Plug-shaped implant for replacing and regenerating biological tissue and method for producing the same

Abstract

The present invention describes a non-biodegradable plug-shaped implant (1) for replacing and regenerating biological tissue. The implant comprises: a base segment (2) and a top segment (4), wherein the base segment is configured to anchor in bone tissue and the top segment is configured to allow cartilage tissue to grow onto and into the top segment. The top segment comprises a porous thermoplastic elastomer material. The thermoplastic elastomer material comprises a linear block copolymer comprising carbamate groups and urea groups, and the thermoplastic elastomer material may be substantially free of additional peptide compounds having cartilage regeneration properties. The base segment material further comprises one of the following: a biocompatible metal, a ceramic, a mineral such as a phosphate mineral, and a polymer, optionally a hydrogel polymer, and combinations thereof, wherein the thermoplastic elastomer material further comprises carbonate groups.

Description

Technical Field

[0001] The present invention relates to a plug-shaped implant for replacing and regenerating biological tissue. In particular, the present invention relates to a plug-shaped implant for replacing and regenerating osteochondral structures. The present invention further relates to a method for preparing the implant and an osteochondral structure comprising the implant. Background Art

[0002] Osteochondral structures refer to structures that include cartilage and bone. Typical osteochondral structures can be found in the thighbone (femur), shinbone (tibia) and kneecap (patella). Because the bone surface is covered with a relatively thick layer of articular (hyaline) cartilage, such structures fit tightly together and move smoothly. (Osteo) cartilage defects are any type of damage to the articular cartilage and optionally to the underlying (subchondral) bone. Typically, (osteo) cartilage defects occur at specific bearing points at the ends of the thighbone and tibia, as well as at the back of the kneecap. (Osteo) cartilage defects can range from rough cartilage, small bone fragments and cartilage fragments that hinder movement, to complete loss of cartilage.

[0003] Trauma to joint surfaces is common among active young people playing sports or as a consequence of accidents. The lesion may involve only the cartilage layer, but often also the underlying subchondral bone. The healing potential of articular cartilage is very low, and the repaired tissue is inferior in quality to the original tissue. Over time, this invariably leads to the development of osteoarthritis (OA), a major cause of disability and decreased quality of life in the elderly. The standard treatment for this condition is ultimately joint replacement with an artificial joint. While clinically effective, non-biological implants have a lifespan of no more than 10-20 years, and revision surgery is far less effective and expensive. For this reason, much research is focused on developing lifelong, durable bioregenerative therapies. However, despite promising in vitro results, no solution has yet proven more effective than the current standard of care over extended periods under real-life conditions.

[0004] Because the cartilage layer lacks nerve fibers, patients are often unaware of the severity of the injury. During the final stages, the affected joint is filled with bone-on-bone friction, which can cause severe pain and limited mobility. By the time patients seek medical treatment, surgical intervention may be required to relieve the pain and repair the cartilage damage. To avoid or postpone such surgical intervention, implants have been developed for joints. These implants can be implanted into the bone structure at an early stage of cartilage damage and can therefore provide preventative treatment, avoiding the joint from degenerating unnoticed.

[0005] There are many treatments available for 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 drugs in the early stages. While these anti-inflammatory drugs can relieve pain, they have a limited effect on arthritis symptoms and do not repair joint tissue. Cartilage repair methods such as arthroscopic debridement attempt to at least delay tissue degeneration. However, these methods are only partially effective in repairing soft tissue and do not restore joint spacing or improve joint stability. When all other options for relieving pain and restoring mobility have failed or are no longer effective, joint replacement (arthroplasty) is considered the ultimate solution. While arthroplasty can be effective, the procedure is extremely invasive, technically challenging, and may affect future treatment options. Cartilage regeneration has also been attempted, more specifically through tissue engineering techniques. The combination of cells, genes, and growth factors with scaffolds plays a fundamental role in regenerating functional and viable articular cartilage. All of these approaches are based on stimulating the body's normal healing or repair processes at the cellular level. Many of these compounds are delivered on various carriers or matrices including woven polylactic acid-based polymers or collagen fibers.Despite various attempts to regenerate cartilage, there is currently no reliable, proven treatment to repair defects in articular cartilage.

[0006] Another standard of care consists of the following: for smaller lesions (≤2 cm 2 ), microfracture surgery (MFx), and for larger lesions (>2cm 2 ), using autologous chondrocyte implantation (ACI). However, the cartilage tissue regenerated using these techniques cannot withstand the biomechanical challenges of the joint and begins to degenerate within 18 months. Therefore, it is impossible to significantly delay joint replacement by artificial joint, let alone prevent it. Summary of the Invention

[0007] An object of the present invention is to provide a plug-shaped implant for replacing and regenerating biological tissue, which has improved load distribution and cartilage regeneration properties. Another object is to provide such a plug-shaped implant for replacing and regenerating osteochondral structures. Yet another object is to provide a method for preparing such an implant. The present invention further aims to provide an implant that can repair articular cartilage lesions in a durable manner and at least delay, and preferably prevent, joint replacement with an artificial joint.

[0008] The above and other objects are provided by a plug-shaped implant according to claim 1. Specifically, the plug-shaped non-biodegradable implant comprises a base segment and a top segment, the base segment being configured to be anchored in bone tissue, the top segment being configured to replace cartilage tissue in a mid-deep region of a cartilage layer and to allow cartilage tissue to grow onto and into the top segment, thereby regenerating a superficial region of the cartilage layer, wherein the top segment comprises a porous thermoplastic elastomeric material, wherein the thermoplastic elastomeric material comprises a linear block copolymer comprising urethane groups and / or urea groups, and wherein the base segment material comprises one of the following: a biocompatible metal, a ceramic, a mineral such as a phosphate mineral, and a polymer, optionally a hydrogel polymer, and combinations thereof.

[0009] In cartilage, a relatively thin superficial (tangential) zone protects the deeper layers from shear stresses and accounts for approximately 10% to 20% of the thickness of articular cartilage. The collagen fibers in this zone (mainly type II and type IX collagen) are tightly packed and aligned parallel to the joint surface (Fig. 2). The superficial layer contains a relatively large amount of flat chondrocytes, and the integrity of this layer is extremely important for the protection and maintenance of the deeper layers. This zone contacts the synovial fluid and is responsible for most of the tensile properties of the cartilage, enabling the cartilage to resist the shearing, tensile and compressive forces applied by the joint.

[0010] Immediately beneath or beneath the superficial zone lies the middle (intermediate or transitional) zone, which provides an anatomical and functional bridge between the superficial and deep zones. The middle zone accounts for 40% to 60% of the total volume of the cartilage and contains proteoglycans and thicker collagen fibers. Within this layer, collagen is organized obliquely, and chondrocytes are spherical and low in density. Functionally, the middle zone is the first line of resistance to compressive forces.

[0011] Because collagen fibers are arranged perpendicular to the joint surface, the deep zone of cartilage is responsible for providing the greatest resistance to compressive forces. This zone contains radially arranged collagen fibers of the largest diameter, the highest proteoglycan content, and the lowest water concentration. Chondrocytes are typically arranged in a columnar orientation parallel to the collagen fibers and perpendicular to the joint line. The deep zone accounts for approximately 30% of the volume of articular cartilage.

[0012] The porous top section of the non-biodegradable implant of the present invention replaces at least the middle and deep regions of the cartilage.

[0013] Preferably, the thermoplastic elastomeric material is substantially free of additional peptide compounds having cartilage regenerative properties.Even more preferably, the thermoplastic elastomeric material is substantially free of any additional compounds having cartilage regenerative properties.

