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Strategies for improving mechanical and biochemical interfaces between medical implants and tissue

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Strategies for improving mechanical and biochemical interfaces between medical implants and tissue

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STRATEGIES FOR IMPROVING MECHANICAL AND BIOCHEMICAL
INTERFACES BETWEEN MEDICAL IMPLANTS AND TISSUE

by

Curtis Dixon Lee

A Dissertation Presented to the
FACULTY OF THE USC GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the
Requirements for the Degree
DOCTOR OF PHILOSOPHY
(BIOMEDICAL ENGINEERING)

August 2015

Copyright 2015 Curtis Dixon Lee
i

To Jamie

ii

ACKNOWLEDGEMENTS
In addition to the joys of discovery and collaboration, the Ph.D. process can be a
long and challenging process. I am grateful for those who joined me in my successes and
supported me in my challenges.
First, I would like to thank Dr. Ellis Meng who has been an inspirational mentor
and advisor. Her incredible work ethic, intelligence, and organizational skills have
motivated me to be a better researcher and showed me how to think clearly about
research goals and methods. Her enthusiasm for MEMS and implantable devices is
contagious.
I would also like to thank my committee members and collaborators. Dr. James
Weiland and Dr. Malancha Gupta of my doctoral committee made valuable comments
and suggestions as I discussed my research with them. From my qualification exam
committee, Dr. Andrea Hodge gave me a material science perspective on my work (as
well as access to her profilometer), and Dr. Dong Song taught me about neuroengineering
both as my committee member and as a professor whose excellent neuroengineering class
I had the opportunity to be a TA for. I also wish to thank Dr. Victor Pikov of the
Huntington Medical Research Institute for performing our animal surgeries and sharing
his great knowledge of electrophysiology. Additionally, I’d like to thank Dr. Jerry Loeb
who taught me how to think about neural recording and stimulation in his
electrophysiology course, and whose presentation at the University of Utah in 2008
convinced to me apply to the bioengineering program at USC.
Also integral to my research were the facilities used to fabricate and image
devices. My thanks to Dr. Donghai Zhu for maintaining the Keck Photonic clean room
despite its neverending need for repairs and attention, and for offering me valuable
fabrication advice. I am grateful to Dr. J Provine for using his expertise in ALD to coat
my mechanical test samples with Al
2
O
3.
In addition, my thanks to Dr. John Curulli at the
Center for Electron Microscopy and Microanalysis who helped train me on the SEM and
iii

maintained the center. Also, my thanks to Dr. Tuan Hoang for managing the lab, offering
suggestions, and handling many of the financial aspects of the lab.
Research could not have been performed without a superb office staff. I’d like to
thank Sandra Johns, Daisy Rusli, and Karen Johnson for tirelessly coordinating countless
orders for laboratory supplies and handling reimbursements and stipends. Mischal
Diasanta was my first official contact with the USC BME department and has expertly
guided me through the administrative and paper work side of the Ph.D. process.
I would like to give a special thanks to Alycen Hall and Dr. Krishna Nayak who
financially supported me during my penultimate year through the Body Engineering Los
Angeles (BELA) program. Their excitement and enthusiasm for science and science
teaching helped me to be a better communicator and gave me a chance to rekindle my
excitement for the scientific method. BELA also allowed me to collaborate with Ben
Patapoff and his 2013-2014 6
th
grade science and math classes. Teaching with Ben at the
Foshay Learning Center was a highlight of my Ph.D. career as it gave me a chance to be
warmly greeted each time I visited class and be inspired by the students’ genuine interest
in my research and other topics that I presented. Additionally, much thanks to Craig
Gross, also from the Foshay Learning Center, for helping me film the mask making
process and for letting me borrow lab supplies for my demonstrations.
I’m grateful for being able to work with the other members of the Biomedical
Microsystems Laboratory (BML). This laboratory has been an ideal environment for
research. In addition to the technical expertise and assistance that were always available,
the laboratory has also been a kind of family, supporting members in their challenges and
celebrating with them in their successes. Dr. Christian Gutierrez oriented me in lab,
guided me during my first project, and was an invaluable resource for engineering
information in general and MEMS fabrication information in particular. Dr. Heidi Tu
welcomed me into lab, organized hikes and other events, and was an example of hard
work and conscientiousness. Dr. Jonathan Kuo taught me microfabrication, was a
resource for the history of the Biomedical Microsystems Laboratory and the comic book
iv

universe, and had a calming attitude. Dr. Seth Hara joined the lab at the same time as I
did. We took classes together, passed our qualifying exams at similar times, and defended
our dissertations within the same month. We even had weddings nearly within a year of
each other. The Ph.D. process was a much more enjoyable with someone as nice and
intelligent as Seth to undertake it with. Dr. Brian Kim shared desk space next to mine and,
in addition to keeping me up to date with the latest viral Youtube videos, was a valuable
sounding board for new ideas, shared his great eye for designing figures and graphs,
helped me with Labview programming, and was ever willing to show me how to solder
tiny electronics. Dr. Roya Sheybani cultivated and preserved lab culture through planning
events and remembering birthdays, she brought me newspapers at lunch, and, along with
Seth, was a resident expert of electrochemistry and wireless electronics. I would also like
to thank Lawrence Yu for maintaining the computers, helping me set up the
spectrophotometer, and being the lab photographer, and Angelica Cobo Wood for
organizing lab racquetball and always being in a good mood.
Additionally, the newer members of the lab have also already had a positive
influence on the group dynamic. Many thanks to Alex Baldwin who for explaining
electrical engineering concepts and helping to program my Arduino, Ahuva Weltman for
helping me complete my drug release study, Jessica Ortigoza Diaz for being willing to
carry on some of my work after me, and to Dr. Kee Sholten who has been an ideal post-
doctoral scholar for our lab with his encyclopedic knowledge of physics and
microfabrication.
I also would like to thank Eric Welder and Louis Jug for their help in
characterizing strain sensors for my first paper, as well as all the other undergraduate and
summer students that have helped around the lab through the years.
Lastly, I would like to thank my family for their unwavering support and
encouragement. My uncle Douglas Lee welcomed me to California, provided home
cooked meals, and let me stay at his home until I found a place closer to USC. Two of my
siblings, Dr. William Scott Lee and Dr. Alison Lee Eldredge, were in a Ph.D. program
v

concurrently with me and empathized with me during my challenges, offered practical
advice, and congratulated me on my accomplishments. My sister Nicole Shelley was
always quick to offer encouragement, a listening ear, and a fresh perspective from outside
my field. My parents have been models of patience, encouragement, and optimism from
start to finish.
Finally, I thank my wife Jamie to whom this dissertation is dedicated. I thank her
for her encouragement and support. She was not only the first person to hear my
complaints, but also the first person to celebrate my successes. She read over drafts of
papers to correct grammar and organization, listened to endless discussions about my
research, took me out to dinner after reaching milestones in my program, ate late dinners
with me, brought food to my defense, and participated in lab events. She is my love and
my best friend.

vi

TABLE OF CONTENTS
ACKNOWLEDGEMENTS ....................................................................................................... ii
TABLE OF CONTENTS ........................................................................................................ vi
LIST OF TABLES ............................................................................................................... viii
LIST OF FIGURES ............................................................................................................... ix
ABSTRACT............... ........................................................................................................ xvii
CHAPTER 1 INTERFACES BETWEEN FLEXIBLE IMPLANTABLE ELECTRONICS AND
BIOLOGICAL TISSUES ................................................................................. 1
1.1 IMPLANTABLE MEDICAL DEVICES ................................................................................ 1
1.2 BIOMATERIALS ........................................................................................................... 2
1.3 BIOCOMPATIBILITY OF ELECTRONIC MEDICAL DEVICES ............................................ 3
1.4 FLEXIBLE ELECTRONICS ............................................................................................. 4
1.5 COATINGS FOR IMPROVED BIOCOMPATIBILITY OF INTERFACES .................................. 6
1.6 METHODS FOR EVALUATING THE MECHANICAL PERFORMANCE OF IMPLANTABLE
THIN FILM DEVICES ................................................................................................... 8
1.7 STRATEGIES FOR IMPROVING MECHANICAL AND BIOCHEMICAL INTERFACES
BETWEEN MEDICAL IMPLANTS AND TISSUES ........................................................... 11
CHAPTER 2 FLEXIBLE, HIGH–STRAIN SENSORS MADE FROM BIOCOMPATIBLE
MATERIALS ............................................................................................... 26
2.1 BACKGROUND ........................................................................................................... 26
2.2 CPDMS STRAIN SENSOR FABRICATION ................................................................... 32
2.3 CHARACTERIZATION OF CPDMS STRAIN SENSORS .................................................. 35
2.4 CONCLUSION ............................................................................................................. 40
CHAPTER 3 MATRIGEL COATINGS FOR PARYLENE SHEATH NEURAL PROBES ......... 47
3.1 MATRIGEL AND PENETRATING NEURAL ELECTRODES .............................................. 47
3.2 PREPARATION OF MG-BASED COATINGS ................................................................... 52
3.3 COATING OF PARYLENE SHEATH ELECTRODE ........................................................... 54
3.4 CHARACTERIZATION OF MG COATINGS .................................................................... 56
3.5 MEASURING DRUG LOADING AND RELEASE FROM PROBE IN VITRO ............................ 56
3.6 IN VIVO ELECTROPHYSIOLOGICAL EVALUATION FROM 1-MONTH STUDY .................... 58
3.7 CHARACTERIZING IMMUNOHISTOCHEMICAL EFFECT OF DRUG RELEASE FROM PROBE
IN VIVO FROM 1-MONTH STUDY ............................................................................... 60
3.8 IN VIVO EIS AND ELECTROPHYSIOLOGICAL EVALUATION FROM 3-MONTH STUDY ..... 60
3.9 MG ON PARYLENE AND ELECTRODE SURFACES ........................................................ 61
3.10 DRUG RELEASE FROM COATING ............................................................................... 65
3.11 ELECTROPHYSIOLOGICAL MEASUREMENTS FROM 1-MONTH STUDY ........................ 68
vii

3.12 IMMUNOHISTOCHEMICAL EFFECT OF DRUG RELEASE FROM PROBE IN VIVO FROM 1-
MONTH STUDY ........................................................................................................ 69
3.13 ELECTROPHYSIOLOGICAL MEASUREMENTS FROM 3-MONTH STUDY ........................ 71
3.14 CONCLUSION ........................................................................................................... 73
CHAPTER 4 MECHANICAL PROPERTIES OF PARYLENE-METAL-PARYLENE THIN
FILM DEVICES ........................................................................................... 80
4.1 THIN FILM POLYMERIC DEVICES ................................................................................ 80
4.2 AL
2
O
3
AS A BARRIER MATERIAL ................................................................................ 84
4.3 PEEL STRENGTH ........................................................................................................ 90
4.4 MINIMUM BENDING DIAMETER .................................................................................. 97
4.5 BENDING FATIGUE .................................................................................................. 104
4.6 DISCUSSION ............................................................................................................ 116
4.7 CONCLUSIONS ......................................................................................................... 120
CHAPTER 5 CONCLUSION ........................................................................................... 127
APPENDIX A: DEVELOPMENT OF COATING METHOD FOR PARYLENE-SHEATH
ELECTRODE ............................................................................................ 130
APPENDIX B: RECIPE FOR MECHANICAL TEST DEVICES ............................................... 153
APPENDIX C: MASKS FOR MECHANICAL TESTING DEVICES ......................................... 155

viii

LIST OF TABLES
Table 1-1. Material properties of traditional electronics materials compared to
traditional implanted biomaterials and biological tissues ..................................4
Table 2-1. MWNT Properties [18] ....................................................................................28
Table 2-2. Graphene Nanoplatelets Properties [28] ..........................................................29
Table 3-1. Composition of Coatings Used in Study .........................................................53
Table 4-1. Material combinations investigated for flexible Parylene devices. .................85

ix

LIST OF FIGURES
Figure 2-1. Photograph of screen-printed MWNT/GNP strain sensor. (Scale bar is 5
mm.) Reprinted with permission from [5]. Copyright 2013, AIP
Publishing LLC. ..........................................................................................26
Figure 2-2. 2D chemical drawing of PMDS. .................................................................27
Figure 2-3. MWNTs (scale bar is 500nm) shown in their pre-processed state.
Reprinted with permission from [5]. Copyright 2013, AIP Publishing
LLC. ............................................................................................................28
Figure 2-4. GNPs (scale bar is 1µm) in their preprocessed state. Reprinted with
permission from [5]. Copyright 2013, AIP Publishing LLC. ......................29
Figure 2-5. Cross sectional SEM image of strain sensor thick film prepared by freeze
fracture. (Scale bar is 1 µm.) Reprinted with permission from [5].
Copyright 2013, AIP Publishing LLC. ........................................................30
Figure 2-6. Schematic of percolation mechanisms in CNT/Polymer composites.
Composites begin non-conductive at low concentrations (a). At slightly
higher concentrations, some tubes will be close enough for tunneling
effect to allow electrons to pass between networks of tubes (b). At a
certain threshold, contiguous networks are formed (c) (networks are
shown in red), the conductivity will then increase following a percolation
power law (eq. 1) as more networks are formed (d) (additional network
shown in blue). (Adapted from [31].) ..........................................................31
Figure 2-7. Expected microstructure change of a composite under straining; (a)
percolation before straining (electrically conductive) and (b) non-
percolating under strain (non-conductive). Reprinted from [35],
Copyright (1998), with permission from Elsevier. ......................................32
Figure 2-8. (a) Mask design used in screen printing along with (b) brass screen, and
(c) process diagram of screen printing. .......................................................34
Figure 2-9. Conductivity data comparing experimental results against those for
similar composites found in literature (Lu et al.[12] used 40-60 nm OD
MWNTs and Liu and Choi [13] used 20-40 nm OD MWNTs). The fitted
line follows Eq. 1 and is
 
02 . 3
499 . 0 95    
MWNT
. Reprinted with
permission from [5]. Copyright 2013, AIP Publishing LLC. ......................36
Figure 2-10. Zero current resistance (mean ± SE, n=6-12). Reprinted with permission
from [5]. Copyright 2013, AIP Publishing LLC. ........................................37
x

Figure 2-11. TCR (mean ± SE, n=3-5). Reprinted with permission from [5]. Copyright
2013, AIP Publishing LLC. .........................................................................38
Figure 2-12. (a) Strain pattern applied to device. (b) Raw data of resistance versus time.
(c) Resistance versus strain (with first 2 cycles removed). Gauge factor
obtained by averaging slope of linear region. (d) Gauge factors for
various compositions including MWNT and/or GNP fillers (mean ± SE,
n=4-6). Reprinted with permission from [5]. Copyright 2013, AIP
Publishing LLC. ..........................................................................................40
Figure 3-1. (a) Micrograph (scalebar = 100 µm) and (b) three dimensional illustration
of PSE. Image from [6]. © 2015 Wiley Periodicals, Inc. ............................48
Figure 3-2. Simplified schematic of fabrication steps. (Note: fabrication steps are
drawn on a single cut-off probe, however probes are actually etched out
process step (c).) (a) Parylene (5 µm) was deposited onto a Si carrier
wafer. Pt electrodes (200 nm) were defined using e-beam evaporation and
lift-off. Parylene insulation (1 μm) was deposited and the electrode sites
exposed using O
2
plasma etching. (b) A sacrificial photoresist structure
was used to form the microchannel structure. (c) An overlying Parylene
(5 μm) layer completed the enclosed microchannel. The ends of the
microchannel were etched open and, in the same etch step, the individual
probes were cut out. Perforations (optional) were sequentially etched at
each previous etch step and completely etched through during this final
etch step. (d) After carefully removing each probe, a microwire mold was
inserted into the sheath to form the 3D sheath structure. (e) The final
mechanically stable sheath shape defined by the microwire was retained
after a thermoforming process. (Figure modified from [7]. © IOP
Publishing. Reproduced by permission of IOP Publishing. All rights
reserved.) .....................................................................................................50
Figure 3-3. Schematic of the sonication method. (1) Probe is placed in coating
mixture which is placed in sonication bath. (2) Coating is gelled and
dehydrated at room temperature. Image from [6]. © 2015 Wiley
Periodicals, Inc. ...........................................................................................55
Figure 3-4. Schematic of the surface modification method. (1) Probe is soaked in
PDL solution for 1 h. (2) Probe placed inside of 10 µL droplet for 5
minutes covered in a polystyrene container. (3) Probe is taken out of
container and is dehydrated for 5 minutes. ..................................................56
Figure 3-5. Optical images of fractal ‘‘trees’’ made from dried MG on a Parylene
coated silicon wafer under (a) unpolarized and (b) polarized light. (Scale
bars are 50 µm.) ...........................................................................................61
xi

Figure 3-6. SEM image of the top portion of a perforated PSE coated using surface
modification. The MG pooled towards the center of the cable near the
large opening of the sheath and filled in the perforations which were
added to the PSE to improve cellular communication across the Parylene
substrate. Image from [6]. © 2015 Wiley Periodicals, Inc. .........................62
Figure 3-7. (a-c) SEM images of recessed platinum electrode on Parylene probe (a)
before coating, (b) after PDL surface treatment, and (c) after coating with
MG. (Scale bars are 10 µm. PSE was sputtered with Au prior to
imaging.) Image from [6]. © 2015 Wiley Periodicals, Inc. ........................63
Figure 3-8. EIS data (mean ± SE) from (a-c) MG coated (n=7) and (d-f) MG and
MG/DEX coated (n=3) neural electrodes (using sonication method). (a &
b; d & e) Magnitude and phase data were measured pre and post coating.
(c & f) Impedance measured at 1 kHz from before the probe was coated
to 3 days after coating. Image from [6]. © 2015 Wiley Periodicals, Inc. ...64
Figure 3-9. Kinetic scan of microwells with MG/DEX and MG coated on microwell
sidewalls. (a) Schematic of test setup showing MG/DEX coating
diffusing from bottom left corner of microwell. (b) Kinetic scan showing
peak absorbance at 0.5 hour (mean ± SD, n=3-5) (c) Calibration curve of
known dilutions of MG/DEX coating. Image from [6]. © 2015 Wiley
Periodicals, Inc. ...........................................................................................65
Figure 3-10. DEX released from MG deposited on Parylene coupons using the PDL
surface modification method before and after sterilization (n = 3, mean
±SE). Image from [6]. © 2015 Wiley Periodicals, Inc. ...............................66
Figure 3-11. Cumulative release of (a) NGF and (b) BDNF from MG loaded with
neurotrophins and deposited on Parylene using the PDL surface
modification method before and after EtO sterilization (n=3, mean ±SE).
Image from [6]. © 2015 Wiley Periodicals, Inc. .........................................67
Figure 3-12. Drug loading and release from PSE coated using surface modification
method. (a) Schematic of testing method: 1) probe coated in DEX loaded
MG, 2) coated probe inserted into microwell, 3) probe soaked for time t
n
,
4) probe removed from microwell and absorbance of eluate measured
using spectrophotometer at a wavelength of 242 nm. Probe is reinserted
into well after measurement for subsequent measurements. (b)
Absorbance of eluate from coated with DEX/MG. (Absorbance of probe
coated with MG and PBS filled well is also provided to show ability of
spectrophotometer to selectively measure DEX concentrations using a
242 nm light source. See Figure 3-9(c) for calibration curve.) Image from
[6]. © 2015 Wiley Periodicals, Inc. .............................................................68
xii

Figure 3-13. (a) 1kHz impedance, (b) SNR, (c) noise, and (d) event rate measurements
and results obtained in comparing different coatings at 28 days (mean ±
SD, n = 121 recording sites in 9 animals, 16 probes). The probes had the
following coatings: MG/DEX (n = 2), MG (n = 4), and MG/NT (n = 4).
(Figure reproduced from [7]. © IOP Publishing. Reproduced by
permission of IOP Publishing. All rights reserved.) ...................................69
Figure 3-14. (a) Sample microphotographs of the GFAP immuno-stained sections of
cerebral cortex through the probe tips coated with MG (left) and
MG/DEX (right); (b) quantification of GFAP immunoreactivity in the
sections shown in (a) at increasing distances from the probe. The
immunoreactivity data are normalized to the level at the distance of 270-
300 µm. Image from [6]. © 2015 Wiley Periodicals, Inc. ..........................70
Figure 3-15. In vivo EIS and electrophysiological data from 3 month chronic study. (a)
In vivo EIS data taken during the first two weeks comparing three probes:
one non-coated, one coated with MG, and one coated with MG
supplemented with DEX, NGF, and BDNF (denoted MG +) at 1 week
and 2 weeks after implantation (n = 13-35, mean ± SD). (b) Comparison
of SNR for neural activities recorded with three probes (n=6-12, mean ±
SD). The p values are for ANOVA comparisons among three coating
groups at each time point. Error bars indicate the standard deviation for
sites within the probe. Asterisk indicates significant difference of non-
coated group vs. two other groups in the post-hoc test. Image from [6]. ©
2015 Wiley Periodicals, Inc. .......................................................................72
Figure 4-1. (a) A cable being bent between two plates with a force “F” creates
regions of (b) tensile and compressive stresses, which are defined by the
bending stress (σ
b
). The bending stress is linearly proportional to the
distance from the neutral axis (b), which is a plane in the middle of the
cable where the compressive and tensile forces cancel out. ........................82
Figure 4-2. Schematic of processing steps for (a) PMP devices and (b) PAMAP
devices. (1) Parylene C is deposited over a silicon wafer. (2) A standard
liftoff procedure is used to pattern e-beam deposited platinum. (3) A
second layer of Parylene C is deposited over the patterned platinum. (4)
Contact pads are exposed and devices are cut out using O
2
plasma etch.
(5) Devices are released from silicon wafer. For PAMAP devices a
conformal barrier layer of Al
2
O
3
was deposited before step ii and before
step iii. For PMAP devices, Al
2
O
3
was only deposited before step iii. .......86
Figure 4-3. (a) Pattern of intended alignment marks. (b) Alignment marks resulting
after depositing metal on Al
2
O
3
for a second time (scale bar is 150 µm).
(c) Overlap between two alignment marks produce image in (b). ..............87
xiii

Figure 4-4. Etch rate of Al
2
O
3
in dilute NaOH developer (~0.1 M) was determined by
breaking a wafer coated in Al
2
O
3
into (a) separate wafer shards. The
wafer shards were then (b) placed in the dilute NaOH developer and (c)
removed at different soak times. The thickness of Al
2
O
3
on the shards
removed from the developer were then (d) measured using ellipsometry. .88
Figure 4-5. Thickness of Al
2
O
3
after different etch times in 1:4, DI H
2
O:AZ 340
developer as determined by ellipsometry using 3 different wavelengths.
(5 nm is approaching resolution of ellipsometer.) Etch rate is determined
to be 3.3 to 3.8 Å/s. .....................................................................................89
Figure 4-6. Simplified schematic of peel test sample fabrication. (1) Sacrificial
photoresist (PR) layer and bonded interface are patterned next to each
other on Parylene substrate. (2) Sacrificial PR and interface are coated
with a second layer of Parylene. (3) O
2
plasma is used to etch through
both layers of Parylene on either side of device. (4) Device is released
and soaked in acetone to dissolve PR. (5) Flaps are peeled back and
attached to a load cell to perform peel test. (Note: For samples with Al
2
O
3
,
Al
2
O
3
was deposited either before and after the sacrificial layer patterned
in (1), or only after the sacrificial layer was patterned in (1).) ....................92
Figure 4-7. (a) Flaps of peel test sample separated with tweezers after removing
sacrificial PR. Bonded metal region (3 × 4.5 mm) is circumscribed by
dotted line. (b) Peel test sample placed over mounting posts on motorized
stage with load cell before top clamp is secured. White arrow in (b)
points to bonded region of peel test device. (Scale bars for (a) and (b) are
5 mm.) .........................................................................................................93
Figure 4-8. Representative peel test raw data from (a) PP and (b) PMP test devices.
The edge effect observed for the PP device was not observed for the PMP
device, most likely due to the force magnitude being dominated by the
larger PMP average peel strength, (~50 times larger). Note the two plots
are on different scales. .................................................................................94
Figure 4-9. Peel test data for all conditions. (a) Combined raw data (mean ± SE, n =
3-8) from various trials, and (c) mean values of average peel strength
from various interfaces (mean ± SE, n = 3-8). Parylene/Parylene
interfaces had the lowest peel strength and were regraphed separately (in
b & d) on a different scale. The slash symbol (used in c & d) indicates the
interface at which delamination occurred. ..................................................95
Figure 4-10. SEM images of all peel test surfaces. For each individual panel, the left
image is the bottom (closest to the silicon wafer during processing) layers
of Parylene, metal, and Al
2
O
3
and the right image is the top layer of
Parylene. Note that images are at different scales. (Surfaces were
sputtered coated with Pt before imaging.) ...................................................96
xiv

