 A microfabricated TEENI device with 4 threads. (Courtesy of JudyLab, UF) ...... 35

3-5 A basic circuit for a potentiostat. ......................................................................... 36

3-6 Experimental set-up for electrochemical impedance spectroscopy (EIS). (Courtesy of E.Patrick, UF) ................................................................................. 36

4-1 EIS experimental set-up to probe the contact pads of TEENI electrodes. (Courtesy of Author) ........................................................................................... 40

4-2 Impedance magnitude and phase plot for each of the electrodes found on a TEENI with three threads ................................................................................... 41

4-3 Impedance magnitude and phase plot of continuity test structure ...................... 42

4-4 Impedance magnitude and phase plot for each of the electrodes found on a TEENI with four threads ..................................................................................... 43

4-5 Comparison of impedance behavior between unbonded and ACA bonded TEENI. ................................................................................................................ 44

4-6 Experimental and simulated impedance data for blocking circuit using Helmholtz capacitance. ...................................................................................... 45

4-7 Experimental and simulated impedance data for blocking circuit using Gouy-Chapman capacitance. ....................................................................................... 46

5-1 Voltage waveform for cyclic voltammetry ........................................................... 53

9

5-2 A CV response for large platinum wire. .............................................................. 54

10

LIST OF ABBREVIATIONS

CV Cyclic Voltammetry

EIS Electrochemical Impedance Spectroscopy

PI Polyimide

PCB

TEENI

Printed Circuit Board

Tissue-Engineered Electronic Nerve Interface

ZIF Zero Insertion Force

11

Abstract of Thesis Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Master of Science

ELECTROCHEMICAL CHARACTERIZATION OF

TISSUE ENGINEERED ELECTRONIC NERVE INTERFACES

By

Sruthi Natt

August 2017

Chair: Jack Judy Major: Biomedical Engineering

Neural prostheses are biomedical systems that restore lost sensory and motor

function after an injury by providing a communication pathway between neural tissue

and prosthetic devices. Despite technological advancements in the field of peripheral

neural interfaces over the years, all existing methods still suffer from critical problems

that limit their performance and/or reliability, e.g., low signal-to-noise ratio, low channel

count, low selectivity, high mechanical stiffness, etc. To fully address the unmet needs

of the amputee population, a new approach is needed.

One strategy to overcome all the major problems limiting the utility and long-term

performance of the neural prostheses is to combine cutting-edge approaches from the

field of MEMS and tissue engineering to develop Tissue Engineered Electronic Nerve

Interfaces (TEENI). This is achieved by integrating polymer based multielectrode

interfaces with high microelectrode density and small cross-sectional area into tissue

engineered scaffolds that facilitate natural regenerative healing process instead of

strong foreign body response evoked by conventional implantation.

With any electrode based neural interfaces, it is essential to have a firm

understanding of the electrochemical mechanisms involved during recording and

12

stimulation. Furthermore, the dependence of these mechanisms on geometry and

materials is vital for optimizing the neural interface design. The impedance of fabricated

electrodes should be managed to achieve good outcomes with both stimulation and

recording. A key challenge is to demonstrate the long-term stability and reliability of

electrodes that are tested under in-vitro conditions over long periods of time.

The objective of this work is to characterize the electrochemical properties of

TEENI electrodes at various stages of fabrication and integration using two main

evaluation techniques: impedance spectroscopy and cyclic voltammetry.

13

CHAPTER 1 INTRODUCTION

This thesis is part of an ongoing project to develop a novel tissue engineered

electronic nerve interface (TEENI) for upper limb amputees. The purpose of this device

is to provide reliable bi-directional interfacing between peripheral nerves and prosthetic

devices for natural sensation and movement.

Amputation is the surgical removal of all or part of a limb that is performed mainly

due to severe injuries, trauma or diabetes[1]. This procedure results in permanent

damage to the nervous system, thus greatly affecting the mobility and quality of life of

the patient. There is compelling evidence that amputees retain significant residual

connectivity and function in their nerves for many years post amputation[2]. In response,

much research has been focused on the development of advanced neural interface and

prosthetics technology, with the ultimate goal of restoring functional disabilities resulting

from limb loss. However, providing amputees with effortless and natural control of

prosthetic limbs is not a near-term solution due to significant challenges. Specifically,

bioelectronic neural interfaces are still limited by issues such as biocompatibility, low

number of independent channels for communication between motor control and sensory

feedback, decline in signal detection and recording abilities with time, relative motion

between implant and tissues and low selectivity to differentiate between nerve types.

Therefore, it is imperative to engineer a reliable neural interface that is scalable to high

channel counts while ensuring biocompatibility and longevity.

One strategy to overcome these challenges is to develop Tissue Engineered

Electrode Nerve Interfaces (TEENI) for the amputee population. The rationale behind

this novel technology is that a natural regenerative healing process can be enabled by

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integrating flexible polymer based multi-electrode “threads” into biodegradable tissue-

engineered hydrogel scaffolds[3]–[5].

The TEENI approach would greatly benefit from advanced understanding of

electrode impedance. When aiming for the best signal-to-noise ratio with maximum

selectivity, it is desirable to record from electrodes that are small enough to be able to

record from a single nerve fiber [6]. However, small electrodes invariably suffer from

higher impedance and hence high thermal noise, ultimately deteriorating the quality of

the detected signal. Since there exists a, trade-off between signal selectivity and

sensitivity, it would be prudent to characterize the electrochemical impedance of any

newly developed electrode interface. Furthermore, recording impedance data to study

the effect of corrosion of electrode sites under accelerated in-vitro soak conditions will

aid in demonstrating their long-term stability[7]. Therefore, a well-characterized

electrode impedance is vital for optimizing its design.

