HYDROGEL ARTHROPLASTY DEVICE
An arthroplasty device is provided having an interpenetrating polymer network (IPN) hydrogel that is strain-hardened by swelling and adapted to be held in place in a joint by conforming to a bone geometry. The strain-hardened IPN hydrogel is based on two different networks: (1) a non-silicone network of preformed hydrophilic non-ionic telechelic macromonomers chemically cross-linked by polymerization of its end-groups, and (2) a non-silicone network of ionizable monomers. The second network was polymerized and chemically cross-linked in the presence of the first network and has formed physical cross-links with the first network. Within the IPN, the degree of chemical cross-linking in the second network is less than in the first network. An aqueous salt solution (neutral pH) is used to ionize and swell the second network. The swelling of the second network is constrained by the first network resulting in an increase in effective physical cross-links within the IPN.
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This application is a continuation of U.S. application Ser. No. 15/206,060, filed Jul. 8, 2016, which is a continuation of U.S. patent application Ser. No. 14/188,257, filed Feb. 24, 2014, now U.S. Pat. No. 9,387,082, which is a continuation of U.S. patent application Ser. No. 13/418,294, filed Mar. 12, 2012, now U.S. Pat. No. 8,679,190, which is a continuation of U.S. patent application Ser. No. 12/148,534, filed Apr. 17, 2008, now abandoned, which claims the benefit of U.S. Provisional Patent Application No. 60/923,988, filed Apr. 17, 2007, and is a continuation-in part of U.S. patent application Ser. No. 12/070,336, filed Feb. 15, 2008, now U.S. Pat. No. 8,821,583, the disclosures of all of which are incorporated herein by reference.
U.S. patent application Ser. No. 12/070,336 claims priority from U.S. Provisional Application No. 60/901,805, filed Feb. 16, 2007, and is a continuation-in part of U.S. patent application Ser. No. 11/243,952, filed Oct. 4, 2005, now U.S. Pat. No. 7,857,849, which claims the benefit of U.S. Provisional Application Nos. 60/616,262, filed Oct. 5, 2004 and 60/673,172, filed Apr. 20, 2005.
U.S. patent application Ser. No. 12/070,336 is also a continuation-in-part of U.S. application Ser. No. 11/636,114, filed Dec. 7, 2006, now U.S. Pat. No. 7,857,447, which claims the benefit of U.S. Provisional Application Nos. 60/843,942, filed Sep. 11, 2006, and 60/783,307, filed Mar. 17, 2006.
U.S. patent application Ser. No. 12/070,336 is also a continuation-in-part of U.S. application Ser. No. 11/409,218, filed Apr. 20, 2006, now abandoned, which claims the benefit of U.S. Provisional Application No. 60/673,600, filed Apr. 21, 2005.
U.S. patent application Ser. No. 12/070,336 is also a continuation-in-part of U.S. application Ser. No. 11/639,049, filed Dec. 13, 2006, now U.S. Pat. No. 7,909,867, which claims the benefit of U.S. Provisional Application No. 60/843,942, filed Sep. 11, 2006.
FIELDThe present invention relates generally to interpenetrating polymer network hydrogels. More particularly, the present invention relates to devices and materials useful for orthopaedic prostheses.
BACKGROUNDWith disease or damage, the normally smooth, lubricious cartilage covering joint surfaces progressively deteriorates, exposing bone and leading to arthritic pain that is exacerbated by activity and relieved by rest. Today, patients with osteoarthritis are faced with only one of two choices: either manage their pain medically, or undergo an effective but highly bone-sacrificing surgery. Medical management includes weight loss, physical therapy, and the use of analgesics and nonsteroidal anti-inflammatories. These can be effective at reducing pain but are not curative. Other options include drugs like glucosamine or hyaluronan to replace the “lost” components of cartilage, but despite their extensive use in the U.S., their efficacy is still questioned. When medical intervention fails and a patient's joint pain becomes unbearable, surgery is advised. Total joint arthroplasty is a surgical procedure in which the diseased parts of a joint are removed and replaced with new, artificial parts (collectively called the prosthesis). In this highly effective but invasive procedure, the affected articular cartilage and underlying subchondral bone are removed from the damaged joint. A variety of replacement systems have been developed, typically comprised of ultra-high molecular weight polyethylene (UHMWPE) and/or metals (e.g. titanium or cobalt chrome), or more recently, ceramics. Some are screwed into place; others are either cemented or treated in such a way that promotes bone ingrowth. These materials have been used successfully in total joint replacements, providing marked pain relief and functional improvement in patients with severe hip or knee osteoarthritis.
A large number of patients undergo total hip arthroplasty (THA) in the U.S. each year, which involves implanting an artificial cup in the acetabulum and a ball and stem on the femoral side. The goals of THA are to increase mobility, improve hip joint function, and relieve pain. Typically, a hip prosthesis lasts for at least 10-15 years before needing to be replaced. Yet despite its success as a surgical procedure, THA is still considered a treatment of last resort because it highly “bone-sacrificing,” requiring excision of the entire femoral head. It is this major alteration of the femur that often makes revision replacement difficult. While this procedure has a survival rate of 90% or more in the elderly (who usually do not outlive the implant), implant lifetimes are significantly shorter in younger, more active patients. As a result, younger patients face the prospect of multiple, difficult revisions in their lifetime. Revisions are required when implants exhibit excessive wear and periprosthetic bone resorption due to wear particles, as well as aseptic loosening of the prosthesis resulting from stress shielding-induced bone resorption around the implant.
The aforementioned limitations of THA have prompted the industry to seek less bone-sacrificing options for younger patients, with the hope that a THA can be postponed by at least five years or more. One approach towards improving treatment has been to develop less invasive surgical procedures such as arthroscopic joint irrigation, debridement, abrasion, and synovectomy. However, the relative advantage of these surgical techniques in treating osteoarthritis is still controversial. An alternative to THA is hip “resurfacing,” has now re-emerged because of new bearing surfaces (metal-on-metal, rather than metal-on-polyethylene). While many patients can expect to outlive the procedure's effectiveness, hip resurfacing preserves enough bone stock on the femoral side to allow for later total hip replacement. Unfortunately, there are enough potential drawbacks that doctors offering hip resurfacing say that the procedure should still be deferred as long as possible. In metal-on-metal resurfacing, the femoral head is shaped appropriately and then covered with a metal cap that is anchored by a long peg through the femoral neck. It requires a more precise fit between the cap and cup, and the procedure generally sacrifices more bone from the acetabulum compared to conventional replacements due to the larger diameter of the femoral component. Furthermore, a resurfacing operation has a steep learning curve and takes longer than a THA. Femoral neck fractures caused by bone resorption around the peg have been reported, and the long-term impact of metal ion release from the bearing surfaces is also not yet known in humans. As a result of these complications, today's resurfacing devices are still only indicated in patients for whom hip pain is unbearable, as is the case for THA.
The present invention addresses the needs in the art and provides an interpenetrating polymer network hydrogel that is strain-hardened by means of swelling that forms the basis of an arthroplasty device and a method for making this device.
