Characterizing Hydrogel Imposed Strain Fields on Brain Tissue Phantom for
Use in Neural Implant Device Coatings in Presence of Micromotion
By
Maxwell Ethan Plaut
Submitted to the
Department of Materials Science and Engineering
in Partial Fulfillment of the Requirements for the Degree of
Bachelor of Science
at the
Massachusetts Institute of Technology
MASSACHUSETTS INSTITUTE
OF TECHNOLOGY
JUN 0
June 2014
14
LIBRARIES
©2014 Maxwell Plaut
All rights reserved
The author hereby grants to MIT permission to reproduce and to
distribute publicly paper and electronic copies of this thesis document in whole or in part
in any medium now known or hereafter created.
Signature of Aut hor
Signature redacted
Department of Materials Science and Engineering
9,2014
4 ,May
ool
Certified by
Signature redacted
Dai H. Koc Prfso ofEgnern
Michael Cima
V
David H. Koch Professor of Engineering
Thesis Supervisor
Accepted by
Signature redacted
f IVV/
Jeffrey Grossman
Associate Pofessor of Materials Science and Engineering
Chairman, Undergraduate Committee
Characterizing Hydrogel Imposed Strain Fields on Brain Tissue Phantom for
Use in Neural Implant Device Coatings in Presence of Micromotion
By
Maxwell E. Plaut
Submitted to the Department of Materials Science and Engineering on May 9, 2014 in
Partial Fulfillment of the Requirements for the Degree of Bachelor of Science in
Materials Science and Engineering
ABSTRACT
Glial scar tissue forms in the brain as a response to the implant injury and hampers the
effectiveness of the implant treatment. Constant relative micromotion between the
mechanically mismatched neural implant and brain tissue is thought to play a key role
glial scar formation. This study investigated the effects of poly(ethylene glycol) (PEG)
hydrogel coatings for glass brain implant devices on strain fields imposed by those
devices to brain tissue due to micromotion in the brain. PEG hydrogels were created
using macromers of 2000-8000 M, and 5-20 wt.% in solution. The moduli of the
hydrogels were calculated via Hertzian analysis of force-deflection curves produced
using an AFM tip as a nanoindenter. The moduli of the samples did not change
significantly with change in macromer M., but did change with solution concentration.
20% gels had moduli of 120-180 kPa and 5-10% gels had moduli of 0-20 kPa. The strains
imposed by the coated devices were found to be lower at the surface by ~30% as
compared to uncoated and the strain field dropped off much more quickly.
2
Introduction
Many neurological diseases, termed circuit disorders, are the result of errors in
communication between anatomically distinct regions of the brain which make up a
neural circuit' 2 . Circuit disorders can be severely debilitating to the persons with them
and some, such as anxiety and depression, are rather prevalent today. Some of these
disorders have been shown to be associated with activity at particular foci within the
brain
. While the understanding of these diseases, and how they affect brain function
has increased tremendously in recent times, there is a lack of technology that enables
effective treatment of these disorders. Treatments containing direct electrical stimulation
to the specific foci in the brain have shown that targeted treatment at the neural circuit
level can be effective at treating these disorders 34,6
''. Members of the Cima lab have
developed a novel drug delivery implant device, named the injectrode, which is capable
of electrically stimulating and recording specific neural areas with an electrode while
simultaneously delivering nanoliter volumes of drug solutions on demand to targeted
regions of the brain through a borosilicate capillary. This targeted approach is expected to
be more effective in regulating circuit activity than electrical stimulation or drug delivery
alone.
There are several materials issues associated with the use of a chronic drug delivery
device within the central nervous system (CNS). Glial scar tissue forms in the brain as a
response to the implant injury and hampers the effectiveness of the implant treatment7.
Many current implant devices and treatment efforts result in significant glial scarring
around the implant device-.
