U.S. patent application number 12/495688 was filed with the patent office on 2010-12-30 for biphasic implant device transmitting mechanical stimulus.
Invention is credited to Robert L. McDade, Timothy A. Ringeisen.
Application Number | 20100331979 12/495688 |
Document ID | / |
Family ID | 56291120 |
Filed Date | 2010-12-30 |
![](/patent/app/20100331979/US20100331979A1-20101230-D00000.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00001.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00002.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00003.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00004.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00005.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00006.png)
![](/patent/app/20100331979/US20100331979A1-20101230-D00007.png)
United States Patent
Application |
20100331979 |
Kind Code |
A1 |
McDade; Robert L. ; et
al. |
December 30, 2010 |
Biphasic implant device transmitting mechanical stimulus
Abstract
Tissue implants prepared for the repair of tissues, especially
avascular tissues such as cartilage. One embodiment presents an
electric potential capable of receiving and accumulating desirable
factors or molecules from surrounding fluid when exposed to dynamic
loading. In another embodiment the implant promotes tissue
conduction by retarding, restricting and controlling cellular
invasion through use of gradients until competent tissue forms.
Further embodiments of the tissue implants may be formed into a
multi-phasic device that provides deep tissue mechanical stimulus
by conduction of mechanical and fluid forces experienced at the
surface of the implant.
Inventors: |
McDade; Robert L.;
(Downingtown, PA) ; Ringeisen; Timothy A.; (Exton,
PA) |
Correspondence
Address: |
KENSEY NASH CORPORATION
735 PENNSYLVANIA DRIVE
EXTON
PA
19341
US
|
Family ID: |
56291120 |
Appl. No.: |
12/495688 |
Filed: |
June 30, 2009 |
Current U.S.
Class: |
623/14.12 |
Current CPC
Class: |
A61F 2002/30014
20130101; A61F 2250/0018 20130101; A61F 2002/30332 20130101; A61L
27/50 20130101; A61F 2250/0051 20130101; A61F 2250/0043 20130101;
A61L 27/48 20130101; A61F 2002/4495 20130101; A61L 27/52 20130101;
A61F 2310/00365 20130101; A61F 2220/0033 20130101; A61L 2430/06
20130101; A61L 27/56 20130101; A61F 2002/30052 20130101; A61F 2/28
20130101; A61F 2210/008 20130101; A61F 2002/30028 20130101; A61F
2250/0056 20130101; A61L 27/24 20130101; A61F 2/30756 20130101;
A61F 2002/30031 20130101; A61F 2002/3092 20130101; A61F 2002/30088
20130101; A61F 2002/30971 20130101 |
Class at
Publication: |
623/14.12 |
International
Class: |
A61F 2/08 20060101
A61F002/08 |
Claims
1. A multiphasic device for repair or replacement of articular
cartilage and the underlying bone, said device comprising a bone
region and a malleable cartilage region said bone region comprising
a porous material and a rigid penetrating force conductive material
for conducting pressure forces experienced by said cartilage region
through said bone region and into said external underlying bone
tissue, thereby preventing bone tissue loss in said external
underlying bone tissue.
2. The device of claim 1 wherein the pressure forces are conducted
as hydrostatic pressure pulses along rigid conductive elongated
channels running through the porous material of the bone phase and
radiating out and away from the cartilage phase.
3. The device of claim 1 wherein the rigid conductive material is
in the form of one or more vertical rigid columns oriented
perpendicular to the top and bottom surfaces that flare out as they
reach the underlying external bone, said vertical rigid columns
transforming the fluid forces into kinetic pressure pulses that are
carried through the porous material and into the underlying
bone.
4. The device of claim 1 wherein the rigid conductive material is a
rigid multi-facetted web structure oriented perpendicular to the
top and bottom surfaces, said web structure transferring the fluid
forces into kinetic pressure pulses that are carried through the
porous material and into the underlying bone.
5. The device of claim 1 wherein the rigid conductive material is
in the form of one or more vertical rigid cones or wedges oriented
perpendicular to the top and bottom surfaces with the base located
towards the cartilage phase, said vertical rigid columns
transforming the fluid forces into kinetic pressure pulses that are
carried through the porous material and into the underlying bone
while at the same time applying an outward force to the porous
material.
6. The device of all the above claims wherein the rigid conductive
material extends partially into the malleable cartilage phase.
7. The device of all the above claims wherein the rigid conductive
material extends partially into the external underlying bone.
8. The device of all the above claims wherein at least one of the
materials is bioresorbable
9. The device of all the above claims wherein the rigid conductive
material is porous.
10. The device of all the above claims wherein the rigid conductive
material is selected from the groups consisting of metals,
ceramics, polymers, glasses, or combinations thereof.
11. The device of all the above claims wherein the porous material
is a polymer.
12. The device of all the above claims wherein the porous material
contains an additive in the form of a particulate or biologically
active agent.
13. The device of all the above claims wherein the rigid conductive
material contains an additive in the form of a particulate or
biologically active agent.
14. A multiphasic device for repair or replacement of articular
cartilage and the underlying bone, said multiphasic device
comprising a bone region and a malleable cartilage region, said
bone region comprising at least a first material in the form of at
least two porous rigid scaffolds separated by at least a second
material in the form of a malleable elastic hydrogel, said hydrogel
being capable of transferring hydrostatic pressure pulses through
the bone region, said bone region preventing bone voids from
forming in external underlying bone tissue by conducting pressure
forces experienced in said malleable cartilage phase through said
bone region and into said underlying bone tissue.
15. The multiphasic device of claim 14 wherein the porous rigid
materials are in the form of disks having a thickness of 1000
microns to 4000 microns, separated by a hydrogel having a thickness
of no less that 5 microns.
16. The multiphasic device of claim 14 wherein the porous rigid
materials are in the form of porous particles having an approximate
diameter in the range of 1000 microns to 4000 microns, suspended in
a hydrogel, and separated from each other by no less than 5
microns.
17. The multiphasic device of claim 14 wherein at least one of the
materials is bioresorbable.
18. The multiphasic device of claim 14 wherein the device becomes
malleable upon hydration.
19. The device of claim 14 wherein the porous rigid material is
selected from the groups consisting of metals, polymers, ceramics,
glasses or combinations thereof.
20. The device of claim 14 wherein the matrix contains an additive
in the form of a particle or biologically active agent.
21. The device claim 14 wherein the porous rigid material contains
an additive in the form of a particulate or biologically active
agent.
22. The device claim 14 wherein the hydrogel contains an additive
in the form of a particulate or biologically active agent.
Description
BACKGROUND OF THE INVENTION
[0001] What is disclosed is a device for repairing and replacing
lost or damaged tissue. Particularly, one embodiment is directed to
a multi-phasic prosthetic device for repairing or replacing
cartilage or cartilage-like tissues. Said prosthetic devices are
useful as articular cartilage substitution material and as a
scaffold for regeneration of articular cartilaginous tissues.
[0002] Cartilage is found throughout the body, such as in the
supporting structure of the nose, ears, ribs (elastic cartilage),
within the meniscus (fibrous cartilage), and on the surfaces of
joints (hyaline cartilage or articular cartilage). A joint is a
bending point where two bones meet. The knee, hip, and shoulder are
the three largest joints.
[0003] The specialized covering on the ends of bones that meet to
form an articulating joint is called hyaline or articular
cartilage. It is the cartilage that is damaged and wears as one
ages, or sustains an injury. Articular cartilage is unique amongst
the body tissues in that it has no nerves or blood supply. This
means that damage will not be felt until the covering wears down to
bare underlying bone. Bone is very sensitive and the sharp pain of
arthritis often comes from irritation of bone nerve endings and
since human tissue has a very limited capacity to heal without a
blood supply, articular cartilage cannot repair itself
effectively.
[0004] Articular cartilage tissue covers the ends of all bones that
form diarthrodial joints. The resilient tissues provide the
important characteristic of friction, lubrication, and wear in a
joint. Furthermore, it acts as a shock absorber, distributing the
load to the bones below. Without articular cartilage, stress and
friction would occur to the extent that the joint would not permit
motion. As stated above, articular cartilage has only a very
limited capacity to regenerate. If this tissue is damaged or lost
by traumatic events, or by chronic and progressive degeneration, it
usually leads to painful arthrosis and decreased range of joint
motion.
[0005] Articular cartilage repair following injury or degeneration
represents a major clinical problem, with treatment modalities
being limited and joint replacement being regarded as appropriate
only for the older patient.
[0006] Current treatments for articular cartilage damage are varied
and include anti-inflammatory medication, viscosupplementation,
arthroscopic chondroplasty, autogenous articular cell implantation,
microfracture and osteochondral articular transplantation.
[0007] Anti-inflammatory medication: Aspirin was the first
anti-inflammatory medication in the world. This was followed in
1950 by cortisone (steroidal medication) used orally or by
injection. (Extensive use of cortisone not only has a wide variety
of harmful effects, but is also believed to harm cartilage.) Later
the non-steroidal drugs such as Motrin came along. These were safer
than Aspirin and cortisone but had potent side effects, especially
causing bleeding within the stomach and intestinal ulcers. These
complications led to the development of the COX-2 inhibitor drugs,
Celebrex and Vioxx. While much safer and seemingly more effective,
Vioxx was found to have significant cardiac side effects and is no
longer available. With certain precautions, Celebrex is still
widely used. However, these anti-inflammatory medications only
treat the symptoms of cartilage damage and arthritis and do not
promote repair.
[0008] Viscosupplementation: Viscosupplementation is a procedure
that involves the injection of gel-like substances (hyaluronates)
into a joint to supplement the viscous properties of synoval fluid.
Currently, hyaluronate injections are approved for the treatments
of osteoarthritis of the knee in those who have failed to respond
to more conservative therapy, Once again, this procedure only
treats the symptoms of cartilage damage and arthritis and does not
promote repair.
[0009] Arthroscopic chondroplasty: Chondroplasty is a term
referring to the arthroscopic smoothing of unstable articular
surfaces either with mechanical shaving or thermal devices. While
not a restorative measure, so called debridement can be useful in
reducing irritating cartilage debris that breaks off in the joint
or causes catching or grinding sensations. The resulting
improvement in the control of inflammation can last for several
years. But this is not a final solution as the degenerative process
continues to wear away at the articular cartilage.
[0010] Autogenous articular cell implantation (ACI): Autogenous
cell implantation can be used for large, shallow defects, which do
not involve the subchondral bone. In this procedure, cartilage
cells collected from the patient and grown to many millions through
cell culture techniques are injected into the joint, under a
membrane that has been attached to the cartilage surface. Although
successful, the window of opportunity for this procedure is often
missed, as the few clinical symptoms showing the need for this
treatment are not evident until the defect deepens to involve the
underlying bone, thus the damage encountered upon detection is
frequently too extensive for repair through ACI.
[0011] Microfracture: The goal of this arthroscopic technique is to
improve the blood supply to the bare areas of the joint by creating
tiny perforations in the underlying bone. The resulting bone marrow
bleeding carries powerful growth stimulating factors found in
platelets as well as stem cells to the damaged area creating what
is referred to as a super-clot. Healing and repair follow over
several weeks. Studies have shown that microfracture techniques do
not fill in the chondral defect fully and the repair material
formed is fibrocartilage. The fibrocartilage tissue can temporarily
return function for activities such as running and a sport play,
but ultimately fails, as fibrocartilage is unable to mechanically
share and disipate loading forces as effectively as the original
hyaline cartilage. Fibrocartilage is much denser and isn't able to
withstand the demands of everyday activities as well as hyaline
cartilage and is therefore at higher risk of breaking down.
[0012] Osteochondral articular transplantation: Osteochondral
transplantation (i.e. mosaioplasty) involves transportation of
tissue plugs from one location of the knee to another. Special
instrumentation has been devised to harvest plugs of articular
cartilage and its supporting bone from the patient's own joint. The
harvested tissue is then transported to the damaged site where it
is inserted into prepared holes. Several plugs can fill up rather
larger defects and will grow to re-supply a new joint surface.
Unfortunately, this procedure leaves defects of equal or worse
proportions elsewhere and often the harvested tissue is not viable
due to the traumatic harvesting procedure.
[0013] Due to the problems associated with current state of the art
treatments, much work has been done to produce a synthetic
off-the-shelf scaffold to be used in place of the harvested
osteochondral plug.
