Polymer composition and method for making medical implants

Abstract

The present invention relates to a polymer composition for making a bone anchoring component of an orthopedic implant for repairing damaged cartilage tissue, the composition comprising a biostable thermoplastic polyurethane and 15-70% by mass of micron-sized inorganic particles, the particles comprising a biocompatible transition metal compound. Studies have found that the relatively rigid thermoplastic polyurethane composition comprising the inorganic particles allows the production of an anchoring component that can be inserted or press-fit into a pre-drilled bone hole to form a firm and lasting connection with bone tissue, the connection being visualized with, for example, X-rays or MRI methods. The polymer composition shows good performance and can provide freedom in designing and manufacturing implants using common techniques such as injection molding. In other aspects, the present invention relates to a method for manufacturing the polymer composition, and using the polymer composition to manufacture an orthopedic implant comprising a bone anchoring component, such as forming an implant using a multi-component injection molding process, the process comprising molding the bone anchoring component with the polymer composition.

Description

Technical Field

[0001] The disclosed invention relates to a polymer composition suitable for making a bone anchoring component of an orthopedic implant, a method for making the polymer composition, and a method for making an orthopedic implant comprising the polymer composition. Background Art

[0002] Orthopedic implants are medical implants used in orthopedic surgery for conditions involving the musculoskeletal system of a person or animal. This system provides contours, stability, and movement for the body and is composed of the body's bones (skeleton), muscles, cartilage, tendons, ligaments, joints, and other connective tissues (tissues that support and hold tissues and organs together). The primary functions of the musculoskeletal system include supporting the body, allowing movement, and protecting vital organs. The joints and musculoskeletal tissues of the human body may be subject to trauma, disease, and degenerative processes over a period of time, which can lead to deterioration or failure of the joints, resulting in severe pain or immobility. Typically, the ability of a joint to provide painless joint movement and to withstand loads depends on the presence of healthy bone, cartilage, and associated musculoskeletal tissue that provide a stable joint. In conjunction with the present disclosure, orthopedic surgery also relates to maintaining movement in the various joints of the human body. Orthopedic implants include devices used in component or total joint replacements, knee and hip prostheses, and osteochondral implants. Examples of orthopedic implants include bone anchors, bone plugs, and bone screws, which are applied to secure implants such as artificial ligaments and tendons, meniscus or labrum replacement devices, spinal implants such as interbody fusion cages, and cartilage replacement devices.

[0003] Cartilage is a smooth connective tissue on the surface of the ends of bones (where the ends of the bones meet to form a joint) that protects and cushions the bones and absorbs forces transmitted throughout the body. Cartilage is an elastic tissue that allows smooth movement of the joints, but it has no direct blood supply and has limited self-repair capabilities in the event of wear or trauma. A frequent and significant cartilage injury that causes pain and / or immobility is damage to the articular cartilage in the knee, i.e., in the joint formed between the femur and the tibia. Long-term such initial local defects—if left untreated—may lead to further degeneration and damage of the cartilage in the joint and may require surgery using an artificial prosthetic joint; such as partial or total knee replacement (UKR / TKR, also known as hemi / total knee arthroplasty or HKA / TKA). However, such total replacement surgery can be problematic because most artificial joints have limited durability and the subsequent revision procedures required are associated with longer operating and hospital stays and may induce complications, especially in elderly patients. In order to delay and possibly even avoid the need for total joint replacement such as TKR, orthopedic implants have been developed to partially replace damaged cartilage; thereby forming a new smooth joint surface at the site of damage. Such implants are often referred to as cartilage plugs.

[0004] Such known cartilage plugs, also known as osteochondral constructs, cartilage replacement devices, or resurfacing implants, are typically made of metal (e.g., titanium). However, the use of metal implants has led to high revision surgery rates, which may be related to the significant difference in mechanical properties (such as stiffness and deformability) between metal and the cartilaginous (undercartilaginous) bone and cartilage tissue. Alternative devices made of natural and / or synthetic materials have been described or proposed in many publications. Cartilage plugs often have a cylindrical or mushroom-like shape and may include at least two parts; a cartilage replacement component and a bone anchoring component. The cartilage replacement component may typically be made of a flexible, elastic, and wear-resistant biocompatible material that mimics some of the properties of natural cartilage, while the bone anchoring component may be made of a more rigid and harder material, including metal.

[0005] One approach is to create plugs from natural materials, including harvesting bone and cartilage from the patient (autologous grafts or autografts), and using tissue from genetically dissimilar donors of the same species (allografts or homografts). For example, in US5782835, devices and methods for implementing this allograft approach are described. However, the use of such grafts carries the risk of infection or disease transmission.

[0006] Alternatively, orthopedic implants such as cartilage plugs can be made of synthetic materials, such as biocompatible polymers that can be biodegradable or biostable. The use of synthetic polymers presents advantages over metal implants because polymers offer a number of properties, can cause less damage to contact tissue, and are non-magnetic; therefore, they are more compatible with medical imaging techniques such as MRI.

[0007] US2008 / 0249632A1 describes a cartilage plug having a stepped shoulder geometry comprising four or more layers, intended to better distribute the load applied to the implanted plug and surrounding tissue and to reduce undesirable movement. The different layers of the plug can be made of the same or different materials, which can be selected from a variety of natural and synthetic materials and can be porous or non-porous.