SUMMARY OF THE DISCLOSUREThe present invention provides a bone-sparing arthroplasty device based on an interpenetrating polymer network hydrogel that is strain-hardened by means of swelling that mimics the molecular structure, and in turn, the elastic modulus, fracture strength, and lubricious surface of natural cartilage. Emulating at least some of these structural and functional aspects of natural cartilage, the hydrogel forms the basis of a novel, bone-sparing, “biomimetic resurfacing” arthroplasty procedure. Designed to replace only cartilage, this material is fabricated as a set of flexible, implantable devices featuring lubricious articular surfaces and osteointegrable bone-interfaces. In principle, the device can be made for any joint surface in the body. For example, a device to cover the tibial plateau will require an analogous bone-preparation and polymer-sizing process. For a device to cover the femoral head in the hip joint, the analogy to a male condom is appropriate in which a cap shaped hydrogel device fits snugly over the contours of the femoral head. For a device to line the acetabulum, the analogy to a female condom is appropriate. A polymer dome stretches over the lip and can be snapped into place to provide a mating surface with the femoral head. In this way, both sides of a patient's hip joint can be repaired, creating a cap-on-cap articulation. However, if only one of the surfaces is damaged, then only one side can be capped, creating a cap-on-cartilage articulation. To create a cap-shaped hydrogel device for the shoulder joint (also a ball-and-socket joint), a process similar to that of the hip joint is used. For instance, a “female condom” can be created to line the inner aspect of the glenoid. Furthermore, devices for other joints in the hand, fingers, elbow, ankles, feet, and intervertebral facets can also be created using this “capping” concept. In one embodiment in the distal femur, the distal femur hydrogel device volume follows the contours of the bone while sparing the anterior and posterior cruciate ligaments.
More specifically, the present invention provides an arthroplasty device having an interpenetrating polymer network hydrogel that is strain-hardened by swelling and is adapted to be held in place in a mammalian joint by conforming to a naturally or artificially prepared geometry of a bone in the mammalian joint. The strain-hardened interpenetrating polymer network hydrogel is based on two different networks. The first network is a non-silicone network of preformed hydrophilic non-ionic telechelic macromonomers chemically cross-linked by polymerization of its end-groups. The second network is a non-silicone network of ionizable monomers. The second network has been polymerized and chemically cross-linked in the presence of the first network and has formed physical cross-links with the first network. Within the interpenetrating polymer network, the degree of chemical cross-linking in the second network is less than the degree of chemical cross-linking in the first network. An aqueous salt solution having a neutral pH is used to ionize and swell the second network in the interpenetrating polymer network. The swelling of the second network is constrained by the first network, and this constraining effect results in an increase in effective physical cross-links within the interpenetrating polymer network. The strain-induced increase in physical cross-links is manifested as a strain-hardened interpenetrating polymer network with an increased initial Young's modulus, which is larger than the initial Young's modulus of either (i) the first network of hydrophilic non-ionic telechelic macromonomers swollen in pure water or in an aqueous salt solution, (ii) the second network of ionized monomers swollen in pure water or in an aqueous salt solution, or (iii) the interpenetrating polymer network hydrogel formed by the combination of the first and second network swollen in pure water. The observed increase in stiffness modulus as a result of strain (induced herein by swelling) is caused by an increase in the number of physical cross-links within the interpenetrating polymer network. For the purposes of the present invention, strain-hardening is defined as an increase in the number of physical cross-links and stiffness modulus with applied strain.
The device arthroplasty has a bone-interfacing region and a bearing region opposite to the bone-interfacing region. The bone-interfacing region is characterized by conforming and capable of fixating to the naturally or artificially prepared geometry of the bone in the mammalian joint.
The device and strain-hardened interpenetrating polymer network hydrogel of the present invention could be varied according to the following embodiments either by themselves or in any combinations thereof. For example, the device can be implanted on one side of the mammalian joint forming a hydrogel-on-cartilage articulation in the mammalian joint. The device could further have a second mating component (i.e. another arthroplasty device as taught in this invention) implanted on the opposing joint surface from the implanted device forming a hydrogel-on-hydrogel articulation. The bone-interfacing region is capable of binding to calcium-containing and phosphate-containing bone-matrix constituents of the bone. In another example, the bone-interfacing region is characterized by having a porosity or surface roughness on the order of 10 to 1000 microns to accommodate bone formation. The bone-interfacing region could also be pre-coated with calcium-containing and phosphate-containing constituents. In still another example, biomolecules could be chemically or physically bonded to the bone-interfacing region.
Instead of having the bone-interfacing region be made of the strain-hardened interpenetrating polymer network hydrogel, the bone-interfacing region could, in one example, be made of a polymeric material chemically bonded to the bearing region. In this example, the bearing region is made of the strain-hardened interpenetrating polymer network hydrogel. In another example, the bearing region and the bone-interfacing region could have different compositions at either side of the device and are physically or chemically and physically integrated with each other within the device.
An adhesive material (biodegradable or non-biodegradable) could be bonded to the bone-interfacing region and would then be capable of bonding the device via the bone-interfacing region to the bone. In another example the device could include a calcium-containing inorganic coating that is chemically or physically bonded to the bone-interfacing region.
In still another example, it is a desire to approximately match the thickness profile of the device to the natural thickness profile of an original cartilage layer. The device can be adapted to fit over a primarily convex or concave three-dimensional bone-receiving surface. In one example, the device is undersized to fit over a primarily convex bone-receiving surface to create an elastic contraction fit over the convex three-dimensional bone-receiving surface. The device is capable of swelling to a swollen equilibrium volume in a fluid and temperature other than body fluids and body temperature prior to implantation and capable of de-swelling to a smaller equilibrium volume, compared to the swollen equilibrium volume, upon implantation and exposure to body fluids or/and body temperature, whereby at the smaller equilibrium volume, the device contracts against or physically grips said primarily convex three-dimensional bone receiving surface.
In another example, the device is oversized to fit against a primarily concave three-dimensional bone-receiving surface to accommodate an elastic expansion fit against the primarily concave bone-receiving surface. The device is capable of at least partially drying or de-swelling to a dried or de-swollen equilibrium volume in a fluid and temperature other than body fluids and body temperature prior to implantation and capable of swelling to a larger equilibrium volume, compared to the dried or de-swollen equilibrium volume, upon implantation and exposure to body fluids and/or body temperature, whereby the larger equilibrium volume expands the device against a primarily concave three-dimensional bone receiving surface.
The hydrophilic non-ionic macromonomer in the first network has a molecular weight between about 275 Da to about 20,000 Da, about 1000 Da to about 10,000 Da, or about 3000 Da to about 8000 Da. In another example, the molar ratio between the ionizable monomers and the hydrophilic non-ionic telechelic macromonomers is greater than or equal to 1:1 or greater than 100:1. In one example, the hydrophilic non-ionic telechelic macromonomer in the first network is a derivative of poly(ethylene glycol), and the ionizable monomers are acrylic acid monomers.
In still another example, the aqueous salt solution has a pH in the range of about 6 to 8. In still other examples, the first network has at least about 50%, at least 75% or at least 95% by dry weight telechelic macromonomers. In still another example, the first network has hydrophilic monomers grafted onto the first network. In still another example, the second network further has hydrophilic macromonomers grafted onto the second polymer network. In still another example, the strain-hardened interpenetrating polymer network hydrogel has a tensile strength of at least about 1 MPa. In still another example, the strain-hardened interpenetrating polymer network hydrogel has an initial equilibrium tensile modulus of at least about 1 MPa. In still another example, the strain-hardened interpenetrating polymer network hydrogel has an equilibrium water content of at least 25%, 35% or 50%. In still another example, the strain-hardened interpenetrating polymer network hydrogel is permeable to the aqueous salt solution and the hydrogel has a permeability coefficient ranging from 1e-17 to 1e-13 m4/Nsec.