[0014] The base section material can be formed from any suitable material that provides an appropriate level of mechanical support to the surrounding bone and preferably allows for bone formation. Suitable materials, including the thermoplastic elastomer material of the top section of the implant, are biocompatible, which means that these materials are able to coexist with living tissue or an organism without causing harm to said living tissue or said organism. Further, the implant according to the present invention is essentially non-biodegradable and combines cartilage replacement with cartilage regeneration. In the context of the present invention, non-biodegradable materials are meant to be materials that are not broken down into less complex compounds or compounds with fewer carbon atoms by the environment in which the implant is implanted. The weight average molecular weight of the essentially non-biodegradable material is reduced by at most 20%, more preferably at most 10%, still more preferably at most 5%, and still more preferably at most 1%, relative to the original weight average molecular weight one year after implantation.

[0015] Suitable metals as base segment materials include, but are not limited to, titanium, zirconium, chromium, aluminum, stainless steel, hafnium, tantalum or molybdenum and alloys thereof or any combination thereof. Optionally, the surface layer of the metal may be oxidized, nitrided, carburized or boronized to form a coated metal base segment.

[0016] Suitable ceramics and minerals as base segment materials include, but are not limited to, oxides, nitrides, carbides, or borides, or any combination thereof. Suitable examples include bioactive glass, calcium phosphates such as β-tricalcium phosphate (TCP), biphasic calcium phosphate, and apatites such as hydroxyapatite, fluorapatite, chlorapatite, and / or calcium-deficient apatite, and combinations thereof.

[0017] Suitable (hydrogel) polymers as base segment materials include, but are not limited to, collagen, poly(lactic-co-glycolic acid) (PLGA), polylactic acid (PLA), polycaprolactone (PCL), polyvinyl alcohol (PVA), polyvinylpyrrolidone (PVP), polyacrylamide, polyurethane, polyethylene glycol (PEG), chitin, poly(hydroxyalkyl methacrylate), water-swellable N-vinyl lactams, starch graft copolymers, and derivatives and combinations thereof.

[0018] Other preferred materials for the base segment include polyaryletherketone (PAEK) polymers. PAEK polymers comprise semicrystalline thermoplastic polymers containing alternating ketone (R-CO-R) and ether groups (ROR). The linking group R between the functional groups comprises a 1,4-substituted aryl group. The PAEK polymers used in the base segment may include, in particular, PEK (polyetherketone), PEEK (polyetheretherketone), PEKK (polyetherketoneketone), PEEKK (polyetheretherketoneketone), and PEKEKK (polyetherketoneetherketoneketone). The polyaryletherketone polymers of the base segment are advantageously used in the implants of the present invention due to their excellent hydrolysis resistance. The polyaryletherketone polymers do not decompose during sterilization or when implanted in the body for extended periods of time. They have been shown to bond particularly well with the elastomeric material of the top segment.

[0019] The material used in the base section of the implant of the present invention may be used as is, or in one embodiment, may include a reinforcing material selected from the group consisting of fibrous or particulate polymers and / or metals.

[0020] The base section of the implant of the present invention may also include a contrast agent for medical imaging that absorbs radiation, such as a radiocontrast agent or an MRI contrast agent, or a radiopharmaceutical that itself emits radiation. The base section may also include a small solid object or body, such as a bead, which may, for example, include a refractory metal such as tantalum.

[0021] The base section of a plug-shaped implant serves as a bone anchor, while the top section serves as a partial replacement for damaged cartilage and a scaffold for cartilage regeneration. In a plug-shaped implant, the top section is the section closest to the cartilage phase when implanted. The base section is the section farthest from the cartilage phase when implanted.

[0022] In some embodiments, the plug-shaped implant may be of a substantially rectangular shape, such as a hexagon, an octagon, or a decagon. In some embodiments, the plug-shaped implant may be of a substantially rectangular shape, such as a hexagon, an octagon, or a decagon. In some embodiments, the plug-shaped implant may be of a substantially rectangular shape, such as a hexagon, an octagon, or a decagon. In some embodiments, the plug-shaped implant may be of a substantially tapered shape, such that the plug-shaped implant is shaped as a truncated cone structure. Preferably, the cross section of the base section of the implant is less than the cross section of the top section. Cross section (or diameter in the case of cylindrical implants) can continuously vary between the base section and the top section, or can for example illustrate discontinuity at the interface between the sections.

[0023] In one embodiment, the implant has a tapered profile, wherein the angle of the taper is preferably between 1 ° and 45 °. In certain embodiments, the taper is between approximately 3 ° and 30 °, more preferably between 5 ° and 30 °, even more preferably between 10 ° and 15 °. The tapered profile can help implant to be inserted into the osteochondral defect and can further reduce the damage that may be caused to host tissue. Described implant preferably uses and remains in the osteochondral structure by its geometry and surrounding tissue structure under the situation that there is no attachment mode. Described implant can be used in the knee joint, but also can be used for other joints, as temporomandibular joint, ankle joint, hip joint, shoulder joint etc.

[0024] According to the present invention, the plug-shaped implant includes a top segment located on top of a base segment, the top segment having a dual function: the top segment is used to replace cartilage tissue and is configured to grow cartilage tissue onto and into the top segment. The thermoplastic elastomer material of the top segment is porous and includes a linear block copolymer comprising urethane and / or urea groups. Furthermore, in one embodiment, the thermoplastic elastomer material is substantially free of additional peptide compounds having cartilage regeneration properties. It has been surprisingly discovered that the implant of the present invention is capable of regenerating cartilage tissue, thereby avoiding the use of any functional compounds exhibiting cartilage regeneration properties. Specifically, it has been discovered that the implant according to this embodiment does not require the use of peptides, such as those comprising the RGD sequence. These compounds are said to be able to bind to integrins and thereby stimulate cell adhesion.

[0025] The linear block copolymers of the present invention are segmented copolymers with elastic properties derived from hydrogen bonding interactions between molecular chains. Such copolymers include "hard" crystalline blocks of polyurethane segments and / or polyurea segments, and may also include "hard" crystalline blocks of polyester and / or polyamide between the "soft" blocks. At room temperature, the low-melting-point "soft" blocks may be incompatible with the high-melting-point "hard" blocks, which may cause phase separation through crystallization or liquid-liquid demixing. These copolymers exhibit reversible physical crosslinking derived from the crystallization of the "hard" blocks of the segmented copolymers. Thermoplastic elastomers can be formed into any shape at higher temperatures, more specifically at temperatures above the melting point of the "hard" blocks. On the other hand, 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 materials for human or animal cartilage.

[0026] The composition of thermoplastic elastomers can generally include three structural units: a long-chain diol, for example with a polyether, polyester or polycarbonate backbone; a difunctional diisocyanate; and finally, a chain extender such as water, another (sometimes short-chain) diol or a diamine. The latter chain extender is preferred because it leads to diurea units in the thermoplastic elastomer.

[0027] Embodiments of the implant in which the thermoplastic elastomer material is aliphatic are preferred. This means that all structural units of the thermoplastic elastomer are free of aromatic groups and contain only aliphatic groups. The thermoplastic elastomers of the present invention can be prepared in a one-pot procedure in which a long-chain diol is first reacted with an excess of diisocyanate to form an isocyanate-functionalized prepolymer. The isocyanate-functionalized prepolymer is then reacted with a chain extender, such as a preferred diamine, which results in the formation of a higher molecular weight thermoplastic elastomer polymer containing urethane groups. If a diamine is used as a chain extender, the thermoplastic elastomer will also contain diurea groups, which is preferred.

[0028] The synthesis procedure for preparing thermoplastic elastomers may result in a distribution of "hard" block lengths. Therefore, the phase separation of these block copolymers may be incomplete because some of the "hard" blocks, especially the shorter "hard" blocks, dissolve in the soft phase, resulting in an increase in the glass transition temperature. This is less desirable for the low-temperature flexibility and elasticity of the thermoplastic elastomer material in the top section. The polydispersity in the "hard" blocks is shown as a rubbery plateau in a wide melting range and temperature-dependent dynamic mechanical thermal analysis (DMTA). Therefore, a preferred embodiment includes an elastomeric block copolymer containing "hard" blocks of substantially uniform length. These elastomeric block copolymers can be prepared by fractionating a mixture of "hard" block oligomers and then copolymerizing a uniform "hard" block oligomer of a specific length (or length variation) with the prepolymer mentioned above.

