Novel Polypyrrole Derivatives to Enhance Conductive Polymer-Tissue
Interactions
By Paul M. George
BSE, Biomedical Engineering, Tulane University
MSE, Biomedical Engineering, Johns Hopkins University
SUBMITTED TO THE HARVARD-MIT DIVISION OF HEALTH SCIENCES AND
TECHNOLOGY IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE
DEGREE OF
DOCTOR OF PHILOSOPHY IN MEDICAL AND ELECTRICAL ENGINEERING
AT THE
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MASSACSTTS INSituTE
OF TECHNOLOGY
JULY 2005
0CT
9 2005
LIBRARIES
( 2005 Paul M. George. All rights reserved.
The author hereby grants MIT permission to reproduce and to distribute publicly paper
and electronic copies of this thesis document in whole or in part.
Signature of Author:
J
Harvard-MIT Division of Health Sciences and Technology
,, y , 2005
Certified by:
Robert Langer, Sc.D.
Institute Professor
Harvard Division of Health Sciences and Technology
Thesis Supervisor
Accepted by
Mart a L.
ray, Ph.D.
Edward Hood Taplin Professor of Medical and Electrical Engineering
Co-Director, Harvard-MIT Division of Health Sciences and Technology
ARCHIVES
Abstract
Developing materials that interact effectively with surrounding tissue is a major
obstacle in sensor and drug delivery research. The body's natural immune response
prevents foreign objects from easily integrating with an organism. Without an intimate
link between a biomedical device and the proximate environment, reliable measurements
or delivery of molecules is not possible. Many of the current materials used for
biomedical applications are centered on inert substances and polymers that degrade in the
body but have limited functional capabilities. This thesis work addresses the need to
develop materials that are capable of interacting in biological environments.
Polypyrrole (PPy) is a conducting polymer that is a promising biomaterial for
drug delivery and sensing applications. Because PPy is a polymer that can be made in
degradable forms and because it can be stimulated electrically, it is an interactive
platform for biomedical applications. By accomplishing the following research
objectives, this thesis work could help develop an effective polymeric paradigm for tissue
interactions:
1) Develop a new method to effectively micro-pattern electrodeposited polymers and
metals for in vivo devices
2) Determine the optimal synthesis conditions of the conductive polymer, PPy, for
sensor and implant applications
3) Fabricate PPy tubes to be used as nerve guides to promote nerve regeneration
4) Modify PPy for neurotrophic factor drug delivery devices and antibody-based
sensing applications
Through the use of standard microfabrication techniques, the patterning template
upon which PPy is electrodeposited can be controlled precisely. By utilizing the growth
mechanism of PPy on these templates, three-dimensional polymer objects can be created.
Being able to micropattern the PPy and release the polymer generates the ability to create
implants and devices that are completely erodible in the body.
To develop the optimum conditions for sensor and drug delivery applications, PPy
implants were fabricated and implanted into rat cortical tissue. Compared to similar
Teflon implants, the electrically conductive PPy had preferable characteristics for
material integration in the cortex. Additionally, PPy tubes have been designed and
promoted peripheral nerve growth after tissue injury. By controlling the shape and
morphology of PPy, the polymer implants formed an interactive bridge with their
biological environment.
By incorporating bioactive molecules into the PPy matrix, materials for externally
controlled drug release and sensing devices can be designed. Drug delivery was
demonstrated through the integration of nerve growth factor (NGF), a neurotrophic
factor, into the PPy followed by triggered pulsatile release. Such neurotrophic factors
can be used to promote neural growth in peripheral and central nervous system injury.
Because PPy is easily modifiable through the use of dopants and control of its shape, PPy
provides a flexible platform for novel polymeric-tissue interactions.
2
Acknowledgments
Any body of research truly involves a team effort, and I have been blessed by the
numerous people who have made this work possible. I would like to sincerely thank my
advisor, Robert Langer, for providing guidance and mentorship throughout my doctoral
work. His insights into how research can make an impact in the lives of others and his
integrity as a researcher and a person have shaped my graduate studies and experiences.
David LaVan is another mentor that has helped to direct my research, and his friendship
and advice have been invaluable in my development as a researcher. I would also like to
thank the other members of my thesis committee, Dr. Joel Voldman and Dr. Alan
Grodzinsky, for their insight and assistance during my thesis work. I have also had the
privilege of working with Mriganka Sur whose collaborations have brought the cerebral
aspect into much of my research. The work with his lab has been some of the most
enjoyable of my doctoral work. Jason Burdick is another friend and mentor who has
contributed to my growth as a researcher and helped to direct my research to this point.
Without the interactions and guidance of all of these friends and colleagues, my graduate
work would not have been possible, and the knowledge that I have gained would have
been limited.
Equally important to me has been the support of my family and friends. My wife,
Sierra, has kept me grounded to what is truly important in life. Her belief in me, her love,
and her laughter have inspired me throughout my research and life. I have also been
blessed with parents, a mother-in-law, and an uncle-in-law who have provided the love
and encouragement to fulfill my dreams for which I am very grateful. My wonderful
friends have also kept me balanced and lighthearted throughout my graduate experience.
Many members of the Health Sciences and Technology community have also helped
along the way. Catherine Modica and Patty Cunningham have both been so supportive
and caring as I navigated through my graduate studies.
I am indebted to numerous other lab members have contributed along the way.
Dan Kohane, Steve Chen, and Yadong Wang have provided valuable insight, and I have
also worked with many great undergraduate researchers, Ellen Liang, Moira Kessler,
Anita Hegde, Phillip Alexander, and others, whose assistance has been tremendous.
Gwen Donahue has also provided important microfabrication expertise throughout my
work for which I am extremely thankful.
I would also like to thank the Whitaker Foundation, the Dupont-MIT Alliance,
and the NIH for their funding and support. Without them and other funding agencies,
research and the search for knowledge would not progress.
My graduate experience has been a wonderful one. I truly have been blessed to
have worked with so many caring, giving, and intelligent people.
3
Table of Contents
Abstract .......................................
...............................................................2
Acknowledgments ........................................
..............................................
3
Table of Contents ..............
1.
2.
4
Introduction...........................
..
.
...........................................................
6
1.1.
Biopolymers .........................................
1.2.
Conductive polymers .................................................................................. 8
1.3.
Polypyrrole................................................................................................ 15
1.4.
PPy applications........................................................................................ 25
1.5.
Specific aim s .........................................
31
1.6.
References .................................................................................................
32
Patterning ......
.............................
6
.
......... 45
2.1.
PPy growth................................................................................................
45
2.2.
Single-step 3-D electrodeposition.............................................................
46
2.3.
LIGA .........................................
47
2.4.
Micropatterning......................................................................................... 48
2.5.
3-D structures............................................................................................ 50
2.6.
M orphology ...............................................................................................
53
2.7.
References ..........
57
................................................................
3. Biocompatibility......................
4.
59
3.1.
Polypyrrole/Tissue interaction .................................................................. 59
3.2.
Experimental overview............................................................................. 60
3.3.
Electrodeposition ......................................................................................61
3.4.
In vitro study techniques ........................................................................... 63
3.5.
In vivo study techniques ............................................................................ 65
3.6.
In vitro study results.................................................................................. 68
3.7.
In vivo study results .........................................
73
3.8.
Conclusion ................................................................................................
77
3.9.
References ...........
79
.........
................... .......................................
PPy tubes ...........................................................................................
4
83
4.1.
Peripheral nerve regeneration ................................................................... 83
4.2.
PPy tubes................................................................................................... 88
4.3.
Formation of PPy tubes............................................................................. 89
4.4.
Sciatic nerve study .................................................................................... 93
4.5.
Future work....................................
100
4.6.
References ........................................
102
5. Drug delivery................................................................................... 109
6.
7.
5.1.
Mechanisms of drug delivery..................................................................
109
5.2.
Experimental overview ........................................
109
5.3.
Electrodeposition
111
5.4.
Stability studies ....................................................................................... 111
5.5.
Drug release studies ................................................................................ 112
5.6.
PC-12 cell studies ........................................
113
5.7.
Dopant stability.........................................
114
5.8.
Biotin/NGF attachment ........................................
115
5.9.
Molecule release ........................................
117
5.10.
PC-12 neurite outgrowth ........................................
118
5.11.
PPy drug delivery.................................................................................... 119
5.12.
Hydrogel drug delivery........................................
120
5.13.
tPA ........................................
120
5.14.
tPA release studies ........................................
120
5.15.
References ........................................
125
....................................................................................
Future directions.............................................................................
130
6.1.
PPy modification.....................................................................................
130
6.2.
RF applications of PPy...................................
130
6.3.
PPy neural scaffold ................................................................................. 133
6.4.
References ........................................
Appendix..........................................................................................
Biographical Information ........................................
5
135
136
141
1. Introduction
1.1.
Biopolymers
The versatility of polymers and their ability to interact with tissue have made them a
key fixture in the development of biomedical sensors and drug delivery devices
18-24
Another advantage of polymer-based biomaterials is the ease of control over degradation
properties, surface properties, and the mechanical properties of the materials
25.
With
greater control of biomaterial properties, more effective drug delivery and tissue
interactions are possible
26.
Many polymers have been designed to allow for controllable degradation 25.
Depending on the application, the requirements of the polymer can vary greatly. Some
applications require a long lasting polymer and others require a degradable form. The
development of various degradable polymers has allowed for the release of bioactive
molecules that directly effect the surrounding environment. Polyesters such as poly(dllactic a cid) (PLA) and poly(glycolic acid) (PGA) w ere found to degrade in b iological
environments into naturally occurring products
poly(DL-lactic-co-glycolic acid) (PLGA)
28
27.
Other biomedical polymers such as
and rosin-based materials
29
have been
designed to undergo bulk degradation and have good biocompatible properties. Over the
past 20 years, biodegradable polymers such as the ones mentioned above have been
characterized in biological environments 30
Biodegradation of polymers may occur through various means. Many degradable
polymers are broken down through hydrolysis of components of the backbone (scission)
30.
This can occur either enzymatically or randomly without requiring enzymes.
6
Hydrophobic
material
polymers degrade at the surface because water cannot reach inside the
as opposed
degradation.
to hydrophilic
polymers
which can swell resulting
Polymers with strongly bonded backbone
hyrdolyzable groups require long time frames to degrade.
biodegradable
polymers are polylactides
glycolide and polyhydroxybutyrates
30.
in bulk
(ie C-C bonds) with no
Two of the main families of
including copolymers
of polylactides
These degrade though scission and degradation at
the surface.
_a
aa
aa
Hydrogels are another material
that has been developed to carry and
release drugs
21,31.
Hydrogels allow for
greater control of polymer formation
and degradation.
Some systems have
been developed for in situ crosslinking
controlled by ultraviolet (UV) light
32.
IIg
JljJ\
g
Others have been designed to release in
response to pH or electrical signals
The
incorporation
33.
of cells such as
.............................
osteoblasts into these gels has also been
developed
for
tissue
engineering
Figure 1.1. Images of biodegradable
polymers at various degradation time
points. PGS at A) 0 days B) 7 days
C) 14 days D) 21 days E) 28 days F)
35 days. PLGA at G) 0 days H) 7
days I) 14 days J)21 days. The PGS
material maintains its shape and
properties more effectively than
PLGA as it degrades. From 6.
..
34
app IIcatlOns .
Materials
have
also
been
designed to match the environment into
which they are placed.
with
Ideally they
7
would
keep these
properties
throughout their
degradation
(See
Figure
1.1).
Polyl(glycerol sebacate) (PGS) has been shown to maintain its geometry and lose
mechanical stability much more slowly than PLGA as it degrades 6,35.
Polymers for biomedical applications have evolved from mostly inert substances
such as teflon to more interactive degradable polymers such as hydrogels and PGS.
Currently, materials are being developed that can dynamically interact with their
environment. Finding polymers that are able to reduce the foreign-body reactions of the
body's immune system and promote positive polymer-tissue interfaces is important for
effective device applications
36.
MEMS devices can have pulsatile release patterns once
in the body 37, and polymeric systems can also be developed for externally modulated
delivery 38.
Conductive polymers are one material being considered that can be
interactively controlled once in the body. Also, because of their ability to deliver charge,
conductive polymers are ideal for more effective interaction with neural cells.
1.2.
Conductive polymers
1.2.1. Conductivepolymer classes
Two general types of conductive polymers exist.
One group is a composite
material that uses a polymer to hold together conductive filler such as metal flakes or
carbon black.
The second group consists of a set of polymers whose backbone
intrinsically propagates charge, making the polymer itself conductive (See Figure 1.2) 39.
One problem with conductive polymers made with a conductive filler is that the need to
use a large percentage of filler creates poor mechanical properties
8
40.
The ability to
control conductivity, the majority of the charge carriers, and to have a completely
polymeric system makes the intrinsically conductive polymers more appealing for
biomedical applications.
Conductive polymers such as polpyrrole (PPy), polyacetylene, polythiophene, and
polyaniline have experienced much development over the past twenty years. These and
other conductive polymers such as polythiopene have been used in a variety of
applications ranging from sensors to capacitors to light-emitting diodes and batteries
41-46.
Additionally, the ability to alter properties of the polymer by incorporating dopants into
the polymer matrix creates more applications for these materials
47,48.
ppy and other
conducting polymers can also be cycled from the neutral or insulating state to the
conductive or oxidative
Semiconducting and Metallic Polymers
through
state
the
applicationof charge.
The
/
cis-polyacetylene
trans-polyacetylene
poly(p-phenylene)
poly(1,6-heptadiyne)
conductive
polymers include a group
of
conjugated
hydrocarbon and aromatic
heterocyclic
polymers
N
including
poly(p-
N
H
H
poly(propriolic
anhydride)
polypyrrole
phenylene),
phenylene
poly(pvinylene),
poly(p-phenylene
sulfide),
PPy,
s
s
polythiophene
and
poly(quinoline)
Figure 1.2. Chemical structures of various conducting
polymers. From 12.
9
polythiopene. Polyphenylenevinylene (PPV) has a structure between polyacetylene and
polyphenylene. It has garnered interest because it can be developed in oriented forms 4 9.
PPy and polyaniline are two of the most common because of their high
conductivity, ease of modification, and stability
48,50-54.
Polyaniline is not only a very
stable conductive polymer, but it can also be modifiend through dopants and by altering
the pH of the medium in which it is deposited
55,56.
Both PPy and polyaniline can be
polymerized through electrodeposition which allows greater flexibility in the fabrication
of these materials.
1.2.2. Conjugatedpolymers
1.2.2.1. Electron classification
- antibondn
Binding
Electrons are classified into one of four
categories in materials.
essential for conductivity.
The
rgy
'rT-antbonding//
electrons are
Core electrons are
/(non-/b/ond/ing
/////
tightly bound to the nucleus and remain on the
elements nucleus with few exceptions.
a
electrons are found between two bonded nuclei
,
din
and are responsible for keeping the structure
Figure 1.3. Binding energy levels of
together.
s, n, and n electron states for organic
molecules From
Fro 33
molecules.
n electrons
are thought of with
heteroatoms (ie. O, N, S, P, etc) and have an
affect on the reactivity of a bond. nt electrons are involved in binding but form weaker
and less-localized bonds than the internuclear bonds of a electrons. The xrelectrons are
10
thought to be moving in the field created by
the nuclei and the other electrons,
E
and they
require the least amount of energy to jump to
the next energy state (See Figure 1.3).
lnsulalor
The
Semiconductor
double bond between the C units gives rise to
the conjugated bond. The
1t
the
atom
one
electron
per
conduction possible.
1t
electron serves as
that
Conducting
substance
makes
Polymers which contain
electrons are known as conjugated polymers.
Figure 1.4. Energy gap
representation of an insulator,
a semiconductor and a metal.
From 2
logo (S/em)
1.2.2.2. Band theory
PPy
PAN
(CH),
PT
Others
A
(a)
In
band
theory,
an
insulator
(b)
has
(e)
3
completely
filled
and
completely
empty
energy bands with a large gap between
its
2
(d)
1
0
.1
(el
.2
(I)
(~
(g)
(hI
(m)
(nl
(i)
(q)
(0)
(P)
(r)
(s)
{JI (1<1
.3
.1,
energy bands.
A conductive
material has a
.5
.6
(I)
(w)
.]
number
energy
of free electrons
band
(See
in an incomplete
Figure
1.4).
At
.8
.9
.10
(ul
(v)
(x)
(y) (z)
.11
.12
.13
temperatures
higher than 0 K, electrons can
-14
-15
jump into the higher energy bands creating the
possibility of conduction.
Depending on the
size of the electron gap, the material can be an
insulator
or a semiconductor.
In materials
with electron gaps of approximately
leY
or
11
Figure 1.5. The conductivity of various
conducting polymers at 24 °C. (a-e) forms
of [CH(b)]x, (f-k) forms of PAN, (I,m) PPy
doped with PF6, (n) Ppy (TSO), (o,p)
forms ofPT, (q) PPY (H2S04), (r) PPP
(AsFs), (s) 84Kr-implanted
Poly(phenylenebenzobisoazole),
(t-z)
undoped versions of the respective
polymers. From 2.
below, the number of electrons that are excited at room temperature becomes more
significant, and they are thought of as semiconductors. In general the conductivities of
semiconductors can range from 103 to 10-9 S/cm. Metals generally have conductivities
along the lines of 106 S/cm, and insulators are at the other end of the spectrum with
magnitudes around 10-22S/cm. Most semiconductors are inorganic, crystalline solids, but
conjugated polymers also display semiconductor properties without the inorganic,
crystalline structure.
Doping with an anion or cation can be used to increase the
conductivity of these polymers (See Figure 1.5)
2.
Doping for inorganic and crystalline
solids is slightly different than doping in conductive polymers.
For inorganic
semiconductors, the dopant is at the level of parts per million whereas with conducting
polymers, the dopant can form up to 50% of the polymer weight.
1.2.2.3. Band theory for conductive polymers
Rigid band models like those used in semiconductor physics are not completely
accurate for conductive polymer physics. At first it appears that the conductive polymers
have similar properties as inorganic crystalline semiconductors, but the movement of
electrons varies between these two types of materials. In conductive polymers, oxidizing
or reducing the material does not create free electrons or holes at the conduction bands.
This is because structural deformation occurs along the polymer backbone, where the
transfer of charge occurs, creating areas more likely to transmit charge.
Conductive polymers all have It-conjugatedsystems with alternating single and
double bonds along the polymer backbone. They are unusual because they can conduct
without having partially empty or filled bands.
12
For conductive polymers, when an
electron is excited from the valence band, a polaron is created. Unlike traditional band
theory, the hole that the electron leaves is not completely empty.
Instead, partial
delocalization takes place and results in a structural deformity from several of the
surround monomer units to balance the energy level created by the electron, thus
polarizing the nearby material which transforms into a new equilibrium condition; hence
the term polaron. In PPy, bipolarons form and the change effects approximately four
monomer units 57.