[0008] US2011 / 0218647A1 discloses a cartilage plug comprising a hydrogel containing a hydrophilic polymer, a fibrous filler, and 40-80% by mass of water. The polymer is preferably cross-linked polyvinyl alcohol. The implant will have a Young's modulus of 0.75-50 MPa, preferably 23-30 MPa, which allows for initial compression to allow placement of the implant in the opening and subsequent expansion for proper assembly.

[0009] US6626945B2 describes a cartilage plug formed from a laminated structure to match the physiological requirements of the repair site. The plug can be cylindrical and formed from three different materials fused or bonded together. In one embodiment, the first layer, closest to the subchondral bone during implantation, is made of a biostable thermoplastic polyurethane (TPU) with a Shore hardness of 75 ShD. The middle layer of the implant is made of TPU with a hardness of 55 ShD, while the third layer, closest to the surface of the cartilage surrounding the implanted plug, is made of a more flexible material (such as TPU with a hardness of 80 ShA) or a thermoplastic hydrogel. This final layer is indicated to exhibit properties similar to those of hyaline cartilage, a type of cartilage found on the outer surface of articulating joints; and a 75 ShD material will have an elastic modulus similar to that of the subchondral bone.

[0010] WO2011 / 098473A1 describes an orthopedic implant, such as a meniscus or intervertebral disc implant, having two or more distinct segments, each comprising a different but chemically related polymeric material; the segments are attached to each other at contact surfaces through interaction between the materials. Preferably, the material is selected from a block copolymer, such as thermoplastic polyurethane (TPU). The implant can be made using a multi-component molding technique. WO2015 / 0135907A1 also describes a meniscus prosthetic device comprising distinct components made from two non-absorbable polymeric materials having a tensile modulus of at most 100 MPa and at least 101 MPa, respectively. TPU materials with Shore hardnesses of 80 ShA and 75 ShD were used in the experiments. It was shown that the material can be a polymer composition comprising TPU and up to 25% by mass of radiopaque filler particles, such as barium sulfate.

[0011] In WO 2007 / 007062 A2 a cartilage repair implant is described in which an elastic layer is combined with an anchoring layer, the anchoring member being made of a bone cement composition comprising an acrylate-based polymer containing calcium ions to promote bone ingrowth.

[0012] US2004 / 0188011A1 discloses a method for manufacturing a prosthetic bearing element, which includes a rigid backing made of carbon fiber reinforced polymer and supporting a soft elastomeric polyurethane bearing lining, in which method, improved bonding of the backing to the bearing lining is achieved by laser welding; that is, by transmitting a laser through a transparent bearing lining to induce thermal fusion at the interface of the lining and the laser-opaque backing.

[0013] US 2009 / 043398 A1 describes a method for producing an articulating surface implant, such as a replacement plug, in which a density, porosity, and / or concentration gradient is created in the implant by subjecting a viscous material to centrifugal forces, such as by spin casting. The viscous material may be a composite material comprising a polymer matrix and particles or fibers dispersed therein; the method produces an article having a concentration gradient of particles or fibers, and therefore a stiffness gradient.

[0014] US2008 / 0081061A1 discloses an orthopedic device comprising a biocompatible polymer-based composite material having ceramic particles dispersed therein, and may further comprise a non-ceramic polymer integrally attached to the composite material. The ceramic material may be in particle or fiber form and, like the polymer, may be selected from a wide range of materials. In one embodiment, the polymer is ultra-high molar mass polyethylene (UHMWPE) and the ceramic comprises hydroxyapatite (HA) particles. This composite part is made by mixing UHMWPE powder with (optionally coated) HA particles and compression molding the mixture.

[0015] Geary et al. describe thermoplastic polyurethanes, such as commercially available Polycarbonate urethane grades with excellent hydrolysis and aging resistance are suitable biostable materials for use in the manufacture of in vivo biomedical devices, such as devices for use in replacing diseased or damaged joints. Incorporating carbon fibers or hydroxyapatite (HA) particles into such polyurethanes via melt compounding can lead to improved mechanical properties of the polymer material. However, it has also been shown that this compounding step enhances polymer degradation, resulting in a significant reduction in the molar mass of the polymer in the polymer composition. In particular, when compounding TPU with HA, the degradation appears to be more prominent than during melt processing steps such as injection molding. Polyurethane compositions based on TPU and HA particles do not show improved tensile properties compared to carbon fibers.

[0016] CN1215890C discloses polyurethane compositions containing 0.1-30% by mass of zinc oxide, titanium dioxide, or zirconium dioxide nanoparticles, wherein 80-99% of the particles are less than 50 nm in size. The solvent-based compositions are prepared by high-speed mixing of the nanoparticles into a polyurethane solution and can be used to coat substrates or cast films. These films or coated substrates are said to exhibit anti-coagulation properties and are suitable for use in medical products that come into contact with blood, such as catheters and vascular prostheses. Summary of the Invention

[0017] Although various synthetic polymer materials for the production of bone anchoring components of orthopedic implants have been proposed or described in the literature, there still appears to be a need for implant materials that allow the production of implants for more permanent replacement of localized cartilage defects, for example in joints, and that can be monitored during and after implantation using common medical techniques.

