U.S. patent application number 15/981761 was filed with the patent office on 2019-01-17 for hydrogel arthroplasty device.
This patent application is currently assigned to The Board of Trustees of The Leland Stanford Junior University. The applicant listed for this patent is The Board of Trustees of The Leland Stanford Junior University. Invention is credited to Dennis R. CARTER, Curtis W. FRANK, Stuart B. GOODMAN, Laura HARTMANN, Lampros KOURTIS, David MYUNG.
Application Number | 20190015211 15/981761 |
Document ID | / |
Family ID | 40509268 |
Filed Date | 2019-01-17 |
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United States Patent
Application |
20190015211 |
Kind Code |
A1 |
MYUNG; David ; et
al. |
January 17, 2019 |
HYDROGEL ARTHROPLASTY DEVICE
Abstract
An arthroplasty device is provided having an interpenetrating
polymer network (IPN) hydrogel that is strain-hardened by swelling
and adapted to be held in place in a joint by conforming to a bone
geometry. The strain-hardened IPN hydrogel is based on two
different networks: (1) a non-silicone network of preformed
hydrophilic non-ionic telechelic macromonomers chemically
cross-linked by polymerization of its end-groups, and (2) a
non-silicone network of ionizable monomers. The second network was
polymerized and chemically cross-linked in the presence of the
first network and has formed physical cross-links with the first
network. Within the IPN, the degree of chemical cross-linking in
the second network is less than in the first network. An aqueous
salt solution (neutral pH) is used to ionize and swell the second
network. The swelling of the second network is constrained by the
first network resulting in an increase in effective physical
cross-links within the IPN.
Inventors: |
MYUNG; David; (Santa Clara,
CA) ; KOURTIS; Lampros; (San Francisco, CA) ;
HARTMANN; Laura; (Berlin, DE) ; FRANK; Curtis W.;
(Cupertino, CA) ; GOODMAN; Stuart B.; (Los Altos,
CA) ; CARTER; Dennis R.; (Stanford, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The Board of Trustees of The Leland Stanford Junior
University |
Stanford |
CA |
US |
|
|
Assignee: |
The Board of Trustees of The Leland
Stanford Junior University
Stanford
CA
|
Family ID: |
40509268 |
Appl. No.: |
15/981761 |
Filed: |
May 16, 2018 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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15206060 |
Jul 8, 2016 |
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15981761 |
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14188257 |
Feb 24, 2014 |
9387082 |
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15206060 |
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13418294 |
Mar 12, 2012 |
8679190 |
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14188257 |
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12148534 |
Apr 17, 2008 |
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13418294 |
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12070336 |
Feb 15, 2008 |
8821583 |
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12148534 |
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11243952 |
Oct 4, 2005 |
7857849 |
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12070336 |
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11636114 |
Dec 7, 2006 |
7857447 |
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12070336 |
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11409218 |
Apr 20, 2006 |
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12070336 |
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11639049 |
Dec 13, 2006 |
7909867 |
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12070336 |
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60923988 |
Apr 17, 2007 |
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60901805 |
Feb 16, 2007 |
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60616262 |
Oct 5, 2004 |
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60673172 |
Apr 20, 2005 |
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60843942 |
Sep 11, 2006 |
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60783307 |
Mar 17, 2006 |
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60673600 |
Apr 21, 2005 |
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60843942 |
Sep 11, 2006 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61F 2/3872 20130101;
A61F 2210/0061 20130101; A61L 27/50 20130101; A61F 2/4202 20130101;
A61F 2002/30754 20130101; A61F 2/34 20130101; A61F 2/36 20130101;
A61F 2002/30075 20130101; A61L 2430/24 20130101; A61F 2310/00293
20130101; A61F 2220/005 20130101; A61L 27/18 20130101; A61F
2002/30593 20130101; A61F 2/30756 20130101; A61F 2/4405 20130101;
A61F 2/4241 20130101; A61F 2/3603 20130101; A61F 2002/30448
20130101; A61L 27/34 20130101; A61L 27/52 20130101; A61L 27/3843
20130101; A61F 2/4225 20130101 |
International
Class: |
A61F 2/30 20060101
A61F002/30; A61L 27/52 20060101 A61L027/52; A61F 2/36 20060101
A61F002/36; A61F 2/38 20060101 A61F002/38; A61F 2/42 20060101
A61F002/42; A61F 2/44 20060101 A61F002/44; A61L 27/18 20060101
A61L027/18; A61L 27/34 20060101 A61L027/34; A61F 2/34 20060101
A61F002/34; A61L 27/50 20060101 A61L027/50; A61L 27/38 20060101
A61L027/38 |
Claims
1. (canceled)
2. An orthopedic implant comprising a hydrogel and having a first
surface and a second surface, the hydrogel having an equilibrium
water content of at least 15% and a modulus of at least about 1
MPa, and the first surface having a coefficient of friction in
aqueous solution of less than 0.2.
3. The orthopedic implant of claim 2, wherein the orthopedic
implant is at least 1 mm in thickness.
4. The orthopedic implant of claim 2, wherein the first and second
surfaces are different.
5. The orthopedic implant of claim 2, wherein the implant is
configured for cartilage replacement in a joint.
6. The orthopedic implant of claim 5, wherein the joint is selected
from a hip, a shoulder, a knee, an elbow, a finger, a joint of the
hand and a joint of the foot.
7. The orthopedic implant of claim 2, wherein the implant is a
cartilage replacement implant and wherein the first surface is a
bearing surface that is adapted to articulate with an additional
bearing surface upon implantation.
8. The orthopedic implant of claim 7, wherein the additional
bearing surface is a natural cartilage surface or an implanted
device surface.
9. The orthopedic implant of claim 7, wherein the second surface is
adapted for anchoring the implant to bone.
10. The orthopedic implant of claim 7, wherein the second surface
is adapted for anchoring to an acetabular cavity, a tibial plateau,
an inner aspect of a glenoid, a femoral head, or a distal
femur.
11. The orthopedic implant of claim 7, wherein the first surface
has a primarily concave shape and the second surface has a
primarily convex shape.
12. The orthopedic implant of claim 11, wherein the primarily
convex shape of the second surface is adapted to mate with an
acetabular cavity or an inner aspect of a glenoid.
13. The orthopedic implant of claim 7, wherein the first surface
has a primarily convex shape and the second surface has a primarily
concave shape.
14. The orthopedic implant of claim 13, wherein the primarily
concave shape of the second surface is adapted to mate with a
femoral head or distal femur.
15. The orthopedic implant of claim 9, wherein the second surface
is adapted to interact with adjacent bone to allow for anchoring
via osteointegration over time.
16. The orthopedic implant of claim 2, wherein the hydrogel has a
permeability coefficient ranging from 1 e.sup.-18 to 1 e.sup.-12
m.sup.4/Nsec.
17. The orthopedic implant of claim 2, wherein the hydrogel is
immersed in an aqueous salt solution having a neutral pH.
18. The orthopedic implant of claim 2, wherein the hydrogel is
negatively charged at neutral pH.
19. The orthopedic implant of claim 2, wherein the hydrogel
comprises sulfonic acid groups.
20. An arthroplasty procedure, wherein an orthopedic implant in
accordance with claim 2 is implanted in a joint.
21. The method of claim 20, wherein the orthopedic implant is
implanted on one side of the joint in a hemi-arthroplasty
procedure.
22. The method of claim 20, wherein the orthopedic implant is
implanted on both sides of the joint in total arthroplasty
procedure.
23. The method of claim 22, wherein the joint is selected from a
hip, knee, shoulder, elbow, a finger, a joint of the hand, or a
joint of the foot.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation of U.S. application Ser.
No. 15/206,060, filed Jul. 8, 2016, which is a continuation of U.S.
patent application Ser. No. 14/188,257, filed Feb. 24, 2014, now
U.S. Pat. No. 9,387,082, which is a continuation of U.S. patent
application Ser. No. 13/418,294, filed Mar. 12, 2012, now U.S. Pat.
No. 8,679,190, which is a continuation of U.S. patent application
Ser. No. 12/148,534, filed Apr. 17, 2008, now abandoned, which
claims the benefit of U.S. Provisional Patent Application No.
60/923,988, filed Apr. 17, 2007, and is a continuation-in part of
U.S. patent application Ser. No. 12/070,336, filed Feb. 15, 2008,
now U.S. Pat. No. 8,821,583, the disclosures of all of which are
incorporated herein by reference.
[0002] U.S. patent application Ser. No. 12/070,336 claims priority
from U.S. Provisional Application No. 60/901,805, filed Feb. 16,
2007, and is a continuation-in part of U.S. patent application Ser.
No. 11/243,952, filed Oct. 4, 2005, now U.S. Pat. No. 7,857,849,
which claims the benefit of U.S. Provisional Application Nos.
60/616,262, filed Oct. 5, 2004 and 60/673,172, filed Apr. 20,
2005.
[0003] U.S. patent application Ser. No. 12/070,336 is also a
continuation-in-part of U.S. application Ser. No. 11/636,114, filed
Dec. 7, 2006, now U.S. Pat. No. 7,857,447, which claims the benefit
of U.S. Provisional Application Nos. 60/843,942, filed Sep. 11,
2006, and 60/783,307, filed Mar. 17, 2006.
[0004] U.S. patent application Ser. No. 12/070,336 is also a
continuation-in-part of U.S. application Ser. No. 11/409,218, filed
Apr. 20, 2006, now abandoned, which claims the benefit of U.S.
Provisional Application No. 60/673,600, filed Apr. 21, 2005.
[0005] U.S. patent application Ser. No. 12/070,336 is also a
continuation-in-part of U.S. application Ser. No. 11/639,049, filed
Dec. 13, 2006, now U.S. Pat. No. 7,909,867, which claims the
benefit of U.S. Provisional Application No. 60/843,942, filed Sep.
11, 2006.
FIELD
[0006] The present invention relates generally to interpenetrating
polymer network hydrogels. More particularly, the present invention
relates to devices and materials useful for orthopaedic
prostheses.
BACKGROUND
[0007] With disease or damage, the normally smooth, lubricious
cartilage covering joint surfaces progressively deteriorates,
exposing bone and leading to arthritic pain that is exacerbated by
activity and relieved by rest. Today, patients with osteoarthritis
are faced with only one of two choices: either manage their pain
medically, or undergo an effective but highly bone-sacrificing
surgery. Medical management includes weight loss, physical therapy,
and the use of analgesics and nonsteroidal anti-inflammatories.
These can be effective at reducing pain but are not curative. Other
options include drugs like glucosamine or hyaluronan to replace the
"lost" components of cartilage, but despite their extensive use in
the U.S., their efficacy is still questioned. When medical
intervention fails and a patient's joint pain becomes unbearable,
surgery is advised. Total joint arthroplasty is a surgical
procedure in which the diseased parts of a joint are removed and
replaced with new, artificial parts (collectively called the
prosthesis). In this highly effective but invasive procedure, the
affected articular cartilage and underlying subchondral bone are
removed from the damaged joint. A variety of replacement systems
have been developed, typically comprised of ultra-high molecular
weight polyethylene (UHMWPE) and/or metals (e.g. titanium or cobalt
chrome), or more recently, ceramics. Some are screwed into place;
others are either cemented or treated in such a way that promotes
bone ingrowth. These materials have been used successfully in total
joint replacements, providing marked pain relief and functional
improvement in patients with severe hip or knee osteoarthritis.
[0008] A large number of patients undergo total hip arthroplasty
(THA) in the U.S. each year, which involves implanting an
artificial cup in the acetabulum and a ball and stem on the femoral
side. The goals of THA are to increase mobility, improve hip joint
function, and relieve pain. Typically, a hip prosthesis lasts for
at least 10-15 years before needing to be replaced. Yet despite its
success as a surgical procedure, THA is still considered a
treatment of last resort because it highly "bone-sacrificing,"
requiring excision of the entire femoral head. It is this major
alteration of the femur that often makes revision replacement
difficult. While this procedure has a survival rate of 90% or more
in the elderly (who usually do not outlive the implant), implant
lifetimes are significantly shorter in younger, more active
patients. As a result, younger patients face the prospect of
multiple, difficult revisions in their lifetime. Revisions are
required when implants exhibit excessive wear and periprosthetic
bone resorption due to wear particles, as well as aseptic loosening
of the prosthesis resulting from stress shielding-induced bone
resorption around the implant.
[0009] The aforementioned limitations of THA have prompted the
industry to seek less bone-sacrificing options for younger
patients, with the hope that a THA can be postponed by at least
five years or more. One approach towards improving treatment has
been to develop less invasive surgical procedures such as
arthroscopic joint irrigation, debridement, abrasion, and
synovectomy. However, the relative advantage of these surgical
techniques in treating osteoarthritis is still controversial. An
alternative to THA is hip "resurfacing," has now re-emerged because
of new bearing surfaces (metal-on-metal, rather than
metal-on-polyethylene). While many patients can expect to outlive
the procedure's effectiveness, hip resurfacing preserves enough
bone stock on the femoral side to allow for later total hip
replacement. Unfortunately, there are enough potential drawbacks
that doctors offering hip resurfacing say that the procedure should
still be deferred as long as possible. In metal-on-metal
resurfacing, the femoral head is shaped appropriately and then
covered with a metal cap that is anchored by a long peg through the
femoral neck. It requires a more precise fit between the cap and
cup, and the procedure generally sacrifices more bone from the
acetabulum compared to conventional replacements due to the larger
diameter of the femoral component. Furthermore, a resurfacing
operation has a steep learning curve and takes longer than a THA.
Femoral neck fractures caused by bone resorption around the peg
have been reported, and the long-term impact of metal ion release
from the bearing surfaces is also not yet known in humans. As a
result of these complications, today's resurfacing devices are
still only indicated in patients for whom hip pain is unbearable,
as is the case for THA.
[0010] The present invention addresses the needs in the art and
provides an interpenetrating polymer network hydrogel that is
strain-hardened by means of swelling that forms the basis of an
arthroplasty device and a method for making this device.
SUMMARY OF THE DISCLOSURE
[0011] The present invention provides a bone-sparing arthroplasty
device based on an interpenetrating polymer network hydrogel that
is strain-hardened by means of swelling that mimics the molecular
structure, and in turn, the elastic modulus, fracture strength, and
lubricious surface of natural cartilage. Emulating at least some of
these structural and functional aspects of natural cartilage, the
hydrogel forms the basis of a novel, bone-sparing, "biomimetic
resurfacing" arthroplasty procedure. Designed to replace only
cartilage, this material is fabricated as a set of flexible,
implantable devices featuring lubricious articular surfaces and
osteointegrable bone-interfaces. In principle, the device can be
made for any joint surface in the body. For example, a device to
cover the tibial plateau will require an analogous bone-preparation
and polymer-sizing process. For a device to cover the femoral head
in the hip joint, the analogy to a male condom is appropriate in
which a cap shaped hydrogel device fits snugly over the contours of
the femoral head. For a device to line the acetabulum, the analogy
to a female condom is appropriate. A polymer dome stretches over
the lip and can be snapped into place to provide a mating surface
with the femoral head. In this way, both sides of a patient's hip
joint can be repaired, creating a cap-on-cap articulation. However,
if only one of the surfaces is damaged, then only one side can be
capped, creating a cap-on-cartilage articulation. To create a
cap-shaped hydrogel device for the shoulder joint (also a
ball-and-socket joint), a process similar to that of the hip joint
is used. For instance, a "female condom" can be created to line the
inner aspect of the glenoid. Furthermore, devices for other joints
in the hand, fingers, elbow, ankles, feet, and intervertebral
facets can also be created using this "capping" concept. In one
embodiment in the distal femur, the distal femur hydrogel device
volume follows the contours of the bone while sparing the anterior
and posterior cruciate ligaments.
[0012] More specifically, the present invention provides an
arthroplasty device having an interpenetrating polymer network
hydrogel that is strain-hardened by swelling and is adapted to be
held in place in a mammalian joint by conforming to a naturally or
artificially prepared geometry of a bone in the mammalian joint.
The strain-hardened interpenetrating polymer network hydrogel is
based on two different networks. The first network is a
non-silicone network of preformed hydrophilic non-ionic telechelic
macromonomers chemically cross-linked by polymerization of its
end-groups. The second network is a non-silicone network of
ionizable monomers. The second network has been polymerized and
chemically cross-linked in the presence of the first network and
has formed physical cross-links with the first network. Within the
interpenetrating polymer network, the degree of chemical
cross-linking in the second network is less than the degree of
chemical cross-linking in the first network. An aqueous salt
solution having a neutral pH is used to ionize and swell the second
network in the interpenetrating polymer network. The swelling of
the second network is constrained by the first network, and this
constraining effect results in an increase in effective physical
cross-links within the interpenetrating polymer network. The
strain-induced increase in physical cross-links is manifested as a
strain-hardened interpenetrating polymer network with an increased
initial Young's modulus, which is larger than the initial Young's
modulus of either (i) the first network of hydrophilic non-ionic
telechelic macromonomers swollen in pure water or in an aqueous
salt solution, (ii) the second network of ionized monomers swollen
in pure water or in an aqueous salt solution, or (iii) the
interpenetrating polymer network hydrogel formed by the combination
of the first and second network swollen in pure water. The observed
increase in stiffness modulus as a result of strain (induced herein
by swelling) is caused by an increase in the number of physical
cross-links within the interpenetrating polymer network. For the
purposes of the present invention, strain-hardening is defined as
an increase in the number of physical cross-links and stiffness
modulus with applied strain.