In the doping process, defects are generated that form radical cations or anions
which are also polarons.
A polaron consists of two defects:
a charged defect
accompanied by a neutral defect, also known as an ion and a radical.
Within this
framework, two types of conductive polymers exist: those with a degenerate ground state
(ie. trans-polyacetylene) and those with a non-degenerate ground state (ie. PPy). For the
polymers with degenerate
ground states, the initial
charge forms a polaron,
and a subsequent charge
b)
will
create
polaron.
polarons,
another
The
N
two
however,
H
c)
H
N
H
N
H
H
H
N
N
H
H
Deformatio
Coordinate
degenerate to form two
charge solitons. For nondegenerate
polymers,
Figure 1.6. A schematic of a) a polaron and b) a bipolaron on
a PPy chain. c) The bottom schematic represents the
deformation of the polymer lattice created from the charge
defects. From
16,17
however, solitons are not
13
formed with two charges, but pairs of defects are created called bipolarons. With the
degenerate polymers the energy level for the distorted state is equivalent to that of the
original structure. For the non-degenerate materials the energy level of the distorted state
is not equivalent which causes the bipolaron to form with two charges to maintain
balance
7.
For non-degenerate systems (PPy, polythiophene, etc.) at low doping levels,
charges are stored as polarons and bipolarons. The polaron is the radical cation or anion
accompanied by the lattice distortion resulting from the charge (See Figure 1.6). PPy is a
highly disordered polymer with as many as one defect for every 3 rings 17. These charge
defects create an electrically conductive partially filled band. Bipolarons are formed
when two polarons form on the same polymer chain. Another method of transport occurs
when the polarons and bipolarons hop to nearby chains to carry the electric current.
When PPy is oxidized and becomes more conductive, it is a polycation with many of
these delocalized positive charges on its backbone which are countered by dopant anions.
The function
of
polarons and bipolarons
can also be considered
with band gap analysis
(See Figure 1.7). When a
polaron is formed it forms
two
energy
states
from
the
center of the gap.
An
equidistant
+9
-
-
®
e
___
1////////////////1///IIiA
///
F/77///////7/I///
b e
a
Figure 1.7. Energy levels of a) soliton (neutral, positively,
negatively charged states); b) polaron (neutral, positively,
negatively charged states); c) bipolaron (postively and
negatively charged states). From 7.
14
electron or hole polaron can be formed with the electron polaron having the lower energy
state occupied by two electrons with opposite spins. In bipolaron formation, both energy
states are occupied by two electrons with opposite spin; or if it is a hole formation, both
energy states are empty. In the case of a soliton, only one energy state is formed in the
center of the gap 7.
To maintain conductivity valence electrons must move to conduction band
through the gain of energy. The product of the carrier mobility (), the charge (q), and
the concentrations of carrier (n) is the conductivity (a).
o=
*q * n
For conductive transport the polarons and bipolarons must be able to overcome the
energy barrier and hop from chain to chain.
This interchain hopping is the second
component of charge transport in conductive polymers.
1.3.
Polypyrrole
1.3.1. Background
Finding a conductive substrate with positive tissue interactions is an essential step
for advancing polymeric sensor and drug delivery designs.
Pyrrole is a 5-member
heterocyclic compound (See Figure 1.8). It can be found in heme and chlorophyll and is
produced by either of two production methods: 1) reacting furan with ammonia or 2)
dehydrogenation of pyrrolidine 58. PPy was first synthesized in 1916 where it prepared
by the oxidation of pyrrole to a powder known as "pyrrole black". In 1968 it was first
15
Figure 1.8. PPy with counter ions (A-) to balance charge. From .
electrochemically deposited 59. PPy (structure seen in Figure 1.8) is an electrodeposited
polymer that can be doped with various agents to alter its physical, chemical and
electrical properties
18,60-65.
Additionally, the properties of PPy can be controlled by
plating under various conditions
66-68.
One of the main advantages of PPy is its stability.
Its conductivity decreases only 20% a year in an unprotected environment. It can also
withstand temperatures of 100-200°C depending on dopant and is stable in acids 16
The ability to control PPy's surface properties such as wettability and charge
density creates the potential for modifying tissue interactions with the polymer 69. The
power to alter the properties of PPy through its dopants also adds versatility not seen in
other conductive polymers and makes it appealing for biomedical applications.
Additionally, PPy can be used to electrically depolarize neurons which has been shown to
modify signal transduction pathways and maintain signaling activity over time 70
providing another method to interact with surrounding neural tissue. PPy has been
studied extensively for biomedical purposes, and the ability to form an erodible form
makes PPy an attractive possibility for sensor and drug delivery applications
71,72.
The
erodible forms that have been developed further increase the scope of biomedical
applications including polymeric devices and neural scaffolds
applications, an erodible or stable form may be desirable.
16
72,73.
Depending on the
1.3.2. Conductivity
One property of PPy that can be modified by altering the dopant anion is
conductivity. Conductivity of PPy can range from those of insulators with almost no
conductivity (10-5 Q-Icm-I) to 100 n-Icm-I
17.
Doping ions help to decrease the band gap
between the energy levels (See Figure 1.9). PPy is conductive because of the ability for
electrons to hop along the polymer chains and across interchains due to the x-conjugating
bonds. By using smaller counter anions with coplanarity with the polymer chains, the
conductivity can be increased
47.
Dopants such as hydrogen peroxide, polyethylene
oxide, dodecylbenzenesulfonate, and salts containing transition metal ions have all been
used
64,65,68,74.
Studies have found that longer deposition times, lower plating potentials
and temperatures, and higher concentrations of monomer and electrolyte are favorable for
conductivity and stability
75.
Other modifications such as increasing the roughness of the
plating surface has also been shown to increase conductivity
The addition of water
U U
Figure 1.9. Band structure
representation of Ppy and how it is
modified with doping: a) no dopant, b)
intermediate doping level - bipolarons
are non-interacting at this point, c) 33%
dopant per monomer, d) 100 % dopant
per monomer. The material has
changed from an insulator with a band
gap of 4.0 eV to a semiconductor with
full doping at l.4e V. From II.
a)
17
47.
+
~
+
~
•
b)
c)
d)
into the electrodeposition solution also produces a more conductive polymer possibly
because water serves as a better proton scavenger than PPy in the solution 76.
Two of the most common dopants that are co-deposited with PPy are polystyrenesulfonate (PSS) or sodium dodecylbenzenesulfonate (NaDBS)
62.
PSS/PPy and
NaDBS/PPy polymers have been used in many applications ranging from actuators to
neural electrode coatings to neural substrates
63,73,77-79.
PPy's
attractive choice for sensor and drug delivery applications.
properties make it an
The ability to dope the
polymer with various molecules and stimulate it electrically creates novel methods for
drug delivery. The conductivity of PPy also creates the opportunity to sense specific
molecules by monitoring changes in the properties of PPy directly or remotely using
radio frequency (RF) technology.
By further exploring the ability to manipulate and
monitor the properties of PPy, more advanced interactions with its surrounding
environment can be achieved.
1.3.3. Deposition techniques
One advantage of conductive polymers is the myriad of methods to produce them.
PPy can be deposited through electrodeposition, chemical bulk polymerization, or vapour
phase polymerization. One method for chemical deposition is to use an organic solvent
such as m-cresol and deposit the polymer film of PPy doped with dodeclybenzene
sulfonic acid
80.
The coating is achieved through the spreading of the conductive
polymer solution onto the surface while the solvent phase evaporates to leave a
conductive coat. This has enabled PPy to be spin coated onto glass substrates with
controlled thickness
81.
Additionally, a dip coating technique has been developed where
18
a polymer containing electron acceptor/initiators
(ie. FeC13 , CuCI2) is placed on the
substrate prior to the vapor deposition. The dip-coated substrates were then placed in the
presence of a dry saturated vapour of the monomer pyrrole for varying periods of time,
and PPy deposits onto the substrate as a film 82. Other methods such as UV-photoinduced PPy formation and plasma polymerized formation have also been performed to
form PPy films 83,84
Electropolymerization of the polymer from a solution through a redox reaction is
another method for polymer production. For aniline and pyrrole chemical polymerization
can occur with the use of an oxidant such as Fe3 + ions or ammonium sulfate 85. Both
chemical
synthesis
and
electrodeposition
have
advantages
and
disadvantages.
Electrodeposition requires a conductive surface, but through the use microfabrication, the
shape of the polymer can be intricately controlled.
Additionally, electrochemical
polymerization allows for more accurate control of polymer thickness and morphology as
well as producing a more pure polymer 8.
For the electrodeposition technique, the pyrrole monomer is mixed in an aqueous
solution, and a potential is applied between the working and reference electrode. The
polymer can be applied with a potentiostatic technique (voltage held constant), a
galvanostatic technique (current is held constant), or potentiodynamic method(voltage is
cycled)
86.
The polymer forms on the anode as an oxidation-reduction reaction occurs.
The anodic oxidation results in a flux of charge and/or neutral species as the polymer
forms 87. Plated PPy remains on the electrode surface while subsequent deposition
continues which shows that the PPy is conductive enough to participate in further
monomer oxidation.
In electrodeposition, doping involves a redox reaction as the
19
polymer forms.
For conjugated polymers, polymers are oxidized or reduced at lower
potentials than the monomers 44.
During the electrodeposition it is important that the working electrode does not
oxidize as the polymer is forming onto it 59. Various groups have discovered methods to
more common metals (Fe, Al, Zn) and not disturb the electrodeposition 59,88. This has
permitted the deposition of PPy onto metals other than the typically inert electrode
materials such as Au and Pt. Techniques such as pulse profile electrochemical deposition
have also been developed to reduce the concentration gradient at the electrode surface
and to prevent diffusional mass transport hindrances to the electrode surface 89.
All of the techniques (vapour deposition, chemical deposition and electrochemical
polymerization) allow for the creation of thin PPy films; however, through the use of
electrodeposition, thick films can also be created. Film thickness is proportional to the
charge used to plate the polymer. It has been shown that the growth of PPy is under
electron-transfer control and self-quenching protonation of the monomer may occur 76.
Traditional electrodeposited PPy is insoluble and infusible because of strong inter- and
intra- molecular interactions and crosslinkings 45. All of these methods yield a stable PPy
film on the desired substrate.
1.3.4. Formation mechanism of electrodeposited PPy
The structure of PPy is a chain of mainly 2,5 coupled aromatic units. PPy is
classified as an aromatic polymer because electrons can cycle around the alternatively
double and single-bonded, ringed carbon structure.
PPy's polymerization proceeds
through an anodic oxidation. The exact mechanism of PPy formation is still debated, but
the basic paradigm is described here. The formation of the polymer begins with an
20
oxidation step creating a cation radical. The cation radical then joins other cation radicals
through a coupling reaction with two stages. The coupling occurs at the most reactive
sites of the cation radicals, the carbon atoms (See Figure 1.10). The first stage of the
coupling is the joining of pyrrole monomers to form dimer intermediates, and the second
steady state coupling reaction is the linking of the pyrrole monomer with oligomeric and
polymeric pyrrole species
8,90.
The monomers are linked by eliminating the two
hydrogens and linking the carbons from which the hydrogen was removed.
For every
three to four polymer rings there is an anion molecule to counter the positive charge
created by the polymer chain units. In summary, the formation of PPy occurs through an
initial oxidation step, followed by a coupling step and then the elimination of H+ ions
(deprotonation), and the cycle is repeated as the polymer forms. The pyrrole monomer
donates an electron to the anode, and the following series of chemical and electron
transfer
reactions
creates
the
polymer
at
the
anode
surface.
During
the
electropolymerization of PPy, the pyrrole dimer has a lower oxidation potential than that
of the monomer, and because of this, monomer units of pyrrole are coupled through
oxidation to the polymer chain. Additionally, as the polymer chain forms, approximately
every three monomer units creates a negative charge which is balanced by the
incorporation of anion dopants into the PPy.
Alternate electrochemical polymerization processes have been offered as well.
One states that the polymerization is begun with the loss of two electrons and a proton
from the pyrrole molecule, and the intermediate is dimerized by a neutral pyrrole
molecule resulting in the loss of a second proton (See Figure 1.1la) 91. Another possible
mechanism is that a cation radical reacts with a neutral molecule to form a cation dimer,
21
1a -
[7
+ e-
+
~~~2_ f11 +2H
H
H
H
H
NH+
--g
+
2H+
+e~H
H n+l
'~n
Figure 1.10. Most widely accepted proposed mechanisms of PPy
electrodeposition. From 2
b)
a)
'
(2a)
H
N
N
~N
NN
H
;I
H
N
H
H
H;
H
H
i
H
N
X
NI
I
+
i
1N+
H
N
\
H
I
I
I'1n~~~~~~~~
N
(2b)
e
H\
H
N
7
H
H
H
N+-N
H
N
H
H
H
Figure 1.11. Alternative proposed methods for PPy electrochemical polymerization. a)
Polymerization begins with the loss of 2 electrons and a proton. The intermediate is
then dimerized by a neutral pyrrole and loses a 2 nd proton. b) The cation radical reacts
directly with a neutral molecule. From 8.
22
and the cation dimer loses a second electron and two protons when forming the neutral
dimer (See Figure 1.1 b) 92. The fact that the first model (See Figure 1.10) corresponds
with the drop of pH observed during polymerization and is in agreement with the number
of electrons used in the reaction make it the most widely accepted paradigm 93.
1.3.5. PPy modification
Ease of modification through dopants has made PPy a popular material to
fabricate. Apart from the dopants used, electrodeposition conditions and alteration of the
chemical structure have resulted in various forms of PPy. The wide range of properties
that can be controlled by altering aspects of PPy formation make the polymer an ideal
material for many applications. Below are several fabrication factors that can affect the
characteristics of PPy.
The solvent in which the PPy is electrodeposited is important for chemical
properties.
Conductivity and mechanical properties can be altered by using various
solvents and altering the amount of water in the plating solution. Conductivity can also
be changed just by modifying the percentage of water. Although the mechanism is not
understood since the chemical compositions are identical, it is believed to correspond
with changes in the polymer unit chain length 93. The presence of water vapour during
electropdeposition decreases the conductivity of PPy by decreasing hopping of electrons
across the polymer 94
Film formation is also proportional to activity of hydrogen ions, anions and the
monomer concentration. It has been found that agitation of the solution can decreases the
23
plating rate because hydrogen ions on the surface might have a catalytic effect. One
group suggests that the anion forms an intermediate species with the pyrrole 9.
Another method for PPy modification is to functionalize the pyrrole monomer by
substituting a desired group in place of the hydrogen molecule. The addition of the new
group can have a great influence on the properties of the polymer. For example, if the
PPy is formed with N substitutions, the conductivity normally drops 5 to 6 orders of
magnitude. This is believed to be due to the fact that the substituition blocks PPy rings
from residing within the same plane 2
1.3.6. Charge transport
Conjugated polymers such as PPy have various oxidation levels that can be
manipulated through the removal of electrons (oxidation) or the addition of electrons
(reduction) by the application of a voltage.
For approximately every three pyrrole
monomers, a positive charge forms and a counter anion is incorporated into the polymer
to bring about charge neutrality. For large dopants such as NaDBS, when the pyrrole is
reduced (a negative voltage applied between the polymer and reference electrode in a
solution) cations from the solution flow into the polymer to neutralize the charge, and the
polymer expands. If the dopants are smaller anions, when the polymer is reduced the
negatively charged dopants will flow out of the polymer into the solution to equilibrate
charge. The ability to change the volume of the PPy has been used to create actuators 96.
It has been proposed that electrons move through PPy by two methods. The first
method involves movement of current across mobile charge carrying regions, polarons
and bipolarons.
Dopants ionize the polymer chain and create a polaron.
24
As dopant
concentration increases more polarons are formed and eventually bipolarons form.
Electrons can then move across a single chain by movement of these charge carrier
regions.
The second method occurs when bipolarons or polarons from two separate
polymer chains reside in the same plane. Electrons, then, travel from polymer chain to
polymer chain by interchain hopping, producing spinless conductivity 97. These two
methods, 1) transfer of polarons and bipolarons throughout a single polymer chain and 2)
hopping of electrons from chain to chain, account for transport of charge in PPy
17.
The
dominant method of charge transfer depends on the type and the concentration of the
dopant in the film
98.
The conjugated backbone of alternating double and single bonds
allows the charged species to move along the backbone, resulting in an electrically
conductive polymer.
1.4.
PPy applications
1.4.1. Technology applications
Several kinds of metallic/PPy (such as Cu-PPy) blends have been made for
possible technological applications 99. By forming stable PPy blends, conductive
polymers can be more widely used in areas ranging from digital displays to integrated
circuits.
Batteries have been one area of development of PPy.
The polymeric battery
operates by the oxidation and reduction of the PPy backbone. For a PPy/lithium battery,
the battery is charged by applying a positive voltage on the PPy with respect to the
lithum, and the PPy is oxidized. Anions are loaded into the PPy from the electrolyte and
at the same time lithium ions from the electrolyte are electrodeposited on the lithium
25
component. When the battery is used, the electrons flow from the lithium, and the ions
flow back to the electrolyte. These lithium ions pass through the load and go into the PPy
which reduces the oxidized PPy. The cycle can be repeated numerous times 00.
1.4.2. Biomedical applications
Over the past 20 years, biomedical applications of PPy have been more closely
studied. Because of its electrical properties, several studies have looked at its interactions
with neural cells
73,101.
The electrical properties of the polymer have allowed for use as
electrode modification for better tissue interactions
61.
PPy has been used as a sensor
through the immobilization of biomolecules on the polymer surface
102.
Additionally, the
pyrrole monomer, itself, has also been modified to allow for more stable sensing
applications
103
1.4.2.1. Biocompatibility
Biocompatibility studies of various conductive polymers have been performed.
Subcutaneous studies of polyethylene and polyaniline films in Sprague Dawley rats have
shown no provoked inflammatory responses
104.
The tissue reaction to PPy-coated
polyester fabrics has been looked at through subcutaneous implantation in SpragueDawley rats. No large inflammatory response was observed 105. PPy composite materials
have also been studied for biocompatibility. Alkaline and acid phosphatase secretion,
staining of macrophage, and histology of the tissue were the common methods for
examining the tissue response to PPy.
Subcutaneous implantation of PPy/poly(DL-
lactide and PPy/poly(DL-lactide-co-glycolide)
26
resulted in no more inflammatory
response than PLGA alone, which is one of the most widely used biodegradable polymers
in biomedical research 101. The biocompatibility studies have not examined stand-alone
PPy in neural tissue which is one area that needs to be addressed before neural
applications can be developed.
One advantage of PPy is that biodegradable forms have been developed. In one
form of erodible PPY, functional groups were added as a side chain
72.