[0018] The present disclosure provides polymeric materials for use in manufacturing orthopedic implants, such as implants for joint replacement, such as for replacing locally damaged cartilage tissue in a knee joint. More specifically, the present disclosure provides polymeric materials suitable for use in manufacturing bone anchoring components of implants that enable a durable and monitorable connection to bone tissue.

[0019] Aspects and embodiments of the present invention as described below and as characterized in the claims provide a polymer composition suitable for manufacturing a bone anchoring component of an orthopedic implant, the bone anchoring component exhibiting good biocompatibility and mechanical properties compatible with bone tissue, and the implant being able to be monitored during and / or after implantation using medical imaging techniques such as X-ray and MRI; and a method for manufacturing the polymer composition. Thus, one aspect of the present invention is a polymer composition comprising a biocompatible, biostable thermoplastic polyurethane (TPU) and 15-70% by weight of biocompatible inorganic particles containing a transition metal compound, wherein the particle size (D 50 ) is in the range of 0.1-5 μm, more specifically, a polymer composition suitable for manufacturing a bone anchoring component of an orthopedic implant, which implant may be suitable for partial replacement of damaged cartilage tissue in a joint, the polymer composition comprising a biostable thermoplastic polyurethane and 15-70% by mass of inorganic particles containing a biocompatible transition metal compound, wherein the particle size (D 50 ) is in the range of 0.1-5μm.

[0020] Orthopedic implants, such as cartilage replacement plugs, in which the polymer composition is used can have a cylindrical or mushroom-shaped shape and can include at least two parts; the bone anchoring component and the cartilage replacement component, which is typically made of an elastic and wear-resistant biocompatible material. It was found that the relatively rigid anchoring component made of the TPU composition containing the inorganic particles allowed the manufacture of an anchoring component that, for example, can be inserted and / or pressed into a pre-drilled bone hole to form a strong and lasting connection to the bone, and the implant component and its connection to the bone can be visualized or monitored over time using, for example, X-ray or MRI methods. The polymer composition shows favorable crystallization behavior, and the (mechanical) properties of the composition are superior to the (mechanical) properties of similar compositions, for example, based on HA filler particles. Other advantages of using polyurethane compositions include freedom in the design and sizing of implants, and freedom in the manufacture of implants using common techniques such as injection molding, thereby also allowing more complex shapes. In particular, when the implant also comprises a component as a cartilage replacement component made of a thermoplastic material compatible with the polyurethane composition (e.g., a more flexible TPU type), the implant can be manufactured using a two-component injection molding technique; the two-component injection molding technique is a relatively simple method that generally produces sufficient adhesion between the polyurethane composition of the bone anchoring component and the flexible TPU of the cartilage replacement component without the need for an adhesive component.

[0021] In other aspects, the present invention relates to a method of preparing the polymer composition.

[0022] In another aspect, the present invention relates to the use of a polymer composition according to the present disclosure in the manufacture of an orthopedic implant or a bone anchoring component of an orthopedic implant, and a method for manufacturing an orthopedic implant or a bone anchoring component of an orthopedic implant, such as a bolt or a screw, the method comprising the step of injection molding the implant or the bone anchoring component with the polymer composition according to the present disclosure; for example, by forming the implant using a multi-component injection molding process, the process comprising the step of molding the bone anchoring component with the polymer composition. Another aspect relates to an orthopedic implant, the bone anchoring component of which comprises the polymer composition. BRIEF DESCRIPTION OF THE DRAWINGS

[0023] The present invention will be further elucidated by the following illustrative drawings, but is not limited thereto.

[0024] Figure 1 A 2-part cylindrical cartilage plug is schematically depicted.

[0025] 2A and 2B schematically illustrate a mushroom-shaped cartilage plug in two different views.

[0026] Figure 3Schematically depicts a cross-section of a mushroom-shaped plug comprising a bone anchoring member (3a) with two "stems" (3b' and 3b"). Figure 1-3 In the drawings, like numbers represent like elements.

[0027] Figure 4 Shown is the particle size distribution as measured on 3 grades of zirconium oxide.

[0028] Figure 5 Shown are the DSC curves (heating-cooling-reheating) as obtained on unfilled polyurethane and on corresponding polyurethane compositions containing 20%, 40% and 60% by mass of zirconium oxide (SA).

[0029] Figure 6 Bone-to-implant contact scores (BIC, %) are provided as an indicator of osseointegration for implanted plugs M (metal), U (unfilled polyurethane), and I (invention; zirconia-filled polyurethane, BCP-coated) at 6 and 12 months after implantation in goat knees.

[0030] Figure 7 AC show representative photomicrographs of tissue sections taken from plugs M, U and I implanted in goats (6 months later).

[0031] Figure 8 Shown are the results of the Modified Mankin score (MMS) of articulated cartilage quality on the opposing bone surfaces of joints with implanted plugs M or I compared to sham-operated joints (after 6 and 12 months). DETAILED DESCRIPTION

[0032] In the context of the present disclosure, a biocompatible material is biocompatible because it does not produce a toxic, injurious, or immune response when in contact with living tissue. Biodegradable means that the material readily degrades chemically or breaks down into simpler components under physiological conditions or by biological means (e.g., by enzymatic action). Biostable or bioinert means that the material is substantially non-biodegradable under the conditions and duration of intended use in contact with living tissue.