In still another example, the coefficient of friction of the bearing region of the strain-hardened interpenetrating polymer network hydrogel in an aqueous solution is less than 0.2. In still another example, one side of the device is modified with another polymeric material, other functional groups, or biomolecules using bifunctional crosslinkers. In one example, the biomolecules could be used to stimulate bone cell growth and/or adhesion. In yet another example, the device is comprised of stimulus-responsive polymeric materials that allow it to shrink or swell to conform to the convexity or concavity of an adjacent joint surface.
The present invention together with its objectives and advantages will be understood by reading the following description in conjunction with the drawings, in which:
1. The starting material for the hydrogel is a solution of telechelic macromonomers 13 with reactive functional end groups 15 dissolved in water 16. The telechelic macromonomers are polymerized to form a first end-linked polymer network 10 swollen in water 16.
2. Hydrophilic, ionizable monomers 14 mixed with water are added to the first polymer network 10 along with a photoinitiator and a crosslinking agent (not shown). The hydrophilic, ionizable monomers are then photopolymerized and cross-linked in the presence of first polymer network 10 to form the second polymer network 11 in the presence of the first. This results in formation of an IPN hydrogel having an end-linked polymer network 10 interpenetrated with a ionizable second network 11 swollen in water 16.
3. The water-imbibed IPN is then immersed in an aqueous salt-containing solution 12 at a typical pH of 7.4 and is swollen to equilibrium, yielding a simultaneous increase in both the water content and the stiffness modulus of the IPN. This IPN swollen in the aqueous salt solution 12 has a higher tensile elastic modulus compared to the IPN swollen in pure water 16 due to strain hardening induced by swelling of the second network 11 within the constraint posed by the highly crosslinked first network 10.
The present invention is a “biomimetic” bone-sparing hydrogel arthroplasty device (
Illustrated in
This device concept can be applied to nearly any joint in the body. For instance, the types of orthopaedic devices for which this invention is potentially useful includes total or partial replacement or resurfacing of the hip (femoral head and/or acetabulum), the knee (the tibial, femoral, and/or patellar aspect), shoulder, hands, fingers (e.g. carpometacarpal joint), feet, ankle, and toes. It is also useful in replacement or repair of intervertebral discs or facets. In the knee, the hydrogel can also serve as a meniscus replacement or a replacement material for the cartilage or bursae in any joint such the elbow or shoulder, or the labrum in joints such as the hip and shoulder.
This device strategy is guided by the limitations of current arthroplasty approaches, which are either highly bone-sacrificing or limited to only the repair of focal defects. The hydrogel device is put in place of damaged cartilage after the damaged cartilage has been removed by the surgeon—cartilage remains may need to be removed because subsequent overlying by the implant might cause unwanted conditions that lead to the differentiation of the remaining cartilage fibrous tissue.
The device itself is comprised of a “bearing” region 5 on one side, and a “bone-interfacing” region 6, in which the former articulates with another bearing surface (either another arthroplasty device such as the present invention or natural cartilage on an apposing joint surface) and the latter interacts with underlying bone.
The device can be described as “biomimetic” (i.e. imitative of a natural cartilage) in that it is comprised of a material that mimics the structure and function of natural articular cartilage. While natural cartilage is composed of a highly negatively charged network of proteoglycans interpenetrating a neutral, rigid network of collagen with a water content of about 75%. In a preferred embodiment, the hydrogel is composed of a highly negatively charged network of poly(acrylic acid) interpenetrating a neutral, rigid hydrophilic, end-linked network of, for example, poly(ethylene glycol) macromonomers, with a water content of at least 35% and up to 90%, but preferably about 70%. Mimicking these structural details is believed to be critical to the formation of a stiff, yet highly lubricious bearing material that behaves like natural cartilage. Other combinations of hydrophilic, end-linked macromonomers and negatively charged second networks are possible. PEG and PAA are arguably the two most biocompatible, hydrophilic polymers available. For instance, PEG is known widely to be resistant to protein adsorption and PAA has recently been shown to have a protective role against macrophage activity in vivo. Although PEG and PAA are conventionally weak individually, we have developed a way to create “strain hardened” IPNs of these materials that mimic the high mechanical strength and elastic modulus, high water content, and low surface friction of natural cartilage. Like natural cartilage, the high mechanical strength and modulus of the hydrogel enable it to take up and distribute loads. At the same time, its high water content and low surface friction enable it to function as a slippery bearing surface, just like the nascent tissue.
Another innovative aspect of the present invention is the anchoring strategy (
In one embodiment, the bone-interface region 6 of the device is prepared such that it interacts with the adjacent bone to allow for anchoring via osteointegration over time. In a version of this embodiment, illustrated in
The localized use of a curable adhesive that bonds the hydrogel to the bone provides a chemical means to attain robust, intraoperative anchoring. In one embodiment the adhesive can be a dental or orthopedic adhesive such as cement (e.g. zinc carbocylate cement), resin, glue or the like. This adhesive may be of one that provides firm bonding between the bearing region of the device and bone. The adhesive in cured form may be porous or non-porous and may be biodegradable or non-biodegradable. In the case of a degradable adhesive, the adhesive material is gradually broken down as new bone is formed that binds to the bone interface region. This degradation takes place over a period of about one to about twelve weeks after being implanted to coincide with the time it takes for new bone to form. In the case of a non-degradable adhesive, the adhesive itself binds and interdigitates with bone even as it is being remodeled.
In another embodiment, the bone interfacing region is made in part from a non-hydrogel polymer such as polyurethane, silicone rubber, or derivatives or combinations thereof (such as copolymers or interpenetrating networks with other polymers such as hydrogels) with good mechanical properties that allow the material to stretch or compress in response to loads and be physically held in place by tensile or compressive stress on or by the adjacent bone. Such a composite material would have a lubricious hydrogel (such as PEG/PAA) as the bearing region and the non-hydrogel polymer (such as polyurethane or silicone-based materials) as the bone-interface region.
One embodiment of the present invention is application as a hip arthroplasty device. According to this embodiment, the arthroplasty hydrogel device is comprised of a femoral head component (1a) and an acetabulum component (2a) as shown in
The overall device geometry resembles the anatomy of natural cartilage. The femoral head component 1a holds a cap shape and is placed on the femoral head 3a bone after the later has been surgically reamed to remove damaged cartilage and the superficial bone layer. The femoral head component 1a bone interface region 6 has a radius of curvature that is slightly undersized compared to the radius of curvature of the femoral head bone 3a; the femoral component 1a can therefore be held in place by means of a tight fit around the femoral head. More specifically, and by analogy to latex condoms, the hydrogel device femoral head component 1a, being slightly undersized than the bone it is mounted onto, is pulled over the femoral head 3a and is held in place by tension generated by stretching of the hydrogel device 1a material. Because the femoral head component 1a material is stretchable, it can be stretched to fit over the femoral head. In one version of this embodiment, this cap shaped device 1a covers the bone 360 degrees on the lateral plane and as much as 200 degrees on the coronal plane. With the bone now occupying its inside space, the hydrogel device femoral head component 1a cannot completely return to its original dimensions, which causes the device 1a to “hug” the bone 3a it surrounds. The entire process can be facilitated by means of a retractor tool that could open up the device 1a opening.