In this work, a model describing the physical processes of electrode-electrolyte

interface for a nearly polarizable electrode was developed. Theoretical equations were

used to calculate the values of solution resistance and the interfacial double layer

capacitance. Electrochemical impedance spectroscopy was used to electrically

characterize the interface of TEENI electrodes of decreasing electrode sizes: 1600, 800,

400 and 200 µm2. The experimental results were compared to the equivalent-circuit

model, thereby confirming the validity of the equations. Since impedance is strongly

dependent on electrode area, cyclic voltammetry was used to electrochemically

determine the active area of electrodes using a ferricyanide solution [8].

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CHAPTER 2 LITERATURE REVIEW

This chapter conveys the necessary information for understanding the

electrochemical mechanisms of the electrode-electrolyte interface (i.e., the region of

contact between an electrode and the electrolyte solution) in biomedical implantable

devices. The concept of double layer capacitance is presented through models of

interfacial capacitance, followed by a brief discussion on the concept of constant-phase

element. Then, the resistive components of the interface are discussed. At the close of

the chapter, a detailed interpretation of Randles circuit model for the electrode-

electrolyte interface is presented with simplified model variations. The work in this

chapter builds upon a large body of literature that describes the theory and operation of

microelectrodes as bioelectric transducers[9]–[12].

Background on the Electrode-Electrolyte Interface

Mechanism of electrical conductivity in the physiological environment involves

ionic charge carriers. Electrodes function as transducers of ionic and electronic signals.

This transduction takes place by the means of non-faradaic effects, wherein there is no

net charge transfer (e.g., capacitive in nature), or by faradaic process in which charge-

transfer reactions do occur (e.g., resistive in nature) [13].

Capacitance Models

When a metal is placed in solution without the application of an external bias, the

interaction between the metal and ions creates a change in their concentration at the

interface. This localized distribution of charge creates an electric potential, called the

half-cell potential, to be established between the interface and solution bulk[11].

16

Incompatibility between charge carriers of metal (i.e., electrons) and electrolyte solution

(i.e., ions) gives rise to a double-layer capacitance.

Many theories have been developed to predict the variation of capacitance with

physical parameters, such as the potential across the interface and the concentration of

the electrolyte. The primary difference between the models is related to the assumption

for how the charges are distributed in the electrolyte.

The double-layer capacitance tends to be extremely sensitive to physical

parameters, such as electrode material, surface geometry and electrolyte composition.

These electrochemical models can serve as a theoretical reference to analyze observed

variations of electrode-electrolyte capacitance.

Helmholtz Model

The first known model of electrode-electrolyte capacitance, which was developed

by Helmholtz, assumes the space-charge layer in the electrolyte near the electrode to

be a two-dimensional double layer[14]. Specifically, the counter-ions in the solution are

assumed to bind to the electrode surface as a single molecular layer [15]. This

assumption allows it to be modeled simply as a constant parallel-plate capacitor. The

capacitance of the Helmholtz layer, denoted by 𝐶𝐻, is given as

𝐶𝐻 =𝜀𝑜𝜀𝑟𝐴

𝑑

(2-1)

with dielectric permittivity of free space 𝜀𝑜, relative permittivity of the electrolyte solution

𝜀𝑟, the area of electrode 𝐴, and the distance between the two layers 𝑑. For accurate

determination of 𝐶𝐻, one is expected to know both the relative permittivity and distance

of the double layer. Dielectric constant of a single molecular layer is influenced by the

orientation of the molecules dipole and will differ from the commonly quoted value for a

17

bulk solution. The distance between the double layers is often assumed to be the radius

of the solvated ions, but this too varies with surface phenomena (e.g., adsorption).

Roughly estimating the dielectric constant to be 80 for saline and separation distance of

0.5 to 1.0 nm, 𝐶𝐻 is about 140 μF/cm2 [10].

A key weakness of the Helmholtz model is in its assumption that 𝐶𝐻 has a

constant value. It is known that as applied electrode potential across the electrode-

electrolyte interfaces increases, the solvated ions pack in more closely to the metal,

thereby shrinking the double-layer thickness [9]. In order to predict the variation of

capacitance with voltage, a more sophisticated model is needed.

Gouy-Chapman Model

A more detailed and dynamic model of the capacitance of the double layer

model, which was developed by Gouy-Chapman, assumes the charge density of the

space-charge layer in the solution to be three dimensional and takes into account the

diffusion of ions near the electrode [16]. Since the ions in this model are assumed to be

point-charges, their size is neglected. The Poisson-Boltzmann equation is used to

describe the distribution of potential and charge in the double layer[15]. The “diffuse

layer” is modeled as having a thickness 𝐿𝐷 given by

𝐿𝐷 = √𝜀𝑜𝜀𝑟𝑘𝑇

2𝑁𝑖𝑧2𝑞2

(2-2)

with Boltzmann constant 𝑘, temperature 𝑇, ionic concentration of the electrolyte

𝑁𝑖, valency of the ion species 𝑧, and charge of an electron 𝑞. From equation (2-2), it is

18

evident that the double layer thickness will be greatest for the most dilute electrolyte.

Gauss’s law is used to derive the capacitance of the diffuse space-charge layer[17].

𝐶𝐺 = 𝜀𝑜𝜀𝑟𝐴

𝐿𝐷𝑐𝑜𝑠h(

𝑧𝑉𝑞

2𝑘𝑇)

(2-3)

The Gouy-Chapman model yields a value of capacitance that varies more

sharply with applied potentials than do experimental measurements. In reality, solvated

ions cannot get infinitely close to the electrode, as they will be separated by a distance

equal to their radius[11].