[0033] According to one aspect, the present invention provides a polymer composition comprising a biostable thermoplastic polyurethane and 15-70% by mass of biocompatible inorganic particles containing a transition metal compound, the polymer composition being suitable for use in the manufacture of an orthopedic implant or at least a bone anchoring component thereof, for example a biostable implant suitable for repairing damaged tissue, such as cartilage tissue in a joint.

[0034] Polymer compositions suitable for manufacturing bone anchoring components include biostable thermoplastic polyurethanes (TPUs). TPUs are non-crosslinked polyurethanes with essentially linear polymer chains that are soluble in good solvents, can be melted by increasing temperatures, and resolidify upon subsequent cooling, allowing for melt processing, such as by extrusion or injection molding. TPUs are typically block copolymers (also known as multi-segment copolymers).

[0035] Block copolymers are polymers that contain multiple blocks (also called segments) of polymers (including oligomers) that are chemically different and exhibit different thermal and mechanical properties and different solubility. Typically, the blocks in a block copolymer containing two (or more) types of blocks are referred to as "hard" and "soft" polymer blocks, and such (chemically) different blocks result in microphase separation into domains rich in hard blocks or soft blocks. In the case of biomedical applications, the hard blocks in block copolymers typically contain rigid or high modulus polymers with a melting temperature (T m ) or glass transition temperature (T g ) is above the use temperature, which is typically about 35-40° C. The soft block in the block copolymer typically comprises T g A flexible, low modulus, amorphous polymer below 25°C, preferably below 0°C. For most mechanical properties, such as T m and T g Thermal parameters are typically measured on dry samples using well-known techniques such as DSC or DMA. In such phase-separated block copolymers, the hard segments act as physical (and non-permanent, thermoreversible) crosslinks of the flexible soft segments, resulting in materials with properties ranging from fairly rigid and stiff to flexible and elastic, depending on, for example, the ratio and type of hard to soft blocks. Depending on the type and content of the hard blocks, polyurethane block copolymers can exhibit good stability and elasticity over the desired temperature range without the need for chemical crosslinking; and can generally be processed as thermoplastics. The term thermoplastic polyurethanes essentially refers to a family of polymers having a backbone comprising a reaction product of at least three main components; the at least three main components being a diisocyanate, a diol chain extender, and a polymer diol or marcoglycol. Optionally, a monofunctional compound may be used as an additional component that acts as a chain terminator and forms (non-reactive) end groups. In an embodiment, the backbone of the polyurethane or TPU used in the present invention is essentially linear.

[0036] In an embodiment, the TPU comprises hard blocks comprising urethane and optional urea groups in repeating units, which groups are generated by the reaction of a diisocyanate with a diol and optionally a diamine, which serve as respective chain extenders.

[0037] Suitable diisocyanates include aromatic, aliphatic, and cycloaliphatic compounds having an average of 1.9 to 2.1 isocyanate groups per molecule. In one embodiment, the diisocyanate includes 4,4'-diphenylmethane diisocyanate (MDI), 2,4-toluene diisocyanate, 2,6-toluene diisocyanate (TDI), 1,4-phenylene diisocyanate, hexamethylene diisocyanate (HDI), tetramethylene-1,4-diisocyanate, cyclohexane-1,4-diisocyanate, dicyclohexylmethane-4,4'-diisocyanate (CHMDI), isophorone diisocyanate (IPDI), or mixtures thereof. In one embodiment, the diisocyanate includes hexamethylene diisocyanate, dicyclohexylmethane 4,4'-diisocyanate, isophorone diisocyanate, or mixtures thereof. In one embodiment, the diisocyanate is composed of the following substances: hexamethylene diisocyanate, dicyclohexylmethane 4,4'-diisocyanate, isophorone diisocyanate, or a mixture thereof. In another embodiment, the diisocyanate includes 4,4'-diphenylmethane diisocyanate, 2,4-toluene diisocyanate, 2,6-toluene diisocyanate, or 1,4-phenylene diisocyanate. In one embodiment, the diisocyanate is composed of the following substances: 4,4'-diphenylmethane diisocyanate, 2,4-toluene diisocyanate, 2,6-toluene diisocyanate, 1,4-phenylene diisocyanate, or a mixture thereof. In one embodiment, the molar mass of the diisocyanate is 100 g / mol to 500 g / mol. In one embodiment, the molar mass of the diisocyanate is 150 g / mol to 260 g / mol.

[0038] Chain extenders are typically low-molar-mass aliphatic compounds having two or more, preferably two, hydroxyl or amine groups. Difunctional chain extenders produce linear, typically thermoplastic polymers, while polyfunctional chain extenders and / or isocyanates result in branched or crosslinked products. In one embodiment, the molar mass of the difunctional chain extender is at least 60 g / mol, at least 70 g / mol, at least 80 g / mol, at least 90 g / mol, or at least 100 g / mol. In one embodiment, the molar mass of the chain extender is at most 500 g / mol, at most 400 g / mol, at most 300 g / mol, at most 200 g / mol, or at most 150 g / mol. In one embodiment, the chain extender comprises ethylene glycol, diethylene glycol, propylene glycol, dipropylene glycol, 1,3-propanediol, 1,4-butanediol, 1,5-pentanediol, 1,6-hexanediol, or 1,8-octanediol; and / or such corresponding diamines. In an embodiment, the thermoplastic polyurethane comprises only a diol chain extender.