[0013] The device arthroplasty has a bone-interfacing region and a
bearing region opposite to the bone-interfacing region. The
bone-interfacing region is characterized by conforming and capable
of fixating to the naturally or artificially prepared geometry of
the bone in the mammalian joint.
[0014] The device and strain-hardened interpenetrating polymer
network hydrogel of the present invention could be varied according
to the following embodiments either by themselves or in any
combinations thereof. For example, the device can be implanted on
one side of the mammalian joint forming a hydrogel-on-cartilage
articulation in the mammalian joint. The device could further have
a second mating component (i.e. another arthroplasty device as
taught in this invention) implanted on the opposing joint surface
from the implanted device forming a hydrogel-on-hydrogel
articulation. The bone-interfacing region is capable of binding to
calcium-containing and phosphate-containing bone-matrix
constituents of the bone. In another example, the bone-interfacing
region is characterized by having a porosity or surface roughness
on the order of 10 to 1000 microns to accommodate bone formation.
The bone-interfacing region could also be pre-coated with
calcium-containing and phosphate-containing constituents. In still
another example, biomolecules could be chemically or physically
bonded to the bone-interfacing region.
[0015] Instead of having the bone-interfacing region be made of the
strain-hardened interpenetrating polymer network hydrogel, the
bone-interfacing region could, in one example, be made of a
polymeric material chemically bonded to the bearing region. In this
example, the bearing region is made of the strain-hardened
interpenetrating polymer network hydrogel. In another example, the
bearing region and the bone-interfacing region could have different
compositions at either side of the device and are physically or
chemically and physically integrated with each other within the
device.
[0016] An adhesive material (biodegradable or non-biodegradable)
could be bonded to the bone-interfacing region and would then be
capable of bonding the device via the bone-interfacing region to
the bone. In another example the device could include a
calcium-containing inorganic coating that is chemically or
physically bonded to the bone-interfacing region.
[0017] In still another example, it is a desire to approximately
match the thickness profile of the device to the natural thickness
profile of an original cartilage layer. The device can be adapted
to fit over a primarily convex or concave three-dimensional
bone-receiving surface. In one example, the device is undersized to
fit over a primarily convex bone-receiving surface to create an
elastic contraction fit over the convex three-dimensional
bone-receiving surface. The device is capable of swelling to a
swollen equilibrium volume in a fluid and temperature other than
body fluids and body temperature prior to implantation and capable
of de-swelling to a smaller equilibrium volume, compared to the
swollen equilibrium volume, upon implantation and exposure to body
fluids or/and body temperature, whereby at the smaller equilibrium
volume, the device contracts against or physically grips said
primarily convex three-dimensional bone receiving surface.
[0018] In another example, the device is oversized to fit against a
primarily concave three-dimensional bone-receiving surface to
accommodate an elastic expansion fit against the primarily concave
bone-receiving surface. The device is capable of at least partially
drying or de-swelling to a dried or de-swollen equilibrium volume
in a fluid and temperature other than body fluids and body
temperature prior to implantation and capable of swelling to a
larger equilibrium volume, compared to the dried or de-swollen
equilibrium volume, upon implantation and exposure to body fluids
and/or body temperature, whereby the larger equilibrium volume
expands the device against a primarily concave three-dimensional
bone receiving surface.
[0019] The hydrophilic non-ionic macromonomer in the first network
has a molecular weight between about 275 Da to about 20,000 Da,
about 1000 Da to about 10,000 Da, or about 3000 Da to about 8000
Da. In another example, the molar ratio between the ionizable
monomers and the hydrophilic non-ionic telechelic macromonomers is
greater than or equal to 1:1 or greater than 100:1. In one example,
the hydrophilic non-ionic telechelic macromonomer in the first
network is a derivative of poly(ethylene glycol), and the ionizable
monomers are acrylic acid monomers.
[0020] In still another example, the aqueous salt solution has a pH
in the range of about 6 to 8. In still other examples, the first
network has at least about 50%, at least 75% or at least 95% by dry
weight telechelic macromonomers. In still another example, the
first network has hydrophilic monomers grafted onto the first
network. In still another example, the second network further has
hydrophilic macromonomers grafted onto the second polymer network.
In still another example, the strain-hardened interpenetrating
polymer network hydrogel has a tensile strength of at least about 1
MPa. In still another example, the strain-hardened interpenetrating
polymer network hydrogel has an initial equilibrium tensile modulus
of at least about 1 MPa. In still another example, the
strain-hardened interpenetrating polymer network hydrogel has an
equilibrium water content of at least 25%, 35% or 50%. In still
another example, the strain-hardened interpenetrating polymer
network hydrogel is permeable to the aqueous salt solution and the
hydrogel has a permeability coefficient ranging from 1e-17 to 1e-13
m4/Nsec.
[0021] In still another example, the coefficient of friction of the
bearing region of the strain-hardened interpenetrating polymer
network hydrogel in an aqueous solution is less than 0.2. In still
another example, one side of the device is modified with another
polymeric material, other functional groups, or biomolecules using
bifunctional crosslinkers. In one example, the biomolecules could
be used to stimulate bone cell growth and/or adhesion. In yet
another example, the device is comprised of stimulus-responsive
polymeric materials that allow it to shrink or swell to conform to
the convexity or concavity of an adjacent joint surface.
BRIEF DESCRIPTION OF THE FIGURES
[0022] The present invention together with its objectives and
advantages will be understood by reading the following description
in conjunction with the drawings, in which:
[0023] FIG. 1 shows a schematic of the device and anatomical
structures according to an embodiment of the invention. The device
has two components, one version 1 that is placed on the primarily
convex bone side 3 of the joint and another version 2 that is
placed on the primarily concave bone side 4. The bone interface
regions 6 secure bone integration and adhesion. The bearing regions
5 possess a low coefficient of friction and allow for smooth
relative sliding and rolling motion between the two components and
are made of a strain-hardened interpenetrating polymer network
hydrogel of a end-linked first network 10, an ionized second
network 11, and an aqueous salt solution 12.
[0024] FIG. 2 shows a schematic of a cross-section of the device
according to an embodiment of the invention, showing the bearing
region 5 of thickness A and the bone-interfacing region 6 of
thickness C that are integrated by a transition zone 7 of thickness
B. The bearing 5 and bone-interfacing 6 regions could have the same
or different materials, while dimensions A, B, and C vary based on
the materials and device specifications.
[0025] FIG. 3 shows a schematic of an anchoring strategy according
to an embodiment of the invention for a convex (left column, A1-A3)
and a concave (right column, B1-B3) joint surface. An adhesive
layer could initially anchor the hydrogel to bone, but as it
calcifies and allows new bone to grow in, hydroxyapatite binds to
the bone interface region via the intervening scaffold to yield a
calcified bone interface that mimics that found in natural
cartilage.
[0026] FIG. 4 shows according to an embodiment of the invention how
the inorganic constituents of bone 3,4 (calcium and phosphate) can
interact with the bone-interface region of an IPN hydrogel 1,2. In
one embodiment, the carboxylic acid groups on the second network 11
(e.g. poly(acrylic acid)) interact and form complexes with the
divalent calcium ions and negatively charged phosphate ions.
[0027] FIGS. 5A and 5B shows according to an embodiment of the
invention a hip arthroplasty procedure. FIG. 5A shows a dislocated
joint exposing the acetabulum 4a and the femoral head 3a. A male
hydrogel device component 1a is placed on the femoral head 3a and
held in place by means of a stretch-to-fit. Similarly, the
acetabulum device component 2a is placed in the acetabulum bone 4a
and held in place by means of an expansive press-fit. FIG. 5B shows
that after the components are implanted in place, the joint is
reduced.
[0028] FIGS. 6A-6C shows according to an embodiment of the
invention a three dimensional version of the hip arthroplasty. FIG.
6A shows a lateral view of the femoral head hydrogel device
component 1a; a recess 103 that accommodates bone vessels is also
shown. FIG. 6B depicts the femoral head bone 3a and a cross section
of the femoral head device component 1a. FIG. 6C depicts the
acetabulum device component 2a.
[0029] FIG. 7 shows according to an embodiment of the invention a
two-sided (total) or one-sided hemi-arthroplasty. In this
embodiment, the femoral device component 1a is stretched over the
femoral head bone 3a while the acetabulum component 2a is press-fit
in the acetabulum recess 4a. The bone interface regions 6 are
porous and coated with hydroxyapatite to ensure bone ingrowth and
the bearing regions 5 have lubricious properties to facilitate
relative sliding. Furthermore, a depression 100 in the acetabulum
component 4a is present that forms a chamber 101 that is filled
with pressurized synovial fluid 102; the chamber is sealed by the
two device components 1a, 2a.
[0030] FIG. 8 shows according to an embodiment of the invention the
hydrogel device applied to the knee. The distal femur device
component 1b is placed on the distal femur bone 3b like a tight
sock. The device holds openings or recesses for the ligaments; as
such, a lateral opening 110 accommodates the lateral ligament while
a central opening 111 accommodates the cruciate ligaments. The
distal femur device component 1b is initially held in place by
means of tight fit, further enhanced by a hydrogel stimulation
process that is disclosed hereafter. The tibial plateau hydrogel
device component 2b in this embodiment has two distinct parts, one
for the lateral facet and one for the medial facet. The hydrogel
device components hold a porous bone interfacing region 6 that
allows for bone ingrowth to secure fixation.
[0031] FIGS. 9A-9C shows according to an embodiment of the
invention the hydrogel device application to the tibial plateau 4b.
FIG. 9A shows a lateral cross sectional view of the tibial plateau
4b and the facet 112. FIG. 9B shows the depression 113 surgically
made by means of punching the bone; it further depicts the hydrogel
device component 2b before implantation. FIG. 9C shows the tibial
hydrogel device component 2b inserted in the depression of the
facet 113.
[0032] FIG. 10 shows according to an embodiment of the invention a
structure of an interpenetrated polymer network based on an
end-linked macromonomer network 10 and an ionized, monomer-based
network 11 which is swollen and osmotically pre-stressed with a
buffered, aqueous salt solution 12.
[0033] FIGS. 11A-11C shows according to an embodiment of the
invention the steps for synthesis of the IPN hydrogel.
[0034] 1. The starting material for the hydrogel is a solution of
telechelic macromonomers 13 with reactive functional end groups 15
dissolved in water 16. The telechelic macromonomers are polymerized
to form a first end-linked polymer network 10 swollen in water
16.
[0035] 2. Hydrophilic, ionizable monomers 14 mixed with water are
added to the first polymer network 10 along with a photoinitiator
and a crosslinking agent (not shown). The hydrophilic, ionizable
monomers are then photopolymerized and cross-linked in the presence
of first polymer network 10 to form the second polymer network 11
in the presence of the first. This results in formation of an IPN
hydrogel having an end-linked polymer network 10 interpenetrated
with a ionizable second network 11 swollen in water 16.
[0036] 3. The water-imbibed IPN is then immersed in an aqueous
salt-containing solution 12 at a typical pH of 7.4 and is swollen
to equilibrium, yielding a simultaneous increase in both the water
content and the stiffness modulus of the IPN. This IPN swollen in
the aqueous salt solution 12 has a higher tensile elastic modulus
compared to the IPN swollen in pure water 16 due to strain
hardening induced by swelling of the second network 11 within the
constraint posed by the highly crosslinked first network 10.
[0037] FIG. 12A shows according to an embodiment of the present
invention method steps of how an IPN is prepared after monomers 17
are used to make the first network 10. Exposure to UV light in the
presence of a photoinitiator and crosslinker (not shown) leads to
polymerization and crosslinking to form a network 10, depicted by
the transition from (i) to (ii). In (iii) to (iv), the first
network is swollen with the second network precursor monomers 14, a
crosslinking agent (not shown) and a photoinitiator (not shown).
Exposure to UV light initiates polymerization and cros slinking of
the second network 11 in the presence of the first (10) to form the
IPN.
[0038] FIG. 12B shows according to an embodiment of the present
invention method steps of how an IPN is prepared after
macromonomers 13 with reactive endgroups 15 are used to form a
first network 10 in the presence of an existing second network 11
or linear macromolecules and/or biomacromolecules. A mixture of the
first and second polymeric components is made, and then the
telechelic macromonomers 13, 15 are reacted under UV light to form
the first network 10 in the presence of the second 11. If the
second network 11 is crosslinked chemically, then it is a fully
interpenetrating network. If it is not (and only physically
crosslinked), then it is a semi-interpenetrating network.
[0039] FIG. 12C shows according to an embodiment of the present
invention method steps of how an IPN is formed from a first network
10 based on monomers 17 and a second network 11 or linear
macromolecules and/or biomacromolecules. A mixture of the monomers
17 and macromolecules is made, and then the monomers are reacted
under UV light to form the first network in the presence of the
second 11. If the second network 11 is crosslinked chemically, then
it is a fully interpenetrating network. If it is not (and only
physically crosslinked), then it is a semi-interpenetrating
network.
[0040] FIG. 13 shows according to an embodiment of the present
invention a schematic of the synthesis of telechelic PEG-diacrylate
from a PEG-diol macromonomer. To generate PEG-dimethacrylate,
methacryloyl chloride would be reacted with the PEG-diol instead of
acryloyl chloride.
[0041] FIG. 14 shows according to an embodiment of the present
invention a schematic of the synthesis of telechelic
PEG-diacrylamide from a PEG-diol macromonomer. To generate
PEG-dimethacrylamide, methacryloyl chloride would be reacted with
the PEG-diol instead of acryloyl chloride.
[0042] FIG. 15 shows according to an embodiment of the present
invention a schematic of the synthesis of telechelic PEG-allyl
ether from a PEG-diol macromonomer.
[0043] FIGS. 16A-16D shows according to embodiments of the present
invention: (FIG. 16A) an IPN with a first network(10 and second
network 11 based on two different polymers, (FIG. 16B) an IPN with
a graft-copolymer 29 attached to the first network 10 and a
homopolymer in the second network 11, (FIG. 16C) an IPN with a
homopolymer in the first network 10 and a graft-copolymer 30 in the
second network 11, and (FIG. 16D) an IPN with graft-copolymers (29,
30 in both the first and the second networks 10, 11.
[0044] FIGS. 17A-17D shows according to the present invention the
mechanical behavior of a PEG(3.4 k)/PAA IPN prepared with 70%
volume fraction of acrylic acid in the second network: (FIG. 17A)
stress-strain profile under tension, (FIG. 17B) stress-strain under
confined compression, (FIG. 17C) stress-strain profile unconfined
compression, and (FIG. 17D) strain versus time in a tensile creep
experiment.
[0045] FIG. 18A shows according to an embodiment of the present
invention true stress-true strain curves for PEG(8.0 k)/PAA IPN,
PEG(8.0 k)-PAA copolymer, PEG(8.0 k), and PAA networks. FIG. 18BB
shows according to an embodiment of the present invention
normalized true stress-true strain curves for PEG(8.0 k)/PAA IPN,
PEG(8.0 k)-PAA copolymer, PEG(8.0 k), and PAA networks.
[0046] FIG. 19A shows according to an embodiment of the present
invention the effect of the mass fraction of acrylic acid (AA)
monomer in the second network precursor solution on the volume
change in the resultant IPN. The vertical dotted line indicates the
point of equimolar amounts of AA and ethylene glycol (EG) monomer
units in the IPN, while the horizontal dotted line indicates where
the PEG network and the PEG/PAA IPN have the same volume.
[0047] FIG. 19B shows according to an embodiment of the present
invention the dependence of the fracture stress and Young's modulus
of the PEG/PAA IPN on the mass fraction of AA in the IPN. The
vertical dotted line indicates the point of equimolar amounts of AA
and ethylene glycol (EG) monomer units in the IPN.
[0048] FIG. 20 shows according to an embodiment of the present
invention time-dependence of the water content of single network
PEG(8.0 k) hydrogels and PEG(8.0 k)/PAA IPNs with different amounts
of acrylic acid (AA) at the time of polymerization. The hydrogels
were placed in deionized water in the dry state at time=0 and then
weighed at regular intervals.
[0049] FIG. 21 shows according to an embodiment of the present
invention true stress versus true strain curves of the PEG(4.6
k)/PAA IPN in PBS and deionized water, as well as the PEG and PAA
single networks in PBS and deionized water. The PEG(4.6 k) network
is unaffected by the change from water to PBS. The arrow indicates
the shift in the stress-strain profile of the IPN after it has been
strain-hardened by swelling to equilibrium in PBS.
[0050] FIG. 22 shows according to an embodiment of the present
invention the stress-strain profiles of PEG(4.6 k)/PAA IPNs
prepared with three different combinations of crosslinker chemical
end-groups but the same formulations of PEG (MW 4.6 k, 50% by
weight in water) and AA (50% v/v in water) as well as the same
polymerization conditions (photoinitiator and crosslinker
concentration by mole and UV intensity) and swelling conditions
(PBS at pH 7.4). Specimen (A) was prepared from PEG-diacrylamide
first network and a PAA second network crosslinked with
N,N'-(1,2-dihydroxyethylene) bisacrylamide. Specimen (B) was
prepared from PEG-diacrylamide first network and a PAA second
network crosslinked with triethylene glycol dimethacrylate.