Stern Model

A model that combines the Helmholtz and Gouy-Chapman was developed by

Stern[18]. There is a compact plane of high electric field immediately adjacent to the

electrode that firmly holds counter-ions to the electrode. Beyond this plane is the Gouy-

Chapman diffuse layer. The electric potential decreases in an exponential fashion with

the distance from the electrode, and decays to zero in the bulk solution. The effective

double-layer is modeled as a series arrangement of the Helmholtz and Gouy-Chapman

capacitances, as given by

1

𝐶𝐷𝐿 =

1

𝐶𝐻+

1

𝐶𝐺

(2-4)

Figure 2-1 illustrates the variation of voltage across the double layer for each

model described above.

While this model, too, in reality, does not perfectly replicate the interfacial

system, it is most suitable for modeling the electrodes in biomedical systems[9].

19

Constant-Phase Element

The electrical double layer does not behave as a pure capacitor, rather, it

exhibits non-linear behavior especially at high current densities[19]. This phenomenon

is modeled by constant phase element, which can be expressed mathematically by:

𝑍𝐶𝑃𝐸 = 1

(𝑗𝜔)𝛼𝑄

(2-5)

with magnitude of 𝑍𝐶𝑃𝐸 𝑄, constant representing the surface inhomogeneities α, and

angular frequency 𝜔=2πf. The constant-phase element behavior can range from purely

resistive (α= 0) to purely capacitive (α= 1). For α= 0.5, the constant-phase element is

called a Warburg element[20][21].

The physical basis for an electrode to behave as a constant phase element is not

fully obvious and is in fact still a topic of research[19]–[21]. Some researchers attribute

the physical basis to surface roughness, while others think it is due to non-uniformities

at the atomic level.

Series Resistance

In electrochemical impedance spectroscopy, impedance is purely resistive if its

phase angle is zero. The resistive term corresponds to the real part of the circuit

impedance. In general, resistance depends on resistivity and geometric aspects of the

material. In case of metals, the electrical resistivity is very low. Therefore, the source of

series resistance is primarily the passage of current through the electrolyte.

Electrolyte spreading resistance, denoted by 𝑅𝑠, is the resistance encountered

beyond the electrode surface when there is ionic current flow in the electrolyte due to

applied electric potential. If the return electrode is far away, this resistance is dependent

20

only on the geometry of the electrode and electrical conductivity of the electrolyte

solution. Analytical expression for solution resistance for a simple planar disk electrode

was derived by Neuman [15]. The expression for the series resistance of a planar disk

electrode is given by:

𝑅𝑠 =𝜌

4 ∗ 𝑟

(2-6)

with solution resistivity 𝜌 and electrode radius 𝑟. From the above expression, it is seen

that solution resistance is inversely proportional to electrode radius and linearly

proportional to the solution resistivity.

Charge-Transfer Resistance

Electrochemical reactions allow charge transfer between electrode and

electrolyte. In the case of faradaic interfaces, electrochemical reactions take place that

result in a DC current path across the interface[22]. Any reaction that obeys Faraday’s

law, which states that the amount of reaction at the electrode surface is directly

proportional to the amount of current passing through it, is considered faradaic. The

magnitude of this current depends on reaction kinetics and diffusion of ionic reactants

towards or away from the electrode. A charge-transfer resistance can be used to model

the physical hindrance encountered by electrons when they move across the interface.

This charge-transfer resistance, in simplest case, can be derived from the equation that

best describes the kinetic relationships in an electrochemical process, (i.e., Butler

Volmer equation)[15]. It is expressed as:

𝑖 = 𝑖𝑜 exp (𝛼𝑎𝐹

𝑅𝑇𝜂) − exp (−

𝛼𝑐𝐹

𝑅𝑇𝜂)

(2-7)

21

with exchange current density 𝑖𝑜, overpotential 𝜂, coefficients for anodic 𝛼𝑎 and

cathodic reactions, 𝛼𝑐. The Butler-Volmer equation denotes that when the cell is not in

equilibrium, the net current flow is the difference of anodic and cathodic currents.

When the overpotential is small, the Butler Volmer current can be approximated

as a linear function.

If 𝛼𝑎 = 1 − 𝛼𝑐, Equation (2-7) reduces to:

𝑖 = 𝑖𝑜

𝜂𝐹

𝑅𝑇

(2-8)

In this case, the charge-transfer resistance can be expressed as:

𝑅𝑐𝑡 =𝑅𝑇

𝑖𝑜𝐹

(2-9)

Randles Model

Although there is active research on new electrode materials for fuel cells,

supercapacitors, and biomedical sciences, the physical basis of electrode models

remains relatively unchanged. In fact, models proposed in the 1900’s are still being

used to understand the mechanism of electrode-electrolyte interface. In early work,

Randles developed a simple mathematical circuit model that could predict the

impedance behavior of mercury electrodes. (Figure 2-2) [23][24]. The circuit elements of

Randles model are double layer capacitance 𝐶𝑑𝑙, charge transfer resistance 𝑅𝑐𝑡,

solution or spreading resistance 𝑅𝑠 and Warburg impedance 𝑊. All the circuit elements

were discussed in the previous section.

22

Variations of Randles Model

Although Randles model looks relatively simple, it powerfully attempts to model

capture both the chemical and mass transport behavior in the electrical domain. Each

element in the model represents a different electrode process. Under certain

experimental conditions, this model can be simplified to retain only those elements that

correspond to the dominant electrochemical phenomena.

Polarizable Electrode Behavior

An ideally polarizable electrode, which is characterized by the absence of net

charge transfer between the electrode interfacial double layer, behaves like a capacitor

[15], [22]. The solution resistance encountered by the current spreading out in the

solution is modeled as a series resistor, as described earlier in this section. All the other

elements of Randles circuit that denote faradaic current can be assumed to have no

effect on a polarizable electrode and hence can be removed from the model.

The equivalent circuit to represent polarizable electrode behavior is represented

by Figure 2-3. A constant-phase element is used in place of a pure capacitor to account

for surface inhomogeneities.