Glial scarring has an increased impendence as compared to
normal brain tissue, and since many of these treatments include electrical stimulation or
3
recording via implanted electrode, the devices' effectiveness tends towards being
negligible over time'".
Damage resulting from micromotion between the brain tissue
and implant is thought to be a key contributor to glial scar formation . Experimental
attempts to reduce glial scar formation from a materials approach have shown modest
success, but have made use of materials with moduli that are still several orders of
magnitude higher than that of brain tissue '14
Hydrogels are highly swollen cross-linked polymer networks widely used in
biomedical applications due to their similar mechanical properties to biological tissue 5 1' 6
Hydrogel coatings that incorporate conductive polymers have been used to improve the
recording capabilities of neural probes ' 1 . However, the effect that the mechanical
properties of these coatings have on the chronic glial scar response has not been
extensively characterized. The aim of this project will investigate the use of hydrogel
coatings (100 [tm scale), with moduli on the order of brain tissue, to reduce the strain
field resulting from implantation. The results of these experiments will help guide the
design of future neural implants and will serve as a crucial step towards the
implementation of chronic delivery systems to treat neurologic diseases of the brain.
Theory
Glial Scarring. In order to effectively treat circuit disorders, devices such as the
injectrode require chronic operation. The injectrode must also operate safely and
effectively at the time of implantation and continue to do so in the presence of bodily
biological response to the device. The brain's primary immune to response to neural
4
electrodes, astrogliosis, results in the formation of glial scar tissue around the electrodes.
The main cellular contributors in the astrogliosis response to implants or other injury to
the central nervous system are astrocytes and microglia. Astrocytes, characterized by
immunostaining for the intermediate filament protein glial fibrillary acidic protein
(GFAP)19 , are the main component of glial scar tissue. Microglia are macrophages in the
brain that engulf particles and secrete proteins that affect a variety of processes that
support inflammation. Astrogliosis is generally divided into two main phases, the acute
phase and the chronic phase7 .
Prolonged Reactive
Response
Ealy Reactive
Response
(A)
Possible snn
Possile signling
patways
pabWays
Possibl, cell
banstounasons
PossibleOcll
tanstonnatons
(B)
Figure 1. Possible mechanics of cellular response to an implant. Acute phase (a) and chronic
phase (b) of cellular response to implant device injury. Neurons are in pink, microglia in blue,
vasculature in purple and astrocytes in red.7
The short-term response to foreign object implantation in the central nervous system
is what is referred to as the acute response. The acute response occurs over about the first
two weeks of implantation and is characterized by a large amount of microglia activity.
Implantation of the device causes injury to the brain through cell damage, blood vessel
severing and disruption of the extracellular matrix. This injury causes the release of
several protein factors into the area to promote inflammation and clotting20 . The proteins
promoting inflammation cause inflammatory cells and microglia to join and form an
5
envelope of cells around the injury. The microglia then breakdown red blood cells in the
area and any cellular remains from the injury, then release cytokines and reactive oxygen
intermediates in order to promote inflammation. This inflammatory response can
negatively affect device effectiveness in the area of the injury7 .