Specimen (C) was prepared from PEG-diacrylate first network and a
PAA second network crosslinked with triethylene glycol
dimethacrylate.
[0051] FIG. 23A shows according to the invention SEM of a plain
PEG/PAA sample (without hydroxyapatite) showing fractured edge
(dark) and top surface (light). FIG. 23B shows according to the
invention SEM of a hydroxyapatite-coated PEG/PAA sample showing
fractured edge (dark) and top surface (light). FIG. 23C shows
according to the invention energy-dispersive X-ray spectroscopy
(EDX) analysis of the hydroxyapatite-coated PEG/PAA IPN (inset),
showing a Ca/P ratio of roughly 1.5-1.6, similar to that of HAP,
with an inset showing a high-magnification SEM image of HAP-coated
PEG/PAA. FIG. 23D shows according to the invention osteoblast-like
cells growing on PEG/PAA hydrogel coated with 200-nm diameter
HAP
[0052] FIGS. 24A-24C shows according to the invention SEMs of
hydroxyapatite coatings of differing diameter (5 .mu.m, .about.200
nm, and 20 nm) on bare silica (FIG. 24A) and on PEG/PAA IPNs (at
low magnification in FIG. 24B and at high magnification in FIG.
24C).
[0053] FIG. 25A shows according to the invention a bonding process
for an IPN hydrogel 10, 11 bonding to bone (convex 3 or concave 4)
through an intervening polymeric adhesive based on monomers 18. The
monomers react when exposed to UV, photoinitiator, and crosslinker
to form a third network 19 that is physically or physically and
chemically crosslinked to the IPN hydrogel and to bone.
[0054] FIG. 25B shows according to the invention a bonding process
of an IPN hydrogel 10, 11 bonding to bone 3, 4 through an
intervening polymer adhesive based on macromonomers 21 with
reactive end-groups 20. The macromonomers react to form a third
macromonomeric network 22 that is physically or physically and
chemically crosslinked to the IPN hydrogel and to bone.
[0055] FIG. 26 shows according to the present invention a
semi-interpenetrating network in which one of the networks acts as
the anchoring intervening polymer. Telechelic macromonomers 13 with
reactive end-groups 15 and physical network 11 or solution of
linear chains are mixed together and cast over a bone surface 3, 4
that is pre-coated and/or functionalized with UV-sensitive
crosslinkable groups 23. Exposure to an initiating source (e.g. UV
light) in the presence of a photoinitiator leads to free-radical
polymerization and crosslinking of these crosslinkable groups on
both the telechelic macromonomers and the coated/functionalized
bone surface. The result of free-radical polymerization and
crosslinking is shown on the right. The ends of the telechelic
macromonomers have formed a network 10 and have copolymerized and
bonded with the surface of the bone. The linear second network
polymers are physically trapped within this first network, forming
a second, physically crosslinked network 11 interpenetrating the
first chemically crosslinked network.
[0056] FIG. 27A shows according to an embodiment of the invention a
fully interpenetrating network in which a third network is
partially interpenetrated within the pre-existing IPN by
interdiffusion of the third network monomer 24 for a predetermined
time and then polymerizing the monomer in the presence of the IPN
10, 11. This yields what is effectively a third network 25 on one
side of the IPN hydrogel, which may have different properties than
the other side, and are properties that may be useful as a
bone-interface region.
[0057] FIG. 27B shows according to an embodiment of the invention a
fully interpenetrating network in which the second network monomer
14 is interfacially copolymerized with another monomer 26 that when
polymerized acts as the bone-interfacing material. A pre-existing
first network is swollen with the precursor monomers of a second
network. At the bone-interface side of the material is a precursor
solution of another reactive monomer 26. These monomers partially
penetrate the matrix of the first network. Upon exposure to UV, the
monomers co-polymerize, yielding a material with a one type of IPN
10, 11 on the bearing side and another type of IPN (10, 27 on the
bone-interfacing side.
[0058] FIG. 27C shows according to an embodiment of the invention
in which an external stimulus is used to create a composition
gradient in the second network within the first network of the IPN.
A mixture of acrylic acid and non-ionic monomers (e.g. acrylamide,
N-isopropylacrylamide, or hydroxylethylacrylate monomers) is used.
The first network 10 is soaked in a solution of ionizable monomer
14, non-ionic monomer 28, crosslinker and photoinitiator (not
shown) and then an electric field is applied to the gel. Only the
ionizable monomers will move along the electric field due to their
charge. After formation of a ionizable monomer concentration
gradient, the gel is exposed to UV and the gradient is fixed via
second network gel formation. The result is an IPN hydrogel with a
second network localized to the bearing region and a non-ionic
second network localized to the bone-interface region. Thus, when
polyacrylic acid is the ionizable monomer 14, the polyacrylic acid
concentration is at a maximum at the bearing region with the
concentration of polyacrylic acid decreasing with distance from the
bearing region towards the bone-interface region. Thus in FIG. 27C,
when the first network 10 is polyurethane, the polyurethane is
present throughout the prosthesis.
[0059] FIG. 28 shows according to an embodiment of the invention
two examples of other device surface modification strategy. This
strategy involves the acrylation/methacrylation of an
amine-containing or hydroxyl-containing molecule or biomolecule by
reaction with a halogenated (active) acid (e.g. acryloyl chloride)
(Reaction A) or with an active ester (e.g.
acryloxy-N-hydroxysuccinimide) (Reaction B) to make it capable of
copolymerizing with the precursor of one of the networks in the
device. The R-group in the these reaction schemes can be any
amine-containing or hydroxyl-containing synthetic chemical or
polymer, proteins, polypeptides, growth factors, amino acids,
carbohydrates, lipids, phosphate-containing moieties, hormones,
neurotransmitters, or nucleic acids.
[0060] FIG. 29 shows according to an embodiment of the invention a
heterobifunctional crosslinker 118 containing two endgroups 115,
117 joined by a spacer 116 that are used to covalently attach
molecules, macromolecules, and biomolecules 114 to IPN hydrogel
surfaces 119.
[0061] FIG. 30 shows according to an embodiment of the invention
methods steps to attain a different surface chemistry at the
bone-interface than that present in the bearing region. This
approach involves activating the functional groups on the surface
of the hydrogel followed by reaction of these activated function
groups with amine-containing or hydroxyl-containing molecules,
macromolecules, or biomolecules. In a preferred embodiment, the
carboxylic acid groups on poly(acrylic acid) within an IPN are
activated to form an active ester, which subsequently forms an
acrylamide linkages when reacted with an amine-containing or
hydroxyl-containing molecule, macromolecule, or biomolecules.
[0062] FIG. 31 shows specific examples of the method shown in FIG.
30 in which carboxylic acid functional groups on the hydrogel are
activated and subsequently reacted with dopamine hydrochloride to
yield a dopamine-conjugated surface. In Reaction A, a PEG/PAA
hydrogel is soaked in a solution of dicyclohexylcarbodiimide and
triethylamine in ethanol to activate the carboxylic acid groups
present on the PAA. Subsequent reaction with dopamine hydrochloride
and Triethylamine yields a dopamine-conjugated surface. In Reaction
B, the PEG/PAA hydrogel is soaked in solution of
N-hydroxysuccinimide and
N-Ethyl-N'-(3-dimethylaminopropyl)carbodiimide in phosphate buffer
to activate the carboxylic acids in PAA. Subsequent reaction with
dopamine hydrochloride in DMF and triethylamine yields a
dopamine-conjugated hydrogel surface.
[0063] FIG. 32 shows an embodiment of the present invention in
which an external stimulus such as a change in pH, salt
concentration, electric field, or temperature causes the device,
after (A) placement on the bone, to (B) shrink to conform to the
contours of the convex-shaped bone it surrounds. Conversely,
stimulated swelling can be achieved as a result of a change in pH,
salt concentration, electric field, or temperature create an
expansile effect on a concave joint surface. Stimulus-responsive
polymers are incorporated into the bearing and/or bone-interfacing
region of the device by the methods described in the present
invention.
DETAILED DESCRIPTION
[0064] The present invention is a "biomimetic" bone-sparing
hydrogel arthroplasty device (FIG. 1) that is designed to overcome
the limitations of current joint replacement technologies. The
device is comprised of flexible implants made from a novel
cartilage-like hydrogel material that conform to the convex and
concave surfaces of mammalian joints in either a total arthroplasty
(both sides) or a hemi-arthroplasty (one side). The device has the
high compressive strength and lubricity necessary to serve as a
replacement for articular cartilage, intervertebral discs (lumbar
or cervical), bursae, menisci, and labral structures in the
body.
[0065] Illustrated in FIG. 1 are the key device and anatomical
structures of the present invention in a typical diarthroidal
joint. Most joints in the mammalian skeleton have a "male,"
primarily convex 3 cartilage surface and a "female," primarily
concave cartilage surface 4. In this embodiment, the arthroplasty
device is comprised of two components, one component (1) that fits
over the primarily convex bone surface 3 and another component 2
that fits inside the primarily concave surface 4. Each component of
the device holds a bearing surface 5 that comes to contact with the
opposing bearing surface 5 of another other component. Each
component of the device also holds a bone interfacing region 6 that
enables the fixation of the device on the bone. Depending on the
joint that the device is applied to, its shape can have a rather
flat or a rather curved form, for example a device to replace the
cartilage of the femoral head resembles a hemispherical cap while a
device to replace the cartilage of the tibial plateau may resemble
a shallow circular dish. In some cases, only one component of the
device can be implanted as a hemi-arthroplasty so that it
articulates with the natural cartilage that is left intact at the
other side of the joint.
[0066] This device concept can be applied to nearly any joint in
the body. For instance, the types of orthopaedic devices for which
this invention is potentially useful includes total or partial
replacement or resurfacing of the hip (femoral head and/or
acetabulum), the knee (the tibial, femoral, and/or patellar
aspect), shoulder, hands, fingers (e.g. carpometacarpal joint),
feet, ankle, and toes. It is also useful in replacement or repair
of intervertebral discs or facets. In the knee, the hydrogel can
also serve as a meniscus replacement or a replacement material for
the cartilage or bursae in any joint such the elbow or shoulder, or
the labrum in joints such as the hip and shoulder.
[0067] This device strategy is guided by the limitations of current
arthroplasty approaches, which are either highly bone-sacrificing
or limited to only the repair of focal defects. The hydrogel device
is put in place of damaged cartilage after the damaged cartilage
has been removed by the surgeon--cartilage remains may need to be
removed because subsequent overlying by the implant might cause
unwanted conditions that lead to the differentiation of the
remaining cartilage fibrous tissue.
[0068] The device itself is comprised of a "bearing" region 5 on
one side, and a "bone-interfacing" region 6, in which the former
articulates with another bearing surface (either another
arthroplasty device such as the present invention or natural
cartilage on an apposing joint surface) and the latter interacts
with underlying bone. FIG. 2 depicts the cross-sectional area of
the device's composition of matter, where one side contains the
bearing region and the adjacent side contains the bone-interfacing
region. The two regions can be comprised of the same material or
different material. In one embodiment, the two regions are
comprised of one and the same IPN hydrogel, while in another
embodiment, the bearing region is comprised of an IPN hydrogel and
the bone-interface region is comprised of another polymer that is
integrated with the IPN hydrogel in such a way that there is a
smooth transition zone 7 between the two materials. In one
embodiment, the bearing region is made from an IPN hydrogel and the
bone-interface region 6 of the hydrogel device 1,2 is made from a
polymer or such as polyurethane, silicone rubber, derivatives, or
combinations thereof (such as copolymers or interpenetrating
networks with other polymers such as hydrogels with good mechanical
properties that allow the device to stretch or compress in response
to loads and be physically held in place by tensile or compressive
stress on or by the adjacent bone. The relative thicknesses of the
two regions can be varied such that the bearing region can make up
either a large or small proportion of the volume of the device.
[0069] The device can be described as "biomimetic" (i.e. imitative
of a natural cartilage) in that it is comprised of a material that
mimics the structure and function of natural articular cartilage.
While natural cartilage is composed of a highly negatively charged
network of proteoglycans interpenetrating a neutral, rigid network
of collagen with a water content of about 75%. In a preferred
embodiment, the hydrogel is composed of a highly negatively charged
network of poly(acrylic acid) interpenetrating a neutral, rigid
hydrophilic, end-linked network of, for example, poly(ethylene
glycol) macromonomers, with a water content of at least 35% and up
to 90%, but preferably about 70%. Mimicking these structural
details is believed to be critical to the formation of a stiff, yet
highly lubricious bearing material that behaves like natural
cartilage. Other combinations of hydrophilic, end-linked
macromonomers and negatively charged second networks are possible.
PEG and PAA are arguably the two most biocompatible, hydrophilic
polymers available. For instance, PEG is known widely to be
resistant to protein adsorption and PAA has recently been shown to
have a protective role against macrophage activity in vivo.
Although PEG and PAA are conventionally weak individually, we have
developed a way to create "strain hardened" IPNs of these materials
that mimic the high mechanical strength and elastic modulus, high
water content, and low surface friction of natural cartilage. Like
natural cartilage, the high mechanical strength and modulus of the
hydrogel enable it to take up and distribute loads. At the same
time, its high water content and low surface friction enable it to
function as a slippery bearing surface, just like the nascent
tissue.
[0070] Another innovative aspect of the present invention is the
anchoring strategy (FIG. 3). A combination of physical, chemical,
and biological means can be used to anchor the device to bone. To
achieve physical anchoring, the bone interfacing region 6 of the
hydrogel device 1, 2 is made to be rough and porous to match the
micro-topography of either natural or artificially prepared (e.g.,
reamed) subchondral bone, which increases surface area and friction
at this interface to enhance the mechanical interlocking of the
bone by the device. In addition, the device is fabricated to
conform to natural convexities and concavities of a given joint
surface. As illustrated in FIGS. 3 B1-B3 for the case of a concave
joint structure 4 such as the acetabulum 4a, the device is
fabricated as a cap 2a to mate perfectly with or is slightly
oversized to create an expansive fit against the concavity. Also
possible is the presence of a "lip" around the outer edge of the
acetabulum component (4a) which creates a labrum-like structure
around the outer groove of the socket, which would further aid in
the positioning and anchoring of the device. As illustrated in
FIGS. 3 A1-A3 for the case of a convex joint structure 3 such as
the femoral head, the hydrogel device 1a is fabricated as a cap to
mate perfectly with or is slightly undersized to create a snug fit
over the convexity. To supplement the aforementioned physical means
to secure the hydrogel device 1 or 2, a number of strategies can be
used. First, the bone interfacing region 6 encourages adhesion to
the underlying bone, by methods that may include but are not
limited to (a) a roughened surface, (b) a porous surface, (c)
tethering the surface with cell adhesion-promoting biomolecules
(such as cadherins or integrins) or biomolecules (e.g. collagen,
Bone Morphogenetic Proteins (BMPs), bisphosphonates, and Osteogenic
Proteins OP-1, or osteopontin), (d) by surface coating with
osteoconductive substances (such as natural hydroxyapatite, calcium
sulfates or purified collagen), or (e) addition of a bonding agent
such as a cement or glue. Combinations of these are also possible.
The anchoring process is depicted in the other plots in FIG. 3.
[0071] In one embodiment, the bone-interface region 6 of the device
is prepared such that it interacts with the adjacent bone to allow
for anchoring via osteointegration over time. In a version of this
embodiment, illustrated in FIG. 4, the carboxylic acids in
poly(acrylic acid) 11 in a PEG/PAA IPN bone-interface region 6
forms complexes with calcium and phosphates in the bone 3 as it is
being remodeled. In another embodiment, the bone-interface region 6
comes precoated with calcium-containing inorganic constituents
(e.g. tricalcium phosphate or/and hydroxyapatite) prior to
implantation. In still another embodiment, another polymer material
serving as the bone-interface region anchors the device through
bone ingrowth and deposition and/or calcification. Thus, the
biological means of anchoring is accomplished through a calcified
layer. This sets the stage for continual bone growth and deposition
within the pores of bone interface region and, in turn, anchorage
of the device through a calcified, bio-artificial composite
interface. Osteointegration of the device with underlying bone may
enable it to move as one with the bone and function like cartilage
within the joint and provide better adhesion through continuous
bone remodeling.
[0072] The localized use of a curable adhesive that bonds the
hydrogel to the bone provides a chemical means to attain robust,
intraoperative anchoring. In one embodiment the adhesive can be a
dental or orthopedic adhesive such as cement (e.g. zinc carbocylate
cement), resin, glue or the like. This adhesive may be of one that
provides firm bonding between the bearing region of the device and
bone. The adhesive in cured form may be porous or non-porous and
may be biodegradable or non-biodegradable. In the case of a
degradable adhesive, the adhesive material is gradually broken down
as new bone is formed that binds to the bone interface region. This
degradation takes place over a period of about one to about twelve
weeks after being implanted to coincide with the time it takes for
new bone to form. In the case of a non-degradable adhesive, the
adhesive itself binds and interdigitates with bone even as it is
being remodeled.