There has also
been a pyrrole modified with a thiophene for stability with three pyrrole units joined by
ester linkages and a aliphatic linker 06. The formation of biodegradable PPy is desirable
to avoid chronic inflammation and so that a device will not remain in the body after the
desired application is performed.
1.4.2.2. Electrical interactions
PPy has been studied as a possible substrate for cultured cells.
Through
stimulation, it has been shown that PPy can modify the DNA synthesis of adherent cells
69.
This allows for more control over the shape and function of cells.
Another
characteristic that can be modified to enhance tissue interactions is the dopants used
during PPy formation.
By using biomolecules as dopants (laminin nonapeptide,
CDPGYIGSR, and fibronectin fragments, SLPF), better in vitro adhesion to the
electrodes was obtained 5. The ability to control the surface and composition of the PPy
surface could help prevent encapsulation - the formation of a non-conductive organic
layer.
The effect of dopants to promote or deter cell growth is one area that could
provide more insight for biomedical applications of PPy.
27
PPy has been utilized as a coating to obtain better tissue interactions.
its biocompatibility
Because of
and surface texture, it has been coated onto microfabricated
probes to enhance the recording signal (See Figure 1.12)
5,79.
neural
The increased surface area
that the polymer provided helped to decrease the impedance of the electrode and facilitate
neural recording.
Polystyrene sulfate (PSS) has been a common dopant because of its
good biocompatibility
and physical properties of the film
neural microelectrodes
and neural tissue is important to obtain long term recording from
73.
The interface between
neurons.
Figure 1.12. PPy coated on recording sites of microelectrode arrays to enhance
tissue interactions. Various thicknesses have been used to try to find optimal
recording impedance. Thickness corresponds to total charge delivered. From 5.
Electrical stimulation is a method of interacting with cells.
electrical
controlled.
69,73,110
stimulus, cell migration
101,
maturation
108,
By application of an
and DNA synthesis
109
can be
Work has also been done to utilize PPy to modify cells and their behavior
through the application of an electrical stimulus.
shown to enhance neutire outgrowth
Stimulation of the PPy has also
with possible nerve regeneration
28
applications
13.
The ability to dynamically interact with the surrounding environment is a major
advantage when using conductive polymers.
1.4.2.3. Nerve guides
Nerve guides made from synthetic materials have been used to help promoted
nerve growth after nerve injury. Current surgical techniques for nerve repair include
using a sacrificial replacement nervee to repair the damaged nerve. Because of PPy's
electrical conductivity as well as the ability to manipulate its characteristics with various
dopants, it has great promise for enhancing restoration of lost nerve function.
Groundwork for the use of PPy in the repair of nerve injuries began with the study of
neurite outgrowth enhancement of PC-12 with PPy stimulation 73. Because the polymer
itself is conductive, electrical stimulation can occur at a localized site instead of a more
1A
E
M
rue
e·...
.e
1B
1C
e
..... .....
'..,...
.
E = Electrode, M = Membrane, S = Solution, e = electron,
Charged. cation
E = Electrode, M = Membrane, S - Solution, e = electron, G: Charged cationic
polymeric site, 0: Neutral polymeric site,
: Counter-anion (anionic drug), I: Reagent
analyte or chemical trigger, 0o : Product analyte
Figure 1.13. An example of a PPy membrane that can be used for A) chemical
sensing B) electrochemical release of a drug and C) chemical release of a molecule
activated by a reactive analyte. From 4.
29
generalized area.
A silicone tube lined with a PPy membrane has been developed as a nerve guide
'101,1. However, development of a totally degradable polymeric conducting tube would
prove more useful for nerve guidance because of the greater flexibility to modify the PPy
tube and, potentially, the eventual degradation of the entire scaffold.
1.4.2.4. Drug delivery
The ability to force dopant molecules out of the polymer allows for possible drug
delivery applications with PPy 67. Dopamine has been delivered by utilizing this property
of PPy
112;
properties.
however, to date, drug delivery applications have been limited by dopant
Additionally, membranes consisting of PPy have been developed for the
release of adenosine triphospate (ATP) (See Figure 1.13) 4.
The molecules being
delivered still have the size limitations because they have to be incorporated into the
polymer. A method that does not rely on dopant characteristics for drug incorporation
will provide a more general platform for molecule delivery from PPy substrates.
Additionally, the electrical properties of the polymer allow for external interaction to
control drug release.
30
1.5.
Specific aims
Many developing in vivo technologies depend upon the ability of a device to
effectively interact with the surrounding environment. Controlling these polymer-tissue
interactions is an essential component of sensing and drug delivery technologies.
Developing a biomaterial that can be modified to facilitate integration into its proximate
settings opens the door for more effective biomedical devices.
One such material, PPy, is a conductive polymer with unique charge carrying
characteristics that make it an ideal material for biological applications. Previous studies
have examined PPy and its interactions with surrounding tissue.
It has been used to
enhance neuronal interactions with electrodes as well as to stimulate neuronal growth.
The following thesis work addresses the following areas of research:
1) Development of a new method to effectively micro-pattern electrodeposited
polymers and metals for in vivo devices
2) Determining the optimal synthesis conditions of the conductive polymer, PPy, for
sensor and implant applications
3) Fabrication of PPy tubes to be used as nerve guides to promote nerve regeneration
4) Modification of PPy for neurotrophic factor drug delivery devices
31
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44
2. Patterning
2.1.
PPy growth
In order to create polymeric, implantable devices, it is necessary to be able to control
the shape of the polymer. PPy has been used in electroplating applications for many
years 90. Various methods have been used to control PPy growth
113.
Because PPy is
electrodeposited, the shape can be controlled by patterning the counter electrode upon
which it is plated. The polymer grows vertically and horizontally with different time
constants in the different directions
67,
and this property can be harnessed to control the
shape of the polymer that is deposited. Microfabricated templates with specified gaps in
the conductive traces can be used to obtain the desired form. The ability to control the
shape of an electrodeposited material creates many possibilities.
The exact mechanism of PPY electropolymerization is still being investigated, but
generally, the polymer is formed through the oxidation of the pyrrole monomer followed
by a series of electron and chemical transfer steps leading to the incorporation of the
doping molecule and the formation of the polymer 8. Under the control of the electron
transfer, PPy electropolymerisation occurs in a three-dimensional (3-D) nucleation and
growth pattern on the plating electrode.
The ability to control the surface morphology is another important aspect for
tissue interactions. Through electropolymerization of PPy, the surface characteristics on
the cellular level can be manipulated through conditions of the plating environment such
as temperature and current density 75,114. Understanding the effect that changes in plating
45
current density and temperature have on film morphology is important for implant
design.
2.2.
Single-step 3-D electrodeposition
Current methods for making 3-D micromachined objects are difficult and
complex involving multiple masks and process steps. A new, simplified approach that
requires only a single mask would be a vast improvement over existing methods.
Through the use of traditional microfabrication techniques, a conductive template can be
designed and after subsequent electrodeposition, the 3-D polymer structure is produced.
Conductive traces of gold on an insulating surface are fabricated by e-beam deposition
and patterned using standard lithographic lift-off techniques. Gaps are left in the gold
template that determine the changes in height in the electrodeposited polymer. Because
the polymer grows both horizontally and vertically, as the polymer expands across the
gap it gains in height as well. Once the polymer spans the gap and forms a connection
with the new conductive trace, a new level of polymer will begin to plate while the
existing layer continues to grow as well.
In this manner, multi-level substrates of
polymer can be formed (See Figure 2.1). The gap spacing corresponds to the difference
in heights of each level.
The proposed technique can be applied not only for PPy
deposition but any electrodeposited metal or polymer.
46
u)
h)
c)
d)
Figure 2.1. Schematic of PPy electrodeposition technique (the same method for
nickel but the plating polarity is reversed). a) The plating begins at isolated
region(s) connected to the anode; b) the film grows both horizontally and
vertically from the initial conductive trace; c) over time the PPy bridges to new
conductive regions and plating continues over the larger surface; d) another
conductive trace is bridged with the relative heights of each region determined by
the spaces between the conductive regions and total plating time. Modified from 9.
2.3.
LIGA
Nickel plating is used for electroforming
acronym,
Lithography,
Galvanoformung,
and molding (known by the German
Abformung
(LIGA)).
L1GA is a common
process used to create multilevel 3-D structures through the irradiation
methacrylate
(PMMA) molds ] 15.
Forming these structures
of polymethyl
is a multi-step
task; and
therefore, nickel will also be electroplated to determine if the use of gaps in the template
can provide a direct method for producing these multi-level structures.
47
2.4.
Micropatterning
2.4.1. Microfabrication of template
A gold pattern was fabricated that formed the shape-determining conductive
template for electrodeposition. The layouts were generated with AutoCAD software, and
DXF files were converted to a chrome-on-glass mask (International Phototool Company).
The plating template was formed with a gold lift-off process on 4-inch silicon wafers
with 3000A LPCVD silicon nitride as an insulating layer. A Standard Lift-off process
was used to pattern the gold electrodes: photoresist was patterned onto the wafer, 100A
of titanium was deposited as an adhesion layer, after which 3000A of gold was deposited.
The photoresist was removed, leaving behind only the gold regions deposited directly
onto the wafer. The wafers were cut into die using a flood cooled die saw. The patterns
were protected using an additional layer of photoresist during die sawing; after sawing,
the individual die were cleaned in acetone, ethanol and then DI water before use.
2.4.2. Electrodepositionofpolypyrrole
The electrodeposition of PPy occurred at 250 C using a constant current power
supply (HP 6614C) in a two electrode configuration. The current density was 3 mA/cm2
to insure that a smooth film of polymer is deposited. The temperature and current density
were varied to find the conditions with the smoothest deposited films. Other conditions
can produce particularly rough surfaces. Solutions of polypyrrole (0.2 M) with NaDBS
48
dopant (0.2 M) were prepared
dissolution of the constituents.
more than 24 hours in advance
to ensure
They were stored at 4°C under nitrogen.
complete
A platinum wire
mesh cathode was used with an enclosed area equal to that of the patterned area of the
die. A mesh was utilized because a single platinum wire cathode resulted in non-uniform
deposition across the surface.
with moderate stirring.
The electrodeposition
occurred under a nitrogen blanket,
After deposition, the devices were ultrasonically
cleaned in OJ
water to remove loosely adherent particles.
2.4.3. Plating of nickel
The nickel was electroplated using published procedures
was the cathode for nickel deposition.
bath was a commercially
available
116,117.
The gold pattern
A pure nickel anode was used, and the plating
nickel
sulfamate
solution
(Mechanical
Nickel
Figure 2.2. Stepwise fatterns formed by the (a) electro-polymerization
ofpoly(pyrrole)
for 1 hour at ImA/cm and ~b) the electro-plating of Nickel from Nickel Sulfamate
solution 6 hours at 1mA/cm . Original gold patterns can be seen on the silicon nitride
background (purple). The fine pattern contains circles spaced 10Jlm apart, the coarser
spaced patterns are 50 Jlm apart.
49
Sulfamate, Technic Inc). Current density was regulated to 3 mA/cm 2. These die were
also ultrasonically cleaned in DI water after deposition.
2.5.
3-D structures
2.5.1. Polypyrrole and nickel electrodeposition characteristics
To test the hypothesis that an electrodeposited material can effectively bridge the
template gaps and form a multi-level 3-D structure, various test patterns were
microfabricated. PPy as well as nickel were plated, and the deposition properties of both
materials were compared.
Figure 2.2 shows the expansion of the electrodeposited
material front as it crossed a gap to connect surrounding circles. The two materials
showed similar properties in forming 3-D structures with the main differences being
nickel having a lower deposition rate (for nickel: 105 nm/min, both vertically and
horizontally compared to PPy with a growth rate horizontally 780 nm/min vertically and
1000 nm/min horizontally). For nickel, the equal growth rates in all directions result in a
radial expansion for nickel rather then a faceted surface as seen with PPy (See Figure
2.3). The surface of the deposited nickel was also much rougher.
Nickel LIGA wafers
are normally lapped to remove any roughness, but in multi-leveled structures it is not
possible to do so. Depending on the application, the rounded front of nickel or the
faceted front of PPY from 4:3 lateral-to-vertical growth ratio could be desired. If a more
vertical side wall is desired, a sacrificial boundary surrounding the conductive patterns
such as photoresist could be used.
50
l.
Figure 2.3. 3-D test patterns. The original circles are 200 Jlm in diameter, and 10 Jlm
apart. The decreasing height of each circle in the pattern is prominent. a.) Top view
of the PPy pattern grown at 3 mAlcm2 for 48 minutes. b.) Top view of the nickel
structure electroplated at 3mAlcm2 for 14 hours. c) Cross-sectional view of plated
PPy. d) Cross-sectional view of plated nickel. Modified from 9.
2.5.2. Polypyrrole structures
Because of the flexibility of this approach we have been able to form a variety of
structures including tapered lines, branched structures, and concave and convex features
(See Figure 2.3).
One application for this method is the production of molds for the
creation of microfluidics or microvasculature for tissue engineering.
Poly(dimethyl
siloxane) (PDMS) has been used extensively to create microfluidic devices from
51
Figure 2.4. Three dimensional structures created from a two-dimensional template.
This device is a template to cast a soft microfluidic vascular network. The gaps in the
original pattern determine the height of each section. a) Original electrode. b)
Resulting three-dimensional structure pattern - the smallest lines are 10 J..lmhigh, the
tallest are 80 f.lm. The deposition started from the left side ofthe image. Modified
from 9.
micromachined patterns
118-120.
By using this technique, a complex system of narrowing
microfluidic channels can be formed by using PPy to form the master pattern (See Figure
2.4). This method removes the necessity of relying on layer-by-Iayer formation of 3-D
structure as has been utilized in the past
121-123.
Additionally, by combining additional
layers and spacers formed by photoresist during the plating process, multilevel structures
such as electrodes, interconnects, gratings and photonic lattices could be formed. Tall
features such as mechanical barriers or sealing rings around a critical region of a device
could be produced as well. The NaDBS doped PPy is impervious to many etchants and
solvents used in microfabrication which could prove useful depending on the application.
Concave and convex patterns can be created by series of concentric patterns - the
curvature is a function of the spacing of the patterns, and whether the deposition initiates
from the inner or outer pattern (See Figure 2.5). For PPy, a series of finely spaced
patterns creates a more uniform deposition. This is the result of the electrodeposited
front being smoothed each time the PPy connects from one conductive trace to the next.
52
By varying the width of a line pattern, a 3-D structure can be produced that varies in
thickness and width (ratios of thickest and thinnest structures of 50: 1 have been
fabricated), or the pattern can be designed so the final structure maintains a constant
width while varying only in thickness. All of these parameters allow for flexibility in the
design of 3-D polymer implants or sensors for increased tissue interactions.
Figure 2.5. Circular 3-D patterns. a) A pattern forming a convex shape by a series of concentric
circles deposited from the inside to the outside. b) pattern forming a concave shape by a series of
concentric circles deposited from the outside to the inside.
2.6.
Morphology
2.6.1. Plating conditions
The morphology of a surface has a great influence on its tissue interactions. It has
been found that a rougher surface has better contact with surrounding tissue than a flat
surface such as gold
79.
One method for controlling PPy's surface characteristics is by
modifying the plating environment. To gain a better understanding of the effect that
53
2 mAlcm2
15 mA/cm2
Figure 2.6. PPy (O.2M) doped with NaDBS (O.2M) plated at various
temperatures and current densities.
54
temperature and current density have on the polymer, these factors were varied, and the
resulting films were characterized with a scanning electron microscope. Samples were
plated in the same manner as noted previously for 15 minutes at current densities of 2
mA/cm2 and 15 mA/cm2 with temperatures of 4 °C, 24 °C, and 44 C. A .2M PPY/.2M
NaDBS solution was used for electrodeposition.
2.6.2. Polypyrroletextures
By increasing the current density, the driving force for PPy polymerization is
increased. This increase the rate at which the polymer forms, and as seen from Figure
2.6, it also increases the granularity of the two high temperature conditions and changes
the texture of the 4 C sample. The temperature affects the solubility of the NaDBS
dopant in the aqueous solution. This may alter the availability of the dopant in solution at
the plating electrode's surface and cause the less compact film and the rougher surface.
At the low plating temperature, more nucleation points appear to be present producing the
more varied surface.
Overall, lower current densities and higher temperatures, with
NaDBS as the dopant, create a more compact and smoother surface. When the current
density is increased, more granular surfaces appear at the higher temperature, and larger
features occur at the 4 °C plating temperature. Additionally, the higher current density
produces a thicker polymer layer because more charge is passed through the solution in
the 15 minutes of plating.
The rougher surface characteristics produced by plating with higher current
densities or lower temperatures creates more surface area, creating a lower impedance.
This can be used to enhance electrode recording. Also, the intricate surface produced
55
with lower temperature as seen in Figure 2.7 may allow for a more intimate interface with
cells or surrounding tissue.
Figure 2.7. High magnification
of textured surface created by
plating with a current density of
15 mA/cm2 at 4°C. A rougher
surface may lead to greater
tissue interactions. Scale bar:
20 ~m
56
2.7.
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Cui, X., Hetke, J. F., Wiler, J. A., Anderson, D. J. & Martin, D. C.
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58
3. Biocompatibility
3.1.
Polypyrrole/Tissue interaction
As neurodegenerative diseases become a more pressing concern in society, the
need for effective treatment methods increases. Therapeutic possibilities range from
electrical interactions with the damaged neuronal circuits to the use of stem cells to
replace injured tissue
124-126.
One challenge is finding biocompatible materials that
effectively interact with neural tissue for these applications.
The stability and
biocompatibility of different polymers have been studied by examining their effect on the
surrounding tissue after implantation
29,127-130
Various conducting polymers have been examined for use in biomedical applications
71,72,85.
ppy has emerged as a promising candidate material that has been effective as a
coating in both in vitro and in vivo neural studies
5,73,79,131.
PPy also has shown promise
l.
as a scaffold material for nerve regeneration
PPy is an electrodeposited polymer that can be doped with various agents to alter
its physical, chemical and electrical properties
61-63,94.
The ability to control PPy's
surface properties such as wettability and charge density creates the potential for
modifying neural interactions with the polymer
that
are
co-deposited
with
PPy
are
69.
Two of the most common dopants
polystyrene-sulfonate
(PSS)
or
sodium
dodecylbenzenesulfonate (NaDBS). PSS/PPy and NaDBS/PPy polymers have been used
in many applications ranging from actuators to neural electrode coatings to neural
substrates
5,63,73.
Another strength of PPy is that erodible forms have been developed
which increase the scope of biomedical applications including polymeric devices and
59
neural scaffolds 71-73. The ease of deposition and the ability to control growth in both the
horizontal and vertical dimensions 9 enables flexibility in the three-dimensional design of
polymer implants.
To insure biocompatibility with standard PPy (those doped with NaDBS and
PSS), in vitro studies have been performed with both endothelial cells and with neurons.