[0029] Although thermoplastic elastomers can be prepared by chain extension reactions of isocyanate-functionalized prepolymers with diamines, they can also be prepared by chain extension reactions of amine-functionalized prepolymers with diisocyanates. Examples of suitable commercially available diamines and diisocyanates include alkylenediamines and / or diisocyanates, arylene diamines and / or diisocyanates. Amine-functionalized prepolymers are also commercially available or can be prepared from (readily available) hydroxyl-functionalized prepolymers by cyanoethylation and subsequent reduction of the cyano group, by the Gabriel synthesis (halogenation or tosylation followed by modification with phthalimide and finally by deprotection of the phthalimide groups to form the primary amine), or by other methods known in the art. Isocyanate-functionalized prepolymers can be prepared by reacting hydroxyl-functionalized prepolymers with diisocyanates, such as isophorone diisocyanate (IPDI), 1,4-diisocyanatobutane, 1,6-diisocyanatohexane or 4,4'-methylenebis(phenyl isocyanate). Alternatively, isocyanate-functionalized prepolymers can be prepared, for example, by reacting with tert-butyl tricarbonate from amine-functionalized prepolymers. Hydroxy-functionalized prepolymers having a molecular weight generally in the range of about 500 g / mol to about 5000 g / mol of various compositions are also advantageously used. Examples include the following prepolymers: polyethers such as polyethylene glycol, polypropylene glycol, poly(ethylene-to-propylene) glycol and poly(tetrahydrofuran), polyesters such as poly(caprolactone) or polyadipate, polycarbonates, polyolefins, hydrogenated polyolefins such as poly(ethylene-butylene), etc. Polycarbonates are preferred.

[0030] Particularly preferred are prepolymers of polycarbonate. According to one embodiment, such prepolymers produce an implant in which, in addition to urethane groups and / or urea groups, the thermoplastic elastomer material further comprises carbonate groups. This implant has been shown to achieve the objectives of the present invention better than other implants. Specifically, it has been shown that the implant is beneficial because its mechanical properties are well-suited to those of human or animal cartilage. Surprisingly, when this embodiment is used in an implanted implant, cartilage regeneration is improved.

[0031] A particularly preferred embodiment of the present invention provides an implant wherein the thermoplastic elastomeric material comprises poly-urethane-bisurea-alkylene carbonate, more preferably poly-urethane-bisurea-hexylene carbonate.

[0032] In addition to preferably abandoning the peptide compound with cartilage regeneration properties, and in other embodiments abandoning any compound with cartilage regeneration properties, in linear block copolymers, implants can include agents that promote cell migration, integration, regeneration, proliferation and growth into implant or patch composition or around the implant or patch composition and / or into damage or defect or around the damage or defect, and / or promote the healing of damage or defect, and / or are chondrogenic and osteogenic, i.e., respectively build, grow and produce cartilage and bone. These agents include but are not limited to cytokine compounds, chemokine compounds, chemical attractant compounds, antimicrobial compounds, antiviral compounds, anti-inflammatory compounds, proinflammatory compounds, bone or cartilage regeneration agent molecules, cells, blood components (e.g., whole blood and platelets) and combinations thereof. The implant can also include agents that increase strength and promote attachment.

[0033] The thermoplastic elastomer material of the top section is porous. A porous material comprises pores which are defined as tiny 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 this means that the pores are internally connected or that there is continuity between the parts or elements. A non-porous material in the context of the present invention does not mean a material which is impermeable to molecules of any size, and some small molecules may indeed be able to pass through a non-porous material. Rather, a non-porous material in the context of the present invention means a material which is impermeable to synovial fluid and / or blood. In the context of the present invention, a substantially non-porous material means a material having a porosity of less than 20% of the total volume of the material, preferably at most 10%, more preferably at most 5%, and still more preferably at most 1%, relative to the total volume of the material.

[0034] The pore size in the porous portion of the implant may be selected from 100-1000 microns, more preferably 100-500 microns, and most preferably 300-500 microns.

[0035] The thermoplastic elastomer used in the top section of the implant is particularly advantageous because it allows its mechanical properties to be adapted to those of human and animal cartilage. In one embodiment of the present invention, an implant may be provided wherein the porous elastomeric material of the top section has an elastic modulus at room temperature of less than 8 MPa, more preferably less than 6 MPa, less than 5 MPa, less than 4 MPa, less than 3 MPa, less than 2 MPa, and most preferably less than 1 MPa.

[0036] In the context of this application, room temperature means a temperature in the range of 20-30°C, more preferably 25°C.

[0037] Embodiments having the preferred mechanical properties of the top section disclosed above tend to promote cartilage regeneration. This is believed to be due to a favorable stress (re)distribution of the osteochondral structure comprising the implant during (dynamic) loading.

[0038] The elastic modulus may be affected by changing the porosity of the material of the top section or by changing the physical properties of the material, such as by changing the weight average molecular weight of the material of the top section.

[0039] The average porosity of the elastomeric material of top section can be selected in a wide range.The preferred average porosity of the elastomeric material of top section is selected from 20-80 volume %, more preferably 30-70 volume %, even more preferably 40-60 volume %, and most preferably 45-55 volume %.The porosity of the elastomeric material in the top section can be substantially the same across the top section.Alternatively, the porosity of the elastomeric material in the top section can vary to some extent across the top section.The porosity of the elastomeric material in the top section can vary to some extent in the transverse direction of plug-shaped implant and / or the longitudinal direction of plug-shaped implant.Preferred embodiments relate to a kind of implant, wherein the porosity of the elastomeric material in the top section increases from the low value at the center of plug-shaped implant towards the higher value at the outside of described implant in the transverse direction of plug-shaped implant.In another preferred embodiment, the porosity of the elastomeric material in the top section increases from the low value at the bottom surface of the top section towards the higher value at the top surface of the top section in the longitudinal direction of plug-shaped implant. The low value of the porosity can be, for example, between 20 vol.% and 45 vol.%, preferably between 25 vol.% and 45 vol.%, even more preferably between 30 vol.% and 45 vol.%, and most preferably between 35 vol.% and 45 vol.%. The high value of the porosity can be, for example, between 45 vol.% and 70 vol.%, more preferably between 45 vol.% and 65 vol.%, even more preferably between 45 vol.% and 60 vol.%, and most preferably between 45 vol.% and 55 vol.%.

[0040] In an implant according to the invention, the base section is in direct contact with the porous top section. A useful embodiment of the invention provides an implant in which the base section comprises a core consisting of a non-porous base section material and a circumferential shell, preferably consisting of a porous base section material, wherein the thickness of the shell is less than 10% of the maximum diameter of the base section. Other useful embodiments provide an implant in which the thickness of the (circumferential) shell is less than 9% of the maximum diameter of the base section, less than 8% of the maximum diameter of the base section, less than 7% of the maximum diameter of the base section, less than 6% of the maximum diameter of the base section, less than 5% of the maximum diameter of the base section, less than 4% of the maximum diameter of the base section, less than 3% of the maximum diameter of the base section, less than 2% of the maximum diameter of the base section or less than 1% of the maximum diameter of the base section. Alternatively, the cross-sectional area of ​​the (circumferential) shell covers at most 35% of the maximum 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% of the maximum cross-sectional area of ​​the base segment, less than 25% of the maximum cross-sectional area of ​​the base segment, less than 20% of the maximum cross-sectional area of ​​the base segment, less than 15% of the maximum cross-sectional area of ​​the base segment, less than 10% of the maximum cross-sectional area of ​​the base segment, less than 5% of the maximum cross-sectional area of ​​the base segment, less than 3% of the maximum cross-sectional area of ​​the base segment, or less than 1% of the maximum cross-sectional area of ​​the base segment.