Figure 4-11. (a) Cartoon and (b) photograph of minimum bend diameter test devices
(scale bar is 5 mm). In cartoon (a), the thickness of the traces have been
increased and number of turns decreased for illustrative purposes. The
metal and hole on the left side of device are for clamping the device to
the test fixture. The contact pads on the right side allow electrical
resistance measurements. ............................................................................98
Figure 4-12. Pictures of crush test setup. (a) Overview of setup with acrylic block
connected to the load cell attached to a motorized stage. (The crush test
device is obscured behind acrylic block.) (b) Close-up of folded end of
crush test device between two acrylic blocks with one acrylic block
attached to a load cell (scale bar is 5 mm). (c) Close-up of crush test
device folded around holder (scale bar is 5mm). ........................................99
Figure 4-13. Schematic of bending failure test setup with (a) contacts pads out and (b)
contact pads in. Setup is the same except that in (b) the device is turned
inside out and the ZIF connector is turned upside-down to accommodate
downward facing contact pads. .................................................................100
Figure 4-14. Raw force (a, c, and e) and resistance (b, d, and f) data from minimum
bending diameter tests. Force data (a, c, and e) were used to determine
where the acrylic blocks made full contact (i.e. where Force = 4.5 N).
Minimum bending diameter was defined as the resistance measurement
(in b, d, and f) closest to zero not measured as an open circuit. ................101
Figure 4-15. Summary of crush test data. PMAP (in) had a significantly larger
minimum diameter than either PMP (in) or PMP (out). (Mean ± SE, n =
3-4) ............................................................................................................102
Figure 4-16. SEM of cracks formed after bending (a & b) PMP devices and (c & d)
PMAP devices away from the contact pads (‘out’). The top layers of
Parylene (a & c) were the outside of the bent device, and the bottom layer
(b & d) were from the inside of the bent device. (Images on the right side
of each panel are zoomed in views of images on the left with scale bars of
the left and right images being 50 and 10 µm, respectively. Parylene
sputtered coated with Pt before imaging.) .................................................103
Figure 4-17. SEM of cracks formed afer bending (a & b) PMP devices and (c & d)
PMAP devices towards the contact pads (‘in’). The bottom layers of
Parylene (a & c) were the outside of the bent device, and the top layer (b
& d) were from the inside of the bent device. (Images on the right side of
each panel are zoomed in views of images on the left with scale bars of
the left and right images being 50 and 10 µm, respectively. Parylene
sputtered coated with Pt before imaging.) In the PMAP sample,
delamination at the bend was observed. ....................................................104
xv

Figure 4-18. (a) Cartoon and (b) photograph of bend test devices (scale bar is 5 mm).
(In cartoon (a), the thickness of the traces have been increased and
number of traces decreased for clarity.) The metal and hole on the left
side of device are for gripping the device. The traces are exposed on the
right side to make contact pads to connect the device to a potentiostat. ...106
Figure 4-19. (a) Schematic and (b) picture of bend testing setup. Design was based on
the cable flex test as described by the ASTM standard D4565-10 for
evaluating insulation on wires. ..................................................................107
Figure 4-20. Electrochemical impedance spectroscopy (EIS) was performed on bend
testing devices between (a) adjacent traces (lateral impedance) and (b)
from traces on the device to a platinum counter electrode off the device
(transverse impedance). .............................................................................108
Figure 4-21. Variations of a simplified Randle’s circuit modeling (a) a fully insulated
electrode, (b) a deinsulated electrode, and (c) a failing electrode, where
C
dl
= double layer capacitance, R
s
= solution resistance, R
ct
= charge
transfer resistance, C
c
= coating capacitance, and R
p
= pore resistance.
Note that the model is the same for the insulated and deinsulated model,
but the capacitive and resistive elements will be different. Figure adapted
from [56]. ...................................................................................................109
Figure 4-22. Representative transverse (a-c) and lateral (d-f) impedance from a PMP
device released from a wafer that peel testing revealed had uneven
adhesion of the top layer of Parylene. The device shows clear change in
impedance over 5k bend cycles represented by downward shifts in
magnitude (a & d), which are emphasized by normalizing the data to the
impedance values at 0 bends (b & e), and shifts in phase (c & f) at high
and low frequencies (mean ± SE, n = 8). ..................................................110
Figure 4-23. Transverse impedance from (a & b) PMP, (c & d) PMAP and (e & f)
PAMAP Parylene cables (mean ± SE, n = 16). Magnitude is normalized
with respect to the impedance at zero cycles. (PMAP and PAMAP
devices were only bent for 10k cycles.) ....................................................111
Figure 4-24. Lateral impedance from (a & b) PMP, (c & d) PMAP, and (e & f)
PAMAP Parylene cables (mean ± SE, n = 16). Magnitude of impedance
is normalized to the impedance of the device at zero cycles. (PMAP and
PAMAP were only bent for 10k cycles.) ..................................................113
Figure 4-25. Nomarski differential interference contrast images of fatigue cracks in the
top Parylene layer covering three traces of a PAMAP devices after (a) 2k
bend cycles and (b) 5k bend cycles. Metal regions are wide, light gray
strips and Parylene appears as thin, dark gray strips. The majority of
xvi

cracks appear after 5k bend cycles; however, the dashed circle marks
cracks present after 2k cycles (circle is 200 µm in diameter). ..................114
Figure 4-26. (a) Optical and (b) SEM images of cracked Parylene over and between
metal traces on cable bent for 100k cycles. The metal traces are on the
left. (Parylene sputtered coated with Pt before imaging in SEM.) ............114
Figure 4-27. Micrograph of two traces near the edge (marked by dotted white line) of a
PAMAP device after 10k bend cycles (scale bar is 250 µm). Evidence of
water intrusion can clearly be seen as large Newton rings in the Parylene
next to the edge and smaller Newton rings on the right edge at the bottom
of the right trace. Discoloration and pock marks on the traces and may be
evidence of Al
2
O
3
etching by water that has penetrated between Parylene
layers. ........................................................................................................115
Figure 4-28. (a) Micrograph of all 8 traces of a PMAP device showing discoloration
and branched patterns occurring on outside traces (scale bar is 500 µm).
(b) The transverse impedance at 1 kHz for the traces shown in (a) versus
their proximity to the edge. Before bend testing, traces closest to the edge
have the lowest impedance, but this difference decreases as the device
undergoes more bend cycles. .....................................................................116

xvii

Implantable medical devices hold great promise to treat and prevent chronic
conditions, but are limited by their biocompatibility. Immune responses due to
mechanical mismatches and surface properties such as topology and biochemistry, can be
dangerous to the patient, degrade the device, or prevent the device from performing its
functions correctly. The biocompatibility of electronic devices is particularly challenging
because they typically contain circuits patterned on silicon wafers, which makes them
prone to be much harder than tissue. They also require hermetic seals to protect the
circuitry from liquid and intimate contact with tissues to monitor tissue properties and
safely deliver therapeutic amounts of current. Even moderate immune responses can
degrade protective coatings and increase the impedance of electrodes. This work presents
strategies to improve the interface between implanted medical devices and biological
tissues.
The first strategy presented is decreasing the mechanical mismatch of implantable
electronics. This is demonstrated by the development of a strain sensor that could be used
to measure the extension of muscles such as the bladder as part of a prosthetic to
determine bladder fullness. The device is made from polydimethylsiloxane, a material
already used in many implantable devices because of its inert chemical structure and soft
tissue-like mechanical properties, mixed with carbon nanotubes which make the rubber
piezoresistive at relatively low weight fractions.
The next strategy is treating an already developed flexible electronic device with a
bioactive coating. In this strategy, a flexible thin film implantable intracortical electrode
is coated with biologically derived extracellular matrix proteins mixed with bioactive
molecules. The extracellular matrix proteins alters the topological, mechanical and
chemical structure of surface and acts as a substrate for cells to bind to in order to make
intimate contact with the electrodes on the device. Bioactive molecules known to reduce
ABSTRACT
xviii

the immune response (dexamethasone) and encourage cell growth and differentiation
(neurotrophins) were also added to the extra cellular matrix proteins to further improve
integration.
The last strategy is to use a combination of mechanical tests to test and screen thin
film flexible electronic architectures, which show promise as implantable medical
devices. By measuring the strength of adhesion forces between the layers of the thin films
along with their flexibility and resistance to fatigue, one can compare different materials
and manufacturing methods to quickly compare and improve the properties of new thin
film flexible electronics architectures.
Through these strategies, the interfaces between tissues and medical devices can
be improved and enable implantable devices to function better and deliver novel therapy
for chronic conditions.

1

1.1 Implantable medical devices
Medical devices are tools that use thermal, mechanical, or electrical means to
diagnose, treat, or prevent disease. For chronic conditions, medical devices are often
implanted into the body to be next to the affected tissue, or to deliver continuous
treatment underneath the protection of the skin. Even though implanted medical devices
are common with nearly 200,000 artificial hips [1] and over half million stents [2] are
implanted each year in the US, many of these devices fail. After 10 years, 12% of hips
need revision surgery [3] and 7% of stents experience restenosis after 2 years [4].
Electronic medical devices implanted in the body have unique challenges.
Electronics short if they are not hermetically sealed and often require very close contact
with tissues to be safe and effective. Nearly 200,000 pacemakers are being implanted
each year in the US [5] and have a major complication rate of 4% [6]. The use of some
experimental devices, such as intracortical electrode arrays, has been limited to
experimental trials because of unknown efficacy and safety of chronically implanted
devices [7]. Improving the interface between implantable electronic medical devices will
help reduce the complication rate of current devices, while providing a means to use
electronics to monitor and deliver novel therapy to tissues.
Medical devices fail either because the materials in the devices break down in the
harsh environment of the body or cause an immune response due to their chemical,
mechanical, or morphological properties that reduce the functionality of the device and/or
is dangerous to the patient. Therefore, medical devices are made from inert materials that
CHAPTER 1
INTERFACES BETWEEN FLEXIBLE IMPLANTABLE ELECTRONICS AND
BIOLOGICAL TISSUES
2

do not cause severe immune responses. These types of materials are often referred to as
“biomaterials.”
1.2 Biomaterials
Inside the body is a harsh saline environment of elevated temperatures and
humidity, which can passively dissolve or corrode many materials. In addition, the
immune system will respond to the injury of the implant surgery, shear forces created by
mechanical mismatches, and the proteins absorbed by the surface of the biomaterial by
creating a region of fibrous encapsulation and chronic inflammation where foreign body
giant cells release degradation products such as reactive oxygen intermediates (i.e. free
radicals) or acid to hasten the degradation of the foreign body [8]. Currently, there are
only a relative few number of materials that can withstand such harsh environments.
Among these materials, the polymers Parylene C and polydimethylsiloxane (PDMS), and
the metals titanium and platinum are some of the most commonly used.
Parylene C is a clear, flexible polymer consisting of subunits containing a
benzene ring with a single chlorine atom. Since a method for vapor depositing Parylene
was developed [9], Parylene C has been widely used as a coating for implantable
materials because of its ability to form conformal, stable, pinhole free coatings [10], and
reduce the coefficient of friction. Recently, Parylene C has been investigated for its utility
as a substrate for making micromachinable devices [11-14].
PDMS is a clear, elastic, inert polymer that consists of a backbone of alternating
silicon and oxygen atoms with methyl groups branching off of the Si groups. Because of
its low modulus and elastic properties, it is often used in reconstruction and plastic
surgery and as a coating for electronics to eliminate sharp edges and provide a weak
water barrier [15]. It is also used in trancheal and bronchial stents [16], vascular grafts
[17], and catheters.
Titanium is a low density, high strength transition metal with excellent corrosion
resistance due to a stable oxide that forms on its surface. This inertness allows titanium to
3

be used as a hermetic case for pacemakers. Titanium is also used for dental and
orthopedic implants because of its inherent ability to allow for osseointegration, the
process of bone integrating onto the surface of titanium [18].
Platinum is a noble metal with good electrical properties and excellent corrosion
resistance. It is often alloyed with iridium to increase its strength and used to make
electric contact with tissues in devices such as pacemaker leads [19], and contacts for
spinal cord stimulators [20], cochlear implants [21], and deep brain stimulation [22].
1.3 Biocompatibility of Electronic Medical Devices
Medical devices that use electricity to monitor tissues or treat symptoms of
diseases have specific biocompatibility challenges at their interfaces. In addition to
requiring a hermetic seal to avoid moisture and prevent short circuits, there is generally a
large mechanical mismatch between electronics and tissue (Table 1-1). The stiffer
material of the device can cause damage the tissues and make it difficult for the device to
maintain close contact with tissues. Close contact with the tissues is required because of
the small magnitude of electric fields being emitted by the tissues (such as neural tissue
[23]), which can also blocked by scarring caused as a result of the immune response to
the foreign object [24]. The present work focuses on improving the biocompatibility of
electronic devices by making them flexible and by modifying the surface with a bioactive
coating. The present work also presents methods for testing the mechanical properties of
flexible electronics were also developed and presented in order to compare processing
methods and materials in order to improve their performance.

4

Table 1-1. Material properties of traditional electronics materials compared to traditional
implanted biomaterials and biological tissues
Material Young’s Modulus (GPa) Elongation (percent)
Thickness of
elongation samples
Silicon 190 [25] 0.7 [26] NA - whole device
Polyimide 2.57 [27] 76 [28] 125 µm
Titanium 102.7 [29] 24 [29] NA
Platinum 164.6 [30] 40 [31] NA
PMMA 3.25 [32] 21 [33] 3.2 mm
PDMS 0.750 x 10
-3
[34] 470 [35] 2.6 mm
Parylene C 4.75 [36] 200 [10] < 1 mm
Skin 6-22 ×10
-3
[37] 65.2† [38] 1.28 mm
Bladder 1-4.2×10
-3
[39] 85 ‡[40] NA - whole organ
Neural Tissue 3.2×10
-6
[41] NA* NA*
†Rat skin,
‡Porcine organ,
*Treated as a viscoelastic solid [42],>6% elongation shows irreversible damage in peripheral nerves [43]

1.4 Flexible Electronics
Electronics are inherently rigid because they are made from adding or subtracting
material away from a relatively thick (~300 µm) rigid wafer of single crystal silicon. The
rigidity of electronics limits the biocompatibility because of the difficulties of attaching
or otherwise interfacing a device to soft, flexible tissue. As living tissues bend and flex,
the electronics may move, or break, which—in addition to possibly making the device
non-functional—can irritate or damage the moving tissue. Making electronics flexible is
one strategy to protect and maintain intimate contact with tissue. Flexibility can be
achieved by reducing the thickness of the conductive components [44], using geometries
that allow bending [45], using compliant conductive polymers [46-48], or incorporating
conductive filler material into a flexible substrate [49].
Silicon becomes flexible when it is less than 50 µm thick (and becomes
transparent when it is less than 20 µm [50]). Thin silicon wafers can be made by building
transistors on thin layers of silicon, suspended on micro-pillars and broken off during
post processing, through a process called ChipFilm processing [51], or etching away the
oxide layer after building a chip on a silicon on insulator (SOI) wafer and transferring it
5

onto a pre-strained flexible polymer that causes the silicon to become “wavy” during
relaxation [45].
Thin films on polymer substrates will also be flexible. Microcontact printing [52]
involves stamping conductive solutions onto flexible substrates. In another process large
area-reduced graphene oxide, which is similar to graphene, is deposited onto a polymer
[53]. Nano-wires (of most materials) are also flexible and can be grown in a specific
direction and transferred [54]. Thin films (< 200 nm) of metal ((e.g. Pt, Au, Ti) can be
deposited onto thin polymers and remain flexible [55, 56].
Flexible polymers used as a substrate for flexible electronics include polyimide
[57, 58], Parylene C [55, 59, 60], liquid crystal polymers [61] and PDMS methods [45,
62]. PDMS is also biocompatible and transparent [44, 63]. Some polymers, such as
polypyrrole, are already conductive and can be used as an electrical component [47, 64-
67].
Instead of depositing material on top of flexible substrates, small conductive
particles can be incorporated into the substrates themselves. Non-conductive polymers
become conductive as the concentration of filler material is high enough for the particles
to form conductive networks (the percolation threshold), which means they are either
directly touching one another or close enough for electrons to tunnel from one conductive
particle to the next [68]. Nanoparticles such as carbon nanotubes [68-70] and graphene
[71, 72] have been widely used as filler in polymers because of its low percolation
threshold [71, 73], favorable electric properties [74, 75] and ability to enhance strength
[76, 77].
In addition to the benefits of flexible electronics to making medical devices that
have better interfaces with tissue, flexible electronics also have the potential to make
electronics more robust by avoiding fracture and allow for “roll-to-roll” manufacturing,
whereby electronics can be produced continuously on rolls of flexible polymers, instead
of in small batches using silicon wafers [78].
6

1.5 Coatings for Improved Biocompatibility of Interfaces
In addition to using materials that have similar mechanical bulk properties to the
biological tissue they are in contact with, the interface between the implant and the tissue
can be improved by altering the surface properties of the implant by the application of a
coating. The goal of coating an implanted device is either to encourage integration with
the surrounding tissue or inhibit the immune response. Coatings achieve these goals by
altering the topography [79] or surface chemistry [80-85], providing a carrier for drug
delivery [86-88], or combining these strategies [89, 90].
Hip replacement surgery and dental implants need to be integrated into the bone
of the patient and use surface treatments and coatings to accomplish this. Titanium has
been shown to have the rare quality of promoting osseointegration [91], which can be
enhanced by surface roughening on a micro [79] or nanoscale [92]. Porous, hydroxyl-
apatite (a mineral found in bone) coatings and other surface treatments of the titanium
femoral component in total hip arthroplasty are used as scaffolds to encourage bone
integration to prevent wear and corrosion [93, 94].
Other devices, particularly ones with sensors or orifices that need to be kept open,
aim to prevent biofouling by presenting surfaces with high surface energies. Hydrophobic
surfaces, and hydrophobic regions on proteins, have been shown to cause non-specific
binding of bacteria and initiation of the immune response [95]. Polyethylene glycol
(PEG) prevents adhesion with long chains that sterically block cell and protein adhesion
[84, 85, 96]. Surfaces are coated with moieties with different ions [97] can also provide
an antifouling barrier.
Coatings can also be loaded with bioactive molecules to modulate the response of
the tissue to the implant. Drug eluting coatings consist of a polymer loaded with drug,
which is either released by diffusion through the polymer or erosion of the polymer [98].
The eluting drugs are generally designed to prevent immunogenesis and thrombogenesis.
The most common example of a coating that is used for drug delivery is the drug eluting
stent, which was developed two decades ago to prevent the recurrence of stenosis after
7

angioplasty and stenting [99] and have since been widely accepted clinically as an
effective means to reduce restenosis and major adverse cardiac events [100, 101]. Drug
eluting coatings have also been developed for titanium for preventing infection [86, 102],
and neural electrodes for encouraging integration and mitigating the immune response
[46, 90, 103].
1.5.1 Coatings for neural electrodes
Designing coatings for neural electrodes are particularly challenging because the
electrodes need to be close to the neurons they are recording or stimulating, as well as
maintain an electrochemical path. When the brain encounters foreign materials, however,
the immune response is similar as it is for most biological tissues, which is to encapsulate
the material with scar tissue and separate the material from the surrounding tissue by
initiating a retraction of electrically active neurons [24]. Encapsulation and isolation is
particularly problematic for neural electrodes because the current needed to stimulate
neurons increases by the distance squared to the neurons and can easily exceed the limit
non-galvanic reactions, and recording electrodes cannot detect action potentials from
neurons that are more than ~50 µm [104]. The goal of coatings for neural electrodes is to
limit the immune response of neurons while discouraging retraction of neurons from the
surface of the electrode. The strategies for coating neural electrodes are similar to coating
other implants including modulating the surface chemistry, and creating drug eluting
coatings.
The goal of modifying the surface chemistry of neural electrodes has been to
reduce the number of reactive astrocytes and activated microglia near the electrode
surface, while maintaining the density of active neurons. To this end, anti-inflammatory
molecules, such as alpha-MSH peptide, has been grafted onto the surface of an electrode,
which has led to a reduced number of astrocytes near the insertion sight of the probe [82].
Counter intuitively, a coating made from laminin which increased the acute immune
response has also been applied to a neural probe because it caused a decrease in the long
term immune response [80, 81].
8

A wide variety of drug eluting coatings have also been developed for neural
electrodes to attenuate the immune response and maintain neuronal density. Hydrogels,
such as pHEMA [105], are often used as a carrier because their high porosity and
hydrophilic nature maintain a conductive path from electrode surface to neuronal tissue
while providing a low modulus interface. To decrease the elution rate of drugs from
highly porous hydrogels, the drug is loaded within slowly dissolving, biodegradable
microparticles such as poly(propylene sulfide) [106] or poly(lactic-co-glycolic) acid
[107]. This type of coating has been used to deliver dexamethasone for suppressing
inflammation and inhibiting the immune response around cortical recording electrodes
[106, 107], as well as neurotrophins to maintain neuronal density and encourage
differentiation and outgrowth of neurons near stimulating neural electrodes. These
neurotrophins include nerve growth factor (NGF) [108] and brain derived neurotrophic
factor (BDNF) [109] for retinal prostheses, and neurotrophin-3 (NT-3) [110] for cochlear
implants. A combination of methods have also been used including DEX loaded PLGA
fibers electrospun around an electrode, and coated with an alginate hydrogel, to provide a
matrix for growing a layer of conductive polymer (poly(3,4-ethylenedioxythiophene))
near the electrode site to decrease impedance [46].
1.6 Methods for evaluating the mechanical performance of implantable
thin film devices
Thin film polymeric devices have great potential for chronically implantable
medical devices. Because they are manufactured using microfabrication, they have
micron scale resolution or better features and allow sensors such as transistors [111] or
force sensors [14] to be integrated directly onto the flexible device. Their thin, flexible
form factor allows them to be used as implantable devices. They are thin enough to be
rolled up and delivered via catheter [112] or incorporated into the catheter itself [113].
Thin film polymeric devices for cochlear stimulation [114], spinal cord stimulation [13],
as well as electrocardiography (ECG) [115, 116], electrocorticography (ECOG) [117],
and intracortical neural recordings [118] have been demonstrated.
9

However, inside the body is a harsh, dynamic environment that can cause thin
film devices to degrade or break. Aqueous environments can cause delamination, which
can short traces in the device, and forces in body can be great enough to bend and crack
the device. In addition, repeated stresses, as would be found if the device is attached to
the heart or neck muscles, could cause fatigue failure of devices. Peel testing, bending
failure testing, and fatigue testing are three tests that can be used to compare and evaluate
how well thin film device may perform in the body.

1.6.1 Peel testing
Delamination is a common concern for thin film polymeric devices, especially in
aqueous solutions. In particular, the bond between untreated Parylene and Parylene has
been shown to be particularly weak [119]. The bond between Parylene and Parylene may
be enhanced by adhesion promoters, such as silane A-174, which can bond covalently
both with Parylene free radicals during the deposition process as well as an underlying
substrate [120]. Also, a combination of heat and compression have been shown to bond
Parylene layers together by causing the polymer chains on the surface of the touching
films to become entangled [121]. Models have shown that other interfaces in Parylene
thin film devices, such as Parylene and metal also shown that metals such as platinum
and titanium have the potential to form stable, covalent bonds when deposited on
Parylene [122].
Because the in vivo survival of Parylene devices can be connected to the adhesion
of the various layers in Parylene, measuring the adhesion can help evaluate different
materials and processes. There are several ways that have been used to measure the
adhesion force of thin films. The “scotch test tape” measures the adhesion of a film
deposited on a substrate by sticking a piece of pressure sensitive tape to the film and
pulling it off [123, 124]. If the top layer comes off, it can be considered to have “failed”
the scotch test tape. The adhesion force of the film can also by creating a flap in the film
(usually with a sacrificial layer) that can be gripped and connected to a load cell to
10

measure the force required to pull the film off of the substrate [125]. One such test is
called the “T” test because of the shape that the test sample makes when the two ends are
pulled apart at 180°.
T-peel test samples consist of two flexible rectangular substrates bonded together
at one end. The T-peel test is performed by gripping each end of the non-bonded regions
and pulling them apart at 180°. In T-peel tests, peel strength is defined as the average
load per unit width of bondline required to progressively separate the two members of the
test. T-peel tests are used for their simplicity of sample design and the ability to compare
different materials and bonding methods.