[0039] In other embodiments, the TPU comprises hard blocks having both urethane and urea linkages. This has the advantage of enhancing the interactions between the hard blocks, leading to higher softening temperatures and / or allowing for higher levels of soft blocks, resulting in block copolymers that exhibit enhanced flexibility and elasticity, as well as excellent flex life or fatigue resistance. Depending on the diol / diamine ratio, the polyurethane chains may exhibit such strong interactions that thermal degradation may occur at melt processing temperatures to the point where solution processing is preferred for optimal performance. Commercially available examples of such polyurethanes containing both urethane and urea linkages (also known as polyurethane ureas) include Product (available from, for example, DSM Biomedical BV, Sittard-Geleen NL).

[0040] In a further embodiment, the thermoplastic polyurethane comprises as a soft block a polymer derived from at least one aliphatic polymer diol or polyol selected from the group consisting of polyethers, polyesters, polyacrylates, polyolefins, and polysiloxanes (also known as silicones); the polymer being difunctional having hydroxyl (or amine) end groups. Such polymer diols used for the soft block are understood to include oligomers, homopolymers, and copolymers, and polyesters are considered to include polycarbonates. Generally known polyurethane block copolymers and methods for preparing these copolymers are described, inter alia, in US Pat. Nos. 4,739,013, 4,810,749, 5,133,742, and 5,229,431.

[0041] In embodiments of the present disclosure, the thermoplastic polyurethane comprises at least one polymer diol selected from aliphatic polyester diols, aliphatic polycarbonate diols, aliphatic polyether diols, poly(isobutylene) diols, and polysiloxane diols in the soft block. As with chain extenders, some amine-functional soft blocks may also be used to create additional urea linkages. Such polyurethane block copolymers have demonstrated biocompatibility and biostability in the human body.

[0042] The mechanical properties and other characteristics of thermoplastic polyurethanes can be adjusted by changing the chemical composition and / or the molar mass of the block. The molar mass of the hard block of the thermoplastic polyurethane included in the composition of the present invention can be about 160Da to 10,000Da, more preferably about 200Da to 2000Da. The molar mass of the soft segment can be generally about 200Da to 100,000Da, preferably at least about 400Da, 600Da, 800Da or 1000Da and at most about 10,000Da, 7500Da, 5000Da, 4000Da, 3000Da or 2500Da. In the context of the present disclosure, the molar mass of the polymer and the oligomer forming the block in the polymer refers to the number-average molar mass (M n), for example, from GPC measurements. The ratio of soft blocks to hard blocks can be selected to result in a certain stiffness or hardness of the polyurethane. Typically, the hardness of the polyurethane, as measured by the Shore durometer hardness test using the A or D scale, can be 40 ShA up to 90 ShD, typically representing a flexural modulus range of about 10 MPa to 2000 MPa. In embodiments, the thermoplastic polyurethane included in the composition has a hardness of 45 ShA to 90 ShA, preferably at least 50 ShA, 55 ShA, or 60 ShA. An advantage of using a relatively low hardness TPU can be higher toughness of the resulting composition further comprising a transition metal compound such as zirconium oxide. In other embodiments, the TPU in the polymer composition has a hardness of 90 ShA to 90 ShD, as this results in a higher stiffness of the composition. In further embodiments, the TPU has a hardness of at least 40 ShD, 50 ShD, or 60 ShD and at most 85 ShD or 80 ShD, to achieve a good balance between stiffness and processing behavior.

[0043] In another embodiment of the present invention, the thermoplastic polyurethane comprises an aliphatic polyether, an aliphatic polyester or an aliphatic polycarbonate, more particularly an aliphatic polycarbonate, in the soft block. The composition of the soft block is preferably selected so as to produce a T g Substantially amorphous oligomers or polymers below 10°C, 0°C, preferably below -10°C. Suitable aliphatic polyethers include poly(propylene oxide) glycol, poly(tetramethylene oxide) glycol, and copolymers thereof. Suitable aliphatic polyesters are typically made from at least one aliphatic dicarboxylic acid and at least one aliphatic diol. Aliphatic polycarbonate diols are based on aliphatic diols similar to those used for polyester diols and can be synthesized via different routes known in the art. Suitable examples include poly(hexamethylene carbonate) glycol and poly(tetramethylene carbonate) glycol. Such polycarbonate-based TPUs exhibit good biocompatibility, such as blood compatibility, and enhanced biostability. In one embodiment, the soft block of the TPU is based on poly(hexamethylene carbonate) glycol, poly(tetramethylene carbonate) glycol, or a mixture thereof. In a preferred embodiment, the soft block of the TPU comprises or is substantially based on poly(hexamethylene carbonate) glycol.