Low-Frequency Behavior

If the electrode is operated at a low frequency range, it is possible to ignore one

or more model elements. Since capacitance and frequency have an inverse

relationship; at low enough frequencies, the double-layer capacitor can be replaced by

an open circuit. The impedance of Warburg element increases with 𝜔−1/2; the inverse

square root of angular frequency, 𝜔 [8]. The equivalent circuit model for low-frequency

behavior of electrodes is illustrated in Figure 2-4.

23

High-Frequency Behavior

At high enough frequencies, mass transport will no longer be a limiting

factor[15]. As a result, the Warburg element can be ignored and replaced by a short

circuit. The equivalent circuit for this case is illustrated by Figure 2-5.

Traditionally, electrodes for biomedical applications are characterized at a

frequency of 1 KHz, since it matches the dominant frequency component of neural

action potentials [7].

Since at high frequencies, diffusing reactants do not have to move very far, the

Warburg impedance will have little to no effect on the electrode-electrolyte interface.

Under this operating condition, the effect of Warburg impedance is less pronounced and

can be neglected without losing accuracy.

Figure 2-1. Potential distribution across double layer for various models. A- Helmholtz theory, B-Guoy Chapman theory, C- Stern theory

24

Figure 2-2. Randles circuit model

Figure 2-3. Equivalent circuit to represent ideal polarizable electrode behavior

Figure 2-4. Equivalent circuit to represent low-frequency electrode behavior

25

Figure 2-5. Equivalent circuit to represent high-frequency electrode behavior

26

CHAPTER 3 EXPERIMENTAL METHODS

To optimize neural interface designs, experiments are needed to understand the

dependence of electrochemical mechanisms on geometry and materials A related

challenge is that, electrochemical measurements can be highly sensitive to ambient

experimental conditions and electrode fabrication techniques. The design and

fabrication process of the TEENI electrodes, the electrochemical instrumentation, and

the analysis techniques are discussed in this chapter.

TEENI Design and Layout

As shown in Figure 3-1, each TEENI thread-set was designed to have an implant

region consisting of the following: a set of individual polyimide threads with four metal

micro-electrodes per thread, a large polyimide wing on both ends of the implant region,

a 1-mm-diameter hole in each polyimide wing for convenience during handling, four

200-μm-diameter suture holes along the polyimide lead-body to secure thread sets in

the implanted position, a large reference electrode outside the implant region to

attenuate common EMG signals surrounding the implant, and electrode contact pads to

integrate with external PCB for data acquisition [3].

Two different implant regions were designed for TEENI. The thread-sets of the

implant region had either three or four threads in a thread set, as shown in Figure 3-2.

Each thread was 10 μm thick, 86 μm wide, and 5-mm-long with a 170-μm-wide edge-to-

edge gap between threads in a thread-set. Each thread of a thread-set had 4 recording

electrodes, each with a different surface area: 200, 400, 800, and 1,600 μm2. The

center-to-center spacing between individual electrodes was 350 μm. Thus, a 3-thread

TEENI device has 12 recording electrodes and the 4-thread-set TEENI device has 16

27

recording electrodes. To evaluate the effect of longitudinal recording surfaces in

comparison with conventional circular electrodes, the recording electrodes also varied in

shape from circular to elliptical. In addition to recording electrodes, the 3-thread TEENI

had an on-thread reference electrode and a stimulation electrode with surface areas of

16,000 μm2 and 3,200 μm2 respectively. Also, test structures were included in the 3-

thread TEENI design to evaluate the electrical continuity of the metal traces and

dielectric integrity of the polyimide layer. The metal traces from the electrodes and test

structures, which had a width of 6 μm, led to a 20-pad connector array.

TEENI Fabrication

The TEENI microfabrication process, which is illustrated in Figure 3-3, began

with a 100-mm-diameter single-crystal silicon wafer, which was used as the carrier

substrate (Fig. 3-3A). A bottom structural polyimide (PI) layer was spun onto a thickness

of 5 μm and cured in an inert atmosphere (Fig. 3-3B). A layer of photoresist (AZ5214)

was photolithographically exposed, image reversed, and developed to obtain inward-

leaning retrograde sidewall slopes (Fig 3-3C). An oxygen plasma etch was

subsequently performed to remove any residual photoresist and to roughen the top of

the first PI layer. To promote adhesion between PI film and the following metal stack, 50

nm of Ti was deposited by sputtering (Fig. 3-3D). Immediately after deposition, the

desired metal layer stack of Pt/Au/Pt (100 nm each) was deposited onto the wafer using

an electron-beam evaporator (Fig. 3-3E). The top of the metal stack was then coated

with Ti (50 nm) to promote adhesion of Pt to the upcoming second PI layer (Fig. 3-3F).

The photoresist was removed to lift-off all the layers deposited on it, leaving behind the

patterned stack of metal (Fig. 3-3G). A second PI layer of 5 μm was spun on to the

wafer to mitigate any stress variations between dissimilar layers (Fig. 3-3H). A thick

28

positive photoresist layer (AZ9260) was then patterned to define the contacts and the

connector pads (Fig. 3-3I). Reactive ion etching (RIE) was used to etch through the

exposed PI and capping Ti layer to reveal the surface of Pt (Fig. 3-3J). The thick

photoresist layer was then stripped with acetone (Fig. 3-3K). A second thick photoresist

layer was patterned to define the shape of the TEENI device. After a second REI step,

the photoresist was stripped in acetone (Fig. 3-3L) and the individual TEENI samples or

thread-sets were mechanically peeled off the wafer. An example of a TEENI device

resulting from this microfabrication process is shown in Figure 3-4.

PCB-TEENI integration

A PCB-based integration system was used to relay neural signals from the

implanted electrodes to an external instrumentation system for data processing. In

collaboration with Tucker Davis Technologies, a manufacturing process was developed

to integrate the released TEENI thread-sets onto a custom-made PCB using conductive

silver-epoxy. The TEENI-PCB assembly was attached to a zero-insertion force (ZIF)

connector via a soldered wire bundle.