Implantation times longer than two weeks usually mark the end of the acute phase
and lead to the chronic phase of the astrogliosis. The chronic phase is characterized by
the formation and densification of glial scar tissue. Astrocytes isolate the injury site from
the rest of the neural tissue even further than was done in the acute phase and increase
production of certain extracellular molecules, including chondroitin sulfate
proteoglycans, that inhibit axon regeneration 7,21,22. The glial tissue, typically a few
hundred microns thick, reaches its full size 4-6 weeks after the initial injury and remains
generally stable for the duration of the implantation (Figure 2)1o,23
2 WAeks
6 Wanks
4 Wmeek
12 WeekR
Figure 2 Time course of glial scar formation around a neural implant. Image was made
by GFAP immunostaining. At 2 and 4 weeks the glial scar falls back into void left by the
electrode. At 6 and 12 weeks the glial scar is a dense sheath which does not collapse into the
void2 1
This process of the chronic phase is analogous to the fibrous encapsulation observed
around implants in other parts of the body 24-26. Additionally, microglia typically remain
at the surface of the implant injury throughout the chronic phase forming a dense layer of
cells on the implant surface that is surrounded by the glial scar 27
6
Glial scar is formed with the intent to protect the brain and the rest of the central
nervous system from the implanted foreign body within the tissue as well as the reactive
proteins and molecules released to the injury site in the process7,28,29. However, while
generally beneficial, glial scar formation is believed to be a key contributor to implant
device and electrode failure. Nearly half of recording electrode implants fail 6 months
". Neurons typically need to be
after implantation despite initially operating properly 11,12
less than 100tm from the recording electrode to be recorded3 0 , and the formation of glial
scarring forces neurons near the implant away from the injury site, decreasing the
potential signal that can be recorded by these devices and eventually leading to complete
failure of the electrode implant 26 . Glial scarring increases the impedance of the tissue,
thus reducing the volume of tissue activated by electrical stimulation by up to 50%".
This can be partially overcome to get the desired therapeutic effect on a given neural
circuit through use of a higher current but a higher current can lead to neuron damage31.
Sykova, et al. used ion selective electrode measurements to investigate how diffusion
properties of cells change following astrogliosis in response to a stab wound 21,323, 3 and it
was found that the astrogliosis leads to an increase in tortuosity, increase in the volume
fraction of extracellular space, and decreased cellular uptake. This hindered diffusion
environment likely affects the extra synaptic transmission of neurons, hinders diffusion of
glucose and oxygen to cells from circulation, and contributes to the barrier to axon
regeneration around the injury. It should be noted that the glial scar resulting from stab
wounds should have slightly different properties than those resulting from chronic
implants due to constant presence of brain micromotion in the case of chronic implants.
7
There have been many studies conducted with the goal of reducing the scarring
response to implanted electrodes with results of varying success' 35'36 . These studies
provide insight into the mechanisms of scar formation, and provide suggestions regarding
device properties that are important to consider when designing chronic neural implants.
One contributor, mechanical mismatch, will be a focus of this project and is discussed
below.
Modulus Mismatch and Micromotion. Brain tissue is constantly undergoing
micromotions up to 40 pm in magnitude in rats due to respiration, vascular pulses, and
rotational accelerations37,38 . Typical neural implant device materials have moduli that are
many orders of magnitude higher than that of brain tissue. Constant relative
motion between the neural implant and tissue is
thought to play a key role in the chronic tissue
Table 1. Moduli of common
implant materials 49' 50 and brain
tissue 4 1
response through constant aggravation of local
inflammatory cells and damage to local
vasculature 39 . This theory is supported by the
finding that implants that are tethered to the skull
result in a much greater glial scar response than
Elastic
MaterialModulus
Material
(GPa)
69
Borosilicate Glass
Gold
80
200
Silicon
Titanium
107
Polyimide
Brain
2
.000005
untethered implants which are free to move with the brain's micromotion 4. Tethered
implants are fixed in place relative to the skull and thus result in a greater relative motion
between the implant and tissue. This results in greater injury and aggravation of tissue
surrounding the implant and more extensive scar formation. Finite element analysis
simulations have been conducted to investigate the effect that micromotion and
mechanical mismatch have on the surrounding tissue' 41 . These simulations estimate the
8
amount of strain that the brain tissue experiences as a result of brain micromotion in the
presence of neural implants of various mechanical properties. It was found that a probe
composed of a hypothetical soft material with Young's modulus of 6MPa results in a
strain two orders of magnitude less than that of a silicon probe with a modulus 200GPa41 .
Poor tissue-device adhesion also was found to contribute to elevated strains. A tangential
tethering force from tissue adhesion to the device reduced strains near the tip of the
electrode by
94%41.