1.6.2 Minimum bend diameter testing
The bend test performance of a thin film device is important both because the
device may bend while it is being used in the body or it might be part of a larger system
that requires the device to be bent. Thin film nerve cuff electrodes [126], for example,
need to bend around a nerve which may, in turn bend as muscles contract. Also, thin film
electromyocardiograph electrodes connected to directly to muscles [127], which contracts
and relaxes. Devices where thin films are attached to catheters [128], or made into tubes
[113] are limited to what size of catheter can be connected to, by what the minimum
diameter the device can be bent into before failing.
A common method for evaluating minimum bending diameters is wrapping thin
film devices around cylinders of a given diameter [126, 129-132], but these tests are
limited to discreet cylinder sizes. One method that measures a continuous drop in
diameter until failure is a two plate bend test, where a loop of the folded device is placed
between two plates and the plates are brought together until the devices fails. This
method was first used in measuring the strength of optic cables [133], but has also been
used to measure the minimum diameter of thin conductive films [134]. The two plate
11

bend test is well suited for thin film polymeric devices because it does not rely on a
specific grip of the cable and requires only a small length of cable.
1.6.3 Fatigue testing
In addition to forces that exceed the tensile force of thin films that they might
encounter in the body and strains caused by rolling the device into a small diameter that
also might fracture the device, implanted devices also may encounter fatigue by repeated
exposure to stresses below their tensile strength. The heart muscles contracts over 100k
times per day and the neck moves an average of 10 million times in each direction each
year [135, 136]. Devices implanted near or on these organs must be designed to
withstand these repeated stresses.
Fatigue is measured by applying multiple cycles of stress (e.g. uniaxial tension
[137, 138] or bending [139-141]) and measuring the effects of fatigue by either looking
for cracks after a certain number of cycles, or continuing to cycle the material until it fails
physically or electrically.
1.7 Strategies for improving mechanical and biochemical interfaces
between medical implants and tissues
This work describes the investigation of a flexible piezoresistive material with the
potential to be used as an implantable strain sensor, a coating for flexible neural
electrodes, and methods developed for evaluating components of flexible electronics. In
Chapter 2, the properties of a piezoresistive conductive silicone are investigated and a
flexible strain sensor capable of high strain was developed as a proof of concept. In
Chapter 3, a coating based on the biologically derived extracellular matrix material
Matrigel is used to improve the performance of the Parylene sheath recording electrode is
developed and characterized. Chapter 4 presents methods for testing various interfaces
found in Parylene-metal-Parylene based flexible electronics. The adhesion of the various
layers are tested with peel testing and electrochemical impedance spectroscopy, and the
minimum diameter before failure is measured by crushing Parylene-metal-Parylene
cables until they break.
12

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[134] Y. Leterrier, L. Medico, F. Demarco, J.-A. Månson, U. Betz, M. Escola, et al.,
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frequency and magnitude of neck motion in healthy individuals," Spine, vol. 33,
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25

used as implantable electrodes," Materials Science and Engineering: A, vol. 486,
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1594-1597.

26

2.1 Background
Biological systems, in contrast to traditional electronics, are flexible and undergo
large displacements. Muscles, for instance, have a theoretic maximum displacement of a
muscle is 50% of its resting length [1]. Bladders, in particular, can range from an ideal
internal volume of 0 mL [2] to a full bladder volume of 600 mL [3]. In the current work,
high performance strain sensors were achieved featuring simple, low-cost construction
involving the screen printing of the novel combination of multi-walled carbon nanotube
(MWNT) and graphene nano-platelet (GNP) nano-composites on biocompatible and
flexible polymer substrates capable of measuring strains up to 40% (Figure 2-1). This is
an improvement over a previous design composed of conductive epoxy as the
piezoresistive element that cracked at strains > 2% [4].

Figure 2-1. Photograph of screen-printed MWNT/GNP strain sensor. (Scale bar is 5
mm.) Reprinted with permission from [5]. Copyright 2013, AIP Publishing LLC.

CHAPTER 2
FLEXIBLE, HIGH–STRAIN SENSORS MADE FROM BIOCOMPATIBLE
MATERIALS
27

2.1.1 Polydimethylsiloxane
PDMS is the most common type of silicone, which is a general category of
synthetic polymers with a backbone of repeating silicon to oxygen bonds (Figure 2-2).
The silicon to oxygen bonds are more stable than carbon-carbon bonds, which give the
polymer long term stability in high and low temperatures and in aqueous solutions,
whereas the methyl groups on the sides of the backbone create a non-reactive surface.
PDMS also has many properties that are more similar to tissues and organs than
traditional materials used for electronics (Table 1-1). Because of these unique properties,
PDMS has come to be used in many medical applications such as catheters, protective
coatings for electronic devices such as pacemakers, wound dressing materials, contact
lenses, and synthetic soft tissue used in plastic surgery [6, 7].

Figure 2-2. 2D chemical drawing of PMDS.

2.1.2 Carbon nanotubes
Carbon nanotubes (CNTs) (Figure 2-3) possess favorable mechanical, thermal and
electrical properties [8], including intrinsic piezoresistivty, and are widely used as a
conductive filler material to in PDMS-based electronics such as a transistors [9] and
strain sensors [10-13] (
Table 2-1). Vacuum filtration [10] and chemical assisted vapor deposition [11]
can create thin films of CNTs for transfer onto PDMS. CNTs were also incorporated into
PDMS using sonication [12], shear mixing [14], magnetic stirring [15], or by hand [16] to
create a conductive PDMS (CPDMS) composite patterned using molds [12], films [14],
28

micromolds [16], or microcontact printing [13]. CNTs are preferred over traditional
conductive filler materials such as carbon black and Ag powder because of their low
percolation threshold in PDMS (0.2 wt% to 5 wt% for carbon nanotubes [9, 12] versus 10
wt% for carbon black and 83 wt% for 1-2µm Ag particles [17]).

Table 2-1. MWNT Properties [18]
Outer Diameter 20-30 nm
Inside Diameter 5-10 nm
Length 10-30 µm
Electrical Conductivity > 100 S/cm
Bulk Density 0.28 g/cm
2

Figure 2-3. MWNTs (scale bar is 500nm) shown in their pre-processed state. Reprinted
with permission from [5]. Copyright 2013, AIP Publishing LLC.

2.1.3 Graphene
Graphene (Table 2-2) is an alternate low cost [19] filler material that has been
used with PDMS to form flexible electronics. Graphene is available commercially in the
form of platelets which are few layers of carbon atoms thick and diameters on the order
of microns (Table 2-2). Graphene shares similar electronic properties as carbon
nanotubes [20, 21], but is more amenable to patterning and bulk manufacturing through
29

solution based exfoliation [22], epitaxial growth on metals [23], or epitaxial
decomposition of SiC [21]. Graphene-based flexible strain sensors are realized using a
number of techniques to create thin piezoresistive layers. One approach is to deposit
graphene on thin layer [24] or 3D matrix [25] of metal (either copper or nickel),
encapsulate the layer in PDMS, and then etch away the metal. A layer of graphene
platelets was achieved by filtration followed by transfer of the resulting layer onto a
polymer substrate [26] . Similarly, a dispersion of graphene oxide platelets was combined
with a polymer precursor and mixed vigorously for 30 min [27].

Table 2-2. Graphene Nanoplatelets Properties [28]
Thickness (avg.) 8 nm
Diameter <2 µm
Surface area 600-750 m
2
/g

Figure 2-4. GNPs (scale bar is 1µm) in their preprocessed state. Reprinted with
permission from [5]. Copyright 2013, AIP Publishing LLC.

2.1.4 Conductive polydimethylsiloxane
Carbon nanotubes and/or graphene can be combined with PDMS to make a
conductive polymer (CPDMS) (Figure 2-5). Conductivity in CPDMS is achieved when
the concentration of filler is sufficient to form contiguous networks within the non-
conductive matrix either by direct contact of filler particles (network formation) (Figure
30

2-6(c)) and/or by electron tunneling through the polymer between closely spaced
particles (Figure 2-6(b)). This percolation threshold concentration is a function of the
uniformity of distribution (which is in turn related to the viscosity of the polymer and the
preponderance of aggregation of the filler material) [29], aspect ratio, and degree of
curvature of the filler material. SWNT are curved [30], however, MWNT are relatively
straight and therefore, the degree of curvature can be ignored.

Figure 2-5. Cross sectional SEM image of strain sensor thick film prepared by freeze
fracture. (Scale bar is 1 µm.) Reprinted with permission from [5]. Copyright 2013, AIP
Publishing LLC.

31

Figure 2-6. Schematic of percolation mechanisms in CNT/Polymer composites.
Composites begin non-conductive at low concentrations (a). At slightly higher
concentrations, some tubes will be close enough for tunneling effect to allow electrons to
pass between networks of tubes (b). At a certain threshold, contiguous networks are
formed (c) (networks are shown in red), the conductivity will then increase following a
percolation power law (eq. 1) as more networks are formed (d) (additional network
shown in blue). (Adapted from [31].)

The piezoresistivity of CPDMS, useful for strain and other sensing applications,
arises primarily from (1) strain-induced break up of conductive networks (Figure 2-7), (2)
the intrinsic piezoresistivity of the filler, and (3) the change in resistance due to electron
tunneling, which changes dramatically as a function of filler separation. Simulations
indicate that applied strain does not sufficiently disrupt conductive nanonetworks to
significantly alter bulk resistance of material [32], suggesting that the other two
mechanisms are responsible. Piezoresistivity contributed by CNTs in conductive
composites is dependent on the diameter of tube used. For thicker MWNT tubes having a
diameter ~65 nm, there is poor load transfer and piezoresistivity is dominated by inter
tube distance changes during strain [33]. However, for thinner tubes with diameters
around 10 nm, the intrinsic piezoresistivity of the tubes become significant [30]. The
32

nanotubes used in the current study were 20-30 nm in diameter and therefore the intrinsic
piezoresistivity of the MWNTs are expected to contribute, but not dominate the overall
piezoresistivity of the device. The contribution of electron tunneling is dependent on the
distribution and geometry of the conductive networks formed by the filler material, which
is in large part determined by the matrix material and processing parameters [34]. The
electron tunneling effects are expected to dominate at low filler concentrations and under
high strain when low density of conductive networks is present.

Figure 2-7. Expected microstructure change of a composite under straining; (a)
percolation before straining (electrically conductive) and (b) non-percolating under strain
(non-conductive). Reprinted from [35], Copyright (1998), with permission from Elsevier.

2.2 CPDMS Strain Sensor Fabrication
A strain sensor was made by sandwiching a screen printed pattern of CPDMS
between two sheets of non-conductive PDMS (Figure 2-1) as described below.
2.2.1 Manufacturing conductive polydimethylsiloxane
CPDMS was created by measuring with a precision scale a predetermined amount
of carbon filler material, either graphene nano-platelets (GNP) and/or multiwalled carbon
nanotubes (MWNT) (cheaptubes.com, Brattleboro, VT), and PDMS (MDX4-4210,
Factor II Inc., Lakeside, AZ ) base into a glass beaker to create a CPDMS prepolymer. To
this was added 10-15mL of a solvent mixture consisting of a 30:70 ratio of IPA and
Stoddard solvent (by volume). The mixture was then placed in an ultrasonic bath
33

(Bransonic 3510, Branson Ultrasonic Corp., Danbury, CT) for 15-18 hours to achieve
thorough mixing of the filler material with the PDMS base and evaporate the solvent.
Using this method, CPDMS base with a range of concentrations of MWNT and GNP
were prepared. Crosslinker (in a ratio of 1:10 to the base) was added to the
nanocomposite conductive filler-PDMS base mixture using a planetary mixer (Thinky
Corp., Laguna Hills, CA).
2.2.2 Manufacturing CPDMS based strain sensor
A strain sensor was made from the conductive prepolymer by screen printing the
CPDMS prepolymer onto a 500 µm thick layer of non-conductive medical grade PDMS
using a lithographically defined brass screen prepared using a previously developed
method [36]. Briefly, CPDMS prepolymer was spread across a custom stencil (Figure
2-8(a & b)) using a plastic squeegee (Figure 2-8(c)). The stencil was made by etching
away a lithographically defined pattern in a 76 µm thick brass shim (4.5” × 6”)
sandwiched between two sheets of a negative, dry-film photoresist. The photoresist was
exposed using a high resolution transparency mask (Mikacolor, Los Angeles, CA)
(Figure 2-8(a)) and UV light source (45 mJ/cm
2
) and developed in a dilute sodium
hydroxide bath. The exposed brass regions were etched away using ferric chloride. The
parts for this etching kit were purchased from MicroMark (Berkeley Heights, NJ).

34

Figure 2-8. (a) Mask design used in screen printing along with (b) brass screen, and (c)
process diagram of screen printing.

The screen-printed CPDMS was then placed under a vacuum for 3-4 hours to
remove any remaining solvent and then cured at 80 °C. A second layer of non-conductive
medical grade PDMS was added to fully encapsulate the CPDMS. Robust electrical
connections were made by threading fine wires through the contact pads several times or
making a connection using conductive epoxy (Microcircuit Silver Type O, Transene
Company, Inc., Danvers, MA).
35

2.3 Characterization of CPDMS Strain Sensors
The properties of the conductive PDMS and the strain sensors were evaluated in
terms of conductivity, resistance, thermal coefficient of resistance, and gauge factor to
determine their usefulness as a material to measure strain.
2.3.1 Conductivity of CPDMS
Conductivity of the CPDMS was calculated from sheet resistance measurements
using a four point probe (R
s
= 4.53(V/I) when probes have equal spacing, films are <40%
of probe spacing, and if the edges of the film are more than four times the spacing of the
probes [9, 12]). Initially, 35 × 45 mm rectangular patterns of CPDMS were printed onto a
non-conductive PDMS substrate to ensure that the edges effects were minimized.
However, subsequent four point probe measurements performed directly on un-
encapsulated sensors (1.8 × 30 mm) yielded identical resistance values, therefore four
point measurements were later made directly on exposed, unencapsulated sensors.
The percolation threshold was 5 wt% for MWNT-only and MWNT-GNP
composites. The percolation threshold for GNP composites was 12.5 wt% which can be
explained by the smaller aspect ratio of the platelets compared and possible non-uniform
dispersion. Once the percolation threshold is crossed, traditional percolation theory [37]
states that conductivity is related to the concentration of the filler material by the
following equation:
 
t
c com
     
0
for
c
   (1)
where σ
com
is the conductivity of the composite, σ
0
is the conductivity of the filler
material,  is the concentration of filler material, 
c
is the concentration of the filler
material at the percolation threshold, and t is an empirically determined exponent for each
system that is related to the aspect ratio, curvature, and degree of aggregation of the filler
material. This exponential trend has been confirmed in literature and agrees with
experimental results (Figure 2-9). Above concentrations of 13% filler material, the
36

MWNT composites were too viscous to screen print and the GNP composites did not
cure.

Figure 2-9. Conductivity data comparing experimental results against those for similar
composites found in literature (Lu et al.[12] used 40-60 nm OD MWNTs and Liu and
Choi [13] used 20-40 nm OD MWNTs). The fitted line follows Eq. 1 and is
 
02 . 3
499 . 0 95    
MWNT
. Reprinted with permission from [5]. Copyright 2013, AIP
Publishing LLC.

2.3.2 Zero current resistance
Zero current resistance was calculated by recording the resistances of the devices
from 1-10 µA using a precision multimeter (Keithley 2400; Keithley Instruments,
Cleveland, OH) and extrapolating the resistances back to zero current (Figure 2-10).
Measuring ZCR in addition to conductivity revealed the current dependency of
conductivity which was higher at lower conductivities. The dependency of resistance
with current has been suggested in literature to be associated with the increased number
of shells contributing to conductance at higher currents in MWNTs [38] and the linear
increase in prevalence of electron tunneling with applied voltage [33, 39]. The screen
printing process allowed three devices to be printed at one time and among each batch,
the standard error of the ZCR was as low as 3 % (data not shown). However, between
batches and across different filler concentrations, the standard error of the ZCR was
37

~20 %, with the exception of 69 % standard error for the 5 % MWNT devices. These data
suggest improvements in controlling the thickness of screen printed conductive layers
(e.g. automating the process) could reduce the standard error.

Figure 2-10. Zero current resistance (mean ± SE, n=6-12). Reprinted with permission
from [5]. Copyright 2013, AIP Publishing LLC.

2.3.3 Thermal coefficient of resistance
The temperature sensitivity of the composite, or temperature coefficient of
resistance (TCR), was determined by measuring the device resistance under a constant
current bias (10 µA) at different temperatures from 30 to 80 °C (10 °C increments) and
taking the slope of the resulting curve (Figure 2-11). The TCR for composites made from
MWNTs were shown to have a weak dependence on composition and be negative,
whereas the TCR for GNP devices were positive. Both of these results are consistent with
literature [40, 41], where it is suggested that filler materials with lower aspect ratios
(GNPs) will have a positive TCR due to polymeric expansion with temperature that
causes increased tunneling resistance which do not affect fillers with high aspect ratios
where contact resistance dominates [41].

38

Figure 2-11. TCR (mean ± SE, n=3-5). Reprinted with permission from [5]. Copyright
2013, AIP Publishing LLC.

2.3.4 Gauge factor
Gauge factor (GF) is a useful measure of merit for piezoresistive strain sensors
and is equal to the normalized change in resistance divided by the strain:
(6)
where R is the nominal as-fabricated, undeformed strain gauge resistance, ΔR is
the total change in resistance, L is the original unstrained strain gauge length, ΔL is the
change in length following application of strain, and ε is the applied strain. Metal and
semiconductors are typical materials used for strain gauges. Metal strain gauges typically
have relatively low gauge factors (GF~2). Although semiconductors have improved
gauge factors (GF~100), mechanical properties limit their use to low strain (< 0.1%)
applications. Thick film materials for strain sensing have reported GFs in the range of 2 –
12 [12, 13, 42].
Gauge factors of the strain devices were obtained using a motorized stage (Z812;
Thorlabs, Newton, NJ) that allowed for user defined strain patterns and a precision
multimeter (Keithley 2400; Keithley Instruments, Cleveland, OH). The sensors were
mounted to the motorized stage using custom acrylic clamps with a known separation
RR
RR
GF
L
L




39

distance and subjected to uniaxial strain using LabVIEW (National Instruments, Austin,
TX) to control the motor (Figure 2-12(a & b)). Gauge factor was calculated by averaging
the linear fit to relative resistance versus uniaxial strain data during the stretching phase
of 3-5 strain cycles (Figure 2-12(c)). (The coefficient of determination (R
2
) value of the
linear regressions ranged from 0.93 to 0.99.) Sensors exhibited a “breaking in” cycle, in
which the first 1 or 2 cycles had either a higher or lower gauge factor compared to
subsequent cycles ((Figure 2-12(b)), and these cycles were not used when calculating
gauge factor. This phenomena is likely due to the differing stiffness between the filler
materials and PDMS; the initial strain cycles likely displace some filler relative to the
PDMS. Two operation regimes were observed: a smaller GF at low strain, and a larger
value at higher strains ((Figure 2-12(c)). This phenomenon was previously observed in
other CNT-polymer composites in which it was hypothesized that at lower strains the
change in resistance is dominated by breaking electrical contacts between overlapping
CNTs, but at high strain the change is resistance is dominated by increasing the tunneling
distance between tubes that are already separated [43]. The present study focused on the
device’s maximum GF in preparation for high strain biomedical applications and
therefore, only the GF at higher strain was calculated.
Gauge factor for our composites ranged from 2 to 22 for high concentrations of
filler material (10wt% MWNT and 12.5wt% GNP) and up to 100 for lower filler
concentrations closer to the percolation threshold (5wt% MWNT and 3.5wt% MWNT +
1.5wt% GNP mixture) (Figure 2-12(d)). This phenomenon is attributed to the dominance
of electron tunneling at filler concentrations close to the percolation threshold; the
distance between filler particles increases when the sensors are stretched while still
allowing conduction through tunneling. For higher concentrations of filler material, the
connections are made mostly by direct contact, which are less effected by strain than
tunneling electrons. This result is similar to other reported strain sensors made from
graphene where it was found that the GF was dependent on initial resistance [44], and
suggests that high sensitive strain sensors can be made from cPDMS near the percolation
threshold of the filler material. The repeatability of GF varied based on concentration.
40

Lower filler concentrations were associated with much higher variabilities (standard error
values approaching ~20% of the GF), whereas the standard error of the GF of the higher
concentrations were within ~10% of the GF, suggesting that cPDMS at lower
concentrations were much more sensitive to processing than higher concentrations.

Figure 2-12. (a) Strain pattern applied to device. (b) Raw data of resistance versus time.
(c) Resistance versus strain (with first 2 cycles removed). Gauge factor obtained by
averaging slope of linear region. (d) Gauge factors for various compositions including
MWNT and/or GNP fillers (mean ± SE, n=4-6). Reprinted with permission from [5].
Copyright 2013, AIP Publishing LLC.

2.4 Conclusion
In summary, strain sensors featuring simple, low-cost construction involving the
screen printing of MWNT and GNP nanocomposites on biocompatible medical grade
41

polymer substrates was achieved. A gauge factor of over 100 was achieved with the 5%
MWNT material. The biocompatibility of the substrate, low cost of the GNPs, and strain
sensitivity and performance, including high strain operation and gauge factors, of these
sensors uniquely enable applications in wearable and implantable sensors in health
diagnostics and monitoring. One potential application is strain sensing to measure bladder
fullness where the bladder wall can experience strains up to 75% [45]. Traditional
semiconductor strain sensors fail at 0.7% [46] strain and the stiffness of silicon could
damage the tissue. Another application is strain sensors on the skin to measure activity
levels and joint movement.
42

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[35] M. Taya, W. Kim, and K. Ono, "Piezoresistivity of a short fiber/elastomer matrix
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[36] C. A. Gutierrez and E. Meng, "Low-cost carbon thick-film strain sensors for
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[37] G. Deutscher, "Percolation and Superconductivity," in Percolation, Localization,
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[38] R. Hobara, S. Yoshimoto, T. Ikuno, M. Katayama, N. Yamauchi, W.
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47

3.1 Matrigel and Penetrating Neural Electrodes
Penetrating multi-electrode intracortical probes provide a means to record both
extracellular field potentials and individual action potentials, thereby enabling study of
neural networks. There is great interest in using these neural recordings to drive
prostheses that restore lost functions to patients with spinal cord injuries,
neurodegenerative diseases, or other neural deficits [1, 2]. However, reliability of chronic
recordings are plagued by gradual signal degradation which is in part attributed to glial
scar formation around the probe, the associated increase in electrode impedance, and
retraction of dendrites away from the probe. The initial probe insertion may disrupt the
blood brain barrier and further contribute to the chronic inflammation associated with the
continued presence of foreign material in the brain [3].
While traditional neural probe formats take on the form of solid shanks with
electrode sites at the tip or shank surface (such as arrays of metal microwires that are
deinsulated at their tips and held together with an adhesive [4] or micromachined silicon
shanks decorated with exposed metal electrode sites [5, 6]), an alternate approach that
fundamentally changes the format and may provide access to reliable chronic recordings
has been proposed. The “neurotrophic cone electrode” concept introduced by Kennedy
consists of a small glass cone filled with a sciatic nerve segment containing neurotrophins
or Matrigel to attract neuronal growth into the cone where neural signals are measured
using the de-insulated tips of microwires in the cone interior [6, 7]. It was reported that
the neurotrophins eluted from the cone encouraged differentiation and neurite outgrowth
of neurons near the openings of the cone for repeatable ingrowth of myelinated axons and
CHAPTER 3
MATRIGEL COATINGS FOR PARYLENE SHEATH NEURAL PROBES
48

occasional cell bodies [8] from which action potentials were recorded for up to 4 years in
human subjects [9].
Despite the success of the neurotrophic cone electrode concept, the labor-
intensive manual fabrication and assembly process has limited widespread use of the
technology (low production yield and fragile construction) and the achievable electrode
density. The cone is constructed from glass (E ~100 GPa) which is significantly stiffer
than neural tissue (3.2×10
-6
GPa [4]) and may induce further damage during micromotion
experienced during blood pulsing or head motion. To overcome these limitations while
preserving the overall neurotrophic cone electrode concept, the Parylene sheath electrode
(PSE) was introduced. The PSE consists of a Parylene sheath structure with 4 platinum
electrodes lining the inside of the sheath and 4 electrodes on two thin wings located on
both sides of the sheath (Figure 3-1) [5]. Parylene C is flexible polymer and two orders
of magnitude softer than glass (E~4 GPa), thereby partially alleviating the mechanical
mismatch between the tissue and probe substrate.

