Patent Application
20220257382
The invention relates to an implant for the replacement and regeneration of biological tissue in the shape of a plug. The invention in particular relates to an implant for the replacement and regeneration of an osteochondral structure in the shape of a plug. The invention further relates to a method for the preparation of the implant, and to an osteochondral structure comprising the implant.
An osteochondral structure refers to a structure comprising cartilage and bone. Typical osteochondral structures can be found in the thighbone (femur), shinbone (tibia), and kneecap (patella). Such structures fit tightly together and move smoothly because the bone surface is covered with a relatively thick layer of articular (hyaline) cartilage. An (osteo)chondral defect is any type of damage to articular cartilage and optionally to underlying (subchondral) bone. Usually, (osteo)chondral defects appear on specific weight-bearing spots at the ends of the thighbone and shinbone and the back of the kneecap for instance. They may range from roughened cartilage, small bone and cartilage fragments that hinder movement, to complete cartilage loss.
Trauma of joint surfaces is common in young active people practicing sports, or as a sequel to accidents. Lesions may comprise the cartilage layer only, but often the underlying subchondral bone too. Articular cartilage has a very low tendency for healing and the repair tissue is qualitatively inferior to the original tissue. This invariably leads to the formation of osteoarthritis (OA) over the years, which is a major cause of disability and loss of quality of life in elderly people. The standard treatment for this condition is ultimately joint replacement by artificial joints. Whilst clinically effective, the non-biological implants do not last longer than 10-20 years and revision surgery is much less effective and very costly. For this reason, much research is dedicated to developing biological regenerative therapies that would be life-long lasting. However, despite promising in vitro results, until now not a single solution has proven to be more effective than the current standard of care over a longer period in real life conditions.
Because the cartilage layer lacks nerve fibers, patients are often not aware of the severity of the damage. During the final stage, an affected joint consists of bone rubbing against bone, which leads to severe pain and limited mobility. By the time patients seek medical treatment, surgical intervention may be required to alleviate pain and repair the cartilage damage. Implants have been developed for the joint in order to avoid or postpone such surgical interventions. These may be implanted in a bone structure at an early stage of cartilage damage, and may thus be provided for preventive treatment, in order to avoid unnoticed degeneration of the joint.
A number of treatments is available to treat articular cartilage damage in joints, such as the knee, starting with the most conservative, non-invasive options and ending with total joint replacement if the damage has spread throughout the joint. Currently available treatments include anti-inflammatory medications in the early stages. Although these may relieve pain, they have limited effect on arthritis symptoms and further do not repair joint tissue. Cartilage repair methods, such as arthroscopic debridement, attempt to at least delay tissue degeneration. These methods however are only partly effective at repairing soft tissue, and do not restore joint spacing or improve joint stability. Joint replacement (arthroplasty) is considered as a final solution, when all other options to relieve pain and restore mobility have failed or are no longer effective. While joint arthroplasty may be effective, the procedure is extremely invasive, technically challenging and may compromise future treatment options. Cartilage regeneration has also been attempted, more in particular by tissue-engineering technology. The use of cells, genes and growth factors combined with scaffolds plays a fundamental role in the regeneration of functional and viable articular cartilage. All of these approaches are based on stimulating the body's normal healing or repair processes at a cellular level. Many of these compounds are delivered on a variety of carriers or matrices including woven polylactic acid based polymers or collagen fibers. Despite various attempts to regenerate cartilage, a reliable and proven treatment does not currently exist for repairing defects to the articular cartilage.
Another standard of care consists of Microfracture (MFx) for smaller lesions (≤2 cm2) and Autologous Chondrocyte Implantation (ACI) for bigger lesions (>2 cm2). The cartilaginous tissue regenerated with these techniques however is not able to withstand the biomechanical challenges in the joint and starts to degenerate within 18 months already. Substantial delay in joint replacement by artificial joints, let alone preventing it, therefore is not possible.
It is an object of the present invention to provide an implant for the replacement and regeneration of biological tissue in the shape of a plug having improved load distribution as well as cartilage regenerating properties. Another aim is to provide such a plug-shaped implant for the replacement and regeneration of an osteochondral structure. Yet another aim is to provide a method for the preparation of the implant. The invention further aims to provide an implant which is able to repair articular cartilage lesions in a durable fashion, and which at least postpones and, preferably, prevents joint replacement by artificial joints.
The above and other aims are provided by a plug-shaped implant in accordance with claim 1. The plug-shaped non-biodegradable implant in particular comprises a base section configured for anchoring in bone tissue, a middle section configured for replacing cartilage tissue of an intermediate and deep zone of the cartilage layer and having a thickness of at least 0.2 mm, and a top section configured for growing cartilage tissue onto and into, thus regenerating a superficial zone of the cartilage layer, wherein the middle and top section comprise the same thermoplastic elastomeric material, which is porous in the top section, and non-porous in the middle section, wherein the thermoplastic elastomeric material comprises a linear block copolymer comprising urethane and/or urea groups, and is substantially free of an added peptide compound having cartilage regenerative properties, and wherein the base section material comprises one of a biocompatible metal, ceramic, mineral, such as phosphate mineral, and polymer, optionally a hydrogel polymer, and combinations thereof. Preferably, the thermoplastic elastomeric material is substantially free of any added compound having cartilage regenerative properties.