This further corroborates the theory that tissue aggravation causes
glial scar formation. Neural implant designs could incorporate materials with lower
mechanical strengths or coatings that promote adhesion with neural tissue to reduce the
extent of glial scar formation.
Hydrogel Coating. It would follow, therefore, that implanting devices with a hydrogel
coating would result in a reduced reactive gliosis response compared to uncoated glass
devices following micromotion. Hydrogels are already widely used in biological
applications and have mechanical properties similar to that of brain tissue15 16 . Although
the presence of a hydrogel coating would likely not completely prevent scar formation
due to the central nervous system's foreign body response, a mechanically matched
coating would reduce injury to the brain tissue due to micromotion caused strain. This
decrease in strain imposed on the brain would reduce the astrogliosis response and the
overall size of the glial scar would be expected to be reduced, increasing the effectiveness
of the device.
Hydrogel Mechanics. The key properties in considering a hydrogel coating for a neural
implant device are its mechanical properties. Due to the numerous controllable variables
associated with the gel formation process, it is possible to tune to the mechanical
9
properties of the formed gel to match that of brain tissue, which has a modulus of-5kPa.
The parameters that are controllable include the molecular weight of the gel precursor
polymer chains and the concentration of those monomers in solution.
The microstructural origin of gel elastic response is the entropic spring effect.
Stretching polymer chains, or otherwise mechanically forcing them out of a high-entropy
coil conformation, reduces the number of available microstructural conformations
available to that chain, decreasing the entropy of the system. Since the potential entropic
contribution to the free energy of the chains is so high, the stretched state is
thermodynamically unfavorable. When the external stretching force is removed from the
system the polymer chain will snap back to its equilibrium, unstretched state.
Hydrogels are formed by linking together many smaller chains of a water-soluble
polymer. One methods of crosslinking these chains is through radiation crosslinking
using ultraviolet light. Irradiation of polymer containing solution causes the creation of
free radicals which recombine to form chemical crosslinks between chains. The resulting
gel can then swell in the presence of a solvent, which in the case of hydrogels, is water.
The Young's modulus of these gels is highly dependent on the total crosslink density
which includes both chemical and physical crosslinks. Increasing the crosslink density
restricts chain movement within the network. As a consequence, this prevents
deformation of the overall structure in response to applied stress.
Crosslinking density is primarily tunable through selection of gel precursor. Free
radical formation, the first step in chemical crosslinking, occurs at the ends of polymer
chains. Increasing the density of polymer chains ends before gelation would therefore
increase the resulting chemical crosslink density. The simplest way of increasing chain
10
ends is using small molecular weight chains as a precursor. Smaller molecular weight
means fewer monomers per chains and a shorter chain, meaning an increase in chain ends
per monomer. Figure 3 shows two different gels, one formed from high molecular weight
chains and one formed from low molecular weight chains.
a)
b)
Figure 3. Schematic of ideal networks. Schematic shows the difference in number of crosslinks,
circled in red, between network polymers (in green) formed from high molecular weight
polymeric chains (a) and low molecular weight chains (b). The chemical crosslink density is
much higher in the low molecular weight case. This diagram does not take physical crosslinks
into account
Physical crosslinks occur when the polymer chains get entangled and restrict
movement This is purely a topological constraint imposed on the system when two
polymer chain intertwine and are relatively fixed in place by steric effects42 . These
entanglement effects begin to occur at low chain length and have rather significant effects
on the mechanical properties of the bulk polymer. Even in melt form, due to increasing
entanglement, the viscosity of a polymer is related to the cube of the molecular weight 42.
At longer chain lengths (>100 monomers) this entanglement effect is often the dominant
crosslinking effect, especially at only small variations in degree of chemical crosslinking.