Specifically, custom-designed jigs were machined to facilitate the mechanical

alignment and subsequent assembly of the TEENI device with the PCB. The PCB was

positioned using 1-mm-diameter pins on a base jig, while the wires were soldered to the

through-holes. Thereafter, conductive silver epoxy was precisely dispensed on the

circular contact pads of the PCB using a Nordon Asymtek dispensing system. The

TEENI thread-set was positioned on another jig, which was then flipped over to mate

with the base jig. The TEENI-PCB-jig assembly was cured in an oven at 150°C.

29

To facilitate in-vivo operation of the TEENI-PCB assembly with the soldered wire

bundle hub, the assembly was potted with medical grade epoxy (Dymax Corporation,

USA). Figure 3-5 shows the final TEENI-PCB assembly.

The bonding of the TEENI with the PCB was a challenging aspect of the entire

device integration process. Although the use of a flip-chip bonder allowed for precise

lateral alignment of opposing electrodes, the application of finite, spatially resolved

conductive epoxy was difficult due to the narrow lateral spacing (250 μm) between

adjacent electrodes. Electrical shorts were common due to contact between epoxied

pads leading to a low yield of 100% isolated channels. Furthermore, the use of

conductive epoxy produced trapped air pockets between bonded pads that could

commonly result in shorting after implantation due to the penetration of water. Although

underfill can be used to electrically isolate channels in such situations, we wanted a

better alternative.

Anisotropic conductive adhesive (ACA), such as that produced by Creative

Materials Inc, was studied due to its reported advantage over conductive pastes and

underfill, particularly in situations involving fine-pitch contact pads. Since ACA materials

provide superior vertical conduction, they have been used in a variety of applications

with fine-pitch contact pads (e.g., liquid-crystal display (LCD) manufacturing and

electrical attachment of surface mounted devices) [25].

Unfortunately, initial tests on ACA-bonded TEENI-PCB assemblies showed low

channel conductivity, with only about 40% of all channels shown to be electrically

conductive. Upon analysis, geometric constraints associated with TEENI, PCB, and

conductive particles within the ACA were believed to be the cause of the low functional

30

yield. Optical profilometry showed that the PCB pads were recessed by 4 to 6 μm below

the solder mask. It was observed that the conductive particles of ACA were too small

(~2.5 μm diameter) to reliably bridge the vertical gap between the recessed pads on

both the TEENI and the PCB.

A new ACA variant with larger conductive particle size (10 to 15 μm diameter)

was used to ensure reliable electrical conduction with high-channel count. To further

increase electrical reliability, the contact pads of the PCB were bumped up so that they

protrude from the surface. This was done by electroless plating nickel and gold on top of

the PCB pads to build up 18 μm of metal, which results in pads that extend > 10 μm

above the surface of the adjacent solder mask. Experimental results from the bonding

process using bumped-up contact pads are discussed in the following chapter.

Experimental Set-up

This section describes the electrochemical cells, instrumentation and

measurement equipment used for the process of data collection and interpretation.

Electrochemical Cell

Since electric potential is not an absolute quantity, the electrical behavior of the

working electrode (i.e., the electrode of interest) cannot be measured in isolation. For

electrodes in solution that interface between electronic and ionic conduction regions,

another electrode is needed to complete the electrochemical circuit through the

electrolyte. This secondary electrode is called the auxiliary or counter electrode.

Experiments with a working and counter electrode are called two-electrode cells.

Counter electrodes are designed so that they can source or sink the current

needed without causing a change in the potential at the interface. To achieve this,

counter electrodes have a large surface area to minimize the current density at their

31

surface. The counter electrode must also be corrosion resistant and inert. Platinum,

being a noble metal, is the common choice for counter electrode as it meets all the

requirements.

A three-electrode cell is formed by including a reference electrode, so that the

potential of the working electrode can be more accurately monitored. The reference

electrode has a stable and well-defined electrochemical potential that remains constant

during the measurement [15]. Reference electrodes are carefully selected so that their

equilibrium state is maintained over long time periods.

Common reference electrodes include saturated calomel electrodes, silver-silver

chloride electrodes, and normal hydrogen electrodes. Calomel electrodes are based on

the reaction between elemental Hg and Hg2Cl2. Ag/AgCl electrodes operate on

reduction reaction of silver chloride to form silver metal and chloride ion. The standard

reference for measuring redox potential is the normal hydrogen electrode (NHE).

However, due to difficulty of its operation, NHE is rarely used and hence all other

electrodes are referenced to this standard.

Instrumentation

Electrochemical measurement techniques can either be current-controlled or

voltage-controlled. In the case of current-controlled experiments, a galvanostat applies

a well-defined current waveform to the working electrode and the resulting potential is

recorded. For voltage-controlled techniques, a potentiostat applies a well-defined

potential waveform to the working electrode and the resulting current is recorded. Since

all the electrochemical measurements described in this work were done in a voltage-

controlled mode, the basic construction of potentiostat will be described next [15].

32

A potentiostat is used to control the potential between the working and the

reference electrode, while the current flow between the working and the counter

electrode is recorded. A simple potentiostat circuit is shown in Figure 3-6. There are two

operational amplifiers in this circuit: a voltage follower and an inverting amplifier. Ideally,

no current flows in and out of the inputs of an op-amp.

The first op-amp controls the current flow through the counter electrode, while

maintaining the applied potential Vbias between the reference and the working electrode.

The second op-amp converts this current to a voltage, Vout.

It is easy to build a simple potentiostat. However, a commercial instrumentation

is valuable to obtain accurate measurements.