In cartilage, a relatively thin superficial (tangential) zone protects deeper layers from shear stresses and makes up approximately 10% to 20% of articular cartilage thickness. The collagen fibers of this zone (primarily, type II and IX collagen) are packed tightly and aligned parallel to the articular surface (
Immediately deep or below to the superficial zone is the middle (intermediate or transitional) zone, which provides an anatomic and functional bridge between the superficial and deep zones. The middle zone represents 40% to 60% of the total cartilage volume, and it contains proteoglycans and thicker collagen fibrils. In this layer, the collagen is organized obliquely, and the chondrocytes are spherical and at low density. Functionally, the middle zone is the first line of resistance to compressive forces.
The deep zone of cartilage is responsible for providing the greatest resistance to compressive forces, given that collagen fibrils are arranged perpendicular to the articular surface. The deep zone contains the largest diameter collagen fibrils in a radial disposition, the highest proteoglycan content, and the lowest water concentration. The chondrocytes are typically arranged in columnar orientation, parallel to the collagen fibers and perpendicular to the joint line. The deep zone represents approximately 30% of articular cartilage volume.
The base section material may be formed of any suitable material which provides an appropriate level of mechanical support to the surrounding bone and preferably allows osteogenesis. Suitable materials, including the thermoplastic elastomeric material of the middle and top section of the implant, are biocompatible, by which is meant that these materials are capable of coexistence with living tissues or organisms without causing harm to them. Further, the implant in accordance with the invention is substantially non-biodegradable and combines cartilage replacement with cartilage regeneration. With a non-biodegradable material in the context of the present invention is meant a material that is not broken down into less complex compounds or compounds having fewer carbon atoms by the environment of the implanted implant. The weight-average molecular weight of a substantially non-biodegradable material is reduced by at most 20%, relative to the original weight-average molecular weight after one year of implantation, more preferably at most 10%, still more preferably at most 5%, and more preferably still at most 1%.
Suitable metals as base section material include but are not limited to titanium, zirconium, chromium, aluminum, stainless steel, hafnium, tantalum or molybdenum, and their alloys, or any combination thereof. Optionally, a surface layer of the metal may be oxidized, nitrided, carburized or boronized to form a coated metal base section.
Suitable ceramics and minerals as base section material include but are not limited to oxides, nitrides, carbides or borides, or any combination thereof. Suitable examples include bioactive glass, calcium phosphates, such as beta-tricalcium phosphate (TCP), biphasic calcium phosphate and apatite such as hydroxylapatite, fluorapatite, chlorapatite, and/or calcium deficient apatite, and combinations thereof.
Suitable (hydrogel) polymers as base section material include but are not limited to collagen, poly(lactic-co-glycolic acid) (PLGA), polylactic acid (PLA), polycaprolactone (PCL), polyvinyl alcohol (PVA), polyvinyl pyrrolidone (PVP), polyacrylamide, polyurethane, polyethylene glycol (PEG), chitin, poly(hydroxyalkyl methacrylate), water-swellable N-vinyl lactams, starch graft copolymers, and derivatives and combinations thereof.
Other preferred materials for the base section comprise a polyaryletherketone (PAEK) polymer. A PAEK polymer comprises a semi-crystalline thermoplastic polymer containing alternately ketone (R—CO—R) and ether groups (R—O—R). The linking group R between the functional groups comprises a 1,4-substituted aryl group. The PAEK polymer used in the base section may inter alia comprise PEK (polyetherketone), PEEK (polyetheretherketone), PEKK (polyetherketoneketone), PEEKK (polyetheretherketoneketone) and PEKEKK (polyetherketoneetherketoneketone). Due to its excellent resistance to hydrolysis, the polyaryletherketone polymer of the base section is advantageously used in the invented implant. It does not break down when sterilized, nor when implanted in the body for an extended time. It also turns out to bond particularly well to the elastomeric material of the middle and top sections.
The material used in the base section of the invented implant may be used as such, or, in an embodiment, may comprise a reinforcing material selected from the group consisting of fibrous or particulate polymers and/or metals.
The base section of the invented implant may also comprise a contrast agent for medical imaging that absorbs radiation, such as a radiocontrast or MRI contrast agent, or a radiopharmaceutical agent that itself emits radiation. The base section may also comprise a small solid object or body, such as a bead, that may for instance comprise a refractory metal such as tantalum.