[0041] Another embodiment of the present invention provides an implant wherein the base segment extends between a top surface and a bottom surface and comprises a layer of porous base segment material, wherein the layer is adjacent to the top surface and has a thickness less than 10% of the maximum height of the base segment, and wherein the pores of the base segment material in the layer comprise a biocompatible elastomeric material, preferably all pores. In other embodiments, the thickness of the layer adjacent to the top surface is less than 10% of the maximum height of the base segment, less than 8% of the maximum height of the base segment, less than 6% of the maximum height of the base segment, less than 5% of the maximum height of the base segment, less than 4% of the maximum height of the base segment, less than 3% of the maximum height of the base segment, less than 2% of the maximum height of the base segment, or less than 1% of the maximum height of the base segment. All of the above embodiments can improve the adhesion of the top segment to the base segment to varying degrees. At the same time, the mechanical properties of the base segment and the support provided by the base to the implant remain at an adequate level.

[0042] Another embodiment of the present invention is directed to an implant comprising a substantially non-porous poly(aryletherketone) polymer having a porosity of less than 20% relative to the total volume of the poly(aryletherketone) polymer.

[0043] Yet another embodiment provides an implant wherein the base segment comprises a non-porous polyaryletherketone polymer.

[0044] In another embodiment of the present invention, the top surface of the base section of the implant comprises irregularities or a wavy profile. The irregularities may, for example, comprise ridges having a sawtooth shape. The wavy profile may be irregular or regular, such as those having a sinusoidal shape.

[0045] Another useful embodiment relates to an implant in which the base section includes a centrally located cavity comprising a biocompatible elastomeric material. Such a cavity can further enhance the adhesion of the top section to the base section. The cavity can be cylindrical, or the cross-section of the cavity can be square or polygonal. The walls of the cavity can also be provided with irregularities or wavy contours, or can include sections having a cross-sectional area greater than the average cross-sectional area of ​​the walls. Multiple such cavities can be provided at different heights of the base section to form a mechanical locking structure.

[0046] Another embodiment provides a kind of implant, wherein base section comprises the outer surface with irregularity or wavy profile.Such outer surface irregularity can for example comprise the ridge with sawtooth shape, and described ridge for example extends circumferentially above the (part) outer surface of base section.Wavy profile can be irregular or regular, as those wavy profiles with sinusoidal shape.Wavy profile can also extend circumferentially above the (part) outer surface of base section.Irregularity and wavy profile can be provided by casting material in the mould of suitable shaping, or alternatively, can be provided by machining, for example, by the rotary milling of moulded implant provides.

[0047] The height of the plug-shaped implant can be selected according to the specific application in vivo. For example, the height can vary from 3mm to 18mm. According to a useful embodiment of the present invention, an implant is provided in which the height of the base section and the height of the porous top section are selected so that when the implant is implanted, the top surface of the implant is located below the top surface of the cartilage present on the osteochondral structure, preferably within a distance of 0.1mm and 1mm. This embodiment promotes the growth of cartilage tissue into the top section and onto the top section, thereby establishing a firm fixation between the top section and the newly formed cartilage. It has been shown that chondrocytes from the host cartilage have a strong affinity to the segmented elastomer of the top section and are therefore easy to colonize on the surface of the elastomer to produce new hyaline cartilage tissue on the top of the implant.

[0048] Another embodiment provides an implant wherein the height of the base section and the height of the porous top section are selected such that when the implant is implanted, the bottom surface of the top section is substantially flush with the bottom surface of cartilage present on the osteochondral structure.

[0049] Another embodiment of the present invention provides top section, the top surface of described top section is slightly curved.The preferred radius of curvature of the top surface of top section in sagittal plane is selected to be 15-150mm, more preferably 17-125mm, even more preferably 19-100mm, even more preferably 21-75mm, even more preferably 23-50mm, and most preferably 25-30mm.This embodiment can regenerate new cartilage layer of approximately equal thickness across top surface on the top surface of the top section of implant.The radius of the top surface of the regenerated cartilage is approximately identical with the radius of the surrounding natural cartilage layer of the adjacent implant, thus illustrating the continuity of radius.The top surface of the top section of implant also can bend in medial-lateral plane, and preferably its radius of curvature has the scope disclosed above for sagittal plane.In practical embodiment, the top surface of the top section of implant has equal radius of curvature in sagittal plane and medial-lateral plane.This embodiment therefore comprises spherical top surface.

[0050] Another aspect of the present invention provides a method for preparing the implant. A method for preparing an implant is provided, the method comprising the following steps:

[0051] a) providing a base segment in a mold at room temperature, the base segment comprising a base segment material comprising one of the following: a biocompatible metal, a ceramic, a mineral such as a phosphate mineral, and a polymer, optionally a hydrogel polymer, and combinations thereof; and providing particles consisting of a thermoplastic elastomeric material on top of the base segment, the thermoplastic material comprising a linear block copolymer comprising urethane groups and urea groups;

[0052] b) closing the mold and heating the assembly to a temperature between 100° C. and 250° C. at a pressure between 1 GPa and 2 GPa such that the thermoplastic elastomer material melts and fuses with the base segment; and

[0053] c) cooling the assembly to room temperature to solidify the thermoplastic elastomer material and opening the mold;

[0054] d) providing a top section of the thermoplastic elastomer material with holes before or after opening the mold.

[0055] A preferred embodiment provides a method wherein step a) comprises providing a base segment material comprising a substantially non-porous polyaryletherketone polymer in a mold at room temperature, the substantially non-porous polyaryletherketone polymer having a porosity of less than 20% relative to the total volume of the polyaryletherketone polymer.

[0056] A preferred embodiment provides a method wherein the thermoplastic elastomeric material is substantially free of additional peptide compounds having cartilage regenerative properties, and even more preferably free of any compounds having cartilage regenerative properties.

[0057] Another embodiment of the present invention provides a method wherein, after step b), the mold is opened and further pellets of the thermoplastic elastomeric material are added to the mold, and step b) is repeated. In the two-step embodiment of the method, the amount of material added can be selected within a wide range. Increasingly better results are achieved when the ratio between the first addition of pellets of thermoplastic elastomeric material and the second addition 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.

[0058] Another embodiment of the present invention provides a process, 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. The preferred pressure at all mentioned temperature ranges is between 1.1 GPa and 1.8 GPa, and more preferably between 1.2 GPa and 1.6 GPa.

[0059] Yet another aspect of the present invention relates to a method for preparing a thermoplastic elastomeric material comprising a linear block copolymer containing urethane and urea groups, and substantially free of additional peptide compounds having cartilage regeneration properties. According to the present invention, the method comprises:

[0060] - preparing an isocyanate-terminated prepolymer by reacting a diol with a diisocyanate,

[0061] - polymerizing the isocyanate-terminated prepolymer by chain extension with a diamine;

[0062] The above steps are performed excluding the peptide compound having cartilage regeneration properties.

[0063] Preferably, the above-described method for preparing a thermoplastic elastomeric material is performed in the absence of any compound having cartilage regenerative properties.

[0064] In a preferred method according to an embodiment, the diol is selected from polyester diols, polyether diols and preferably carbonate diols and combinations thereof.

[0065] Another preferred embodiment provides a process wherein the diisocyanate comprises an n-alkylene-diisocyanate.