[0073] In another embodiment, the bone interfacing region is made
in part from a non-hydrogel polymer such as polyurethane, silicone
rubber, or derivatives or combinations thereof (such as copolymers
or interpenetrating networks with other polymers such as hydrogels)
with good mechanical properties that allow the material to stretch
or compress in response to loads and be physically held in place by
tensile or compressive stress on or by the adjacent bone. Such a
composite material would have a lubricious hydrogel (such as
PEG/PAA) as the bearing region and the non-hydrogel polymer (such
as polyurethane or silicone-based materials) as the bone-interface
region.
[0074] One embodiment of the present invention is application as a
hip arthroplasty device. According to this embodiment, the
arthroplasty hydrogel device is comprised of a femoral head
component (1a) and an acetabulum component (2a) as shown in FIGS.
5A-5B, 6A-6C and 7. Both components are comprised of a PEG/PAA
interpenetrating network hydrogel with properties described in
Table 1 and made by processes described hereafter.
TABLE-US-00001 TABLE 1 PEG(3.4k)/PAA physical properties (averages)
in PBS, pH 7.4 Water Content 65% Tensile Modulus 12 MPa Tensile
Fracture Strength 12 MPa Aggregate Equilibrium 1.6 MPa Compressive
Modulus Unconfined Compressive Strength 18 MPa Hydraulic
Permeability (K) 2.4 .times. 10.sup.-14 m.sup.4/N/sec Dynamic
Coefficient of Friction 0.05 (gel-on-gel) Linear Wear Rate
(gel-on-gel) ~0.75 microns/3.0M cycles
[0075] The overall device geometry resembles the anatomy of natural
cartilage. The femoral head component 1a holds a cap shape and is
placed on the femoral head 3a bone after the later has been
surgically reamed to remove damaged cartilage and the superficial
bone layer. The femoral head component 1a bone interface region 6
has a radius of curvature that is slightly undersized compared to
the radius of curvature of the femoral head bone 3a; the femoral
component 1a can therefore be held in place by means of a tight fit
around the femoral head. More specifically, and by analogy to latex
condoms, the hydrogel device femoral head component 1a, being
slightly undersized than the bone it is mounted onto, is pulled
over the femoral head 3a and is held in place by tension generated
by stretching of the hydrogel device 1a material. Because the
femoral head component 1a material is stretchable, it can be
stretched to fit over the femoral head. In one version of this
embodiment, this cap shaped device 1a covers the bone 360 degrees
on the lateral plane and as much as 200 degrees on the coronal
plane. With the bone now occupying its inside space, the hydrogel
device femoral head component 1a cannot completely return to its
original dimensions, which causes the device 1a to "hug" the bone
3a it surrounds. The entire process can be facilitated by means of
a retractor tool that could open up the device 1a opening.
[0076] The acetabulum component 2a is placed on the acetabulum bone
4a after the later has been surgically reamed to remove damaged
cartilage and the superficial bone layer. The acetabulum hydrogel
device component 2a holds a hemispherical shell shape and its bone
interface region 6 has a radius of curvature that is slightly
oversized compared to the radius of curvature of the acetabulum
bone 4a socket; the acetabulum component 2a can be held in place by
means of a tight press-fit inside the acetabulum 4a. The hydrogel
device acetabulum component may also have a thickness profile that
matches that of natural acetabular cartilage and is in the range of
1 mm-5 mm. The dimensions of the hydrogel devices are in accordance
with the dimensions of the reamers employed by the surgeon. In
addition, the edges of the devices may be rounded to prevent edge
stress concentration.
[0077] A library of different size devices 1,2 may cover the wide
range of joint sizes so that every patient would have a nearly
perfect fit. At the time of surgery, the physician would choose and
implant the device of the appropriate dimensions. The thickness can
be adjusted, if necessary, to accommodate variations in joint
surface area and/or the patient's weight, as well as joint
conformity factors (i.e. the less conforming the joint, the higher
the thickness needs be).
[0078] The bone interface region 6 of the device is porous with a
pore size in the range of 10-1000 microns. The bone interface
region is coated with a layer of soluble or insoluble
hydroxyapatite that is chemically deposited by taking advantage of
the bonds created due to the negative charges of the hydrogel and
the calcium ions contained in the hydroxyapatite crystals as
demonstrated in FIG. 4. Two to twelve weeks after implantation, the
pores are filled with new bone tissue achieving an interdigitation
of the bone and the hydrogel device.
[0079] The surface of the bearing region 5 of the femoral head
component 1a has the same radius of curvature as the surface of the
bearing region 5 of the acetabulum component 4a to achieve a
dimensionally matched ball-in-socket mechanism and thus yield an
even distribution of the contact stresses. Furthermore, the bearing
region 6a of the acetabulum component may hold in its central
region a depression 100 so that a chamber 101 is formed between the
bearing sides of the acetabulum component 2a and the femoral
component 1a. The chamber 101 is filled with fluid 102 at times of
non bearing joint load, said fluid 102 gets pressurized once joint
loads are applied since the chamber 101 is effectively sealed by
the bearing region 5 surfaces; the pressurized fluid 102 can take
up significant portions of the joint load.
[0080] The femoral component 1a may have a variable shell thickness
profile as shown in FIG. 6B and in FIG. 7; the device thickness may
vary from 1 mm to 5 mm. As such, the thickest shell region is at
the superior side of the component 4, where the contact stresses
are higher, while it gradually tapers out towards the edges 5 to
increase range of motion of the joint and protect the device from
impingement. The femoral component 1a may also hold a recess 103 on
the superior side to accommodate any vessels that supply the
femoral head bone. The acetabulum component 2a may hold a
protrusion on its convex side that can fit inside the acetabular
fossa, after the later is surgically reamed to remove any soft
tissue; the said protrusion secures the initial placement of the
hydrogel device acetabulum component 4a so that in combination with
the continuous compression the joint is subjected to, implant
migration is prevented.
[0081] In another embodiment, the hydrogel device can be applied to
the knee joint. The device is comprised of a distal femur component
1b and a tibial plateau component 2b as shown in FIG. 8. The distal
femur component 1b resembles in overall shape that of natural
distal femur cartilage. It can be premade to have a generic
adaptable shape or a patient specific geometry through reverse
engineering methods. The component is placed on the bone like a
sock. After the knee joint is exposed and damaged cartilage layer
is surgically removed, the distal femur component 1b can be placed.
Special openings in the device allow ligament insertion; as such a
lateral opening 110 and a central opening 111 accommodate the
lateral ligament and the cruciate ligaments respectively. The
device can be tightly held in place by means of hydrogel
stimulation and subsequent shrinking, either because of a change in
the pH, a change in salt concentration or a change in the
temperature, as also discussed in FIG. 32. For example, the
component 1b can be equilibrated in a pH 9 environment
pre-surgically which leads to increased swelling as discussed later
in this application. Upon equilibrium with the body fluids and
subsequent lowering of the pH, the component 1b will shrink, and
thus conform to the particular geometry of the distal femur 3b.
Alternatively the hydrogel can be pre-surgically equilibrated with
a low (compared to body fluids) salt concentration solution, for
example 0.01 M-0.05 M pre-surgically; upon implantation and salt
equilibrium with the body's salt concentration, for example 0.15 M,
the component conforms to the particular geometry of the distal
femur 3b taking advantage of the material's sensitivity to salt
concentration. In this way, an initial fixation of the component 1b
is secured on the distal femur 3b.
[0082] The tibial plateau component 2b can have a curved disk shape
and can be either unilateral or bilateral, that is it can cover
both tibial plateau 4b facets, or simply either the lateral or the
medial facet depending on the extent of the cartilage damage. One
way the tibial plateau component 2b can be fixated in the bone is
by surgically creating a depression 113 on the facet surface as
shown in FIGS. 9A-9C. The depression 113 can be made by either
reaming or by locally crushing the subchondral bone 112, for
example with a punch. The depression 113 has such dimensions so
that the implant can be press fit in it; for example, a circular
depression 113 can have a diameter that is one or two millimeters
smaller than that of a circular component 2b.
[0083] The bone interfacing region 6 of both components is porous,
with bone morphogenic proteins tethered on the surface to promote
bone adhesion and/or ingrowth as discussed in FIG. 29.
Microfractured or reamed bone exhibits regenerative properties; the
interdigitation between bone and the hydrogel device takes up to
twelve weeks post surgically.
Material Specifications
[0084] Current materials used in arthroplasty function well as
mechanical "bearings" but suffer from key material property
differences compared to natural cartilage. Because plastics,
metals, and ceramics are not hydrated, they solely rely on
serum/synovial fluid lubrication; the bearing function relies on
the tolerances as well as on the surface roughness. Interfacial
wear ultimately produces wear debris by means of abrasion. The
products of wear are typically in particulate form (e.g.
polyethylene particles) or in the form of ions (e.g., metal ions).
Both of these have been shown to be promoters of inflammation in
synovial joints and have been found to migrate into internal
organs. Moreover, because metals are significantly stiffer than
bone, they alter the stress transfer to the bone leading to bone
resorption or fibrous tissue formation and ultimately loosening
around the implants. One way that researchers have been exploring
to avoid problems associated with conventional orthopaedic
"hardware" is to use "software" (soft materials). One such approach
available in the U.S. is "Carticel" autologous cartilage grafting.
This has been shown to be effective in "filling in" focal defects
in knee cartilage with regenerated cartilage from a patient's own
chondrocytes. There are a number of other approaches under
development that are related to tissue engineered cartilage, cell
transplantation, and autologous grafting. To date, the simultaneous
combination of cartilage-like stiffness and a hydrated, lubricious
surface has been an elusive pair of properties to attain in
materials engineering.
[0085] The present invention provides a hydrogel device 1 having an
interpenetrating polymer network (IPN) hydrogel network based on a
neutral cross-linked network of end-linked macromonomers 13 as the
first network 10 and an ionized crosslinked polymer in the second
network 11 depicted in FIG. 10. In one of the embodiments, the
first network 10 is composed of end-linked poly(ethylene glycol)
macromonomers with defined molecular weight. The second network 11
is, in contrast, a loosely crosslinked, ionizable network of
poly(acrylic acid) (PAA). Furthermore, the hydrogel is comprised of
an aqueous salt solution 12. This PEG/PAA IPN has high tensile
strength, high compressive strength, and a low coefficient of
friction when swollen in phosphate buffered saline at a pH of 7.4,
as detailed in Table 1.
[0086] Homopolymer networks of PEG and PAA are both relatively
fragile materials (the former is relatively brittle, the latter is
highly compliant). However, the two polymers can form complexes
through hydrogen bonds between the ether groups on PEG and the
carboxyl groups on PAA. This inter-polymer hydrogen bonding
enhances their mutual miscibility in aqueous solution, which, in
turn, yields optically clear, homogeneous polymer blends. By
loosely cross-linking (instead of densely cross-linking) the
ionizable network (PAA, pKa=4.7), large changes in its network
configuration can be induced by changing the pH of the solvent
without affecting the neutral PEG network. In salt-containing
buffers of pH greater than 4.7, the PAA network becomes charged and
swells; at a pH lower than 4.7, the PAA network is protonated and
contracts.
[0087] FIGS. 11A-11C shows the steps required for synthesis of an
IPN hydrogel according to the present invention. The starting
material for the hydrogel is a solution of telechelic macromonomers
13 with functional end groups 15 dissolved in water 16. The
telechelic macromonomers are polymerized (FIG. 11A) to form a
first, water-swollen polymer network 10. Next, (FIG. 11B)
hydrophilic, ionizable monomers 14 mixed with water 16 are added to
the first polymer network 10 along with a photoinitiator and a
crosslinking agent. The hydrophilic, ionizable monomers 14 are then
photopolymerized and cross-linked in the presence of first polymer
network 10 to form second polymer network 11 in the presence of the
first 10. This results in formation of a water-swollen IPN hydrogel
(FIG. 11B, right). The water-imbibed IPN is then immersed in a
salt-containing solution 12 at pH 7.4 (FIG. 11C), and is swollen to
equilibrium, yielding a simultaneous increase in both the water
content and stiffness modulus of the IPN. The IPN on the right in
FIG. 11C has a higher stiffness modulus compared to the IPN on the
left. This increase in modulus as a result of strain (induced in
this case by swelling) is believed to be caused by an increase in
the number of physical crosslinks within the IPN. For the purpose
of the present invention, "strain hardening" is defined as an
increase in physical crosslinks (entanglements) and an increase in
the stiffness modulus with applied swelling induced strain. The end
material is an internally osmotically pre-stressed IPN that
exhibits increased stiffness and strength.
[0088] FIG. 12A i-iv shows according to an embodiment of the
present invention method steps of how an IPN is prepared after
monomers 17 are used to make the first network 10. Exposure to UV
light in the presence of a photoinitiator and crosslinker (not
shown) leads to polymerization and crosslinking to form a network
10, depicted by the transition from (i) to (ii). In (iii) to (iv),
the first network is swollen with the second network precursor
monomers 14, a crosslinking agent (not shown) and a photoinitiator
(not shown). Exposure to UV light initiates polymerization and cros
slinking of the second network 11 in the presence of the first 10
to form the IPN. FIG. 12B shows according to an embodiment of the
present invention method steps of how an IPN is prepared after
macromonomers 13 with reactive endgroups 15 are used to form a
first network 10 in the presence of an existing second network 11
or linear macromolecules and/or biomacromolecules. A mixture of the
first and second polymeric components is made, and then the
telechelic macromonomers 13, 15 are reacted under UV light to form
the first network 10 in the presence of the second 11. If the
second network 11 is crosslinked chemically, then it is a fully
interpenetrating network. If it is not (and only physically
crosslinked), then it is a semi-interpenetrating network. FIG. 12C
shows according to an embodiment of the present invention method
steps of how an IPN is formed from a first network 10 based on
monomers 17 and a second network 11 or linear macromolecules and/or
biomacromolecules. A mixture of the monomers 17 and macromolecules
is made, and then the monomers are reacted under UV light to form
the first network in the presence of the second 11. If the second
network 11 is crosslinked chemically, then it is a fully
interpenetrating network. If it is not (and only physically
crosslinked), then it is a semi-interpenetrating network.
[0089] In one embodiment of the present invention, grafted polymers
are used to form the IPN. FIG. 16A shows a standard IPN according
to the present invention, with first polymer network 10 and second
polymer network 11. FIG. 16B shows an IPN in which first polymer
network 10 is grafted with a hydrophilic polymer 29. Any of the
aforementioned macromonomers, monomers, or combinations of
macromonomers and monomers may be used to get a grafted structure.
FIG. 16C shows an IPN in which the second polymer network 11 is
grafted with another hydrophilic macromonomer 30. FIG. 16D shows an
IPN in which first polymer network 10 is grafted with a hydrophilic
monomer 29 and the second polymer network 11 is grafted with
another hydrophilic macromonomer 30. The grafted networks are made
by polymerizing aqueous mixtures of the two components in ratios
that yield a network that is predominantly made from one polymer
but has grafted chains of the second polymer.
[0090] Any hydrophilic telechelic macromonomer 13 may be used to
form the first polymer network 10. In a preferred embodiment,
preformed polyethylene glycol (PEG) macromonomers are used as the
basis of the first network (10). PEG is biocompatible, soluble in
aqueous solution, and can be synthesized to give a wide range of
molecular weights and chemical structures. The hydroxyl end-groups
of the bifunctional glycol can be modified into crosslinkable
end-groups 15. End-group or side-group functionalities to these
macromolecules and biomacromolecules may include, but are not
limited to, acrylate (e.g. PEG-diacrylate), methacrylate, vinyl,
allyl, N-vinyl sulfones, methacrylamide (e.g.
PEG-dimethacrylamide), and acrylamide (e.g. PEG-diacrylamide). For
instance, PEG macromonomers can be chemically modified with
endgroups such as diacrylates, dimethacrylates, diallyl ethers,
divinyls, diacrylamides, and dimethacrylamides. Examples of the
end-group functionalization reactions to yield telechelic,
crosslinkable PEG macromonomers are shown in FIGS. 13, 14, 15.
These same endgroups can be added to other macromonomers, such as
polycarbonate, poly(N-vinyl pyrrolidone), polyurethane, poly(vinyl
alcohol), polysacchrarides (e.g. dextran), biomacromolecules (e.g.
collagen) and derivatives or combinations thereof. The first
network 10 can also be copolymerized with any number of other
polymers including but not limited to those based on acrylamide,
hydroxyethyl acrylamide, N-isopropylacrylamide, polyurethane,
2-hydroxyethyl methacrylate, polycarbonate, 2-hydroxyethyl acrylate
or derivatives thereof.
[0091] Preferably, the hydrophilic monomer 14 in the second network
11 is ionizable and anionic (capable of being negatively charged).