The acetabulum component 2a is placed on the acetabulum bone 4a after the later has been surgically reamed to remove damaged cartilage and the superficial bone layer. The acetabulum hydrogel device component 2a holds a hemispherical shell shape and its bone interface region 6 has a radius of curvature that is slightly oversized compared to the radius of curvature of the acetabulum bone 4a socket; the acetabulum component 2a can be held in place by means of a tight press-fit inside the acetabulum 4a. The hydrogel device acetabulum component may also have a thickness profile that matches that of natural acetabular cartilage and is in the range of 1 mm-5 mm. The dimensions of the hydrogel devices are in accordance with the dimensions of the reamers employed by the surgeon. In addition, the edges of the devices may be rounded to prevent edge stress concentration.
A library of different size devices 1,2 may cover the wide range of joint sizes so that every patient would have a nearly perfect fit. At the time of surgery, the physician would choose and implant the device of the appropriate dimensions. The thickness can be adjusted, if necessary, to accommodate variations in joint surface area and/or the patient's weight, as well as joint conformity factors (i.e. the less conforming the joint, the higher the thickness needs be).
The bone interface region 6 of the device is porous with a pore size in the range of 10-1000 microns. The bone interface region is coated with a layer of soluble or insoluble hydroxyapatite that is chemically deposited by taking advantage of the bonds created due to the negative charges of the hydrogel and the calcium ions contained in the hydroxyapatite crystals as demonstrated in
The surface of the bearing region 5 of the femoral head component 1a has the same radius of curvature as the surface of the bearing region 5 of the acetabulum component 4a to achieve a dimensionally matched ball-in-socket mechanism and thus yield an even distribution of the contact stresses. Furthermore, the bearing region 6a of the acetabulum component may hold in its central region a depression 100 so that a chamber 101 is formed between the bearing sides of the acetabulum component 2a and the femoral component 1a. The chamber 101 is filled with fluid 102 at times of non bearing joint load, said fluid 102 gets pressurized once joint loads are applied since the chamber 101 is effectively sealed by the bearing region 5 surfaces; the pressurized fluid 102 can take up significant portions of the joint load.
The femoral component 1a may have a variable shell thickness profile as shown in
In another embodiment, the hydrogel device can be applied to the knee joint. The device is comprised of a distal femur component 1b and a tibial plateau component 2b as shown in
The tibial plateau component 2b can have a curved disk shape and can be either unilateral or bilateral, that is it can cover both tibial plateau 4b facets, or simply either the lateral or the medial facet depending on the extent of the cartilage damage. One way the tibial plateau component 2b can be fixated in the bone is by surgically creating a depression 113 on the facet surface as shown in
The bone interfacing region 6 of both components is porous, with bone morphogenic proteins tethered on the surface to promote bone adhesion and/or ingrowth as discussed in
Current materials used in arthroplasty function well as mechanical “bearings” but suffer from key material property differences compared to natural cartilage. Because plastics, metals, and ceramics are not hydrated, they solely rely on serum/synovial fluid lubrication; the bearing function relies on the tolerances as well as on the surface roughness. Interfacial wear ultimately produces wear debris by means of abrasion. The products of wear are typically in particulate form (e.g. polyethylene particles) or in the form of ions (e.g., metal ions). Both of these have been shown to be promoters of inflammation in synovial joints and have been found to migrate into internal organs. Moreover, because metals are significantly stiffer than bone, they alter the stress transfer to the bone leading to bone resorption or fibrous tissue formation and ultimately loosening around the implants. One way that researchers have been exploring to avoid problems associated with conventional orthopaedic “hardware” is to use “software” (soft materials). One such approach available in the U.S. is “Carticel” autologous cartilage grafting. This has been shown to be effective in “filling in” focal defects in knee cartilage with regenerated cartilage from a patient's own chondrocytes. There are a number of other approaches under development that are related to tissue engineered cartilage, cell transplantation, and autologous grafting. To date, the simultaneous combination of cartilage-like stiffness and a hydrated, lubricious surface has been an elusive pair of properties to attain in materials engineering.
The present invention provides a hydrogel device 1 having an interpenetrating polymer network (IPN) hydrogel network based on a neutral cross-linked network of end-linked macromonomers 13 as the first network 10 and an ionized crosslinked polymer in the second network 11 depicted in
Homopolymer networks of PEG and PAA are both relatively fragile materials (the former is relatively brittle, the latter is highly compliant). However, the two polymers can form complexes through hydrogen bonds between the ether groups on PEG and the carboxyl groups on PAA. This inter-polymer hydrogen bonding enhances their mutual miscibility in aqueous solution, which, in turn, yields optically clear, homogeneous polymer blends. By loosely cross-linking (instead of densely cross-linking) the ionizable network (PAA, pKa=4.7), large changes in its network configuration can be induced by changing the pH of the solvent without affecting the neutral PEG network. In salt-containing buffers of pH greater than 4.7, the PAA network becomes charged and swells; at a pH lower than 4.7, the PAA network is protonated and contracts.
In one embodiment of the present invention, grafted polymers are used to form the IPN.
Any hydrophilic telechelic macromonomer 13 may be used to form the first polymer network 10. In a preferred embodiment, preformed polyethylene glycol (PEG) macromonomers are used as the basis of the first network (10). PEG is biocompatible, soluble in aqueous solution, and can be synthesized to give a wide range of molecular weights and chemical structures. The hydroxyl end-groups of the bifunctional glycol can be modified into crosslinkable end-groups 15. End-group or side-group functionalities to these macromolecules and biomacromolecules may include, but are not limited to, acrylate (e.g. PEG-diacrylate), methacrylate, vinyl, allyl, N-vinyl sulfones, methacrylamide (e.g. PEG-dimethacrylamide), and acrylamide (e.g. PEG-diacrylamide). For instance, PEG macromonomers can be chemically modified with endgroups such as diacrylates, dimethacrylates, diallyl ethers, divinyls, diacrylamides, and dimethacrylamides. Examples of the end-group functionalization reactions to yield telechelic, crosslinkable PEG macromonomers are shown in
Preferably, the hydrophilic monomer 14 in the second network 11 is ionizable and anionic (capable of being negatively charged). In a preferred embodiment, poly(acrylic acid) (PAA) hydrogel is used as the second polymer network, formed from an aqueous solution of acrylic acid monomers. Other ionizable monomers include ones that contain negatively charged carboxylic acid or sulfonic acid groups, such as methacrylic acid, 2-acrylamido-2-methylpropanesulfonic acid, hyaluronic acid, heparin sulfate, chondroitin sulfate, and derivatives, or combinations thereof. The second network monomer 14 may also be positively charged or cationic. The hydrophilic monomer may also be non-ionic, such as acrylamide, methacrylamide, N-hydroxyethyl acrylamide, N-isopropylacrylamide, methylmethacrylate, N-vinyl pyrrolidone, 2-hydroxyethyl methacrylate, 2-hydroxyethyl acrylate or derivatives thereof. These can be copolymerized with less hydrophilic species such as methylmethacrylate or other more hydrophobic monomers or macromonomers. Crosslinked linear polymer chains (i.e. macromolecules) based on these monomers may also be used in the second network 11, as well as biomacromolecules such as proteins and polypeptides (e.g. collagen, hyaluronic acid, or chitosan).