[0014] Originally, single-phase scaffolds of uniform construction
were contemplated for use as implants. However, these single-phase
scaffold implants proved unsuccessful in healing of the complex
multiphasic articular cartilage along with the underlying bone.
Soon biphasic and then gradient devices were developed that were
either mechanically or anatomically specific for the tissues
involved. While these showed an improvement over single phase
devices, it is evident that these devices do not take into
consideration how cells will be migrating into the scaffolds as
well as how their presence influences the surrounding, uninvolved
tissue. Additionally, prior art scaffolds did not take into
consideration the joint fluid and how it impacts maturation and
maintenance of healthy hyaline cartilage. Although prior art
synthetic scaffolds, whether single phase, multi-phase, or of
gradient construction have proven suitable for growth and
maturation of cells within a bioreactor, these prior art devices
are unsuitable for direct implantation, for at least the reasons
that follow.
[0015] Applicants have made the surprising discovery that in
effecting the repair of cartilage defects, prior art synthetic
implants and synthetic bi-phasic implant devices failed to
recognized the need to ignore the normal histological and
mechanical gradient of the articular cartilage, and instead focused
on the limited cell population surrounding the defect and its slow
rate of tissue formation within the devices resulting from this
sparse population of cells. The prior art synthetic implants
mistakenly focused on speeding up the rate of cell migration within
the scaffold in hopes of getting tissue to form rapidly throughout
the device prior to collapse of the scaffold. This increased rate
of cell migration was done using chemotactic ground substances such
as hyaluronic acid, cell seeding or biologics. All this served to
do was to spread out the cell population and reduce the rate of
hyaline cartilage tissue formation, and as a result, biased any new
tissue growth of cartilage towards the fibrocartilage lineage.
Although some success in establishing hyaline cartilage can be seen
in small defects of 5 millimeters or less, larger defects show tell
tale signs of collapse or dimpling in the center of a repair plug,
as the less desirable fibrocartilage, which has grown within the
prior art devices, succumbs to the forces within the joint.
Additionally, prior art devices show a halo or ring of collapsed
tissue around the periphery of the device do to lack of intimate
contact with the uninvolved tissue that has retracted away from the
defect site.
[0016] Another discovery of applicants is that prior art devices do
not address the instantaneous articular cartilage tissue
contraction that occurs when the surface of hyaline cartilage is
cut or torn. Upon damage, the cartilage retracts way from the
defect site forming a funnel. Thus prior art devices, upon
implantation, do not make contact with the surrounding uninvolved
cartilage.
[0017] The uninvolved host tissue, that is, the normal tissue
adjacent to and surrounding the defect site that is not involved
with the defect, is able to influence the activities of cells that
migrate into and establish themselves at the periphery of a
scaffold placed into the defect. The cells of the uninvolved
tissue, along with the extracellular matrix of the uninvolved host
tissue adjacent to the periphery of the implanted scaffold are
already established as hyaline cartilage and thus mechanically and
chemically react to stresses appropriately. Through a process,
sometimes referred to as mechanical signal transduction, the
established host tissue is able to influence the phenotype and
extracellular matrix produced by the adjacent cells in the scaffold
thus producing the desired hyaline cartilage. Specifically,
cartilaginous tissues perform specialized functions under normal
physiological conditions. Anomalous mechanical loading of these
tissues often leads to pathology. For example, the lack of
mechanical stimulation of a joint leads to suppression of
proteoglycan synthesis and release of mediators responsible for
degradation of cartilage matrix components. This is believed to be
the cause of collapse or dimpling of the newly formed cartilage
seen with prior art devices.
[0018] The molecular mechanisms controlling the response of
cartilaginous tissues to their mechanical environment are not
completely understood. Furthermore, there is a dearth of knowledge
about the modes of mechanical signal transduction in chondrocytes.
Several theories concerning the molecular mechanisms through which
mechanical stimuli modulate the expression of cartilage
extracellular matrix (ECM) components have been proposed, some of
which are: 1) receptor mediated cell-ECM adhesion contributes to
the transduction of mechanical signals in chondrocytes, 2)
mechanical signal transduction in chondrocytes requires activation
of the phosphoinositol and/or cyclic AMP (also known as Cyclic
adenosine monophosphate or cAMP) signaling pathways, and 3)
mechanical stimulation of the expression of aggrecan is mediated
through activation of specific cis-acting elements of the promoter
and/or UTRs (untranslated regions) of the aggrecan gene. No matter
the specific mechanism through which it happens, applicants believe
that the influence uninvolved host tissue has over the cells in the
scaffold matrix extends approximately 2.5 millimeters. Thus, this
places a limit of success for prior art devices having a matrix
equal to, or less stiff than the surrounding host tissue to 5
millimeters in diameter. However, any device having a cartilage
scaffold matrix greater in stiffness than the surrounding host
tissue will not be properly influenced by mechanical signal
transduction and will either form calcified tissues or disorganized
fibrocartilage that collapses as the matrix degrades and the tissue
experiences stress loading.
[0019] In order to prevent the observed central collapse or
dimpling within the cartilage layer of prior art implants,
applicants have discovered that a new type of scaffold must be made
that retards rapid migration of cells across the entire diameter of
the device, thereby concentrating cells and cell activity at the
edges of the device, promoting rapid and systematic tissue
conduction and maturation, moving from the outer edge of the device
towards the interior. Additionally, the area within the cartilage
region of the scaffold where cell activity is occurring must be
less rigid than the surrounding uninvolved tissue, to ensure that
it is subject to the mechanical influences of the adjacent
uninvolved tissue.
[0020] Within the bone layer, known prior art devices failed to
recognize the impact a rigid scaffold has on the surrounding
uninvolved tissue. Whereas malleable elastic scaffolds (scaffolds
that can be deformed and then return to their original shape) are
desirable for the cartilage layer, rigid stable scaffolds
(scaffolds that resist deformation) are required for proper
migration and attachment of bone forming cells. However, nearly the
opposite conditions are required for stability of existing bone, as
micromotion is beneficial to healthy bone structure. Micro-motion
and/or stresses are necessary to keep healthy bone from becoming
osteopenic. Osteopenia refers to bone mineral density that is lower
than normal. Bone mineral density has been shown to drop in healthy
individuals who are bedridden, as well as in astronauts who have
reduced stress on their skeletal system due to the effects of
reduced gravity while in space. As this occurs, the bones lose
minerals, heaviness (mass), and structure, making them weaker and
increasing their risk of collapse and or breaking. Localized bone
mineral density loss has been witnessed due to stress shielding
caused by orthopedic rods and plates. During repair of damaged
cartilage with prior art devices, voids and osteopenic zones have
been observed to form below implanted tissue scaffolds. The theory
behind this pathology formation is that stress shielding, caused by
the presence of porous tissue scaffolds, results in bone density
loss. The scaffolds dampen vibrations that would normally be
transferred through the malleable elastic articular cartilage to
the calcified region and then conducted deeper into the bone. These
conductive forces are necessary for normal bone biology. The
conducted forces in normal bone located below an articulating joint
travel not only through the bone trabecula, but also through the
viscous gel of bone marrow surrounding the bone trabecula. This is
because the bone trabecula located under the cartilage of a joint
shows a general histologic pattern of elongated channels radiating
out from the calcified region into the subchondral bone. Thus
forces are not only transmitted down the rigid walls of the
channels formed by the trabecula, but are also transmitted by the
gelatinous bone marrow contained within the channels. Two
functional problems identified with rigid porous scaffolds of prior
art devices are as follows. First these rigid devices do not
contain elongated channels and thus they tend to dissipate and
dampen the hydrostatic pressure pulses that would normally flow
through viscous fluids. Secondly these devices are too rigid
through the cartilage region thus not allowing for initial
compression to establish a pressure wave through the bone
marrow.
[0021] In order to prevent undesirable bone voids from forming in
uninvolved tissues adjacent to the repair device, what is needed is
a scaffold capable of transferring forces through the device, and
into the tissue. This deep bone mechanical stimulation is due to
compression of the articular cartilage region generating mechanical
and fluidic forces during normal movement in the joint.
[0022] Concerning the synovial fluid, prior art devices fail to
recognize the role this substance plays in maintaining healthy
articular cartilage. Synovial fluid is a thick, stringy fluid found
in the cavities of synovial joints. Synovial fluid reduces friction
between the articular cartilage surfaces as well as providing
cushioning during movement. The inner membrane of synovial joints
is called the synovial membrane and it secretes synovial fluid into
the joint cavity. This fluid forms a thin layer (about 50 microns
thick) at the surface of cartilage and seeps into the
micro-cavities and irregularities in the articular cartilage
surface, filling all empty space thus presenting a uniform, smooth
surface. The fluid in the articular cartilage effectively serves as
a synovial fluid reserve, during movement; the synovial fluid held
in the cartilage is squeezed out mechanically to maintain a layer
of fluid on the cartilage surface. This so called weeping
lubrication ensures that increased friction does not occur as some
of the lubrication fluid is swept away during joint movement.
[0023] Synovial tissue is composed of vascularized connective
tissue that lacks a basement membrane. Two cell types (type A and
type B) are present: Type B cells produce synovial fluid. Synovial
fluid is made of hyaluronic acid and lubricin, proteinases, and
collagenases. Synovial fluid exhibits non-Newtonian flow
characteristics. The viscosity coefficient is not a constant, the
fluid is not linearly viscous, and its viscosity increases as the
shear rate decreases.
[0024] Almost all of the protein constituents of synovial fluid are
derived from plasma. The passage of plasma proteins to synovial
fluid is related to the size and shape of the protein molecule.
Most proteins with molecular weights less than 100,000 daltons are
readily transferred from one fluid space to another. Thus synovial
fluid is a plasma dialysate modified by constituents secreted by
the joint tissues. The major difference between synovial fluid and
other body fluids derived from plasma is the high content of
hyaluronic acid (mucin) in synovial fluid. Normal synovial fluid
contains 3-4 mg/ml hyaluronan (hyaluronic acid), a polymer of
nonsulfated polysaccharides composed of D-glucuronic acid and
D-N-acetylglucosamine joined by alternating beta-1,4 and beta-1,3
glycosidic bonds. Hyaluronan is synthesized by the synovial
membrane and secreted into the joint cavity to increase the
viscosity and elasticity of articular cartilage and lubricates the
surfaces between synovium and cartilage. Both fibroblasts beneath
the synovial membrane intima and synovial membrane-lining cells
produce this mucopolysaccharide constituent of synovial fluid.
[0025] Synovial fluid is believed to have two main functions: to
aid in the nutrition of articular cartilage by acting as a
transport medium for nutritional substances, such as glucose, and
to aid in the mechanical function of joints by lubricating the
articulating surfaces. Articular cartilage has no blood, nerve, or
lymphatic supply. Glucose for articular cartilage chondrocyte
energy is transported from the periarticular vasculature to the
cartilage by the synovial fluid. Synovial fluid contains lubricin
secreted by synovial cells. Synovial fluid is chiefly responsible
for so-called boundary-layer lubrication, which reduces friction
between opposing surfaces of cartilage. There is also some evidence
that synovial fluid helps regulate synovial cell growth. Synovial
fluid serves many functions including: reducing friction by
lubricating the joint; absorbing shocks; and supplying oxygen and
nutrients to, as well as removing carbon dioxide and metabolic
wastes from, the chondrocytes within articular cartilage.
[0026] Normal synovial fluid does not clot but may exhibit
thixotropy, the property of certain gels to become fluid when
exposed to shear forces such as shaking. On standing at room
temperature, normal synovial fluid may assume gelatin-like
appearance, characterized by higher viscosities. When shaken it
will assume a normal fluid nature. Many enzymes have been found in
the normal synovial fluid. Alkaline phosphatase, acid phosphatase,
lactic dehydrogenase, and other enzymes are present in detectable
quantities. Enzymes enter the synovial fluid directly from the
plasma or may be produced locally by the synovial membrane or
released by synovial fluid macrophages. Synovial fluid also
contains phagocytic cells that remove microbes and the debris that
results from normal wear and tear in the joint.
[0027] Some prior art devices utilize fluid impermeable layers at
the cartilage surface, the bone/cartilage interface, or both
locations, or have rigid articular cartilage regions resistant to
receiving fluid from the synovial space. These types of structures
serve as barriers that prevent the normal transfer of essential
elements from the synovial fluid, into and out of the cartilage
region. What is needed is a device capable of facilitating joint
fluid therapy to the chondrocytes within the defect. Joint fluid
therapy encompasses delivering, receiving, accumulating and
controlling the location of desirable factors or molecules present
in the synovial fluid while also delaying or preventing destructive
factors, such as digestive enzymes, from prematurely degrading the
matrix. These desirable factors or molecules can be those naturally
occurring within the synovial fluid or biologically active agents
administered into the synovial fluid.