Past groups have also shown that various proteins and molecules can be incorporated into
PPy, but the effect on neural cells has not been studied 5,132-134. To study this further, PPy
has been doped with various pro-neuron (laminin, NGF, BDNF) and anti-neuron (PEG,
Poly-lysine) molecules as well as other charged salts to determine the growth of neurons
on the conductive polymer.
3.2.
Experimental overview
The following in vitro and in vivo studies show the ability of PPy to interact with
neural tissue from the mammalian cerebral cortex. The biocompatibility of the PPy
implants is compared to sham stab wounds (where an implant-sized incision is made with
no implant left behind) and Teflon implants with similar size and features, and these
results demonstrate the positive surface interactions at the PPy implant-cortical interface.
Bovine endothelial cells have been used to test the biocompatibility of PPy with
standard PSS and NaDBS dopants. A live/dead assay ascertained the viability of the cells
after a week of culture with NaDBS and PSS dopants.
Following these studies,
dissociated primary cerebral cortical cells were cultured on the various PPy samples, and
fluorescent labeling was used to distinguish glial growth (GFAP), neuronal growth
(MAP2), and synapse formation (VGLUT) and to assess cell viability.
60
In vivo testing followed using the most promising PPy dopants from the in vitro
studies. Fabricated PPy implants along with Teflon implants were surgically placed in the
cerebral cortex of the rat. The implants were designed with several apertures so that
neural growth through the polymer windows could be observed. Immunofluorescence
was used to study the tissue surrounding the PPy implants. Quantification of the intensity
and extent of gliosis at 3- and 6-week time points was analyzed to determine the
biocompatibility of the PPy surfaces.
3.3.
Electrodeposition
After the templates were cleaned, various forms of PPy (Aldrich Chemicals) were
electrodeposited onto the gold surface using a constant-current power supply (HP
6614C). A current density of 1 mA/cm2 was applied between the gold template and a
platinum wire mesh reference electrode. The electrodeposition chamber was perfused
with N2 5 minutes prior to the start of deposition as well as throughout the
electrodeposition
process.
By varying dopant composition and electroplating
temperature, multiple types of culture substrates were made for the in vitro studies, and
five types of implants were made for in vivo studies.
3.3.1. In vitro samples
For the in vitro testing, each dopant was mixed with a .2 M PPy solution both
with and without .2 M NaDBS. Each dopant combination was also plated at 4 C and 24
°C (See Appendix for table of dopants and biocompatibility results).
61
Additionally,
varying concentrations (2 mg/ml, 0.2 mg/ml, 0.02 mg/ml) for the majority of dopants
were tried to determine the threshold for the effect of the dopant. Laminin peptides
(MIT, Biopolymer lab), NGF (Santa Cruz), BDNF (Santa Cruz), poly-lysine, PEG, polyethyl, poly-aspartic acid, and sodium acetate (all Sigma-Aldrich unless otherwise labeled)
were used with or without 0.2 M NaDBS. 0.2 M NaDBS and 0.2M PSS were also used
as dopants by themselves to compare to other work.
Table 3.1. Implants for in vivo experiments
ImplantType
3 week
6 week
Stab
4
4
Teflon
4
4
4 0 C PSS/PPy
4
4
240 C PSS/PPy
4
2
4 0 C NaDBS/PPy
4
240 C NaDBS/PPy
4
240 C PSS/PPy in PBS
2
4
3.3.2. In vivo samples
The in vivo studies used aqueous solutions of 0.2M PPy plus 0.2M PSS (Aldrich),
and 0.2M PPy plus 0.2M NaDBS (Aldrich). Surface texture was controlled by varying
the temperature during electrodeposition: 4°C was to create a more macroscopic/course
surface, while 240C was to create a fine-textured surface. Finally, a fifth formulation,
0.2M PPy plus 0.2M PSS in PBS, was electrodeposited at 24°C to create the fifth type of
implant to evaluate solvent conditions on the electrodeposition product (See Table 3.1).
62
3.4.
In vitro study techniques
3.4.1. Bovine endothelial cells
Bovine endothelial cells were plated onto 0.2M NaDBS/0.2M PPy
and 0.2M
PSS/02.M PPy samples in a 6 well plate (VWR). The cells were cultured in Dulbecco's
Modified Eagle Medium (DMEM) at 37 °C for 3 days and then stained in the following
manner with a Live/Dead Viability Cytoxicity Kit (Molecular Probes). On the third day,
media was removed from the cells and the cells were washed with sterile PBS. The cells
then were placed in a solution of 1 ,gM Calcein AM, which enters living cells, and 1 igM
Ethidium Homodimer-1, which enters cells with disrupted membranes, and incubated for
30 minutes at room temperature. The cells were then washed with sterile PBS and left in
PBS for imaging.
3.4.2. Cortical cell harvesting
For the in vitro experiments, the polymer remained on-chip during the
experiments.
All animal procedures were performed in accordance with protocols
approved by the MIT Committee on Animal Care (IACUC and conformed to NIH
guidelines. All reagents are from Sigma-Aldrich (unless noted otherwise). Dissociated
cortical neurons were placed on lx2.5 cm squares of the various types of PPy with four
samples of each substrate being tested. Brains were removed from 1-3 day old SpragueDawley rat pups (Charles River).
The cerebral cortices were dissected out in 80%
Ca/Mg-free Hank's balanced salt solution (HBSS) containing NaHCO3, (4mM), HEPES
(5mM) and 20% fetal bovine serum (FBS). The meninges were removed, and the cortex
63
was cut into millimeter sections with a scalpel. The cortex was washed 3 times in HBSS
and incubated for 5 minutes at 37 °C in digestion medium (trypsin type XI (5mg/ml),
DNase type IV (0.5mg/ml), 137 mM NaCI, 5mM KCI, 7mM Na 2HPO 4, and 25mM
HEPES, pH 7.2).
Trypsin was neutralized with FBS.
The cells were chemically
dissociated in 12 mM MgSO4 *7H 20in HBSS containing DNase type IV (0.5mg/ml) and
then physically dissociated by triturating through 2 glass pipettes of decreasing size. Cells
were harvested by centrifugation (1000 rpm, 4 °C, 10 min); plated onto the various PPy
substrates in plating medium (90% of 28mM Glucose 2.5mM NaHCO3, lmg/l0ml of
transferrin (Calbiochem), 30mM glutamine, 0.73mM HCI, and 2.5 mg/100ml in
Minimum Essential Medium with Earle's salts, without L-glutamine or phenol red
(MEM) (Gibco) and 10% FBS); and, cultured at 370 C with 5% CO2 . The PPy culture
substrates were sterilized using ultraviolet light exposure for 12 hours prior to the
addition of cells. After 24 hours, plating medium was replaced with feeding medium
(28mM Glucose, 2.5mM NaHCO3, lmg/l0ml
of transferrin (Calbiochem), 30mM
glutamine, lml/100ml B27 50x supplement (Gibco), and 0.84mM cytosine arabinoside in
MEM).
Cultures were fed weekly.
After 21 days in vitro, culture were fixed and
prepared for immunofluorescence.
3.4.3. Immunofluorescence
The cells were fixed in 4% paraformaldehyde in PBS for 20 minutes; washed 3
times in PBS; permeabilized in 0.2% Triton X in PBS; the cells were again rinsed in PBS.
A blocking solution of 5% goat serum (Vector Labs) in PBS was applied for lh. The
primary antibodies [guinea pig anti-vesicular glutamate transporter 1 (VGLUT1) 1:500
64
(Chemicon); mouse monocolonal anti-glial fibrillary acidic protein (GFAP) 1:100
(Sigma); and, rabbit anti-neuronal class P3-tubulin 1:250 (Covance)] were diluted in PBS
with 5% goat serum and left at 4°C overnight. The cells were rinsed three times for five
minutes each in PBS and the secondary antibodies [Alexafluor 647 goat anti-mouse IgG
1:500; Alexafluor 546 goat anti-guinea pig IgG 1:400; and, Alexafluor 488 goat antirabbit IgG 1:500 (all Molecular Probes)] were diluted in PBS with 5% goat serum, and
applied for 30min. DAPI was added to label cell nuclei. After thorough rinsing, the
samples were mounted on slides for imaging.
3.5.
In vivo study techniques
3.5.1. Implant release
In vivo experiments required the PPy to be released from the chip before
implantation.
Stand-alone polymer implants provide the advantages of flexibility and
potential biodegradability.
The implants were released using a variety of methods
depending on the PPy dopant.
The PSS/PPy implants of both temperatures can be
removed from the gold template by a gentle mechanical force. The removal of the
NaDBS/PPy implants required chemical etching. The silicon nitride was etched by a
6:100 mixture of Fluoroboric Acid:Phoshoric acid at 105 C (US patent number
3,859,222) for 12 hours. Upon removal of the silicon nitride layer, the PPy implants and
the template die were placed in KOH. After approximately h the PPy implants floated
off of the template or were removed by a gentle mechanical force. After removal from
the template, the implants were separated by a razor into individual implants and soaked
in 4 separate baths of filtered deionized water for 1h each.
65
3.5.2. Surgical implantation
We
shaped
fabricated
PPy
specially
implants
with
dimensions of approximately 2 x 3 x
0.25 mm (sufficiently
long to span
the cerebral cortex and sufficiently
thick
to
be
manipulated
during
surgery) using PSS and NaDBS as
PPy
dopants
implanted
them
and
surgically
into
rat cerebral
cortex (See Figure 3.1).
also
contained
Implants
three,
500~m-
diameter apertures to permit neural
tissue to cross the lesion through the
implant.
Each side of the cerebral
cortex received one implant.
monomer
Figure 3.1. PPy implants a.) An example ofa
typical PPy implant. Scale bar: 1mm) b.) Two
PPy implants placed in the rat's cortex. Scale bar:
2mm c.) A histological slice at 6 weeks postimplantation with the remnants of the PPy
implant. Scale bar: 200um. From 10.
One of the implants was fabricated to determine
solvent (H20 or PBS) had any affect on the implant properties.
(McMasterCarr)
implants with this same design (including
implanted into the rat cerebral cortex.
the apertures)
if the
Teflon
were also
Finally, stab wounds were performed as a further
control in lieu of inserting an implant.
All procedures
were performed
on 150g male Sprague-Dawley
River) with sterile technique under a surgical microscope.
rats (Charles
Implants were sterilized with
ethylene oxide. Anesthesia was induced with 3% isoflurane with oxygen (lVmin) for 10
66
min and maintained with 1% isoflurane with oxygen for the duration of the procedure (20
min). The anterior scalp was incised along the midline. A 1.5 mm-diameter craniotomy
was made 2mm posterior to bregma and 4mm from the midline on each side. Implants
were inserted with fine forceps into the cerebral cortex. The scalp was closed with
Ethilon suture. Once fully alert, rats were returned to home cages. Buprenex (0.1 mg/kg,
i.m.) was given for analgesia 2 times daily for 2 days post-implantation.
3.5.3. Immunofluorescence
At the 3- or 6- week time-points, rats were terminally anesthetized and perfused
with PBS and 4% paraformaldehyde. Brains were dissected from the skull, postfixed in
4% paraformaldehyde overnight at 4C, and then transferred to (30%) sucrose in PBS
until they sunk. Brains were frozen with dry ice and sectioned at 50gm.
Sections were processed free-floating for immunofluorescence.
Blocking and
antibody incubation steps were performed in 0.5 % triton X-100 in PBS with 5% donkey
serum. The sections were incubated overnight in the primary antibody (rabbit anti-GFAP
1:100, mouse monoclonal anti-microtubule associated protein 2 (MAP2) 1:1500, mouse
anti-macrophage 1:1500, mouse anti-laminin 1:1500) solution at 4°C. After three washes
in PBS for 10 minutes each, the secondary antibodies (Alexafluor 594 donkey anti-rabbit
IgG 1:400 and Alexafluor 488 donkey anti-mouse IgG 1:400) were added to the sections.
After incubation of 1 hr in the secondary solution, the sections were rinsed 3 times in
PBS and mounted on slides for imaging.
67
Figure 3.2. Bovine endothelial cells on Ppy. a) Bovine cells stained with DAPI and
FITC dyes are contluent on PPy and silicon nitride substrate b). Live/dead assay of
bovine endothelial cells on NaDBSIPPy substrate. Circles are the only dead cells.
3.6.
In vitro study results
3.6.1. Bovine endothelial cells
Bovine endothelial cells grew to contluency on the PPy films with PSS or NaDBS
as a dopant and at 4 °C or 24 °C (See Figure 3.2a). The live/dead viability test showed
almost all living cells after 3 days with <1% dead cells remaining on the polymer (Figure
3.2b shows a typical surface).
Based on the compatibility of PPy with endothelial cells,
neural cells were then examined.
3.6.2. Dissociated cortical neurons
3.6.2.1. Polypyrrole resistivity
To examine biocompatibility
with standard PPy (doped with NaDBS or PSS), in vitro
studies were performed with dissociated cortical neurons.
were used to produce differing surface characteristics
68
Different plating conditions
of the PPy films as described in
Chapter
2.
We
varied
the
plating
temperature to modifY the texture of the
PPy surfaces. The surface texture of
PPy/NaDBS
electrodeposited
at
4°C
appeared course and irregular, while at
25°C, a smoother surface was obtained as
noted in the last chapter (Fig. 3.3a,b).
There was a similar but less dramatic
effect of temperature on the surface
texture of the PPy/PSS samples, but both
4°C- and 25°C-samples appeared relatively
smooth and were comparable to the
PPy/NaDBS-25°C sample (not shown, see
Fig. 3.3a).
Using four-point impedance
measurements,
the
resistivity
of
the
different PPy samples on their gold
Figure 3.3. PPy surfaces. The surface texture
of the PPy can be controlled through plating
conditions from a.) smooth, PPy/NaDBS plated
at 24°C b.) to rough, PPy/NaDBS plated at
4°C. Scale bars: 200mm c.) Fluorescently
labeled explanted cortical neurons growing and
forming networks on a PPy/NaDBS surface
after 21 days. Scale bar: 50mm. Green:
neurons. Red: glia. Blue: nuclei. From 10
templates was measured (Table 3.2).
NaDBSIPPy had lower resistivity than the
PSS/PPy samples (p < 0.05, n
respective
indicating
=
temperatures,
greater
surface
4) at the
possibly
elaboration
and/or greater intrinsic conductivity of the
PPy polymer doped with NaDBS.
69
3.6.2.2. Bioactive dopants
In vitro tests were performed to first study the response of cortical cells to PPy
interactions.
Complex neural networks, whose cellular components (glia, neurons, and
synapses) were identified by immunofluorescence
biocompatibility
on the various PPy mixtures.
was determined to be neuronal network formation (neurons and glia
with synapse formation) on the polymer sample.
For poor biocompatibility,
the sample
had no neurons after 21 days or had a minimal number of glial cells remaining
Appendix for full results).
PEG, poly-aspartic
(See
The basic 0.2M PSS/ 0.2M PPy and 0.2M NaDBS/ 0.2M
NaDBS samples were found to have good biocompatibility
biocompatibility.
Good
acid, sodium acetate,
in all conditions.
and even the laminin
peptides
Poly-lysine,
had poor
When NaDBS was mixed with these various dopants (ie. laminin
nonapeptide), it was ab Ie to mask the effects of the dopants and produce neuronal growth.
This might be due to the fact that the NaDBS was more easily incorporated
during
Figure 3.4. Scanning electron microscope images of poly-lysine doped-PPy. a) The polylysine dopant causes a rippling effect of the polymer surface b) A magnified view of the
surface showing round projections.
70
electrodeposition and outcompeted the other dopant and it reduced the amount of the
other dopant present so that the cytotoxic effects could not be appreciated.
The
neurotrophic dopants (NGF and BDNF) were found to have mixed results which could
result from too much of the factor being present or a modification to the factor during
plating. The morphology of the polymer varied depending on the dopants as seen in
Figure 3.4, where poly-lysine (2mg/ml) is plated with the 0.2 M PPy at room
temperature.
Table 3.2. Average Resistivity of PPy samples (n=4).
Dopant
Temperature (C)
Resistivity (lcm)
NaDBS
4
15.5
NaDBS
25
22.4
PSS
4
22.6
PSS
25
38.6
Neuronal circuits formed on both the PSS/PPy and NaDBS/PPy culture substrates at
both temperatures. Figure 3.3c shows a culture that is representative of what was seen on
the 4°C and 24°C PSS and NaDBS samples. Similar neural networks grew regardless of
its surface structure or if the dopant was PSS or NaDBS. Thus, neurons are capable of
extending axonal and dendritic processes, and of forming putatively functional synapses,
on these PPy substrates examined in vitro.
Based upon these positive PPy/neuronal
interactions, in vivo studies began using PSS and NaDBS as dopants (See Table 3.1).
71
Cell Nu:lei
"
... / -,.,;.. ;'
Stabwound 3 week
. ,f. '- •..
."
.
~,~
.
.~
".'
Stabwound 6week
.
-
I
..
~. :.;
..
i
Macrophages
Neurons
Glia
;
" . i.
"
.
..
.
;.
,~
..
..
:-
I._
.
,
I
~~.;(}.
,1
.'" ~
~ ."..
Tefbn3week
~
~
...
...
Tefbn6week
I
'.
t
•
.
~~I
Po Iypyno Ie 3 week
,
. (.\
Po Iypyno Ie 6 week
,., J
.
"
'
.
>
t
•
,
"
7?~
.
)
: «\'i'.
,
-.,..
'U'~~.I'
1
\
Figure 3.5. Representative fluorescently labeled sections of3 and 6 week time points with
the various types of implants. All of the implants are on the right side of the images, The
arrow indicates increased macrophage presence in the 3-week Teflon implant. Scale bar:
200um. From 10
72
3.7.
In vivo study results
3.7.1. Implants
We examined the implants, stab
wounds and the surrounding cerebral
cortical
tissues
at 3- and 6-week
time points to determine the extent
of rejection or integration
between
the chips and surrounding
tissue,
as well
neural
as the effect
of
insertion per se (via stab wound).
Immunofluorescence
cell
bodies
surrounding
revealed
that
of
and
the
neuronal
synapses
implant
site
surrounding
tended
to envelope
Figure
3.5 shows
cortex
the implants.
a typical
stab
wound, Teflon, and PPy implant site
for the 3- and 6-week time points.
Staining
for macrophages
the expected
showed
increased presence of
Figure 3.6. A fluorescently labeled section of
neural tissue in an implant lumen. a) Neural tissue
in the lumen of the Teflon implants b) Neural
tissue in the PPy lumen where the glia has
reformed and neurons are present.
Scale bar:
IOOum, Green: glia. Red: neurons. From 10
this cell type around the implant site
73
at the 3-week time point (Fig. 3.5). At the 6-week time point all of the implants had little
macrophage activity, and in most cases, the implant was clear of macrophage activity
after 6 weeks. Overall, the neural response to all three types of surgeries appeared
qualitatively similar. The neural tissue tended to reform after the stab wounds leaving a
scar
demarcated
by
laminin
immunofluorescence.