Figure 3-1. (a) Micrograph (scalebar = 100 µm) and (b) three dimensional illustration of
PSE. Image from [6]. © 2015 Wiley Periodicals, Inc.

Parylene C can be structured using microfabrication, which permits flexibility in
PSE design and precision PSE manufacture in large batches. Details of the
microfabrication process were previously described [11, 12] and summarized here for
completeness. A 5 µm layer of Parylene C (Specialty Coating Systems, Indianapolis, IN)
49

was first deposited on a bare silicon wafer (Figure 3-2(a)). Platinum metal (200 nm)
deposited using e-beam evaporation was patterned using a lift-off process. An overlying
Parylene insulation layer (1 µm) was applied and etched with O
2
plasma to expose the
electrodes. Following this, a thick photoresist layer (~10 µm) was spun on and patterned
to form the microchannel that defined the inner lumen of the of the 3D Parylene sheath
(Figure 3-2(b)). Another final layer of Parylene (5 µm) was deposited on top of the
photoresist structures to complete the sheath structure, and O
2
plasma was used to etch
open both ends of the microchannel, expose the surface of the electrodes and contact pads,
create perforations through the sheath, and cut out the device (Figure 3-2(c)). After the
devices were gently removed from the wafer, the microchannel was thermoformed into a
cone shape by inserting a tapered microwire mold into the microchannel and heating the
device to 200° under vacuum (Figure 3-2(d & e)).

50

Figure 3-2. Simplified schematic of fabrication steps. (Note: fabrication steps are drawn
on a single cut-off probe, however probes are actually etched out process step (c).) (a)
Parylene (5 µm) was deposited onto a Si carrier wafer. Pt electrodes (200 nm) were
defined using e-beam evaporation and lift-off. Parylene insulation (1 μm) was deposited
and the electrode sites exposed using O
2
plasma etching. (b) A sacrificial photoresist
structure was used to form the microchannel structure. (c) An overlying Parylene (5 μm)
layer completed the enclosed microchannel. The ends of the microchannel were etched
open and, in the same etch step, the individual probes were cut out. Perforations
(optional) were sequentially etched at each previous etch step and completely etched
through during this final etch step. (d) After carefully removing each probe, a microwire
mold was inserted into the sheath to form the 3D sheath structure. (e) The final
mechanically stable sheath shape defined by the microwire was retained after a
thermoforming process. (Figure modified from [7]. © IOP Publishing. Reproduced by
permission of IOP Publishing. All rights reserved.)

For both the cone electrode and the PSE, it is desirable to have extracellular
membrane proteins and bioactive molecules, such as neurotrophins, to encourage
neuronal integration. The initial source of these molecules in the cone electrode was an
autologous section of sciatic nerve [8], but due to potential complications and the
complexity of the surgery, the nerve segment was replaced by Matrigel (MG) [9] and
nerve growth factor (NGF) [10]. MG is a biologically derived collection of basement
membrane proteins and growth factors, which gives MG unique biological properties that
have made it useful in a variety of applications, such as a surface treatment to promote
differentiation in cell cultures, a thick gel to induce specific differentiation in tissue
explants (such as aortic rings), and a plug injected in vivo to measure the effect of drugs
51

and growth factors on the ingrowth of surrounding tissues [11]. When used as a filler in
the neurotrophic cone electrode, MG supported neuronal growth into the cone next to the
internal wire electrodes [12]. Additionally, MG has been used to support transported
neuronal stem cells in the brains of rats [13] and mice [14]. MG was chosen as a coating
material because its unique composition has been shown to promote neuronal cell
attachment and differentiation in vitro [15] as well as promote neuronal survival in vivo
[13, 14].
MG used was used either by itself or supplemented with an immunosuppressant
and combination of neurotrophins to improve neuronal integration onto the surface of the
probe. The surfaces of neural probes may be modified to modulate the immune response
and improve signal quality recorded by implanted electrodes. Dexamethasone (DEX)
incorporated into a biodegradable coating has been shown to suppress inflammation and
inhibit the immune response [16, 17]. Similar biodegradable coatings loaded with
neurotrophins (including NGF [18, 19], brain-derived neurotrophic factor [20], or
neurotrophin-3 [21]) have been shown to maintain neuronal density and encourage
differentiation and outgrowth of neurons. Electrode surface modifications with anti-
inflammatory peptides [22], adhesion molecules such as laminin [23], or conductive
polymers [24] have also been shown to reduce chronic inflammation and enhance tissue
integration.
Although the direct application of the coatings developed is for implantable multi-
electrode arrays, there are also implications on patterned cell cultures that use Parylene C
(with or without electrodes) and other electronic devices that are coated with Parylene
prior to implantation. Parylene C is a USP class VI biocompatible polymer that is inert,
non-biodegradable, and vapor deposited at room temperature to form pinhole free,
micromachinable, conformal coatings; these thin film polymer coatings have been used
extensively for insulating implantable biomedical devices [25] and have recently been
used as a structural material as well [26-28]. Because of its biocompatibility and
compatibility with microfabrication, there has also been interest in using Parylene C for
52

cell cultures and other applications that require cell adhesion; however, cells will not
adhere to hydrophobic Parylene C without first modifying the surface. Parylene C surface
treatments include exposure to O
2
plasma to render the surface hydrophilic [29], texturing
the surface to create variable wettability [30], or soaking in horse serum [31] or other
proteins [29] to present natural binding sites to cells.
Parylene C coupons and PSEs were coated with MG to promote cell adhesion and
encourage neurite ingrowth in vivo in order to improve neural recordings for neural
probes such as the PSE. MG was chosen as a coating material because its unique
composition has been shown to promote neuronal cell attachment and differentiation [15].
The modulus of MG (0.034-0.440 kPa [32-34]) is also much closer to the modulus of
brain tissue (3.2 kPa [4]), and thus could provide an additional mechanical buffer from
the more rigid Parylene (4,750 kPa [35]). The MG coated surfaces were characterized for
wettability, uniformity of adhesion, thickness, and effect on impedance and stability over
time. On a subset of probes, MG was loaded with DEX, NGF, and BDNF and release rate
was measured. The PSE was used to record neural events from rat motor cortex for 3
months and the SNR from signals obtained from coated (with either MG or MG
supplemented with an immunosuppressant and neurotrophins) and uncoated probes were
compared.
3.2 Preparation of MG-based coatings
Coatings consisted of MG (BDbiosciences, San Jose, CA) with either phosphate
buffered saline (PBS) (EMD Millipore, Billerica, Massachusetts) or a combination PBS
and the bioactive molecules: water soluble DEX (D2915 Sigma Aldrich, St. Louis, MO,
USA), NGF (N2513 Sigma Aldrich, St. Louis, MO, USA), and BDNF (amsBio,
Abingdon, UK) mixed together using a vortex mixer at 4 °C. All coatings contained a
final composition 75 vol% MG and 25 vol% PBS, which was loaded with bioactive
molecules. The concentration of MG was kept above 50% per manufacturer’s guidance to
retain its ability to gel. The composition of the different coatings used for the studies is
summarized in Table 3-1.
53

Table 3-1. Composition of Coatings Used in Study
Coating Name
(abbreviation)
MG
(vol%)
PBS
(vol%)
DEX
(mg/mL)
NGF
(µg/mL)
BDNF
(µg/mL)
MG 75 25 x x x
MG/DEX 75 25 1.6 x x
MG/NT 75 25 x 100 10
MG+ 75 25 1.6 100 10

Concentrations of bioactive molecules were based on their solubility. DEX is
soluble in PBS at ~20 mg/mL and NGF at 1 mg/mL. For the final coating, 100 µL of PBS
filled with 16 mg/mL DEX and 1 mg/mL of NGF was added to MG to obtain final
concentrations of 1.6 mg/mL and 100 µg/mL, respectively. The concentration of BDNF
was based on its availability of 10 µg, which was diluted into 50 µL of PBS and added to
the coating mixture. The final concentration of DEX is similar to other slow-release DEX
coatings that have reduced astrocytic responses [36]. Others reported using 1 mg/mL
NGF solution added to hydrogel to form coatings [37] or reported a release of 8 ng of
NGF from a hydrogel [19], which would be available from the amount of coating
deposited on our probes. A hydrogel coating containing 100 µg/mL of BDNF was used to
support the growth of neurons; the eluate from the hydrogel was reported to have a
similar affect as adding 100 ng/mL solution directly to a cell culture [20].
Parylene C test coupons were 5 × 10 mm pieces, cut from a silicon wafer coated
with 8 µm of Parylene C (Specialty Coating Systems, Indianapolis, Indiana), which
required pretreatment of the wafer with the adhesion promoter silane A-174 (Momentive
Specialty Chemicals, Columbus, OH) to retain the deposited Parylene film. Additionally,
600 × 5000 × 10 µm Parylene C coupons were cut using a razor blade from 10 µm thick
freestanding Parylene sheets. These smaller coupons were coated using the same methods
used to coat the PSEs (described in next section) and were used to measure the cross
sectional thickness of the MG coating.
54

3.3 Coating of Parylene Sheath Electrode
The Parylene sheath on the PSE is in the shape of a truncated cone with a 94 µm
base tapering down to a tip with a radius of 32 µm over a height of 652 µm (Figure 3-1).
The sheath is designed not only as a space to carry immunosuppressant and neurotrophic
drugs to the insertion site in order to mitigate inflammation and scar tissue, but also
provides a surface that encourages the integration of surroundings neurons into the sheath
[5, 7].
The amount of drug carried in the interior volume of the sheath is limited
compared to the 3 mm
2
of surface area available to be coated on the wings, outside edges
of the sheath, and the implanted portion of the cabling for a total of 8.9 nL. In order to
maximize drug loading and provide surfaces for neuronal integration, it was desirable
that not only the lumen be filled, but that the surrounding edges be coated as well.
The hydrophobicity of the Parylene sheath requires strategies designed to
incorporate water-soluble drugs into the lumen of the sheath [38]. Two different
approaches were used to overcome the low surface energy of Parylene. In the first
method, the probe tip was immersed into coating solution contained in a centrifuge tube
and then submersed in an ultrasonic bath (Branson Ultrasonics, Danbury, CT, USA).
After 5 minutes of sonication at 4 °C, the probe was removed from the tube and bath,
after which the coating was allowed to gel for 5 minutes followed by air drying for at
least 5 minutes to reduce the profile of the cone and prevent the coating from wicking
onto other surfaces as it was prepared for implantation (Figure 3-3).

55

Figure 3-3. Schematic of the sonication method. (1) Probe is placed in coating mixture
which is placed in sonication bath. (2) Coating is gelled and dehydrated at room
temperature. Image from [6]. © 2015 Wiley Periodicals, Inc.

In the second method, the Parylene was first treated with the positively charged
molecule poly-D-lysine (PDL) (P6407, Sigma Aldrich, St. Louis, MO) in order to
increase the surface energy. Polylysine (a group of molecules that includes PDL) is a
polycationic molecule that easily adsorbs to solid surfaces (including Parylene C [39])
and creates cationic binding sites [40]. Polylysine was chosen because it has been widely
studied as a molecule used in biomaterials (mainly in the microencapsulation of
exogenous cells [41]). The cationic binding sites increase the surface energy, which
makes the surface hydrophilic. Once the PSE was coated with PDL, the probe was
inserted into a 10 µL droplet of the coating mixture and incubated at 50 °C inside a
polystyrene container for 5 minutes. After the 5 minutes, the probe was removed from the
container and held at 50 °C for an additional 5 minutes to dehydrate the coating (Figure
3-4).

56

Figure 3-4. Schematic of the surface modification method. (1) Probe is soaked in PDL
solution for 1 h. (2) Probe placed inside of 10 µL droplet for 5 minutes covered in a
polystyrene container. (3) Probe is taken out of container and is dehydrated for 5 minutes.

3.4 Characterization of MG coatings
The coating was visualized to determine its structure and distribution on the
Parylene substrate using optical and scanning electron microscopy (SEM; 7001, JEOL,
Peabody, MA). Optical microscopy showed the structure of MG on a Parylene substrate,
whereas SEM was used to determine the distribution of MG over the PSE, as well as the
coverage of MG on the platinum and Parylene surfaces of the probe.
Electrochemical impedance spectroscopy (EIS) is widely used to assess the
recording capability of microelectrodes [42]. Here, EIS was used here to determine the
change in impedance at the electrode surface caused by coating the electrodes with MG
and how this change varied over time as the coating was soaked in PBS. EIS was
performed in 1× PBS as 37 °C with an amplitude of 10 mV
rms
and a frequency range of
1-100,000 Hz. A large area Pt plate was used as a counter electrode and an Ag/AgCl (3M
NaCl) electrode was used as a reference. EIS data was taken once a day for 3 days. In
between tests, probes were soaked in 1× PBS at 37 °C.
3.5 Measuring drug loading and release from probe in vitro
To obtain a continuous measurement of drug release from MG, 12.5 µL of
MG/DEX coating was placed on a bare microwell sidewall and, after heating, uncovered
for 15 minutes at 55 °C (Sun Systems, Titusville, FL) to remove the liquid from the
57

coating. 100 µL of PBS was added to the microwell. The microplate was then
immediately covered with Parafilm to avoid evaporation and then was placed into a
spectrophotometric microplate reader (Epoch, Biotek, Winooski, VT) and scanned at
various time intervals for several hours. The absorbance was compared to a calibration
curve that was made by measuring the absorbance (at 242 nm) of known concentrations
of DEX in PBS. In this manner, the mass of DEX released from the MG could be
calculated.
To measure the effect of EtO sterilization on the release of DEX from MG
deposited using the surface modification method, 5 µm of Parylene was deposited
directly on a 96 well plate. The wells were then coated with PDL by placing 100 µL of
100 µg/mL PDL solution into the wells and soaking for 1 hour at 4° C. Excess PDL was
then removed by rinsing the wells 3 times with deionized water. 12.5 µL droplet of MG
loaded with DEX (1.6 mg/mL) were then deposited into the bottom of the microwell
which was then covered with Parafilm and heated to 55 °C for 5 minutes. The Parafilm
was removed and the well was held at 55° for an additional 5 minutes. The coated wells
were then subjected to a 24 h EtO sterilization using a room-temperature sterilizing
system (Anprolene, AN74i, Andersen Products, Haw River, NC). After sterilization, the
wells were filled with 100 µL of PBS at room temperature. At various time intervals, the
PBS was removed and replaced with fresh PBS. The eluate was placed in a UV
transparent microwell scanned at 242 nm and a calibration curve was used to determine
the amount of eluted DEX. The results were compared to wells that had not been
sterilized.
Parylene-coated microwells treated with PDL were also used to measure the
amount of neurotrophic factors released from MG deposited onto Parylene using surface
modification. For these tests, 1 µL droplet of MG/NT coating (100 µg/mL NGF, 10
µg/mL BDNF) was deposited into the bottom of 96 well plates previously coated with 5
µm of Parylene and treated with PDL. The droplet was then covered with Parafilm and
heated to 55 °C for 5 minutes. The Parafilm was removed and the wells were held at 55°
58

for an additional 5 minutes. The wells were then either filled with 100 µL of PBS at room
temperature or sterilized for 24 hours using a room-temperature sterilizing system
(Anprolene, AN74i, Andersen Products, Haw River, NC) and then filled with 100 µL of
PBS at room temperature. At various time intervals, the PBS was removed and replaced
with fresh PBS. The eluate was then diluted and the concentration determined using
enzyme linked immunosorbent assay (ELISA).
To determine the amount of drug that could be loaded on a probe and to confirm
that drugs were released from the probe surface the same as the coupons, probes coated
with MG/DEX (1.6 mg/mL DEX) using the surface modification method were submersed
into 100 µL of PBS contained in a UV-transparent microwell (96 well microplate) and
removed at various intervals to measure the absorbance of the PBS. The absorbance of
the eluate was compared to a calibration curve to back calculate the mass of DEX
released from the coated probe.
3.6 In vivo electrophysiological evaluation from 1-month study
Eight male Sprague Dawley rats (> 320g) were implanted with dual-probe arrays
coated with either MG (n = 3), MG/NT (n = 3) or MG/DEX (n = 2) (using the sonication
method) for 1 month. All procedures for the animal experiments were in accordance with
the animal protocol approved by the Huntington Medical Research Institutes Institutional
Animal Care and Use Committee (HMRI IACUC) and in compliance with the Animal
Welfare Act. The implantation procedure is as described earlier [7]. Briefly, after coating
the probes with their respective coatings, the coated probes were additionally coated with
an overcoat of PEG 8000 (PEG; MW 8000, Sigma Aldrich, St. Louis, MO, USA) to
stiffen the probe during insertion. These probes were then sterilized for 24h at room
termpeture using an ethylene-oxide system (Anprolene, AN74i, Andersen Products, Haw
River, NC, USA) and stored in sterile packaging at 4 °C until the time of implantation
(~1-2 days). At the time of implantation, probes were manually inserted in the motor
cortex with a micromanipulator until reaching the required depth of 2 mm.
59

To assess the in vivo functionality of the probes, weekly EIS and neural
recordings were taken and used to calculate impedance at 1 kHz, unresolved neural noise,
signal-to-noise ratio (SNR), and event rate. During the electrochemical and
electrophysiological recordings, the rats were anesthetized with Ketamine/Xylazine
(90/10 mg/kg, IP). A PC4/300 potentiostat system (Gamry Instruments, Warminster, PA)
was used to take 2 point impedance measurements with the reference and counter
electrode being connected to a titanium headplate with 6 stainless steel screws on the
rat’s skull which held the probes in place. The probes were scanned with 10 mV
rms

sinusoids ranging from 1 Hz to 1 kHz.
Neural recordings were made using a 64 channel data acquisition system
(OmniPlex; Plexon Inc., Dallas, TX) which measured signals at 16 bit and 40 kHz per
channel. The data was high-pass filtered at 300 Hz to remove low frequencies variations
from the baseline. A nonlinear energy operator (NEO) was used for spike detection,
which is more accurate than amplitude thresholding at low signal to noise ratios [43]. The
spike detection threshold was 14 times the standard deviation of the NEO values recorded
during 120 s of measurement. Neuronal noise was equal to the standard deviation of the
recorded data following the removal of 0.8 msec segments that contained detected spikes.
Spike amplitude was defined as the average absolute value of the spike’s peak height and
average spike amplitude was the average of all detected spike like events. Because the
animal was anesthetized and in a Faraday cage (along with the headstage and
preamplifier) during the recordings, spike sorting to remove false positives was not
performed as it was assumed that electromyographic artifacts and electromagnetic
interference was eliminated. SNR was calculated as the average spike amplitude divided
by the average noise. The event rate was defined as all spike-like events during 120 s
recording session and was only calculated (along with SNR) if more than 10 spikes were
detected during that time period.
60

3.7 Characterizing immunohistochemical effect of drug release from
probe in vivo from 1-month study
At one month post-implantation, following the last electrochemical and
electrophysiological measurements, the animals were transcardially perfused with PBS
followed by phosphate buffered 4% paraformaldehyde solution. The cerebral cortex was
dissected and the cortical tissue block containing the probe tracks (and occasionally the
probes) was embedded into paraffin. The tissue was sectioned perpendicular to the probe
tracks. The sections were then immunostained for GFAP (marker for reactive astrocytic
processes) and visualized using Vector nickel-DAB. The sections were subjected to the
semi-automated image analysis using software, custom-written in Visual Basic 6.0
(Microsoft Co., Redmond, WA) using National Instruments Image ActiveX component
(National Instruments Co., Austin, TX). Within the software, a rounded rectangle was
placed on the image for use during analysis; its placement, width and length were
adjusted to match the perimeter of the probe track. Then, six larger concentric rounded
rectangles were automatically drawn at 50 μm incremental distances from the inner one
and the GFAP density was automatically calculated in these six areas and normalized to
the average density in outer-most area (from 270 to 300 μm).
3.8 In vivo EIS and electrophysiological evaluation from 3-month study
Three male Sprague Dawley rats (> 320g) were implanted with four-probe arrays
for 3 months (one array with uncoated probe, one with MG coating, and one with MG
coating supplemented with DEX, NGF, and BDNF). To assess the functionality of the
coated PSE, in vivo EIS and neural recordings were carried out weekly from week 1 to 14
post-implantation, and the collected data were used to calculate the EIS and signal-to-
noise ratio (SNR). For the EIS and electrophysiological measurements, rats were
anesthetized with Ketamine/Xylazine (90/10 mg/kg, IP). EIS measurements were
obtained using a PC4/300 potentiostat system (Gamry Instruments, Warminster, PA) in a
two-electrode configuration, with the reference and counter connected to the titanium
headplate and its six transcranial stainless steel screws. The data were collected with 10
mV
rms
sinusoids at the frequencies from 1 Hz to 100 kHz, and 1 kHz values were selected
61

for analysis. The electrophysiological data were acquired at 16 bit and 40 kHz per
channel using a 64-channel data acquisition system (OmniPlex; Plexon Inc., Dallas, TX)
and high-pass filtered at 300 Hz to remove the low-frequency fluctuations from the
baseline. In the 120 second data records, spike detection was performed using the NEO
algorithm. The neuronal noise was calculated as the standard deviation of the data after
removal of 0.8-msec-long segments containing the detected spikes. The spike amplitude
was calculated as the absolute value of spike’s peak height, and the SNR was calculated
as the ratio of average spike amplitude to noise.
3.9 MG on Parylene and Electrode surfaces
Untreated Parylene (deposited on silicon coupons) dip coated in MG yielded an
uneven distribution of thick and thin regions of MG similar to the coating appearance on
an untreated PSE (as described in [38]). The thicker regions of MG on the coupon
revealed fractal patterns (Figure 3-5), which is a behavior predicted by diffusion limited
aggregation theory applied to collagen [44, 45] and shown in many other natural and
synthetic polymer systems [46-49]. These fractal patterns also demonstrate MG’s ability
to self-assemble into microstructured topography, which has been shown to be conducive
to neuronal attachment [48].

Figure 3-5. Optical images of fractal ‘‘trees’’ made from dried MG on a Parylene coated
silicon wafer under (a) unpolarized and (b) polarized light. (Scale bars are 50 µm.)

62

The 600 × 5000 × 10 µm Parylene coupons coated with MG by sonication or
surface treatment with PDL both exhibited uneven distribution across the surface. Pools
of MG formed along the center of the coupon similar to what was seen on the PSE cable
just above the sheath (Figure 3-6). For the PDL treated coupons, the average thickness of
the coating was 12.9 ± 5.2 µm (mean±SD, n=3) in a representative sample and ranged
from a minimum of 3.4 µm near the edge of the cable to 33.2 µm in the center where the
foliated texture of the pooled MG is apparent. Coupons coated using sonication (no PDL)
did not retain enough MG to measure cross sectional thickness using optical microscopy.

Figure 3-6. SEM image of the top portion of a perforated PSE coated using surface
modification. The MG pooled towards the center of the cable near the large opening of
the sheath and filled in the perforations which were added to the PSE to improve cellular
communication across the Parylene substrate. Image from [6]. © 2015 Wiley Periodicals,
Inc.

SEM images of electrodes on the PSE reveal platinum at the bottom of a recess
formed by etching away Parylene insulation (Figure 3-7(a)). When coated with PDL,
crystals appeared that cross from Parylene down to the platinum surfaces (Figure 3-7(b)).
After coating with MG, MG was present on both the Parylene and platinum surfaces as
evidenced by the change of surface roughness around the edges of the Parylene insulation
and the appearance of characteristic MG fractal patterns (shown in Figure 3-5). Other
63

work shows electrodes on PSEs coated using sonication showed similar coverage of
platinum and Parylene surfaces [38].