Data Acquisition and Analysis

A Metrohm Autolab PGSTAT128N with the frequency response analysis module

(FRA32) was used for performing electrochemical impedance spectroscopy (EIS)

experiments. Impedance responses of the TEENI microelectrodes were measured with

respect to a large Pt wire as the counter electrode and a Ag/AgCl reference electrode in

0.01M phosphate buffered NaCl (Sigma) at room temperature (pH 7.4). The potentiostat

has a system limit of 10nA at low frequencies. To avoid stability issues during

measurement, the frequency scan range was limited to 10-2 to 105 Hz. The perturbation

voltage was 10mV. Simple Matlab scripts were used to plot the results such as

impedance magnitude, phase and Nyquist graphs from the experiments. Most of the

analysis was later done on Excel.

Interpreting the behavior of an electrode-electrolyte interface from the

experimental EIS data is extremely challenging without invoking the physical basis of

33

electrochemistry and making justifiable assumptions. Modeling the data will be helpful in

understanding the experimental results.

Electrochemical Impedance Spectroscopy

Electrochemical impedance spectroscopy (EIS) is a technique that measures the

response of an electrode to a sinusoidal voltage or current applied at different

frequencies [26], [27]. The mathematical approach of this measurement technique is

based on the fact that-, impedance is the frequency dependent resistance to the flow of

current as a function of an applied voltage or current. A Typical experimental set-up is

as shown in Figure 3-7[22]. During the experiment, a potential is applied between the

reference and the working electrodes, while the current flow from the working electrode

to the counter electrode is measured.

For meaningful mathematical analysis of EIS results, it is important to ensure that

the perturbation signal is small enough to elicit a linear current-voltage response. Yet,

the signal must also be large enough to be detected by the instrumentation. Therefore,

the amplitude of the signal must be adjusted to achieve the best possible compromise

between signal-to-noise ratio and linearity.

The results from impedance spectroscopy are typically displayed as bode plots

and Nyquist plots. In a bode plot, the magnitude and phase of the impedance are

plotted against frequency. In a Nyquist plot or complex plane plot, the imaginary part of

the impedance is plotted against the real part. Nyquist plots often reveals the model

parameters that dominate the behavior of the electrode-electrolyte interaction.

34

Figure 3-1. TEENI design.

Figure 3-2. Implant regions of TEENI thread-sets with 3 threads or 4 threads.

35

Figure 3-3. TEENI microfabrication process flow.

Figure 3-4. A microfabricated TEENI device with 4 threads. (Courtesy of JudyLab, UF)

36

Figure 3-5. A basic circuit for a potentiostat.

Figure 3-6. Experimental set-up for electrochemical impedance spectroscopy (EIS).

(Courtesy of E.Patrick, UF)

37

CHAPTER 4 RESULTS AND DISCUSSION

The goal of the work described in this thesis was to characterize the

microfabricated electrodes in the tissue-engineered electronic nerve interfaces and to

compare the experimental results with simple circuit models that represent the physical

basis of the electrode-electrolyte interface. Using the methods described in the Chapter

3, electrochemical impedance spectroscopy (EIS) data was collected from different

TEENI thread-sets at different stages of microfabrication and integration processes.

Simple models were then developed to enable meaningful interpretations of the EIS

data and to draw conclusions.

The experimental set-up used to capture the EIS data is as shown in Fgure 4-1.

The implant region of the TEENI was soaked in a phosphate-buffered saline solution

(PBS) while the connector pads of the TEENI remained well outside the solution and

dry. A micro-positioner with a tungsten probe tip was used to make sequential electrical

contact with the pads on the TEENI. Leads integrated into the TEENI device connected

each pad to a corresponding microelectrode. A potentiostat was used to probe the

TEENI connector pads and measure their impedance. The reference electrode was an

Ag/AgCl electrode and the counter electrode was a platinum wire. EIS measurements

were carried out in a faraday cage to reduce external electrical noise.

EIS of TEENI with 3 Threads

As mentioned in Chapter 3, the TEENI with three threads has a total of 12

recording electrodes, one on-thread reference electrode, one stimulation electrode a

pair of interdigitated electrodes used to evaluate the dielectric integrity, a continuity test

structure, and one large EMG reference electrode outside the implant region. All EIS

38

measurements of recording electrodes that had the same surface area were combined

to compute the mean and standard deviation of the channel impedance. Bode

magnitude and phase plot are shown in Figure 4-2, with error bars representing one

standard deviation of the mean. Table 4-1 shows impedance magnitude and phase

values of all the microelectrodes at a frequency of 1 KHz.

Since microelectrodes reach steady-state more quickly than larger electrodes,

they can be very useful for electrochemical sensing[28]. All TEENI microelectrodes

show a primarily capacitive behavior even at the highest measurement frequency

(100KHz). This result is consistent with the fact that the electrode charging time

constant, τ, is proportional to electrode radius.

The impedance spectra of the continuity test structure, which was sandwiched

between the polyimide layers, was measured and shown to be purely capacitive, as

expected (Figure 4-3) [29].

EIS of TEENI with 4 threads

TEENI with four threads have a total of 16 recording electrodes and a continuity

test structure. Results from a single device are shown in Figure 4-4. Table 4-2 lists the

impedance magnitude and phase values of all electrodes at the frequency of 1 KHz.

ACA Bonded TEENI-PCB Assembly

The TEENI connector pads were align-bonded with the pads on the PCB to

establish direct electrical connections between the electrodes and wires leading to the

outside the body. Assembled TEENI were also characterized by EIS using the same

method as described above, but the PCB vias were probed instead of the TEENI

contact pads. Figure 4-5 shows a bode plot for impedance magnitude that compares the

recording electrodes of ACA bonded TEENI vs unbonded TEENI. It can be observed

39

from the plots that the bonding process does not appreciably change the electrical

characteristics of the TEENI.