Our
histological
and
immunofluorescence studies show that the neural parenchyma completely enveloped the
implants and was intimately associated with the surface of the PPy implants, including
the implant lumens, demonstrating a very high degree of tolerance of neural tissue for
PPy. The neural tissue tended to bridge more completely the PPy lumen than the lumen
of the Teflon implants (Figure 3.6).
3.7.2. Quantitativeanalysis
To quantify differences in the degree of tissue reactions to the various implants,
ImageJ software was used to create a single-pixel column region-of-interest (ROI)
running parallel to the lesion site. Columns varied in height according to the length of
the lesion in a given section. We plotted the average intensities of all ROIs moving
perpendicularly away from the lesion site versus distance from the lesion, until the
average intensity reached background (Fig. 3.7a). Gliosis, as measured from ROIs using
GFAP-immunofluorescence, was expressed as the peak intensities of the averaged
column ROIs and as the slope of the logarithmic regression of the ROI versus distance
plots. All intensity values were normalized to a background value that was obtained from
an area of tissue positioned away from the implant site.
74
Figure 3.7 shows a
comparison
of
gliosis
intensity and slope for the
various
wounds
implants.
had
Stab
the
least
amount of gliosis and the
sharpest decrease in gliosis
b
site in all cases, except for
,
4.50
extending from the implant
4.00
g
3.50
j
3.00
f
oCi ~50
the
gliosis
gradient
of
Q/
•
~ ~.oO
-:l1.50
PPy/PSS
electrodeposited
a:
1.00
r-
Pl'ak Intcnstiy of Gliosis
•
,
I
!
3 w.8k.
• 6 weeki
t
I
050
0.00
in PBS at 25°C at 6 weeks
nab
lIIflm
nadb •• adb. p .....
He
4C
lID
and the
peak
gliosis
at 25°C at 6
(p<O.05)
Gradient
!
which
~
~.2S
~
~.20
Q/
were as biocompatible
as
the stab wounds at these
~
\3l ~.lS
~
t
the
gliosis
than
the
Teflon
implant
3-week
(p<O.05).
At the 6-week time point
t
• +
f
I
!
0.00
3-week
time point had less peak
,. 3 w..k.~1
• 6 veeka
t
• lab
at
of Gliosis
~.lO
~.o5
time points. All of the PPy
implants
psS24C-pbs
~.n
~30
weeks
p ••• ~CC'
of
C
PPylNaDBS
c
clan tTVDe
tafbl
~. :~I).
p..... cp ••
lI1I plimtType
-"'CC'
pss
24C-pbs
Figure 3.7. Quantification of gliosis from implant site.
a) Typical slice with red line indicating a column (width:
I pixel) where the pixel intensities would be summed to
obtain an intensity value for that column. The red line
parallels the lesion site and moves away from the
implant site (in the direction ofthe arrows), an intensity
value was obtained for each column. b) Peak gliosis
intensity values and c) a gradient of the intensity values
were obtained from these intensity values at 3 and 6
week time points with indicated standard errors. From 10.
75
the differences had lessened, except for PPy/PSS at 40 C, which had a sharper decrease in
gliosis than Teflon (p<0.05), and PPy/NaDBS at 25°C had less peak gliosis than Teflon
(p<0.05). Also, PPy/PSS electrodeposited in PBS at 25°C had less peak gliosis as well as
a sharper decrease in gliosis than the Teflon implant (p<0.05) at the 6-week time point.
The PPy/PSS 6-week sample electrodeposited at 25°C in PBS has a greater decrease in
gliosis than the PPy/PSS and the PPy/NaDBS at 25°C electrodeposited in deionized water
and less peak gliosis than the PPy/PSS sample electrodeposited at 40 C in deionized water
(all p<0.05). The PPy/NaDBS sample electrodeposited at 40 C had less peak gliosis at 3
weeks than the same sample type at 25C (p<0.05) suggesting that the rougher surface
fosters greater implant integration with the surrounding tissue, and less gliosis.
These data taken together tend to indicate that surfaces that are highly inert and
relatively unreactive for the host parenchyma (e.g., Teflon) may achieve less physical
integration and more inflammatory response (e.g., gliosis) than surfaces (e.g., PPy) that
appear to be well tolerated by the host parenchyma and that foster intimate physical
interaction between substrate and tissue elements.
Because PPy is conductive, bio-
electrical circuits could be fabricated that integrate electrical and neural signals. PPy
could also serve as a tissue scaffold to support neural cells for placement into areas of
neuronal loss in injuries such as stroke and Parkinson's disease 135. Reducing the amount
of gliosis surrounding the implant should enhance the polymer's ability to interact with
the normal brain parenchyma.
76
3. 7.3. Bioactive dopants
Pilot studies have been perfonned to incorporate neurotrophic factors molecules
such as nerve growth factor (NGF) and brain-derived neurotrophic factor (BDNF) into
the polymer matrix (Fig. 3.8). Compared to Figure 3.6, more neural tissue is present in
the
lumen
of the
PPy
implants
containing NGF than those without
PPy.
Future studies will elucidate
whether
the
bioactive
promote
neuronal
molecules
adhesion
and
interactions with the PPy implants.
3.8.
Conclusion
We have demonstrated the
manufacture
of
three-dimensional,
free-standing PPy substrates that can
have
a
progressively
biocompatibility
positive
profile with eNS
parenchyma in vivo.
These results
support future investigations aimed at
using
PPy
in
the
design
and
manufacture of neural prosthetics that
Figure 3.8. NGF implants a) A stained slice
at 6 weeks post-surgery with neurons (blue)
and glia (green). Scale bar: 100 urn b) A
fluorescently stained slice at 6 weeks around
the implant site showing glial and neuronal
cell bodies (blue) with the neurons
highlighted (red). No severe gliosis is seen
surrounding the implant. Scale bar: 200um
Arrows indicate region where cells are
extending into the PPy lumen and black is
the space occupied by PPy implant. From 10.
77
are capable of integrating with CNS tissues based on specific chemical and physical
properties of the PPy polymer. Such prosthetics should enable reliable transmission of
external and internal electrical signals for significant postoperative periods. Moreover,
they may, if properly formulated, stimulate damaged neural tissues to repair and
reconnect.
78
3.9.
1.
References
George, P. M., Lyckman, A. W., LaVan, D. A., Hegde, A., Leung, Y., Avasare,
R., Testa, C., Alexander, P. M., Langer, R. & Sur, M. Fabrication and
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Lehle, K., Lohn, S., Reinerth, G., Schubert, T., Preuner, J. G. & Birnbaum, D. E.
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conducting polymers for potential biomedical applications. Angewandte Chemie
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82
4. PPy tubes
4.1.
Peripheral nerve regeneration
4.1.1. Peripheral nerve repair
In 1995, over 50,000 nerve repair surgeries occurred . Two techniques currently
exist for repairing peripheral nerve injury:
If the lesion is small (<4mm) the two
damaged ends can be sutured together. If the injury is larger and the nerve cannot be
sutured together without causing tension which results in further nerve damage, a nerve
autograph (most commonly the sural nerve) can be used to bridge the two segments 5
Unfortunately, functional tests of nerve reinnervation in rats have not been very
successful
6.
Clinical reports show that only 1-3% of patients with autograph repair of
the median nerve recover normal sensation after 5 years and only 25% regain normal
motor function
7.
Disadvantages of nerve autografts include 1) the need for another
surgical step, 2) loss of the original nerve graft function, 3) the low numbers of suitable
donor nerves, and 4) mismatch of donor nerve dimension with the transected nerve 8
Partial or total transection of a nerve leads to the degeneration of the axonal
segment.
Additionally, Schwann cells also degenerate in the distal segment of a
transected nerve; however, along with the new growth of axons, Schwann cells can
develop and release laminin to promote additional axonal outgrowth.
The axons of
peripheral nerves do have the capacity to regenerate under the proper conditions 910
Multiple factors influence the recovery of a peripheral nerve function after an injury: 1)
neuronal survival to reinnervate distal areas 2) axonal emission of growth cones and
regeneration through the connective scar 3) accurate reinnervation of distal targets such
83
as muscle and sensory areas 4) the stability of the spinal circuit involved with the
peripheral nerve control 1 . For surgical repair of neurotmesis (total nerve transection) it
is important to align the nerve fascicles. If the gap is sufficiently large, a bridge is
needed to prevent neuroma formation 6. Cell adhesion molecules, lipid carrier and
myelin proteins, extracellular matrix proteins and integrins as well as neurotrophins,
cytokine and other growth factors are expressed as the axons regenerate
12.
Producing an
artificial microenvironment that provides similar factors would help to enhance nerve
regeneration.
The shortcomings of nerve autographs for large areas of damage have led to the
design of tubular nerve guides. Nerve guides not only provide mechanical guidance and
support but also reduce invasions from connective tissue, are a simpler surgical procedure
than autograph placement, and allow control of an isolated repair site to create a more
ideal regenerative environment 7
4.1.2. Nerve guides
Nerve guides that promote regeneration utilize contact guidance to direct the
growth of the damaged nerve. The first work with nerve guides was done with natural
materials using decalcified bone as a guide tube in animal experiments '0. Nerve guides
have consisted of other natural materials such as blood vessels, muscle tissue and
collagen
12,13.
Additionally, chitosan tubes from crab tendon have also been used 4.
Many of these natural materials, however, do not have ideal properties (ie. they are too
weak or flexible) to serve as guides. Muscle tissue is not ideal because the nerves can
redirect into the muscle which causes a neuroma 15.
84
To develop a nerve guide
with more ideal properties,
degradable
polymers
non-
such
as
silicone rubber, acrylic polymers,
and
polyethylene
4
studied
been
Non-degradable
microelectrode
been
have
arrays
investigated
regeneration
have
for
also
nerve
(See Figure 4.1)
Non-degradable
hydrogels,
4.
such
Figure 4.1. Designs of various electrodes to
enhance nerve regeneration. From 2-4.
poly(2-hydroxyethyl
as
methacry late-co-methy I
methacrylate),
have been shown to promote nerve growth
16.
Because of their stable and
inert properties in vivo, silicone tubes have also been used as nerve guides.
shown that increasing the wettability
and the absorption
surface by carbon negative ion-implantation
acquire hydrophilic
properties
It has been
of proteins onto the silicone
improves nerve regeneration
by the creation of hydroxyl
functional
17.
Silicone can
groups on the
silicone surface.
One of the main problems with non-degradable
polymers is that these materials
would remain in the body permanently, which eventually leads to foreign body reactions
or can cause other mechanical
degradable
damage.
This drawback
led to the development
guides that would erode away after the nerve was regenerated.
materials have been used to form biodegradable
85
of
Many
nerve guides including poly(L-lactide-
co-s-caprolactone) 8,18,19,poly(phosphoester)
hyalunronic acids 6.
20,21,
gelatins 22 and benzyl esther of
All of these materials have shown various degrees of success;
however, functional recovery was still limited.
Other structures besides tubular guides have been designed to address nerve
injury. Degradable scaffolds have also been used to repair spinal cord injuries through
contact guidance 23. Braided systems of poly(L-lactide-co-glycolide) (PLGA) have been
fabricated that are less rigid than solid tubes of PLGA to enhance adaption to the
implanted environment 15. Degradable peptide scaffolds have also been studied for
neurite outgrowth and differentiation
24
4.1.3. Fillings for nerve guides
Axonal growth relies on more than contact guidance cues, and therefore, nerve
guides that emit various growth promoting molecules have shown enhanced nerve
regeneration. Modifications have been made to the nerve guides themselves and to the
lumens to incorporate bioactive molecules.
Lumens of these tubes have been filled with saline, supportive cells, or
neurotrophic factors to improve regeneration
5,15.
Multiple methods have also been used
to increase fibrin development in order to support nerve growth. Additionally, hyaluronic
acid has been used for nerve guides and is believed to arrange the extracellular matrix
through its interactions with fibrin into a hydrated open lattice through which
regenerating axons can traverse 25. Type I collagen and fibrin gels are other materials
that have been used to enhance regeneration 26. Multiple factors have been placed in
silicone tubes to promote nerve regrowth. Recombinant MDP77 which promoted neurite
86
outgrowth in cultured spinal neurons was used in a nerve guide to promote sciatic nerve
regeneration 26. Factors such as nerve growth factor (NGF), brain-derived neurotrophic
factor (BDNF), fibroblast growth factor, and glial growth factor have all been added to
the nerve guides to promote axonal growth 27. Cells to promote nerve growth have also
been used. Schwann cells have been incorporated into various guides to encourage nerve
growth as well as neuronal progenitor cells 128.
Modifications of the nerve guide material have also advanced the field of
peripheral nerve regeneration. Matrices that release neurotrophic factors such as nerve
growth factor (NGF) have been used
27.
Hydrogel graphs releasing laminin-l and NGF
improved nerve growth in Sprague-Dawley rats 5. The use of laminin in humans is
questionable because it is a possible teratogen so peptides of laminin are used and have
also been shown to promote nerve regeneration 29
4.1.4. Electricfield and nerve regrowth
Electrical charge is another factor that has been shown to influence axonal
regeneration.
In vitro studies have shown that nerves grow faster cathodically than
anodically 30. The electric fields are believed to play a role in growth cone guidance for
the growing axons 3.
Moreover, increased protein absorbtion has been observed in
electrically stimulated nerve cells as well as increased growth of neurites in PC-12 cells
32,33.
Additionally, in vivo studies of rat sciatic nerves have shown to be effected by
electromagnetic fields 34
87
4.1.5. PPy nerveguides
The conductive properties of PPy make it a favorable substrate to promote nerve
regeneration. Electrical stimulation of nerves has been shown to have a positive effect on
in vitro axonal regeneration
30.
Additionally, PPy's hydrophilicity and surface charge
properties should promote nerve regeneration similar to what is seen with the negativeion implanted silicon tubes 17.
PPy can also be easily modified to change the
microenvironment for nerve growth.
Because bioerodible forms of PPy exist, a
biodegradable, conductive nerve guide can be designed.
Previous attempts to produce PPy tubes have used a composite PPy/ silicone
structure which uses silicone tubing as the backing for PPy synthesis3
536 .
Because the
nerve guide is not required after nerve regeneration, a method that can create a standalone and potentially degradable PPy tube would be better suited for nerve regeneration.
A method for PPy tube formation that also allows for the production of any tube size
would also allow for optimal nerve guide channel selection during surgery.
4.2.
PPy tubes
The ability to control the three dimensional shape and surface texture of PPy
creates the ability to control tissue interactions. By creating PPy tubes, the conductive
material can be used not only as nerve conduits for peripheral nerve and spinal cord
injuries but also in a wide variety of applications such as serving as microfluidic
channels.
Additionally, a method where the tube dimensions and properties can be
precisely controlled allows for a flexible fabrication process.
88
Tubular PPy with diameters in the hundreds of nanometers have been formed
using various methods. PPy tubes are another option for molecular wires because of their
high conductivity and stability in air. The nanoscale dimensions cause the polymers to
exhibit appealing electrical properties, and because PPy is easily modified through
control of the plating environment, the PPy nanotubes were created as an alternative to
carbon nanotubes
37.
Methods for tubular formation of PPy include template-synthesis
methods where pores in a membrane are used to direct the polymer growth. Once the
tubes are formed, the membrane is dissolved away
38.
Nanotubes can also be fabricated
through electrochemical polymerization on a scanning micro-needle electrode
39.
Other
nanotubes have been formed through the chemical polymerization of pyrrole doped with
P-naphthalene sulfonic acid or dodecyl-benzene sulfonic acid 40,41
These methods make it difficult to isolate single tubes and to control the length
and dimensions of the tubes. Additionally, many of these methods cannot be used to
create conductive PPy tubes with diameters with dimensions of tens of microns to several
millimeters which is the size needed for nerve guidance. Because of these limitations,
another method for the fabrication of PPy tubes is needed and was developed.
4.3.
Formation of PPy tubes
4.3.1. Sacrificial core
A new method to form PPy tubes was developed that takes advantage of the
properties of the electrodeposition of PPy. When electrodeposited, PPy assumes the
shape of the object upon which it is plated. A wire was chosen as the template to achieve
89
a tubular PPy structure.
After deposition, the conductive core is sacrificed, and an
entirely polymeric PPy tubes remains.
However, removing the core is frequently a
limitation for processes like this.
Originally, aluminum wires with diameters of 25 and 100 ~m (California Fine
Wire) tested as the anode for deposition, and the sample was placed in an aluminum
etchant ( 80% phosphoric acid, 5% acetic acid, 5% nitric acid (all Aldrich) and 10%
deionized water) after deposition in an attempt to dissolve the aluminum core. The 25
and I00 ~m wires were plated at 4°C or 24°C at 0.1 mA or 0.2 mA for 40 minutes
(EG&G Princeton Applied Research Potentiostat/Galvanostat).
The PPy coated wires
were shaken, vortexed, or left stationary for I week in the aluminum etchant, but because
of the transport limitations of the solution, the wire inside the PPy tube could not be
dissolved. Because of these limitations, using an etchant to remove the sacrificial core
with a simple etch release was not possible because the ends of the wire would be etched,
but the center of the wire remained (See Figure 4.2).
Figure 4.2. PPy tubes with attempted removal of wire template with aluminum etch. a) The
etch has removed the aluminum wire at the exposed end of the tube. b) When the tube is cut in
half, the aluminum wire core still remains were the acid could not reach.
90
4.3.2. Reverse potential
A new method was developed which utilized the ionic swelling of the doped PPy.
Reversing the potential on the cathode and anode after plating disrupts the polymerconductor interface due to a force of the anion dopant being repelled from the plating
Figure 4.3. Scanning electron microscope images ofPPy tubes with a 25 Jim inner diameter. a) PPy
tube plated at 4°C. b) Higher magnification of the rough tube lumen. c) PPy tube plated at 24°C. d)
Higher magnification of smooth tube lumen. For the images the tubes were cut in halfand pictures
were taken from the middle of the tube. No wire is left after mechanical removal of the PPy from the
wire core.
91
wire and allows removal of the PPy from the core, leaving a long PPy tube.
The
dimensions of the tube using this method are only limited by the size of the wire. The
flexibility in size allows for interaction at the cellular level with 25 J.1mtubes or at the
tissue level with 1.6 mm inner diameter tubes.
The wall thickness is also easily
controlled. The thickness is a function of deposition times with longer times resulting in
thicker tube walls.
The new method allows for the creation of stand-alone,
biodegradable, electrically conducting PPy tubes without the constraints of carrier
material.