Figure 3-7. (a-c) SEM images of recessed platinum electrode on Parylene probe (a)
before coating, (b) after PDL surface treatment, and (c) after coating with MG. (Scale
bars are 10 µm. PSE was sputtered with Au prior to imaging.) Image from [6]. © 2015
Wiley Periodicals, Inc.

EIS revealed that impedance did not significantly increase after being dip coated
via sonication (Figure 3-8(a & d)). For probes coated with only MG, the impedance
decreased over time indicating the gradual removal of MG on the surface of the electrode
(Figure 3-8(c)). In contrast, the impedance electrodes having MG loaded with DEX
increased over time (Figure 3-8(f)). EIS data measured in vivo from probes coated using
the surface modification method similarly showed a change in impedance that lasted less
than 2 weeks.
64

Figure 3-8. EIS data (mean ± SE) from (a-c) MG coated (n=7) and (d-f) MG and
MG/DEX coated (n=3) neural electrodes (using sonication method). (a & b; d & e)
Magnitude and phase data were measured pre and post coating. (c & f) Impedance
measured at 1 kHz from before the probe was coated to 3 days after coating. Image from
[6]. © 2015 Wiley Periodicals, Inc.

As mentioned earlier, EIS measurements have been correlated with recording
performance of electrodes in vivo, with impedances of less than 1 MΩ being preferred for
recording electrodes. The lack of a significant increase in impedance associated with the
coating process is expected from a hydrogel with large pores (2 µm at 50% MG [33]) and
is similar to results from other hydrogels deposited on neural electrodes, such as sodium
alginate [24]. Lower impedances also portend well for in vivo recordings. The decrease in
impedance over time associated with the MG coating suggests gradual removal of bulk
coating from the probe surface. Visual inspection after soaking for 1 day also suggests a
reduction in the amount of MG present on probe surfaces. The increase of impedance
over time associated with DEX loaded MG is likely due to the water soluble DEX used in
this study that was incorporated into cyclodextrin, which has been shown to swell in
65

aqueous solutions [50]. Even with the increase, however, the 1 kHz impedance never
exceeded 1 MΩ.
3.10 Drug release from coating
DEX/MG deposited onto the sidewall of the microwell (with no surface
treatment) with a pipette was hydrated with PBS solution and continuous kinetic
absorbance scans measured DEX release (at 242 nm) (Figure 3-9(a)). These scans show
the maximum absorbance occurring within 35 min after the PBS was added to the well
(Figure 3-9(b)). The calibration curve, of known concentrations of DEX at 242 nm, is
shown in Figure 3-9(c).

Figure 3-9. Kinetic scan of microwells with MG/DEX and MG coated on microwell
sidewalls. (a) Schematic of test setup showing MG/DEX coating diffusing from bottom
left corner of microwell. (b) Kinetic scan showing peak absorbance at 0.5 hour (mean ±
SD, n=3-5) (c) Calibration curve of known dilutions of MG/DEX coating. Image from [6].
© 2015 Wiley Periodicals, Inc.

DEX being released within the first 35 min after the PBS was added to the well
(Figure 3-9(b)) corresponds to a diffusion coefficient of 9.4×10
-10
m
2
/s which is faster
than predicted by Einstein-Stokes theory for diffusion of DEX in water (6.82 ×10
-10
m
2
/s
[51]) and suggests that MG does not hinder the diffusion of DEX into solution.
To verify that coating does not modify DEX release, the rate of release of DEX
from MG coated onto a Parylene coupon using the PDL surface modification method was
measured using spectrophotometry at 242 nm. DEX was released within 1 day of soaking
66

in PBS (Figure 3-10). Sterilizing the coating using the EtO gas process did not change
the release rate, but slightly decreased the amount of DEX released (9%).

Figure 3-10. DEX released from MG deposited on Parylene coupons using the PDL
surface modification method before and after sterilization (n = 3, mean ±SE). Image from
[6]. © 2015 Wiley Periodicals, Inc.

The rate of NGF and BDNF release from MG deposited on a Parylene coupon
using the surface modification method was measured using ELISA. These measurements
indicate that the NGF eluted from the MG within 4 hours. (Figure 3-11(a)), whereas the
BDNF had eluted out of the MG within 1 day (Figure 3-11(b)). This is only a fraction of
the total amount of NGF and BDNF initially deposited onto the Parylene (100 ng and 10
ng, respectively). After sterilization, the amount of NGF and BDNF was reduced further
(46% and 35%, respectively) from the amount released from the non-sterilized coating.

67

Figure 3-11. Cumulative release of (a) NGF and (b) BDNF from MG loaded with
neurotrophins and deposited on Parylene using the PDL surface modification method
before and after EtO sterilization (n=3, mean ±SE). Image from [6]. © 2015 Wiley
Periodicals, Inc.

NGF and BDNF eluted more slowly from MG than DEX (within 4 h and 1 day,
respectively), which is likely due to their larger sizes (NGF and BDNF are 27 kDa
compared to DEX, which is 392 Da), but this rate is likely too fast to have an effect on
chronic recording quality. The loss of factors during the elution process could be due to
protein adhesion to the sidewalls of the wells and denaturing of the proteins. During
sterilization, the extra time at room temperature and exposure to EtO could cause further
denaturing of the proteins. During actual recordings in vivo, the proteins will not be
exposed to the large hydrophobic surface area of the microwell and therefore more will
be available to neurons. Sterilizing the probes prior to coating and coating under sterile
conditions could eliminate the need for post-coating sterilization and preserve a larger
portion of the neurotrophic factors.
The rate of DEX release from a probe coated using PDL surface modification was
also verified using spectrophotometry. For each measurement, the coated probe was
dipped into a microwell containing PBS for a given time period and then removed to
allow spectrophotometric scanning (Figure 3-12(a)). The absorbance from the eluate of
probes coated using the surface treatment method at 242 nm was 0.3 after 2 hours
68

(Figure 3-12(b)) and corresponds to 1.2 µg of DEX being released from each probe. The
maximum absorbance measured occurred during the first time period, which showed that
the majority of the DEX eluted out before the second measurement (i.e. within 30
minutes). These results verify that the release profile measured by kinetic scans of DEX
being released from MG deposited with no surface treatment is similar to the release
profile from probes coated using surface modification.

Figure 3-12. Drug loading and release from PSE coated using surface modification
method. (a) Schematic of testing method: 1) probe coated in DEX loaded MG, 2) coated
probe inserted into microwell, 3) probe soaked for time t
n
, 4) probe removed from
microwell and absorbance of eluate measured using spectrophotometer at a wavelength
of 242 nm. Probe is reinserted into well after measurement for subsequent measurements.
(b) Absorbance of eluate from coated with DEX/MG. (Absorbance of probe coated with
MG and PBS filled well is also provided to show ability of spectrophotometer to
selectively measure DEX concentrations using a 242 nm light source. See Figure 3-9(c)
for calibration curve.) Image from [6]. © 2015 Wiley Periodicals, Inc.

3.11 Electrophysiological measurements from 1-month study
Electrophysiological measurements showed that the addition of DEX or
neurotrophic factors to the MG significantly increased the event rate of spikes detected,
but did not have an effect on the 1kHz impedance, SNR, or noise. The small size of this
improvement could be due to such factors as the quick release of factors from the coating,
the short duration of the trial, or denaturing of the coating during sterilization.
69

Figure 3-13. (a) 1kHz impedance, (b) SNR, (c) noise, and (d) event rate measurements
and results obtained in comparing different coatings at 28 days (mean ± SD, n = 121
recording sites in 9 animals, 16 probes). The probes had the following coatings:
MG/DEX (n = 2), MG (n = 4), and MG/NT (n = 4). (Figure reproduced from [7]. © IOP
Publishing. Reproduced by permission of IOP Publishing. All rights reserved.)

3.12 Immunohistochemical effect of drug release from probe in vivo
from 1-month study
The MG-coated and MG/DEX-coated probes (coated using the sonication
method) were implanted in the rodent cerebral cortex for 1 month in order to evaluate the
in vivo response to DEX elution from the probe. The results from this study are shown in
Figure 3-14. The anti-astrocytic GFAP antibody was used to evaluate the immune
response to a chronic probe presence in the cerebral cortex. Since the probes with MG
and DEX coatings were implanted in different animals, three-step normalization was
performed. First, the GFAP immunoreactivity was calculated from the actual edge of the
probe rather than from the center of the probe track in order to reduce the confounding
effect of different track diameter (which is due to a tapering shape of the probe). Second,
a section of the image containing any artifacts (e.g., tissue folding or presence of a large
70

blood vessel) was removed from quantification. Third, the GFAP immunoreactivity was
normalized based on the level remote to the probe (at the 270-300 µm distance), in order
to reduce animal-to-animal and section-to-section variability in the overall level of GFAP
staining. The immunohistochemistry did not reveal a significant difference in astrocytic
density between the probes coated with only MG compared to the probes coated with MG
supplemented with DEX.

Figure 3-14. (a) Sample microphotographs of the GFAP immuno-stained sections of
cerebral cortex through the probe tips coated with MG (left) and MG/DEX (right); (b)
quantification of GFAP immunoreactivity in the sections shown in (a) at increasing
distances from the probe. The immunoreactivity data are normalized to the level at the
distance of 270-300 µm. Image from [6]. © 2015 Wiley Periodicals, Inc.

The immunohistochemical results indicate no difference between the normalized
GFAP intensity around the MG-coated probe as compared to the MG/DEX-coated probe,
suggesting that the DEX was released too quickly to produce any significant effect on the
chronic astrocytic response. These results are consistent with electrophysiological
recordings, which showed little difference in SNR and noise levels between probes
coated with only MG and probes coated with MG supplemented with bioactive molecules
[7]. However, the histology fails to explain why, in the same study, both MG
supplemented with DEX and MG supplemented with neurotrophins showed a statistically
higher (p < 0.001, and p = 0.001, respectively) event rate [7].
71

3.13 Electrophysiological measurements from 3-month study
Probes coated with either MG, MG supplemented with DEX and neurotrophic
factors (MG+) (coated using the surface modification method) or non-coated probes were
implanted in the rodent cerebral cortex for 3 months in order to evaluate the functionality
of the coated versus uncoated probes. The results from this study are shown in Figure
3-15. EIS data taken at 1 and 2 weeks show that the presence of the hydrogel
significantly lowers the impedance of the electrode during the first week. By the second
week, however, these differences are no longer significant (Figure 3-15(a)).
Electrophysiological data taken at 2 and 14 weeks show a significant improvement in
SNR over time for two MG-coated probes, but not for a non-coated probe. While the
SNR was not significantly different among the probes at week 2, by 14 weeks the probes
coated with MG and with MG supplemented with DEX, NGF, and BDNF exhibited a
significantly better SNR than the uncoated probe (Figure 3-15(b)).

72

Figure 3-15. In vivo EIS and electrophysiological data from 3 month chronic study. (a) In
vivo EIS data taken during the first two weeks comparing three probes: one non-coated,
one coated with MG, and one coated with MG supplemented with DEX, NGF, and
BDNF (denoted MG +) at 1 week and 2 weeks after implantation (n = 13-35, mean ± SD).
(b) Comparison of SNR for neural activities recorded with three probes (n=6-12, mean ±
SD). The p values are for ANOVA comparisons among three coating groups at each time
point. Error bars indicate the standard deviation for sites within the probe. Asterisk
indicates significant difference of non-coated group vs. two other groups in the post-hoc
test. Image from [6]. © 2015 Wiley Periodicals, Inc.

The absence of a significant difference in electrical impedance between coated
and uncoated probes at two weeks post-implantation corroborates the conclusions from
the in vitro data that the bulk of the hydrogel is removed from the electrodes during the
first two weeks. The observation of lower impedance in MG-coated groups at one week
in vivo instead of higher initial impedance seen in the bench top saline test is likely due to
the hydrogel being more conductive than brain tissue, but less conductive than saline.
The SNR measured in vivo for 14 weeks was improved over time for probes that
were coated with either MG by itself or MG supplemented by DEX, NGF, and BDNF.
There was no measurable benefit from supplementing the MG with DEX, NGF, and
BDNF, consistent with the drug elution data that showed that all drugs loaded into MG
are eluted out within the first day. However, both MG-coated probes (supplemented and
non-supplemented) had considerably better functional performance than uncoated probes.
The stability and long term efficacy of MG were previously demonstrated in literature
a
b
73

where MG has been used as a scaffold to support survival and differentiation of
transplanted neural precursor cells [13, 14].
3.14 Conclusion
These experiments show that MG can be coated onto a Parylene based neural
probe to improve performance. Previous studies have shown that Parylene surfaces can
be modified to improve its performance for cell culturing [29, 31] and used as a stencil
for co-culturing [52] and that MG can be used for creating realistic cell morphology [53]
in vitro, whereas this study evaluates the behavior of MG on Parylene C for the purpose
of implanting the coated Parylene surface in vivo. The immunosuppressant DEX and a
combination of neurotrophic factors were also loaded into MG to modulate the immune
response and improve neuronal integration into the probe. Although 1 month in vivo
electrophysiological recordings showed that probes coated in MG supplemented with
bioactive molecules recorded an increased number of spike-like events, in the 3-month
study, probes coated with MG that included bioactive molecules did not perform any
better than probes coated with only MG. However, probes coated in either MG or MG
supplemented with bioactive molecules had a higher SNR than non-coated probes.
Elution studies showed that the bioactive molecules are all released from MG within a
day and therefore would have little effect during a chronic implantation. Other groups
that have successfully shown decreases in astrocytic growth around probes using other
coatings that released DEX over the course of days [54], or weeks [55]. In order to obtain
the benefits of the bioactive molecules while maintaining the improved performance of
MG, the bioactive molecules could be encapsulated in slow-releasing, biodegradable
microparticles that could be loaded into the MG. The microparticles could be small
enough to be released from the MG, but large enough remain in the insertion site, similar
to other microparticles loaded into hydrogels [16].

74

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80

4.1 Thin film polymeric devices
Implanted electronic devices are used to sense electronic signals such heart rate
and neural activity, and deliver therapeutic electric current such as in pacemakers, deep
brain stimulation, cochlear implants, and spinal cord stimulation. Because these devices
target soft, moving tissues, a degree of mechanical flexibility is required to maintain good
contact (important for sensing weak electrical signals and reducing the amount of current
required for a therapeutic effect) and avoid damaging the tissue through tethering effects.
Current flexible devices have large form factor and are made from bulk materials such as
insulated wires for pacemaker leads or laser cut metal foils embedded in silicone for
cochlear implants and spinal cord stimulation. Thus, the resolution and density of the
features and mechanical flexibility of sensors that can be integrated onto these devices
are limited.
Thin film polymeric devices use photolithography to pattern electrodes or other
sensors or structures onto the surfaces of materials such as Parylene [1-3], polyimide [4-
6], or polydimethylsiloxane [7, 8]. Based on microfabrication technology used to make
transistors or micro-sensors, thin film polymeric devices have micron scale resolution or
better and allow sensors such as transistors [9] or force sensors [10] to be integrated
directly onto the flexible device. These devices are thin enough to be rolled up and
delivered via catheter [11] or incorporated into the catheter itself [12]. Thin film
polymeric devices for cochlear stimulation [13] spinal cord stimulation [14], as well as
electrocardiography (ECG) [6, 15], electrocorticography (ECOG) [16], and intracortical
neural recordings [17] have been demonstrated.
CHAPTER 4
MECHANICAL PROPERTIES OF PARYLENE-METAL-PARYLENE THIN
FILM DEVICES
81

While thin film polymeric devices are flexible, biocompatible, and allow for high
resolution features, they are prone to delamination [18-20] and there are concerns about
their mechanical robustness [18]. Delamination is a common concern for thin film
polymeric devices, especially in aqueous solutions [19]. Any interface in a thin film
polymer device that delaminates could potentially create paths to water intrusion and so
determining the adhesion strength between layers is important. Thin films can bond to
underlying substrates physically by filling in pores or surface topography, through
covalent bonds, or by van der Waals forces. In addition, interfaces that typically have
poor adhesion, such as Parylene deposited on silicon can be improved with adhesion
promoters, such as the silane A-174, which bonds covalently both with Parylene free
radicals during the deposition process and provide a strong linkage with the underlying
substrate [20]. Also, a combination of heat and compression can bond Parylene layers
together by causing the polymer chains on the surface of the touching films to become
entangled [21]. Computational models suggest that metals such as platinum and titanium
can form stable, covalent bonds when deposited on Parylene [22]. Water intrusion (both
from the edges of thin film devices and through the polymer layers themselves) can
weaken adhesion by penetrating between physical bonds or dissolving intermediate layers.
Delamination occurs when upward forces perpendicular to the substrate are stronger than
adhesion forces the top film.
Several methods have been used to measure the adhesion force of thin films. The
“scotch test tape” qualitatively tests the adhesion of a film deposited on a substrate by
applying a piece of pressure sensitive tape to the film and pulling the tape off [23, 24]. If
the top layer of the deposited film comes off, it is said to have “failed” the scotch test
tape. The adhesion force of the film can also be determined using a load cell connected to
a free end of a film to measure force required to separate the film from a substrate [25].
Thin film polymeric devices can also fail under the application of large and
persistent mechanical forces due to muscle contractions or acceleration. When a material
is subjected to bending, it experiences compression forces on the inside surface and
82

tensile forces on the outside surface Figure 4-1. These stresses follow a linear relationship
according to the following equation for bending stress (σ
b
):
𝜎 𝑏 =
𝑀𝑦
𝐼

(4-1)

where M is the bending moment, y is the vertical distance from the neutral axis,
and I is the moment of inertia around the neutral axis. The neutral axis within the sample
is where the transition from compressive to tensile forces occurs (Figure 4-1(b)). Because
materials generally fail due to cracks propagating perpendicular to tension forces, cracks
in bent materials tend to propagate from the outside surface to the inside surface.

Figure 4-1. (a) A cable being bent between two plates with a force “F” creates regions of
(b) tensile and compressive stresses, which are defined by the bending stress (σ
b
). The
bending stress is linearly proportional to the distance from the neutral axis (b), which is a
plane in the middle of the cable where the compressive and tensile forces cancel out.

The effect of bending stress on thin film devices can be assessed by wrapping
films around cylinders of specific diameters ranging from hundreds of microns to several
millimeters and then either measuring the resistance across the traces or the testing the
83

performance [26-30]. Alternatively, films can be wrapped around spheres [14, 26] and
inflatable balloons [31]. Some devices are self-curling and were shown to maintain
function while being bent 90° [32] or down to a radius of 2 mm or less [2, 33]. A
convenient and automated setup involves bending between two plates [34], a versatile
technique that was developed for optical fibers [35], but has also been used to determine
the smallest diameter that thin conductive films on flexible substrates can be bent before
registering an open circuit [36, 37]. There are several advantages to using this type of
bend test rather than a traditional tensile test, including eliminating the influence of the
grips on the test and simple setup; however, the test is limited by difficulty in determining
the exact position of zero separation and only being able to test the portion of the sample
between the plates.
Traces can also break due to fatigue when they are repeatedly stressed at forces
less than their tensile strength. This can easily happen in the body where the heart
muscles contracts over 100k times per day and the neck moves an average of 10 million
times in each direction each year [38, 39]. Fatigue occurs because dislocations and voids
move and coalesce under bending stress. When a sufficient number of dislocations
aggregate, crack propagation sites form that weaken the material. Fatigue in thin film
devices has been measured in test in vitro by continuously measuring resistance across
polyimide [5] or Parylene [40] test cables bent to 90° angles for thousands or millions of
cycles, and in vivo by measuring ECG signals for ~10,000 heart beats [6].
Although mechanical testing of individual flexible thin-film devices has been
reported, there has been little work to done in applying multiple tests to the same types of
devices manufacturing methods to obtain a more complete understanding of the potential
and limitations of devices made using a specific structure. Also, many of these tests (such
as bend testing across cylinders of multiple diameters) were discrete and not done to
failure. Destructive and continuous peel/bend tests can provide specific limits to where
devices fail. The present work seeks to combine multiple mechanical testing methods to
one type of device manufacturing method (Parylene/metal/Parylene) to determine the
84

potential and limitations of current Parylene/metal/Parylene devices, find implications for
improved future designs, and present a general set of tests that can be done to screen and
compare different manufacturing methods of thin film devices.
This chapter presents methods that were developed to measure the peel test
strength, the minimum bending diameter, and the effects of bending fatigue on the
various interfaces of Parylene/metal/Parylene devices. In addition, Al
2
O
3
was deposited
on either both sides or one side of the metal traces in the Parylene/metal/Parylene device
using atomic layer deposition (ALD) to determine how this material, used for its water
barrier properties [41], affects the mechanical properties of Parylene/metal/Parylene
devices.
4.2 Al
2
O
3
as a barrier material
Al
2
O
3
has been shown to be an excellent barrier to water vapor [41] and for this
reason was deposited with Parylene to improve the performance of implanted neural
electrodes [42]. Parylene serves as a barrier for bulk water, ions and chemicals, while the
Al
2
O
3
serves as a barrier for water molecules that are able to penetrate the Parylene.
Plasma enhanced ALD is a process that sequentially deposits monolayers of molecules
onto a substrate using self-limiting chemical reactions. ALD deposited Al
2
O
3
is preferred
over sputtered or plasma enhanced chemical vapor deposition (PECVD) because of argon
sputtered Al
2
O
3
can trap argon into the coating making it porous, and PECVD can lead to
less dense films [41]. ALD coatings are also highly conformal and pinhole free, thus
avoiding gaps which can be major sources of corrosion in encapsulated electronics in
contact with water [43].
In the present study, ALD was used to incorporate Al
2
O
3
in PMP devices to create
a variety of interfaces summarized in Table 4-1.

85

Table 4-1. Material combinations investigated for flexible Parylene devices.
Abbreviation Interfacial layers (from bottom to top)
PP Parylene/ Parylene
PMP Parylene/platinum/Parylene
PAAP Parylene/Al
2
O
3
/Parylene
PAMAP Parylene/Al
2
O
3
/titanium/platinum/Al
2
O
3
/Parylene
PAAP Parylene /Al
2
O
3
/Parylene
PMAP Parylene /titanium/platinum/Al
2
O
3
/Parylene
Thickness of layers: Parylene C (top and bottom layers) 12-13 µm; Al
2
O
3
(top and bottom layers) 17-25 nm,
Ti 20nm; Pt 200 nm.

4.2.1 Processing thin film devices with Al
2
O
3

To incorporate Al
2
O
3
into PMP devices, ALD was used to deposit 17-24 nm of
(250 cycles on a Fiji 2 System (Ultratech, MA, USA)) either before and after metal
deposition (for PAMAP devices (Figure 4-2(b)) or only after metal deposition (for PMAP
devices). Al
2
O
3
is inert and can be etched by water [44] and KOH-based developers. As
such, modifications to standard Parylene surface micromachining processes were
required.

86

Figure 4-2. Schematic of processing steps for (a) PMP devices and (b) PAMAP devices.
(1) Parylene C is deposited over a silicon wafer. (2) A standard liftoff procedure is used
to pattern e-beam deposited platinum. (3) A second layer of Parylene C is deposited over
the patterned platinum. (4) Contact pads are exposed and devices are cut out using O
2

plasma etch. (5) Devices are released from silicon wafer. For PAMAP devices a
conformal barrier layer of Al
2
O
3
was deposited before step ii and before step iii. For
PMAP devices, Al
2
O
3
was only deposited before step iii.

4.2.1.1 Metal deposition onto Al
2
O
3

For PMP devices, a liftoff process is used to pattern platinum metal onto Parylene
substrates, which consists of using a negative photoresist layer to create a negative image
of the desired pattern, using e-beam evaporation to deposit platinum over the entire wafer,
and then stripping the photoresist by submersing the wafer in acetone to reveal the final
metal pattern. When platinum was deposited onto surfaces with exposed Parylene and
Al
2
O
3
, the platinum peeled off the regions of exposed Al
2
O
3
before the wafer could be
placed in acetone for liftoff. After stripping the platinum coated photoresist, the wafers
were re-patterned with fresh photoresist (with a near blind attempt to align the new
patterns with the previous pattern) and platinum was again deposited on the wafer using a
20 nm titanium adhesion layer.
87

In this process, the metal successfully adhered to the Al
2
O
3
regions. Because there
were no clear alignment marks after the metal had been removed during the first
deposition step, the patterns in the second deposition step were shifted resulting in partial
alignment marks (Figure 4-3(b)).