In a preferred embodiment, poly(acrylic acid) (PAA) hydrogel is
used as the second polymer network, formed from an aqueous solution
of acrylic acid monomers. Other ionizable monomers include ones
that contain negatively charged carboxylic acid or sulfonic acid
groups, such as methacrylic acid,
2-acrylamido-2-methylpropanesulfonic acid, hyaluronic acid, heparin
sulfate, chondroitin sulfate, and derivatives, or combinations
thereof. The second network monomer 14 may also be positively
charged or cationic. The hydrophilic monomer may also be non-ionic,
such as acrylamide, methacrylamide, N-hydroxyethyl acrylamide,
N-isopropylacrylamide, methylmethacrylate, N-vinyl pyrrolidone,
2-hydroxyethyl methacrylate, 2-hydroxyethyl acrylate or derivatives
thereof. These can be copolymerized with less hydrophilic species
such as methylmethacrylate or other more hydrophobic monomers or
macromonomers. Crosslinked linear polymer chains (i.e.
macromolecules) based on these monomers may also be used in the
second network 11, as well as biomacromolecules such as proteins
and polypeptides (e.g. collagen, hyaluronic acid, or chitosan).
[0092] Adding a photoinitiator to an aqueous solution of the
end-linkable macromonomers 13 in water and exposing the solution to
UV light results in the crosslinking of the PEG macromonomers,
giving rise to a PEG hydrogel that serves as the first network 10.
Polymerizing and crosslinking a second network 11 inside the first
network will give rise to the IPN structure. Preparing IPN
hydrogels through free-radical polymerization has the additional
advantage that it enables the use of molds to form hydrogels of
desired shape such as the ones depicted in FIGS. 7, 8. Preferably,
the first polymer network contains at least 50%, more preferably at
least 75%, most preferably at least 95% of the telechelic
macromonomer 13, 15 by dry weight. Other solutions including
buffers and organic solvents (or mixtures thereof) may also be used
to dissolve the first network macromonomers 13 or second network
monomers 14.
[0093] Any type compatible cross-linkers may be used to crosslink
the second network 11 in the presence of any of the aforementioned
first networks 10 such as, for example, ethylene glycol
dimethacrylate, ethylene glycol diacrylate, diethylene glycol
dimethacrylate (or diacrylate), triethylene glycol dimethacrylate
(or diacrylate), tetraethylene glycol dimethacrylate (or
diacrylate), polyethylene glycol dimethacrylate, or polyethylene
glycol diacrylate, methylene bisacrylamide,
N,N'-(1,2-dihydroxyethylene) bisacrylamide, derivatives, or
combinations thereof. Any number of photoinitiators can also be
used. These include, but are not limited to,
2-hydroxy-2-methyl-propiophenone and
2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone.
Examples of First Network Telechelic Macromonomers
[0094] Telechelic PEG macromonomers 13 with acrylate or
methacrylate endgroups can be synthesized in the following manner.
PEG was dried from Toluene, redissolved in THF (550 mL per 100 g)
and kept under Nitrogen. Distilled triethylamine (2.5 eq per OH
group) was added slowly to this solution. Acryloyl chloride (or
methacryloyl chloride) was then added via a dropping funnel
(diluted with THF) over 30 min at room temperature. The reaction
(FIG. 13) was allowed to proceed overnight. Filtration was carried
out to remove the formed salt. The volume of the solvent was
reduced using a Rotavap, and precipitation was carried out in
diethylether. As an alternative to extraction, filtration via a
cellulose membrane has also been performed. The raw product was
dried after precipitation from diethylether in a vacuum, then
dissolved in MeOH and dried in a Rotavap. It is then dissolved in
water and filtrated through a membrane, and was finally
freeze-dried.
[0095] Networks have also been formed from PEG-diacrylamide.
PEG-diol was converted to PEG-diacrylamide (FIG. 14) using the
following procedure. PEG mol wt 3400 (100 g, 58.8 mmol --OH) was
azeotropically distilled in 700 mL toluene under nitrogen and
removing about 300 mL of toluene. The toluene was then evaporated
completely and then the PEG re-dissolved in anhydrous
tetrahydrofuran. Triethylamine was distilled prior to use. The
solution was cooled in a room temperature bath under Nitrogen and
then cooled in an ice bath. Anhydrous dichloromethane was added
until the solution became clear (about 100 mL). Triethylamine (24.6
mL, 176.5 mmol) was then added dropwise with stirring, followed by
the dropwise addition of 13.65 mmol mesyl chloride (176.5 mmol, an
excess of 3 eq per OH endgroup). The reaction proceeded overnight
under argon. The solution was filtered through paper under vacuum
until clear, followed by precipitation in diethyl ether. The
product was then collected by filtration and dried under vacuum.
The PEG-dimesylate product was added to 400 mL 25% aqueous ammonia
solution in a 1 L bottle. The lid was tightly closed and sealed
with Parafilm, and the reaction was vigorously stirred for 4 days
at room temperature. The lid was then removed and the ammonia
allowed to evaporate for 3 days. The pH of the solution was raised
to 13 with 1 N NaOH, and the solution was extracted with 100 mL
dichloromethane. For the extraction with dichloromethane, NaCl was
added to the water-phase (.about.5 g) and the water-phase was
extracted several times with 150 mL of dichloromethane. The
dichloromethane washes were combined and concentrated in vacuo. The
product was precipitated in diethyl ether, and dried under vacuum.
PEG-diamine mol wt 3400 (20 g, 11.76 mmol amine) was then
azeotropically distilled in 400 mL of toluene under Nitrogen,
removing about 100 mL of toluene. The toluene was then evaporated
completely and then the PEG re-dissolved in anhydrous
tetrahydrofuran. The solution was cooled in a room temperature bath
under Nitrogen and then cooled in an ice bath. Triethylamine (2.46
mL, 17.64 mmol) was added dropwise with stirring, followed by the
dropwise addition of 1.43 mL of acryloyl chloride (17.64 mmol). The
reaction (FIG. 14) proceeded overnight in the dark under Nitrogen.
The solution was then filtered through paper under vacuum until
clear, followed by precipitation in diethyl ether. The product was
collected by filtration and dried under vacuum. The product was
then dissolved in 200 mL of deionized water, with 10 g of sodium
chloride. The pH was adjusted to pH 6 with NaOH and extracted 3
times with 100 mL of dichloromethane (with some product remaining
in the water phase as an emulsion). The dichloromethane washes were
combined and the product was precipitated in diethyl ether, and
dried under vacuum. Alternatively, PEG-diacrylamide has been
precipitated from Diethylether once, redissolved in MeOH, dried
from MeOH and then purified by centrifugal filtration in water
through a cellulose membrane (MWCO:3000). Freeze drying was used to
attain the desired product.
[0096] PEG macromonomers containing diols have also been converted
into allyl ethers. Difunctional allyl ether macromonomers were
synthesized from PEG using the following procedure (FIG. 15). Fresh
anhydrous tetrahydrofuran (THF) (100 mL) was added to every 10 g of
PEG. This mixture was gently heated until the PEG dissolved and
then cooled in an ice bath before sodium hydride was slowly added
in multiple portions (1.05 molar equiv. NaH for the PEG ReOH
groups). After the release of H2 gas ceased, the system was purged
with argon and allyl chloride or allyl bromide (1.1 molar equiv.
per PEG OH-group, diluted 1:10 in THF) was added dropwise using an
addition funnel, after which the reaction mixture (FIG. 15) was
transferred to an 85 degrees Celsius oil bath and refluxed
overnight. Vacuum filtration was used to remove the sodium bromide
side products and rotary evaporation was used to reduce the
concentration of THF before the PEG-allyl ether products were
precipitated from solution using iced diethyl ether (10:1 v:v
diethyl ether:THF solution).
EXAMPLES
[0097] The following description refers to an exemplary embodiment
of a strain-hardened interpenetrating polymer network hydrogel with
PEG as a first network 10 polymer and PAA as a second network 11
polymer. The IPN hydrogel is synthesized by a (two-step) sequential
network formation technique based on UV initiated free radical
polymerization. A precursor solution for the first network is made
of purified, telechelic PEG dissolved at a typical concentration of
50% w/v in phosphate buffered saline (PBS) solution, water, or an
organic solvent with either 2-hydroxy-2-methyl-propiophenone or
2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone as
the UV sensitive free radical initiator. The types of telechelic
PEG macromonomers used were PEG-diacrylate, PEG-dimethacrylate,
PEG-diacrylamide, and PEG-diallyl ether. In other embodiments,
either network can be synthesized by free radical polymerization
initiated by other means, such as thermal-initiation and other
chemistries not involving the use of ultraviolet light. In the case
of UV polymerization, the precursor solution is cast in a
transparent mold and reacts under a UV light source at room
temperature. Upon exposure, the precursor solution undergoes a
free-radical induced gelation and becomes insoluble in water. The
mold is fabricated in such a way that yields hydrogels at
equilibrium swelling desired dimensions.
[0098] To incorporate the second network 11, the PEG-based hydrogel
is immersed in the second monomer 14 solution, such as an aqueous
solution of (10-100% v/v) acrylic acid containing a photo-initiator
and a cross-linker, from 0.1% to 10% by volume triethylene glycol
dimethacrylate (TEGDMA), triethylene glycol divinyl ether,
N,N-methylene bisacrylamide, or
N,N'-(1,2-dihydroxyethylene)bisacrylamide, for 24 hours at room
temperature. The swollen gel is then exposed to the UV source and
the second network 11 is polymerized and crosslinked inside the
first network 10 to form an IPN structure in which the degree of
crosslinking in the second network is less than that of the first
network. Preferably, the molar ratio of the first network
telechelic macromonomer to the second network monomer ranges from
about 1:1 to about 1:5000. Also preferably, the weight ratio of the
first network to the second network is in the range of about 10:1
to about 1:10. In another embodiment of the present invention, the
IPNs have a molar ratio of the second monomer ingredient to the
first macromonomer ingredient higher than 100:1.
[0099] Key characteristics of hydrogels such as optical clarity,
water content, flexibility, and mechanical strength can be
controlled by changing various factors such as the second monomer
type, monomer concentration, molecular weight and UV exposure time.
The experimental focus of the ensuing section is on the swelling
induced strain hardening observed in this system by testing how it
manifests through uniaxial tensile tests under various conditions
of first 10 and second 11 network crosslinking and swelling.
Swelling data were used to calculate the equilibrium water and
polymer content of the networks, which were correlated with
stiffness modulus, true stress-at-break, and true strain-at-break.
The results indicate that strain hardening is derived from physical
entanglements between the PEG and PAA networks that are intensified
by bulk deformation. Under conditions that promote hydrogen bonding
(when the pH is at or below 4.7, the pKa of PAA), these
entanglements are reinforced by interpolymer complexes between PEG
and PAA, leading to an increase in the fracture strength of the
IPN. Under conditions that promote ionization of PAA (when the pH
is above 4.7 and salt is added), increased steric interactions
(i.e. physical crosslinks) between the swelling PAA network and
static, telechelic PEG macromonomer network lead to an increase in
the stiffness modulus.
[0100] In particular embodiment, an array of IPNs with varying
molecular weights of PEG in the first network 10 and varying PAA
polymer content in the second network 11 were fabricated based on
diacrylate crosslinking in the first network 10 and triethylene
glycol dimethacrylate crosslinking in the second network 11. All
hydrogels were formed by photopolymerization with UV light using
the photoinitiator, 2-hydroxy-2-methyl-propiophenone at a
concentration of 1% v/v with respect to the monomer 14 or
macromonomer 15. Before the IPNs were prepared, single network
hydrogels based on PEG and PAA were synthesized separately to
confirm the formation of gels of each composition and to
investigate the physical properties of the single networks. For the
PEG single network, a range of hydrogels that varied between 275
and 14000 for the MW of the PEG macromonomer was synthesized. It
was found that low MW PEG macromonomers gave rise to gels that were
brittle, whereas the hydrogels made from higher molecular weight
PEG-DA (3400) were transparent and flexible when swollen in
deionized water. Based on these results, a range of different MWs
of PEG (3400, 4600, 8000, and 14000) were chosen as macromonomers
for the first hydrogel network. A series of IPNs was synthesized by
polymerizing and crosslinking a PAA network within each type of PEG
network. The resultant IPNs had significantly better mechanical
properties compared with single network hydrogels.
[0101] To explore the effect of the molecular weight of the
telechelic PEG-DA macromonomer on IPN mechanical strength, PEG
chains with MWs 3400 Da, 4600 Da, 8000 Da, and 14000 Da were used
in the first network while keeping the acrylic acid polymerization
conditions constant (50% v/v in deionized water with 1% v/v
crosslinker and 1% v/v photoinitiator with respect to the monomer).
The resulting IPNs were characterized in terms of their water
content, tensile properties, and mesh size in deionized water.
Changing the MW of the PEG-DA macromonomer led to a change in the
moduli of the PEG-DA single networks, as shown in Table 2. This
effect was magnified in the PEG/PAA IPNs, where the IPNs initial
and final moduli get increasingly higher as the networks are
prepared from lower molecular weight PEG-DA macromonomers. Of note,
there was little increase in strength when the PEG MW is increased
above 8000, indicating that a contrast between the molecular weight
between cros slinks of the PEG and PAA networks is important for
strength enhancement. Moreover, the molecular weight of the PEG
macromonomer was strongly correlated to the critical strain
(.epsilon..sub.crit) at which the stress-strain curve makes the
transition from the initial modulus to the strain-hardened final
modulus. The .epsilon..sub.crit was smaller for the IPNs prepared
from lower MW PEG macromonomers, meaning that these networks
strain-harden more rapidly in response to deformation.
TABLE-US-00002 TABLE 2 Physical properties of PEG/PAA IPNs under
different PEG crosslinking and swelling conditions specimen
swelling solution WC** (%) q*** .sigma..sub.max (MPa)
.epsilon..sub.break E.sub.o (MPa) E.sub.f (MPa) PEG(3.4k) dH.sub.2O
79.3 .+-. 2.1 4.6 0.33 .+-. 0.09 0.23 .+-. 0.089 1.49 .+-. 0.05 --
PEG(3.4k)/PAA dH.sub.2O 56.3 .+-. 3.3 2.3 8.94 .+-. 0.97 0.62 .+-.
0.03 2.32 .+-. 0.09 36.2 .+-. 2.9 PEG(3.4k)/PAA pH 7.4, I = 0.15
68.7 .+-. 1.6 3.2 8.94 .+-. 1.08 0.50 .+-. 0.11 3.58 .+-. 0.001
PEG(4.6k) dH.sub.2O 84.5 .+-. 0.4 6.5 0.65 .+-. 0.14 0.67 .+-. 0.13
0.85 .+-. 0.002 -- PEG(4.6k)/PAA dH.sub.2O 57.0 .+-. 0.6 2.3 5.98
.+-. 2.31 0.77 .+-. 0.11 1.15 .+-. 0.20 20.5 .+-. 5.0 PEG(4.6k)/PAA
pH 7.4, I = 0.15 77.0 .+-. 1.2 3.0 6.28 .+-. 1.98 0.62 .+-. 0.07
3.50 .+-. 0.28 15.1 .+-. 2.0 PEG(8.0k) dH.sub.2O 90.5 .+-. 1.2 10.5
0.27 .+-. 0.04 0.63 .+-. 0.04 0.20 .+-. 0.05 -- PEG(8.0k)/PAA
dH.sub.2O 80.2 .+-. 1.5 5.1 4.83 .+-. 1.09 1.18 .+-. 0.09 0.38 .+-.
0.04 11.4 .+-. 0.79 PEG(8.0k)/PAA pH 7.4, I = 0.15 90.9 .+-. 0.1
11.0 1.98 .+-. 0.24 0.75 .+-. 0.05 0.53 .+-. 0.12 6.1 .+-. 0.01
PEG(8.0k)/PAA pH 7.4, I = 0.30 89.5 .+-. 0.4 9.5 1.74 .+-. 0.20
0.73 .+-. 0.05 0.49 .+-. 0.07 5.25 .+-. 0.01 PEG(8.0k)/PAA pH 7.4,
I = 0.75 83.1 .+-. 0.6 5.9 2.15 .+-. 0.40 0.80 .+-. 0.07 0.47 .+-.
0.03 6.6 .+-. 0.01 PEG(8.0k)/PAA pH 7.4, I = 1.5 77.7 .+-. 0.2 4.5
3.16 .+-. 0.97 0.84 .+-. 0.09 0.53 .+-. 0.11 8.98 .+-. 0.01
PEG(8.0k)/PAA pH 3, I = 0.05 76.5 .+-. 2.1 4.3 8.18 .+-. 1.76 1.20
.+-. 0.01 0.52 .+-. 0.03 24.0 .+-. 3.6 PEG(8.0k)/PAA pH 4, I = 0.05
86.4 .+-. 1.5 7.4 5.48 .+-. 1.44 1.01 .+-. 0.12 0.56 .+-. 0.04 15.1
.+-. 1.8 PEG(8.0k)/PAA pH 5, I = 0.05 94.5 .+-. 1.1 18.2 1.26 .+-.
0.05 0.63 .+-. 0.02 0.62 .+-. 0.08 3.99 .+-. 0.29 PEG(8.0k)/PAA pH
6, I = 0.05 95.6 .+-. 1.0 22.7 0.86 .+-. 0.15 0.53 .+-. 0.02 0.68
.+-. 0.005 3.10 .+-. 0.30 PEG(14.0k) dH.sub.2O 95.1 .+-. 1.2 20.4
0.07 .+-. 0.007 0.70 .+-. 0.02 0.062 .+-. 0.005 -- PEG(14.0k)/PAA
dH.sub.2O 84.3 .+-. 1.7 6.4 0.25 .+-. 0.05 0.82 .+-. 0.07 0.18 .+-.