Tubes with diameters ranging from 25 J.1mto 1.6 mm have been produced (See
Figure 4.3). Tubes fabricated on wires of25 J.1m(AI, California Fine Wire), 100 J.1m,and
200 J.1m(both Pt, Sigma-Aldrich) were plated using an EG&G Princeton Applied
Research PotentiostatlGalvanostat (Model 263A) at a constant current of 0.1 mA for 40
minutes. All electrodeposition occurred under a blanket of nitrogen gas. As described in
previous chapters, the morphology of the PPy could be controlled with temperature and
Figure 4.4. PPy tubes with 500 J-lminner diameter with rough and
smooth surface. a) Plated at 24°C for half of the time. The thinner
wall gives the structure more flexibility as seen by the bend in the
tube wall. b) Plated at 4 °C giving the tube a more textured surface.
92
current densities (See Figure 4.4).
Thickness of the tube could also be precisely
controlled by limiting the plating time.
All of these variables control the physical
properties of the tube produced.
4.4.
Sciatic nerve study
4.4.1. PPy nerve guide
Tubes with diameters of 1.6 mm were fabricated for in vivo study in a rat model.
Copper wire (14 AWG, McMaster Carr) was placed in a .2 M PPy/ .2M NaDBS aqueous
solution covering approximately 30 mm of the copper wire. The wire was then attached
to the anode of a Hewlett Packard 6614 Power supply, and a current of 10 mA was
applied between the copper wire and a platinum mesh attached to the cathode for 60
minutes at 24 °C. After plating, the wire was placed on the cathode of a Hewlett Packard
33120A function generator, and a DC potential of 10
OV was applied between the wire and
a platinum mesh attached to the anode for 2 minutes. The PPy was removed by gentle
mechanical dissociation from the wire and cut with a razor blade into 15 mm sections.
The tubes were then immersed in ethanol for 5 minutes for sterilization and washed 3
times with sterile PBS. The tubes were filled with sterile PBS and stored in PBS until
surgery at room temperature.
4.4.2. Tube implantation
Five 15mm PPy tubes were implanted and compared to 5 15mm silicone tubes
with a 1.6 mm inner diameter (VWR).
The tubes were implanted for 4 weeks to
93
determine the mechanical stability and biocompatibility of the PPy tubes compared to the
standard silicone tubes. Four PPy tubes were also placed subcutaneously in 2 Sprague
Dawley rat to monitor tissue response.
The total procedure was performed according to institutional and NIH guidelines
on animal experimentation. Male Sprague Dawley rats (350-400 g, Charles River) were
used for the implantation study.
anesthesia.
The surgery was carried out under isofluorane/0
2
The rats were also given a dose of buprenorphine analgesia prior to the
surgery. An approximately 3 cm long incision was made on the right thigh of the rat.
The gluteus maximus muscle was retracted and the sciatic nerve was freed from the
surrounding tissue, and a 1 cm transection was performed at mid-thigh, proximal to the
tibial and peroneal bifurcation. The two ends of the nerve were inserted mm into the
tube ends and secured with 9-0 suture (Fine Science Tools). The muscle layers were
closed with 4-0 sutures, and the skin was fastened with Michel clips that were removed
10 days post-surgery. At 4 weeks after surgery, the rats were euthanized with carbon
dioxide gas, and the tubes were dissected and immediately placed in fixative.
4.4.3. Histology
A hematoxylin and eosin (H&E) stain was used to look at the biocompatibility
and the surrounding tissue response,.
For this stain, the samples were fixed in 10%
formalin for 24 hours and embedded in paraffin after a series of dehydration steps in
ethanol and xylenes. The paraffin was sliced and placed onto slides, and the tissue slices
stained with H&E.
94
A trichrome stain was used to view the nerve mylenation.
prepared
for
glutaraldehyde,
histology
2.0%
by
placing
them
paraformaldehyde
in
Karnovsky's
(Electron
The samples were
KIl
Microscopy
Solution
Supplies),
(2.5%
0.025%
calcium chloride in a O.1M sodium cacodylated buffer (all others Aldrich) with pH 7.4)
after dissection.
The samples were fixed at room temperature
for 24 hrs. The tissue was
then placed in warm 2.0% agar and centrifuged in Eppendorf tubes.
When the agar had
hardened, the agar blocks were prepared for electron microscopy.
The samples were
post-fixed
dehydrated
in osmium tetroxide
in graded
ethanol
and stained with uranyl
solutions
and infiltrated
acetate.
They were then
with propylene
oxide/Epon
mixtures.
Finally, the samples were flat embedded with pure Epon and stored overnight
at 60°C.
Subsequently,
Representative
one micron sections were cut and stained with toluidine blue.
area were selected for electron microscopy and thin sections of these areas
were cut with an LKB 8801 ultramicrotome
and diamond knife.
Sato's lead was used to
Figure 4.5. PPy (right) and silicone (left) nerve guides in the right sciatic nerve of
Sprague Dawley rats at 4 weeks post-implantation. No gross inflammation is present at
dissection and a fibrous capsule has surrounded both implants.
95
stain the samples and they were examined with a Phillips 30 I (Eindhoven, Netherlands)
transmission electron microscope (Similar described in 36).
4.4.4. Gross pathology
Upon dissection
and removal,
all of the PPy tubes had remained
intact, and no apparent
inflammatory
macroscopic
reaction was present at
the nerve guide site (See Figure 4.5
and Figure 4.6). The implant area was
also free from debris from the PPy
implants.
Additionally,
a thin fibrous
layer had covered most of the tubes.
All of the PPy tubes had remained
intact, and no significant
cracks had
formed in the tubes during the 4 weeks
Figure 4.6. Dissected 15mm PPy (top) and
silicone (bottom) nerve guides with attached
sciatic nerve, surrounding fibrous capsule,
and adjacent muscle prior to sectioning for
fixation and histology.
of implantation.
4.4.5. In vivo histology
The H&E stains revealed a thin fibrous capsule around the PPy implant at 28
days. No inflammatory response was seen at the 28 day time point, and the tissue was in
close contact with the PPy tube indicating favorable biocompatibility
96
(See Figure 4.7).
Figure 4.7. H&E stains of PPy tubes at 4 weeks. a) Subcutaneous implant ofPPy tube in
Sprague-Dawley rat. b) PPy nerve guide in rat sciatic nerve. Both sites have minimal
inflammatory response to the implant. Scale bar: 50 11m
The trichrome
stain showed the presence
proximal portion of the PPy tube at 4 weeks.
of myelinated
nerve fibers in the
Figure 4.8 shows the cross section of the
nerve proximal to the tube and a portion of the nerve distal to the resection site. Figure
4.9 shows a proximal segment of the nerve in the tube and a distal segment of the nerve
in the tube. The portion of the nerve proximal to the tube had numerous myelinated
Figure 4.8. Trichrome staining of sciatic nerve. a) Proximal portion of nerve with
evident myelinated nerve fibers (arrow pointed to an example). b) Distal portion to
resection demonstrating necrosis with a few nerve fibers with myelin. Scale bars: 50 11m
97
Figure 4.9. Trichrome stain of rat sciatic nerve in PPy nerve guide. a) Sciatic nerve in
proximal portion of PPy nerve guide. Scale bar: I00 ~m b) Higher magnification of
proximal portion of sciatic nerve. Arrow points to an example of a myelinated nerve fiber
which are present. Scale bar: 50 ~m c) Sciatic nerve in the distal portion ofPPy nerve guide.
Arrow points to PPy tube. Scale bar: 5011m d) Higher magnification of distal portion of
nerve in PPy nerve guide. Fewer myelinated neurons because atrophy has occurred in area
of nerve distal to lesion. Scale bar: 50~m
98
Figure 4.10. Osmium tetroxide stain of sciatic nerve at 8 weeks. a) Cross-section of
nerve inside PPy tube b) Nerve with myelnated axons at higher magnification. No
nerve was seen in the middle of the tubes at 4 weeks. Arrow points to mylenated
nerve, Scale bars: 50J.lm
Figure 4.11. Comparison of sciatic nerve distal to the resection point in the PPy tube
at a) 4 weeks and b) 8 weeks. Necrosis of the nerve is seen at the 4 wk time point
while myelinated nerve fibers (arrow points to an example of myelinated fiber) are
seen at the 8 week time point. Scale bars: 50 J.lm.
nerve fibers.
The myelinated fibers were also seen in the nerve located in the proximal
portion of the tube. The myelinated fibers decrease in the portion of the tube distal to the
resection point. The distal end of the nerve in the tube and outside the tube showed some
necrosis, but this is to be expected because the proximal nerve end has not reached the
distal segment at this time point.
The results indicate that the PPy tubes can support
99
nerve growth and that nerve tissue is present in the PPy tube at 4 weeks postimplantation. The segments of nerve in the tube are not appreciably different histochemically from segments outside of the tube.
Later time points will be needed to
elucidate the degree to which PPy promotes nerve regeneration.
4.5.
Future work
4.5.1. 8-week studies
To form a more complete picture, an eight week time point will be obtained with
both PPy and silicone tubes.
Studies lasting two months have been shown to be
necessary for 1 cm resections in the rat sciatic nerve to reach the distal end 9. At 8 weeks
post-surgery, the number of mylenated fibers in the regenerated nerves will be compared
between the PPy and silicone tubes.
Preliminary results from 8 week studies of PPy have been obtained. Nerve with
myelinated fibers was present in the middle section of the nerve conduit which had not
been seen in any of the 4 week samples (See Figure 4.10). Additionally, the portion of
the nerve distal to the nerve resection had myelinated fibers present (See Figure 4.11).
None of the 4-week time points had significant myelinated axons present in the nerve
distal to the resection. Further results of PPy and silicone tubes at 8 weeks post-surgery
will help to elucidate the effectiveness of the PPy nerve guides.
100
4.5.2. PPy
nerve
guide
modifications
Future studies could also incorporate
neurotrophic factors such as NGF into the
PPy tubes as well as include electrical
stimulation via the polymer.
Neural cells
could also be used in the tube as seen in
Figure 4.12 to help promote nerve growth.
The erodible form of PPy
42
could also be
added to determine the difference between it
and the non-degradable PPy implants.
Figure 4.12. A scanning electron microscope
image of C 17.2 neural stem cells (murine) in a
PPy tube. The cells were cultured for 3 days prior
to imaging.
101
4.6.
1.
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108
5. Drug delivery
5.1.
Mechanisms of drug delivery
Incorporating release of molecules into PPy would create a more interactive
material. Because it is possible to incorporate dopants and certain charged moieties into
PPy during electrodeposition, PPy has promising possibilities as a drug delivery scaffold.
With larger dopants such as NaDBS, PPy is also known to have piezoelectric effects
When a negative potential is applied to the NaDBS/PPy, cations from the solution flow
into the film causing expansion. This behavior has extended PPy's applications to valves
and other mechanical devices
1,2.
The flux of ions also could allow for drug delivery
applications. If the dopant is a smaller anion, the negative potential applied to the
polymer forces the dopant out of the PPy into the solution, thus acting as a chargecontrolled drug delivery platform.
5.2.
Experimental overview
An externally controlled, polymeric drug delivery system would allow the release
profile to be tailored to match physiologic processes
3.
Current implantable electronic
delivery systems are not biodegradable and often require multiple components, while
extended or controllable release polymeric systems that have been used do not allow for
switchable release profiles 4,5. Conducting polymers (e.g., polypyrrole (PPy)) offer the
possibility of controllable drug administration through electrical stimulation 6. However,
the use of conductive polymers in delivery systems has been restricted due to limitations
in the choice of dopant and the molecular weight of the delivered drug. To circumvent
these barriers, we have developed a method for attaching molecules to the surface of PPy
109
through biotin/streptavidin coupling. After attachment of the desired molecule to the
biotin dopant, drug release is triggered through electrical stimulation. This method
provides a novel platform for controlled drug delivery from a conductive polymer
substrate.
Because of PPy's beneficial chemical properties and ease of preparation, it is
often chosen for biological applications 7-9. PPy's favorable biocompatibility also makes
it an ideal electroactive polymer for drug delivery applications
10-15.
In previous
approaches to deliver drugs from PPy, the drug was either incorporated into the PPy film
or transported through a voltage-controlled polymer membrane
6,16,17.
These factors limit
the range of drugs that can be administered from PPy. We have shown that biotin can be
used as a co-dopant in the PPy film which allows greater flexibility in designing a release
system. This system is not limited to those compounds with the charge and size needed
for incorporation into the PPy during deposition. Also, the molecular release profile is
consistent between drugs because release is dependent on the properties of biotin and the
stimulation conditions instead of being dominated by the drug properties. Other groups
have incorporated biotin into the PPy monomer, making it a part of the PPy backbone 1820.
This approach works for sensor applications
21,22,
but is not ideal for drug delivery
because the biotin is part of the pyrrole monomer preventing release. By incorporating
biotin as a dopant, electrical stimulation results in reduction of the PPy backbone which
triggers the release of the biotin and the attached drug.
110
5.3.
Electrodeposition
PPy was deposited by electro-oxidation onto micropatterned 5 mm by 8 mm gold
templates (3000A of gold with 200A of titanium for adhesion, see Chapter 2 for methods)
on a <100> silicon substrate (Wafernet, Inc)
12.
The dies were cleaned by washing with
acetone, isopropanol, and water before use. Each of the compositions of PPy were
electrodeposited onto the gold surface using a constant-current power supply (HP
6614C). A current density of 2 mA/cm2 was applied between the gold template and a
platinum wire mesh reference electrode (Aldrich). The reaction was maintained at 4°C
and perfused with N2 throughout the process. For the biotin-doped PPy, the solution
consisted of 0.1 M pyrrole (Aldrich), 0.02 M sodium dodecylbenzenesulfonate (NaDBS)
(Aldrich), and 8.2 mM biotin (Molecular Probes) in an aqueous solution. For the control
without biotin, the solution consisted of only 0. IM pyrrole and 0.02M NaDBS.
5.4.
Stability studies
The biotin-doped PPy samples were stored in PBS at 4°C until the specified time
point. Streptavidin tagged with rhodamine (Molecular Probes) at a concentration of 0.1
mg/ml was added to the sample at 1, 3, 7, or 14 days. All samples (n=3 for each time
point) were imaged at 20x magnification using a Zeiss Axiovert and Hamamatsu Digital
CCD camera. ImageJ software was used to quantify the average pixel intensity (See
Figure 5.1 for summary).
111
Ele-ctrodeposition
of PPylBiotin
Polymer
polYIDe-r
+
aft(>1"addition
of rhodamine-
tagged stret;n;din
(}- Biotin
•
•
StreptaVl'd"In
with Rhodamine tag
Figure 5.1. Schematic of biotin stability tests. The biotin was plated with the PPy as a dopant.
At the stated time point, streptavidin tagged with rhodamine was added to the solution, and the
polymer was imaged.
5.5.
Drug release studies
After the PPy samples were prepared as described above, 0.2mg/ml of bare
streptavidin (Molecular Probes) in PBS was added to the sample and incubated for 15
minutes at 24°C with gentle agitation. After washing the sample once with PBS, 16
J.1g/mlof biotinylated NGF was added and incubated for 15 minutes at 24°C with gentle
agitation. NGF (R&D Systems) was biotinylated as in
23.
Briefly, 4 mg of EZ-link
biotin-PEO-Amine (Pierce) and 0.1 mg of carrier free NGF (R&D Systems) were added
to 142 J.11
ofMES-NaOH buffer (Pierce), pH 5. Then, 8 fll of.5 M l-ethyl-3-(3-dimethyl
aminopropyl) carbodimide (EDC) (Pierce) were added to the solution. After 2 hours with
vigorous stirring, the reaction was quenched with water, and the biotinylated-NGF was
dialyzed with 500 MW tubing (Spectrum) to purify the NGF.
The degree of
biotinylization was determined to be an average of 15.1 mol biotin per mol NGF through
a competitive displacement of HABA dye from avidin (Pierce). The biotinylated-NGF
solution was then lyophilized, and the biotinylated-NGF was dissolved in PBS to a
112
concentration of 16jgg/ml. After biotinylated NGF removal, the samples were washed
three times with PBS and incubated with PBS overnight at 4C.
The supernatant was
removed and fresh PBS added before each time point.
Stimulation was applied with a constant-voltage power supply (HP 6614C). For
the stimulated samples, three volts was applied between the PPy sample and a platinum
wire mesh at 24 °C. Each cycle lasted for 5 minutes. The stimulation occurred for the
prescribed time (30 or 150 seconds), after which the samples remained in the solution for
the remainder of the 5 minute cycle. Aliquots were also sampled for a five minute period
after the stimulation cycles were completed. Two stimulation periods were applied for
each of the samples. At the end of each period, the supernatant was removed from the
sample and stored at -40 C for ELISA quantification. NGF presence was quantified with
Human P-NGF DuoSet ELISA Development kit (R&D Systems, Inc.).
Direct incorporation of NGF as a dopant for release was also tried. For these
studies, 10
g/ml of NGF was added to the .02M NaDBS, .1M PPy solution for
electrodeposition. The polymer was stimulated as described above but no release was
seen. It is believed that this is due to the fact that the NGF could not free itself from the
polymer matrix.
5.6.
PC-12 cell studies
PC-12 cells (ATCC) were grown in growth medium [85% RPMI 1640 medium
with 2mM glutamine, 12.5% horse serum, 2.5% fetal bovine serum (Gibco)] with 5ng/ml
of unmodified NGF, biotinylated NGF, stimulated NGF, or no NGF. Stimulated NGF
was dialyzed (500 MW, Spectrum) and lyophilized before addition. Cells were cultured
113
on Vitrogen (Cohesion,
Inc.) coated plates at low density (6250 cellslcm
neurite measurement.
After 4 days of incubation at 37°C,
parafonnaldehyde
(Electron Microscopy
2
)
to allow for
cells were fixed in 4%
Sciences) in PBS for 20 minutes.
The cells
were then washed 3 times for 5 minutes each in PBS. Neurites of length greater than a
cell body were measured using Axiovision software, and the lengths ofthe measurements
on each cell were averaged for comparison.
5.7.
Dopant stability
Stability
of
the
iI
dopant in PPy is essential
for
many
applications
including
controlled
drug
delivery.
To measure the
dopant's
stability,
was
used
to
verify the concentration
of
biotin in PPy films over a
two
week
time
course.
After deposition, the biotindoped
PPy was stored
phosphate-buffered
b
c
fluorescently-tagged
streptavidin
d
in
saline
Figure 5.2. Stability study of biotin-doped PPy surface
using fluorescently-tagged streptavidin to indicate biotin
incorporation in PPy film. a) PPy without biotin with tag
introduced. b) One day time point with tagged streptavidin
attached to biotin in the PPy . c) Two week time point
showing biotin remaining in the PPy surface. d) Intensity
of fluorescence in the samples nonnalized by the
fluorescent intensity ofthe control sample with no biotin
incubated for various time points. There was no statistical
difference (p<.O 1) between the samples at the various time
points. Scale bar: 500 J.1m
114
(PBS) for periods up to 14 days (1, 3, 7, or 14
days).