[0066] Yet another preferred embodiment of the present invention relates to a process wherein the diamine comprises a primary diamine, preferably an n-alkylene-diamine. BRIEF DESCRIPTION OF THE DRAWINGS

[0067] The present invention will now be further illustrated by the following figures and examples, but the present invention is not limited thereto. In the accompanying drawings:

[0068] Figures 1A to 1D Schematic side views of four embodiments of exemplary implants according to the present invention are shown;

[0069] Figure 2A shows a schematic perspective view of a base segment according to an embodiment of the present invention;

[0070] Figure 2B Shown Figure 2A A schematic cross section of an embodiment of;

[0071] Figure 2C and 2D Shown Figure 2B Schematic detail views of parts B and C of an embodiment of the present invention;

[0072] Figure 3 shows a schematic representation of a possible synthesis route for a thermoplastic polycarbonate material according to an embodiment of the present invention;

[0073] Figure 4 Thermoplastic polycarbonate material according to an embodiment of the present invention is shown. 1 H-NMR spectroscopy;

[0074] Figures 5A to 5C shows DSC thermograms of thermoplastic polycarbonate materials according to embodiments of the present invention at different heating rates;

[0075] Figures 6A to 6C shows a schematic representation of a defect in an osteochondral structure (6A) comprising an implant according to an embodiment of the invention (6B) and the same osteochondral structure after cartilage growth onto / into the implant (6C); Figures 7A to 7D shows schematic side views of four embodiments of an implant according to yet another embodiment of the present invention; and finally

[0076] Figures 8A to 8CShown are schematic representations of a defect in an osteochondral structure (8A) comprising an implant according to another embodiment of the invention (8B) and the same osteochondral structure after cartilage growth onto / into the implant (8C).

[0077] Referring to FIG1 , a side view of an embodiment of an exemplary implant according to the present invention is shown. The implant 1, which is in the shape of a plug, includes a base section 2 and a porous top section 4, wherein the base section is configured to anchor in bone tissue and the porous top section is configured to displace cartilage tissue and allow cartilage tissue to grow onto and into the porous top section. The top section 4 comprises a thermoplastic elastomer material in a porous form. The thermoplastic elastomer material in this embodiment comprises poly-urethane-bisurea-hexamethylene carbonate, the preparation and properties of which will be further explained below. The base section 2 comprises a non-porous polyaryletherketone polymer, which in the embodiment shown is a non-porous PEKK polymer. The implant 1 is cylindrical and has a diameter 10 of 6 mm. The height 20 of the base section 2 and the height 40 of the top section 4 add up to a total height of 6 mm.

[0078] Figure 1B A side view of another embodiment of an implant according to the present invention is schematically shown. The specific implant 1 in the shape of a plug again includes a base segment 2 and a top segment 4, wherein the base segment is configured to anchor in bone tissue and the top segment is configured to replace cartilage tissue and allow cartilage tissue to grow onto and into the top segment. The top segment 4 includes the same porous poly-urethane-bisurea-hexamethylene carbonate material. The base segment 2 includes a substantially non-porous PEKK polymer having a porosity of less than 20% relative to the total volume of the PEKK polymer. Specifically, the base segment 2 of this embodiment includes a core 21 composed of a non-porous PEKK polymer and a circumferential shell 22 composed of a porous PEKK polymer. The thickness 23 of the shell 22 is approximately 8% of the diameter 10 of the base segment 2 (and the implant 1). The base segment 2 further extends between a top surface 24 and a bottom surface 25 and includes a layer 26 of porous PEKK polymer, the layer 26 being adjacent to the top surface 24 and having a thickness 27 of about 8% of the height 20 of the base segment 2. The pores of the PEKK polymer in the layer 26 include a biocompatible poly-urethane-bisurea-hexamethylene carbonate, which is derived from the top segment 4 and infiltrated into the pores of the PEKK polymer in the layer 26 during manufacture. A method for manufacturing the implant will be further explained below. Figure 1AAs in the embodiment of the present invention, the implant 1 is cylindrical and has a diameter 10 of 6 mm. The height 20 of the base section 2 and the height 40 of the top section 4 add up to a total height of 6 mm.

[0079] Figure 1C A side view of yet another embodiment of an implant according to the present invention is schematically shown. The specific implant 1, in the shape of a plug, again includes a base segment 2 configured to anchor in bone tissue and a top segment 4 configured to displace and grow cartilage tissue onto and into the top segment. Top segment 4 comprises a polyurethane-bisurea-hexamethylene carbonate material that is porous in top segment 4. Base segment 2 comprises a substantially non-porous PEKK polymer having a porosity of less than 20% relative to the total volume of the PEKK polymer. Specifically, base segment 2 of this embodiment further extends between a top surface 24 and a bottom surface 25 and includes a layer 26 of porous PEKK polymer, adjacent to top surface 24 and having a thickness 27 of approximately 8% of the height 20 of base segment 2. The pores of the PEKK polymer in layer 26 include biocompatible poly-urethane-bisurea-hexamethylene carbonate, which originates from the top section 4 and penetrates into the pores of the PEKK polymer in layer 26 during manufacture. Figure 1A and 1B The sizes and shapes of the embodiments are the same.

[0080] Figure 1D A side view of yet another embodiment of an implant according to the present invention is schematically shown. A specific implant 1 in the form of a plug corresponds to Figure 1C. In addition, the porosity of the elastomeric material in the top segment 4Ap increases in the transverse direction 30 of the plug-shaped implant 1 from a low value of about 35 vol.% at the centerline 3 of the plug-shaped implant to a higher value of about 55 vol.% at the outside of the implant 1. Further, the porosity of the elastomeric material in the top segment 4 increases in the longitudinal direction 31 of the plug-shaped implant 1 from a low value of about 35 vol.% at the bottom surface of the top segment 4 (and the top surface 24 of the base segment 2) to a higher value of about 55 vol.% at the top surface 41 of the top segment 4. In addition, the base segment 2 includes a layer 26 of porous PEKK polymer, the layer 26 being adjacent to the top surface 24 and having a thickness 27 of about 5% of the height 20 of the base segment 2. The pores of the PEKK polymer in layer 26 include biocompatible poly-urethane-bisurea-hexamethylene carbonate, which originates from the top segment 4 and infiltrates into the pores of the PEKK polymer in layer 26 during manufacture. The base segment 2 further includes a core 21 composed of a non-porous PEKK polymer and a circumferential shell 22 composed of a porous PEKK polymer. The thickness 23 of the shell 22 is approximately 5% of the diameter 10 of the base segment 2 (and the implant 1). Finally, the base segment 2 also includes a layer 28 composed of a porous PEKK polymer, which is adjacent to the bottom surface 25 and has a thickness 29 of approximately 5% of the height 20 of the base segment 2. Size and shape are similar to Figures 1A to 1C The sizes and shapes of the embodiments are the same.

[0081] Please note that in Figure 1B 、 1C In 1D and 1D, the circumferential shells (22, 32) are shown in cross-section to illustrate the respective thicknesses (23, 33) of the circumferential shells. In side view, the circumferential shells would extend over the full diameter 10 of the implant 1.

[0082] refer to Figure 7A , shows a side view of another embodiment of an implant according to the present invention. The implant 1 in the shape of a plug comprises Figure 1A The materials and sections are the same as those shown in the . Figure 7A The size of the implant and Figure 1A Instead of having a flat top surface 41 of the top section 4 (and implant 1), as Figure 1A , the top surface 41a of the top section 4 is spherical (not drawn to scale) with a radius of curvature R of approximately 28 mm.

[0083] refer to Figure 7B , shows a side view of another embodiment of an implant according to the present invention. The implant 1 in the shape of a plug comprises Figure 1BThe materials and sections are the same as those shown in the . Figure 7B The size of the implant and Figure 1B Instead of having a flat top surface 41 of the top section 4, as Figure 1B , the top surface 41a of the top section 4 is spherical (not drawn to scale) with a radius of curvature R of approximately 28 mm.

[0084] refer to Figure 7C , shows a side view of another embodiment of an implant according to the present invention. The implant 1 in the shape of a plug comprises Figure 1C The materials and sections are the same as those shown in the . Figure 7C The size of the implant and Figure 1C Instead of having a flat top surface 41 of the top section 4, as Figure 1C , the top surface 41a of the top section 4 is spherical (not drawn to scale) with a radius of curvature R of approximately 28 mm.