Figure 4-3. (a) Pattern of intended alignment marks. (b) Alignment marks resulting after
depositing metal on Al
2
O
3
for a second time (scale bar is 150 µm). (c) Overlap between
two alignment marks produce image in (b).

Two possible reasons why metal would adhere to Parylene coated with Al
2
O
3
on
the second time metal is evaporated onto the wafer, but not the first time are: one, the
photolithographic patterning, metal evaporation, and metal removal process modified the
surface of the Al
2
O
3
to make it able to accept metal ions during the second evaporation
attempt, and two, performing the photolithography step twice (particularly the
development step in which the wafers are submersed in a dilute NaOH solution)
completely removed the Al
2
O
3
, thus exposing the underlying Parylene, which is known to
bond to evaporated titanium and platinum. NaOH has been used to etch Al
2
O
3
[45] in
making nanowires, but higher concentration of NaOH was used (0.4 M) and the wires
were made from anodic alumina and not ALD deposited Al
2
O
3
. To determine how
quickly ALD Al
2
O
3
etches in dilute NaOH developer (~0.1 M), silicon wafer shards
coated with 20 nm of Al
2
O
3
were dipped into developer for different periods of time and
the thickness of the Al
2
O
3
after soaking was measured using a multiwavelength
ellipsometer (Gaertner, Skokie, IL, USA) (Figure 4-4).
88

Figure 4-4. Etch rate of Al
2
O
3
in dilute NaOH developer (~0.1 M) was determined by
breaking a wafer coated in Al
2
O
3
into (a) separate wafer shards. The wafer shards were
then (b) placed in the dilute NaOH developer and (c) removed at different soak times.
The thickness of Al
2
O
3
on the shards removed from the developer were then (d)
measured using ellipsometry.

The average etch rate of the Al
2
O
3
in the developer solution was found to be 3.3-
3.8 Å/s (Figure 4-5), with the etch rate being higher at the beginning of the etch and
decreasing over time. At this rate, a 30 second development done twice would remove 24
nm and remove the majority of the 25-29 nm thick layer of Al
2
O
3
on top of the Parylene.
If the bottom layer of Al
2
O
3
was removed, the devices with two layers of Al
2
O
3
(PAMAP
devices) should behave similarly to wafers with one layer of Al
2
O
3
(PMAP devices).

89

Figure 4-5. Thickness of Al
2
O
3
after different etch times in 1:4, DI H
2
O:AZ 340
developer as determined by ellipsometry using 3 different wavelengths. (5 nm is
approaching resolution of ellipsometer.) Etch rate is determined to be 3.3 to 3.8 Å/s.

4.2.1.2 Parylene adhesion to Al
2
O
3

A top layer of Al
2
O
3
was applied to the wafers after metal and sacrificial
photoresist structures had been patterned and deposited on the wafers (the second Al
2
O
3

layer for PAMAP devices and the first layer for PMAP devices). Once this layer had been
applied, a top layer of Parylene C was deposited onto the Al
2
O
3
. Photoresist was spun
onto this layer of Parylene and photolithography was used to expose Parylene where
contact pads would be etched out. However, when the wafers were put under vacuum
during reactive ion etching using O
2
plasma, the top layer of Parylene delaminated from
the Al
2
O
3
and was completely removed from the wafer.
To improve adhesion of Parylene onto Al
2
O
3
, the wafer was first treated with
silane A-174 (Momentive Performance Materials, Columbus, OH, USA) prior to
Parylene deposition, which has been shown to improve the adhesion of Parylene to
oxides [46]. 3 mL of A-174 was placed in a small reservoir and placed, uncovered, in a
dessicator with the wafers. The silane was allowed to evaporate for 2.5 hours under
vacuum, after which, the wafers were removed and immediately placed in the Parylene
90

deposition chamber. After silane treatment, the top Parylene layer was patterned and
etched using O
2
plasma without delamination.
4.2.1.3 Etching Al
2
O
3

The last steps of making the devices were to etch out contact pads and release the
devices. O
2
plasma was used to etch through the Parylene down to the contact pads and
underlying layer of Al
2
O
3
, which acted as an etch stop.
To etch through the bottom layer of Parylene, the Al
2
O
3
was removed by placing
in dilute AZ 340 developer (di H
2
O:developer, 4:1). Although an etch rate of 3.8 Å/s was
determined earlier (Figure 4-5), the underlying Parylene was not affected by the plasma
until after soaking for the wafers for 6 minutes. The longer time required perhaps being
due to a decrease in etch rate over time (Figure 4-5). (An attempt was made to determine
the presence of Al
2
O
3
by measuring the resistance across the traces, but it was found that
multimeter tips easily broke through the thin Al
2
O
3
layer and made contact with the
metal.).
Once the devices were released, if electrical contact could not be made by the
zero insertion force (ZIF) connector breaking through the Al
2
O
3
and making contact with
the metal, they were soaked in a 20% KOH solution for 2-4 minutes at room temperature
to remove Al
2
O
3
from the contact pads.

4.3 Peel strength
T-peel tests were performed on Parylene test samples to measure the adhesion
force between the different material interfaces combinations found in the
Parylene/metal/Parylene devices, including: PP, PMP, PMAP, PAP, PAMAP and PAAP
(Table 4-1). T-Peel tests performed were based on ASTM standard D1876_08 [47], but
modified to be used on thin film devices. T-peel test samples consist of two flexible
rectangular substrates bonded together at one end. The T-peel test is performed by
91

gripping each substrate at the non-bonded regions and pulling apart at 180°, a “T” shape
results (Figure 4-6(5), Figure 4-7(b)). The peel strength is defined as the average load
per unit width of bondline required to progressively separate the two materials. T-peel
tests are used for their simplicity of sample design and testing to compare different forms
of bonding or adhesives [48]; however, these tests are sensitive to the thickness of the
substrates which require energy to deform during the test and must be accounted for if
direct measurement of cohesion strength is desired [49].
4.3.1 Fabrication of peel test devices
Peel test samples consisted of a bonded interface with a clearly defined area to be
tested connected to two flaps that can be clamped to a load cell and pulled to separate the
joined interface. The flaps were created by patterning sacrificial photoresist (PR) between
two ~13 µm layers of Parylene (Specialty Coating Systems, Indianapolis, IN) to enable
them to be separated after the device was released (Figure 4-6).
4.3.2 Peel testing and characterization
The devices were tested by attaching one flap to a stationary clamp and the other
to a 50 g load cell (Omega, Stamford, CT, USA), which was attached to a motorized
stage (Z812, Thorlabs, Newton, NJ). A custom Labview program was used to control the
stage and measure the voltage output from the load cell (Figure 4-7(b)). PMP devices
were tested both dry and after a 24 hour soak in 1 × PBS solution at 37 °C for 24 hours.
Scanning electron microscope images were taken of PMP peel test samples after they had
been sputter coated in Pt.

92

Figure 4-6. Simplified schematic of peel test sample fabrication. (1) Sacrificial
photoresist (PR) layer and bonded interface are patterned next to each other on Parylene
substrate. (2) Sacrificial PR and interface are coated with a second layer of Parylene. (3)
O
2
plasma is used to etch through both layers of Parylene on either side of device. (4)
Device is released and soaked in acetone to dissolve PR. (5) Flaps are peeled back and
attached to a load cell to perform peel test. (Note: For samples with Al
2
O
3
, Al
2
O
3
was
deposited either before and after the sacrificial layer patterned in (1), or only after the
sacrificial layer was patterned in (1).)

4.3.3 Results
Peel test samples of identical geometries, but different material interfaces (Table
4-1) were fabricated. Test samples included Parylene flaps that could be gripped and
pulled apart (Figure 4-6, Figure 4-7).

93

Figure 4-7. (a) Flaps of peel test sample separated with tweezers after removing
sacrificial PR. Bonded metal region (3 × 4.5 mm) is circumscribed by dotted line. (b)
Peel test sample placed over mounting posts on motorized stage with load cell before top
clamp is secured. White arrow in (b) points to bonded region of peel test device. (Scale
bars for (a) and (b) are 5 mm.)

Peel testing resulted in characteristic T-peel curves with an increase in force once
the samples become taut, and a relatively continuous force as the bonded area is
separated by peeling (Figure 4-8). Overall, the average peel test strength of the PP
devices was low with spikes observed in the force plots at the beginning and end of the
peel tests (representative plot in Figure 4-8(a)).The spike at the beginning of the peel test
could be due to crack initiation forces being higher than crack propagation forces, as well
as residual photoresist from the sacrificial layer not being completely removed. The spike
at end of the test could have been caused by a reorientation of the samples as the device is
separating so that the pulling force is acting on more of the bonded surface area than
bondline being pulled during the steady state phase of the test. In contrast, the PMP
interfaces were much stronger than the PP interfaces and dominated any additional
bonding that may have taken place at the edges (Figure 4-8(b)).

94

Figure 4-8. Representative peel test raw data from (a) PP and (b) PMP test devices. The
edge effect observed for the PP device was not observed for the PMP device, most likely
due to the force magnitude being dominated by the larger PMP average peel strength,
(~50 times larger). Note the two plots are on different scales.

The raw force data provided in Figure 4-9(a & b) indicated PMP devices
exhibited the highest adhesion, followed by the PMAP, PAP, PAMAP, and PAAP
devices (Figure 4-9 (c)). The force required to pull apart PP and PP wet devices was the
lowest (Figure 4-9(b)). In all cases, the top layer of Parylene was separated from the
underlying layers.
Mean values of average peel strength are presented in Figure 4-9(c & d).
Significant differences were found between PAP and PMAP and the PMAP and the
PAMAP devices. However, there is no significant difference between the PAMAP and
PAAP devices, likely due to defects introduced because of the difficulty of processing the
device with 2 layers of Al
2
O
3
which caused the results of the peel test to be inconsistent
and have a larger standard deviation than the other devices. The cumulative mean average
peel test bar graphs also show that samples with either 1 or 2 layers of Al
2
O
3
showed
much smaller variation in peel strength than samples without Al
2
O
3
(Figure 4-9(c)). This
is likely due to the uniform Al
2
O
3
layer that coated the entire surface before the second
layer of Parylene was deposited.
PP samples soaked overnight at 37 °C in 1 M phosphate buffered saline solution
had significantly lower peel strength than devices that were no exposed to soaking
95

(Figure 4-9(d)), likely due to water intrusion between the layers which resulted in
decreased adhesion strength.

Figure 4-9. Peel test data for all conditions. (a) Combined raw data (mean ± SE, n = 3-8)
from various trials, and (c) mean values of average peel strength from various interfaces
(mean ± SE, n = 3-8). Parylene/Parylene interfaces had the lowest peel strength and were
regraphed separately (in b & d) on a different scale. The slash symbol (used in c & d)
indicates the interface at which delamination occurred.

SEM images were acquired for all conditions (Figure 4-10). The PMP surfaces
(Figure 4-10(a)) contained cracked and delaminated Pt on the bottom layer of Parylene
(left image) and plastically deformed Parylene on the top layer (right image). In contrast,
the PP devices (Figure 4-10(b)) were relatively smooth and clean except for some tiny
strands of plastically deformed Parylene that appeared on the bottom layer (left image).
The devices with Al
2
O
3
with or without a layer of Pt were similar (Figure 4-10(c-f)). The
96

bottom layer of these images (left images), which included the bottom layer of Parylene
and one or two layers of Al
2
O
3
, contained parallel ridges topped with plastically
deformed flaps of Parylene. The top layers (Figure 4-10(c-f), right images), contained
clusters of very thin strings of plastically deformed Parylene.

Figure 4-10. SEM images of all peel test surfaces. For each individual panel, the left
image is the bottom (closest to the silicon wafer during processing) layers of Parylene,
metal, and Al
2
O
3
and the right image is the top layer of Parylene. Note that images are at
different scales. (Surfaces were sputtered coated with Pt before imaging.)

97

4.4 Minimum bending diameter
4.4.1 Measuring minimum bending diameter before failure
Often, thin film devices are folded around biological tissues [30] or catheters [28]
in their given application and therefore, it is important to quantify the bending diameter,
or radius, at which failure occurs to define the usable bending range. To determine the
minimum bending diameter, the electrical resistance of a Pt trace across cable was
monitored as a loop of cable was bent between two plates. Bend testing with the
specimen being crushed between two plates was developed for optical fibers [35], but
have also been used to test the minimum diameter that thin conductive films on flexible
substrates can be bent [36, 37]. There are several advantages to using this type of bend
test rather than a traditional tensile test, including: eliminating the influence of the grips
on the test and simple setup; however, the test is limited by difficulty in determining the
exact position of zero separation and only being able to test the part of the sample
between the plates. Also, if the strength of the material is being tested, the strength of the
device being tested is influenced by the length of the material between the plates and
needs to be considered.
To test the minimum bending diameter of Parylene/metal/Parylene devices,
rectangular cables with single serpentine trace along the length of the device were
fabricated using standard photolithographic techniques (Figure 4-11). Briefly, a 13 µm
layer of Parylene C (Specialty Coating Systems, Indianapolis, IN) was deposited on a
bare silicon wafer. A lift-off process was then used to pattern of platinum metal (200 nm)
deposited onto the Parylene layer using e-beam evaporation. Another 13 µm layer of
Parylene was applied on top of the platinum and the ends of the traces were exposed to
form contact pads using O
2
plasma. Finally, a protective layer of photoresist was spun on
to cover the contact pads and O
2
plasma was once again used to cut out the devices
before they were released. For devices with Al
2
O
3
, ALD was used to deposit a ~20 nm
layer of Al
2
O
3
either before and after the platinum deposition and liftoff steps (for the
98

PAMAP devices) or only after the liftoff step (for the PMAP devices) (Figure 4-2).
Details of Al
2
O
3
fabrication included in section 4.2.1.

Figure 4-11. (a) Cartoon and (b) photograph of minimum bend diameter test devices
(scale bar is 5 mm). In cartoon (a), the thickness of the traces have been increased and
number of turns decreased for illustrative purposes. The metal and hole on the left side of
device are for clamping the device to the test fixture. The contact pads on the right side
allow electrical resistance measurements.

To measure the minimum bend diameter, a setup was designed to controllably
compress a folded device between two acrylic blocks while monitoring the resistance
(Figure 4-12(a)). One of the acrylic blocks was attached to a 50 g load cell (Omega,
Stamford, CT, USA) (Figure 4-12 (a & b)), which, in turn, was connected to a motorized
stage (Z812, Thorlabs, Newton, NJ) for measuring axial force for the purpose of
determining complete contact between the blocks. The devices were supported by a
holder (Figure 4-12(c)) and connected to a precision multimeter (Keithley 2700, Keithley,
Cleveland, OH, USA). The cable bend was positioned between two acrylic blocks
(Figure 4-12(b)). Using a custom Labview program (National Instruments, Austin, TX),
the motorized stage was advanced, bringing the two acrylic blocks together and reducing
the bend diameter. Failure was defined as an open circuit measurement. The minimum
diameter was measured with the devices folded towards (‘in’) and away from (‘out) the
contact pads (Figure 4-13).
99

Figure 4-12. Pictures of crush test setup. (a) Overview of setup with acrylic block
connected to the load cell attached to a motorized stage. (The crush test device is
obscured behind acrylic block.) (b) Close-up of folded end of crush test device between
two acrylic blocks with one acrylic block attached to a load cell (scale bar is 5 mm). (c)
Close-up of crush test device folded around holder (scale bar is 5mm).

100

Figure 4-13. Schematic of bending failure test setup with (a) contacts pads out and (b)
contact pads in. Setup is the same except that in (b) the device is turned inside out and the
ZIF connector is turned upside-down to accommodate downward facing contact pads.

A diameter of 0 mm was defined as the point in which the force measurement on
the acrylic clamping blocks was 4.5 N and all data was normalized to this setting. Full
contact of acrylic blocks at 4.5 N was confirmed using video images. The force data
along with the resistance of the traces as the bending diameter was reduced was graphed
(Figure 4-14) and the minimum bending diameter just prior to the formation of an open
circuit was recorded (Figure 4-15).

4.4.2 Results
Raw force data from the crush tests shows little increase in force until the acrylic
blocks are within 0.2 mm of each other at which time the force begins to increase rapidly
(Figure 4-14(a, c, e)) indicating that the blocks are beginning to make contact with each
other. Video images of the devices, however, reveal that complete contact between the
blocks at the position of the crush test device does not take place until the sensor reads
4.5 N, which is taken to be the zero diameter for the crush tests (Figure 4-14(b, d, f)).
101

Figure 4-14. Raw force (a, c, and e) and resistance (b, d, and f) data from minimum
bending diameter tests. Force data (a, c, and e) were used to determine where the acrylic
blocks made full contact (i.e. where Force = 4.5 N). Minimum bending diameter was
defined as the resistance measurement (in b, d, and f) closest to zero not measured as an
open circuit.

The minimum bending diameter of PMP devices was measured to be ~130 µm
with no significant difference between bending the device towards the top layer of
Parylene (in) or towards the bottom layer of Parylene (out). Contrarily, the PMAP device
with one layer of Al
2
O
3
underneath the top layer of Parylene failed much earlier than the
PMP devices when bent towards the Al
2
O
3
layer, but did not break at all when bent away
from the Al
2
O
3
layer. The low yield of PAMAP devices was attributed to difficulty in
102

establishing electrical contact, and therefore PAMAP devices were only tested in the “in”
condition (the devices were bent towards the contact pads).

Figure 4-15. Summary of crush test data. PMAP (in) had a significantly larger minimum
diameter than either PMP (in) or PMP (out). (Mean ± SE, n = 3-4)

The variation in minimum diameter between the PMP and the PMAP devices is
likely due to the adhesion strength between the two layers. Because materials generally
fail due to cracks forming and propagating due to tensile forces, cracks form on the
outside edge of a bend sample and then propagate towards the inner edge. Therefore, the
cables will fail by a crack forming in the outside layer of Parylene, which will propagate
inwards towards the trace. If there is poor adhesion between the ALD Al
2
O
3
then the
crack may have difficulty propagating between the two layers of Parylene. When a
PMAP device is oriented “out”, the trace will be better adhered to the inside layer of
Parylene. When the outside layer of Parylene cracks, the crack cannot propagate over the
interface. However when oriented “in”, the trace is bonded to the outside layer of
Parylene and cracks when the outside layer of Parylene cracks. When there is good
adhesion on both sides, such as with the PMP devices, the traces lie along a neutral plane
103

(Figure 4-1), which will reduce the forces on the traces and results in breaking at the
same point, whether being bent “in” or “out”.
SEM images were taken of both the top and bottom surfaces of devices that had
been bent away from (Figure 4-16) and toward (Figure 4-17) the contact pads. Cracks
that formed due to tension on the outside of the bend (Figure 4-16(a & c) Figure 4-17(a &
c)) appeared to be deeper with cleaner edges compared to cracks formed due to
compression on the inside of the bent device (Figure 4-16(b & d) Figure 4-17(b & d)),
which are characterized by rope-like ridges of Parylene accompanied by plastically
deformed shark-tooth shaped fringes.

Figure 4-16. SEM of cracks formed after bending (a & b) PMP devices and (c & d)
PMAP devices away from the contact pads (‘out’). The top layers of Parylene (a & c)
were the outside of the bent device, and the bottom layer (b & d) were from the inside of
the bent device. (Images on the right side of each panel are zoomed in views of images on
the left with scale bars of the left and right images being 50 and 10 µm, respectively.
Parylene sputtered coated with Pt before imaging.)

104

Figure 4-17. SEM of cracks formed afer bending (a & b) PMP devices and (c & d)
PMAP devices towards the contact pads (‘in’). The bottom layers of Parylene (a & c)
were the outside of the bent device, and the top layer (b & d) were from the inside of the
bent device. (Images on the right side of each panel are zoomed in views of images on the
left with scale bars of the left and right images being 50 and 10 µm, respectively.
Parylene sputtered coated with Pt before imaging.) In the PMAP sample, delamination at
the bend was observed.

4.5 Bending fatigue
4.5.1 Fabrication of bend test devices
Fatigue occurs when a material cracks due to repeated exposure to stresses less
than their yield stress and occurs due to the accumulation of dislocations during strain,
which causes microscopic cracks to form around areas of stress concentration such as
surfaces, grain boundaries, and inclusions. The microscopic cracks will grow until the
material is weak enough to fail. Bulk materials are tested for fatigue by exposing them to
cyclic uniaxial stresses until failure [50], which is also a method used in thin films [51].
Fatigue in cables (both traditional cables such as FFC’s [52] and pace maker leads [53],
as well as thin film cables [5]) is studied by measuring the resistance across cables as
they are cyclically bent to a given angle until an open circuit is measured. In this work, a
105

method for measuring fatigue in telecommunication cables called a “flex test” [54] was
modified to test thin film cables. In this test, a specimen is connected to a rotating stage
and suspended between 2 mandrels on either side of the cable. A weight is connected to
the bottom of the specimen and then the specimen is bent ± 90° for a given number of
cycles (Figure 4-19). The sample is then inspected for evidence of fracture. In addition to
microscopic inspection, the insulation was inspected using electrochemical impedance
spectroscopy, which is a method used to measure the performance of implanted
electrodes [55].
To measure the effect of repeated deformation on thin film Parylene devices, bend
test samples with fully insulated traces were fabricated (Figure 4-18). Similar to the
crush test devices, a 13 µm layer of Parylene C (Specialty Coating Systems, Indianapolis,
IN) was deposited on a bare silicon wafer. A liftoff process was then used to pattern
platinum metal (200 nm) traces deposited on top of the Parylene substrate using e-beam
evaporation. The traces were insulated with another 13 µm layer of Parylene applied on
top of the platinum. Finally, O
2
plasma was used to expose the ends of the traces to form
contact pad, and cut out the devices. For devices with Al
2
O
3
, ALD was used to deposit a
~25 nm layer of Al
2
O
3
either before and after the platinum deposition and liftoff steps
(for the PAMAP devices, Figure 4-2(b)) or only after the liftoff step (for the PMAP
devices). Details of Al
2
O
3
fabrication included in section 4.2.1.

106

Figure 4-18. (a) Cartoon and (b) photograph of bend test devices (scale bar is 5 mm). (In
cartoon (a), the thickness of the traces have been increased and number of traces
decreased for clarity.) The metal and hole on the left side of device are for gripping the
device. The traces are exposed on the right side to make contact pads to connect the
device to a potentiostat.

4.5.2 Testing
A custom cable flex tester (based on ASTM standard D4565-10 for testing cable
insulation) was constructed consisting of a Sparkfun stepper motor (ROB-09238)
(Sparkfun, Boulder, CO, USA), an Arduino UNO (Arduino, Torino, Italy) controller, and
mounting fixtures (Figure 4-19). Devices were subjected to multiple bending cycles in
which one cycle one cycle is defined by bending between -90° to +90° around a mandrel
(1.25 mm diameter).

107

Figure 4-19. (a) Schematic and (b) picture of bend testing setup. Design was based on
the cable flex test as described by the ASTM standard D4565-10 for evaluating insulation
on wires.

Electrochemical impedance spectroscopy (EIS) was used identify signs of failure
(e.g. delamination, water infiltration) by measuring the impedance between traces (the
lateral impedance (Figure 4-20(a)) and the impedance between a trace and an external
counter electrode (the transverse impedance (Figure 4-20(b)) while immersed in
phosphate buffered saline (PBS).
108

Figure 4-20. Electrochemical impedance spectroscopy (EIS) was performed on bend
testing devices between (a) adjacent traces (lateral impedance) and (b) from traces on the
device to a platinum counter electrode off the device (transverse impedance).

EIS was performed in 1× PBS as 37 °C with amplitude of 10 mV
rms
and a
frequency range of 1-100,000 Hz. A platinum wire was used as a counter electrode and
an Ag/AgCl (3M NaCl) electrode was used as a reference (transverse impedance). The
impedance magnitude was normalized to absolute impedance of the magnitude at zero
bends and plotted.
The impedance was expected to behave according to the simplified Randall’s
model, with the values for the capacitive and resistive components changing as the
insulation fails (Figure 4-21, [56]). Delamination of the Parylene film is expected to
cause a drop in measured impedance starting at lower frequencies, which has been shown
to be more sensitive to electrode failure [57].

109

Figure 4-21. Variations of a simplified Randle’s circuit modeling (a) a fully insulated
electrode, (b) a deinsulated electrode, and (c) a failing electrode, where C
dl
= double layer
capacitance, R
s
= solution resistance, R
ct
= charge transfer resistance, C
c
= coating
capacitance, and R
p
= pore resistance. Note that the model is the same for the insulated
and deinsulated model, but the capacitive and resistive elements will be different. Figure
adapted from [56].