Adding a photoinitiator to an aqueous solution of the end-linkable macromonomers 13 in water and exposing the solution to UV light results in the crosslinking of the PEG macromonomers, giving rise to a PEG hydrogel that serves as the first network 10. Polymerizing and crosslinking a second network 11 inside the first network will give rise to the IPN structure. Preparing IPN hydrogels through free-radical polymerization has the additional advantage that it enables the use of molds to form hydrogels of desired shape such as the ones depicted in
Any type compatible cross-linkers may be used to crosslink the second network 11 in the presence of any of the aforementioned first networks 10 such as, for example, ethylene glycol dimethacrylate, ethylene glycol diacrylate, diethylene glycol dimethacrylate (or diacrylate), triethylene glycol dimethacrylate (or diacrylate), tetraethylene glycol dimethacrylate (or diacrylate), polyethylene glycol dimethacrylate, or polyethylene glycol diacrylate, methylene bisacrylamide, N,N′-(1,2-dihydroxyethylene) bisacrylamide, derivatives, or combinations thereof. Any number of photoinitiators can also be used. These include, but are not limited to, 2-hydroxy-2-methyl-propiophenone and 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone.
Examples of First Network Telechelic MacromonomersTelechelic PEG macromonomers 13 with acrylate or methacrylate endgroups can be synthesized in the following manner. PEG was dried from Toluene, redissolved in THF (550 mL per 100 g) and kept under Nitrogen. Distilled triethylamine (2.5 eq per OH group) was added slowly to this solution. Acryloyl chloride (or methacryloyl chloride) was then added via a dropping funnel (diluted with THF) over 30 min at room temperature. The reaction (
Networks have also been formed from PEG-diacrylamide. PEG-diol was converted to PEG-diacrylamide (
PEG macromonomers containing diols have also been converted into allyl ethers. Difunctional allyl ether macromonomers were synthesized from PEG using the following procedure (
The following description refers to an exemplary embodiment of a strain-hardened interpenetrating polymer network hydrogel with PEG as a first network 10 polymer and PAA as a second network 11 polymer. The IPN hydrogel is synthesized by a (two-step) sequential network formation technique based on UV initiated free radical polymerization. A precursor solution for the first network is made of purified, telechelic PEG dissolved at a typical concentration of 50% w/v in phosphate buffered saline (PBS) solution, water, or an organic solvent with either 2-hydroxy-2-methyl-propiophenone or 2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone as the UV sensitive free radical initiator. The types of telechelic PEG macromonomers used were PEG-diacrylate, PEG-dimethacrylate, PEG-diacrylamide, and PEG-diallyl ether. In other embodiments, either network can be synthesized by free radical polymerization initiated by other means, such as thermal-initiation and other chemistries not involving the use of ultraviolet light. In the case of UV polymerization, the precursor solution is cast in a transparent mold and reacts under a UV light source at room temperature. Upon exposure, the precursor solution undergoes a free-radical induced gelation and becomes insoluble in water. The mold is fabricated in such a way that yields hydrogels at equilibrium swelling desired dimensions.
To incorporate the second network 11, the PEG-based hydrogel is immersed in the second monomer 14 solution, such as an aqueous solution of (10-100% v/v) acrylic acid containing a photo-initiator and a cross-linker, from 0.1% to 10% by volume triethylene glycol dimethacrylate (TEGDMA), triethylene glycol divinyl ether, N,N-methylene bisacrylamide, or N,N′-(1,2-dihydroxyethylene)bisacrylamide, for 24 hours at room temperature. The swollen gel is then exposed to the UV source and the second network 11 is polymerized and crosslinked inside the first network 10 to form an IPN structure in which the degree of crosslinking in the second network is less than that of the first network. Preferably, the molar ratio of the first network telechelic macromonomer to the second network monomer ranges from about 1:1 to about 1:5000. Also preferably, the weight ratio of the first network to the second network is in the range of about 10:1 to about 1:10. In another embodiment of the present invention, the IPNs have a molar ratio of the second monomer ingredient to the first macromonomer ingredient higher than 100:1.
Key characteristics of hydrogels such as optical clarity, water content, flexibility, and mechanical strength can be controlled by changing various factors such as the second monomer type, monomer concentration, molecular weight and UV exposure time. The experimental focus of the ensuing section is on the swelling induced strain hardening observed in this system by testing how it manifests through uniaxial tensile tests under various conditions of first 10 and second 11 network crosslinking and swelling. Swelling data were used to calculate the equilibrium water and polymer content of the networks, which were correlated with stiffness modulus, true stress-at-break, and true strain-at-break. The results indicate that strain hardening is derived from physical entanglements between the PEG and PAA networks that are intensified by bulk deformation. Under conditions that promote hydrogen bonding (when the pH is at or below 4.7, the pKa of PAA), these entanglements are reinforced by interpolymer complexes between PEG and PAA, leading to an increase in the fracture strength of the IPN. Under conditions that promote ionization of PAA (when the pH is above 4.7 and salt is added), increased steric interactions (i.e. physical crosslinks) between the swelling PAA network and static, telechelic PEG macromonomer network lead to an increase in the stiffness modulus.
In particular embodiment, an array of IPNs with varying molecular weights of PEG in the first network 10 and varying PAA polymer content in the second network 11 were fabricated based on diacrylate crosslinking in the first network 10 and triethylene glycol dimethacrylate crosslinking in the second network 11. All hydrogels were formed by photopolymerization with UV light using the photoinitiator, 2-hydroxy-2-methyl-propiophenone at a concentration of 1% v/v with respect to the monomer 14 or macromonomer 15. Before the IPNs were prepared, single network hydrogels based on PEG and PAA were synthesized separately to confirm the formation of gels of each composition and to investigate the physical properties of the single networks. For the PEG single network, a range of hydrogels that varied between 275 and 14000 for the MW of the PEG macromonomer was synthesized. It was found that low MW PEG macromonomers gave rise to gels that were brittle, whereas the hydrogels made from higher molecular weight PEG-DA (3400) were transparent and flexible when swollen in deionized water. Based on these results, a range of different MWs of PEG (3400, 4600, 8000, and 14000) were chosen as macromonomers for the first hydrogel network. A series of IPNs was synthesized by polymerizing and crosslinking a PAA network within each type of PEG network. The resultant IPNs had significantly better mechanical properties compared with single network hydrogels.
To explore the effect of the molecular weight of the telechelic PEG-DA macromonomer on IPN mechanical strength, PEG chains with MWs 3400 Da, 4600 Da, 8000 Da, and 14000 Da were used in the first network while keeping the acrylic acid polymerization conditions constant (50% v/v in deionized water with 1% v/v crosslinker and 1% v/v photoinitiator with respect to the monomer). The resulting IPNs were characterized in terms of their water content, tensile properties, and mesh size in deionized water. Changing the MW of the PEG-DA macromonomer led to a change in the moduli of the PEG-DA single networks, as shown in Table 2. This effect was magnified in the PEG/PAA IPNs, where the IPNs initial and final moduli get increasingly higher as the networks are prepared from lower molecular weight PEG-DA macromonomers. Of note, there was little increase in strength when the PEG MW is increased above 8000, indicating that a contrast between the molecular weight between cros slinks of the PEG and PAA networks is important for strength enhancement. Moreover, the molecular weight of the PEG macromonomer was strongly correlated to the critical strain (εcrit) at which the stress-strain curve makes the transition from the initial modulus to the strain-hardened final modulus. The εcrit was smaller for the IPNs prepared from lower MW PEG macromonomers, meaning that these networks strain-harden more rapidly in response to deformation.