0.01 0.57 .+-. 0.17 PAA dH.sub.2O 90 .+-. 1.7 10.0 0.14 .+-. 0.03
0.89 .+-. 0.09 0.14 .+-. 0.03 -- PAA pH 7.4, I* = 0.15 95.5 .+-.
1.7 22.2 0.07 .+-. 0.01 0.65 .+-. 0.10 0.050 .+-. 0.001 -- PAA pH
3, I = 0.05 80.4 .+-. 1.0 5.1 0.38 .+-. 0.08 1.23 .+-. 0.05 0.09
.+-. 0.01 -- PAA pH 4, I = 0.05 90.0 .+-. 0.7 10.0 0.35 .+-. 0.11
1.19 .+-. 0.15 0.090 .+-. 0.001 -- PAA pH 5, I = 0.05 96.2 .+-. 0.2
26.3 0.04 .+-. 0.007 0.50 .+-. 0.11 0.05 .+-. 0.008 -- PAA pH 6, I
= 0.05 96.6 .+-. 0.1 30.3 0.05 .+-. 0.01 0.66 .+-. 0.08 0.050 .+-.
0.002 -- *I = ionic strength **water content = (swollen weight -
dry weight)/(swollen weight) ***average swelling ratio = (swollen
weight)/(dry weight)
[0102] The significance of forming an interpenetrating structure
rather than a copolymeric structure was explored by synthesizing a
PEG-co-PAA copolymer hydrogel and testing its tensile properties.
Its stress-strain profile was then juxtaposed with those of the IPN
and the PEG and PAA single networks. In FIG. 18A, a representative
true stress (.sigma..sub.true) versus true strain
(.epsilon..sub.true) profile of the PEG(8.0 k)/PAA IPN is compared
to those of the PEG(8.0 k)-PAA copolymer and their component
PEG(8.0 k) and PAA networks. The IPN exhibits strain-hardening
behavior with a stress-at-break that is greater than four times
that of the copolymer and single network. However, since each of
the materials tested has different water content, the stress data
were normalized on the basis of polymer content to determine the
true stress per unit polymer in each hydrogel. In FIG. 18B, the
true stress per unit polymer (.sigma..sub.true per unit polymer) is
plotted against true strain for PEG(8.0 k)-DA, PAA, PEG(8.0 k)/PAA,
and the PEG(8.0 k)-PAA copolymer. The initial moduli of the PEG
single network, the copolymer, and IPN are identical (E.sub.o per
unit polymer=0.91 MPa), while that of the PAA single network is
lower (E.sub.o per unit polymer=0.55 MPa). Near the break point of
the PEG network, .epsilon..sub.true.about.0.6, the copolymer
continues to be elongated with a modulus that is intermediate
between the PEG and PAA single networks, of which it is equally
composed by weight. Ultimately, it fails at a strain that is also
intermediate between the .epsilon..sub.break values of the two
single networks. In stark contrast, just beyond the failure point
of the PEG network, the PEG/PAA IPN manifests a dramatic strain
hardening effect in which its modulus increases by 30 fold, and
breaks at .epsilon..sub.true.about.1.0 under a mean maximum stress
per unit solid of 10.6 MPa. Without normalization for polymer
content, .sigma..sub.break for the IPN (20% solid) and copolymer
(51% solid) are 3.5 MPa and 0.75 MPa, respectively.
[0103] To explore the role of interpolymer hydrogen bonding, the pH
of the hydrogel swelling liquid was varied to change the ionization
state of the PAA network. Since the equilibrium swelling of PAA is
sensitive to variations in pH, a change in the pH was expected to
have an impact on the mechanical properties of PEG/PAA IPNs. After
synthesis, the water-swollen PAA single networks and PEG(8.0 k)/PAA
IPNs were placed in buffers of pH 3-6 and constant ionic strength
(I) of 0.05. In both the PAA network and the IPN, the equilibrium
water content increased as the pH was increased from 3 to 6 (Table
2). In the case of the PAA networks, those at pH 3 and 4 were
moderately swollen, while those at pH 5 or 6 were highly swollen
due to ionization of PAA above its pKa (4.7). The IPNs also
achieved different levels of swelling depending on the pH; those at
pH 3 and 4 were moderately swollen, while those at pH 5 or 6 were
highly swollen due to ionization of PAA above its pKa (4.7). Of
note, at both pH 3 and 4, the IPN achieved a lower equilibrium
water content than PAA alone. This can be explained, in part, by
the fact that PEG and PAA complex with each other via hydrogen
bonds in an acidic environment, leading to a more compact, less
hydrated interpenetrating network structure. At pH above 4.7, the
PEG and PAA chains dissociate as the PAA becomes ionized and
counterions (along with water) enter the hydrogel to maintain
charge neutrality, leading to a high degree of swelling.
Nevertheless, the IPNs swell to a slightly lower extent (1.0-1.5%)
than the PAA single networks due to the constraint that the PEG
network places on PAA swelling. Table 2 also shows that the maximum
stress (.sigma..sub.max), or tensile strength, of the PEG/PAA IPN
is nearly an order of magnitude greater in its less-swollen state
at pH 3 (.sigma..sub.max=8.2 MPa) than in its more swollen state at
pH 6 (.sigma..sub.max=0.86 MPa). A similar phenomenon is observed
in the PAA network, but the absolute values for .sigma..sub.max are
0.38 MPa at pH 3 and 0.05 MPa at pH 6. At every pH, then, the IPN
has greater tensile strength than the PAA network, and this
difference is intensified at lower pH. In contrast to the
differences in the stress-at-break, the trends in the
strain-at-break values of the IPN and PAA networks are roughly
equivalent, changing from .epsilon..sub.break values of .about.1.2
at pH 3 to .about.0.55 at pH 6. This result confirms the
observation made in FIGS. 18A-18B, in which the extensibility of
the IPN seems to be due to the presence of the PAA network, which
has a higher .epsilon..sub.break (0.9) than PEG (0.6). The mere
presence of the PAA network in the IPN appears to enhance the
uniaxial extensibility of the network, a property that enables the
IPN hydrogel to be used to support joint loads. In the context of
the maximum stress data (Table 2), however, the load-bearing
capacity at higher extensions is greater in the presence of
hydrogen bonding at low pH than it is in the absence of hydrogen
bonding at high pH. In contrast, pH dependence of the initial
stiffness moduli (E.sub.o) of the IPN and PAA networks is less
straightforward. The modulus of the PAA network exhibits a small
drop from 0.09 MPa to 0.05 MPa as the pH is increased from 3 to 6.
On the other hand, the modulus of the IPN does not decrease at all,
but rather increase when the pH is changed from 3 to 6. Of note,
the pH-dependence of the IPN does not follow the trend exhibited by
the PAA single network, in which the modulus drops by approximately
one-half when transitioning from pH 4 to pH 5. This decrease in
modulus is correlated with an increase in water content of the PAA
single network. In addition, the dependency of water content and
subsequently of the hydrogel volume or surface on the pH, enables a
(pH) stimulus sensitive hydrogel arthroplasty device that takes
advantage of the shrinking or swelling to adapt and secure fixation
inside or around a bone as described in a previous section.
[0104] To investigate the consequence of relative network moduli
even further, the swelling of PAA within the IPNs was maximized.
The experimental data shown in Table 2 indicated that the modulus
of the IPN was not negatively affected by the increased swelling.
The PEG network acts as a constraint on the swelling of PAA in a
way that leads to additional interpolymer interactions and a
corresponding increase in the IPN modulus. In particular, the
increase in the constraining effect of the neutral PEG network on
PAA swelling would increase the intensity and number of physical
entanglements in the IPN and, in turn, lead to the strain hardening
behavior observed in the IPN. To test this hypothesis, the IPNs
with first network MW PEG 3400, 4600, and 8000 and constant PAA
network conditions were placed in phosphate buffered saline (PBS,
pH 7.4, I=0.15) in order to induce maximal swelling under
physiologic conditions. Table 2 also shows the equilibrium water
content and corresponding swelling ratios for networks prepared
from PEG macromonomers with each of these molecular weights,
juxtaposed with the water content of the water-swollen and
PBS-swollen IPNs. Increasing the size of the first PEG network from
3400 Da to 4600 Da and 8000 Da increases the degree to which the
IPN is able to swell. Specifically, while the PEG(3.4 k)/PAA IPN
swells to only 70% water when ionized, the PEG(4.6 k)/PAA IPN
swells to 77% water and the PEG(8.0 k)/PAA IPN swells to 90% water
(nearly the same water content as the PEG(8.0 k) single network)
when ionized. Of note, the equilibrium water content values of the
PEG(3.4 k) and PEG(4.6 k)-based IPNs do not approach those of their
component PEG-DA networks (79.3% and 84.5%, respectively).
[0105] The time-dependent water content of the hydrogels was
evaluated in terms of the swollen-weight-to-dry-weight ratio. The
dry hydrogel was weighed and then immersed in water as well as
phosphate buffered saline. At regular intervals, the swollen gels
were lifted, patted dry, and weighed until equilibrium was
attained. The percentage of equilibrium water content (WC) was
calculated from the swollen and dry weights of the hydrogel:
WC = W s - W d W s .times. 100 ##EQU00001##
[0106] where W.sub.s and W.sub.d are the weights of swollen and dry
hydrogel, respectively.
[0107] FIG. 20 shows the time-dependent swelling behavior of an IPN
hydrogel composed of PEG and two different amounts of acrylic acid
in the second network (25% and 50%). The single network IPN gels
were dried in a desiccator, placed in deionized water, and then
weighed at regular time intervals. In both hydrogels, the majority
of swelling took place within 5-10 minutes and equilibrium swelling
was achieved within 30-40 minutes. The parameters varied to obtain
hydrogels with differing water content were the molecular weight of
the PEG macronomonomer, the weight fraction of PAA in the second
network, as well as the amount of crosslinking agent (e.g.
triethylene glycol dimethacrylate, or low molecular weight PEG-DA)
added to the first or second networks.
[0108] Table 3 shows the effect of varying the concentration of
acrylic acid monomer used to prepare the second network on the
equilibrium water content of PEG/PAA IPNs in PBS. In general,
higher concentrations of acrylic acid monomer leads to hydrogels
with lower equilibrium water content and higher stiffness (tensile
modulus) and tensile strength for a given set of cros slinking
conditions. IPN hydrogels according to the present invention made
from these constituents, preferably have an equilibrium water
content of between about 15%-95% and more preferably between about
50%-90%.
TABLE-US-00003 TABLE 3 Physical properties of PEG(3.4k)/PAA IPNs
with varying AA content in PBS WC (%) Tensile Modulus Tensile
Strength PEG(3.4k)/PAA[0.5] 69% 3.6 MPa 4.0 MPa PEG(3.4k)/PAA[0.7]
65% 12.0 MPa 12.0 MPa PEG(3.4k)/PAA[0.8] 62% 19.6 MPa 13 MPa
[0109] Because different MWs of PEG and different starting
concentrations of acrylic acid result in different amounts of
equilibrium water content, the final amount of PEG and PAA in the
hydrogel varies depending on the MW of the starting PEG used and
the concentration of acrylic acid used. Examples of compositions of
varying weight ratios of PEG and PAA that have been made according
to the present invention are shown in Table 4. The compositions in
this table were all made using a starting concentration of 50% PEG
macromonomers of molecular weight 8000 Da swollen in pure
water.
TABLE-US-00004 TABLE 4 Compositions of PEG(8.0k)/PAA IPNs with
varying preparation concentration of AA monomer Concentration of AA
in Dry Wt. % Dry Wt. % (Dry Wt. PEG)/ the preparation state PEG in
IPN PAA in IPN (Dry Wt. PAA) 30% 23.5% 76.5% 0.30 40% 17.5% 82.5%
0.20 50% 13.0% 87.0% 0.15
[0110] Swelling of the PAA network within the confines of a more
densely crosslinked PEG network (by lowering the MW of the PEG
macromonomer) has dramatic consequences on the resulting IPN
modulus. Specifically, FIG. 21 shows that the accelerated strain
hardening due to elevated pH, as demonstrated in FIG. 18B, is
accentuated even further when a PEG network with lower MW (4600
rather than 8000) is used to constrain PAA. These more tightly
crosslinked IPNs were placed in phosphate buffered saline to
examine them under physiologic conditions (pH 7.4, ionic
strength=0.15) where the PAA network is greater than 99% ionized.
The PEG(4.6 k)/PAA IPN was first swollen to equilibrium in pure
deionized water (pH 5.5, salt-free); it was then switched to the
ionizing conditions of phosphate buffered saline (pH 7.4, I=0.15)
and again swollen to equilibrium. The increase in the pH to 7.4 and
the addition of salt caused the PAA network (but not the PEG
network) to swell. The result of this differential swelling within
the IPN was a dramatic upward shift in the stress-strain profile
that included the initial portion of the curve. In other words,
there was an increase in not only the rate of strain hardening, but
also in the initial modulus. The strain-hardened PEG/PAA hydrogel
therefore demonstrates a compatible set of material properties
(stiffness, strength) in physiologic pH, rendering it an
appropriate selection for the arthroplasty device.
[0111] FIG. 22 shows according to an embodiment of the present
invention the stress-strain profiles of PEG(4.6 k)/PAA IPNs
prepared with three different combinations of crosslinker chemical
end-groups but the same formulations of PEG (MW 4.6 k, 50% by
weight in water) and AA (50% v/v in water) as well as the same
polymerization conditions (photoinitiator and crosslinker
concentration by mole and UV intensity) and swelling conditions
(PBS at pH 7.4). Specimen (A) was prepared from PEG-diacrylamide
first network and a PAA second network crosslinked with
N,N'-(1,2-dihydroxyethylene) bisacrylamide. Specimen (B) was
prepared from PEG-diacrylamide first network and a PAA second
network crosslinked with triethylene glycol dimethacrylate.
Specimen (C) was prepared from PEG-diacrylate first network and a
PAA second network crosslinked with triethylene glycol
dimethacrylate. These results demonstrated that alternate
crosslinking strategies can be employed to create the
strain-hardened IPNs based on telechelic macromonomer-based first
networks and ionized second networks without deviating from the
essence of the present invention.
[0112] PEG/PAA IPNs were swollen to equilibrium in a series of PBS
solutions of varying ionic strength (0.15 M, 0.30 M, 0.75 M, and
1.5 M) and their equilibrium water content and stress-strain
properties were measured. Table 2 shows that the water content of
the IPN is reduced with higher salt concentration in the swelling
medium, from over 90% at I=0.15 to less then 78% at I=1.5. This is
caused by the fact that increased salt in the buffer screens the
negative charges on the PAA chains, reducing electrostatic
repulsion and, in turn, swelling of the networks.
[0113] Ionic strength had a modest effect on the stress-strain
properties. Table 2 shows that the stress-strain properties of IPNs
in I=0.15 to I=0.75 were roughly equivalent. The IPN swollen in
buffer with I=1.5 showed a slight enhancement in fracture stress at
higher strains. This result is consistent with the water content
data, since the hydrogels with higher solids content (the IPN at
higher ionic strength conditions) should have greater mechanical
strength. Of note, the final modulus of the IPN in the solution
with the highest ionic strength (I=1.5) appeared to be higher than
those at lower ionic strength. However, the difference was small
and was not found to be statistically significant.
[0114] To increase the quantity of topological interactions between
the PAA and PEG networks, the polymer content of PAA was varied
inside of a PEG(3.4 k) first network. The volume fraction of
acrylic acid in solution at the time of the second network
polymerization was varied between 0.5 and 0.8 prior to
polymerization. After polymerization, the IPNs were swollen to
equilibrium in PBS. The resultant hydrogels had different water
content, from 62% in the PEG(3.4 k)/PAA[0.8] IPN to 65% in the
PEG(3.4 k)/PAA[0.7] IPN and 77% in the PEG(3.4 k)/PAA[0.5] IPN. Of
note, the IPNs with increased acrylic acid concentration had lower
water content, which in light of the super-absorbency of PAA is a
counterintuitive result. The water content and tensile properties
of these IPNs are shown in Table 3. The IPN with the highest PAA
content had the highest stress-at-break and modulus, while the one
with the lowest PAA content had the lowest stress-at-break and
strain-at-break. Notably, the initial modulus values for these
samples varied significantly, from 3.6 MPa in the PEG(3.4
k)/PAA[0.5] to 12 MPa in the PEG(3.4 k)/PAA[0.7] IPN and 19.6 MPa
in PEG(3.4 k)/PAA[0.8] IPN.
Effect of PAA Content on IPN Swelling in Pure Water
[0115] PEG(4600) single networks were prepared and imbibed with
varying concentrations of AA in the second network in the presence
of the photoinitiator and crosslinker. IPNs based on these
AA-swollen PEG networks were then formed by UV-initiated
polymerization. The IPNs were then removed from their molds,
immersed in deionized water, and allowed to reach equilibrium. The
volume of the IPNs relative to the PEG single networks were then
measured and compared. The results are plotted in FIGS. 19A-19B.