At each end point, fluorescently-tagged
streptavidin was incubated with thorough rinsing
(See Figure 5.2a-c). As shown by the fluorescent
intensity of each sample, the amount of biotin in
the PPy remained constant without a statistically
significant change over time (p<.OI) (See Figure
Figure 5.3. Scanning electron
microscope image of PPy
doped with biotin and NaDBS
at 4°C showing crystalline
roughness on the surface.
5.2d). All intensity values are normalized to the fluorescent intensity of a control sample
of PPy that did not have biotin incorporated. These results indicate that the biotin is not
being released without activation, minimizing passive drug release. Electron microscope
images of the biotin-doped PPy surface were also acquired.
Crytalline projections
emanated from the surface creating a larger surface area which may have allowed for
greater streptavidin binding to the biotin sites on the surface (See Figure 5.3). Future
studies to determine the composition of the crystals would be important to understand the
nature of the biotin-doped PPy film.
5.8.
Biotin/NGF attachment
Biotin-doped PPy film was electrodeposited in the same manner as above, and
then incubated with streptavidin to form a surface capable of attaching any biotin-labeled
compound (Fig. 5.3a). Because streptavidin has four binding sites for biotin, it can
adhere to the biotin dopant found at the polymer's surface and still have open sites for the
addition of a biotinylated species. The fact that biotin/streptavidin coupling is possible
with the biotin incorporated in the polymer suggests that antibodies and their antigens
115
be
also
could
+
a
incorporated
matrix
into the PPy
and
because
1-
tf
released
binding was not inhibited.
b
illustrate
controlled
used
to
6.
*
D SI'Oll Slomlhlbon
(30
C LoI'Q 511mllatJOn ('~
C No SumUlalion
.NaBool"~
i5
delivery
mesh
factor
was
drug
1
•
6
.& •
Streptavidin
Platinum
PBS
(NGF)
NGF
\
In this study, the release
growth
Biotinylated
_
the site-specific
of nerve
Biotin
seconds)
&eooncl6)
I I
'I
I'
for neurological
... 2 ,
applications.
NGF
IS
a
i
,.
o.
member
of
the
nme
neurotrophin
family
which
neural
influences
growth,
differentiation,
survival and death in the
central
and
peripheral
nervous systems
has
many
24.
polnt
Figure 5.3. Stimulated NGF release. a) Voltage applied
across the polymer in PHS solution causes the release of the
biotin from the PPy surface. b) A 3V stimulation ofthe PPy
showed an increase in the amount ofNGF released from the
surface of the conductive polymer. Short stimulations of30
seconds did not result not result in as much release as long
stimulations of 150 seconds. * indicates statistical difference
versus short stimulation, no stimulation, and no biotin
dopant (p<0.05).
NGF
possible
applications because of its ability to stimulate nerve growth as well as its involvement
Alzheimer's
and other neurodegenerative
Additionally,
release of a neurotrophic
diseases
25-27
and can be biotinylated
factor can be utilized
116
to promote
in
23.
neural
interactions in many tissue engineering applications. This system is also a good model
for a controllable, erodible drug delivery device.
5.9.
Molecule release
The PPy film was washed repeatedly (3x with PBS) to ensure removal of all NGF
that simply adsorbed to the surface instead of binding to the streptavidin. Samples were
placed in PBS and release was triggered by applying a potential between a reference
electrode and the PPy for either 30 or 150 seconds (Fig. 5.3a). The supernatant was
removed and fresh solution was added. After five minutes, the supernatant was replaced
and the samples were stimulated again. The supernatant was again replaced and the films
were incubated for an additional five minutes.
An Enzyme-Linked Immunosorbent
Assay (EL ISA) kit (R&D Systems) was used to quantify the release of NGF at each time
point. Negative control samples were also tested in two different conditions - one group
included NGF but was not stimulated, and the other was exposed to NGF but the polymer
did not contain the original biotin dopant. For the control group with NGF but without
stimulation, there was minimal NGF release, which we attributed to ion exchange.
Additionally, in the control materials produced without biotin as a dopant, there was
essentially no NGF release. As seen in Figure 5.3b, the first stimulation of 3V for 30
seconds resulted in greater NGF release than with unstimulated controls.
When the
stimulation was extended to 150 seconds, there was a statistically significant (p<.05)
increase in the amount of NGF release compared to all of the controls. This stimulation
released nearly all of the NGF on the PPy surface, and the second stimulation did not
117
result in further significant delivery.
Voltages much higher than 3V resulted in
hydrolysis and ineffective delivery of NGF.
5.10. PC-12 neurite outgrowth
important
Another
aspect of a drug delivery
Il
system is to maintain the
chemical integrity of the
drug
throughout
release.
There was some concern as
to whether the electrical
stimulation process could
cause
inactivation
Fg
Unmodd
r
lfonla
Amusad
No
Figure 5.4. Average neurite outgrowth of PC-12
cells after the addition of NGF. Unmodified NGF,
biotinylated NGF, and stimulated biotinylated NGF
all resulted in similar neurite outgrowth in PC-12
cells. Cells that were just exposed to media showed
no outgrowth. * indicates statistically significant
difference (p<.01) versus cells with no NGF.
or
degradation of the released NGF. PC-12 pheochromocytoma cells, which originate from
a tumor line, are known to express functional TrkA receptors which bind to NGF and
cause the extension of neurites
activity of NGF.
15,28,30.
28,29,
and thus have been used to assay the functional
The released, biotinylated NGF (5 ng/ml) was introduced to the
PC-12 cells and compared to regular NGF (5 ng/ml), biotinylated NGF (5 ng/ml), and
plain media to determine the activity of the released NGF and the stability of the release
system. As seen in Figure 5.4, the stimulated, released NGF remained active and caused
neurite outgrowth with no statistical difference when compared to unstimulated,
biotinylated-NGF or normal NGF. No outgrowth was seen from any of the PC-12 cells
exposed to media alone (i.e., no NGF), and all three types of NGF had statistically
118
significant outgrowth when compared to the controls (p<.01). These results indicate that
the stimulated, released moiety (NGF in this case) remains intact and functional.
5.11. PPy drug delivery
Electrically triggered drug release from PPy provides a new platform for
controlled drug delivery. Although we used the delivery of NGF as a model system, a
wide range of compounds, biomolecules and drugs could be released with this system.
Multiple compounds could be released by selectively attaching compounds to an array of
PPy electrodes; the same approach could also be used for controlled release at multiple
time points. Additionally, the conductive properties of PPy open the door for remotely
controlling an entirely polymeric drug delivery device through the use of inductive
coupling or radio frequency techniques. These same approaches could also be translated
to degradable PPy materials
31.
The ability to control release externally provides more
flexibility than other release systems that rely on polymer degradation for delivery and
allows for fully customized release profiles specific to a certain disease or patient.
The incorporation of biotin as a dopant also provides a new method to selectively
tune the surface properties of PPy films through the attachment of hydrophilic and
hydrophobic moieties. Because PPy is frequently used to improve tissue interactions, the
additional ability to control both release and local surface chemistry provides a means for
flexible control of the tissue/PPy interface.
119
5.12. Hydrogel drug delivery
The combination of the conductive polymer with a biodegradable substrate allows
for greater range of drug release profiles applications. Hydrogels are one possible
degradable polymer that could be combined with PPy. Degradable hydrogels have been
developed to deliver various molecules in biomedical applications
32,33.
They are highly
biocompatibily and can be polymerized by light which allows for incorporation with a
PPy structure.
5.13. tPA
One application for a bimodal delivery system would be to improve neural
recovery following a stroke. Recovery from the large loss of neural cells associated with
stroke will involve the breakdown of the necrotic tissue and a regrowth of new cells.
Tissue plasminogen activator (tPA) has recently been shown to increase plasticity in the
brain through initiation of the tPA/plasmin cascade 3435. The use of tPA to initiate the
degradation of neural tissue to allow for new cell growth is one possible method to
achieve functional recovery. Combining externally controlled release of a neurotrophic
factor, such as NGF, from PPy, and release of tPA from a hydrogel upon degradation
would be a powerful tool for creating a nurturing environment for nerve regrowth.
5.14. tPA release studies
The first step to achieve tPA delivery is to determine the release rates over a
period of time. Degradable hydrogels were loaded with tPA, and the release profile was
120
studied over a 14 day time course. After loading with tPA, the hydrogels were placed in
0.5 ml PBS and rotated at 37 °C. Each day the supernatant was removed and the tPA
concentration was determined with an ELISA kit.
5.14.1. Hydrogel synthesis
The degradable PLA-b-PEG-b-PLA macromers were made as described in 32,33
In short, synthesis was as follows: a ring opening polymerization of d,l-lactide
(Polysciences) enabled the addition of lactic acid units to the hydroxyl end groups of
4000 MW PEG (Polysciences). This occurred under vacuum while mixing the
components at 140 °C with the addition of stannous 2-ethyl-hexanoate (Sigma) for 6
hours. After the product reached room temperature it was dissolved in methylene
chloride. Acrylate groups were then added by stirring the methylene chloride solution on
ice and adding triethylamine followed by acryloyl chloride in ration of 1:10 to the methyl
chloride. The stirring continued for 36 hours and the product was precipitated in cold
ethyl ether. Following filtering and drying under vacuum, the macromer was obtained.
5.14.2. tPA incorporation
To incorporate the tPA into the macromer, tPA (Molecular Innovations) (.04
jgg/gl single stranded, 0.66 gg/gl of single and double stranded) was added to a PBS
solution with 10 wt% macromer and 0.05wt% 2-hydroxy-1-[4-(hydroxyethoxy)phenyl]2-methyl-I -propanone ( a photoinitiator, Ciba Geigy). A mold that held 50 ,glof the
121
Average tPA Release
Total tPA conained in gel- 2 I-Ig
a)
2
E
2
1.5
t:
I~IPA
• Con
.5:!
~
0.5
i:
CIl
o.
U
t:
o
U
-0.5
ay1
day3
day5
day?
day9
day11
day13
Day
Average
b)
tPA Release
TotailPA contained in gel- 33 I-Ig
40
E
:3
30
c:
20
~c:
10
.5:!
CIl
u
c:
0
u
-10
Time Point
Figure 5.5. Average tPA release over a 14 day time
course measured by ELISA (in lU/ml). a) A small
loading dose 2 flg is placed in the hyrdogel. The
control is a degrading hydrogel without any loaded
tPA. b) A larger loading dose of33 flg is placed in
each gel (both single and double straaanded). Doublestranded tPA has a higher activity rate than an equal
amount of single-stranded. The first day release
concentrations are most likely greater than the range of
the ELISA.
122
I
solution was placed under approximately
10m W/cm2 ultraviolet light (365 nm, Blackray)
for 10 minutes.
5.14.3. tP A release
A tPA ELISA kit (Molecular Innovations)
was used to measure the functionally active tPA.
The ELISA reports the concentration of active
tPA in international units (IU/ml).
Double-
stranded tPA has a higher activity level than the
naturally occurring
single-stranded
form. [For single stranded tPA: I
tPA IU/ml= 1.67 ng tPA. For double stranded
tPA: 1 tPA IU/ml=1.38 ng tPA.]
Figure 5.6. Polypyrrole (black
regions) on a biodegradable
substrate (PGS). An all polymer
microdevice.
Over a 14 day period, the tP A was
released as the hydrogels degraded (See Figure 5.5). As shown by Figure 5.5, the amount
released follows an exponential decay. Also, the amount released scales approximately
linearly with the amount loaded in the gel. The average amount of single-stranded
tPA
released on Day 2 from the hydrogel loaded with 2 J.lgof tPA is about 15 times less than
the amount of single-stranded
tPA released from the hydrogel loaded with 33 J.lgof tPA.
Because of the approximately linear relationship with tPA released and amount loaded,
the amount of tP A can be tailored to match the desired tP A delivery rate. These results
show the plausibility of using the hydrogel as a tPA delivery platform.
123
Using an additional polymer with PPy, such as a hydrogel or poly(glycerol
sebacate) (PGS) as seen in Figure 5.6, allows for bimodal drug delivery. The ability to
use the hydrogel as a substrate for PPy and combine the drug delivery capabilities of both
polymers creates novel opportunities for stroke treatment as well as others. One
paradigm would be to utilize the initial release of tPA from the hydrogel to break down
the extracellular matrix of the neural scar formed from the stroke. After the tPA-initiated
breakdown, a triggered dose of a neurotrophic factor could be released from the PPy to
encourage neuronal growth. Additionally, the ability to synthesize an erodible form of
PPy 31 along with the degradable hyrdogel creates the possibility for complete
degradation of the polymers after the desired molecules are delivered.
124
5.15. References
1.
Berdichevsky, Y. & Lo, Y. H. in Material Research Society (Boston, MA, 2003).
2.
Smela, E. & Gadegaard, N. Volume change in polypyrrole studied by atomic
force microscopy. Journal of Physical Chemistry B 105, 9395-9405 (2001).
3.
Langer, R. New methods of drug delivery. Science 249, 1527-1533 (1990).
4.
Grayson, A. C. R., Shawgo, R. S., Li, Y. & Cima, M. J. Electronic MEMS for
triggered delivery. Adv. Drug Del. Rev. 56, 173-184 (2004).
5.
Shershen, S. & West, J. Implantable, polymeric systems for modulated drug
delivery. Adv. Drug Del. Rev. 54, 1225-1235 (2002).
6.
Miller, L. L. Electrochemically controlled release of drug ions from conducting
polymers. Mol. Cryst. and Liq. Cryst. 160, 297-301 (1988).
7.
Shi, G., Rouabhia, M., Wang, Z., Dao, L. H. & Zhang, Z. A novel electrically
conductive and biodegradable composite made of polypyrrole nanoparticles and
polylactide. Biomat. 25, 2477-2488 (2004).
8.
Peppas, N. A. & Langer, R. New Chal. in Biomat. Science 263, 1715-1720
(1994).
9.
Wong, J. Y., Langer, R. & Ingber, D. E. Electrically conducting polymers can
noninvasively control the shape and growth of mammalian cells. Proc. Nat. Acad.
Sci. 91, 3201-3204 (1994).
10.
Wang, X., Gu, X., Yuan, C., Cen, S., Zhang, P., Zhang, T., Yao, J., Chen, F. &
Chen, G. Evaluation of biocompatibility of polypyrrole in vitro and in vivo. J.
Biomed. Mat. Res. 68A, 411-422 (2004).
125
11.
Jiang, X., Marois, Y., Traore, A., Tessier, D., Dao, L. H., Guidoin, R. & Zhang,
Z. Tissue reaction to polypyrrole-coated polyester fabrics: an in vivo study in rats.
Tissue Eng. 8, 635-647 (2002).
12.
George, P. M., Lyckman, A. W., LaVan, D. A., Hegde, A., Leung, Y., Avasare,
R., Testa, C., Alexander, P. M., Langer, R. & Sur, M. Fabrication and
biocompatibility of polypyrrole implants suitable for neural prosthetics. Biomat.
26, 3511-3519 (2005).
13.
Cui, X., Hetke, J. F., Wiler, J. A., Anderson, D. J. & Martin, D. C.
Electrochemical deposition and characterization of conducting polymer
polypyrrole/PSS on multichannel neural probes. Sens. andAct. A 93, 8-18 (2001).
14.
Kamalesh, S., Tan, P., Wang, J., Lee, T., Kang, E.-T. & Wang, C.-H.
Biocompatibility of electroactive polymers in tissue. J. of Biomed. Mat. Res. 52,
467-478 (2000).
15.
Schmidt, C. E., Shastri, V. R., Vacanti, J. P. & Langer, R. Stimulation of neurite
outgrowth using an electrically conducting polymer. Proc. Nat. Acad. Sci. 94,
8948-8953 (1997).
16.
Pernaut, J.-M. & Reynolds, J. R. Use of conducting electroactive polymers fro
drug delivery and sensing of bioactive molecules. a redox chemistry approach. J.
of Phys. Chem. B 104, 4080-4090 (2000).
17.
Davey, J. M., Ralph, S. F., Too, C. O., Wallace, G. G. & Partridge, A. C.
Electrochemically controlled transport of metal ions across polypyrrole
membranes using a flow-through cell. React. & Funct. Poly. 49, 87-98 (2001).
126
18.
Cosnier, S. Biomolecule immobilization on electrode surfaces by entrapment or
attachment to electrochemically polymerized films. A review. Biosens. &
Bioelectr. 14, 443-456 (1999).
19.
Cosnier, S., Stoytcheva, M., Senillou, A., Perrot, H., Furriel, R. P. M. & Leone, F.
A. A biotinylated conducting polypyrrole for the spatially controlled construction
of an amperometric biosensor. Analyt. Chem. 71, 3692-3697 (1999).
20.
Torres-Rodriguez, L. M., Billon, M., Roget, A. & Bidan, G. A polypyrrole-biotin
based biosensor: elaboration and characterization. Synth. Met. 102, 1328-1329
(1999).
21.
Dupont-Filliard, A., Roget, A., Livache, T. & Billon, M. Reversible
oligonucleotide immobilisation based on biotinylated polypyrrole film. Anal.
Chim. Acta 449, 45-50 (2001).
22.
Dupont-Filliard, A., Billon, M., Livache, T. & Guillerez, S. Biotin/avidin system
for the generation of fully renewable DNA sensor based on biotinylated
polypyrrole film. Anal. Chim. Acta (2004).
23.
Bronfman, F. C., Tcherpakov, M., Jovin, T. M. & Fainzilber, M. Ligand-induced
internalization of the p75 neurotrophin receptor: a slow route to the signaling
endosome. J. of Neurosci. 23, 3209-3220 (2003).
24.
Kalb, R. The protean actions of neurotrophins and their receptors on the life and
death of neurons. Trends in Neurosci. 28, 5-11 (2005).
25.
Counts, S. E., Nadeem, M., Wuu, J., Ginsberg, S. D., Saragovi, H. U. & Mufson,
E. J. Reduction of cortical TrkA but not p75NTR protein in early-stage
alzheimer's disease. Ann. of Neurol. 56, 520-531 (2004).
127
26.
Lee, A. C., Yu, V. M., Lowe III, J. B., Brenner, M. J., Hunter, D. A., Mackinnon,
S. E. & Sakiyama-Elbert, S. E. Controlled release of nerve growth factor
enhances sciatic nerve regeneration. Exp. Neurol. 184, 295-303 (2003).
27.
Micera, A., Lambiase, A., Aloe, L., Bonini, S., Levi-Schaffer, F. & Bonini, S.
Nerve growth factor involvement in the visual system: implications in allergic and
neurodegenerative disease. Cyto. and Gr. Fact. Rev 15, 411-417 (2004).
28.
Greene, L. A. & Tishcler, A. S. Establishment of a noradrenergic clonal line of rat
adrenal pheochromocytoma cells which respond to nerve growth factor. Proc. of
the Nat. Acad. of Sci., USA 73, 2424-2428 (1976).