The base section of the plug-shaped implant functions as a bone anchor, whereas the combination of middle and top sections functions as partial replacement for the damaged cartilage and as scaffold for cartilage regeneration. In the plug-shaped implant, the top section refers to the section that is closest to the cartilage phase, when implanted. The base section refers to the section that is furthest from the cartilage phase, when implanted. The middle section is situated in between the top and base sections.
The cross-section of the plug-shaped implant through a horizontal or a vertical plane may have any suitable shape. The cross-section may be circular, square or may be polygonal, such as hexagonal, octagonal, or decagonal. In some embodiments, the plug-shaped implant may be tapered such that it is shaped as a truncated cone structure. Preferably, the implant has a smaller cross-section at the base section than at the top section. The cross-section (or diameter in case of a cylindrical implant) may vary continuously between the base and top section, or may show discontinuities, for instance at the interface between sections.
When the implant has a tapered profile, the angle of the taper is preferably between 1° and 45°. In some embodiments, the taper is between about 3° and 30°, more preferably between 5° and 30°, even more preferably between 10° and 15°. A tapered profile may facilitate insertion of the implant into an osteochondral defect and may further reduce possible damage to host tissue. The implant is preferably used without any means of attachment and remains in the osteochondral structure by its geometry and the surrounding tissue structure. The implant may be used in the knee, but may also be used for other joints, such as a temporal-mandibular joint, an ankle, a hip, a shoulder, and the like.
According to the invention, the plug-shaped implant on top of the base section further comprises a middle section configured for replacing cartilage tissue, and a top section configured for growing cartilage tissue onto and into, wherein the middle and top section comprise the same thermoplastic elastomeric material. By this is meant that at least its building blocks are chemically the same. As mentioned herein below, some physical properties may differ, for instance their weight averaged molecular weight. The thermoplastic elastomeric material is porous in the top section, and non-porous in the middle section, and comprises a linear block copolymer comprising urethane and/or urea groups. Moreover, the thermoplastic elastomeric material is substantially free of an added peptide compound having cartilage regenerative properties. It has surprisingly been found that the implant of the invention is able to regenerate cartilage tissue, thus avoiding the use of any functional compound exhibiting cartilage regenerative properties. In particular, it has been found that the implant according to this embodiment does not need the use of peptides, for instance those comprising an RGD-sequence. These compounds have been said to enable binding integrin's and thereby stimulating cell adhesion.
The linear block copolymers of the invention are segmented copolymers with elastic properties that originate from hydrogen bonding interaction between molecular chains. Such copolymers comprise ‘hard’ crystallized blocks of polyurethane and/or polyurea segments, and may also comprise ‘hard’ crystallized blocks of polyester and/or polyamide between ‘soft’ blocks. At room temperature, the low melting ‘soft’ blocks may be incompatible with the high melting ‘hard’ blocks, which induces phase separation by crystallization or liquid-liquid demixing. These copolymers exhibit reversible physical crosslinks that originate from crystallization of the ‘hard’ blocks of the segmented copolymer. The thermoplastic elastomers may be formed into any shape at higher temperatures, more in particular at temperatures above the melting point of the ‘hard’ blocks. On the other hand, the thermoplastic elastomers provide mechanical stability and elastic properties at low temperatures, i.e. at typical body temperatures. This makes these materials particularly suitable as replacement material for human or animal cartilage.
The constituents of the thermoplastic elastomer may generally comprise three building blocks: a long-chain diol, for example with a polyether, polyester or polycarbonate backbone, a bifunctional di-isocyanate, and, finally, a chain extender, such as water, another (sometimes short-chain) diol, or a diamine. The latter chain extender is preferred since this leads to bisurea units in the thermoplastic elastomer.
An embodiment of the implant wherein the thermoplastic elastomeric material is aliphatic is preferred. This means that all building blocks of the thermoplastic elastomer are devoid of aromatic groups and contain aliphatic groups only. The thermoplastic elastomer of the invention may be prepared in a one pot procedure, in which a long-chain diol is first reacted with an excess of a di-isocyanate to form an isocyanate-functionalized prepolymer. The latter is subsequently reacted with a chain extender, such as the preferred diamine, which results in the formation of a higher molecular weight thermoplastic elastomeric polymer containing urethane groups. If a diamine is used as the chain extender, the thermoplastic elastomer will also contain bisurea groups, which is preferred.
The synthetic procedure to prepare the thermoplastic elastomers may lead to a distribution in the ‘hard’ block lengths. As a result, the phase separation of these block copolymers may be incomplete, in that part of the ‘hard’ blocks, in particular the shorter ones, are dissolved in the soft phase, causing an increase in the glass transition temperature. This is less desired for the low temperature flexibility and elasticity of the thermoplastic elastomeric material of the top and middle sections. The polydispersity in ‘hard’ blocks shows as a broad melting range, and a rubbery plateau in dynamic mechanical thermal analysis (DMTA) that is dependent on temperature. Preferred embodiments therefore comprise elastomeric block copolymers containing ‘hard’ blocks of substantially uniform length. These may be prepared by fractionation of a mixture of ‘hard’ block oligomers, and subsequent copolymerization of the uniform ‘hard’ block oligomers of a specific length (or length variation) with the prepolymer, mentioned above.