[0085] refer to Figure 7D , shows a side view of another embodiment of an implant according to the present invention. The implant 1 in the shape of a plug comprises Figure 1D The materials and sections are the same as those shown in the . Figure 7D The size of the implant and Figure 1D Instead of having a flat top surface 41 of the top section 4, as Figure 1D , the top surface 41a of the top section 4 is spherical (not drawn to scale) with a radius of curvature R of approximately 28 mm.

[0086] Note again that in Figure 7B 、 7C In 7D and 7D, the circumferential shells (22, 32) are shown in cross-section to illustrate the respective thicknesses (23, 33) of the circumferential shells. In side view, the circumferential shells would extend over the full diameter 10 of the implant 1 (not drawn to scale).

[0087] refer to Figures 2A to 2D , schematically shows an embodiment of the base section 2 of the implant 1 of the present invention. The base section 2 shown is essentially cylindrical with a diameter of 10 and a height of 20. The top surface 24 of the base section has a circumferential flat rim portion 240 that gradually extends into a centrally located cavity 241. The cavity 241 is provided with a locking portion 242 having a diameter greater than the diameter of the cavity 241. As shown in Figure 2CAs shown in detail in , the locking portion 242 of the cavity 241 is disc-shaped, whereby the outer edge of the disc makes an angle 246 of between 1° and 20°, more preferably between 5° and 15°, with the longitudinal direction 247 of the base segment 2. During manufacture of the implant, the cavity 241 (and portion 242) is filled with a portion of a biocompatible elastomeric material to provide adequate locking of the top segment 4 to the base segment 2. As discussed above, the base segment 2 comprises a PEKK polymer, which can be non-porous or substantially non-porous, the latter embodiment including the examples disclosed above. It is further seen that the base segment 2 comprises an outer surface having irregularities or wavy contours. In this embodiment, these irregularities or wavy contours include circumferential ridges 243, which are sawtooth-shaped in cross-section, such as Figure 2D The angle 244 at which the serrated flanks extend relative to a transverse direction 245 of the base section 2 is preferably between 70° and 85°, more preferably between 75° and 80°.

[0088] Preparation of the elastomeric material of the top section

[0089] Example 1: Polycarbonate-Aliphatic: Poly(hexamethylene carbonate urethane)-bisurea biomaterial MVH313, see Table 1 below.

[0090] This one-pot, two-step biomaterial, MVH313, was prepared by functionalizing 1.0 molar equivalent of poly(hexamethylene carbonate) diol (MW=2000) with 2.0 molar equivalents of 1,6-diisocyanatohexane (step 1) and subsequently chain extending using 1.0 molar equivalent of 1,6-diaminohexane (step 2).

[0091] Specifically, the aliphatic poly-urethane-urea-hexamethylene carbonate biomaterial of the top section 4 was manufactured as follows (refer to Figure 3). Poly(hexamethylene carbonate) diol (MW=2000; 23.9 g, 11.9 mmol) was weighed into a 500 mL 3-necked flask and dried under vacuum by heating to 75 ° C overnight, after which it was allowed to cool to room temperature. Under an argon atmosphere, 1,6-diisocyanatohexane (4.1 g, 23.9 mmol), DMAc (20 mL) and a drop of Sn(II) bis(2-ethylhexanoate) were added, and the mixture was heated and stirred for 3 hours, with the viscosity increasing. The mixture was allowed to cool to room temperature, diluted with DMAc (100 mL), and a solution of 1,6-diaminohexane (1.4 g, 11.9 mmol) in DMAc (50 mL) was added in one portion with thorough mixing. A gel formed immediately after addition and mixing. The mixture was further diluted with DMAc (150 mL) and heated in an oil bath at 130 ° C to obtain a homogeneous viscous slurry. After cooling to room temperature, the mixture was precipitated in a water / brine mixture (2.75 L water + 0.25 L saturated brine) to give a soft white material. This material was cut into smaller pieces and stirred in a 1:5 mixture of methanol and water (3 L) for 64 hours. After decanting the supernatant, the resulting solid was stirred in a 2:1 mixture of methanol and water (0.75 L) for 6 hours. The supernatant was decanted, stirred in a 2:1 mixture of methanol and water (0.75 L) for 16 hours, the supernatant was decanted, and the solid was dried in a vacuum at 70°C to give a flexible, tough elastomeric polymer.

[0092] The resulting polymers were analyzed at 298 K using a Varian 200, Varian 400 MHz, or 400 MHz Bruker spectrometer. 1 H NMR spectroscopy. DSC was performed using a Q2000 machine (TA Instruments). Melting temperature (Tm) and glass transition temperature (Tg) were evaluated using heating scan rates of 10°C / min and 40°C / min, respectively. Tm was determined by the peak melting temperature, and Tg was determined based on the inflection point.

[0093] All reagents, chemicals, materials and solvents were obtained from commercial sources and used without further purification.The average molecular weight of the poly(hexamethylene carbonate) diol used was approximately 2 kg / mol. Figure 4 and 5 show the obtained polymers respectively. 1 H NMR spectrum and DSC thermogram. 1 The H NMR spectroscopy results can be summarized as follows: 1H NMR (400 MHz, HFIP-d2): δ = 4.23 (m, n*4H, n~14.3), 4.10 (m, 4H), 3.17 (m, 12H), 1.87-1.32 (multiple signals of aliphatic CH2 methylene groups) ppm. The average molecular weight of the repeating hard / soft block segments is approximately 2.5 kDa. DSC results can be summarized as follows: DSC (10°C / min, Figure 5A ): Tm (top) = 20.9°C (soft block melting); DSC (40°C / min, Figure 5B ): Tg = -38.0 ° C. No second melting point of the hard block was observed up to 200 ° C. However, when the final heating was performed at 10 ° C / min to 250 ° C ( Figure 5C ), a small and broad melting transition is observed at about 227° C. In the DSC graph, the endothermic melting peak is plotted downward, while the exothermic crystallization peak is plotted upward.

[0094] The elastic modulus of the non-porous aliphatic poly-urethane-urea-hexylene carbonate biomaterial according to ASTM D638 is 3.6±0.03 MPa.

[0095] Example 2 : Polyether-aromatic: Poly(tetramethylenetetrahydrofuran carbamate)-bisurea biomaterial MVH309B, see Table 1 below.

[0096] Biomaterial MVH309B was also produced in a one-pot, two-step experimental procedure similar to that described in detail for biomaterial MVH313. Specifically, biomaterial MVH309B was prepared by functionalizing 1.0 molar equivalent of polytetramethylene glycol (MW=2000) with 1.33 molar equivalents of bis(4-isocyanatophenyl)methane (MDI) (step 1) and subsequently chain-extending it with 0.33 molar equivalents of 1,6-diaminohexane (step 2). Biomaterial MVH309B was isolated as a white, flexible, tough elastomeric polymer.

[0097] Example 3: Polyether-aliphatic: Poly(tetrahydrofuranyl carbamate)-bisurea biomaterial MVH312, see Table 1 below.

[0098] Biomaterial MVH312 was also produced in a one-pot, two-step experimental procedure similar to that described in detail for biomaterial MVH313. Specifically, biomaterial MVH312 was prepared by functionalizing 1.0 molar equivalent of poly-tetramethylene glycol (MW=2000) with 2.0 molar equivalents of 1,6-diisocyanatohexane (step 1) and subsequently chain-extending it with 1.0 molar equivalent of 1,6-diaminohexane (step 2). Biomaterial MVH312 was isolated as a flexible, tough, elastomeric polymer.

[0099] Example 4: Polycarbonate-Aromatic: Poly(hexamethylene carbonate urethane)-bisurea biomaterial MVH311, see Table 1 below.

[0100] Biomaterial MVH311 was also produced in a one-pot, two-step experimental procedure similar to that described in detail for biomaterial MVH313. Specifically, biomaterial MVH311 was prepared by functionalizing 1.0 molar equivalent of poly(hexamethylene carbonate) diol (MW=2000) with 1.33 molar equivalents of bis(4-isocyanatophenyl)methane (MDI) (step 1) and subsequently chain-extending it with 0.33 molar equivalents of 1,6-diaminohexane (step 2). Biomaterial MVH311 was isolated as a flexible, tough, elastomeric polymer.