29.
Howe, C. L. Depolarization of PC 12 cells induces neurite outgrowth and
enhances nerve growth factor-induced neurite outgrowth in rats. Neurosci. Let.
351, 41-45 (2003).
30.
Levi, A., Eldridge, J. D. & Paterson, B. M. Molecular cloning oof a gene
sequence regulated by nerve growth factor. Science 229, 393-5 (1985).
31.
Zelikin, A. N., Lynn, D., Farhadi, J., Martin, I., Shastri, V. & Langer, R. Erodible
conducting polymers for potential biomedical applications. Ange. Chem. 41, 141144 (2002).
32.
Sawhney, A. S., Pathak, C. P. & Hubbell, J. A. Bioerodible hydrogels based on
photopolymerized (poly)ethylene glycol)-co-poly(alpha-hydroxy acid) diacrylate
macromers. Macromolecules 26, 581-587 (1993).
33.
Burdick, J. A., Mason, M. N., Hinman, A. D., Thorne, K. & Anseth, K. S.
Delivery of osteoinductive growth factors from degradable PEG hydrogels
128
influences osteoblast differentiation and mineralization. Journal of Controlled
Release 83, 53-63 (2002).
34.
Oray, S., Majewska, A. & Sur, M. in Societyfor Neuroscience (New Orleans, LA,
2004).
35.
Oray, S., Majewska, A. & Sur, M. Dendritic spine motility in visual cortex is
regulated by brief monocular deprivation and extracellular degradation. Neuron
(2004).
129
6. Future directions
6.1.
PPy modification
Devices made from modified PPy have the ability to interact with the
surrounding environment through the delivery of charge, surface topography,
morphology, and its ability to release molecules. Additionally, the interactions with
neural cells enable PPy to serve as a cell scaffold that can be used as a possible neural
prosthesis. The minimal neural tissue response to its presence coupled with efficient drug
delivery and sensing, should facilitate new methods to treat trauma and stroke. The
ability to create 3-D objects from PPy also allows for the fabrication of implants and
nerve guides to enhance neuronal regeneration. Another option is to combine PPy with
other polymers to create a completely polymeric bimodal system creating a unique drug
delivery system. The conductive properties of PPy also allow for the possibility of
external control of the polymer and its actions once it is implanted.
6.2.
RF applications of PPy
Because PPy is conductive, the ability to control the polymer remotely for sensing
or delivery applications also exists.
Development of a completely polymeric,
biodegradable sensor will provide a new paradigm for in vivo sensing. Being able to
pulse PPy externally also enables externally controlled drug delivery without the need for
a complex telemetry and power system. One could envision a sensor or drug delivery
device that utilized the conductive and bioerodible properties of PPy in a radio frequency
(RF) driven inductance-capacitance (L-C) coil design.
130
For sensing application, PPy's conductive properties could be made sensitive to a
molecule in the body (in the case of a prostate cancer sensor, prostate specific antigen or
PSA). By sensing a change in the PPy coil's RF signature due to the presence of the
molecule, the concentration of the molecule could be determined. The RF coils will be
powered externally which eliminates the need for an internal power supply. Since there
would be no metallic components, the coils could fully degrade after use. For drug
delivery, the RF coil would be designed with a certain resonant frequency that when
stimulated would deliver a specific molecule.
6.2.1. RF Coils
Small coils that function as L-C resonators have been tested. Once the change in
impedance of these coils is characterized, the L-C sensor can be developed. The L-C
sandwich could be made through the use of standard microfabrication techniques.
To
start development of the RF system without the complications of the polymer system,
standard microfabricated materials were first used - gold for the conductive layers and
silicon dioxide as the insulator. Silicon wafers were insulated by thermally grown silicon
nitride. Lift-off techniques were used to pattern evaporated gold to form the base coils.
Next a silicon dioxide layer was thermally grown over the surface of the wafer and a
reactive ion etch was used to remove unwanted silicon dioxide. Finally, a final layer of
gold was laid down and patterned using the standard lift-off technique described earlier to
complete the gold-silicon dioxide-gold coils (See Figure 6.1).
131
6.2.2. RF system
These coils will be used to characterize the
RF system. RF devices have been developed which
sense changes in the capacitance
1-3.
The RF system
is composed of a high frequency sweep generator
that reflects a signal onto a group of coils. Each coil
creates an L-C circuit that has a characteristic
resonant frequency.
A second receiving RF coil
collects the reflected signal from the coils, and a
peak is seen at the resonant frequency through the
use of a network analyzer. The peak will shift if a
property of the coil changes. When the PPy device
Figure 6.1. RF coils a) SEM
of gold-silicon dioxide-gold
coils produced using standard
microfabrication techniques.
b) Edge of coil showing two
gold layers with silicon
dioxide in between.
is made sensitive to a protein marker of a disease, as
in the case of PSA for prostate cancer, the resonance frequency will shift in the protein's
presence and can be detected by a change in the reflected signal.
6.2.3. Polymeric coils
After characterization of the RF system with the standard coils, bioerodible coils
could be made using PPy as the conductor and a hydrogel as the insulator. The polymer
coils can be made by plating the antibody-doped PPy and then spin coating a hydrogel
onto the surface and cross-linking with UV light. Then an exact duplicate of the PPy
coils will be aligned on top of the original coils, and the PPy will be released from the
gold surfaces, remaining attached to the hydrogel. The hydrogel will then be cut into
individual coil "sandwiches".
132
An alternative method to create a greater RF shift in the presence of PSA would
be to modify the hydrogel used as the insulator to form the capacitor in the system.
Peptide sequences targeted by PSA have been determined
4-8.
These peptide sequences
could be incorporated into a cleavable hydrogel matrix. When PSA begins to cleave the
substrate, the gel would break down, and the deterioration of the insulator would be
reflected as a shift in the resonant frequency of the coil.
By injecting large numbers of the erodible coils into the prostate at several time
points close together, the PSA levels can be monitored accurately without requiring an
extensive procedure. The coils could be tested with known concentrations of purified,
active PSA, then in vitro with a cell line that produces active PSA (i.e. LNCaP cells), and
finally in vivo in animal models with a PSA secreting tumor that has been established in
our lab.
The coils could also serve as minature drug delivery systems that when stimulated
by an RF signal release a pulse of drug to the surrounding tissue. By having an injectible
system, the coils can be delivered directly to the site and external control of release can
be achieved. Multiple types of coils with various resonant frequencies could be used to
pulse different drugs with different RF signals.
6.3.
PPy neural scaffold
PPy can also serve as scaffold for neural cells to replace damaged tissue in stroke,
spinal cord injury, or various neurodegenerative disorders. It could be used as a standalone implant similar to those developed in Chapter 3 with neural stems cells seeded on
top. With cells cultured on the implant, the PPy could be placed in a region of necrotic
133
.
Conductive polymer electrodes
.....-..
------.........
Neural cells to
be patterned by
array
Non-conductive polymer
(ie. biorubber, hydrogel)
Figure 6.2. A polymeric array for patterning of neural cells before implantation to
foster neural regeneration
tissue such as that caused by stroke to replace cells that have been loss. Because of its
conductive properties and the ability to intricately pattern the polymer, PPy could also be
electrodeposited
on a polymeric substrate in the form of an array of electrodes.
The
neural cells could then be cultured on the substrate prior to implantation, and electrical
stimulus through the PPy could prime the cells before insertion (See Figure 6.2).
Neurons subjected to such stimulus should form stronger connections which could
possibly provide a survival advantage to the stimulated cells. The cells, on the polymeric
platform, could then be implanted.
The PPy and substrate could enhance the ability of
the implanted cells to integrate with the surrounding tissue through release of molecules,
such as a neurotrophic factor, as they degrade, finally leaving only the native tissue and
the implanted cells.
134
6.4.
1.
References
Akar, O., Akin, T. & Najafi, K. A wireless batch sealed absolute capacitive
pressure sensor. Sensors and Actuators A 95, 29-38 (2001).
2.
Husak, M. One-chip integrated resonance circuit with a capacitive pressure
sensor.Journal of Micromechanicsand Microengineering7, 173-178(1997).
3.
Liu, Y., Cui, T. & Varahramyan, K. All-polymer capacitor fabricated with inkjet
printing technique. Solid-State Electronics 47, 1543-1548 (2003).
4.
Yang, C. F., Porter, E. S., Boths, J., Kanyi, D., Hsieh, M. & Cooperman, B. S.
Design of synthetic hexapeptide substrates for prostate-specific antigen usiong
single-position minilibraries. Journal of Peptide Research 54, 444-448 (1999).
5.
Wu, P., Leinonen, J., Koivunen, E., Lankinen, H. & Stenman, U.-H. Identification
of novel prostate-specific antigen-binding peptides modulating its enzyme
activity. European Journal of Biochemistry 267, 6212-6220 (2000).
6.
Takayama, T. K., Carter, C. A. & Deng, T. Activation of prostate-specific antigen
precursor (pro-PSA) by prostin, a novel human prostatic serine protease identified
by degenerate PCR. Biochemistry 40, 1679-1687 (2001).
7.
Niemela, P., Lovgren, J., Karp, M., Lilja, H. & Pettersson, K. Sensitive and
specific enzymatic assay for the determination of precursor forms of prostatespecific antigen after an activation step. Clinical Chemistry 48, 1257-1264 (2002).
8.
Coombs, G. S., Bergstrom, R. C., Pellequer, J.-L., Baker, S. I., Navre, M., Smith,
M. M., Tainer, J. A., Madison, E. L. & Corey, D. R. Substrate specificity oof
prostate-specific antigen (PSA). Chemistry and Biology 5, 475-488 (1998).
135
7. Appendix
Table 1. Dissociated Cortical Cell Testing of PPylDopant Biocompatibility
•
•
PPy and
Dopant(s)
Green - good biocompatibility defined as the formation of neural circuits
(neurons, glia, and synapse formation) after 21 days
Red - oor biocom atibilit defined as absent of cells or minimal lial resence
Concentration
PPy
Concentration Concentration
Dopant 1
Dopant 2
.2 M
.2 M
.2 M
.2 M
.2 M
(2 mg/mL) = 14.320 uM
(.2mg/mL) = 1.432 uM
.2 M
Deposition
Temperature
C
24
4
24
4
24
24
24
4
=
.2M
(0.02 mg/mL)
.143-2 uM
.2 M
(2 mg/mL) = 14.320 uM
.2M
24
.2M
(.2mg/mL) = 1.432uM
.2M
4
24
.2 M
(0.02 mg/mL) =
.143-2 uM
.2M
4
24
4
.2 M
(2mg/mL)=20uM
24
.2 M
(.2mg/mL) = 2 uM
4
24
4
136
.2 M
24
(0.02 mg/mL) = .2
uM
4
.2 M
(2mg/mL)=20uM
.2M
24
.2 M
(.2mg/mL) = 2 uM
.2M
4
24
.2 M
(0.02 mg/mL) = .2
uM
.2M
4
24
4
.2 M
.2M
.2M
(2mg/mL)=2000
uM
24
4
24
(.2mg/mL) = 200
uM
4
24
(0.02 mg/mL) =
20 uM
4
.2 M
.2M
.2M
(2mg/mL)=2000
uM
(.2mg/mL) = 200
uM
(0.02 mg/mL) =
20 uM
.2M
24
.2M
4
24
.2M
24
4
.2 M
(2mg/mL) = 3.14
mM
.2 M
(.2 mg/mL)= .314
mM
.2M
(.02mg/mL)=
.0314 mM
.2M
(1 mg/mL) = 3.14
mM
137
24
4
24
4
24
4
24
4
.2 M
(2mg/mL) = 3.14
mM
.2M
24
.2M
(.2 mg/mL)= .314
mM
.2M
4
24
.2 M
(.02mg/mL)=
.0314 mM
.2M
4
24
4
.2 M
(2mg/mL)= 1.98
mM
.2 M
(.2 mg/mL) = .198
mM
.2 M
(.02mg/mL) =
.0198 mM
24
4
24
4
24
4
.2M
(2mg/mL)= 1.98
mM
.2M
24
.2M
(.2 mg/mL) = .198
mM
.2M
4
24
.2M
(.02mg/mL) =
.0198 mM
.2M
4
24
4
.2 M
(2mg/mL) = 133
uM
.2 M
(.2 mg/mL) = 13.3
uM
138
24
4
24
.2 M
(.02mg/mL) =
1.33 uM
24
.2M
(2mg/mL) = 24.38
mM
24
.2 M
(.2mg/mL) =
2.438 mM
.2 M
(.02/mg/mL)
.2438 mM
4
24
4
24
4
.2 M
(2mg/mL) = 24.38
mM
.2 M
24
.2 M
(.2mg/mL) =
2.438 mM
.2M
4
24
4
.2 M
(2 mg/mL)= 133400 uM
.2M
24
.2 M
(.2 mg/mL) =
13.30-40 uM
.2M
4
24
.2 M
(.02mg/mL) =
1.33-4 uM
.2M
4
4
.2M
100u 110ml
24
4
.2M
(1 ug/ml)
24
.1m1/10ml
4
139
24
_:
===='2=M=====1<1 ug/ml).1 ml/1om'I=====,2=M====
140
4
BIOGRAPHICAL INFORMATION
PAUL M. GEORGE
HOME:
LAB: 45 CARLETON ST. E25-342
CAMBRIDGE, MA 02142
(617) 258-9489
E-MAIL:
P(, )I
,
(
I.
1203 BOYLSTON ST. APT30
BOSTON, MA 02215
617-852-6114
:,1) U
EDUCATION
Tulane University
Bachelor of Science in Engineering, Biomedical Engineering
* Electrical Engineering Minor, Mathematics Minor
·
Summa Cum Laude with Departmental Honors
New Orleans, LA
1999-2001
Johns Hopkins University
Masters of Science in Engineering, Biomedical Engineering
Baltimore, MD
1995-1999
2001-2005
Massachusetts Institute of Technology
Boston, MA
Ph.D., Medicaland ElectricalEngineering
·
--
----
----
Thesis Defense: July 15, 2005
'--
------I--------I
--------
2001- August 2005 Massachusetts Institute of Technology
Boston, MA
PhD Thesis,Novel Polypyrrole Derivatives to Enhance Conductive PolymerTissue Interactions
·
·
Develop polymeric based sensor and drug delivery devices
Advised by Robert Langer, Sc.D.
Spring, 2003
Massachusetts Institute of Technology
TeachingAssistant
*
*
Boston, MA
6.002 Circuits and Electronics - Taught weekly tutorials and labs
Professor Paul Gray
Johns Hopkins University
1999-2001
MSE Thesis,Microfabrication of Electrical and Electrochemical
Baltimore, MD
Sensors for Neural Encoding
Designed microfabricated electrical and neurochemical sensors to study injury in rat
hippocampal tissue slices
* Advised by Nitish Thakor, Ph.D.
*
Summer 1999 University of New Orleans
New Orleans, LA
ResearchAssistant
* Developed theoretical model of autoreceptor feedback of the dopamine neuron
·
Advised by Carmen Canavier, Ph.D.
1998-1999
Tulane University
New Orleans, LA
Senior Thesis,Computational Analysis of the Dopamine Neuron and
Its Application to Parkinson's Disease
·
·
Developed C++ code for theoretical model of dopamine neuron
Advised by Carmen Canavier, Ph.D.
141
Summer 1998,97 Mayo Clinic
Rochester, MN
Summer Undergraduate Research Fellowship
*
Utilized virtual programming languages to create user interface with virtual reality
workbench
*
Advised by Richard Robb, Ph.D.
PUBLICATIONS. PATENTS, AND CONFERENCES
P. George, D. LaVan, J. Burdick, C.Y. Chen, E. Liang, and R. Langer. ElectricalyControlled
Drug Deliveryfrom a Biotin-DopedConductivePoypyrole. Advanced Materials, submitted.
P. George, A. Lyckman, D. LaVan, A. Hegde, Y. Leung, R. Avasare, C. Testa, P.
Alexander, R. Langer, M. Sur. Fabricationand biocompatibiliyofpo/pyrroleimplants suitablefor
neuralprosthetics.Biomaterials, Vol. 26 (17), 3511-3519, 2005.
P. George, D. LaVan, S. Chen, R. Langer. Electrodepositionof biotin-dopedpopyrrole on
microfabricated
electrodes.Poster. Material Research Society Spring meeting; 2005.
P. Alexander, P. George, A. Nashat, R. Langer, D. LaVan, A. Lyckman, M. Sur, N. Wilson.
ImplantableNeural Networks. US Provisional Patent 0492611-0523, Full Patent pending.
D. LaVan, P. George, and R. Langer. Simple, Three-DimensionalMicrofabrication of
Electrodeposited
Structures. Angewandte Chemie International Edition, Vol. 42(11), 1262-1265;
2003.
P. George, R. Langer, and D. Lavan. Methods and Usesfor Three-DimensionalMicrofabrication.
US Provisional Patent 60452891, Full patent pending.
P. George, J. Muthuswamy, J. Curie, N. Thakor, and M. Paranjape. Fabricationof ScreenPrinted Carbon ElectrodeArrays for SensingNeuronal Messengers.Biomedical Microdevices, Vol.
3(4), 307-313; 2001.
N.V. Thakor, A. Bandyopadhyay, P. George, G. Mulliken, and T. Wong. Microfabricated
NeurochemicalSensorsand IntegratedVLSI Interface
for ImplantedSystems. Presentation. Annals of
Biomedical Engineering, S-119, Vol.29, Oct-2001.
M. Meyer, P. George, A. Bandyopadhyay, J. Muthuswamy, and N. Thakor. A Microfabricated
Drug Delivery and ElectrochemicalSensing Devicefor Neural Recordings.Presentation. World
Congress on Medical Physics and Biomedical Engineering, 2000.
N. Thakor, Y. Tsai, M. Meyer, A. Bandyopadhyay, and P. George. "Medical Microsystems"
in Medical Diagnostic Techniques and Procedures, M. Singh et. al eds., Narosa, New Delhi,
2000.
FELLOWSHIPS AND AWARDS
Whitaker Graduate Fellowship (99-August, 04), Sandia National Laboratory Student
Internship (2002-present), Dean's Honor Scholarship, Tulane University (95-99), Tylenol
Scholarship (95-96), Beta Club Scholarship (95-96), Eagle Scout
PROFESSIONAL MEMBERSHIPS AND DUTIES
Materials Research Society: Symposium Assistant and Referee (Fall 2002, 2003); IEEE
Engineering in Medicine and Biology Society; Biomedical Engineering Society; Tau Beta Pi;
Alpha Eta Mu Beta Biomedical Engineering Honor Society
HONORS AND ACTIVITIES
HST Student Leadership Award (05); HST Joint council, PhD Chair (03-04); HST PhD
Admission Hosting Coordinator (03); HST GSC Representative (02-03);
142
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