FIG. 19A shows that the volume of the IPN is increased with
increased amount of AA monomer in the second network. This is
consistent with the understanding that PAA absorbs water, and
therefore increased PAA content in the IPN should lead to increased
water absorption. Of note, however, is the fact that the IPN
deswells relative to the PEG single network when the AA:EG monomer
ratio is less than unity, and swells relative to the PEG network
when AA is in excess to EG monomers.
[0116] The same PEG/PAA IPNs of varying AA monomer content were
tested by uniaxial tensile measurements. The results are shown in
FIG. 19B. In this figure, both the fracture stress and Young's
modulus are plotted as functions of AA mass fraction at the time of
polymerization. Young's modulus exhibited a modest monotonic
increase as the AA concentration increased. In contrast, the
fracture stress exhibited a dramatic increase in magnitude when the
AA:EG ratio was increased beyond unity. As the AA monomer
concentration increased, however, the fracture stress exhibited a
monotonic decline. Finally, the photoinitiator (2-hydroxy-2-methyl
propiophenone) and crosslinker (triethylene glycol dimethacrylate)
concentrations of the PAA second network were varied during
polymerization within PEG(4.6 k) networks and the resulting PEG(4.6
k)/PAA IPNs were studied in terms of their mechanical properties in
both pure water and in PBS. The results are shown in Table 5.
TABLE-US-00005 TABLE 5 * Effect of crosslinker and photoinitiator
concentrations on the mechanical properties of PEG(4.6k)/PAA IPNs
Crosslinker Photoinitiator Sample Swelling Medium (vol. %) (vol. %)
E.sub.0 (MPa) .sigma..sub.max (MPa) .epsilon..sub.max 1 dH.sub.2O
0.1 1.0 1.0 .+-. 0.1 3.9 .+-. 1.2 0.63 .+-. 0.07 2 dH.sub.2O 1.0
1.0 1.4 .+-. 0.3 9.7 .+-. 0.4 0.91 .+-. 0.53 3 dH.sub.2O 10.0 1.0
0.8 .+-. 0.0 5.6 .+-. 3.7 1.07 .+-. 0.41 4 PBS, pH 7.4, I = 0.15
0.1 1.0 5.3 .+-. 0.3 0.5 .+-. 0.2 0.12 .+-. 0.03 5 PBS, pH 7.4, I =
0.15 1.0 1.0 8.4 .+-. 0.5 4.3 .+-. 0.8 0.44 .+-. 0.03 6 PBS, pH
7.4, I = 0.15 10.0 1.0 6.9 .+-. 0.7 1.1 .+-. 0.2 0.20 .+-. 0.03 7
dH.sub.2O 1.0 0.1 0.9 .+-. 0.2 5.2 .+-. 2.4 1.11 .+-. 0.08 8
dH.sub.2O 1.0 1.0 1.4 .+-. 0.3 9.7 .+-. 0.4 0.91 .+-. 0.53 9
dH.sub.2O 1.0 10.0 0.9 .+-. 0.0 4.2 .+-. 0.0 0.67 .+-. 0.00 10 PBS,
pH 7.4, I = 0.15 1.0 0.1 8.8 .+-. 0.0 3.3 .+-. 1.1 0.35 .+-. 0.10
11 PBS, pH 7.4, I = 0.15 1.0 1.0 8.4 .+-. 0.5 4.3 .+-. 0.8 0.44
.+-. 0.03 12 PBS, pH 7.4, I = 0.15 1.0 10.0 7.8 .+-. 0.2 1.9 .+-.
0.6 0.34 .+-. 0.06 * Samples 2 & 8 and 5 & 11 provided
repeated data to aid visual comparison between experimental
conditions
[0117] To demonstrate that an ionizable monomer is important in the
second network, a series of IPNs were prepared under conditions
that disrupted the degree of ionizability in the second network.
The first method used was copolymerization of the second network
with non-ionic monomers. AA monomers in the second network were
mixed in three different concentrations relative to the HEA
monomers: 10:1, 3:1, and 1:1. Uniaxial tensile testing experiments
of the hydrogels swollen in deionized water showed that the
PEG/P(AA-co-HEA) IPNs with the highest ratio of AA:HEA in the
second network exhibited enhanced mechanical strength.
Specifically, changing tensile strength of the IPNs decreased from
9 MPa to 6 MPa and then to 3.5 MPa when the AA:HEA ratio decreased
from 10:1 to 3:1 to 1:1, respectively. In other words, IPNs with
higher relative HEA content exhibited almost no enhancement in
mechanical properties. This result demonstrates that the presence
of ionizable carboxyl acid groups in PAA is an important element in
the present invention.
[0118] In another set of experiments, PEG networks were immersed in
AA solutions (containing photoinitiator and crosslinker) that were
partially neutralized to pH 5.5 by titration with sodium hydroxide.
The monomer-swollen PEG networks were then exposed to UV light to
form a partially neutralized PAA network within the PEG network.
These "pre-neutralized" PEG/PAA IPNs were then washed in PBS and
subjected to uniaxial tensile tests. It was found that neutralizing
the AA solution prior to polymerization and then forming the second
network leads to an IPN with the same elastic modulus, but with
dramatically reduced fracture strength. The stress-at-break is
reduced from nearly 4 MPa--in the case of the IPNs prepared under
acidic conditions and then neutralized in PBS buffer--to roughly
0.5 MPa. This demonstrates the importance of the fabrication
process in creating these strain-hardened IPNs; that is, in the
preferred embodiment, ionization and swelling of the second network
with buffered, aqueous salt solution should be carried out after
the IPN is fully formed.
[0119] These results demonstrate that the PEG/PAA IPN system
strain-hardens and, in turn, becomes "pre-stressed" with high
values for initial stiffness moduli when swollen in buffers of
physiologic pH and salt concentrations (e.g. phosphate buffered
saline). The strain hardening under these conditions is the result
of the constraining effect that the tightly crosslinked, neutral
PEG network has on the swelling of the ionized PAA network. This
constraining effect leads to additional physical crosslinks between
the two networks and manifests as an increase in the initial
Young's modulus of the IPN. The tensile modulus values that the
hydrogel can attain (12 MPa, but tunable between about 1 to about
20 MPa) exceed those reported in the art. Of note, the hydrogel's
modulus (12 MPa) is in the range of values reported for natural
healthy human cartilage.
[0120] Natural cartilage is, in effect, an avascular "IPN hydrogel"
comprised of collagen and negatively charged proteoglycans. By
comparison, the IPN hydrogel comprised of PEG and negatively
charged PAA. PEG acts as the analog of collagen while PAA acts as
the analog of proteoglycans. This fundamental structural similarity
of these IPNs to natural cartilage is believed to the reason for
their functional similarity: the osmotic pressure created by the
polyelectrolyte, coupled with the steric constraint posed by the
first network, yields a "pre-stressed" material that, like
cartilage, is stiff, yet flexible, and exhibits a highly lubricious
surface. To explain the low friction coefficient that cartilage
exhibits, a number of scientific approaches have been developed:
the fluid-solid stress sharing described by the biphasic theory and
the "weeping lubrication" theory are some representative examples.
According to these theories, it is important that the material is
permeable for low friction to occur; the combination of the
permeability coefficient and the equilibrium modulus need to be
such so that to allow for the so called "weeping lubrication" but
at the same time prevent excessive fluid loss under continuous or
repeated dynamic loading. Based on the fact that the
strain-hardened IPN has similar permeability, negative charge,
water content and stiffness to natural cartilage, we hypothesize
that the IPN exhibits a low surface friction coefficient for the
same reasons natural cartilage does through any of the
aforementioned mechanisms.
[0121] We have shown that one of the defining features of the
PEG/PAA IPN is its high (compared to state-of-the-art existing
hydrogels) tensile stiffness modulus. The tensile stress-strain
behavior of the PEG(3400)/PAA(70%) hydrogel material is shown in
FIG. 17A from which the elastic tensile modulus is found to be 12
MPa. FIG. 17B presents the confined compression behavior of the
above-mentioned hydrogel from which the biphasic constants can be
determined. From the time-strain curve, the aggregate equilibrium
modulus is found to be Ha=1.56 MPa and the permeability coefficient
is K=2.4.times.10.sup.-14 m.sup.4/N/sec. In a preferred embodiment,
the strain-hardened interpenetrating polymer network hydrogel has a
permeability coefficient ranging from 1e-18 to 1e-12 m.sup.4/Nsec.
The hydrogel unconfined compression behavior is presented in FIG.
17C from which the unconfined compressive strength was found to be
18 MPa, with a failure strain under compression of over 80%. The
tensile creep behavior of the hydrogel is also depicted in FIG.
17D. Comparison of the set of hydrogel material properties to those
of cartilage shows a marked similarity.
[0122] Through pin-on-disc tribometer experiments, the wear rates
of PEG/PAA hydrogel in PBS and in synovial fluid under physiologic
contact stresses were determined; the hydrogel was tested for
3,000,000 cycles at .about.1 Hz loading frequency and the linear
wear rate was found to be 0.2 .mu.m/million cycles equivalent to
about 0.2 .mu.m/year, suggesting that based on the thickness of the
bearing region 5 wear life of the device suffices for a lifetime.
The material was also tested in a gel-on-cartilage configuration
under dynamic physiologic loading conditions. The test was carried
out for 150,000 cycles at a sliding frequency of 1 Hz, and a
0.5-1.5 MPa dynamic loading in a synovial fluid and bovine serum
solution. Gross observation showed that neither the cartilage nor
PEG/PAA showed any macroscopically discernible fibrillation or
wear.
Anchoring Specifications
[0123] Initial anchoring of the device is made possible by the
stretch-to-fit fixation provided by the slight size difference
between the hydrogel device and the underlying bone. The polymer
cap is placed over the femoral head, creating a snug, compressive
fit over the bone. In the case of a concave joint such as the hip
socket, a slightly oversized female-type implant creates an
expansion fit against the walls of the joint.
[0124] Biological anchoring of the device is achieved by means of
osteointegration with the inorganic constituents of bone. In the
present invention, calcium and phosphate ions are bound to PEG/PAA
IPNs through their affinity for the PAA component of the hydrogel
as illustrated in FIG. 3. Hydroxyapatite (HAP) is the major
inorganic component of bones and teeth comprised of calcium and
phosphate ions and is a known promoter of osteoblast growth. In the
dental industry, polycarboxylate cements are used to adhere
artificial substrates (e.g. dental caps) to enamel. The basis of
these cements is electrostatic interaction between the carboxylic
acid groups of PAA chains and the calcium phosphate matrix that
makes up HAP. Two mechanisms have been proposed, one in which the
carboxylic acid groups displace calcium phosphate in HAP and
essentially "insert" into the matrix, and the other (which may work
synergistically) in which calcium crosslinks HAP and PAA by
ion-bridging. In experiments to show that calcium-containing bone
constituents can bind to the PEG/PAA IPN, hydroxyapatite (HAP), a
known osteo-conductive bone mineral, was coated onto the surface
PEG/PAA IPNs. A variety of hydroxyapatite particle sizes were able
to bind to PEG/PAA. PEG/PAA hydrogels were incubated in 10% w/v
aqueous suspensions of HAP in deionized water; this led to visible
binding of HAP particles on the surface of the hydrogels.
Incubation of the hydrogels in aqueous suspensions of HAP particles
of different diameters (ranging from 20 nm to 5 .mu.m), yielded a
thick, opaque surface layer on the hydrogels. The samples were then
prepared for scanning electron microscopy (SEM) analysis by
processing them in graded ethanol solutions. Immersion in ethanol
removed the physisorbed, visible layer of HAP. SEM revealed
differences in the surface morphology of uncoated hydrogels (FIG.
23A) versus hydroxyapatite-coated hydrogels (FIG. 23B). Energy
dispersive x-ray (EDX) spectroscopy (FIG. 23C) revealed the
presence of calcium and phosphate on the surface of the hydrogel in
a ratio of approximately 1.5-1.6, which is characteristic of
hydroxyapatite. SEM coated hydrogel (inset) showed that the HAP
(200 nm diameter, shown) was localized to its surface. The
biological response to the particles was also studied by seeding
osteoblast-like cells (MG-63 cell line) on the
hydroxyapatite-coated hydrogels (FIG. 23D). The osteoblast-like
cells exhibited evidence of spreading and growth on HAP coatings of
200 nm diameter and higher.
[0125] Three different sized particles (20 nm, 200 nm, and 5 .mu.m)
of HAP were investigated to determine the effect of particle size
on surface coverage on the hydrogel as well on the biological
response by osteoblast-like cells. FIGS. 24A-24C shows SEM images
of the three types of HAP used on both bare silica (FIG. 24A) and
the PEG/PAA hydrogels, shown in the FIG. 24B (center) and FIG. 24C
(bottom) rows at low and high magnification, respectively. These
images demonstrate that surface coverage of the hydrogels was
inversely related to the particle diameter: the smaller the
particle, the more evenly and thoroughly distributed it is on the
hydrogel. This surface modification strategy takes advantage of
electrostatic interactions between inorganic hydroxyapatite and the
negative charge density of PAA. The hydroxyapatite can either be
pre-coated on the device prior to implantation in the body, or be
coated in vivo as the bone adjacent to the device is remodeled.
Chemical Anchoring
[0126] FIGS. 25A-B show according to the present invention an IPN
network bonded to bone through a separate polymeric adhesive. A
pre-existing IPN hydrogel 10, 11 is placed over bone 3, 4 that is
either functionalized with UV-sensitive crosslinkable groups or not
treated at all. At the interface between the hydrogel and the bone
is a precursor solution of reactive monomers 18 or macromonomers
21. These monomers or macromonomers partially penetrate the matrix
of the interpenetrating polymer network. Upon initiation of
polymerization, the monomers or macromonomers polymerize and
crosslink, yielding an intervening polymer that is bonded to the
underlying surface and physically entangled and/or chemically
bonded with the hydrogel.
[0127] In one example of this anchoring approach, the
heterobifunctional crosslinking agent,
3-trimethoxysilylpropylmethacrylate at a concentration of 0.1% w/v
in 95% ethanol in deionized water (with pH-adjusted to 4.5) was
brushed onto the surface of previously cleaned and dried bovine
bone and allowed to dry for 15 minutes and react with the
phosphates in the inorganic matrix of the bone. A 25% w/v solution
of PEG-dimethacrylate (MW 1000 Da) was then prepared along with 1%
v/v 2-hydroxy-2-methyl propiophenone as the photointiator and then
spread over the bone-interface surface of a PEG/PAA IPN hydrogel.
The PEG-dimethacrylate solution was then allowed to diffuse into
the IPN hydrogel for 1 hour. Bone was then placed on top of the
PEG-dimethacrylate solution on the IPN hydrogel, and then the bone
and the hydrogel were clamped together using a binder clip and
glass slide (1.0 mm thick) placed on top of the hydrogel to attain
even clamping pressure. The specimen was then placed under a UV
light source (350 nm) for 45 seconds to cause the
PEG-dimethacrylate to cure. The result was a PEG/PAA IPN hydrogel
bonded to the bovine bone specimen through a PEG-dimethacrylate
adhesive that is interpenetrated within the bone-interface of the
IPN (FIG. 25B). Because of the presence of methacrylate groups on
the bone through reaction of the trimethoxypropylsilyl methacrylate
to the bone, the PEG-dimethacrylate adhesive not only filled in the
pores of the bone but also is chemically bonded to the surface.
Another example of a "bone-primer" is isocyanatotrimethoxysilane,
which after reacting with the inorganic part of bone yields
reactive isocyanate groups on the surface, which are available to
react with functional groups (such as hydroxyl, amine, or
carboxylic acid) on either the bone-interface of the device itself
or an adhesive. This method can be used with or without silane
functionalization of the underlying bone, as well as with other
crosslinkable polymers.
[0128] FIG. 26 shows according to the present invention a
semi-interpenetrating network in which one of the networks acts as
the anchoring intervening polymer. Telechelic macromonomers 13, 15
and second network polymer 11 are mixed together in solution and
cast over a bone surface that is pre-coated and/or functionalized
with UV-sensitive crosslinkable groups 23. Exposure to an
initiating source (e.g. UV light) in the presence of a
photoinitiator leads to free-radical polymerization and
crosslinking of these crosslinkable groups on both the telechelic
macromonomers and the coated/functionalized bone surface. The
result of free-radical polymerization and cros slinking is shown on
the right. The ends 15 of the telechelic macromonomers have
polymerized and have formed physical and/or chemical bonds with the
surface of the bone. The linear second network polymers 11 are
physically trapped within this first network, forming a second,
physically crosslinked network interpenetrating the first
chemically crosslinked network 10.
Chemical Surface Modification
[0129] An embodiment of the device according to the present
invention comprises a bearing region and bone-interfacing region
with two different polymeric compositions. In general, this
approach leads to a composition gradient within the device as
described in FIG. 2. FIG. 27A shows an embodiment of the present
invention a fully interpenetrating network in which a third network
precursor is partially interpenetrated within the pre-existing IPN
by interdiffusion of the monomer for a predetermined time and then
polymerized and crosslinked in the presence of the IPN. This yields
what is effectively a triple network on one side of the IPN
hydrogel that can serve as a bone-interfacing region, which has
different properties than the other side containing only two
networks. The transition zone between the two sides is determined
by the diffusion depth of the third network monomers prior to
polymerization of the third network.