Equivalent Circuit Modeling

As mentioned in Chapter 2, it is best to start by modeling the electrochemical

results using an equivalent circuit that is conceptually simple and relevant to the

frequency range of interest. Gradually, this simple model can be improved by further

building up a more detailed circuit.

In case of TEENI recording electrodes, the Randles circuit model and its variants

is a suitable place to begin to see how the experimental data compares to the model

parameters built using theoretical equations.

Kovacs proposed that neural recording electrodes are operated in linear or small-

signal mode which involves capacitive properties, while the operation of stimulating

electrodes occurs in large-signal mode with appreciable current flow[9]. Therefore, to

characterize the platinum recording microelectrodes of TEENI in saline, the model for a

blocking system with a constant-phase element was assumed[22].

Values from literature are adopted for platinum as electrode material and

physiological saline as electrolyte as shown in Table 4-3 to calculate the theoretical

values of the double layer capacitance, 𝐶𝑑𝑙 and solution resistance, 𝑅𝑠 [8] , [10].

Theoretical values of 𝑅𝑠 and 𝐶𝑑𝑙 are calculated for circular recording electrodes of

different areas and listed in Table 4-4. Experimental results for recording electrodes of

different surface areas are compared to the corresponding theoretical results. Assuming

that the voltage across the double layer is equal to the applied voltage, Stern model

gives the best fit.

40

Figure 4-1. EIS experimental set-up to probe the contact pads of TEENI electrodes.

(Courtesy of Author)

41

Figure 4-2. Impedance magnitude and phase plot for each of the electrodes found on a TEENI with three threads

42

Figure 4-3. Impedance magnitude and phase plot of continuity test structure

1.E+03

1.E+04

1.E+05

1.E+06

1.E+07

1.E+08

1.E+09

1.E+10

1.E+01 1.E+02 1.E+03 1.E+04 1.E+05

Imp

edan

ce, O

hm

Frequency, Hz

Impedance Magnitude Plot

0

20

40

60

80

100

120

140

1.E+01 1.E+02 1.E+03 1.E+04 1.E+05

-Ph

ase,

deg

rees

Frequency, Hz

Phase Plot

43

Figure 4-4. Impedance magnitude and phase plot for each of the electrodes found on a

TEENI with four threads

44

Figure 4-5. Comparison of impedance behavior between unbonded and ACA bonded

TEENI.

45

Figure 4-6. Experimental and simulated impedance data for blocking circuit using Helmholtz capacitance.

46

Figure 4-7. Experimental and simulated impedance data for blocking circuit using Gouy-

Chapman capacitance.

47

Figure 4-8. Experimental and simulated impedance data for blocking circuit using Stern capacitance.

48

Table 4-1. Impedance magnitude and phase values of all electrodes at 1KHz, from largest electrode to smallest

Electrode Area (μm2) Impedance (Ω) -Phase (degrees)

EMG Ref (160,000) 18 KΩ 57.8

On-thread Ref (16,000) 62 KΩ 63.0

Stimulation Electrode (3200) 306 KΩ 65.1

1600 575 KΩ 65.6

800 1.1 MΩ 66.6

400 2 MΩ 66.1

200 3.7 MΩ 66.6

Table 4-2. Impedance magnitude and phase values of all electrodes at 1KHz

Electrode Area (μm2) Impedance (Ω) Phase (degrees)

1600 5.59E+05 71.03

800 1.10E+06 71.79

400 2.23E+06 73.55

200 4.74E+06 76.75

EMG Ref- 160,000 7.09E+03 55.84

Table 4-3. Literature values of various parameters to calculate the double-layer capacitance and solution resistance

Parameter Notation Value

Absolute Permittivity 𝜀𝑜 8.85 p𝐹/𝑚

Relative Permittivity 𝜀𝑟 79.4

Outer Helmholtz Distance 𝑑𝑂𝐻𝑃 0.5 nm

Valency 𝑧 4

Boltzmann Constant k 1.38E-23 𝑚2𝐾𝑔𝑠−2𝐾−1

Temperature 𝑇 295 𝐾

Elementary Charge 𝑞 1.6E-19 𝐶

Ionic Concentration of Saline

𝑁𝑖 9.27E+25 𝑖𝑜𝑛𝑠/𝑚3

Resistivity ρ 0.72 Ω m

49

Table 4-4. Theoretical double-layer capacitance, series resistance and fitted frequency- dependent exponent values for different electrode sizes.

Electrode Area (μm2)

Helmholtz Capacitance

(pF)

Gouy-Chapman

Capacitance (pF)

Stern Capacitance

(pF)

Solution Resistance

(kΩ)

Frequency dependent exponent

(α)

1600 22.5 76.9 17.3 8.0 0.8

800 11.2 38.4 8.69 11.2 0.8

400 5.6 19.2 4.34 15.9 0.8

200 2.8 9.6 2.17 22.5 0.8

50

CHAPTER 5 CONCLUSION AND FUTURE WORK

The overall goal of this thesis was to better understand the electrochemical

behavior of microfabricated electrodes for use in upper-limb prosthetic devices. Despite

the many experimental trials, theoretical calculations and model development, there is

much to be explored. This closing chapter will discuss the implications of the results and

highlight some important experiments that could be pursued in follow-on work.

EIS Data and Modeling

As previously described, it is challenging to interpret the electrochemical

behavior of microelectrodes from experimental EIS data without a basic theoretical

background in electrochemistry. To support the design and fabrication of

microelectrodes, accurate equivalent electrical circuit models are necessary. Using an

imperfect circuit will only lead to difficulty in interpretation of model parameters. Fitting

programs are available with many electrochemical instrumentation software systems

and are simple to use. However, the resulting circuit models tend to do a poor job in

predicting the impedance parameters for various electrode sizes. The major reason for

this is the non-linear behavior of the microelectrodes. One way to overcome this

challenge is to simply limit the frequency range of analysis.