[0101] Mechanical properties of the elastomeric material of the non-porous top section

[0102] Stress relaxation test is carried out on two kinds of aromatic polymers and two kinds of aliphatic polymers of example 1-4 and three horse cartilage samples obtained from Utrecht Medical Centre (Utrecht Medical Centre). The description of the size of sample (for example, polymer category) and described sample is listed in Table 1. Using Instron Electropulse E10000, each sample is compressed to the strain of 0.05mm / mm with the strain rate of 0.005s-1, and described strain keeps constant for 1800 seconds. All tests are carried out in triplicate. During the test, load, displacement and time are recorded, and then stress relaxation curve is obtained from the data. Stress relaxation is shown by using the following equation to determine the stress relaxation modulus G (t) when stress relaxation starts (G (0)) and after stress relaxation starts 1800 seconds (G (1800)): G (t) = σ (t) / ε 0, wherein σ (t) is compressive stress, and ε 0 is setting (constant) strain.

[0103]

[0104] Table 1 : Overview of stress relaxation testing. All tests were performed in triplicate.

[0105] The results are shown in Table 2 below.

[0106]

[0107] Table 2 : Stress relaxation modulus of the material at and after 1800 seconds after the onset of stress relaxation9.

[0108] Preparation of biomaterial-covered PEKK bone anchors

[0109] Implant 1 is manufactured by attaching a top segment 4 to a PEKK base segment 2, which serves as a bone anchor. In a method according to an embodiment of the present invention, a PEKK bone anchor is covered with a polyurethane-urea-hexylene carbonate biomaterial by pressing small particles composed of an aliphatic polycarbonate polymer onto the top of the PEKK anchor and into the PEKK anchor. For this purpose, a custom pressing setup was used. Various temperatures (100°C to approximately 150°C), compression forces (2 kN to approximately 4 kN), and methods were tested. The best results were achieved using a two-step procedure, employing a temperature of 150°C and a compression force of 40 kN (4 tons or 4000 kg; corresponding to a pressure of 1.4 GPa). Temperatures below 150°C appear to result in less homogeneous pressing of the polyurethane-urea-hexylene carbonate biomaterial layer (segments 3 and 4), while higher temperatures are less desirable because the urea groups in the polyurethane-urea-hexylene carbonate biomaterial may subsequently degrade to some extent. In the first step, approximately 50 mg of polymer 12 was pressed onto the PEKK bone anchor and into the bone anchor for 15 minutes, while in the second step, approximately 2 mg of polymer 12 was added to the setup and the sample was pressed under the same conditions (150°C and 40 kN) for another 15 minutes. The sample was then removed from the compression setup and allowed to cool. After the second pressing step, the surface of the polyurethane-urea-hexylene carbonate biomaterial layer (segments 3 and 4) on top of base segment 2 appeared essentially flat. The biomaterial was almost transparent and colorless. The edges of the biomaterial showed some streaks or wear, which were removed using a scalpel.

[0110] The center hole (241, 242) of the base section 2 is about 4.5 mm deep and about 2 mm in diameter. The hole is essentially filled with poly-urethane-urea-hexamethylene carbonate biomaterial, and the attachment of the biomaterial to the PEKK base section 2 appears to be quite firm and robust. Practice has shown that it is impossible to remove the biomaterial from the PEKK base section or to loosen the connection at the PEKK-biomaterial interface with force. All used equipment and accessories intended to come into contact with the PEKK base section 2 and / or with the elastomeric biomaterial are rinsed with ethanol or isopropyl alcohol and then dried. After pressing and cutting off the wear, the PEKK-biomaterial plug implant is rinsed with isopropyl alcohol and dried. If necessary, the plug can also be produced in a sterilized environment.

[0111] As assessed by measurement, the diameter of the PEKK base section is 6 mm, and the base section is 6 mm high (6 mm in height). The diameter of the central cavity in the base section is approximately 2 mm and the depth is 4.5 mm. The diameter of the elastomeric biomaterial (aliphatic polycarbonate) positioned on the PEKK base section is approximately 6 mm and the height is approximately 1 mm. Therefore, the total height of the PEKK-biomaterial plug implant is approximately 7 mm.

[0112] The top section 4 was provided with pores by drilling holes with an average diameter of 300 μm, resulting in a final porosity of 50 vol.%. The elastic modulus of the non-porous aliphatic poly-urethane-urea-hexamethylene carbonate biomaterial of the top section 4 was 0.9±0.2 MPa according to ASTM D638.

[0113] like Figures 6A to 6C As shown in FIG, the implant 1 can be implanted into an osteochondral defect 8. In a typical method, a cartilage defect extending into the subchondral bone is drilled ( Figure 6A ), and insert the plug-shaped implant 1 into the drill hole under a certain pressure ("press fit"), as shown Figure 6B Bone is then grown onto the PEKK base segment 2, and in some embodiments, into the base, thereby anchoring the implant 1. The surrounding natural cartilage 5 grows onto the top side 41 of the top segment 4, and new cartilage 5a is generated on top of the implant 1, as shown in FIG. Figure 6C As shown in . Figure 6CAs shown in FIG, the height 20 of the base section 2 and the height 40 of the porous top section 4 are selected so that when the implant is implanted, the top surface 41 of the implant 1 is located below the top surface 50 of the cartilage 5 present on the osteochondral structure (5, 6), preferably within a distance 51 between 0.1 mm and 1 mm. In this case, this distance is about 0.5 mm. The osteochondral structure (5, 6) includes subchondral bone 6 and a cartilage layer 5 on top of the subchondral bone. A synovial cavity 7 is also typically present.

[0114] Also like Figure 6B and 6C As shown in , the height 20 of the base segment 2 and the height 40 of the porous top segment 4 are selected so that when the implant is implanted, the bottom surface 24 of the top 4 (or the top surface 24 of the base segment 2) is roughly flush with the bottom surface 51 of the cartilage layer 5 of the osteochondral structure (5, 6).

[0115] Finally, you can also Figures 7A to 7D The implant of the embodiment shown in FIG. Figures 8A to 8C . Due to the spherical top surface 41a of the top layer 4, this embodiment can regenerate a new cartilage layer 5a on the top surface 41a of the top section 4 of the implant 1, the cartilage layer having approximately equal thickness across the top surface 41a. As a result, the radius of the top surface 50 of the regenerated cartilage 5a may be approximately the same as the radius of the surrounding natural cartilage layer 5 immediately adjacent to the implant, thereby illustrating the continuity of the radius.

[0116] Preparation of metal bone anchors covered with biomaterials

[0117] Another embodiment of an implant 1 was manufactured by attaching a top section 4 to a titanium base section 2 serving as a bone anchor. The titanium used was the commercially readily available alloy Ti6A14V. The titanium base section was provided with pores having an average pore size of approximately 300 micrometers. In a method according to an embodiment of the present invention, the titanium bone anchor was coated with a polyurethane-urea-hexylene carbonate biomaterial by pressing small particles composed of an aliphatic polycarbonate polymer onto the top of the titanium anchor and into the pores of the titanium anchor. For this purpose, a custom pressing setup identical to that used in the previous example was used. Optimal results were again achieved using a two-step procedure, employing a temperature of 150°C and a compressive force of 40 kN (4 tons or 4000 kg; corresponding to a pressure of 1.4 GPa). In the first step, approximately 50 mg of elastomeric polymer was pressed onto and into the titanium bone anchor for 15 minutes, while in the second step, approximately 2 mg of 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 sample was then removed from the compression setup and allowed to cool. After the second pressing step, the surface of the polyurethane-urea-hexylene carbonate biomaterial layer (segments 3 and 4) on top of base segment 2 appeared essentially flat. The biomaterial was nearly transparent and colorless. Some edges of the biomaterial showed streaks or wear, which were removed using a scalpel.