SUMMARY OF THE INVENTION
[0028] This invention includes implantable biphasic devices for the
repair of tissues of a living being, especially, cartilaginous
tissue defects. In the embodiment of a biphasic device, the device
has a first region and a second region, each being specific for the
growth of a particular tissue type. In an embodiment useful for
repair of cartilage defects, the first region is specific for
cartilage tissue growth, and the second region is specific for bone
growth.
[0029] In one aspect of the invention, the device is an
electro-kinetic implant, in which at least a portion of the device
features two juxtaposed materials that form a malleable matrix,
where the first material presents a positively charged surface, and
the second material presents a negatively charged surface. As the
malleable matrix is deformed under the application of pressures,
such as may occur while implanted in a living being, an electrical
potential is produced as a result of interactions, and
interruptions, between the charged surfaces of the first and second
materials. In one embodiment, the malleable material will be
malleable while hydrated, though it may be rigid, or at least
capable of being handled without deformation, while in a dry state.
In another embodiment, the malleable material may exhibit an
elastic property, tending to return to its original shape after
having been deformed. The first material of the malleable materials
may be a particulate, especially a fibrous particulate, and the
second material of the malleable material may be a hydrogel, such
that the particulate is suspended within the hydrogel, and upon
deformation, the hydrogel and particulate move relative to each
other. The malleable material may be porous. The materials may be
ceramics, natural polymers, synthetic polymers, or combinations
thereof.
[0030] The charges in the charged surfaces may be the result of
exposure of the constituent materials to acidic or basic
environments, plasma gas, or a result of the attachment of charged
substances to the materials.
[0031] In one embodiment, the first and second materials of the
malleable material are collagen, with the first collagen material,
such as a fibrous collagen, presenting a positively charged
surface, and the second collagen material, such as a hydrogel
presenting a negatively charged surface. In this embodiment, the
charged surfaces of the fibrous collagen and hydrogel collagen may
be created by exposing each of the collagens to solutions, where
one collagen is exposed to a solution having a pH above the
isoelectric point of the collagens, and the other collagen is
exposed to a solution having a pH below the isolectric point of the
other collagen.
[0032] Another aspect of the invention provides for the
transmission of forces and loads throughout a malleable matrix
component making up at least a portion of the implantable device.
In one embodiment, the malleable matrix component is created having
a first and second material, where the first material is a hydrogel
and the second material is an interconnected network of fibers. In
this embodiment, the hydrogel component may be collagen, or
hyaluronic acid, and the fibrous component may be collagen or
chitosan. The malleable matrix component is able to provide joint
fluid therapy to the cells or tissue within the implant device as
it is arranged to transmit forces throughout the entire, or at
least substantially the entire volume, of the malleable matrix
component, as forces applied will cause a vortex ring or gyre due
to the interactions of the interconnected fibers pulling on each
other, as they are displaced within the hydrogel material. It is
believed that the three-dimensional transmission of forces
throughout the malleable material will result in the malleable
material, or at least the hydrogel component of the malleable
material, receiving and accumulating desirable factors or molecules
from surrounding fluids, which may be utilized by cells within the
device.
[0033] In one embodiment, the malleable material is one phase of a
biphasic device, and corresponds to the cartilage region, thus the
malleable material may be attached to a rigid base corresponding to
the bone region.
[0034] In another aspect of the invention, the implant provides for
the systematic tissue conduction and growth from the surrounding
cartilage tissue, and retards the formation of tissue in the
interior of the implant. In this manner, it is believed that the
growth of the incorrect type of tissue can be avoided, and better
ensure that only the desirable hyaline cartilage is formed. In an
embodiment, the device may comprise a gradient, where the gradient
is arranged to retard the tissue formation most at or near the
center of the implant (when viewed top down), and transitions to
little or no retardation of tissue formation towards the perimeter
of the implant, adjacent to normal cartilage tissue. The gradient
may be in the form of a circular gradient, and may be uniform
throughout the device from upper surface to lower surface, or
alternatively may vary from top to bottom. The gradient may be a
smooth transition or gradual gradient, or alternatively a stepwise
gradient having well defined regions within the gradient. The
gradient may be a concentration gradient, such as biologically
active agents, additives, or combinations thereof. The gradient may
be a physical gradient, such as porosity, density, expansion,
swelling, elasticity, hardness, compressibility, and combinations
thereof. The gradient may be a material gradient, or chemical
gradient, such as molecular weight, cross-linking, hydrophobicity,
polarity, crystallinity, and combinations thereof. The gradient may
be part of the first phase of a multiphasic device, and corresponds
to the cartilage region, and may be attached to a rigid base
corresponding to the bone region.
[0035] In another aspect of the invention, the multiphasic implant
provides for the transmission or conduction of pressure forces
through the device, down to the underlying bone tissue below the
device; in this manner, bone tissue loss below the device, such as
may occur due to stress-shielding, may be minimized or avoided. One
embodiment of an implant device capable of transmitting such forces
would present a bone region presenting a porous material and a
rigid penetrating force conductive material capable of transmitting
the forces received from a malleable cartilage region to the
underlying tissue. The forces to be transmitted may be hydrostatic
and directed through channels running through the bone region
material, or alternatively force transmission may be in the form of
kinetic pressure pulses through the rigid conductive material
arranged in the bone phase. The rigid conductive material may be in
the form of columns arranged perpendicular to the top and bottom
surfaces of the implant, and may flare out to a wider dimension at
the junction with underlying bone. The rigid conductive material
may be in the form of a rigid multi-facetted web structure oriented
perpendicular to the top and bottom surfaces of the implant. In
another embodiment, the rigid conductive material is a wedge or
cone that transmits the forces through the implant to the
underlying bone, but may also transmit some forces laterally as an
outward force to the porous bone region material.
[0036] In yet another embodiment the multiphasic device capable of
transmitting pressure forces presents at least a first material in
the form of at least two porous rigid scaffolds, where the first
material is separated by at least a second material in the form of
a malleable elastic hydrogel, and where the hydrogel is capable of
transferring hydrostatic pressure pulses through the bone region of
the device in order to prevent bone voids from forming in external
underlying bone tissue.
[0037] The various embodiments described herein may be at least
partially or completely resorbed by the living being. The various
embodiments described herein may also feature drugs, biologically
active agents, or other additives in all or at least a portion of
the device.
[0038] Various medical uses of the above-described invention are
described below. Other features or advantages of the present
invention will be apparent from the following drawings and detailed
description of the invention, as well as from the claims.
DESCRIPTION OF THE DRAWINGS
[0039] FIG. 1 is a perspective depiction of a circular
gradient.
[0040] FIG. 2 is a cross-section depiction of the circular gradient
of FIG. 1.
[0041] FIG. 3 is a perspective depiction of a circular gradient
having a tapered construction from upper surface to lower
surface.
[0042] FIG. 4 is a perspective depiction of a circular gradient
having an hour-glass shape, wherein the gradient zones are wider in
the upper and lower surfaces, and featuring a narrow
mid-section.
[0043] FIGS. 5 and 6 are perspective depictions of multiple
circular gradients within the same device,
[0044] FIG. 7 is a cross-sectional depiction of a biphasic device
as found in the prior art, having a cartilage region comprising a
gel or porous material.
[0045] FIG. 8 is a cross-sectional depiction of a biphasic device,
having a cartilage region arranged as a web or matrix, where the
web or matrix is able to telegraph applied forces through
substantially all of the cartilage region, by the movement of the
web or matrix constituents in a manner analogous to a vortex ring,
or gyres.
[0046] FIG. 9 is a 1-year histology slide of a repair site having
had a prior art biphasic implant device implanted, after the device
has been completely absorbed, wherein stress shielding is
evident.
[0047] FIG. 10 is a cross-sectional depiction of an implant
embodiment that is arranged to transmit forces or loads through the
device to underlying tissue below, using a rigid central
column.
[0048] FIGS. 11a and b are cross-sectional depictions of another
implant embodiment arranged to transmit forces or loads through the
device using a stiff, multifaceted web structure and filler porous
material.
[0049] FIG. 12 is a cross-sectional depiction of an implant
embodiment that is arranged to transmit forces or loads through the
device using a rigid central conically shaped center post.
[0050] FIG. 13 is a depiction of a multi-layered cylinder
comprising various material thicknesses and densities, where the
layers can serve to transmit forces or loads through the device to
underlying tissue below.
[0051] FIGS. 14a and b are depictions of a multi-layered cylinder
material having swellable properties upon hydration, and capable of
transmitting forces hydraulically through the device to the
underlying tissue below.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0052] A device and methods are disclosed for treating tissue
deficiencies, defects, voids and conformational discontinuities
produced by congenital deformities, tissue pathology, traumatic
injuries and surgical procedures, particularly those located in
mammalian bone and cartilage. In one embodiment, the device is to
provide the means by which hyaline cartilage tissue can be
conducted across a tissue specific first scaffold region by
controlled migration of chondrocytes and/or cartilage precursor
cells. Additionally, in an embodiment, the scaffold region can be
designed to affect the concentration, location and activity of
fluids, factors, molecules or other biologically active agents
received from, or delivered to, the extracellular fluids,
especially synovial fluid. Thus, the device provides means to
regenerate a first specific form of tissue.
[0053] A tissue specific second scaffold region may be attached to
the first region for controlled migration of osteoblasts and/or
bone precursor cells. Thus described, an embodiment of the device
is a biphasic device, wherein the device consists of two main
parts, the cartilage region, and the subchondral bone region, which
are joined at an interface surface. Additionally, an embodiment
provides a means for deep bone mechanical stimulus by conduction of
mechanical and/or fluid forces originating in, or being applied to
the cartilage specific scaffold region. These stimuli will be
conducted through the subchondral bone region into the adjacent
uninvolved subchondral bone.
[0054] In a bi-phasic embodiment, the cartilage region can be
joined or bound to the subchondral bone region of the device by a
number of processes, including but not limited to, heat fusion,
heat welding, adhesives, glues or solvent welding. The resulting
union between the two architectural regions is preferably very
strong and can withstand any handling required to package the
device as well as any forces delivered to it as a result of the
implantation technique without permanently distorting the device's
internal architecture of void spaces.
[0055] The interface surface between the two regions may be a
permanent or temporary barrier to the passage of cells, fluids, or
biological components (e.g. growth factors, proteins, cells
signals, etc.) so long as it does not interfere with the
transmission of mechanical stimuli resulting from compression of
the first region.
[0056] In the biphasic device embodiment, the ingrowth or formation
of tissue would be specific to the device region, that is,
cartilage tissue would grow into the cartilage region of the
device, and bone tissue would grow into the bone region of the
device based upon the cells within the immediately adjacent
tissues, as well as mechanical and chemical signals provided by the
individual layers of the device. Furthermore, each of the cartilage
regions and bone regions may provide for a physical structure that
is appropriate to the type of tissue for which it is providing a
substrate. That is, the bone region will provide a stable
substratum for attachment of bone or bone forming cells, while the
cartilage region will provide a malleable elastic substratum
capable of allowing the surrounding uninvolved tissue to mediate,
or affect, the compression and motion of the scaffold adjacent to
the host tissue. Additionally, additives capable of enhancing the
growth of the target tissue are contemplated within the current
invention. Additives in the bone region can include ceramics,
glass, glass-ceramics, bioactive glasses, as well as biologically
active agents. Additives in the cartilage region can include
gelatinous materials, as well as biologically active agents. The
additives may be initially loaded into the cartilage region for
interaction internally within the device and/or for external device
delivery. Additionally, additives can originate within the synovial
fluid and be passively or actively transported into the cartilage
region of the device. Non-limiting examples of materials and
additives useful in construction of the various embodiments of the
devices described herein can be found in Table 2.
[0057] The architecture of each device region may be formed
utilizing established techniques widely practiced by those skilled
in the art of medical grade polymers. These methods may include
injection molding, extrusion and machining, vacuum foaming,
precipitation, sintering, spinning hollow filaments, solvent
evaporation, soluble particulate leaching or combinations thereof.