The significance of forming an interpenetrating structure rather than a copolymeric structure was explored by synthesizing a PEG-co-PAA copolymer hydrogel and testing its tensile properties. Its stress-strain profile was then juxtaposed with those of the IPN and the PEG and PAA single networks. In
To explore the role of interpolymer hydrogen bonding, the pH of the hydrogel swelling liquid was varied to change the ionization state of the PAA network. Since the equilibrium swelling of PAA is sensitive to variations in pH, a change in the pH was expected to have an impact on the mechanical properties of PEG/PAA IPNs. After synthesis, the water-swollen PAA single networks and PEG(8.0 k)/PAA IPNs were placed in buffers of pH 3-6 and constant ionic strength (I) of 0.05. In both the PAA network and the IPN, the equilibrium water content increased as the pH was increased from 3 to 6 (Table 2). In the case of the PAA networks, those at pH 3 and 4 were moderately swollen, while those at pH 5 or 6 were highly swollen due to ionization of PAA above its pKa (4.7). The IPNs also achieved different levels of swelling depending on the pH; those at pH 3 and 4 were moderately swollen, while those at pH 5 or 6 were highly swollen due to ionization of PAA above its pKa (4.7). Of note, at both pH 3 and 4, the IPN achieved a lower equilibrium water content than PAA alone. This can be explained, in part, by the fact that PEG and PAA complex with each other via hydrogen bonds in an acidic environment, leading to a more compact, less hydrated interpenetrating network structure. At pH above 4.7, the PEG and PAA chains dissociate as the PAA becomes ionized and counterions (along with water) enter the hydrogel to maintain charge neutrality, leading to a high degree of swelling. Nevertheless, the IPNs swell to a slightly lower extent (1.0-1.5%) than the PAA single networks due to the constraint that the PEG network places on PAA swelling. Table 2 also shows that the maximum stress (σmax), or tensile strength, of the PEG/PAA IPN is nearly an order of magnitude greater in its less-swollen state at pH 3 (σmax=8.2 MPa) than in its more swollen state at pH 6 (σmax=0.86 MPa). A similar phenomenon is observed in the PAA network, but the absolute values for σmax are 0.38 MPa at pH 3 and 0.05 MPa at pH 6. At every pH, then, the IPN has greater tensile strength than the PAA network, and this difference is intensified at lower pH. In contrast to the differences in the stress-at-break, the trends in the strain-at-break values of the IPN and PAA networks are roughly equivalent, changing from εbreak values of ˜1.2 at pH 3 to ˜0.55 at pH 6. This result confirms the observation made in
To investigate the consequence of relative network moduli even further, the swelling of PAA within the IPNs was maximized. The experimental data shown in Table 2 indicated that the modulus of the IPN was not negatively affected by the increased swelling. The PEG network acts as a constraint on the swelling of PAA in a way that leads to additional interpolymer interactions and a corresponding increase in the IPN modulus. In particular, the increase in the constraining effect of the neutral PEG network on PAA swelling would increase the intensity and number of physical entanglements in the IPN and, in turn, lead to the strain hardening behavior observed in the IPN. To test this hypothesis, the IPNs with first network MW PEG 3400, 4600, and 8000 and constant PAA network conditions were placed in phosphate buffered saline (PBS, pH 7.4, I=0.15) in order to induce maximal swelling under physiologic conditions. Table 2 also shows the equilibrium water content and corresponding swelling ratios for networks prepared from PEG macromonomers with each of these molecular weights, juxtaposed with the water content of the water-swollen and PBS-swollen IPNs. Increasing the size of the first PEG network from 3400 Da to 4600 Da and 8000 Da increases the degree to which the IPN is able to swell. Specifically, while the PEG(3.4 k)/PAA IPN swells to only 70% water when ionized, the PEG(4.6 k)/PAA IPN swells to 77% water and the PEG(8.0 k)/PAA IPN swells to 90% water (nearly the same water content as the PEG(8.0 k) single network) when ionized. Of note, the equilibrium water content values of the PEG(3.4 k) and PEG(4.6 k)-based IPNs do not approach those of their component PEG-DA networks (79.3% and 84.5%, respectively).
The time-dependent water content of the hydrogels was evaluated in terms of the swollen-weight-to-dry-weight ratio. The dry hydrogel was weighed and then immersed in water as well as phosphate buffered saline. At regular intervals, the swollen gels were lifted, patted dry, and weighed until equilibrium was attained. The percentage of equilibrium water content (WC) was calculated from the swollen and dry weights of the hydrogel:
where Ws and Wd are the weights of swollen and dry hydrogel, respectively.
Table 3 shows the effect of varying the concentration of acrylic acid monomer used to prepare the second network on the equilibrium water content of PEG/PAA IPNs in PBS. In general, higher concentrations of acrylic acid monomer leads to hydrogels with lower equilibrium water content and higher stiffness (tensile modulus) and tensile strength for a given set of cros slinking conditions. IPN hydrogels according to the present invention made from these constituents, preferably have an equilibrium water content of between about 15%-95% and more preferably between about 50%-90%.
Because different MWs of PEG and different starting concentrations of acrylic acid result in different amounts of equilibrium water content, the final amount of PEG and PAA in the hydrogel varies depending on the MW of the starting PEG used and the concentration of acrylic acid used. Examples of compositions of varying weight ratios of PEG and PAA that have been made according to the present invention are shown in Table 4. The compositions in this table were all made using a starting concentration of 50% PEG macromonomers of molecular weight 8000 Da swollen in pure water.
Swelling of the PAA network within the confines of a more densely crosslinked PEG network (by lowering the MW of the PEG macromonomer) has dramatic consequences on the resulting IPN modulus. Specifically,
PEG/PAA IPNs were swollen to equilibrium in a series of PBS solutions of varying ionic strength (0.15 M, 0.30 M, 0.75 M, and 1.5 M) and their equilibrium water content and stress-strain properties were measured. Table 2 shows that the water content of the IPN is reduced with higher salt concentration in the swelling medium, from over 90% at I=0.15 to less then 78% at I=1.5. This is caused by the fact that increased salt in the buffer screens the negative charges on the PAA chains, reducing electrostatic repulsion and, in turn, swelling of the networks.
Ionic strength had a modest effect on the stress-strain properties. Table 2 shows that the stress-strain properties of IPNs in I=0.15 to I=0.75 were roughly equivalent. The IPN swollen in buffer with I=1.5 showed a slight enhancement in fracture stress at higher strains. This result is consistent with the water content data, since the hydrogels with higher solids content (the IPN at higher ionic strength conditions) should have greater mechanical strength. Of note, the final modulus of the IPN in the solution with the highest ionic strength (I=1.5) appeared to be higher than those at lower ionic strength. However, the difference was small and was not found to be statistically significant.