[0118] Like the PEKK anchor, the titanium anchor also has a central hole (241, 242) of the same size. The hole is substantially filled with poly-urethane-urea-hexamethylene carbonate biomaterial, and the biomaterial is very satisfactorily attached to the titanium base segment 2.

[0119] The titanium base segment 2 has the same dimensions as the PEKK base segment. Since the same mold was used, the elastomeric biomaterial (aliphatic polycarbonate) positioned on the titanium base segment has a diameter of approximately 6 mm and a height of approximately 1 mm. Therefore, the total height of the titanium biomaterial plug implant is approximately 7 mm.

[0120] The top section 4 was provided with pores by drilling holes with an average diameter of 300 μm, resulting in a final porosity of 50 vol.%. The elastic modulus of the non-porous aliphatic poly-urethane-urea-hexamethylene carbonate biomaterial of the top section 4 was 0.9±0.2 MPa according to ASTM D638.

[0121] like Figures 6A to 6C As shown in FIG, the implant 1 can be implanted into the osteochondral defect 8 as described above. In a typical method, a cartilage defect extending into the subchondral bone is drilled ( Figure 6A), and the plug-shaped implant 1 is implanted into the drill hole, as shown Figure 6B As shown in . Due to the relatively high stiffness of the titanium base segment 2, a press fit is not appropriate. Instead, the size of the drilled subchondral bone is slightly larger than the size of the titanium base segment 2. It can be seen that bone grows onto the titanium base segment 2, thereby anchoring the implant 1. The surrounding natural cartilage 5 grows onto the top side 41 of the top segment 4 and generates new cartilage 5a on the top of the implant 1, as shown in Figure 6C As shown in . Figure 6C As shown in FIG, the height 20 of the base section 2 and the height 40 of the porous top section 4 are selected so that when the implant is implanted, the top surface 41 of the implant 1 is located below the top surface 50 of the cartilage 5 present on the osteochondral structure (5, 6), preferably within a distance 51 between 0.1 mm and 1 mm. In this case, this distance is about 0.5 mm. The osteochondral structure (5, 6) includes subchondral bone 6 and a cartilage layer 5 on top of the subchondral bone. A synovial cavity 7 is also typically present.

[0122] Also like Figure 6B and 6C As shown in , the height 20 of the base segment 2 and the height 40 of the porous top segment 4 are selected so that when the implant is implanted, the bottom surface 24 of the top segment 4 (or the top surface 24 of the base segment 2) is roughly flush with the bottom surface 51 of the cartilage layer 5 of the osteochondral structure (5, 6).

[0123] It will be appreciated that those skilled in the art may make many variations and applications of the present invention within the scope of the appended claims.

Claims

1. A non-biodegradable implant in the shape of a plug for replacing and regenerating biological tissue, the non-biodegradable implant comprising a base segment and a porous top segment, the base segment being configured to anchor in bone tissue, the porous top segment being configured to replace cartilage tissue in a mid-deep region of a cartilage layer and to allow cartilage tissue to grow onto and into the porous top segment, thereby regenerating a superficial region of the cartilage layer, wherein the porous top segment comprises a porous thermoplastic elastomer material, wherein the porous thermoplastic elastomer material comprises a linear block copolymer, the linear block copolymer comprising urethane groups and urea groups, and wherein the base segment material comprises one of the following: a biocompatible metal, a ceramic, a mineral, and a non-hydrogel polymer, and combinations thereof, wherein the porous thermoplastic elastomer material further comprises a carbonate group.

2. The implant of claim 1, wherein the porous thermoplastic elastomeric material is substantially free of additional peptide compounds having cartilage regeneration properties.

3. The implant of claim 1, wherein the porous thermoplastic elastomeric material comprises poly-urethane-bisurea-alkylene carbonate.

4. The implant of claim 1, wherein the porous thermoplastic elastomeric material is aliphatic.

5. The implant of claim 1, wherein the porous thermoplastic elastomeric material has an elastic modulus of less than 8 MPa at room temperature.

6. The implant of claim 1, wherein the base segment comprises a core of a non-porous base segment material and a circumferential shell of a porous base segment material, wherein the shell has a thickness less than 10% of the maximum diameter of the base segment.

7. An implant according to claim 1, wherein the base segment extends between a top surface and a bottom surface and comprises a layer of porous base segment material, wherein the layer is adjacent to the top surface and the thickness of the layer is less than 10% of the maximum height of the base segment, and wherein the pores of the base segment material in the layer comprise a biocompatible elastomeric material.

8. The implant of claim 1, wherein the base section material comprises a metal selected from the group consisting of titanium, zirconium, chromium, aluminum, stainless steel, hafnium, tantalum, or molybdenum and alloys thereof, or any combination thereof.

9. The implant of claim 1, wherein the base segment material comprises a ceramic or mineral selected from the group consisting of oxides, nitrides, carbides, and borides, or any combination thereof.

10. The implant of claim 1, wherein the base segment material comprises a polyaryletherketone polymer.

11. The implant of claim 10, comprising a substantially non-porous poly(aryletherketone) polymer having a porosity of less than 20% relative to the total volume of the poly(aryletherketone) polymer.

12. The implant of claim 10 or 11, wherein the base segment comprises a non-porous polyaryletherketone polymer.

13. The implant of claim 1, further comprising a contrast agent or a radiopharmaceutical for medical imaging.

14. The implant of claim 1, wherein the base segment extends between a top surface and a bottom surface, and wherein the top surface of the base segment includes irregularities or a wavy contour.

15. The implant of claim 1, wherein the base segment comprises a centrally located cavity comprising the porous thermoplastic elastomeric material.

16. The implant of claim 1, wherein the base segment comprises an outer surface having an irregular or undulating profile.

17. The implant of claim 1, wherein the height of the base segment and the height of the porous top segment are selected such that when the implant is implanted, the top surface of the implant is below the top surface of the cartilage present on the osteochondral structure.

18. The implant of claim 1, wherein the height of the base segment and the height of the porous top segment are selected such that when the implant is implanted, the bottom surface of the porous top segment is substantially flush with the bottom surface of the cartilage present on the osteochondral structure.

19. The implant of claim 1 comprising a porous top section having a slightly curved top surface having a radius of curvature in the sagittal plane and / or in the medial-lateral plane in the range of 15 mm to 150 mm.

20. The implant of claim 1, wherein the base section material comprises a reinforcement material selected from the group consisting of fibrous or particulate polymers and / or metals.

21. A method for preparing an implant according to any one of the preceding claims 1 to 20, the method comprising: a) providing an assembly in a mold at room temperature, the assembly comprising a base segment comprising a base segment material comprising one of the following: a biocompatible metal, a ceramic, a mineral, and a non-hydrogel polymer, and combinations thereof; and providing particles comprised of a thermoplastic elastomeric material on top of the base segment, wherein the thermoplastic elastomeric material comprises a linear block copolymer comprising urethane groups and urea groups and further comprising carbonate groups; b) closing the mold and heating the assembly to a temperature between 100° C. and 250° C. at a pressure between 1 GPa and 2 GPa such that the thermoplastic elastomer material melts and fuses with the base segment; and c) cooling the assembly to room temperature to solidify the thermoplastic elastomer material and opening the mold; d) providing a top section of the thermoplastic elastomer material with holes before or after opening the mold.

22. The method of claim 21, wherein the thermoplastic elastomeric material is substantially free of additional peptide compounds having cartilage regeneration properties.

23. A method according to claim 21 or 22, wherein after step b), the mould is opened and further particles of the thermoplastic elastomeric material are added to the mould and step b) is repeated.

24. An osteochondral structure comprising an implant according to any one of claims 1 to 20, wherein a top surface of the implant is located below a top surface of a cartilage layer on the osteochondral structure.

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