[0044] In another embodiment, the soft block of the TPU comprises a polysiloxane diol, such as poly(dimethylsiloxane) diol, polycarbonate diol, or poly(tetramethylene oxide) diol. In one embodiment, the soft block is based on polysiloxane diol, polycarbonate diol, poly(tetramethylene oxide) diol, or a mixture thereof. In one embodiment, the soft block comprises a mixture of two or more of polysiloxane diol, polycarbonate diol, or poly(tetramethylene oxide) diol. Such a mixture allows the manufacture of a biocompatible polyurethane exhibiting a combination of enhanced hydrolytic stability and high toughness. In one embodiment, the soft block is based on a mixture of two or more of polysiloxane diol, polycarbonate diol, or poly(tetramethylene oxide) diol. In one embodiment, the soft block comprises polysiloxane diol and one or more of polycarbonate diol and poly(tetramethylene oxide) diol. In one embodiment, the soft block is based on polysiloxane diol and one or more of polycarbonate diol and poly(tetramethylene oxide) diol.

[0045] In one embodiment, the soft block further comprises C2-C 16 Fluoroalkyl diol or C2-C 16 In one embodiment, the soft block in the polyurethane backbone comprises 1H,1H,4H,4H-perfluoro-1,4-butanediol, 1H,1H,5H,5H-perfluoro-1,5-pentanediol, 1H,1H,6H,6H-perfluoro-1,6-hexanediol, 1H,1H,8H,8H-perfluoro-1,8-octanediol, 1H,1H,9H,9H-perfluoro-1,9-nonanediol, 1 The residue of 1H,1H,10H,10H-perfluoro-1,10-decanediol, 1H,1H,12H,12H-perfluoro-1,12-dodecanediol, 1H,1H,8H,8H-perfluoro-3,6-dioxaoctane-1,8-diol, 1H,1H,11H,11H-perfluoro-3,6,9-trioxaundecane-1,11-diol, fluorinated triethylene glycol, or fluorinated tetraethylene glycol.

[0046] In one embodiment, C2-C 16 Fluoroalkyl diol or C2-C 16 The Mn of the fluoroalkyl ether diol is at least 150 g / mol, at least 250 g / mol, or at least 500 g / mol. In one embodiment, the molar mass of the fluoroalkyl diol or fluoroalkyl ether diol is at most 1500 g / mol, at most 1000 g / mol, or at most 850 g / mol. In one embodiment, based on the total mass of the polyurethane, the C2-C 16 Fluoroalkyl diol or C2-C 16The fluoroalkyl ether diol is present in an amount of at least 1 mass %, at least 2 mass %, or at least 5 mass %. In one embodiment, based on the total mass of the polyurethane, C2-C 16 Fluoroalkyl diol or C2-C 16 The fluoroalkyl ether diol is present in an amount of up to 15 mass %, up to 10 mass %, or up to 8 mass %.

[0047] In an embodiment, the polyurethane may contain one or more hydrophobic or hydrophilic end groups. An end group is typically a non-reactive moiety present at the end of a molecule. In one embodiment, the polyurethane contains an end group at each end of the main chain; that is, it has an average of about 2 end groups. In one embodiment, the end group is a straight chain compound. In another embodiment, the end group is a branched chain. The end group can be formed by reacting a terminal isocyanate group present during or after the formation of the polymer backbone with a co-reactive group on a monofunctional compound (also known as a chain terminator). For example, the formulation for forming the polyurethane can contain a diisocyanate, a polymeric aliphatic diol, a chain extender, and a monofunctional alcohol or amine; such as 1-octanol or octylamine, to form a C8 alkyl end group.

[0048] In an embodiment, the terminal group is a hydrophobic terminal group, for example comprising a C2-C 20 Alkyl, C2-C 16 Fluoroalkyl, C2-C 16 In one embodiment, the hydrophobic poly(alkylene oxide) is poly(propylene oxide), poly(tetramethylene oxide) or a copolymer thereof. In one embodiment, the hydrophobic end group is a polysiloxane such as poly(dimethylsiloxane). In one embodiment, the end group comprises C2-C 20 Alkyl, C2-C 16 Fluoroalkyl, C2-C 16 Fluoroalkyl ethers, or hydrophobic poly(alkylene oxides). Such end groups can be formed with monofunctional alcohols including methanol or the aforementioned amines. It has been found that polyurethane elastomers with such hydrophobic end groups positively affect the properties of polyurethanes and their interactions with other materials, including other polymers such as polyolefins and body tissues and fluids such as blood.

[0049] In one embodiment, the hydrophobic end groups include C2-C 16 Fluoroalkyl or C2-C 16 Fluoroalkyl ethers. Such end groups can be used with C2-C 16 Fluoroalkyl or C2-C 16In one embodiment, the terminal groups are formed from 1H,1H-perfluoro-3,6-dioxaheptan-1-ol, 1H,1H-perfluoro-1-pentanol, 1H,1H-perfluoro-1-hexanol, 1H,1H-perfluoro-3,6,9-trioxadecan-1-ol, 1H,1H-perfluoro-1-heptanol, 1H,1H-perfluoro-3,6 The present invention relates to a 1H,1H-perfluoro-2-dioxadecan-1-ol, 1H,1H-perfluoro-1-octanol, 1H,1H-perfluoro-1-nonanol, 1H,1H-perfluoro-3,6,9-trioxadecan-1-ol, 1H,1H-perfluoro-1-decanol, 1H,1H-perfluoro-1-undecanol, 1H,1H-perfluoro-1-dodecanol, 1H,1H-perfluoro-1-myristyl alcohol, or 1H,1H-perfluoro-1-palmitol.