11
F
V
p
-
py
f 4ha
Win
y
s
Figure 4. An entangled polymer chain forming physical crosslinks
Gel modulus can also be affected by the polymer concentration in solution prior
to crosslinking. Gelation procedures typically dictate that gels form from specific volume
of solution, usually to obtain a particular thickness of shape of the gel. Under this
condition, changing the concentration of polymer in the solution results in a change in
total amount of polymer in the network while the volume of the total gel remains
unchanged. Since solvent occupies the volume of gel not occupied by polymer,
decreasing the polymer amount effectively increases the volume of solvent. This loss of
connectivity allows the structure to yield more under applied stresses, meaning a lower
modulus for those structures.
Materials and Methods
Materials. Poly(ethylene glycol) (PEG) (Mw~ 2000 - 8000 g/mol), methacrylic
anhydride, 2-isocyanatoethyl methacrylate (IEM), ethyl ether, and triethylamine (TEA)
were purchased from Sigma-Aldrich and used as received. Dichloromethane, sulfuric
acid, hydrogen peroxide solution, 3-(Trichlorosilyl)propyl methacrylate (TPM), 2-
12
hydroxy-4'-(2-hydroxyethoxy)-2-methylpropiophenone, 2,2-dimethoxy-2phenylacetophenone and carbon tetrachloride were purchased from Sigma-Aldrich.
Heptane was purchased from Macron Fine Chemicals and used as received. The glass
slides used were 12mm diameter with a thickness of .1 6-.19mm and were purchased from
Electron Microscopy Sciences. Glass capillaries were purchased from VWR. Fluorescent
polystyrene particles with a diameter of 5.9ptm were purchased from Polysciences, Inc.
and used as received.
Equipment. The Cure Spot 50 from ADAC systems controlled the ultraviolet radiation
source during gelation. Strain field measurements were performed using a standard bright
field optical microscope from Micro-Tech Optical, Inc. Force and modulus measurements
were taken using an atomic force microscope from Veeco, the Nanoscope IV with
Multimode and Picoforce
PEG Synthesis. Poly(ethylene glycol) dimethacrylate was prepared by the procedure
described by Gibson et a, 43 . PEGDM was formed from the reaction of various PEGs, MA
and IEM. An example of the synthesis of a 5k PEGDM is as follows. PEG (5 g, Z0.001
mol), 2.2 equiv of MA (0.34 g, 0.0022 mol), and TEA (0.2 mL) were reacted in =15 mL
of dichloromethane over freshly activated molecular sieves (Z3 g) for 4 days at room
temperature. The solution was filtered over alumina and precipitated into ethyl ether. The
43
product was filtered and then dried in a vacuum oven overnight at room temperature
0
0
H
0
N
0
00
0 ,
N
0
Figure 5. Structure of Poly(ethylene glycol) dimethacrylate
13
Surface Functionalization. Glass substrates were functionalized using standard
protocols for surface modification as described by Revzin et al. 44' 45 . Substrates were
cleaned in a 3:1 sulfuric acid:hydrogen peroxide "piranha" solution and then treated with
TPM in a 1mM solution of 4:1 heptane:carbon tetrachloride in a nitrogen atmosphere.
The substrates were washed and dried after each step. This treatment forms a monolayer
of methacrylate groups on the surface on the glass to provide points for the PEG to
covalently bond to the surface to prevent delamination
44
0
/
OH
OH
OH
Si0
SiI
Si
0
0
0
I
OH
OH
Si
SiN
I
0
0
_
_
I
HO-Si
TPM
in 4:1 Heptane:CC
4
OH
0
-0
0
0
Si-OH
LH
19 S,lo
a
OH
10101
I~ II
Figure 6. Surface functionalization of glass substrates. A clean glass surface (left) gets
functionalized to have methacrylate groups (rights) for PEG to bind to by TPM in a 4:1
Heptane:CCl 4 environment.