To increase the quantity of topological interactions between the PAA and PEG networks, the polymer content of PAA was varied inside of a PEG(3.4 k) first network. The volume fraction of acrylic acid in solution at the time of the second network polymerization was varied between 0.5 and 0.8 prior to polymerization. After polymerization, the IPNs were swollen to equilibrium in PBS. The resultant hydrogels had different water content, from 62% in the PEG(3.4 k)/PAA[0.8] IPN to 65% in the PEG(3.4 k)/PAA[0.7] IPN and 77% in the PEG(3.4 k)/PAA[0.5] IPN. Of note, the IPNs with increased acrylic acid concentration had lower water content, which in light of the super-absorbency of PAA is a counterintuitive result. The water content and tensile properties of these IPNs are shown in Table 3. The IPN with the highest PAA content had the highest stress-at-break and modulus, while the one with the lowest PAA content had the lowest stress-at-break and strain-at-break. Notably, the initial modulus values for these samples varied significantly, from 3.6 MPa in the PEG(3.4 k)/PAA[0.5] to 12 MPa in the PEG(3.4 k)/PAA[0.7] IPN and 19.6 MPa in PEG(3.4 k)/PAA[0.8] IPN.
Effect of PAA Content on IPN Swelling in Pure WaterPEG(4600) single networks were prepared and imbibed with varying concentrations of AA in the second network in the presence of the photoinitiator and crosslinker. IPNs based on these AA-swollen PEG networks were then formed by UV-initiated polymerization. The IPNs were then removed from their molds, immersed in deionized water, and allowed to reach equilibrium. The volume of the IPNs relative to the PEG single networks were then measured and compared. The results are plotted in
The same PEG/PAA IPNs of varying AA monomer content were tested by uniaxial tensile measurements. The results are shown in
To demonstrate that an ionizable monomer is important in the second network, a series of IPNs were prepared under conditions that disrupted the degree of ionizability in the second network. The first method used was copolymerization of the second network with non-ionic monomers. AA monomers in the second network were mixed in three different concentrations relative to the HEA monomers: 10:1, 3:1, and 1:1. Uniaxial tensile testing experiments of the hydrogels swollen in deionized water showed that the PEG/P(AA-co-HEA) IPNs with the highest ratio of AA:HEA in the second network exhibited enhanced mechanical strength. Specifically, changing tensile strength of the IPNs decreased from 9 MPa to 6 MPa and then to 3.5 MPa when the AA:HEA ratio decreased from 10:1 to 3:1 to 1:1, respectively. In other words, IPNs with higher relative HEA content exhibited almost no enhancement in mechanical properties. This result demonstrates that the presence of ionizable carboxyl acid groups in PAA is an important element in the present invention.
In another set of experiments, PEG networks were immersed in AA solutions (containing photoinitiator and crosslinker) that were partially neutralized to pH 5.5 by titration with sodium hydroxide. The monomer-swollen PEG networks were then exposed to UV light to form a partially neutralized PAA network within the PEG network. These “pre-neutralized” PEG/PAA IPNs were then washed in PBS and subjected to uniaxial tensile tests. It was found that neutralizing the AA solution prior to polymerization and then forming the second network leads to an IPN with the same elastic modulus, but with dramatically reduced fracture strength. The stress-at-break is reduced from nearly 4 MPa—in the case of the IPNs prepared under acidic conditions and then neutralized in PBS buffer—to roughly 0.5 MPa. This demonstrates the importance of the fabrication process in creating these strain-hardened IPNs; that is, in the preferred embodiment, ionization and swelling of the second network with buffered, aqueous salt solution should be carried out after the IPN is fully formed.
These results demonstrate that the PEG/PAA IPN system strain-hardens and, in turn, becomes “pre-stressed” with high values for initial stiffness moduli when swollen in buffers of physiologic pH and salt concentrations (e.g. phosphate buffered saline). The strain hardening under these conditions is the result of the constraining effect that the tightly crosslinked, neutral PEG network has on the swelling of the ionized PAA network. This constraining effect leads to additional physical crosslinks between the two networks and manifests as an increase in the initial Young's modulus of the IPN. The tensile modulus values that the hydrogel can attain (12 MPa, but tunable between about 1 to about 20 MPa) exceed those reported in the art. Of note, the hydrogel's modulus (12 MPa) is in the range of values reported for natural healthy human cartilage.
Natural cartilage is, in effect, an avascular “IPN hydrogel” comprised of collagen and negatively charged proteoglycans. By comparison, the IPN hydrogel comprised of PEG and negatively charged PAA. PEG acts as the analog of collagen while PAA acts as the analog of proteoglycans. This fundamental structural similarity of these IPNs to natural cartilage is believed to the reason for their functional similarity: the osmotic pressure created by the polyelectrolyte, coupled with the steric constraint posed by the first network, yields a “pre-stressed” material that, like cartilage, is stiff, yet flexible, and exhibits a highly lubricious surface. To explain the low friction coefficient that cartilage exhibits, a number of scientific approaches have been developed: the fluid-solid stress sharing described by the biphasic theory and the “weeping lubrication” theory are some representative examples. According to these theories, it is important that the material is permeable for low friction to occur; the combination of the permeability coefficient and the equilibrium modulus need to be such so that to allow for the so called “weeping lubrication” but at the same time prevent excessive fluid loss under continuous or repeated dynamic loading. Based on the fact that the strain-hardened IPN has similar permeability, negative charge, water content and stiffness to natural cartilage, we hypothesize that the IPN exhibits a low surface friction coefficient for the same reasons natural cartilage does through any of the aforementioned mechanisms.
We have shown that one of the defining features of the PEG/PAA IPN is its high (compared to state-of-the-art existing hydrogels) tensile stiffness modulus. The tensile stress-strain behavior of the PEG(3400)/PAA(70%) hydrogel material is shown in
Through pin-on-disc tribometer experiments, the wear rates of PEG/PAA hydrogel in PBS and in synovial fluid under physiologic contact stresses were determined; the hydrogel was tested for 3,000,000 cycles at ˜1 Hz loading frequency and the linear wear rate was found to be 0.2 μm/million cycles equivalent to about 0.2 μm/year, suggesting that based on the thickness of the bearing region 5 wear life of the device suffices for a lifetime. The material was also tested in a gel-on-cartilage configuration under dynamic physiologic loading conditions. The test was carried out for 150,000 cycles at a sliding frequency of 1 Hz, and a 0.5-1.5 MPa dynamic loading in a synovial fluid and bovine serum solution. Gross observation showed that neither the cartilage nor PEG/PAA showed any macroscopically discernible fibrillation or wear.
Anchoring SpecificationsInitial anchoring of the device is made possible by the stretch-to-fit fixation provided by the slight size difference between the hydrogel device and the underlying bone. The polymer cap is placed over the femoral head, creating a snug, compressive fit over the bone. In the case of a concave joint such as the hip socket, a slightly oversized female-type implant creates an expansion fit against the walls of the joint.
Biological anchoring of the device is achieved by means of osteointegration with the inorganic constituents of bone. In the present invention, calcium and phosphate ions are bound to PEG/PAA IPNs through their affinity for the PAA component of the hydrogel as illustrated in
Three different sized particles (20 nm, 200 nm, and 5 μm) of HAP were investigated to determine the effect of particle size on surface coverage on the hydrogel as well on the biological response by osteoblast-like cells.