[0050] In another embodiment, the end groups are hydrophilic end groups formed from hydrophilic monofunctional alcohol or amine compounds. Such compounds are typically soluble in water and may exhibit surface activity, such as polyethylene oxide or sulfonate-functional compounds. Such hydrophilic end groups can influence interactions or adhesion with other materials, for example, enhancing the dispersion of certain inorganic filler particles.

[0051] In another embodiment, the polyurethane comprises a mixture of hydrophobic end groups and hydrophilic end groups. Such modifications allow control of the hydrophobicity to hydrophilicity balance of the polymer. A general advantage of using TPUs with end groups is the ability to modify and control the properties of polymers and polymer compositions without incorporating additives that could introduce potential problems of migration from polymer compositions and implants.

[0052] In one embodiment, the end groups are monomeric and have a molar mass of 200 g / mol or more, 300 g / mol or more, or 500 g / mol or more; and 1,000 g / mol or less, or 800 g / mol or less.

[0053] In other embodiments, the end groups are polymeric and have a molar mass of 10,000 g / mol or less, 8,000 g / mol or less, 6,000 g / mol or less, or 4,000 g / mol or less. In one embodiment, the end groups are polymeric and have a molar mass of 500 g / mol or more, 1,000 g / mol or more, or 2,000 g / mol or more.

[0054] In one embodiment, based on the total mass of polyurethane, the end group is present in an amount of at least 0.1 mass %, at least 0.2 mass %, at least 0.3 mass % or at least 0.5 mass %. In one embodiment, based on the total mass of polyurethane, the end group is present in an amount of at most 3 mass %, at most 2 mass % or at most 1 mass %. In one embodiment, based on the total mass of polyurethane, the end group is present in an amount of at least 0.1 mass %, at least 0.2 mass %, at least 0.3 mass % or at least 0.5 mass %; and at most 3 mass %, at most 2 mass % or at most 1 mass % is present.

[0055] In one embodiment, the polyurethane comprises less than 0.1 mass % end groups based on the total weight of the polyurethane. In one embodiment, the polyurethane is substantially free of end groups. In one embodiment, the polyurethane is free of end groups.

[0056] The hard blocks in TPU are typically based on aromatic diisocyanates, such as toluene diisocyanate (TDI) or methylene diphenyl diisocyanate (MDI); and low molar mass aliphatic diols, such as 1,4-butanediol. Polyether and polycarbonate polyurethanes may be suitable for biomedical applications given their flexibility, strength, biostability, biocompatibility, and wear resistance. TPUs containing a combination of polyether and polysiloxane or a combination of polycarbonate and polysiloxane in the soft blocks exhibit a unique combination of properties and can be advantageously used as polyurethanes in polymer compositions. Commercially available examples of such polymers include and Product (available from DSM Biomedical BV, Sittard-Geleen NL).

[0057] In further embodiments, the TPU can be a blend of two or more polyurethanes, such as a blend of two biostable, biocompatible TPU grades having at least one different hardness (e.g., a combination of a 50-80 ShA grade and a 70-85 ShD grade). In such a TPU blend, the soft block types of the two polymers may also differ. An example is a blend of a low-hardness TPU containing polysiloxane in the soft block with a rigid TPU containing polycarbonate in the soft block. Such TPU blends can provide a favorable combination of hardness and toughness.

[0058] In other embodiments, in addition to, for example, catalyst residues, the TPU may also contain one or more conventional biocompatible additives, i.e., additives that allow for the targeted use of the TPU and polymer composition in medical implants. Examples of additives include stabilizers, antioxidants, processing aids, lubricants, surfactants, antistatic agents, colorants, and the like. The additives may be present in typical effective amounts known in the art, such as 0.01-5% by mass, preferably 0.1-2% by mass, based on the amount of polyurethane and additives. In another embodiment, the TPU is substantially free of additives.

[0059] In an embodiment, the polymer composition consists essentially of or consists of: a) 30-85 mass % of a biostable thermoplastic polyurethane; and b) 15-70 mass % of biocompatible particles comprising a transition metal compound, wherein the sum of a) and b) is 100 mass %.

[0060] In other embodiments, the polymer composition is composed of: a) 20-85% by mass of a biostable thermoplastic polyurethane (as described above); b) 15-70% by mass of biocompatible particles comprising a transition metal compound; and c) 0-10% by mass of other compounds, wherein the sum of a)-c) is 100% by mass. Examples of other compounds include bioactive compounds, such as antimicrobial or anti-inflammatory agents, active compounds or drugs that reduce pain or improve healing or bone formation; additives commonly used in polymer composite compositions, such as stabilizers, dispersants, or other compounds added intentionally, and residual amounts of solvents that may have been used to prepare the composition, including solvents for cleaning equipment used. In embodiments, the composition and parts made from the composition contain up to 5%, 4%, 3%, or 2% by mass of other compounds, and up to 1000 ppm of solvent, preferably up to 800 ppm, 600 ppm, 500 ppm, or 400 ppm of solvent.