Gel Formation. Functionalized glass substrates were then coated with a solution
containing dissolved PEG macromers at the concentrations of 5,10 and 20% by weight,
with .5% photoinitiator. Samples were coated via either a tube filling process, for
capillaries, or direct pipetting onto the surface. The solutions were exposed to 365 nm
ultraviolet radiation for to crosslink the PEG via free radical polymerization. The solution
was exposed to UV for 90 seconds or, in the case of low concentrations, until gelation
occurred for direct pipetting. The glass capillaries were suspended inside a larger glass
tube for tube filling and solution was dripped into the tube and then pulled across the
14
surface of the capillary through capillary action. This method was utilized to ensure a
uniform thickness of the hydrogel coating across and around the device.
b
Photoinitiator
Light
Figure 7. Schematic representation of the PEGDM network showing (a) cross-linked PEG
chains and network defects including, (b) unreacted acrylate terminuses, (c) PEG cycles and (d)
chain entanglements. Shaded arrow shapes represent reacted acrylates, unshaded arrow shapes
represent unreacted acrylate, and dark lines represent PEG chains esterified to the acrylic acid.
For clarity, short acrylate chains are shown, but in actual gels these chain lengths may be much
longer 6 .
Force Measurements and Modulus Calculations. The atomic force microscope (AFM)
is well suited for probing the local elasticity of small very soft samples 47 when the
cantilver tip is used as a nanoindenter. Samples with swollen PEG hydrogel were placed
in the AFM and probed by a spherical polystyrene tip with a diameter of 45plm and a
cantilever spring constant of 14 N/m. Tip deflection as a function of depth of indentation
was measured. The Young's moduli of the gels were then determined via Hertzian
analysis of the resulting force-displacement curves.
This analysis assumes that the
hydrogels are deforming in a linear elastic fashion and there is negligible adhesion
between the gels and the AFM tip. Using the methods presented by Lin et al., an AFM
15
was used to measure the relationship between tip deflection and depth of indentation into
the gel. The force applied by the tip was calculated by
F = kc(d - do)
(1)
where kc is the spring constant of the cantilever, d is the measured deflection, and do is
the deflection offset at the point of contact 4 7. Using the spherical geometry of the tip it
was then possible to derive the modulus from the relationship
1
F
4 E R22
3(1-V )
z - Zo) - (d - do)] 3/2
(2)
where z is the measured translation of the cantilever, z is the translation of the cantilever
at the contact point, R is the radius of the tip, v is the Poisson's ratio, which was assumed
to be .5, and E is the Young's modulus47 .
quad photodiode
laser
cantilever-tip
sample
Figure 8. Basic schematic of AFM. An AFM tip can be used as a nanoindenter n the sample and
the force required to indent a sample can be calculated by measuring the deflection of the
cantilever.
16
Particle Tracking Analysis. The brain undergoes micromotion in the radial directions
(along device axis) due to vascular (1-3 pm) and respiratory pulsations (2-40 ptm) 37'38 .
Agarose gel with fluorescent particles suspended was formed around the coated device to
prevent shearing during insertion. The device was then displaced 20-40 Pm in the radial
direction to simulate micromotion.
Figure 9. Strain field measurement setup. The coated glass capillary (a) is immersed in
agarose gel in a petri dish (b) under the optical microscope. It is held by a capillary holder (c)
which is connected to a stepper motor capable of simulating micromotion in the brain in the radial
direction
Bright field images before and after deformation were analyzed to obtain the resulting
displacements for each particle. The strain fields in the gel were calculated from the
observed particle displacements using the particle image velocimetry plugin for ImageJ.
Strain fields imposed by different samples were compared by examination of the
17
maximum particle displacement around the implant as well the variation displacement
vectors as a function of distance from the implant 8 . This analysis determined the effect
that hydrogel coatings have on cells directly around the implant, as well as estimate the
volume of brain tissue around the implant that experiences stress resulting from
micromotion respectively.
Figure 10. Optical bright field microscope image of uncoated device. The black
capillary is the uncoated glass device. The device is surrounded by agarose gel and the
multitudes of black spots in the image are the polystyrene particles suspended in the gel
used for strain field analysis.