The magnitude of the impedance was normalized to magnitude at 100,000 Hz,
where the resistive component associated with the trace will have the least effect on the
magnitude. Delamination of the Parylene film is expected to cause a drop in measured
impedance starting at lower frequencies, which has been shown to be more sensitive to
electrode failure [57].
In addition to EIS, devices were examined using optical and Nomarski differential
interference coherence (DIC) microscopy, as well as scanning electron microscopy
(SEM), before and after bending.
4.5.3 Results
EIS of a PMP cable released from a wafer that showed inconsistent adhesion
during peel testing (data not shown) demonstrated the ability to reveal insulation damage
without desctructive testing (Figure 4-22). The magnitude of the transverse impedance
increased at lower frequencies after the smallest number of bend cycles (100) and then
experienced a broad spectrum decrease after 5k bend cycles (Figure 4-22(a)). Similarly,
the phase decreases at lower frequencies during the first 4k bend cycles, and then at
higher frequencies after 5k bend cycles (Figure 4-22(b)). The lateral impedance proved to
be more sensitive, showing a broad spectrum decrease in the magnitude of the impedance
at 4k as well as 5k (Figure 4-22(c)). Before bend testing (0 cycles), the lateral impedance
had a -90° phase at all frequencies, which is indicative of a fully insulated trace. However,
110

the phase dropped at lower frequencies as soon as bend testing began, and there was a
broad spectrum drop after 4k cycles (Figure 4-22(d)). Broad spectrum decreases in
magnitude and phase suggest that the insulation has been damaged and there is a path for
ions to move between traces.

Figure 4-22. Representative transverse (a-c) and lateral (d-f) impedance from a PMP
device released from a wafer that peel testing revealed had uneven adhesion of the top
layer of Parylene. The device shows clear change in impedance over 5k bend cycles
represented by downward shifts in magnitude (a & d), which are emphasized by
normalizing the data to the impedance values at 0 bends (b & e), and shifts in phase (c &
f) at high and low frequencies (mean ± SE, n = 8).

The average transverse impedance from PMP devices (which peel testing showed
good adhesion between the Parylene and metal (Figure 4-9)) showed a small drop in
magnitude at lower frequencies, but remained relatively constant at higher frequencies
(Figure 4-23(a)). However, as the number of bend cycles increased, the impedance at
lower frequencies became lower. The phase of the PMP devices was initially -90°, but
increased at lower frequenices after the first set of bends. At higher frequenices, the phase
111

remains capacitive (around -90°) even when the devices are bent through 100k cycles
(Figure 4-23(b)). In contrast to the PMP devices, the PMAP and PAMAP exhibited larger
fluctuations in magnitude over most frequencies (Figure 4-23(c & e)) and compromised
insulation prior to bending based on the phase data (Figure 4-23(d & f)).

Figure 4-23. Transverse impedance from (a & b) PMP, (c & d) PMAP and (e & f)
PAMAP Parylene cables (mean ± SE, n = 16). Magnitude is normalized with respect to
the impedance at zero cycles. (PMAP and PAMAP devices were only bent for 10k
cycles.)
112

The phase of the PMP devices’ lateral impedance started greater than -90° at
lower frequencies (Figure 4-24(b)). However, similar to the transverse impedance, there
was no broad spectrum shift of magnitude, but only an increase at lower frequencies
(Figure 4-24(a)). Similar to the transverse impedance, the lateral impedance of the
devices with Al
2
O
3
had large shifts in magnitude with large standard errors (Figure 4-24
(c & e)). In addition, the initial phase was greater than -90 ° even before the devices were
bent (Figure 4-24 (d & f)).
The lateral impedance of the PMP devices revealed that even before the devices
were bent, the traces were not fully capacitive at lower frequencies, indicating that there
were ionic paths between traces that could not be detected when measuring transverse
impedance (Figure 4-24(b)). However, similar to the transverse impedance, the absence
of a broad spectrum decrease in magnitude and the presence of an enduring capacitive
phase at higher frequencies imply that at least part of the Parylene insulation remained
intact even up to 100k bend cycles.
Also similar to the transverse impedance, the lateral impedance of the devices
with Al
2
O
3
had low magnitudes and small phase shifts, which suggest open ion paths
between traces (Figure 4-24(c-f)).

113

Figure 4-24. Lateral impedance from (a & b) PMP, (c & d) PMAP, and (e & f) PAMAP
Parylene cables (mean ± SE, n = 16). Magnitude of impedance is normalized to the
impedance of the device at zero cycles. (PMAP and PAMAP were only bent for 10k
cycles.)

Cracks in the Parylene that started at edges (not shown) were visible by optical
microscopy after 2k bend cycles (Figure 4-25(a)) and became pronounced after 5k bend
cycles (Figure 4-25(b)). The cracks were either chevron-like, radiating from a central
point such (top left traces in Figure 4-25(b)) or perpendicular to the traces (right trace in
114

Figure 4-25(b)). Cracks were observed to cross over the metal traces and the Parylene
region between traces (Figure 4-26).

Figure 4-25. Nomarski differential interference contrast images of fatigue cracks in the
top Parylene layer covering three traces of a PAMAP devices after (a) 2k bend cycles and
(b) 5k bend cycles. Metal regions are wide, light gray strips and Parylene appears as thin,
dark gray strips. The majority of cracks appear after 5k bend cycles; however, the dashed
circle marks cracks present after 2k cycles (circle is 200 µm in diameter).

Figure 4-26. (a) Optical and (b) SEM images of cracked Parylene over and between
metal traces on cable bent for 100k cycles. The metal traces are on the left. (Parylene
sputtered coated with Pt before imaging in SEM.)

Optical micrographs of PMAP and PAMAP devices showed noticeable water
intrusion after 100 bend cycles (images were not taken after the initial pre-bend EIS and
before first bend testing cycles) and continued to be present after subsequent bend tests
(Figure 4-27). In Figure 4-27, there are clearly Newton rings in the Parylene next to the
115

edge of the device, but there is also evidence of Newton rings appearing near the bottom
of the right trace. Water intrusion could have been worse on Al
2
O
3
devices because Al
2
O
3
is less hydrophobic than Parylene or due to Al
2
O
3
etching in water [44] to provide paths
for water intrusion.

Figure 4-27. Micrograph of two traces near the edge (marked by dotted white line) of a
PAMAP device after 10k bend cycles (scale bar is 250 µm). Evidence of water intrusion
can clearly be seen as large Newton rings in the Parylene next to the edge and smaller
Newton rings on the right edge at the bottom of the right trace. Discoloration and pock
marks on the traces and may be evidence of Al
2
O
3
etching by water that has penetrated
between Parylene layers.

Evidence of Al
2
O
3
etching can be seen in pock marks on the traces (Figure 4-27)
as well as discoloration and branched patterns that start from the outside traces (Figure
4-28(a)). The impedance of the traces measured at 1 kHz show that before bend testing
begins, the impedance is much lower towards the edges of the device, but as the device is
subjected to further bending and EIS scans (each time an EIS was performed on all traces,
the devices were immersed in PBS for ~45 minutes), the impedance is reduced across all
traces and becomes more uniform (Figure 4-28(b)).
116

Figure 4-28. (a) Micrograph of all 8 traces of a PMAP device showing discoloration and
branched patterns occurring on outside traces (scale bar is 500 µm). (b) The transverse
impedance at 1 kHz for the traces shown in (a) versus their proximity to the edge. Before
bend testing, traces closest to the edge have the lowest impedance, but this difference
decreases as the device undergoes more bend cycles.

4.6 Discussion
4.6.1 Peel Testing
Peel tests revealed that PMP interfaces had the strongest adhesion measuring
nearly 20 times more than for PP interfaces. PP interfaces were further weakened by
soaking for 24 hours in 1 MPBS at 37 °C (Figure 4-9). The reduced adhesion in soaked
samples is likely due to water intrusion between the layers. The Al
2
O
3
-Parylene interface
was weaker than Pt-Parylene but stronger than PP. However, the adhesion between
Al
2
O
3
-Parylene interfaces for different processing conditions was inconsistent. The
PAMAP and PAAP devices had lower strength than the devices with only one layer of
117

Al
2
O
3
. This is most likely due to extra processing steps required for the extra layer of
Al
2
O
3
. These results imply that excellent adhesion is possible between platinum and
Parylene, which has been suggested through the results of modeling [22], and that
platinum could potentially be used as an adhesion promoter between Parylene layers. The
PMAP and PAMAP peel strength being higher than the PP peel strength also suggest that
Al
2
O
3
and A-174 could be used as an adhesion promoter between two layers of Parylene.
In all cases, the top layer of Parylene deposited on Parylene, metal, and Al
2
O
3
,
always separated from the underlying layer after peel testing. This indicates that the
interface between materials (metal and Al
2
O
3
) deposited onto Parylene possess greater
adhesion than the interface formed when Parylene is deposited onto the same materials.
Therefore, strategies to promote adhesion should focus on the substrate-Parylene
interface, such as the use of adhesion promoters [25], surface treatments [58], and
annealing [59].
4.6.2 Minimum Bend Diameter
The variation in minimum diameter between the PMP and the PMAP devices is
likely due to the adhesion strength between the two layers. Because materials generally
fail due to cracks initiating due to tensile forces, cracks will form on the outside edge of a
bent cable and then propagate towards the inner edge. However, if there is poor adhesion
between the Al
2
O
3
and top layer of Parylene, then the crack may have difficulty
propagating between the two layers of Parylene separated by Al
2
O
3
. Conversely, it was
observed that for PMAP devices oriented “in”, the metal trace is bonded to the outside
layer of Parylene and cracked when the outside layer of Parylene cracks. When there is
good adhesion on both sides, such as with the PMP devices, the traces lie along a neutral
plane, which will reduce the forces on the traces and results in breaking below the same
minimum diameter, whether being bent “in” or “out” (Figure 4-15).
These results suggest that if it is desired to make devices with very tight diameters
in one direction, it should be possible to deposit metal on only one side of a material, or
make one layer of insulation thinner than the other and then fold the device towards the
118

thinner layer of insulation. Otherwise, the results suggest that improving adhesion when
the metal traces lie along the neutral plane should improve bidirectional bending of
devices.
4.6.3 Bending fatigue
The average transverse impedance from a PMP devices (which peel testing
showed good adhesion between the Parylene and metal (Figure 4-9)) showed a small
decrease in magnitude at lower frequencies which are more sensitive to device failure
[60], but remained relatively constant at higher frequencies (Figure 4-23(a)). The -90°
phase of the PMP devices before bending indicate proper insulation of traces but that
insulation was in part compromised after the first set of bend tests as indicated by the
increase in phase at lower frequencies. Notably, however, the phase remained capacitive
at higher frequenices even when the devices are bent through 100k cycles (Figure
4-23(b)) indicating that the insulation was at least partially intact.
In contrast to the PMP devices, the PMAP and PAMAP devices showed large
variations in normalized magnitude, indicating large changes in the insulation. Increases
in in the magnitude (Figure 4-23(e)) could be due to cracks that broke the traces which
would increase the impedance. The decrease in impedance back to the baseline after 10k
bends suggested could be due to re-establishing contact between broken traces, or further
deteriation of the lamination between the two layers of Parylene. The increased phase (> -
90°) even at zero bends (Figure 4-23(e & f)) indicates that there was an almost immediate
intrusion of ions across the insulation to the traces in those devices. The large standard
error in the magnitude and phase of the PAMAP devices (Figure 4-23(c & e) show that
there was a wide variation in trace performance, which could be caused by flaws
introduced during processing Al
2
O
3
(addressed in the next section) or the effects of water
intrusion.
The phase of the lateral impedance of the PMP devices was greater than -90° at
lower frequencies before the devices were bent indicating that there were ionic paths
between traces that could not be detected when measuring transverse impedance (Figure
119

4-24(b)). However, similar to the transverse impedance, the absence of a broad spectrum
decrease in magnitude and the presence of an enduring capacitive phase (-90°) at higher
frequencies imply that at least part of the Parylene insulation remained intact up to 100k
bend cycles. Also similar to the transverse impedance, the phase of the lateral impedance
of Al
2
O
3
devices had small phase shifts, which suggested open ion paths between traces
(Figure 4-24(c-f)).
Fatigue cracks in the Parylene were visible in the cables after 2k bends (Figure
4-25(a)). The number and size of the cracks increased with more bends (Figure 4-25(b)).
The cracks appeared to originate in the Parylene layer and not the metal as the first
appearance was on the device edges (in which there was no metal) and the cracks
spanned across areas with and without metal (Figure 4-26(a & b)). In addition, while all
devices showed extensive cracking, PMP devices did not exhibit impedance changes.
This is in contrast to literature where cracking in Parylene due to, being flexed during
accelerated lifetime testing [61] and in long term in vivo testing [62] caused drops in
impedance.

4.6.3.1 Failure of Al
2
O
3
Devices
Optical micrographs of PMAP and PAMAP devices showed noticeable water
intrusion after 100 bend cycles (images were not taken before 100 bend cycles) (Figure
4-27). Water intrusion was more pronounced on Al
2
O
3
devices and could be attributed to
the reduced hydrophobicity of Al
2
O
3
compared to Parylene and the removal of Al
2
O
3

upon exposure to water by etching [44]. Based on the impedance measurements, the
water intrusion related etching of Al
2
O
3
starts on the device edges and propagates inward
towards the center. These results indicate that the open edges on the sides of Parylene
thin film devices Figure 4-2 (steps 5 & v) allow interfaces where Al
2
O
3
is in contact with
and can be etched by the surrounding water. Assuming that Parylene films can
sufficiently prevent water condensation, Al
2
O
3
should be completely encapsulated in
Parylene for devices intended for use in wet applications including implantation.
120

4.7 Conclusions
Peel tests, crush tests and bending tests were developed and used to measure the
mechanical properties of Parylene based, flexible devices with and without Al
2
O
3
. These
studies establish methods for comparing various thin film polymeric device materials and
methods. Peel tests showed that Al
2
O
3
and silane A-174 can be used to increase the
adhesion between two Parylene films, but that adhesion force is less than Parylene/Pt
bonds. Minimum bend diameter tests showed that that adding Al
2
O
3
underneath the top
Parylene allows the devices to be bent into tighter diameters when bent away from the
Al
2
O
3
, but increases the minimum diameter when bent towards the Al
2
O
3
, likely due to
poor adhesion. EIS performed on devices undergoing fatigue tests, showed that PMP can
withstand up to 100k bends without significant evidence of insulation failure. Al
2
O
3

caused insulation to fail before fatigue testing when the devices were placed in water.
Measuring several mechanical properties of a thin film device at once is valuable
for comparing and improving processing techniques. By comparing different materials
and processing methods using peel tests, crush tests, and fatigue tests, the limits of PMP
devices were found, potential methods for improving design were found, and several
methods for processing Al
2
O
3
were developed and screened.

121

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127

Implantable electronic medical devices hold promise in restoring function and
relieving symptoms due to disease, as well as monitoring the health of patients and
functionality of devices for early detection of problems and preventative care. The
performance of implantable electronics can be improved by more closely matching the
mechanical properties of the tissues they interface with and by altering the surface of the
devices to modulate the immune response to the implanted device.
Chapter 2 presented a way to alter the properties of PDMS, a material already
widely used in implantable medical devices with mechanical properties approaching
those of some tissues, to transform it into a piezoresistive device with a high gauge factor
that can be used to measure strain in tissues such as the muscles surrounding the bladder.
By incorporating a relatively small weight percentage of multi-walled carbon nanotubes
to PDMS, the polymer became piezoresistive, with the conductivity decreasing as the
polymer is stretched and the tubes are pulled apart, while maintaining its highly elastic
properties and low modulus.
Chapter 3 included the evaluation of a drug eluting coating used on a Parylene
sheath intracortical electrode designed to reduce the immune response surrounding the
implant and encourage integration into the neural tissue of the brain. The coating
consisted of the biologically derived hydrogel, Matrigel (known to encourage cell
attachment and differentiation) combined with dexamethasone to reduce the immune
response, and a combination of neurotrophins (BDNF, NGF and NT-3) to encourage
neural integration. Electrophysiological data showed that including either dexamethasone
or a combination of neurotrophins moderately, but significantly increased the event rate
measured by the electrodes during a 1-month trial. A longer 3-month trial comparing
CHAPTER 5
CONCLUSION
128

Matrigel by itself or with a combination of dexamethasone and neurotrophins to probes
with no coating did not show any significant difference between Matrigel by itself and
Matrigel combined with bioactive molecules. However, there was a significant decrease
in impedance during the first week of probes coated with Matrigel (both with and without
bioactive molecules) that disappeared during subsequent weeks, likely due to the
dissipation of factors found in the Matrigel. Also, probes coated in Matrigel, with or
without bioactive molecules, showed a significant increase in signal to noise ratio during
the second week that continued to improve over the 14 weeks of the study. These results
indicate that using biologically derived extracellular matrix can improve the interface
between recording electrodes and neurons.
Finally, chapter 4 presented methods designed to compare and screen processes
and materials used in micro-machined flexible electronics. The adhesion of the two layers
was measured using peel tests and bend tests combined with electrochemical testing.
Parylene/metal interfaces were the strongest interface tested requiring the most force to
pull apart, whereas Parylene/Parylene interfaces were the weakest. The addition of Al
2
O
3

treated with a silane adhesion promoter caused the Parylene/Parylene and Parylene/metal
interfaces to be more uniform than without the Al
2
O
3
in the middle and have an adhesion
strength between that of Parylene/metal and Parylene/Parylene. The flexibility of the
materials was measured using a minimum bend diameter test to determine the minimum
diameter that a 24 µm thick, Parylene/metal/Parylene device could be bent to before
failure. The minimum bending diameter test also showed that including one layer of
Al
2
O
3
into the device increased the minimum diameter of the device when bent towards
the Al
2
O
3
, but that it caused the devices to be able to survive being completely creased
when bent away from the Al
2
O
3
. Minimum bend diameter tests also showed that the
difficulty in processing with 2 layers of Al
2
O
3
made the variability of performance too
large to determine whether the minimum diameter was increased or decreased.
Electrochemical impedance measurements of the devices with and without Al
2
O
3
showed
that the Al
2
O
3
created pathways for water intrusion, which compromised the insulation
129

before the cables were bent, but that the Parylene insulation on the
Parylene/metal/Parylene devices without Al
2
O
3
remained intact after 100k bend cycles.
Methods of improving the interfaces between implantable electronic medical
devices and tissues can be extended beyond the examples in this work. Materials found in
FDA approved implantable devices besides PDMS can be modified to be made
conductive and used as components to make flexible electronics with mechanical
properties that more closely match tissues. Coatings made from Matrigel and bioactive
molecules can be adapted to any implanted device that needs to be more fully integrated
into the surrounding tissues, and the methods used to compare flexible electronic
materials and processes can be used to screen new materials and processes to be used as
implanted electronic medical devices. These strategies will help the exponential advances
in electronics that improve so many aspect of human life to more easily be transferred to
promoting health and treating disease.
130

APPENDIX A: DEVELOPMENT OF COATING METHOD FOR
PARYLENE-SHEATH ELECTRODE
Dip Coating
Dip coating without soak
1. 1 ml of MG was placed in an eppendorf tube and kept at 4 °C.
2. Probe was slowly lowered into eppendorf tube and immediately slowly removed.
3. Probe was then hung vertically on a clip at room temperature for at least 10
minutes before imaging.

Figure A-1. MG coating of Parylene sheath electrode after being slowly dipped in and
out of MG. (a) Optical micrograph. (b) Fluorescent image. Scale bars are 150 µm.

Dip coating with 25 minute soak
After slowly dipping the Parylene sheath electrode in and out of MG, it was
returned to MG and soaked for 25 minutes.
1. 1 ml of MG was placed in an eppendorf tube and kept at 4 °C.
2. Probe was slowly lowered into eppendorf tube and immediately slowly removed.
3. Probe was then hung vertically on a clip at room temperature for at least 10
minutes.
4. Probe was again immersed in MG and soaked for 25 minutes.
5. Probe was slowly removed from MG and hung vertically on a clip for 10 minutes
to gel before imaging.

a b
131

Figure A-2. MG coating of Parylene sheath electrode after being slowly dipped in and
out of MG and then returned to MG to soak for 25 minutes. (a) Optical micrograph. (b)
Fluorescent image. Scale bars are 150 µm.

Morphology of dip coated surfaces
To determine the morphology of MG on Parylene C, test coupons were 5 × 10
mm pieces, cut from a silicon wafer coated with 8 µm of Parylene C (Specialty
Coating Systems, Indianapolis, Indiana), which required pretreatment of the wafer
with the adhesion promoter silane A-174 (Momentive Specialty Chemicals,
Columbus, OH) to retain the deposited Parylene film. These coupons were dip
coated and then visualized under an optical microscope.

Figure A-3. Optical images of fractal ‘‘trees’’ made from dried MG on a Parylene coated
silicon wafer (a&b) under unpolarized light and (c&d) with polarized light.

The thicker regions of MG on the coupon revealed fractal patterns (Figure A-3),
which is a behavior predicted by diffusion limited aggregation theory applied to
collagen [1, 2] and shown in many other natural and synthetic polymers systems
a b
132

[3-6]. Having a complex microstructure encourages neuronal attachment as has
been demonstrated in experiments using textured [7].
Dip coating using dyed MG
Red food coloring (FC) (DecACake, Ach Food Companies, Inc., Memphis, TN)
was used in some coating experiments to visualize MG using optical microscope.
Protocol is the same as above, but MG mixture is 3:1, MG:FC.

Figure A-4. “A” cone after being dipped in dyed MG for 10 minutes. Image taken with
reflected and transmitted light. Scale bar is 150 µm.

Dip coating on heat treated Parylene
1. Parylene sheaths that were either minimally heat treated or heat treated for 48
hours were obtained.
2. 1 ml of MG was placed in an eppendorf tube and kept at 4 °C.
3. Probe was slowly lowered into eppendorf tube allowed to soak for 10 minutes and
then slowly removed.
4. Probe was then hung vertically on a clip at room temperature for at least 10
minutes before imaging.
Results

Figure A-5. (a) Minimally heat treated and (b) 48 hour heat treated cones after 10 minute
dip coating. Scale bars are 150 µm.

a
b
133

Physical Disruption of Surface Tension with Natural Filament
1. A 75 µm filament (hair) was obtained and threaded through the sheath. (For cones,
smaller filaments were obtained from a camel hair paint brush.)
2. The sheath and filament are placed onto a glass slide, which was placed on ice
and allowed to chill.
3. Two 15 µl drops of dyed MG are applied to the sheath and filament is slowly
removed.
4. Light from Vision microscope provides enough heat for MG to gel. No hang
drying is needed.

Figure A-6. (a) “C” Probe with filament inserted through sheath. (b) Sheath after being
dipped in MG and sliding out filament to break surface tension and draw MG into sheath.
(c) Sheath that was dipped in MG without using filament. (d) “A” probe with coating
pulled through sheath using filament. Scale bars all 150 µm.

Activating Surface with O
2
Plasma Surface Treatment
1. Sheath electrode is exposed to 1 min of O
2
plasma (273 watt, 400 mTorr).
2. Sheath electrodes are removed from chamber and directly placed in dyed C. Since
processing is done in clean room, dyed C (C:food coloring, 3:1), is held on ice to
keep it liquid.
3. Sheath electrode remains in dyed C for 10 minutes after which it is hang dried for
an additional 10 minutes before being covered and stored in a petri dish.

a
b
c d
134

Figure A-7. Cone dip coated after exposure to O
2
plasma. Scale bar is 150 µm.

Vacuum Soak
A 4:1, C-coating to food coloring mixture was kept at 4 °C by placing eppendorf
tube with mixture in a beaker filled with ice.
2. Parylene Sheath electrode was placed into MG mixture and beaker with ice and
MG mixture were placed in vacuum oven (with heating element turned off).
3. Vacuum oven was pumped down and Parylene sheath electrodes were allowed
to soak for 5 minutes under vacuum.
4. Vacuum was vented and the sheaths were removed from the mixture and
allowed to hang dry for 5 minutes at room temperature to gel.