In one example of this anchoring approach, the heterobifunctional crosslinking agent, 3-trimethoxysilylpropylmethacrylate at a concentration of 0.1% w/v in 95% ethanol in deionized water (with pH-adjusted to 4.5) was brushed onto the surface of previously cleaned and dried bovine bone and allowed to dry for 15 minutes and react with the phosphates in the inorganic matrix of the bone. A 25% w/v solution of PEG-dimethacrylate (MW 1000 Da) was then prepared along with 1% v/v 2-hydroxy-2-methyl propiophenone as the photointiator and then spread over the bone-interface surface of a PEG/PAA IPN hydrogel. The PEG-dimethacrylate solution was then allowed to diffuse into the IPN hydrogel for 1 hour. Bone was then placed on top of the PEG-dimethacrylate solution on the IPN hydrogel, and then the bone and the hydrogel were clamped together using a binder clip and glass slide (1.0 mm thick) placed on top of the hydrogel to attain even clamping pressure. The specimen was then placed under a UV light source (350 nm) for 45 seconds to cause the PEG-dimethacrylate to cure. The result was a PEG/PAA IPN hydrogel bonded to the bovine bone specimen through a PEG-dimethacrylate adhesive that is interpenetrated within the bone-interface of the IPN (
An embodiment of the device according to the present invention comprises a bearing region and bone-interfacing region with two different polymeric compositions. In general, this approach leads to a composition gradient within the device as described in
Another embodiment of the present invention is to use an external stimulus to create a composition gradient in the second network within the first network of the IPN as illustrated in
Another embodiment of the device according to the present invention covalently links molecules or biomolecules to a pre-fabricated device in order to create a bone-interface region with different characteristics than the bearing region. In one such embodiment, any suitable biomolecules may be covalently linked to the IPN hydrogel. In another embodiment, a synthetic polymer is linked to the IPN hydrogel. Preferably, the biomolecules are at least one of proteins, polypeptides, growth factors (e.g. epidermal growth factor) amino acids, carbohydrates, lipids, phosphate-containing moieties, hormones, neurotransmitters, or nucleic acids. Any combination of small molecules or biomolecules can be used, including, but not limited to, drugs, chemicals, proteins, polypeptides, carbohydrates, proteoglycans, glycoproteins, lipids, and nucleic acids. This approach may rely, for example, on (a) photoinitiated attachment of azidobenzamido peptides or proteins, (b) photoinitiated functionalization of hydrogels with an N-hydroxysuccinimide ester, maleimide, pyridyl disulfide, imidoester, active halogen, carbodiimide, hydrazide, or other chemical functional group, followed by reaction with peptides/proteins, or (c) chemoselective reaction of aminooxy peptides with carbonyl-containing polymers. These biomolecules may, for example, promote bone cell adhesion or activity. In one example, a heterobifunctional crosslinker 118 (
Implantation of the device through volume changes in the device can be achieved by taking advantage of the stimulus-responsiveness of certain polymers. In addition, fabricating the device with different polymer compositions in the bearing and bone-interfacing regions makes offers an additional level of control over the implantation of the device via external stimuli while preserving certain advantageous attributes of a non-responsive polymer or by introducing new attributes to the responsive polymer. Stimuli hereafter refers to a characteristic change in a property that regulates hydrogel volume or shape; this change is caused by maintaining the hydrogel pre-surgically in an environment that is different than the environment inside the body. In an embodiment of the present invention, an external stimulus such as a change in pH, salt concentration, electric field, or temperature causes the device, after A being placed on the bone, to B shrink to conform to the contours of the convex-shaped bone it surrounds, as depicted in
The interpenetrating polymer networks could have two or more networks or polymeric components (such as linear chains). Examples include but are not limited to a “triple” or even “quadruple” network or a double network interpenetrated with additional polymer chains as discussed in
As one of ordinary skill in the art will appreciate, various changes, substitutions, and alterations could be made or otherwise implemented without departing from the principles of the present invention. Accordingly, the scope of the invention should be determined by the following claims and their legal equivalents.
Claims
1. (canceled)
2. An orthopedic implant comprising a hydrogel and having a first surface and a second surface, the hydrogel having an equilibrium water content of at least 15% and a modulus of at least about 1 MPa, and the first surface having a coefficient of friction in aqueous solution of less than 0.2.
3. The orthopedic implant of claim 2, wherein the orthopedic implant is at least 1 mm in thickness.
4. The orthopedic implant of claim 2, wherein the first and second surfaces are different.
5. The orthopedic implant of claim 2, wherein the implant is configured for cartilage replacement in a joint.
6. The orthopedic implant of claim 5, wherein the joint is selected from a hip, a shoulder, a knee, an elbow, a finger, a joint of the hand and a joint of the foot.
7. The orthopedic implant of claim 2, wherein the implant is a cartilage replacement implant and wherein the first surface is a bearing surface that is adapted to articulate with an additional bearing surface upon implantation.
8. The orthopedic implant of claim 7, wherein the additional bearing surface is a natural cartilage surface or an implanted device surface.
9. The orthopedic implant of claim 7, wherein the second surface is adapted for anchoring the implant to bone.
10. The orthopedic implant of claim 7, wherein the second surface is adapted for anchoring to an acetabular cavity, a tibial plateau, an inner aspect of a glenoid, a femoral head, or a distal femur.
11. The orthopedic implant of claim 7, wherein the first surface has a primarily concave shape and the second surface has a primarily convex shape.
12. The orthopedic implant of claim 11, wherein the primarily convex shape of the second surface is adapted to mate with an acetabular cavity or an inner aspect of a glenoid.
13. The orthopedic implant of claim 7, wherein the first surface has a primarily convex shape and the second surface has a primarily concave shape.
14. The orthopedic implant of claim 13, wherein the primarily concave shape of the second surface is adapted to mate with a femoral head or distal femur.
15. The orthopedic implant of claim 9, wherein the second surface is adapted to interact with adjacent bone to allow for anchoring via osteointegration over time.
16. The orthopedic implant of claim 2, wherein the hydrogel has a permeability coefficient ranging from 1 e−18 to 1 e−12 m4/Nsec.
17. The orthopedic implant of claim 2, wherein the hydrogel is immersed in an aqueous salt solution having a neutral pH.
18. The orthopedic implant of claim 2, wherein the hydrogel is negatively charged at neutral pH.
19. The orthopedic implant of claim 2, wherein the hydrogel comprises sulfonic acid groups.
20. An arthroplasty procedure, wherein an orthopedic implant in accordance with claim 2 is implanted in a joint.
21. The method of claim 20, wherein the orthopedic implant is implanted on one side of the joint in a hemi-arthroplasty procedure.
22. The method of claim 20, wherein the orthopedic implant is implanted on both sides of the joint in total arthroplasty procedure.
23. The method of claim 22, wherein the joint is selected from a hip, knee, shoulder, elbow, a finger, a joint of the hand, or a joint of the foot.
Type: Application
Filed: May 16, 2018
Publication Date: Jan 17, 2019
Applicant: The Board of Trustees of The Leland Stanford Junior University (Stanford, CA)
Inventors: David MYUNG (Santa Clara, CA), Lampros KOURTIS (San Francisco, CA), Laura HARTMANN (Berlin), Curtis W. FRANK (Cupertino, CA), Stuart B. GOODMAN (Los Altos, CA), Dennis R. CARTER (Stanford, CA)
Application Number: 15/981,761