For some methods, plasticizers may be required to reduce the glass
transition temperature to low enough levels so that polymer flow
will occur without decomposition. Additionally additives such as
plasticizers or particulates can be added to polymers to make them
more or less malleable (malleable materials can be elastic as
defined earlier or plastic wherein they do not return to there
original shape after deformation) in order to provide the desired
mechanical properties for the specific device region they will be
located in. For example, a normally rigid polymer may incorporate a
plasticizer to make it malleable and thus useful in the cartilage
region whereas rigid particles could be added to a malleable
polymer to provide a stable substratum suitable for use in the bone
region.
[0058] In an embodiment, the osteochondral repair device will be
formed as a plug, typically circular in cross-section, and shaped
to fill a void or defect created through the cartilage layer and
into the underlying bone. Additionally, it is recognized that the
plug may have a tapered form, such that one end of the device is
larger than the other. A defect suitable for accepting the device
can be created in a manner known to those skilled in the art, for
example, using the device as described in U.S. application Ser. No.
11/049,410, or alternatively using defect creation techniques known
as the OATS procedure. It is recognized that alternative shapes
other than cylinders, may be utilized, for example joining or
overlapping circular elements together into one larger shape will
allow for larger defect areas to be repaired with coring tool
devices suitable for smaller defects (e.g., approximating an oval,
figure eight or a cloverleaf shape). Additionally, non-circular
shapes may be utilized as well, such as by providing plug devices
with alternative cross-sections, for example, polygonal shapes may
be employed or combined (e.g. rectilinear, triangular, hexagonal,
etc.), as the polygons may be joined alongside other devices to
form a mosaic covering a larger area than could a single
device.
[0059] Once there is a void created in the bone to accept the
implant device (e.g., core created by a coring tool), the implant
device is prepared for implantation. The implantable device may be
directed into the void through arthroscopic means, or alternatively
by hand into the exposed bone void. Preferably, the device is
loaded into an insertion tool. Though any known insertion tool or
mechanism may be employed, it is envisioned that the delivery can
be accomplished with an insertion tool including a
device-containing barrel with a delivery end, and also a plunger
extending into the barrel for ejecting the device out the delivery
end, in a manner similar to a wide mouth syringe. The insertion
tool is then placed adjacent to the opening, or directed into the
opening, and the device is then ejected from the delivery tool,
into the bone void. Preferably, care is taken, both in the creation
of the void, and in the delivery of the device, to avoid damaging
the healthy nearby tissue, particularly the cartilage tissue and
chondrocytes.
[0060] Once cellular tissue is fully established within the defect
repair site, it is expected that normal loads will be fully
supported by the new tissue. For biodegradable devices, the device
degrades and is eventually resorbed or removed from the
implantation site. This occurs as the device is degraded and
provides for the complete transfer of load bearing ability from the
device to the ingrown tissue, prior to the device's load bearing
ability falling below the levels required to aid in tissue
incorporation. Within this document, biodegradable, degradable,
bioresorbable, resorbable, bioerodable and erodeble may be used
interchangeably.
[0061] The various embodiments of a tissue repair device as
described herein may be implanted dry, or hydrated with
biologically relevant fluids, for example, saline, blood, bone
marrow aspirate or Platelet Rich Plasma (PRP). Also, growth
factors, hormones, drugs, cells or other useful biologically active
agents, can be used to hydrate the device. These materials can
provide therapy to the cells migrating into the implant, the
surrounding tissue, or the synovial fluid. Optionally, growth
factors, hormones, drugs, cells or other useful biologically active
agents can be located within the synovial fluid and adsorbed into
the implant by passive or active means. For reference, a
non-exhaustive list of biologically active agents that may be
incorporated into at least a portion or the entirety of the various
embodiments contained herein can be found in Table 1.
[0062] In healthy osteochondral tissues, for example a knee, having
a vertical axis that is in the load bearing direction, and a
horizontal axis that is normal to both the tissue surface and the
load bearing direction, typically, the encountered loads due to
natural movement and gravity are able to be transmitted or
conducted through the soft tissues of the joint, and into the hard
bony tissues. The load transmission is largely vertical, being in
the direction of load application, and creates compression of the
soft tissue, however, due to the interconnectivity of the soft
tissues, particularly across the transverse layer of the articular
cartilage, some portion of the loads are distributed laterally as
well, to adjoining soft tissue. One effect of this lateral
distribution is that a force of a given magnitude, having been
applied at only a small area at the top of the soft tissue, and
being transmitted through the soft tissue, would result in the
force being distributed over a wider area at the bottom of the soft
tissue, and into the bone. Given the wider distribution of the
force over a larger area, a compressive force in only a small area
of the articulating surface can provide deep bone mechanical
stimulus to a large area of subchondral bone, with the peak force
felt directly below the originating compressive force and lesser
amounts of conductive stimulus radiating outward.
[0063] In a similar fashion, where there has been a defect in
osteochondral tissue, and a plug device is implanted, the loads
that would have normally been transmitted by healthy tissue, would
now desirously be transmitted by the plug device as well.
Consequently, not only should a device that is inserted into a
defect beneficially be able to withstand the expected loads in the
defect location, both in the direction of the initial force
application, and also laterally as the force is distributed through
the soft tissue, but should also be able to adequately transmit or
conduct those forces through the device and into healthy adjacent
tissues.
[0064] Where the device is bioresorbable and also supports the
growth of new tissue, it is beneficial to ensure that the
degradation characteristics of the device are such that new tissue
ingrowth is structurally competent, meaning that it is able to
support the expected loads in the defect area, at least
coincidentally, or prior, to the degradation of that portion of the
device being subsumed by the new tissue ingrowth. In this manner,
the device can avoid the dimpling failure mode seen in prior art
devices, as a portion of the device becomes structurally
incompetent, the newly grown and structurally competent tissues can
provide the required weight bearing ability as well as the ability
to transmit mechanical stimulus.
[0065] One embodiment is intended to address the previously
described dimpling failure modes, where, it is believed, a portion
of the repaired defect area collapses prior to the growth of
structurally competent tissue. It is believed that the collapse
manifested as dimpling at the surface of the repair site, is a
result of failure in either, or both of, the remaining structures
of the implanted device, or in the new tissue ingrowth replacing
the device as it degrades. This embodiment alleviates this
occurrence by providing for a resorbable implant structure that
fosters satisfactory and controlled tissue ingrowth, and provides
for the last invaded and absorbed portion of the device to be
degraded after the tissue growth in the device is able to withstand
and transmit the encountered loads, also termed "structurally
competent". This may generally be achieved in one of two broad
manners. One may ensure that the device has adequate structural
competence for a period of time that is long enough to allow
adequate tissue restoration prior to the device becoming
structurally incompetent. Alternatively, one may accelerate the
radial ingrowth of competent tissue into the device, such that
cells are significantly established and forming the morphologically
correct tissue, thereby creating structurally competent tissue in a
shorter time frame, prior to the device losing its structural
competence. There are various techniques that may be employed for
achieving each of these goals, such as controlling porosity,
density, cross-linking, drug delivery, cell seeding, etc. These
techniques will be discussed later. It is recognized that one or
more techniques may be combined into a single device, to create an
ideal solution.
[0066] With reference to the following figures, applicants will
describe various embodiments for presenting a tissue repair device.
FIGS. 1-6 depict various gradient formations that may be employed
within a cartilage region of a biphasic device, that could allow
for competent tissue growth to be achieved as the device is
degraded and ultimately absorbed, thereby avoiding a mechanical
failure of the device caused by collapse or dimpling of the central
portion of the newly established tissue within the cartilage
region.
[0067] In one embodiment, and with reference to FIG. 1, a device is
provided having at least one controlled gradient in the device that
is arranged concentrically around a vertical axis and normal to the
cartilage surface. This circular gradient may provide, for example,
for a longer duration of implant structural competence as tissue
grows in concentrically (or in the form of accelerated tissue
regeneration from the outer zone) and spreads to the inner zone
depicted at the center of the device and whereas the zones depicted
in the outer portion of the device allows for rapid cell invasion
and the inner central zones retard cell invasion, or extracellular
matrix deposition, until such a time as the cells in the outer
zones have laid down the appropriate extracellular matrix,
influenced by the mechanical reaction to loading of the uninvolved
adjacent cartilage, in the form of hyaline cartilage. It is
important to prevent the occurrence of tissue formation as isolated
islands, which are not in contact with the uninvolved normal
articular cartilage, as isolated islands will not receive
appropriate mechanical stimulus from the surrounding uninvolved
tissue.
[0068] The controlled circular gradient in the device of FIG. 1 is
termed a "bull's-eye" gradient. The "bull's-eye" gradient refers to
the way the device appears when viewed from the cross-sectional
direction depicted in FIG. 2. As can be seen, the bull's-eye
gradient consists of a central core region or zone 300, surrounded
by one or more annular rings. In this depiction, there are two
annular rings, 100 and 200, concentrically arranged about the
central core. Though it is recognized that more or less annular
rings could easily be achieved as well. The controlled gradient
depicted in FIG. 1 with the cross section as shown in FIG. 2 is
uniform throughout the length of the device. It is also
contemplated that the gradient could vary along the vertical axis,
for example, differing in dimension to provide non-uniform
cross-sections throughout the length of the device, as can be seen,
for example in FIG. 3 and FIG. 4. In FIG. 3, the gradient has two
annular rings 110, and 210, surrounding a core region 310. In FIG.
4, the gradient has two annular rings 120, and 220, surrounding a
core region 320. It is recognized that the gradients depicted by
the figures may exist within a separate structural element in the
form of a cylinder or disk
[0069] Gradients can fall into many different groups including but
not limited to concentration, chemical, physical and material. The
invention can be provided in a great variety of useful shaped
devices, as will be discussed later, where the gradients of the
invention may be created by varying one or more of a variety of
characteristics, including porosity, density, molecular weight,
cross-linking, hydrophobicity, hydrophilicity, polarity, drug
concentration, drug delivery, material, expansion, swelling,
elasticity, hardness, compressability, crystallinity, cell seeding,
etc. To provide further clarity, select characteristics will be
explored more fully below, with reference to FIG. 1, as the
simplest embodiment, however, it is recognized that other shapes or
gradient forms for practicing the invention could employ similar
characteristic or compositional gradients.
[0070] Controlling the density of specific regions of the device
may be useful to provide greater structural resistance to
compressive loads. In an embodiment, a gradient can be constructed
where the center of the device has a higher density then the outer
edge. The density change may be achieved, for example, by varying
any of the porosity, pore size or pore number in each region of the
device, or by varying the molecular weight of the polymer in
various zones. For the example of a bull's-eye gradient, as
depicted in FIG. 1, the device may provide higher density polymers
or less porous scaffolding at the center zone 300, and then further
removed from the center to the perimeter on the cross-sectional
plain of the device, the material becomes less dense and more
porous. This embodiment with high porosity at the outer zone 100,
allows for the cells to migrate quicker initially at the outer zone
100, but retards their ability to reach the central zone 200 and
inner zone 300, thus concentrating the cells in the outer zone 100.
Central zone 200 will have a porosity or density in between that of
the interior and exterior of the device. This will also extend the
duration of structural competence at the core, as the mechanical
strength of the core is elevated due to the increased density, and
can thus be tailored to stay above a minimum threshold value (as
determined by the expected physical loading in the defect area) for
a longer period of time, as the device goes through biological
degradation. The increased duration of structural competence at the
center zone 300 allows more time for tissue to infiltrate, grow,
and become structurally competent in the core of the device, prior
to the total degradation or structural collapse. Those skilled in
the art will recognize other types of gradients that can be used to
decelerate cellular migration, as will be discussed.
[0071] An embodiment of the device may provide for a gradient by
using biologically active agents (e.g., drugs, cells, growth
factors, etc), ceramics, glass, metals or polymers, all of which
are included in the term "additives" incorporated into the device.
In this embodiment, the outer zone 100 of the device may provide an
elevated additive concentration, relative to the additive
concentration provided at the central zone 300 of the device. For
the specific case of growth factors, or other agents that will
enhance cellular chemotaxis and growth, this high concentration in
the outer zone 100 will help recruit cells to the outer edge of the
device faster and can increase tissue regeneration at the exterior
of the device, resulting in a shorter time period to reach
structural competence as the new tissue continues to grow into the
middle zone 200, and then into the core zone 300.