Although the thermoplastic elastomers may be prepared by a chain extension reaction of an isocyanate-functionalized prepolymer with a diamine, they may also be prepared by a chain extension reaction of an amine-functionalized prepolymer with a di-isocyanate. Examples of suitable, commercially available diamines and di-isocyanates include alkylene diamines and/or di-isocyanates, arylene diamines and/or di-isocyanates. Amine-functionalized prepolymers are also commercially available, or can be prepared from (readily available) hydroxy functionalized prepolymers by cyanoethylation followed by reduction of the cyano-groups, by Gabriel synthesis (halogenation or tosylation followed by modification with phthalimide, and finally formation of the primary amine by deprotection of the phthalimide group) or by other methods that are known in the art. Isocyanate-functionalized prepolymers can be prepared by reaction of hydroxy functionalized prepolymers with di-isocyanates, such as for example isophorone di-isocyanate (IPDI), 1,4-diisocyanato butane, 1,6-diisocyanato hexane or 4,4′-methylene bis(phenyl isocyanate). Alternatively, isocyanate-functionalized prepolymers can be prepared from amine-functionalized prepolymers, for example by reaction with di-tert-butyl tricarbonate. Hydroxy-functionalized prepolymers of molecular weights typically ranging from about 500 g/mol to about 5000 g/mol of all sorts of compositions are also advantageously used. Examples include prepolymers of polyether's, such as polyethylene glycols, polypropylene glycols, poly(ethylene-co-propylene) glycols and poly(tetrahydrofuran), polyesters, such as poly(caprolactone)s or polyadipates, polycarbonates, polyolefins, hydrogenated polyolefins such as poly(ethylene-butylene)s, and the like. Polycarbonates are preferred.
Particularly preferred are prepolymers of polycarbonates. Such prepolymers yield an implant according to an embodiment, wherein the thermoplastic elastomeric material further comprises carbonate groups, besides the urethane and/or urea groups. Such an implant has proven to better fulfill the aims of the present invention than other implants. In particular, it has proven to be beneficial in that its mechanical properties are well adapted to the mechanical properties of human or animal cartilage. Surprisingly, regeneration of cartilage is improved when using this embodiment in an implanted implant.
A particularly preferred embodiment of the invention provides an implant, wherein the thermoplastic elastomeric material comprises a poly-urethane-bisurea-alkylenecarbonate, more preferably a poly-urethane-bisurea-hexylenecarbonate.
Apart from disclaiming a peptide compound having cartilage regenerative properties in the linear block copolymer, the implant may comprise agents that facilitate migration, integration, regeneration, proliferation, and growth of cells into and around the implant or patch composition, and/or the injury or defect, and/or promote healing of the injury or defect, and/or are chondrogenic and osteogenic, i.e., build, grow and produce cartilage and bone, respectively. These agents, include but are not limited to cytokine compounds, chemokine compounds, chemo attractant compounds, anti-microbial compounds, anti-viral compounds, anti-inflammatory compounds, pro-inflammatory compounds, bone or cartilage regenerator molecules, cells, blood components (e.g., whole blood and platelets), and combinations thereof. Agents that increase strength and facilitate attachment can also be included in the implant. In a preferred embodiment, the elastomeric linear block copolymer does not comprise any compound having cartilage regenerative properties.
With a substantially non-porous material in the context of the present invention is meant a material having a porosity of less than 20%, relative to the total volume of the material, preferably up to 10%, more preferably up to 5%, and more preferably still up to 1% of the total volume of the material. A porous material comprises pores, which are defined as minute openings. The pores may be micropores, having a diameter of less than 1 mm, and may be macropores, having a diameter of greater than 1 mm. The pores may be interconnected, which is preferred, and which means that pores are internally connected or there is continuity between parts or elements. A non-porous material in the context of the present invention does not mean a material that is impermeable to molecules of any size, and some small molecules may indeed be able to pass through the non-porous material. Rather, a non-porous material in the context of the present invention represents a material that is impermeable to synovial fluid and/or blood.
Pore sizes in the porous parts of the implant may be chosen from 100-1000 micron, more preferably from 100-500 micron, and most preferably from 300-500 micron.
The thermoplastic elastomer used in the top and middle sections of the implant is particularly advantageous since it allows adapting its mechanical properties to those of human and animal cartilage. In an embodiment of the invention, an implant may be provided wherein the elastomeric material of the middle section has an elastic modulus at room temperature of less than 10 MPa, more preferably of less than 8 MPa, of less than 7 MPa, of less than 6 MPa, of less than 5 MPa, of less than 4 MPa, of less than 3 MPa, or of less than 2 MPa.