Figure A-8. “A” probes dip coated using vacuum. (a&b) Optical and fluorescent
microscope images of outside surfaces of probe after vacuum dip coating. (c&d) Optical
and fluorescent microscope images of inside surfaces of probe after having outer layer
removed. Each row represents the same probe. Scale bars are 150 µm.

a b c d
135

Figure A-9. “B” probe dip coated using vacuum. (a&b) Optical and fluorescent
microscope images of outside surfaces of probe after vacuum dip coating. (c&d) Optical
and fluorescent microscope images of inside surfaces of probe after having outer layer
removed. All images are of same device. Scale bars are 150 µm.

Surface Wetting with Ethanol
A 4:1 C-coating: food coloring mixture was kept at 4 °C in and eppendorf tube.
2. A second eppendorf tube was filled with ethanol.
3. Parylene sheath electrode was slowly dipped in and out of ethanol 5 times and
then placed in MG mixture for 5 minutes.
4. Parylene Sheath was removed from mixture and hung vertically on a clip for 5
minutes at room temperature to gel.

Figure A-10. “B” probes dip coated after being wetted with ethanol (a&b) Optical and
fluorescent microscope images of outside surfaces of probe after vacuum dip coating.
(c&d) Optical and fluorescent microscope images of inside surfaces of probe after having
outer layer removed. Each row represents the same probe. Scale bars are 150 µm.
a b c d
a b c d
136

Figure A-11. Zoomed in image of MG structure on inner electrode at small end of cone.
Scale bar is 30µm.

Manually Filling Cone with Syringe
Double sided tape was used to attach Parylene sheath to a stationary glass slide.
2. A 33 gauge needle was attached to a 3-D positioning stage.
3. Needle was lined up with Parylene sheath under stereo microscope and the
syringe is manually depressed.

Figure A-12. Filling an “A” sheath electrode with a 33 gauge needle. (a) Needle
approaching sheath. (b) Needle making contact with sheath. (c) Filling sheath with DI
water. (d) Needle is retracted, sheath is filled with DI water. Scale bars are 200 µm.

Sonication
Sonication at room temperature
a b
c d
137

1. MG is mixed with food coloring using a 1.5 ml eppendorf tube and vortex mixer
in a ratio of 3:1 and kept at 4 °C to ensure MG remains liquid.
2. While keeping MG mixture at 4 °C, probes are gently slid into eppendorf tube
with sheaths facing the middle of the tube, and eppendorf tubes are placed into a
sonicating bath filled with ice.
3. Probes remain submersed for 1, 5, or 10 min. while sonicating and then are slowly
removed from the eppendorf tube.
4. Probes are hung on a clip at room temperature (23 °C) for at least 10 minutes
before placing probe between two texwipes inside a petri dish. (Texwipes help
prevent static cling to plastic petri dish).

Figure A-13: Parylene Sheath electrodes sonicated for (a,b) 1, (c,d) 5, and (e,f) 10
minutes. (a,c,e) Micrographs taken from optical microscope. (b,d,f) Fluorescent images.
All scale bars are 200 µm.

a
b
c
d
e
f
138

Sonication without dye
1. C-coating is kept at 4 °C inside of an eppendorf tube placed in a sonicating bath.
2. Probes were sonicated 10 minutes.
3. While keeping MG mixture at 4 °C, probes are gently slid into eppendorf tube
with sheaths facing the middle of the tube and the tube is placed in sonicating
bath.
4. Probes remain submersed for 10 min. and then are slowly removed from the
eppendorf tube.
5. Probes are hung on a clip at room temperature (23 °C) for at least 10 minutes
before placing probe between two texwipes inside a petri dish. (Texwipes help
prevent static cling to plastic petri dish).

139

Figure A-14. (a-d) SEM images taken of “B” shaped sheath electrode after being dipped
in MG. (a) Matrigel can be seen gathered on the sides of the sheath, but the top surface is
relatively clean. (b & c) Close up images of coating on the right side of sheath. Complex
micro and nanostructures may encourage neurite attachment. (d) A crack in the MG
coating reveals that the thickness of the coating is ~1µm. (e) Close up of the tip of sheath
shows coating at tip of sheath.

Sonication in cold room
There was a concern that ice in the sonicating bath was absorbing some of the
sonicating energy. Tests were done with probes being sonicated in a cold room
(kept at 4 °C) to eliminate the use of ice.
1. Sonicating bath is filled with DI water in a cold room.
b
a
d
c
e
140

2. C-coating is mixed with food coloring in a 4:1 ratio (C-coating: food coloring)
and mixed using a vortex mixer.
3. Parylene sheath is submersed in MG mixture and placed in sonicating bath.
4. Sheath is sonicated for 5 minutes in cold room.
5. Sheath is removed from sonicating bath, hung vertically on a clip and placed in
a Styrofoam box that has been kept at room temperature to gel for at least 5
minutes.

.
Figure A-15. “B” Probes dip coated using sonication in a cold room. (a&b) Optical and
fluorescent microscope images of outside surfaces of probe after sonication. (c&d)
Optical and fluorescent microscope images of inside surfaces of probe after having outer
layer removed. Each row represents the same probe. Scale bars are 150 µm.

Figure A-16. Close up of inner most electrode covered with MG. Scale bar is 30µm.

a b c d
141

Figure A-17. “B” probes coated in C-coating. (a) Optical and (b) fluorescent
micgrograph of outside of Parylene sheath. (c) Optical and (d) fluorescent micrograph of
Parylene electrode with the sheath removed. Scale bars are 150 µm.

Figure A-18. “B” probes coated in N-coating. (a) Optical and (b) fluorescent
micgrograph of outside of Parylene sheath. (c) Optical and (d) fluorescent micrograph of
Parylene electrode with the sheath removed. Scale bars are 150 µm.

Figure A-19. “B” probes coated in I-coating. (a) Optical and (b) fluorescent micgrograph
of outside of Parylene sheath. (c) Optical and (d) fluorescent micrograph of Parylene
electrode with the sheath removed. Scale bars are 150 µm.

a b c d
a b c d
a b c d
142

Figure A-20. “B” probes coated in E-coating. (a) Optical and (b) fluorescent
micgrograph of outside of Parylene sheath. (c) Optical and (d) fluorescent micrograph of
Parylene electrode with the sheath removed. Scale bars are 150 µm.

Multiple Coatings
1. C-coating was mixed with diluted food coloring (DI water was used to pass food
coloring through filter paper) in a ratio of 4:1 (C-coating: food coloring) and kept
at 4 °C.
2. Parylene sheath was dipped into dyed MG mixture (either while sonicating or
without sonication) for a 5 seconds and then removed
3. Parylene sheath was placed into a clip and held under a stereo microscope until
mixture crystalized (~90 seconds).
4. This process was repeated 10 times.

Figure A-21. “B” type probe during multiple coating experiment without sonication. (a)
After 1
st
coating. (b) After 5
th
coating. (c) After 10
th
coating. Scale bars are 150 µm.

a b c d
a b c
143

Figure A-22. “B” type probe during multiple coating experiment with sonication. (a)
After 1
st
coating. (b) After 5
th
coating. (c) After 10
th
coating. Scale bars are 150 µm. Note
that (c) is taken from the backside of probe.

1. Probe was dipped into 70% ethanol for 3 seconds and then immediately submersed
into a solution of poly-D-lysine (PDL) (P6407, Sigma Aldrich, St. Louis, MO) (100
µg/mL ) and held at 4 °C for 1 hour.
2. The probe is rinsed three times in triple distilled and then inserted into a 10 µL
droplet of MG and held 50 °C inside a closed polystyrene container (to prevent
evaporation) for 5 min.
3. After the 5 min, the cover of the container is removed and the probe held at 50 °C for
an additional 5 minutes to dehydrate the coating.
PLGA-based Coatings
Thickness Calibration
1. Solutions of PLGA:DCM were mixed to various concentrations: 200, 100, 50, 25,
and 12.5 mg/mL.
2. Coupons (0.5 x 5 x 10 mm sections of silicon wafers coated with Parylene C) are
slowly dipped into the solution using tweezers and gradually pulled out.
3. Coated samples are held at atmosphere for 0.5 h and overnight in dessicator to
remove DCM.
4. Samples are scanned using a profilometer to determine coating thickness.
PLGA coating on Perforated Probe
1. A mixture of 9:1, DCM:MeOH was mixed with DEX and PLGA to get a
concentration (of both substances) of 10 mg/mL (of solvent).
2. The probe was dipped into the coating solution (in a 70 µL microwell) and held
for a count of three and then pulled out.
3. Coated probe was held in a vacuum desiccator until ready to SEM.
Absorbance Protocols
Measuring Concentration of Factors (Using Spectrophotometry)
Materials:
 96 well microplates (UV transparent)
a b c
144

 Centrifuge tubes (500 µL and 1.5 mL, lowbind eppendorf tubes)
 Micropipette tips (100 µL, 1000 µL)
 Reagent wells
 Sandwich ELISA kits (Rat NGF/NGF beta kit Protocol, BEK-2077-2P, biosensis,
Thebarton, Austrialia; NGF Rapid ELISA Kit: Rat, BEK-2214-1P, biosensis,
Thebarton, Austrialia; Human neurotrophin-3 ELISA Kit, EK0472, amsbio,
Bioggio-Lugano, Switzerland)

Equipment:
 Vortex mixer
 Spectrophotometer (Epoch Microplate Spectrophotometer, Biotek, Winooski, VT,
USA)
 Micropipettes (p100, p1000, multichannel 50-300 µL)
 Samples

Prepare standards:
Dilute standards from kit according to manufacturer’s protocol using buffer solution,
vortex mixer and micropipettes. Store in eppendorf tubes at 4 °C until scan.

Sample preparation and storage:
1. After sampling from the microwell, samples are diluted by 10x by placing 50 µL
of sample into 450 µL of sample buffer from the kit and stored at 4 °C. (Samples
will need to be diluted more before running ELISA; however, higher
concentrations of samples are more stable for storage. They are not kept at full
concentration because the buffer also has preservatives in it to help the proteins
remain stable.)
2. Before running ELISA, samples are serially diluted to a target concentration in
the middle of the sensitivity range of the ELISA kit (as reported by kit
manufacturer). Samples are diluted using eppendorf tubes, sample buffer (from
kit) and pipettes. The maximum dilution is 30×, which is done by placing 50 µL
of concentrated sample into 1450 µL of buffer. In general, maximize the amount
of concentrate (use no less than 50 µL) and minimize the amount of buffer.

ELISA:
1. ELISA is performed according to protocols provided from the ELISA kit
manufacturer. The output of the ELISA will be absorbance measurements for the
standards solutions as well as absorbance measurements for each sample.
2. A concentration versus absorbance plot is made using the standards, and a 4
parameter logistic regression (4PL) is used to fit a curve to this data. (Regression
software available from platereader software or from www.myassays.com.)
3. The curve generated from the 4PL regression is used to find concentration of
diluted samples, which is then multiplied by the dilution factor to get the original
sample concentration.
145

Measuring Dexamethasone (DEX) Concentration (Using Spectrophotometry)
Materials:
 96 well microplates (UV transparent)
 Water soluble dexamethasone (for standard)
 Centrifuge tubes (500 µL and 1.5 mL tubes)
 Parafilm
 Micropipette Tips (100 µL)
 Samples

Equipment:
 Vortex mixer
 Spectrophotometer (Epoch Microplate Spectrophotometer, Biotek, Winooski, VT,
USA)
 Micropipettes (p100, p1000)

Prepare standard curve:
1. Dilute dexamethasone aliquots (20 mg/mL) to 200 µg/mL by serially diluting the
aliquot by 10× twice by placing 100 µL of aliquot in 900 µL of PBS, vortex
mixing the resulting mixture, and then placing 100 µL of that mixture into 900 µL
of PBS.
2. Serially dilute the 200 µg/mL sample five times (by placing 500 µL of
concentrate into 500 µL of diluent, vortex mixing, and repeating) to get the
following concentrations (or standards): 200 µg, 100 µg, 50 µg, 25 µg, 12.5 µg,
and 6.25 µg.
3. For each concentration, fill two microwells with 100 µL of the standard.
4. Scan microwells at 242 nm using platereader.
5. Plot concentration versus absorbance (average the absorbances for each
concentration) and generate a linear fit for the model. This linear fit will be used
to back calculate concentration from future absorbance measurements of the
sample.

Measuring absorbance of samples (with probes submerged in microwells):
1. At each time point, the probe is removed and the plate is scanned at 242 nm.
2. The absorbance is compared to the standard curve to calculate the concentration
of dexamethasone.

Measuring absorbance of sidewall experiments:
1. Set up plate reader for a kinetic scan to measure absorbance of each sample at 242
nm every 2 min. for the first 20 min, and every 5 min. for the next 2 hours.
2. Place 100 µL of PBS in all the wells with the coating stuck on the sidewalls.
146

3. Cover plate with Parafilm to prevent evaporation. (Note: make sure that Parafilm
is stretched until clear before placing over microwell plate.)
4. Place microwell plate in platereader and start kinetic scan.
5. Compare absorbance to standard curve to get concentration of samples.

Elution of Matrigel from Microwells
1. 30 µL of Matrigel (MG) was loaded onto side wall of well.
2. 150 µL PBS was added to the well, and MG diffusion occurred across the well.
3. An optical density measurement (280 nm) was taken at the center over the course
of one hour. 5 sec. for the first 10 minutes, 10 seconds for the next ten minutes;
30 seconds for the next 40 minutes

Figure A-23. Schematic of experiment showing dissolution of Matrigel showing (a) top
view and (b) sideview.

Elution of DEX from Parylene Coupons
1. 6 Parylene coated silicon coupons (0.5 x 5 x 10 mm)
2. Coupons dip coated by fully submersing in coating solution (9:1,
DCM:MeOH; 10 mg/mL DEX and PLGA)
3. Coupons sat on TEXwipes at room temperature and pressure for 0.5 h and in
a vacuum desiccator overnight to evaporate solvent.
4. The coupons are submersed in PBS (in either a 0.5 or 1 mL eppe tube) for
various time points.
5. After a given time period, the PBS is removed, scanned at 242 nm using a
spectrophotometer, and replaced with fresh PBS (Figure A-24).

(
a
)
(
b
)
147

Figure A-24. Elution rate of DEX from DEX loaded PLGA coupons at discreet time
points.

Elution of DEX from Microwells
1. Place 10 µL of DEX/PLGA in the corner of a 96 well microwell plate and allow
to dry for 5 minutes. (475 µL of DEX in 8.5:1 DCM:MeOH + 25 µL of 200
mg/mL PLGA in DCM)
2. Place 1, 3, or 5 drops of 10 µL of a double concentrated PLGA (50 µL of 200
mg/mL in 450 µL of 8.5:1 DCM) on top of the DEX/PLGA mix.
3. Leave wells in hood overnight to get rid of excess solvent.
4. Do a kinetic scan using spectrophotometer (242 nm).
Elution of NGF from Coupons
1. 1 µL Neurotrophin cocktail (diluted 1:100) was placed onto tip of Polystyrene
coupon
2. The coupon was baked at 37C for 30 min
3. 100 µL PBS was added to each ELISA microwell
4. Coupon was dipped into microwell for 20s, 40s, 60s, 5m, 10m, 20m

148

Figure A-25. Schematic of coupon study with Matrigel loaded with neurotrophic factors.

Elution of NGF and BDNF from Microwells

Figure A-26. Schematic for the deposition of coating on microwells. Modeled after the
surface modification method (Figure 3-4) there are 3 steps: (1) the Parylene coated
surface is exposed to 100 µg/mL PDL solution, (2) a 1 µL droplet of MG is placed on the
bottom of the well which is covered and heated to 50 °C for 5 min, and (3) the cover is
removed and the microwell is again heated to 50 °C for 5 minutes to dehydrate the
coating.

Once the coating was deposited in the microwells the elution of factors was
monitored as follows:
149

1. The wells were filled with 100 µL of PBS and the factors were allowed to
elute into the PBS.
2. At specific time points, the eluate (PBS plus factors) was removed and
replaced with fresh PBS. The eluate was diluted in an assay buffer and
stored at 4 °C.
3. After complete elution time was completed. Sandwich ELISA was used to
measure the concentration of factors at each time point and the amount (in
ng) of factors released from each time point was calculated.

REFERENCES
[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

150

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.
151

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

152

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

[1] J. Parkinson, K. E. Kadler, and A. Brass, "Self-assembly of rodlike particles in
two dimensions: A simple model for collagen fibrillogenesis," Physical Review E,
vol. 50, p. 2963, 1994.

[2] T. A. Witten, Jr. and L. M. Sander, "Diffusion-Limited Aggregation, a Kinetic
Critical Phenomenon," Physical Review Letters, vol. 47, pp. 1400-1403, 1981.

[3] D. Zhang, J. Li, S. Chen, T. Li, J. Zhou, X. Cheng, et al., "Hybrid Self ‐
Assembly, Crystal, and Fractal Behavior of a Carboxy ‐Ended Hyperbranched
Polyester/Copper Complex," Macromolecular Chemistry and Physics, 2012.

[4] H. Gan, Y. Li, H. Liu, S. Wang, C. Li, M. Yuan, et al., "Self-assembly of
conjugated polymers and ds-oligonucleotides directed fractal-like aggregates,"
Biomacromolecules, vol. 8, pp. 1723-1729, 2007.

[5] W. Wang and Y. Chau, "Self-assembled peptide nanorods as building blocks of
fractal patterns," Soft Matter, vol. 5, pp. 4893-4898, 2009.

[6] Q. Zhao, J. Qian, Z. Gui, Q. An, and M. Zhu, "Interfacial self-assembly of
cellulose-based polyelectrolyte complexes: pattern formation of fractal “trees”,"
Soft Matter, vol. 6, pp. 1129-1137, 2010.

[7] H. Craighead, C. James, and A. Turner, "Chemical and topographical patterning
for directed cell attachment," Current opinion in solid state and materials science,
vol. 5, pp. 177-184, 2001.

153

APPENDIX B: RECIPE FOR MECHANICAL TEST DEVICES
Obtain prime 4" Si wafer out of box
Bake wafer (140 °C, for >10 min.)
Deposit Parylene (12 µm, ~28 mg)
Atomic layer deposition of Al
2
O
3
– (25 nm, 0.065 s pulse, 90 s purge, 200 cycles)
performed by J Provine at Stanford (independentj@gmail.com)*
AZ 5214 E –IR Mask 1: Metal
5 sec @ 500 rpm; 45 sec @ 2 krpm
Bake 1:10 min. @ 90 °C
Expose 50 mJ = 10 mW/cm
2
*5 s
IR bake 45 sec @ 120 °C
Global exposure 300 mJ = 10 mW/cm2 * 30 s
Develop: 20-22 sec (1:4 dilution)
Descum 100 W: 100 mT : 30 sec
Ebeam Ti (1 run 200 Å)*
Ebeam Pt (3 runs of 666 Å for a total of 2000 Å)
Liftoff
Soak in acetone until PR is removed
Rinse in IPA, DI H
2
O
AZ 4400 Mask 2: Sacrificial Layer
5 sec @ 500 rpm; 45 sec @ 4krpm (4 µm)
Bake 2 min @ 90° C
Expose 150 mJ = 10 mW/cm2 * 15 sec
Develop 45-50 sec.
Pseudo-hardbake: Bake 3 min. @ 90 °C
Descum 100 W: 100 mT : 30 sec
Deposit Parylene (12 µm, 28 g)
ALD Al
2
O
3
– (~25 nm, 0.065 s pulse, 90 s purge, 200 cycles) performed by J
Provine at Stanford (independentj@gmail.com)**
AZ 4620 Mask 3: Contact Pad
5 sec @ 500 rpm (accl: 5); 45 sec @ 2 krpm (accl: 15) for (~9.6 µm)
Bake 6 min @ 90 °C
AZ 4620 (2
nd
layer)
5 sec @ 500 rpm; 45 sec @ 2 krpm for (~9.6 µm)
Bake 7 min @ 90 °C
Expose 550 mJ = 10 mW/cm
2
* 55 sec
Develop 2-5 min. (make sure that non exposed area is clear)
Hard bake 100 °C, 1 min
RIE insulation layer etch 100 W : 100 mT : 5 min × 10-12 cycles (let wafer sit
and cool after every 5 min. cycle, rotate wafer after every 3
rd
5 min. cycle,
continue until contact pads are clear)
DO NOT CLEAN OFF MASK
154

AZ 4620 Mask 4: Cutout
5 sec @ 500 rpm; 45 sec @ 2 krpm for (~9.6 µm)
Bake 5-6 min @ 90 °C
AZ 4620 (2
nd
layer)
5 sec @ 500 rpm; 45 sec @ 2 krpm for (~9.6 µm)
Bake 5-7 min @ 90 °C
Expose 550 mJ = 10 mW/cm
2
* 55 sec
Develop 3-5 min (make sure that non exposed area is clear)
Develop for additional 6-8 minutes**
RIE insulation layer etch 100 W : 100 mT : 5 min × 8-11 cycles (let wafer sit and
cool after every 5 min. cycle, rotate after every 3
rd
5 min. cycle, continue until
bare silicon wafer is clearly visible)
KOH etch: soak released devices in 20 wt% KOH in diH
2
O, at room
temperature**

* For PAMAP Devices
**For PAMAP and PMAP Devices
155

APPENDIX C: MASKS FOR MECHANICAL TESTING DEVICES

156

157

158

Abstract (if available)

Abstract Implantable medical devices hold great promise to treat and prevent chronic conditions, but are limited by their biocompatibility. Immune responses due to mechanical mismatches and surface properties such as topology and biochemistry, can be dangerous to the patient, degrade the device, or prevent the device from performing its functions correctly. The biocompatibility of electronic devices is particularly challenging because they typically contain circuits patterned on silicon wafers, which makes them prone to be much harder than tissue. They also require hermetic seals to protect the circuitry from liquid and intimate contact with tissues to monitor tissue properties and safely deliver therapeutic amounts of current. Even moderate immune responses can degrade protective coatings and increase the impedance of electrodes. This work presents strategies to improve the interface between implanted medical devices and biological tissues. ❧ The first strategy presented is decreasing the mechanical mismatch of implantable electronics. This is demonstrated by the development of a strain sensor that could be used to measure the extension of muscles such as the bladder as part of a prosthetic to determine bladder fullness. The device is made from polydimethylsiloxane, a material already used in many implantable devices because of its inert chemical structure and soft tissue‐like mechanical properties, mixed with carbon nanotubes which make the rubber piezoresistive at relatively low weight fractions. ❧ The next strategy is treating an already developed flexible electronic device with a bioactive coating. In this strategy, a flexible thin film implantable intracortical electrode is coated with biologically derived extracellular matrix proteins mixed with bioactive molecules. The extracellular matrix proteins alters the topological, mechanical and chemical structure of surface and acts as a substrate for cells to bind to in order to make intimate contact with the electrodes on the device. Bioactive molecules known to reduce the immune response (dexamethasone) and encourage cell growth and differentiation (neurotrophins) were also added to the extra cellular matrix proteins to further improve integration. ❧ The last strategy is to use a combination of mechanical tests to test and screen thin film flexible electronic architectures, which show promise as implantable medical devices. By measuring the strength of adhesion forces between the layers of the thin films along with their flexibility and resistance to fatigue, one can compare different materials and manufacturing methods to quickly compare and improve the properties of new thin film flexible electronics architectures. ❧ Through these strategies, the interfaces between tissues and medical devices can be improved and enable implantable devices to function better and deliver novel therapy for chronic conditions.

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Asset Metadata

Creator Lee, Curtis Dixon (author)

Core Title Strategies for improving mechanical and biochemical interfaces between medical implants and tissue

School Viterbi School of Engineering

Degree Doctor of Philosophy

Degree Program Biomedical Engineering

Publisher University of Southern California (original), University of Southern California. Libraries (digital)

Tag bioMEMS,carbon nanotubes,graphene,implantable electronics,Matrigel,microfabrication,neural electrode,OAI-PMH Harvest,Parylene C,PDMS,thin film electronics

Format application/pdf (imt)

Language English

Contributor Electronically uploaded by the author (provenance)

Advisor Meng, Ellis (committee chair), Gupta, Malancha (committee member), Weiland, James D. (committee member)

Creator Email cleedix@gmail.com,curtisle@usc.edu

Permanent Link (DOI) https://doi.org/10.25549/usctheses-c3-588898

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Access Conditions The author retains rights to his/her dissertation, thesis or other graduate work according to U.S. copyright law. Electronic access is being provided by the USC Libraries in agreement with the a...

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