[0072] Controlling the rate of cross-linking of the polymer in
specific regions of the device may be useful to provide greater
structural resistance to compressive loads. In an embodiment, a
gradient can be constructed where the innermost zone 300 of the
device has a higher percentage of cross-linked polymers than the
outer zone 100 with middle zone 200 having a percentage of
cross-linked polymer somewhere in-between. As a result of the
cross-linking, the polymer will be more stable under loads, and
less subject to biodegradation and bioresorption, resulting in a
longer duration of structural competence in the more extensively
cross-linked regions, relative to the lesser cross-linked regions
of the device. This increased resistance to compressive loads will
protect any cells prematurely gaining access to the core portion of
the cartilage region from receiving incorrect mechanical signals
prior to being influenced by the encroaching tissue. Cells
receiving little to no mechanical stimulus will either attempt to
move down the bone lineage line (i.e., differentiate), or if
isolated from high oxygen content as naturally occurs in the
articular cartilage, will remain relatively dormant while waiting
for mechanical or chemical stimulus. In this way the innermost,
more cross-linked region will not inadvertently allow cells to
commit to the bone or fibrocartilage line, but instead cause the
cells to wait to be influenced by the mechanical properties of the
tissue being conducted through the matrix from the outer zones as
the matrix degrades and becomes softer. With reference to FIG. 1,
the core zone 300 may be a highly cross-linked polymer, and
transition to outer zone 100 that is not cross-linked at all, or
features less cross-linking. As stated previously, mechanical
signal transduction is critical to differentiation of the newly
forming tissue, any device having a cartilage scaffold matrix
greater in stiffness than the surrounding host tissue will not be
influenced by mechanical signal transduction and will either form
calcified tissues or disorganized fibrocartilage that collapses as
the matrix degrades and the tissue is stress loaded. Thus it is
important to initially concentrate the tissue forming cells in the
outer zones where they can be influenced by the surrounding
uninvolved tissue while at the same time preventing premature
collapse of the central zone. The device as described herein is
intended to set up the best circumstances to allow for the
formation of the correct tissue type.
[0073] Controlling the compositional makeup of specific regions of
the device may be useful to provide regions with longer durations
of structural competence. In an embodiment, a gradient can be
constructed by controlling the polymer blend ratio in each of the
zones to provide varying mechanical strength, or degradation rates.
For example, the innermost zone of the device may be manufactured
with a polymer or a blend of polymers that provides enhanced
resistance to degradation, or increased mechanical strength, when
compared to the polymer, or blend of polymers provided in the outer
zone of the device. In this embodiment, the center core of the
embodiment will feature an enhanced duration of structural
competence relative to the outer zone of the device
[0074] As a specific non-limiting example, and with reference again
to FIG. 1, natural polymers such as collagen may be used to create
regions with varying durations of structural competence. The outer
zone 100 of the device can be constructed from soluble collagen
that posses no fibers and is gelatinous by nature. This allows for
more rapid cellular tissue in-growth to the outer region of the
device as the collagen has a low compressive modulus and/or
degrades at a rapid rate allowing the newly recruited cells to be
stimulated by the mechanical forces necessary to lay down the
appropriate tissue matrix. The middle region 200 of the device may
be constructed from fibrillar collagen. Being of a higher
hierarchical structure the fibrillar collagen provides greater
structural integrity and/or greater resistance to degradation. In
the core zone 300 the collagen may be fibrous, thereby providing
even greater mechanical properties and/or greater resistance to
degradation than either of the outer zones. Thus, using the
hierarchical structure of collagen, a gradient can run through the
spectrum of gelatin, soluble collagen, fibrillar collagen, fibrous
collagen and collagen in the form of decellularized tissue, with or
without its extracellular matrix components, some or all of which
can be cross-linked as a tool for further control. Additionally,
the gradient could be based on length and/or thickness and/or
density of fibrils or fibers. For instance a homogenous soluble
collagen disk may contain an additive such as collagen fibers with
the mass or density of said collagen fibers increasing as one
proceeds or travels from outer zone 100 towards inner core 300.
Collagen gradients, as well as other material gradients, may also
be the result of differing animal sources (bovine, porcine, equine,
etc), or use of genetically engineered collagen, for instance from
plant sources.
[0075] Regions with varying durations of structural competence may
also be achieved with different types or species of polymers from
natural or synthetic sources. As an example, outer zone 100 can be
made from hyaluronic acid, which is very easily degradable, while
the middle region 200 can be constructed from natural polymer that
is more resistant to degradation such as collagen. The inner core
300 may contain an even tougher polymer such as chitosan.
Non-limiting examples of materials and additives useful in
construction of devices described herein can be found in Table
2.
[0076] It is recognized that various embodiments of the device may
provide more than one gradient, examples of which are depicted in
FIGS. 5 and 6. In these multi-gradient embodiments, a pair of
gradients are created, a first bull's-eye gradient may extend from
its widest dimension at the upper surface, and as one travels down
the vertical axis, the bull's-eye of the first gradient is shown to
diminish in cross-section, ultimately contracting to a point where
the zones merge. The second gradient may extend from the lower
surface, and diminish in area as one travels up the vertical
axis.
[0077] Specifically with regard to the multiple gradient
embodiment, as depicted in FIG. 5, the first gradient has two
annular rings 130, and 230, surrounding a core region 330, and the
second gradient has two annular rings 140 and 240, surrounding a
core region 340. As can be seen, the second bull's-eye gradient has
its largest dimensional area at the lower surface, and as one
travels up the vertical axis, the bull's-eye gradient forms around
the cone formed by the first bull's-eye gradient described
previously. Thus the outer dimension of the second gradient is
shown to remain uniform, while the inner zone of the second
gradient forms as an annular ring surrounding the cone of the first
gradient. As one nears the upper surface of the device, the second
bull's-eye gradient regions merge into a narrow annulus, preferably
at or near the upper surface of the device.
[0078] Specifically with regard to the multiple gradient
embodiment, as depicted in FIG. 6, the first gradient has two
annular rings 150, and 250, surrounding a core region 350, and the
second gradient has two annular rings 160 and 260, surrounding a
core region 360. As can be seen, the second gradient is created as
an inverse to the first gradient, and has its largest dimension at
the lower surface, and as one travels up the vertical axis, the
bull's-eye of the second gradient is shown to reduce in
cross-section, ultimately reducing to a point where the zones
merge. In this embodiment, the first and second gradients are
composed of unrelated characteristics or materials, and the
presence of one gradient will not necessarily interfere with the
presence of the other, thus they can be seen to overlap or extend
into each other as depicted in FIG. 6. It should be recognized that
these gradients might exist within a device having the gross shape
in the form of a cylinder. For example, a plug device having a
uniform porosity with one of the gradients being a first
biologically active agent, and the other gradient being a second
biologically active agent. In another example, one of the gradients
may comprise a structural gradient (e.g., density, cross-linking,
etc.)
[0079] A potential application of this "reverse-cone" of FIG. 5, or
"inverse-cone" of FIG. 6 is that the gradients can be employed to
optimize the balance required between promoting rapid cell
regeneration and tissue competence, against the required need for
adequate mechanical competence of the device as well as regulating
the rate of device degradation, which is so important to the
success of the device. In this embodiment, it is believed that the
first gradient (upper) could preferably be a density gradient, such
as can be created by controlling the porosity, pore size, pore
density, or polymer molecular weight, and the second gradient
(lower) could preferably be an additive gradient (e.g., growth
factors, drugs, ceramics, cells, etc.)
[0080] Yet another "bull's-eye" design can have a narrow
mid-section creating an hourglass look, as depicted in FIG. 4. (It
should be noted that this depiction only represents the gradient,
such as mechanical integrity resulting from fiber incorporation,
and that the general matrix in which the gradient resides is not
pictured.) The cross-section in this area of the device gradient is
much smaller, relative to the upper and the lower regions of the
device gradient, but still maintains the "bull's-eye" pattern. In
such an embodiment, the mechanical integrity of the device is
maintained by the gradient depicted. Thus the gradient may not
always occupy the entirety of the device to be implanted. This
extra area not pictured can be filled, for example, with highly
porous polymers that aren't required to provide any structural
competence properties, and whose main objective would be to receive
host cells and thus promote more rapid tissue regeneration in the
external regions closest to the uninvolved host tissue. That is,
for the external regions where little or no structural competence
is required to be provided by the device allowing the uninvolved
adjacent host tissue to mechanically influence the region, it is
preferable to provide a material that maximizes the amount and
extent of cellular ingrowth into the exterior of the device, in
order to provide a foothold of structurally competent tissue within
the device as quickly as possible.
[0081] For controlled gradients generally, it is contemplated that
the gradient be formed by altering some material or property within
the device in a manner corresponding to the patterns depicted in
the figures. Starting from the innermost zone at the core and
transitioning through the intermediate zones out to the outer
region, the gradient would provide some characteristic that varies
as one moves further out from the center. For the sake of
simplicity and ease of visualization, much of the explanation in
this application only discusses the example of FIG. 1, however, it
is recognized that the teachings of this application also are
applicable to the other examples and figures contained in this
application as well.
[0082] As depicted in FIG. 1, the gradient may feature zones
delineated by the concentric annular rings that provides a
recognizable or detectable border or interface between each of the
differing zones presented by FIG. 1. Alternatively, it is
recognized that a continuous transitional gradient or gradual
circular gradient could be provided, where there is a gradual
change in the characteristic, from the core region and progressing
out to the outside circumference, and the rings depicted in FIG. 1
are merely representative of the direction of the transition.
[0083] It is envisioned that a device providing for the various
gradient characteristics described herein could be manufactured as
an intact device, using carefully controlled lyophilization
techniques for creating these gradients. Alternatively, a series of
components may be manufactured, each varying in a particular
characteristic. Subsequently, the components may be shaped to a
form, where each component will form one of the zones, and
thereafter be assembled into a final device. For example, and with
reference to FIG. 1, a core piece could be manufactured, and later
inserted into annular rings sized concentrically, where each of the
assembled components will create the gradient desired in the final
device. Alternatively, one component may be provided as a scaffold
material in the manufacture of the other components, thereby
forming a multi-zoned device providing a gradient characteristic.
An example of this manufacturing method would include the injection
molding of a central skeleton followed by the incorporation of
other less dense open-celled matrices whose densities progress from
the central structure outward towards the perimeter of the finished
device.
[0084] It is also envisioned that gradients could be made or
created by compressing a starting porous polymer matrix to collapse
or sacrifice pores and thus develop a device having the various
zones as previously described. In addition, these gradients could
be developed by starting with granulated material, and then through
the use of heat and compression, could yield a finished device
containing varying porosities and physical shapes. For example,
fine granular material having an average diameter less than 50
microns can be placed in the center of a cylindrical mold creating
a central core. Around this can be pored a medium granular material
having an average diameter in the range of 50-100 microns creation
a middle zone. A course granular material having an average
diameter exceeding 100 microns in turn will surround this.
Compression and heat may then be used to fuse this granular
material together to create a bull's-eye gradient device.
[0085] It is also contemplated that that the cartilage region of
the current invention could be made to expand after implantation.
In this manner, the device would provide intimate contact with the
surrounding uninvolved cartilage tissue that has retracted away
from the defect hole, as the removal of a circular defect from
normal articular cartilage has been observed to result in
differential retraction of the edges. Depending on the depth of the
defect, the edges retract more in the superficial zone as compared
to the deeper zones after a circular defect is removed with a
punch. Normal human cartilage, with an intact superficial zone,
curls when removed from the underlying bone. The retraction away
from the defect site, as well as the curling of the removed
cartilage, is the result of the high tension existing within the
superficial zone of articular cartilage. This results in a cone or
funnel shape forming in the articular cartilage portion of a
surgically created defect, narrowing as one moves down towards the
subchondral bone portion of a surgically created defect. The
current invention anticipates this and thus can be capable of
radial expansion in order to ensure a tight fit. For example, a
cylindrical device can be place into a newly created defect and
expand until is has a shape as shown in FIG. 3.
[0086] Applicants have made an additional surprising discovery that
in effecting the repair of cartilage defects, prior art synthetic
implants and synthetic bi-phasic implant devices failed to
recognize the importance of synovial fluid in the maintenance and
repair of articular cartilage. As an additional consideration in
the development of a device for repair of articular cartilage one
needs to understand how friction, cyclic motion, electric potential
and synovial fluid all work together to maintain the articular
cartilage phenotype. Under normal physiological conditions,
articular cartilage provides a nearly frictionless surface between
moving joint. To help lubricate these joints, the body uses
synovial fluid. This fluid component consists primarily of water
with dissolved solutes and mobile ions.