In the context of the present application, room temperature is meant to be a temperature in the range of 20-30° C., more preferably 25° C.
Likewise, preferred embodiments of the implant comprise a top section wherein the porous elastomeric material of the top section has an elastic modulus at room temperature of less than 80% of the elastic modulus of the elastomeric material of the middle section, more preferably of less than 50%, even more preferably of between 10-50%, even more preferably of between 15-40%, and most preferably of between 20-30% of the elastic modulus of the elastomeric material of the middle section. Such a reduced elastic modulus may be achieved by modifying the porosity of the material of the middle section, or by modifying physical properties of the material in the middle section through changing its weight average molecular weight for instance.
The porosity of the elastomeric material of the top section may be chosen within a broad range. Preferred porosities of the elastomeric material of the top section are selected from 20-80% by volume, more preferably from 30-70% by volume, even more preferably from 40-60% by volume, and most preferably from 45-55% by volume.
A useful embodiment of the invention provides an implant, wherein the base section comprises a core of non-porous base section material and a, preferably circumferential, shell of porous base section material, wherein the shell has a thickness that is less than 10% of a largest diameter of the base section. Other useful embodiments provide an implant wherein the (circumferential) shell has a thickness of less than 9%, of less than 8%, of less than 7%, of less than 6%, of less than 5%, of less than 4%, of less than 3%, of less than 2%, or of less than 1% of a largest diameter of the base section. Alternatively, the cross-sectional area of the (circumferential) shell covers at most 35% of a largest cross-sectional area of the base section. Other useful embodiments provide an implant wherein the cross-sectional area of the (circumferential) shell is less than 30%, less than 25%, less than 20%, less than 15%, less than 10%, less than 5%, less than 3%, or less than 1% of a largest cross-sectional area of the base section.
Embodiments having the above-disclosed preferred combinations of mechanical properties of the top and middle section tend to promote regeneration of cartilage. This is believed to be due to a favorable stress (re)distribution of the osteochondral structure including the implant during (dynamic) loading.
Another embodiment of the invention provides an implant, wherein the base section extends between a top surface and a bottom surface, and comprises a layer of porous base section material, wherein the layer is adjacent to the top surface and has a thickness that is less than 10% of a largest height of the base section, and wherein pores of the base section material in the layer comprise the biocompatible elastomeric material, preferably all pores. In other embodiments, the layer that is adjacent to the top surface has a thickness of less than 10%, of less than 8%, of less than 6%, of less than 5%, of less than 4%, of less than 3%, of less than 2%, or of less than 1% of a largest height of the base section. All the above embodiments may improve the adhesion of the middle section (and top section) to the base section to varying degrees. At the same time, the mechanical properties of the base section, and the support offered by the base section to the implant, remain at an adequate level.
Another embodiment of the invention relates to an implant, comprising a substantially non-porous polyaryletherketone polymer with a porosity of less than 20%, relative to the total volume of the polyaryletherketone polymer.
Yet another embodiment provides an implant wherein the base section comprises a non-porous polyaryletherketone polymer.
In another embodiment of the invention, the top surface of the base section of the implant comprises irregularities or undulations. Irregularities may for instance comprise ridges having a saw-toothed shape. Undulations may be irregular or regular, such as those having a sinusoidal shape.
Another useful embodiment relates to an implant, wherein the base section comprises a centrally located cavity that comprises the biocompatible elastomeric material. Such a cavity may further improve the adhesion of the middle section (and top section) to the base section. The cavity may be cylindrical, or its cross-section may be square, or polygonal. The walls of the cavity may also be provided with irregularities or undulations, or may comprise sections of a larger cross-sectional area than its average cross-sectional area. Several of such cavity sections may be provided at different heights of the base section to form mechanical locking structures.
Yet another embodiment provides an implant, wherein the base section comprises an outer surface having irregularities or undulations. Such outer surface irregularities may for instance comprise ridges having a saw-toothed shape, for instance extending circumferentially over (part of) the outer surface of the base section. Undulations may be irregular or regular, such as those having a sinusoidal shape. The undulations may likewise extend circumferentially over (part of) the outer surface of the base section. Irregularities and undulations may be provided by casting the materials in a suitably profiled mold, or, alternatively, may be provided by mechanical machining, for instance by rotary milling of a molded implant.
A useful embodiment of the invention provides an implant, wherein the middle section comprises a core of non-porous elastomeric material and a circumferential shell of porous elastomeric material, wherein the shell has a thickness that is less than 10% of a largest diameter of the middle section. Other useful embodiments provide an implant wherein the circumferential shell has a thickness of less than 9%, of less than 8%, of less than 7%, of less than 6%, of less than 5%, of less than 4%, of less than 3%, of less than 2%, or of less than 1% of a largest diameter of the middle section. The largest diameter is for instance appropriate in an embodiment wherein the plug-shaped implant is tapered and has circular cross-sections. Alternatively, the cross-sectional area of the circumferential shell covers at most 35% of a largest cross-sectional area of the middle section. Other useful embodiments provide an implant wherein the cross-sectional area of the circumferential shell is less than 30%, less than 25%, less than 20%, less than 15%, less than 10%, less than 5%, less than 3%, or less than 1% of a largest cross-sectional area of the middle section. The largest cross-sectional area is for instance appropriate in an embodiment wherein the plug-shaped implant is tapered.