Results and Discussion
Modulus Calculations of Gels. The moduli were visualized by plotting as a function of
both molecular weight of the original chains of PEG and by the concentration of the
polymer in the gel precursor solution.
When comparing the moduli of the samples as a function of molecular weight of
the original chains (figure 11), it seems as though varying the molecular weight of the
polymer has little to no effect on the resulting modulus. Increasing the molecular weight
18
increased the resulting physical crosslink density slightly, but greatly decreased the
chemical crosslink density. A great decrease in chemical crosslink density, without any
physical crosslinking would result in a great decrease in modulus. In these samples,
however, the degree of physical crosslinking seems to be the dominant factor in
determining the modulus of the samples. The change in chemical crosslink density,
modulated through the change in molecular weight, at least on this scale, has no
significant effect on the modulus of samples of similar precursor solution concentration.
Changing the concentration of PEG in the precursor solution does seem to be able
to have a large effect on modulus of the samples. The modulus of the gel was observed to
increase as a function of polymer concentration for all chain lengths (figure 12). This
suggests a greater crosslink density at higher polymer concentrations. The 10% gels have
slightly higher moduli than the 5% gels, but the set of 20% gels has considerably higher
moduli than the other gels. This is likely due to the lower concentration polymer being
too disperse in solution to form a homogenous distribution of clusters. This results in
large defects in matrix, and reduced crosslinking, and thus a reduced modulus. At low
concentration it is also more likely that the reactive terminuses of a single PEG chain
react with themselves, further reducing the connectivity of the resulting gel. Somewhere
in the range between 10%-20% a critical concentration is reached at which homogeneity
is reached and the gel can properly form.
19
Modulus vs. Mw of PEG
200
180
y = -0.0056x + 178.39
160
140
+20Wt.%
120
*5Wt.%
A
100
Wt.%
10
60
401
0 0009
y=- .
rR
20
x + 18.172
= 0 84463
Sy
0
=-
0
2000
1000
3000
4000
Mw of PEG
5000
6000
7000
0 0 0 07
.
x + 7.
3584
R=0.95922
-
--
8000
9000
(g/mot)
Figure 11. Plot of modulus vs. molecular weight. Plot shows the change in modulus of samples
using 5 wt.% (red), 10 wt.% (green) and 20 wt.% (blue) precursor solutions. For each
concentration the modulus seems to change very little in response to changing molecular weight
Modulus vs. Wt. % Polymer In Solution
200
150
2000
--.-,
4000
.8000
100
50.--
0
5%
10%
20%
PEG Conc In Sorn (Wt. %)
Figure 12. Plot of modulus vs. concentration of PEG. Plot shows the change in modulus as a
function of polymer concentration in precursor solution for given molecular weights on original
polymer chains. There seems to be a threshold at which the modulus increases dramatically
somewhere between 10%-20%.
20
Strain Field Analysis. The particle displacement vector fields were visualized by
determining the movement of each polystyrene bead in the gel near the device tip after
simulating micromotion through movement of the device and then plotting that
movement as a vector. Cross sections of the vector plots were taken near the device tips
and graphed showing the change in vector magnitude as a function of distance from the
device.
The strain imposed on the gel was highest right near the devices and decreased
with distance from the device on all sides. Both the particle displacement at the surface of
the device and the displacement at longer distances were smaller when the devices had a
PEG coating as compared to the control.
The polystyrene beads near the surface of the device in the uncoated sample move
about 3Optm, which is the amount of movement applied to the device to simulate the
micromotion. Near the surface of the hydrogel coated devices, which have a reduced
modulus, the beads were found to move 15-20pm. The displacements in the coated
device samples also have a higher rate of decrease as a function of distance from the
device than the uncoated devices. The exception to this was the coated sample made from
PEG-2000. The particle displacement on one side of the device was comparable to the
particle displacement of the uncoated, while the other side has displacement more similar
to the coated device (figure 13). This can be understood by looking at a picture of the
device. The device has a hydrogel coating on one side, while the other side is essentially
uncoated. This further corroborates the difference between imposed displacement fields
from uncoated and coated devices. These finding also support those of the simulations
21
run by Jeyakumar et al. that soft materials imposed lower strain on brain tissue
43
undergoing micromotion than materials with high moduli .