[0087] Solute transport in biological tissues is a fundamental
process of life, providing nutrients to cells and carrying away
waste products. In avascular tissues such as adult articular
cartilage, solute transport occurs primarily across the articular
surface, with synovial fluid mediating exchanges with the synovium
lining the joint capsule. A primary mechanism of solute transport
is through diffusion. The mechanism of passive diffusion in healthy
cartilage has been shown experimentally to be enhanced by cyclical
loading of the cartilage, and by electro-osmotic flow both, of
which mechanisms lead to convective flow within the tissue. Other
avascular tissue types that respond similarly to articular
cartilage include tendon, ligament, meniscus and annulus thus the
techniques described herein for use in cartilage repair by
manipulating the natural fluid and electric potential in the region
may be used on these other tissue types as well. It is also
envisioned that these techniques could be beneficial on
vascularized tissue that are elastic in nature, including but not
limited to blood vessels and skin.
[0088] Within cartilage, it is recognized that the synovial fluid
acts as a transport medium for substances into and out of the
articular cartilage region. This is necessary because the articular
cartilage region is a non-vascular tissue. Substances are
transported into and out of the articular cartilage region due to
repetitive mechanical stimulus followed by a period of rest. During
active mechanical stimuli, molecules located within the synovial
fluid are actively transported into the articular cartilage layer.
This allows the concentration of molecules within the cartilage
tissue to exceed that of the synovial fluid. During rest, the
concentration will return to equilibrium. In this way, necessary
substances located within the synovial fluid are forced into the
cartilage tissue, whereupon the cells can absorb them. Waste
products are excreted by the cells into the interstitial space of
the tissue where they build up. During a period of rest, the system
moves towards equilibrium and thus the waste products move out of
the cartilage tissue and into the synovial fluid wherein they are
ultimately transported into the vasculature and away from the
knee.
[0089] Thus, vital nutrients are supplied to the non-vascular or
avascular tissues from the blood vessels located at the margins of
the tissue. The transport of nutrients through the dense complex
extracellular matrix to the cells making up these tissues relies
mainly on diffusion. Poor nutrient supply has been suggested as a
potential mechanism for degenerative processes that affect the
avascular tissues (i.e.--osteoporosis, disk degeneration, etc.) and
is also suspected in failure of prior art cartilage implants.
[0090] The effects of dynamic compression on chondrocyte
biosynthesis have been well characterized in cartilage explants and
chondrocyte-seeded scaffolds. In explants, continuously applied
dynamic compression and dynamic tissue shear have been found to
increase synthesis of proteins and proteoglycans.
[0091] Studies of articular cartilage metabolism have demonstrated
that static loading, as well as loading below a characteristic
frequency of 0.001 Hz, leads to biosynthetic inhibition, whereas
dynamic loading stimulates tissue synthesis. This enhanced
biosynthetic response is believed to result from an enhanced
nutritional supply, as well as a tissue biosynthetic response under
dynamic loading, and thus resulting in enhanced fluid flow and
changes in cell shape or mechanotransduction. Static compression of
articular cartilage has been shown to reduce the diffusivity of
various solutes within the tissue, and has been implicated in the
altered biosynthetic response of the tissue to static loading.
Growth factors, which have been shown to regulate the biosynthetic
response of articular cartilage, are generally large solutes with
molecular weights on the order of tens of kilodaltons. A further
benefit of dynamic loading is growth-factor uptake. It has been
shown that dynamic compression accelerates the biosynthetic
response of cartilage to free IGF-I and increases the rate of
transport of free IGF-I into the cartilage matrix, suggesting that
cyclic compression may improve the access of soluble growth
factors.
[0092] Dynamic compression, thus, augments the transport of solutes
in avascular tissues such as cartilage. However, the effect of
mechanical compression on the distribution and metabolism of
nutrients is difficult to directly evaluate. To this end, research
has been conducted on synthetic gels in order to answer these
questions.
[0093] Exposing an agarose gel, submerged in a fluid medium
containing target molecules, to repetitive mechanical compression
can crudely simulate the dynamic tissue compression system. It has
been observed that although the target molecules move against the
concentration gradient onto the gel, they are not evenly
distributed throughout the gel. The molecules only move into the
area under direct mechanical stimulus. If it was the case that
cartilage tissue behaved identically, then it would follow that
cells around the edges of the cartilage, would be deprived of
necessary substances. However, as will be discussed, cartilage does
not behave identically to agarose gel, though it does exhibit the
similar phenomenon of increasing the concentration of molecules as
a result of repeated compression. This unequal distribution of
necessary substances is a shortcoming of prior art devices having a
gel-like property within the cartilage region. Normal articular
cartilage overcomes this unequal distribution by having a dense
fibrous layer, known as the transverse layer that absorbs and
distributes mechanical stimulus across the entirety of the tissue
layer by providing a mechanical coupling of the cartilage molecules
to each other. In this way, necessary substances are actively moved
into the entirety of the cartilage tissue layer.
[0094] Similar to the normal cartilage tissue layer, a preferred
form of the current invention allows for uniform incorporation of
necessary target molecules by providing a biodegradable, insoluble
malleable elastic gel or hydrogel like substratum containing a
sufficient concentration of fibers so that they form a penetrating
interconnected phase. The gel or hydrogel can also present an
interconnecting porosity. The fibers, making up a second phase can
be entangled, entwined, interwoven, knitted or in some other
fashion connected in a three-dimensional web or matrix so that
stresses in the form of a push or pull are telegraphed throughout
the entire device. In this way the current invention is capable of
receiving joint fluid therapy throughout its entire volume.
[0095] FIG. 7 represents prior art biphasic device 700 having a
cartilage region of simple pores, or in the form of a gel. When
force 710 is applied to the surface of cartilage region 720,
pressure waves 730 remain focused just below original force 710.
Contrasted with the embodiment of FIG. 8, which depicts a
cross-section of biphasic device 800 wherein connected fibers 810
are located within layer 820. When force 830 is applied to the
surface of layer 820, downward pressure force 830 causes connected
fibers 810 to pull on each other, telegraphing pressure force 830
throughout the entire volume of layer 820, creating a circular
force, vortex, vortex ring, toroid, or gyre, as represented by
arrows 840. It should be noted that although shown in a single
plane, the described circular force occurs three-dimensionally
establishing a vortex ring, that is, multiple vortexes or gyres
within the device. The potential vorticity of fluid within the
device is directly related to the volume of displacement within the
device matrix from the downward pressure force. In the simplest
sense, vorticity is the tendency for elements of the fluid to
"spin." and can be related to the amount of "circulation" or
"rotation" in the fluid contained in the matrix caused by the
gyres. As new host tissue grows into the edges of this embodiment
of the device, forces applied to cartilage tissue distant from the
device will be transmitted through the host tissue and into the
device.
[0096] The cartilage layer of an embodiment of the device will be
composed of at least two phases. This first phase is an insoluble
gel or hydrogel capable of adsorbing and concentrating target
molecules from the synovial fluid when placed under repetitive
compressive forces. The second phase will be a fibrous component
associated with or contained within the gel phase having sufficient
connectivity so that a compressive force applied to one location of
the cartilage layer is transmitted throughout substantially the
entire volume of the cartilage layer. In order to achieve this the
minimum average fiber length for fibers randomly located within the
gel should be approximately equal to the thickness of articular
cartilage, which is from 2-3 millimeters in length. The maximum
average fiber length should not exceed 1.5 times the diameter of
the devices so as to prevent curling or coiling of the fibers
preventing them from being taut within the matrix and thus
dampening the transmission of mechanical stimulus. These same
length restrictions apply to interwoven or knitted type devices in
as much as connecting nodes or knots holding the structure together
should be no closer than 2-3 millimeters apart and no farther apart
than 1.5 times the diameter of the devices. For the example of a
plug implant device having a diameter of 10 mm, and a cartilage
region thickness of 3 mm, the average length of the fibers would be
in the range of 2-15 mm, and the average distance between
connecting nodes or knots would be in the range of 2-15 mm.
[0097] The material phases can be fabricated from natural and/or
synthetic polymers including but not limited to collagen, elastin,
keratin, chitosan, hyaluronic acid, silk, alginate, polyethylene
glycol (PEG) and combinations thereof. (Non-limiting examples of
materials and additives useful in construction of devices described
herein can be found in Table 2.) One or more of the phases can also
contain biologically active agents such as those listed in Table
1.
[0098] The biologic activities of the chondrocyte population are
regulated by genetic, and other biologic and biochemical factors,
as well as environmental factors. It has often been noted that
physical environmental factors, such as stress, fluid flow,
electric fields, etc. are as strong as biologic factors in
regulating cellular activities. There has been much research on the
effects of mechanical and/or hydrostatic/osmotic pressure loading
on cartilage explant metabolism. Such studies have been
specifically aimed at elucidating possible "mechano-signal"
transduction (also referred to as "mechanotransduction")
mechanism(s) that might govern the chondrocytes' biosynthetic
activities in maintaining and organizing the extracellular matrix
(ECM) comprising the tissue. Over decades many researchers have
observed electrical events in cartilage, but few studies have
focused on the details of the electrical potential within the ECM
where the chondrocytes reside. This phenomenon of electromechanical
or electrokinetic cell signaling has also be ignored by prior art
devices. Electromechanical or electrokinetic cell signaling is not
to be confused with mechanotransduction, as mechanotransduction
does not create electrical potential.
[0099] The electromechanical signals that chondrocytes perceive in
situ are the result of stresses, strains, pressures and the
electric fields generated inside the extracellular matrix when the
tissue is deformed. The potential induced by convection in the
presence of a pressure gradient in deformed tissue is the
"streaming potential". The potential induced by diffusion in the
presence of a concentration gradient in static tissue is the
"diffusion potential".
[0100] Avascular tissues such as cartilage are composed of water,
collagen enmeshed in a proteoglycan gel, and various matrix
proteins. The osmotic pressure of these tissues is mainly due to
the high density of charged carboxyl and sulfate groups on the
glycosaminoglycans of the proteoglycans within the tissues. When
avascular tissues are deformed under loading, interstitial fluid
flow occurs, even though the hydraulic permeability of the tissues
is very low. The electrical response of the tissues also changes
when it is compressed due to the effects of diffusion potential and
streaming potential.
[0101] The diffusion potential is the electric potential caused by
the separation between the bulk positive and bulk negative charges
caused by the gradients of mobile ions within the different fluid
regions of the tissue or between the tissue fluid and the synovial
fluid.
[0102] Streaming potential is defined as the difference in electric
potential between a diaphragm, capillary, or porous solid and a
liquid that is forced to flow through it. The definition of
streaming potential can also include the difference in electric
potential caused by the oscillation, separation or flow of a gel in
relationship to a diaphragm, capillary or porous solid.
Specifically, it is the potential that is produced when a liquid or
gel is forced to flow through a capillary or a porous solid. The
streaming potential is one of four related electrokinetic phenomena
that depend upon the presence of an electrical double layer at a
solid-liquid/gel interface. This electrical double layer is made up
of ions of one charge type that are fixed to the surface of the
solid and an equal number of mobile ions of the opposite charge
which are distributed through the neighboring region of the
liquid/gel phase. In such a system the movement of liquid/gel in
relation to the surface of the solid produces an electric current,
because the motion of the liquid/gel causes a displacement of the
mobile counterions with respect to the fixed charges on the solid
surface. The applied potential necessary to reduce the net flow of
electricity to zero is the streaming potential. Streaming potential
is related to zeta potential by factors that include the electrical
conductivity and fluid/gel viscosity. The value of streaming
potential under given conditions of conductivity and pressure can
be used to judge how strongly the tissue will interact with anionic
or cationic molecules. The zeta potential is a good predictor of
the magnitude of electrical repulsive force. A resulting voltage is
measured between electrode probes on either side of a boundary.
This voltage is then compared with the voltage at zero applied
pressure.
[0103] The source of electrical events, as measured on the outside
surface of normal articular cartilage, derives from the fixed,
immobilized or trapped negative charges .about.SO3 and COO2,
distributed along the chondroitin, keratin sulfates and hyaluronan
molecules comprising the aggrecan inside the extracellular matrix
of the tissue. These proteoglycans may be assumed to be
"immobilized and trapped" inside the extracellular matrix, and
therefore considered to be fixed to the extracellular matrix.