The height of the plug-shaped implant may be chosen according to the specific application in the body. Heights may vary from 3 to 18 mm for instance. According to a useful embodiment of the invention, an implant is provided wherein a height of the base section, a height of the non-porous middle section, and a height of the porous top section are selected such that a top surface of the implant comes to lie below a top surface of cartilage present on an osteochondral structure when implanted, preferably over a distance of between 0.1-1 mm. This embodiment promotes growing cartilage tissue into, but also onto the top section, whereby a strong fixation is built between the top section and the newly formed cartilage. It has turned out that cartilage cells from the host cartilage have a strong affinity for the segmented elastomer of the top section, and therefore are prone to colonize the surface thereof to produce new hyaline cartilage tissue on top of the implant.
Another embodiment provides an implant wherein a height of the base section, a height of the non-porous middle section, and a height of the porous top section are selected such that a bottom surface of the middle section comes to lie about level with a bottom surface of cartilage present on an osteochondral structure when implanted.
Yet another embodiment of the invention provides a top section, a top surface of which is slightly curved. Preferred radii of curvature of the top surface of the top section in a sagittal plane are selected to range from 15-150 mm, more preferably from 17-125 mm, even more preferably from 19-100 mm, even more preferably from 21-75 mm, even more preferably from 23-50 mm, and most preferably from 25-30 mm. This embodiment may regenerate a new cartilage layer on the top surface of the top section of the implant of about equal thickness across the top surface. The result may be a radius of a top surface of the regenerated cartilage that is about the same as the radius of the surrounding native cartilage layer next to the implant, thereby showing a continuity in radius. The top surface of the top section of the implant may also be curved in a medial-lateral plane, preferably with a radius of curvature with the ranges disclosed above for the sagittal plane. In a practical embodiment, the top surface of the top section of the implant has a radius of curvature that is equal in the sagittal and the medial-lateral plane. This embodiment thus comprises a spherical top surface.
Another aspect of the invention provides a method for the preparation of the implant. A method for the preparation of an implant is provided, comprising the steps of:
a) providing in a mold at room temperature a base section that comprises base section material comprising one of a biocompatible metal, ceramic, mineral, such as phosphate mineral, and polymer, optionally a hydrogel polymer, and combinations thereof; and granules of a thermoplastic elastomeric material on top of the base section, the thermoplastic material comprising a linear block copolymer comprising urethane and urea groups, and substantially free of an added peptide compound having cartilage regenerative properties;
b) closing the mold and heating the above assembly to a temperature of between 100° C. and 250° C. under a pressure of between 1 and 2 GPa, such that the thermoplastic elastomeric material melts and fuses with the base section; and
c) cooling the assembly to room temperature to consolidate the thermoplastic elastomeric material and opening the mold;
d) providing a top section of the thermoplastic elastomeric material with pores either before or after opening the mold.
A preferred embodiment of the method comprises a step a) wherein a base section that comprises a substantially non-porous polyaryletherketone polymer with a porosity of less than 20% relative to the total volume of the polyaryletherketone polymer; and granules of a thermoplastic elastomeric material on top of the base section are provided in a mold at room temperature.
Another embodiment of the invention provides a method wherein after step b) the mold is opened and additional granules of the thermoplastic elastomeric material are added to the mold, and step b) is repeated. The amount of material added in the two-step embodiment of the method may be chosen within wide ranges. Increasingly good results are obtained when the ratio between the first addition and the second addition of granules of the thermoplastic elastomeric material is selected from 01:99 to 99:01, more preferably from 30:70 to 97:03, and most preferably from 70:30 to 95:05.
Another embodiment of the invention provides a method wherein the heating temperature of step b) is between 110° C. and 225° C., more preferably between 120° C. and 200° C., and most preferably between 130° C. and 175° C. Preferred pressures at all cited temperature ranges are between 1.1 and 1.8 GPa, and more preferably between 1.2 and 1.6 GPa.
Yet another aspect of the invention relates to a method for the preparation of a thermoplastic elastomeric material comprising a linear block copolymer comprising urethane and urea groups, and being substantially free of an added peptide compound having cartilage regenerative properties. According to the invention, the method comprises:
In a preferred method according to an embodiment, the diol is selected from a polyester diol, a polyether diol and, preferably, a carbonate diol, and combinations thereof.
Another preferred embodiment provides a method wherein the di-isocyanate comprises an n-alkylene-diisocyanate.