Figure 13. PEG-2000 coated device. The black part of the device is the glass capillary
and the area below it is the PEG coating. There is very little coating above the device leading to
strain behavior in that region more similar to that of uncoated devices. The red line denotes the
line on which the cross section of the displacement field was taken.
25404
0004MW
200OMw
C(0ntro4
-
.
- -
-
:
25M
"
"40
'
-
-
-
I
I
2000
200
4,500M
4000
-A
0
560
10M0
15011
2440
25M4
0
[5M0'
50
ll
,-"
6
"'P)
2000
2
0
500'
400
1506
'-i-44~4
2000
2500
(PM)
Figure 14. Vector plots of particle displacement field. Particle displacement fields imposed by
device under -30pm motion on the surrounding agarose. Scale of vector magnitude ranges from
0-30pm. Left to right: Uncoated, PEG-2000 coating, PEG-8000 coating
22
Figure 15. Particle displacement field vector plot overlays. Images of the particle displacement
fields associated with uncoated (A), PEG-2000 coated (B), and PEG-8000 coated (C) overlaid on
images of the devices immersed in the agarose.
23
Magnitude,f
Magnitue 'r
Vtor tm)
V-1"or (W")
30
Figure 16. Particle displacement field cross sections. Cross sections of the particle
displacement fields near the tip of the uncoated (left) and PEG-8000 coated (right) devices.
Magnitude of the particle displacement vector is plotted. The red shaded area is the actual device.
for
The coated device is much larger because of the coating. The particle displacements are lower
sample.
the
of
modulus
the PEG-8000 coated device due to the lower
2000Mw
Magnitude of Vector (pm)
0!
25
90
20
9
9
15
Sg
9
9
9
999
*9
9
69
-PI
10
I
500
r
1000
y-position (pmn)
1500
Figure 17. Particle displacement field cross section of PEG-2000 coated device. Cross section
of the particle displacement fields near the tip of the PEG-2000 coated device. This coated device
only has a significant amount of coating on the bottom (corresponds to the left side) and the
agarose on that side has a reduced particle displacement when compared to the nearly uncoated
top (right in this image).
24
Displacment at Surface
40
35
30
25
20
S
R~.
10
10
0
8000 Mw
Uncoated
2000 Mw uncoated
side
2000 Mw Coated
side
Figure 18. Displacement at Surface. Comparison of particle displacement near the surface of
the device. For the PEG 2000 coating the two sides of the device were analyzed separately due to
the loss of hydrogel coating on one side.
Conclusions
The use of PEG hydrogel coatings on glass brain implant devices reduces the
strain field imposed by those devices on tissue due to brain micromotion. The lower
modulus of the hydrogel coating acts as a remedy to the mechanical mismatch between
the high modulus glass (~69 GPa) and the low modulus brain tissue (5 kPa).
The moduli of these coating can be controlled to an extent using varying
concentration of PEG in solution before gelation. It is possible that the modulus can also
be controlled by varying the degree of swelling, or by varying the chain lengths to a
much higher degree. Thickness of the gel may also be a factor in the moduli of the gel if
the gels are thin enough that there are still residual effects of the high modulus glass even
at the surface of the gel.
25
The reduction in strain field should reduce the extent of glial scarring near the
implant area of the device. In effect, the addition of PEG hydrogel coatings should
improve the effectiveness and longevity of the device. More work needs to be done to
understand the effects of the act of implantation on the hydrogel coating to prevent
shearing and the effect of gel swelling on the device.
26
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