Together with the surrounding collagen network, these proteoglycan
macromolecules form the cohesive, strong, porous-permeable,
charged, collagen/proteoglycan solid matrix. By virtue of the
electro-neutrality law, there is always a cloud of counter-ions
(e.g., Ca, Na) and co-ions (e.g., Cl) dissolved in the interstitial
water surrounding the fixed charges in the extracellular matrix.
These ions may move by convection with the interstitial fluid due
to a hydraulic pressure or by diffusion through the fluid due to a
concentration gradient or by conduction, drifting through the fluid
as a current due to an electric field. Forces for the electric
current inside the tissues include the mechano-chemical force
generated by the gradient from movement of ions resulting from
compression and diffusion caused by ion concentration
gradients.
[0104] Within deformable tissues such as articular cartilage,
intervertebal disk, epiphyseal (growth) plate, and cornea, the
electric fields resulting from mechano-chemical forces are
constantly present. Thus, for such tissues, both streaming
potential and diffusion potential must always exist inside the
tissue and in fact they always compete against each other. The
streaming potential arises from the slight separation of the bulk
of the positive charges from that of the negative charges due to
the flow convection effects caused by a pressure gradient from
deformation of the tissue. The diffusion potential arises from the
slight separation of the bulk of positive charges from that of the
negative charges due to diffusion caused by the gradients of mobile
ions. It is believed that electrical events inside the tissue are
important in stimulating chondrocyte biosyntheses. It is also
believed that non-uniform electrical effects resulting from
deformation occurs when a tissue is softened during a disease
process such as osteoarthritis. In osteoarthritic cartilage, with
matrix degradation, the intrinsic compressive stiffness always
diminishes, thus affecting chondrocyte deformation and metabolic
activities as well as the nature of the mechano-electrochemical
events within cartilage when it is deformed.
[0105] Another preferred embodiment of the current invention
presents a cartilage region that takes into consideration both
diffusion potential and streaming potential in its constructions.
The cartilage layer of this preferred device will be composed of at
least two phases. This first phase is an insoluble gel or hydrogel
capable of adsorbing and concentrating target charged molecules
from the synovial fluid when placed under repetitive compressive
forces. The second phase will be a fibrous component contained
within the gel phase having sufficient connectivity so that a
compressive force applied to one location of the cartilage layer is
transmitted throughout the entire volume of the cartilage layer.
This allows creation of a disparity between the overall charges of
the synovial fluid from that of the cartilage layer establishing
the diffusion potential. In addition to this it is desirable for
the first phase to predominantly contain either positive or
negative charges while the second phase will predominantly contain
charges opposite that of the first phase. In this way a pressure
gradient from deformation of the cartilage layer of the preferred
embodiment creates a slight separation between the charges of the
first phase from that of the second phase, as the gel and fibers
flex, thus establishing the streaming potential. If desirable, one
or both phases can be cross-linked. Thus the electric potentials
created by such an embodiment simulate that which occurs in normal
articular cartilage, thus improving and/or stimulating chondrocyte
biosyntheses and thus articular cartilage tissue formation.
[0106] In one possible method for the manufacture of an embodiment
that takes into consideration both diffusion potential and
streaming potential, insoluble collagen fibers are exposed to a
more basic chemical environment (above the pH of the collagen's
isoelectric point) in order to bring the surface of the collagen
above its isoelectric point and thus providing a predominantly
negative charge to the surface of the fibers composing the second
phase of the devices. These negatively charged fibers are embedded
within a collagen gel or hydrogel that was exposed to a more acidic
chemical environment (below the pH of the collagen's isoelectric
point) so as to drive this collagen below its isoelectric point to
provide a predominantly positive charge to this first phase. This
is unlike prior art devices that contain two phases of collagen
wherein both collagens are on the same side of the isoelectric
point.
[0107] In another embodiment, biodegradable polyester fibers (ie
-PLA, PGA, PCL, etc), which have been subjected to surface
modifications, such as exposure to acids, bases, or plasma gas
processes) are used in the second phase of the device.
[0108] In another embodiment, hyaluronic acid gel or hydrogel
having a predominantly negative charge is used as the first phase
that encapsulates and surrounds a second phase of chitosan fibers
having an overall positive charge. When making combinations such as
hyaluronic acid and chitosan, care must be taken so that a
polyelectrolytic complex (PEC) is not formed as this will not allow
the charges to separate during compression and thus no electric
potential will occur.
[0109] In another embodiment, an electrically neutral hydrogel
first phase envelops a charged fibrous second phase, wherein the
gel allows mobile ions to penetrate and take up residence within
the gel thus balancing out the charge of the fiberous second phase.
As described previously, deformation of the combined matrix will
result in charge separation, creating the electric potential. An
example of an electrically neutral hydrogel would be a PEC. Such a
PEC could be manufactured by various techniques known in the art,
incorporating known components. The neutral hydrogel PEC could be
created by the combination of charged components, such as
hyaluronic acid--chitosan, collagen--chitosan, and hyaluronic
acid--collagen.
[0110] It is also recognized that the second phase material can be
composed of particulate materials that are not fibrous or polymeric
in nature so long as they provide the necessary charged surface. A
non-limiting list of materials suitable for this use can be found
in table 2.
[0111] Those skilled in that art will identify other combinations
of positively and negatively charged materials all of which are
embraced by this disclosure for use in creation of an
electro-kinetic tissue repair device.
[0112] As already discussed, in some embodiments, part of the
function of the device is to transfer forces or loads, experienced
by the cartilage layer, through the devices and into the
subchondral bone. This deep bone mechanical stimulus is necessary
to prevent stress shielding that currently results in bone voids
below the device. FIG. 9 shows 12-month histology from a prior art
device that provided stress shielding to the underlying bone. Box
910 shows the approximate location of the implant that has
completely resorbed. Soft tissue void 920 within the bone is the
result of this stress shielding.
[0113] Both cartilage and bone are living tissues that respond and
adapt to the loads they experience. If a joint surface remains
unloaded for appreciable periods of time the cartilage tends to
soften and weaken. Further, as with most materials that experience
structural loads, particularly cyclic structural loads, both bone
and cartilage begin to show signs of failure at loads that are
below their ultimate strength. Research into bone healing has shown
that some mechanical stimulation can enhance the healing response
and it is likely that the optimum regime for a cartilage/bone graft
or construct will involve different levels of loading over time in
order to properly repair a damaged region. This same observation
was concluded by Surgeon Julius Wolff back in the 19.sup.th century
and is still known today as Wolff's law.
[0114] Many prior art implants that are made for use in repairing
damaged bone and cartilage are fabricated from soft materials and
deform when they are implanted into a cored hole in the bone. These
implants do not provide a means for the transfer of loading through
the implant for stimulating the growth of new bone at the bottom or
side walls of the cored hole, or even controlling or preventing
osteopenia or osteoporosis. Other implants that are fabricated as
bone void fillers are made from rather stiff materials such as
ceramics. These devices can provide a means for mechanical
stimulation; however, the implant must be precision fitted to the
bone void in order to create the proper length to match up with the
hole that has been cored into the patient's bone. Since any
protrusion of these devices will result in higher contact pressure,
which may further damage the cartilage in joint areas, it is not
advisable to use these devices for cartilage repair.
[0115] For osteochondral transplantation involving the replacement
of damaged cartilage sites with harvested plugs taken from the
patients' joint, it is also difficult to match the cored hole depth
with the exact implant length. This is a function of the design of
the coring tool as well as the technique utilized by the surgeon.
For some coring tools, the cored hole will exhibit a very uniform
cylindrical shape, however, the bottom surface may be inconsistent
and have a rather jagged and irregular surface. This can create
gaps or void pockets under the implant or create a void between the
top of the implant and the mating rotating bone and prevent any
transfer of forces or pressure during the healing process. In
addition, the surgeon is concerned about protrusion of the
harvested plug creating too much pressure on the transplanted
hyaline cartilage thereby damaging this tissue as the joint moves.
Therefore, the surgeon often creates a deeper recipient site defect
then the length of the harvested plug. This allows the surgeon to
control the final position or height of the implanted device;
however, this is assuming that the frictional forces alone will
provide enough stability for the plug to stay in position. This
also creates a void space under the implant, which prevents contact
from occurring with the subchondral bone.
[0116] Other studies have shown that bottoming out the implant can
provide for better support and stability during the time that the
cells are growing into the newly implanted plug. However, bottoming
out the implant can cause high compressive forces during insertion,
which can also damage the transplanted cartilage during the
surgery. These same studies have also shown that these implanted
plugs are more stable and can be cut to shorter lengths if they are
bottomed out.
[0117] In order to obtain loading through the cartilage/bone region
of any device, contact and pressure are required to exist. As
previously discussed, it may not be possible to create a tight
enough fit between the implanted device and the cored hole in the
patient's bone. Therefore, the implanted device needs to provide
the capabilities to expand and contract to fill this space.
[0118] Based on these requirements, it is envisioned that a device
could be designed so that a portion of it has the ability to expand
and contract like an extension spring. Once the device is implanted
into a cored hole, the expansion and contraction of the implant
would provide the proper functionality. In addition, it is
desirable to also create sufficient contact with the walls of the
cored hole.
[0119] A cartilage/bone repair device is envisioned which takes
into consideration the transfer of structural loads or pressures
that may be seen by the implant once it is installed into a
cored-out hole in the recipient's bone.
[0120] In various embodiments, the implant may be made of different
materials or different forms of the same material. As an example, a
rigid support skeleton can be injection molded from a PLA polymer
and this same polymer can be chemically processed to create an
open-celled foam structure. Both of these materials would act in
completely different ways in regards to their absorption
characteristics, their load transfer characteristics, and their
biological cell attraction characteristics.
[0121] In other embodiments, the implant may include various means
of securing itself within the area of bone repair. These securing
means can include mechanical methods such as teeth or ridges that
are incorporated around the outside surfaces of the device. These
teeth or ridges can also assist with the transfer of forces through
the device and into the surrounding bone.
[0122] In further embodiments, the device could utilize different
characteristics formulated into the structural make up of the
device in order to promote the take up of fluid thereby causing a
hydraulic effect in a portion of the device, which would create a
means of expansion and thereby allow for pressure to be transferred
through the device.
[0123] In another embodiment, the device contains fluid swellable
expansion zones that provide for a tight fit within the void and
allow for micro-motion while other porous stable zones allow for
cell attachment and tissue growth.
[0124] Various methods can be utilized for transferring the forces
or loads through the device in order to provide mechanical
stimulation to the bone interfacing surfaces. As shown in FIG. 10,
device 1000 has center column 1010 positioned under cartilage layer
1020 that transcends down the center and then transitions to a
larger diameter at the bottom to allow the transfer of force or
pressure between the upper surface of the implant and the
implant/bone interface layer at the bottom of the device. Porous
matrix 1030 surrounds center column 1010 and makes contact with the
host tissue. Center column 1010 can be porous, but is rigid and
thus conductive of mechanical stimulus that would be dampened by
porous matrix 1030. It is preferable that porous matrix 1030 swells
shortly after placement into the tissue void so direct contact is
made with the tissue void walls. In addition transitioning to a
larger diameter at the bottom, center column 1010 can also
transition to a larger diameter at the top (not shown), presenting
an hourglass type of shape. Additionally, center column 1010 can be
formed from a small diameter cylinder with a thin flat plate on the
bottom and optionally the top (not shown). The porosity, if
present, in center column 1010 can be random, or in the form of
elongated channels capable of conducting hydraulic forces.
[0125] FIG. 11a shows device 1100 having multi-facetted web
structure 1110 that is oriented perpendicular to the top and bottom
surfaces of device 1100. In this configuration, the web acts as a
stiffener to transfer the load originating in cartilage layer 1120
through the implant. Secondary material 1130 is a less dense, more
porous structure formed in between the spokes of web structure
1110. FIG. 11b shows a top down view with the cartilage layer
removed so that the relationship of the spokes of web structure
1110 and secondary material 1130 can easily be visualized. In this
embodiment, the web would continue to transfer the forces into the
subchondral bone region while bone growth was occurring within the
porous structure of secondary material 1110 found in between the
webs or spokes of web structure 1110. As bone growth completed the
encroachment of this area, it would assist with the load or
pressure transfer while the materials of web structure 1110 started
its degradation and eventual removal. Web structure 1110 can have
holes or slots within its structure to allow intercommunication of
the secondary material 1130.