Yet another preferred embodiment of the invention relates to a method wherein the diamine comprises a primary diamine, preferably an n-alkylene-diamine.
The invention will now be further elucidated by the following figures and examples, without however being limited thereto. In the figures:
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This one-pot two-step produced Biomaterial MVH313 was prepared by functionalization of 1.0 molar equivalent of poly(hexylene carbonate) diol (MW=2000) with 2.0 molar equivalents of 1,6-diisocyanatohexane (step 1), and subsequent chain extension using 1.0 molar equivalent of 1,6-diaminohexane (step 2).
In particular, the aliphatic poly-urethane-urea-hexylene carbonate biomaterial of the middle section 3 and the top section 4 was manufactured as follows (with reference to
1H NMR spectroscopy was performed on the resulting polymer, using a Varian 200, a Varian 400 MHz, or a 400 MHz Bruker spectrometer at 298K. DSC was performed using a Q2000 machine (TA Instruments). Heating scan rates of 10° C./min and 40° C./min were used for the assessment of the melting temperature (Tm) and the glass transition temperature (Tg), respectively. The Tm was determined by the peak melting temperature and the Tg was determined from the inflection point.
All reagents, chemicals, materials, and solvents were obtained from commercial sources and were used without further purification. The used poly(hexylene carbonate) diol had an average molecular weight of approximately 2 kg/mol.
The non-porous aliphatic poly-urethane-urea-hexylene carbonate biomaterial had an elastic modulus according to ASTM D638 of 3.6±0.03 MPa.
In a similar one-pot two-step experimental procedure as described in detail for Biomaterial MVH313, Biomaterial MVH309B was also produced. Particularly, Biomaterial MVH309B was prepared by functionalization of 1.0 molar equivalent of poly-tetrahydrofuran diol (MW=2000) with 1.33 molar equivalents of bis(4-isocyanatophenyl)methane (MDI) (step 1), and subsequent chain extension using 0.33 molar equivalent of 1,6-diaminohexane (step 2). Biomaterial MVH309B was isolated as a white, flexible, tough elastomeric polymer.
In a similar one-pot two-step experimental procedure as described in detail for Biomaterial MVH313, Biomaterial MVH312 was also produced. Particularly, Biomaterial MVH312 was prepared by functionalization of 1.0 molar equivalent of poly-tetrahydrofuran diol (MW=2000) with 2.0 molar equivalents of 1,6-diisocyanatohexane (step 1), and subsequent chain extension using 1.0 molar equivalent of 1,6-diaminohexane (step 2). Biomaterial MVH312 was isolated as a flexible, tough elastomeric polymer.
In a similar one-pot two-step experimental procedure as described in detail for Biomaterial MVH313, Biomaterial MVH311 was also produced. Particularly, Biomaterial MVH311 was prepared by functionalization of 1.0 molar equivalent of poly(hexylene carbonate) diol (MW=2000) with 1.33 molar equivalents of bis(4-isocyanatophenyl)methane (MDI) (step 1), and subsequent chain extension using 0.33 molar equivalent of 1,6-diaminohexane (step 2). Biomaterial MVH311 was isolated as a flexible, tough elastomeric polymer.
Stress Relaxation Testing was performed on the two aromatic and two aliphatic polymers of Examples 1-4, as well as on three equine cartilage specimens obtained from the Utrecht Medical Centre. A description of the specimens (e.g. polymer classes) and their dimensions are listed in Table 1. Using an Instron Electropulse E10000, each specimen was compressed at a strain rate of 0.005 s−1 up to a strain of 0.05 mm/mm which remained constant for 1800 s. All tests were done in triplicate. During the tests, load, displacement and time were recorded and afterwards, stress relaxation curves were obtained from the data. Stress relaxation is shown by determining the stress relaxation modulus G(t) at the onset of stress relaxation (G(0)) and 1800 s after the onset of stress relaxation (G(1800)) using the following equation: G(t)=σ(t)/ε0, where σ(t) is the compressive stress and ε0 is the set (constant) strain.
The results are shown in Table 2 below.
The implant 1 was manufactured by attaching the top and middle sections (4, 3) to a PEKK base section 2 which serves as bone anchor. In a method according to an embodiment of the invention, PEKK bone anchors were capped with the poly-urethane-urea-hexylene carbonate biomaterial by pressing small granules of the aliphatic polycarbonate polymer on top of and into the PEKK anchors. For this purpose, a custom press setup was used. Various temperatures (100° C. to about 150° C.), compressive forces (2 kN to about 4 kN) and methods have been tested. The best results were obtained using a two-step procedure, employing a temperature of 150° C. and using a compressive force of 40 kN (4 tons, or 4000 kg; corresponding to a pressure of 1.4 GPa). Lower temperatures than 150° C. seemed to give less homogenously pressed poly-urethane-urea-hexylene carbonate biomaterial layers (sections 3 and 4), while higher temperatures are less desired as the urea groups in the poly-urethane-urea-hexylene carbonate biomaterial may then degrade to some extent. In the first step, ca. 50 mg of the polymer 12 was pressed onto and into the PEKK bone anchor for 15 minutes, while in the second step, ca. 2 mg of polymer 12 was added to the setup and the sample was pressed for another 15 minutes under the same conditions (150° C. and 40 kN). The samples were subsequently removed from the compression setup and were then allowed to cool. After the second pressing step, the surface of the poly-urethane-urea-hexylene carbonate biomaterial layer (sections 3 and 4) on top of the base section 2 seemed to be substantially flat. The biomaterial was almost transparent and colorless. The edges of the biomaterial showed some fringes or frays, and these were removed using a scalpel.