The model of blocking system with constant-phase-element behavior, which is

represented as a resistor in series with a constant-phase element, best describes the

characteristics of the recording electrodes in the TEENI device over the frequency

range of interest in most nerve-interface applications (I.e., 10 Hz to 10 KHz). Outside of

this frequency range, the observed trend in impedance could not be explained by this

model. To account for this behavior, future work will be needed to understand the cause

51

of these non-linearities and to develop new circuit elements and/or models. One

possibility is to use a simplified non-uniform model that can approximate the

inhomogeneities of the electrode.

Finite-element-modeling (FEM) techniques can be used to simulate and/or

analyze microelectrode behavior. Commercial Multiphysics FEM software packages,

such as COMSOL, can be used with relevant boundary conditions to perform qualitative

analysis of unique geometries.

Measurement Techniques

Impedance spectroscopy was the only measurement technique employed to

characterize the electrochemical behavior of TEENI recording electrodes. Another

important technique that can provide information on electrode stability would be cyclic

voltammetry (CV). Reaction rates, adsorption process, and the nature of electrode

reactions can also be studied using this technique.

In cyclic voltammetry, the potential of the working electrode with respect to the

reference electrode is swept cyclically at a constant rate between two potential limits

and the current flow between the working and the counter electrode is measured. A

typical CV voltage waveform is shown in Figure 5-1. The current resulting from the

applied CV waveform is recorded and plotted against applied potential.

A CV experiment with a platinum wire with known geometric area as the working

electrode was recorded with a scan rate of 50mV/s and using the three-electrode set up

with potassium nitrate as the electrolyte. The resulting graph, which is shown in Figure

5-2, illustrates the characteristics of a reversible redox reaction [15].

One unusual, yet highly informative application of CV is to determine the active

electrochemical surface area of the working electrode. In the case of TEENI

52

microelectrodes, the active area could be different from the as-fabricated electrode area

due several factors (e.g., surface roughness, contamination, deposition of resistive

oxide film, metal corrosion and delamination of underlying polyimide film). Therefore,

finding the true electrochemically active area of electrode can be an indication of any of

the above phenomenon. Since impedance is strongly dependent on active electrode

area, this method also could reveal reasons for any unusual electrode behavior

observed.

To determine the active electrochemical area of electrodes using CV, it is

common to use a standard ferricyanide solution of known concentration at a well-

defined scan rate [8][30].

The relationship between peak current and electrode area is given by

𝐼𝑝 = 2.69 ∗ 108𝑛3/2𝐴𝐷1/2𝐶𝑣1/2, (5-1)

With peak current 𝐼𝑝, number of electrons involved in the redox reaction 𝑛, electrode

area 𝐴, diffusion coefficient of the electrolyte solution 𝐷,concentration of the electrolyte

solution 𝐶, and scan rate 𝑣.

In-vivo Measurements

All the electrochemical measurements listed in this thesis were performed in-vitro

with 0.01M phosphate buffered saline as the electrolyte solution. In this work and on-

going work, the chronic performance of microelectrodes was characterized by

performing long-term soak tests at elevated temperatures [31].

However, in the physiological environment, the same experiment could yield

different results. To study degradation rates, microelectrodes could be tested in an

53

accelerated aging system with hydrogen peroxide, in an attempt to mimic the

aggressive response of the body on the implanted device [7]. Although creating such a

system would require a new more complex experimental set-up, it may provide

invaluable information to change the design and microfabrication process in order to

improve the robustness of the TEENI technology.

Figure 5-1. Voltage waveform for cyclic voltammetry

54

Figure 5-2. A CV response for large platinum wire.

Final Words

The ultimate goal of this project is to develop a novel nerve interface that is

scalable to high independent sensory and motor channels, which can serve the needs

of the upper-limb amputee. The focus of this thesis is on the modeling and

characterization, and improvement of tissue-engineered electronic nerve interfaces. To

design the interface and assess its performance, a detailed understanding of electrode

properties and accurate models of their behavior are needed.

The thesis progressed in three major steps. First, it compared models for

electrode behavior in an electrolyte. Based on published observations and its expected

behavior and limitations, the Randles model was chosen as an appropriate starting

point. Second, the physical basis of the electrode-electrolyte interface and its

parameters were explored and electrochemical interface theories were used to predict

the observed values. Under the assumption that the applied electrode potential is

55

approximately equal to the capacitive double layer potential, the Stern model was used

to best explain the resulting electrochemical behavior. Lastly, the electrochemical

impedance of the microfabricated electrodes were experimentally measured and

compared with theoretical values. Electrodes of various sizes were investigated (i.e.,

ranging from 200 to 160,000). Results from short-term impedance spectroscopy

experiments showed a clear and predictable dependence of impedance magnitude and

phase on microelectrode size and analysis frequency. However, accelerated long-term

soak tests demonstrate that significant changes in electrode impedance occur that are

not yet predicted by available models. Future work should examine developing more

accurate models of accelerated long-term soak tests, which can help determine the

failure mechanisms, and then ultimately leads to improvements to achieve better and

more stable long-term performance.

56

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59

BIOGRAPHICAL SKETCH

Sruthi Natt was born in 1994 in Chennai, India to Sriram Natt and Sai Lakshmi

Sriram. She moved to Bangalore, India and attended The Oxford Senior Secondary

School where she graduated in 2011. She was interested in engineering, and pursued a

degree in instrumentation technology while attending Visvesvaraya Technological

University in India. She graduated as the Best Outgoing Instrumentation Engineer in

May 2015. Her desire to become a Biomedical Engineer brought her to University of

Florida, Gainesville, in August 2015. She became a research assistant with the

Department of Electrical and Computer Engineering in December 2016 where she

began working under Dr. Jack Judy. Sruthi was awarded her master’s degree in August

2017.