[0126] As shown in FIG. 12, device 1200 has conically shaped center
post 1210 sitting below cartilage layer 1220. Center post 1210
wedges into outer cylinder layer 1230 possessing a shaped inner
cavity designed to receive center post 1210. Center post 1210 may
extend completely through outer layer cylinder layer 1230 as
pictured or may instead just come flush to the base of device 1200.
The tapered shape of center post 1210 provides for a means of
seating the implant while also providing a method for transferring
mechanical stimulus to all sections of the subchondral bone region.
When downward force 1250 is applied to device 1200 outer cylinder
layer 1230 is experiences outward force 1240 thus providing
improved seating of device 1200 into a cored bone void. Thus forces
applied to cartilage layer 1220 pass into center post 1210 and are
transferred to the tissue void.
[0127] FIG. 13 is composed of a multi-layered cylinder containing
various material thicknesses and densities. The layers can be
constructed to act to transfer the pressure between the top surface
of the device and the bottom surface. The composition of these
various layers can also be utilized to create hydraulic swelling to
thereby create the spring-like effect previously described.
[0128] FIG. 14a shows a simplified example of a bone region
multi-layered cylinder 1400. To simplify understanding, the
cartilage layer is not pictured. Swellable layers 1420 separate
rigid porous layers 1410. More or less layers are also
contemplated. Upon implantation or exposure to liquid, swellable
layers 1420 imbibe fluid and become an uncompressible, flexible
hydrogel as depicted in FIG. 14b where rigid porous layers 1410 are
now separated by swollen layers 1430. Referring to FIG. 14b, a
force applied to the cartilage layer (not shown) is transferred as
a pressure wave through the device so long as rigid porous layers
1410 do not exceed 4 mm in thickness and have an average porosity
greater than 50 microns and are rigid enough to avoid collapse of
their porosity thus not dissipating the pressure wave prior to it
reaching the bottom layer and finally being conducted into the
underlying bone. Optionally one or more holes, 2 millimeters in
diameter or greater can exist in layers 1410 allowing pillars of
hydrogel to connect swollen layers 1430. It may be I that newly
forming bone needs a stable substratum to attach to so that bone
forming cells can lay down extracellular matrix. Bone forming cells
known as osteoblasts are approximately 50 microns in diameter and
should establish themselves in newly forming islands of bone
approximately 1 mm in diameter, thus the minimum thickness of
porous layer 1410 is 1 millimeter. The thickness of swollen layer
1430 has no maximum, but should be at a minimum of 5 microns with a
preferred thickness of 50 microns to trap a sufficient amount of
fluid and thus function as an incompressible hydrogel capable of
transferring pressure waves.
[0129] In another embodiment (not shown) porous particles having a
minimum approximate diameter of 1 millimeter can be surrounded by a
swellable material wherein the swellable material maintains
connectivity throughout the entire device. In this way, pressure
waves and micro motion, necessary for establishing bone external to
the device, can be conducted through the swellable material matrix
while the porous particles provide a stable platform for attachment
and proliferation of osteoblasts. As a non limiting example, porous
particles composed of ceramic, polymer or composites of the two can
be suspended within a hydrogel forming material such as collagen,
hyaluronic acid, chitosan, alginate, keratin, or PEG. In addition
to being a homogenous material, the hydrogel can be formed into a
porous network presenting fluid swollen struts or partitions
defining fluid containing pores.
[0130] The bone region of all the above devices can be designed so
that they provide the required expansion and transfer of force as
the materials degrade. This transfer of force can occur through the
use of rigid polymeric or ceramic elements, incompressible
hydrogels or combinations thereof. As more cells are stimulated to
grow into the implanted matrix, newly formed tissue will help to
continue the transfer of the mechanical stimulation.
[0131] The inclusion of groups and subgroups in the tables is
exemplary and for convenience only. The grouping does not indicate
a preferred use or limitation on use of any material therein. For
example, in Table 1, the groupings are for reference only and not
meant to be limiting in any way. Additionally, the tables are not
exhaustive, as many other drugs and drug groups are contemplated
for use in the current embodiments. There are naturally occurring
and synthesized forms of many therapies, both existing and under
development, and the table is meant to include both forms.
[0132] Numerous other embodiments and modifications will be
apparent to those skilled in the art and it will be appreciated
that the above description of a preferred embodiment is
illustrative only. It is not intended to limit the scope of the
embodiments contained herein, which are defined by the following
claims. Without further elaboration the foregoing will so fully
illustrate our invention that others may, by applying current or
future knowledge, adopt the same for use under various conditions
of service.
TABLE-US-00001 TABLE 1 Examples of Biologically Active Agents
Adenovirus with or without genetic material Angiogenic agents
Angiotensin Converting Enzyme Inhibitors (ACE inhibitors)
Angiotensin II antagonists Anti-angiogenic agents Antiarrhythmics
Anti-bacterial agents Antibiotics Erythromycin Penicillin
Anti-coagulants Heparin Anti-growth factors Anti-inflammatory
agents Dexamethasone Aspirin Hydrocortisone Antimicrobial
Antioxidants Anti-platelet agents Forskolin Anti-proliferation
agents Anti-rejection agents Rapamycin Anti-restenosis agents
Antisense Anti-thrombogenic agents Argatroban Hirudin GP IIb/IIIa
inhibitors Anti-virus drugs Arteriogenesis agents acidic fibroblast
growth factor (aFGF) angiogenin angiotropin basic fibroblast growth
factor (bFGF) Bone morphogenic proteins (BMP) epidermal growth
factor (EGF) fibrin granulocyte-macrophage colony stimulating
factor (GM-CSF) hepatocyte growth factor (HGF) HIF-1 insulin growth
factor-1 (IGF-1) interleukin-8 (IL-8) MAC-1 nicotinamide
platelet-derived endothelial cell growth factor (PD-ECGF)
platelet-derived growth factor (PDGF) transforming growth factors
alpha & beta (TGF-.alpha., TGF-beta.) tumor necrosis factor
alpha (TNF-.alpha.) vascular endothelial growth factor (VEGF)
vascular permeability factor (VPF) Bacteria Beta blocker Blood
clotting factor Bone morphogenic proteins (BMP) Calcium channel
blockers Carcinogens Cells Adipose cells Bone marrow cells Blood
cells Endothelial Cells Epithelial cells Skeletal muscle cells
Smooth muscle cells Stem Cells Umbilical cord cells Fat cells Bone
cells Mesenchymal stem cells Progenitor cells Cartilage cells
Cellular Material Bone marrow Cells with altered receptors or
binding sites Fibroblasts Genetically altered cells Glycoproteins
Growth factors Lipids Liposomes Macrophages Reticulocytes Vesicles
Chemotherapeutic agents (e.g. Ceramide, Taxol, Cisplatin)
Cholesterol reducers Chondroitin Collagen Inhibitors Colony
stimulating factors Coumadin Cytokines prostaglandins Dentin
Etretinate Genetic material Glucosamine Glycosaminoglycans GP
IIb/IIIa inhibitors L-703,081 Granulocyte-macrophage colony
stimulating factor (GM-CSF) Growth factor antagonists or inhibitors
Growth factors Acidic fibroblast growth factor (aFGF) Autologous
Growth Factors Basic fibroblast growth factor (bFGF) Bone
morphogenic proteins (BMPs) Bovine Derived Growth Factors Cartilage
Derived Growth Factors (CDF) Endothelial Cell Growth Factor (ECGF)
Epidermal growth factor (EGF) Fibroblast Growth Factors (FGF)
Hepatocyte growth factor (HGF) Insulin-like Growth Factors (e.g.
IGF-I) Nerve growth factor (NGF) Platelet Derived endothelial cell
growth factor (PD-ECGF) Platelet Derived Growth Factor (PDGF)
Recombinant NGF (rhNGF) Recombinant Growth Factors Tissue Derived
Cytokines Tissue necrosis factor (TNF) Transforming growth factors
alpha (TGF-alpha) Transforming growth factors beta (TGF-beta) Tumor
necrosis factor alpha (TNF-.alpha.) Vascular Endothelial Growth
Factor (VEGF) Vascular permeability factor (UPF) Growth hormones
Heparin sulfate proteoglycan HMC-CoA reductase inhibitors (statins)
Hormones Erythropoietin Immoxidal Immunosuppressant agents
inflammatory mediator Insulin Interleukins Interlukin-8 (IL-8)
Interlukins Lipid lowering agents Lipo-proteins Liposomes Lipids
Low-molecular weight heparin Lymphocites Lysine MAC-1 Morphogens
Nitric oxide (NO) Nucleotides n-methylpyrrolidinone (NMP) Dimethyl
Sulfoxide (DMSO) Peptides Phosphorylcholine Phospholipids
Polypeptides PR39 Proteins Prostaglandins Proteoglycans Perlecan
Radioactive materials Iodine - 125 Iodine - 131 Iridium - 192
Palladium 103 Radio-pharmaceuticals Secondary Messengers Ceramide
Somatomedins Statins Steroids Sulfonyl Thrombin Thrombin inhibitor
Thrombolytics Ticlid Tyrosine kinase Inhibitors ST638 AG-17
Vasodilator Histamine Forskolin Nitroglycerin Vitamins E C
Yeast
TABLE-US-00002 TABLE 2 Examples of Materials and Additives
Aliphatic polyesters Cellulose Bioglass Chitin Collagen Copolymers
of glycolide Copolymers of lactide Elastin Fibrin
Glycolide/l-lactide copolymers (PGA/PLLA) Glycolide/trimethylene
carbonate copolymers (PGA/TMC) Hydrogel
Lactide/tetramethylglycolide copolymers Lactide/trimethylene
carbonate copolymers Lactide/.epsilon.-caprolactone copolymers
Lactide/.sigma.-valerolactone copolymers L-lactide/dl-lactide
copolymers Methyl methacrylate-N-vinyl pyrrolidone copolymers
Modified proteins Nylon-2 Organic Solvents
PHBA/.gamma.-hydroxyvalerate copolymers (PHBA/HVA) PLA/polyethylene
oxide copolymers PLA-polyethylene oxide (PELA) Poly (amino acids)
Poly (trimethylene carbonates) Polyhydroxyalkanoate polymers (PHA)
Poly(alklyene oxalates) Poly(butylene diglycolate) Poly(hydroxy
butyrate) (PHB) Poly(n-vinyl pyrrolidone) Poly(ortho esters)
Polyalkyl-2-cyanoacrylates Polyanhydrides Polycyanoacrylates
Polydepsipeptides Polydihydropyrans Poly-dl-lactide (PDLLA)
Polyesteramides Polyesters of oxalic acid Polyglycolide (PGA)
Polyiminocarbonates Polylactides (PLA) Poly-l-lactide (PLLA)
Polyorthoesters Poly-p-dioxanone (PDO) Polypeptides
Polyphosphazenes Polysaccharides Polyurethanes (PU) Polyvinyl
alcohol (PVA) Poly-.beta.-hydroxypropionate (PHPA)
Poly-.beta.-hydroxybutyrate (PBA) Poly-.sigma.-valerolactone
Poly-.beta.-alkanoic acids Poly-.beta.-malic acid (PMLA)
Poly-.epsilon.-caprolactone (PCL) Pseudo-Poly(Amino Acids) Starch
Trimethylene carbonate (TMC) Tyrosine based polymers Alginate
Calcium Calcium Phosphate Calcium Sulfate Ceramics Chitosan
Cyanoacrylate Collagen Dacron Demineralized bone Elastin Keratin
Plasticizers Fibrin Gelatin Glass Gold Glycosaminoglycans
Hyaluronic acid Hydrogels Hydroxyapatite Hydroxyethyl methacrylate
Hyaluronic Acid Liposomes Lipids Nitinol Nanoparticles Osteoblasts
Oxidized regenerated cellulose Phosphate glasses Polyethylene
glycol Polyester Polysaccharides Polyvinyl alcohol Platelets, blood
cells Radiopacifiers Salts Silicone Silk Steel (e.g. Stainless
Steel) Synthetic polymers Thrombin Titanium Silica Clay Metals
Silver Aluminum Oxides Ceramics Polymers Metal Oxides Alkyl
methlacrylates Hydrophilic polymer Integrins Paralyne
Polyacrylamide Polyanhydrides Polyethylene acetate Polyethylene
glycol Polyethylene oxide Polyurethane Polyvinyl alcohol Polyvinyl
pyrrolidone Silanes Silicone
* * * * *