A central hole (241, 242) of the base section 2 was about 4.5 mm deep and about 2 mm in diameter. The hole was substantially filled with the poly-urethane-urea-hexylene carbonate biomaterial, and the attachment of the biomaterial to the PEKK base section 2 seemed quite strong and robust. Removing the biomaterial from the PEKK base section by force, or loosening the connection at the PEKK-biomaterial interfaces, proved practically impossible. All used equipment and accessories that were intended to come into contact with the PEKK base section 2 and/or with the elastomeric biomaterial were rinsed with ethanol or isopropanol and were thereafter dried. After pressing, and cutting the frays, the PEKK-biomaterial plug implant was rinsed with isopropanol and dried. The plugs may also be produced in a sterilized environment, if needed.
As assessed by measuring, the PEKK base section was 6 mm in diameter and 6 mm tall (a height of 6 mm). The central cavity in the base section was about 2 mm in diameter and about 4.5 mm deep. The elastomeric biomaterial (the aliphatic polycarbonate) positioned onto the PEKK base section was about 6 mm in diameter and about 1 mm high. Accordingly, the total PEKK-biomaterial plug implant was about 7 mm tall.
The top section 4 was provided with pores by drilling holes in it with an average diameter of 300 micron, to a final porosity of 50 vol. %. The porous aliphatic poly-urethane-urea-hexylene carbonate biomaterial of the top section 4 had an elastic modulus according to ASTM D638 of 0.9±0.2 MPa.
The implant 1 may be implanted into an osteochondral defect 8 as shown in
As also shown in
Another embodiment of the implant 1 was manufactured by attaching the top and middle sections (4, 3) to a titanium base section 2 which serves as bone anchor. The titanium used was alloy Ti6Al4V, which is readily commercially available. The titanium base section was provided with pores having an average pore size of about 300 microns. In a method according to an embodiment of the invention, titanium bone anchors were capped with a poly-urethane-urea-hexylene carbonate biomaterial by pressing small granules of the aliphatic polycarbonate polymer on top of and into the pores of the titanium anchors. For this purpose, the same custom press setup as used in the previous example was used. Optimum results were again obtained using a two-step procedure, employing a temperature of 150° C. and using a compressive force of 40 kN (4 tons, or 4000 kg; corresponding to a pressure of 1.4 GPa). In the first step, ca. 50 mg of the elastomeric polymer was pressed onto and into the titanium bone anchor for 15 minutes, while in the second step, ca. 2 mg of the elastomeric polymer was added to the setup and the sample was pressed for another 15 minutes under the same conditions (150° C. and 40 kN). The samples were subsequently removed from the compression setup and were then allowed to cool. After the second pressing step, the surface of the poly-urethane-urea-hexylene carbonate biomaterial layer (sections 3 and 4) on top of the base section 2 seemed to be substantially flat. The biomaterial was almost transparent and colorless. Some edges of the biomaterial showed fringes or frays, which were removed using a scalpel.
As with the PEKK base anchor, the titanium base anchor was also provided with a central hole (241, 242) with the same dimensions. The hole was substantially filled with the poly-urethane-urea-hexylene carbonate biomaterial, and the attachment of the biomaterial to the titanium base section 2 was satisfactory.
The titanium base section 2 had the same dimensions as the PEKK base section. Since the same mold was used, the elastomeric biomaterial (the aliphatic polycarbonate) positioned onto the titanium base section was about 6 mm in diameter and about 1 mm high. Accordingly, the total titanium-biomaterial plug implant was about 7 mm tall.
The top section 4 was provided with pores by drilling holes in it with an average diameter of 300 micron, to a final porosity of 50 vol. %. The porous aliphatic poly-urethane-urea-hexylene carbonate biomaterial of the top section 4 had an elastic modulus according to ASTM D638 of 0.9±0.2 MPa.
The implant 1 may be implanted into an osteochondral defect 8 as shown in
As also shown in
Finally, the implant according to the embodiment shown in
It will be apparent that many variations and applications are possible for a skilled person in the field within the scope of the appended claims of the invention.
| Number | Date | Country | Kind |
|---|---|---|---|
| 2023398 | Jun 2019 | NL | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/NL2020/050412 | 6/23/2020 | WO |