[0130] FIG. 27B shows another embodiment of the present invention a
fully interpenetrating network in which one of the networks is
interfacially copolymerized with another polymer that acts as the
bone-interfacing material. A pre-existing homopolymer network is
swollen with the precursor monomers 14 of a second network. At the
bone-interface side of the material is a precursor solution of
another reactive monomer 26. These monomers partially penetrate the
matrix of the overlying interpenetrating polymer network. Upon
exposure to UV, the monomers co-polymerize, yielding a material
with a one type of IPN containing 10 and 11 on the bearing side and
another type of IPN containing 10 and 27 on the bone-interfacing
side. The transition between the two sides is determined by the
diffusion depth of the third monomers 26 prior to polymerization of
the third network.
[0131] Another embodiment of the present invention is to use an
external stimulus to create a composition gradient in the second
network within the first network of the IPN as illustrated in FIG.
27C. In one example, instead of just acrylic acid monomers for the
second network precursor solution, a mixture of ionizable monomer
14 (e.g. acrylic acid) and non-ionic monomers 28 (e.g. acrylamide,
N-isopropylacrylamide, or hydroxylethylacrylate monomers) is used.
Any combination of ionizable monomer and non-ionizable monomer can
be used as comonomers to create the gradient so long as they are
capable of copolymerizing with each other. The first network 10 is
soaked in a salt solution of ionizable monomer 14, non-ionic
monomer 28, crosslinker and photoinitiator (not shown) and then an
electric field is applied to the gel (e.g. using electrophoresis
equipment). Only the acrylic acid monomers will move along the
electric field due to their charge. After formation of an acrylic
acid concentration gradient, the gel is exposed to UV and the
gradient is fixed via second network gel formation. The result is
an IPN hydrogel with a poly(acrylic acid) second network localized
to the bearing region and a non-ionic second network (e.g.
poly(N-isopropylacrylamide, a temperature-sensitive polymer)
localized to the bone-interface region. This is an approach that
yields a device that is responsive to both pH and temperature, as
described later in FIG. 32.
[0132] FIG. 28 shows two embodiments of another device surface
modification strategy according to the present invention. This
strategy involves the acrylation/methacrylation of an
amine-containing or hydroxyl-containing molecule or biomolecule by
reaction with a halogenated (active) acid (e.g. acryloyl chloride)
(FIG. 28, Reaction A) or with an active ester (e.g.
acryloxy-N-hydroxysuccinimide) (FIG. 28, Reaction B) to make it
capable of copolymerizing with the precursor of one of the networks
in the device. The R-group in these reaction schemes can be any
amine-containing or hydroxyl-containing chemical or polymer,
proteins, polypeptides, growth factors, amino acids, carbohydrates,
lipids, phosphate-containing moieties, hormones, neurotransmitters,
or nucleic acids. An example of this process is the reaction of
dopamine with acryloyl chloride and subsequent attachment of the
conjugated dopamine molecules to the surface of a PEG/PAA hydrogel
during the second network formation by the process shown in either
FIG. 27B or C. Dopamine hydrochloride (500 mg, 2.6 mmol, 1 eq) was
dissolved in methanol (10 mL) and freshly distilled triethylamine
(362 .mu.L, 1 eq) was added. Acryloyl chloride (210 .mu.L, 1 eq)
was dissolved separately in MeOH and Triethylamine (1.1 mL, 3 eq)
was added. The acryloyl chloride solution was then added dropwise
to the dopamine solution and the resulting mixture was stirred
overnight at room temperature (Reaction A). During the reaction, a
colorless precipitate formed that was removed by filtration.
Precipitation in diethylether lead to the product, an acrylated
dopamine molecule (yield: 85%). In an alternative reaction
(Reaction B) to achieve the same result, dopamine hydrochloride
(500 mg, 2.6 mmol, 1 eq) was dissolved in methanol (10 mL) and
freshly distilled Triethylamine (362 .mu.L, 1 eq) was added.
Acrylic acid N-hydroxysuccinimide ester (440 mg, 1 eq) was
dissolved separately in methanol and triethylamine (1.1 mL, 3 eq)
was added. The acrylic acid N-hydroxysuccinimide ester solution was
then added dropwise to the dopamine solution and the resulting
mixture was stirred overnight at room temperature. During the
reaction a colourless precipitate formed that was removed by
filtration. Precipitation in diethylether lead to the product
(yield: 75%). The resulting conjugated molecule was then
interfacially polymerized with an acrylic acid-based second network
in separate experiments (one using the Reaction A conjugate and one
using the Reaction B conjugate) as shown in FIG. 27B. A 50% v/v
solution of dopamine acrylate containing 1% v/v
2-hydroxy-2-methyl-propiophenone and 1% v/v triethylene glycol
dimethacrylate was spread the surface of a preformed PEG-diacrylate
network that had been dabbed dry after being swollen overnight in a
50% v/v solution of acrylic acid, 1%
2-hydroxy-2-methyl-propiophenone and 1% triethylene glycol
dimethacrylate. After briefly allowing the dopamine-acrylate
monomers to mix with the acrylic acid monomers, the swollen gel was
placed between glass slides and exposed to UV. The result was an
IPN with a PEG/PAA IPN on one side and an IPN of PEG and a
dopamine-conjugated polymer network on the other surface. In the
transition zone between these was an IPN of PEG and a copolymer of
PAA and dopamine-conjugated polymer. This method can be generalized
to attain a variety of types of conjugates of the IPN surface.
[0133] Another embodiment of the device according to the present
invention covalently links molecules or biomolecules to a
pre-fabricated device in order to create a bone-interface region
with different characteristics than the bearing region. In one such
embodiment, any suitable biomolecules may be covalently linked to
the IPN hydrogel. In another embodiment, a synthetic polymer is
linked to the IPN hydrogel. Preferably, the biomolecules are at
least one of proteins, polypeptides, growth factors (e.g. epidermal
growth factor) amino acids, carbohydrates, lipids,
phosphate-containing moieties, hormones, neurotransmitters, or
nucleic acids. Any combination of small molecules or biomolecules
can be used, including, but not limited to, drugs, chemicals,
proteins, polypeptides, carbohydrates, proteoglycans,
glycoproteins, lipids, and nucleic acids. This approach may rely,
for example, on (a) photoinitiated attachment of azidobenzamido
peptides or proteins, (b) photoinitiated functionalization of
hydrogels with an N-hydroxysuccinimide ester, maleimide, pyridyl
disulfide, imidoester, active halogen, carbodiimide, hydrazide, or
other chemical functional group, followed by reaction with
peptides/proteins, or (c) chemoselective reaction of aminooxy
peptides with carbonyl-containing polymers. These biomolecules may,
for example, promote bone cell adhesion or activity. In one
example, a heterobifunctional crosslinker 118 (FIG. 29) with
reactive endgroups 115 and 117 joined by a spacer arm 116 is used
to modify the IPN hydrogel surfaces 119. One such class of
heterobifunctional chemicals are described as azide-active-ester
linkers, such as 5-azido-2-nitrobenzoyloxy-N-hydroxysuccinimide
ester or its derivatives such as its sulfonated and/or its
chain-extended derivatives. However, any coupling strategy can be
used to create strain-hardened IPN hydrogels with bioactive
surfaces. A detailed example of this embodiment is the attachment
of collagen type I to a PEG/PAA IPN surface through the
heterobifunctional crosslinker,
5-azido-2-nitrobenzoyloxy-N-hydroxysuccinimide ester, which has a
phenyl azide group on one end and a protein-binding
N-hydroxysuccinimide group on the other. Substituted phenyl azides
have been shown to react with light (250-320 nm, 5 min) to generate
aromatic nitrenes, which insert into a variety of covalent bonds.
Attachment of the linker to the hydrogel via the phenyl azide group
then allows the N-hydroxysuccinimide groups to react with free
amines on proteins, and in turn, tether them to the hydrogel
surface. The surfaces of the PEG/PAA hydrogels were dabbed dry and
then 100 .mu.L of a 0.5% w/v solution of 5-azidonitrobenzoyloxy
N-hydroxysuccinimide in dimethylformamide was drop-casted onto the
gel and spread evenly over its surface. The solvent was then
allowed to evaporate under a fume hood to ensure deposition of the
crosslinker onto the hydrogel. The air-dried gel surface was then
exposed to UV light for 5 min to react the azide groups to the
hydrogel surface. The surface-functionalized gels were then
incubated in a 0.3% (w/v) collagen type I solution (Vitrogen) in a
37.degree. C. oven for 16 hours to couple reactive protein amine
groups to the N-hydroxysuccinimide moieties on the hydrogel
surface. Finally, the gels were washed extensively in PBS to remove
organic solvent and unreacted monomers. The presence of tethered
protein on the surface was confirmed by X-ray photoelectron
spectroscopy, which showed the presence of amide linkages of the
surface of the hydrogel, confirming the presence of protein. Table
6 shows quantitative amino acid analysis data showing the presence
of collagen on the surface of the gels.
TABLE-US-00006 TABLE 6 Results of quantitative amino acid analysis
on collagen- tethered PEG/PAA hydrogels (in total micrograms).
Residue Reaction 1* Reaction 2** Reaction 3*** Asx 3.18 .+-. 1.01
2.62 .+-. 0.26 2.37 .+-. 0.39 Thr 1.36 .+-. 0.50 1.10 .+-. 0.10
0.97 .+-. 0.15 Ser 1.73 .+-. 0.55 1.40 .+-. 0.14 1.28 .+-. 0.19 Glx
8.56 .+-. 2.75 7.58 .+-. 0.80 6.90 .+-. 1.13 Pro 7.62 .+-. 2.34
6.18 .+-. 0.74 5.64 .+-. 0.97 Gly 11.78 .+-. 3.06 9.75 .+-. 1.10
8.92 .+-. 1.48 Ala 4.84 .+-. 1.47 3.96 .+-. 0.44 3.60 .+-. 0.60 Val
1.55 .+-. 0.50 1.09 .+-. 0.12 0.93 .+-. 0.15 Ile 0.90 .+-. 0.29
0.69 .+-. 0.06 0.60 .+-. 0.10 Leu 1.80 .+-. 0.58 1.38 .+-. 0.14
1.17 .+-. 0.19 Tyr 0.18 .+-. 0.06 0.13 .+-. 0.01 0.11 .+-. 0.01 Phe
0.96 .+-. 0.32 0.77 .+-. 0.09 0.70 .+-. 0.12 His 0.52 .+-. 0.17
0.34 .+-. 0.04 0.29 .+-. 0.05 Lys 1.97 .+-. 0.63 1.70 .+-. 0.17
1.57 .+-. 0.25 Arg 4.97 .+-. 1.62 3.89 .+-. 0.47 3.52 .+-. 0.59 Hy
Pro 6.80 .+-. 1.99 5.91 .+-. 0.71 5.51 .+-. 0.96 Hy Lys 0.78 .+-.
0.12 0.63 .+-. 0.06 0.55 .+-. 0.09 Total 59.50 .+-. 17.94 49.13
.+-. 5.46 44.63 .+-. 7.42 *Reaction 1 involved incubation of the
hydrogels with 0.3% w/v collagen type I *Reaction 2 involved
incubation of the hydrogels with 0.1% w/v collagen type I *Reaction
3 involved incubation of the hydrogels with 0.03% w/v collagen type
I
[0134] FIG. 30 shows another embodiment of the present invention to
attain a different surface chemistry at the bone-interface than
that present in the bearing region. This approach involves
activating the functional groups on the surface of the hydrogel
followed by reaction of these activated functional groups with
amine-containing or hydroxyl-containing molecules, macromolecules,
or biomolecules. In a preferred embodiment, the carboxylic acid
groups on poly(acrylic acid) within an IPN are activated to form an
active ester, which subsequently forms acrylamide linkages when
reacted with an amine-containing molecule, macromolecule, or
biomolecule. In two examples of this strategy, a PEG/PAA IPN
hydrogel according to the present invention was surface modified
with dopamine functional groups. In Reaction A, the PEG/PAA
hydrogel was first washed with ethanol/water mixtures containing
increasing amounts of ethanol up to 100 vol. % ethanol. The
hydrogel was then soaked in a solution of dicyclohexylcarbodiimide
(0.1 M) and Triethylamine (0.2 M) in ethanol for 2 hours. A
solution of dopamine hydrochloride (0.1 M) and triethylamine (0.1
M) was prepared and applied onto the surface of the gel. After one
hour, the hydrogel was washed with ethanol and then with
ethanol/water mixtures containing increasing amounts of water up to
100 vol. % water. The resulting hydrogel had dopamine molecules
attached to the hydrogel surface through amide linkages where the
carboxylic acids once were. In an alternative to this procedure
(FIG. 31, Reaction B), the PEG/PAA hydrogel was soaked in a
solution of N-hydroxysuccinimide (15 mM) and
N-Ethyl-N'-(3-dimethylaminopropyl)carbodiimide (75 mM) in phosphate
buffer (10 mM, pH 6) for one hour. After washing with buffer and
water, the surface of the gel was exposed to a solution of dopamine
hydrochloride in DMF (0.1 M) and triethylamine (0.1 M) for one
hour. The hydrogel was then washed with DMF, ethanol and water to
remove all excess material to yield the hydrogel with dopamine
tethered to its surface. These reactions can be used to tether any
molecule, macromolecule, or biomolecule with accessible amine or
hydroxyl functional groups to the surface of carboxyl-group
containing IPNs. The resulting surface-modification would then be
used as the basis of a bone-interface region of the present
invention, with the unmodified side serving as the bearing
region.
Stimulus-Responsive Hydrogel Arthroplasty Devices
[0135] Implantation of the device through volume changes in the
device can be achieved by taking advantage of the
stimulus-responsiveness of certain polymers. In addition,
fabricating the device with different polymer compositions in the
bearing and bone-interfacing regions makes offers an additional
level of control over the implantation of the device via external
stimuli while preserving certain advantageous attributes of a
non-responsive polymer or by introducing new attributes to the
responsive polymer. Stimuli hereafter refers to a characteristic
change in a property that regulates hydrogel volume or shape; this
change is caused by maintaining the hydrogel pre-surgically in an
environment that is different than the environment inside the body.
In an embodiment of the present invention, an external stimulus
such as a change in pH, salt concentration, electric field, or
temperature causes the device, after A being placed on the bone, to
B shrink to conform to the contours of the convex-shaped bone it
surrounds, as depicted in FIG. 32. For a concave joint, the device
is designed such that the stimulus causes the device to expand
against the concavity. Polyelectrolytes are a class of hydrogel
polymers that swell/deswell to varying degrees in response to
changes in pH, salt concentration, and electric field. Changing pH
and salt to control swelling and hydrogel device size would work in
the following manner. In one example, the device is pre-swollen in
a state where the cap is slightly larger than a convex joint
surface, and then after placement on the joint, it would be
deswelled by the change after equilibrium in the pH or salt
concentration that is present inside the body. The pH/salt
concentration can be changed by external means (such as immersing
the implant/joint in a bath prior to surgery). Alternatively, it
can be implanted and allowed to reach equilibrium swelling in
response to the pH and salt concentration of the surrounding body
fluids (e.g. synovial fluid). Interpenetrating networks with
polyelectrolyte components (e.g. poly(acrylic acid)) such as
poly(ethylene glycol)/poly(acrylic acid) networks would be
particularly useful in this regard. If this material is preswollen
at pH>7.4 and/or salt concentration of less than the osmolarity
of the body and is placed loosely over a joint surface, it will,
after some time equilibrating in the body, shrink in response to
the decrease in pH and/or increase in salt concentration and
conform to the contours of the underlying bone. The dimensions of
polyelectrolyte-based IPNs can also be modulated by application of
an electric field which electrically expands the device. After the
electric field is removed, the device shrinks again over the joint.
Temperature-sensitive hydrogels such as poly(N-isopropylacrylamide)
(NIPAAm) have a lower critical solution temperature that causes
them to contract at temperatures higher than about 32.degree. C.
This makes possible a scenario where a NIPAAm-based device is
placed loosely over a joint at the time of implantation, and after
some time in the body, it shrinks to conform to the contours of the
bone it surrounds, as depicted in FIG. 32. Thus, using stimuli to
alter the hydrogel device size slightly at the time of implantation
facilitates its placement without physically stretching it by hand
or with a tool, enabling less invasive or arthroscopic approaches
for surgical placement.
Variations and Modifications
[0136] The interpenetrating polymer networks could have two or more
networks or polymeric components (such as linear chains). Examples
include but are not limited to a "triple" or even "quadruple"
network or a double network interpenetrated with additional polymer
chains as discussed in FIGS. 25 and 27. In addition, polymeric
tethers (such as poly(ethylene glycol) chains) can be used as
intervening spacer arms between the bone-interface region and
tethered biomolecules or attached polymer materials.
[0137] As one of ordinary skill in the art will appreciate, various
changes, substitutions, and alterations could be made or otherwise
implemented without departing from the principles of the present
invention. Accordingly, the scope of the invention should be
determined by the following claims and their legal equivalents.
* * * * *