[0061] The polymer composition includes biocompatible inorganic particles composed of radiopaque transition metal compounds. Within the scope of this application, transition metal is defined as one of the elements indicated in the periodic table, including the lanthanides. Suitable transition metal compounds are oxides (or other salts) of one or more transition metals that combine inertness to TPU and relevant processing conditions, biocompatibility, radiation intensity, and MRI compatibility. Radiopacity or radiodensity, as well as radiopaque or radiodense, means that the inorganic particles inhibit the passage of radio waves and X-ray portions of the electromagnetic spectrum (or absorb radio waves and X-ray portions of the electromagnetic spectrum) to a certain extent so that sufficient contrast with natural tissue is visible using medical X-ray imaging techniques (also known as radiographic imaging). Factors that contribute to radiopacity are electron density and atomic number. MRI compatibility or compatible means that the material has no known hazards in all MRI environments; that is, it is non-conductive, non-metallic, and non-magnetic. Such radiopaque and MRI-compatible particles can show contrast with natural tissue on MRI images.

[0062] In an embodiment, the polymer composition comprises biocompatible transition metal salt particles, or preferably biocompatible transition metal oxide particles, as inorganic particles.

[0063] In an embodiment, the polymer composition comprises biocompatible inorganic particles comprising a salt, preferably an oxide of at least one of titanium (Ti), zinc (Zn), yttrium (Y), zirconium (Zr), lanthanum (La), ytterbium (Yb), hafnium (Hf), and tantalum (Ta). In other embodiments, the polymer composition comprises inorganic particles comprising a salt, preferably an oxide of at least one of Ti, Zn, Y, Zr, and Ta. Alternatively, the polymer composition comprises inorganic particles consisting essentially of one or more of the salts or oxides. In further embodiments, the polymer composition comprises inorganic particles consisting essentially of or consisting of a salt, preferably an oxide of at least one of Ti, Zn, and Zr.

[0064] In an embodiment, the polymer composition comprises inorganic particles comprising titanium oxide, preferably the polymer composition comprises particles consisting essentially of or consisting of titanium dioxide, also known as titania or TiO. Titanium dioxide exists in different mineral forms, such as rutile and anatase, and is primarily used as a white pigment in paints, plastics, food, toothpaste, and pills.

[0065] In an embodiment, the polymer composition comprises particles comprising zinc oxide, preferably the polymer composition comprises particles consisting essentially of or consisting of zinc oxide (ZnO). Naturally occurring zinc oxide typically contains many impurities and is therefore typically synthesized from metallic zinc. Pure zinc oxide is a white powder used in many different applications, including paints, as a filler in plastics, rubber, and cement, in ceramics, and in dental and skin care products (for its antibacterial properties).

[0066] In an embodiment, the polymer composition comprises particles comprising zirconium oxide, preferably the polymer composition comprises particles comprising zirconia. In an embodiment, the inorganic particles consist essentially of or consist of zirconia. Zirconia, also known as zirconium dioxide or ZrO2, is a white crystalline zirconium oxide.

[0067] Zirconia's primary use is in the production of hard ceramics, for example, through high-temperature sintering. Within the biomedical field, this ceramic is often used in dentistry as crowns and bridges. Other applications include protective optical and thermal barrier coatings, ceramic cutting tools, and diamond simulants in jewelry. Unlike many other mineral particles, zirconium oxide is rarely used as a filler or reinforcement in thermoplastic polymer compositions (also known as polymer compounds).

[0068] Zirconia itself is chemically stable, but during injection molding of the composition of the present invention or during use of an implant made from the composition, it may undergo phase changes at high temperatures, i.e., phase changes at temperatures much higher than during compounding with thermoplastic polyurethane. Commercially available zirconium oxide grades may contain other elements as dopants to thermally stabilize certain phases; for example, MgO, Y2O3, CaO, or Ce2O3 may have been added in amounts ranging from 1 mol% to more than 10 mol%. In addition, zirconium oxide grades may contain small amounts of elements such as Hf, Al, Si, Fe, Na. In the context of the present disclosure, zirconium oxide is therefore understood to include essentially pure ZrO2 as well as mixed oxides comprising ZrO2 and up to about 20% by mass, preferably up to 15%, 10%, or 5% by mass, of other inorganic oxides (as described above).

[0069] Polymer compositions comprising polyurethane and particles such as titanium dioxide, zirconium oxide, and / or zinc oxide are radiopaque and can therefore be distinguished from other materials and tissues using medical X-ray techniques. The particles are non-magnetic, and their addition to polyurethane also enables imaging of parts made from the compositions of the present invention using MRI techniques, which is not possible with metal implants commonly used in orthopedic surgery. Thus, implants made from the polymer compositions of the present invention can be visualized using common medical imaging techniques such as X-rays and MRI; this enables monitoring during surgery to correctly position the implant at the targeted implantation site, as well as postoperative examination of its position relative to surrounding tissue.

[0070] The particle size of the inorganic particles in the polymer composition is generally in the range of 0.1-5 μm. In the present disclosure, the particle size of the inorganic particles is D 50 The particle size is the median diameter or the median of the particle size distribution, as measured by light diffraction according to ISO 13320:2009, for example with a Malvern Mastersizer 2000. The particle size relates to the size of the particles dispersed in water, which may differ from the particle size distribution in the polymer composition, since the primary particles may not deagglomerate or disperse in the same way as during polymer mixing.

[0071] The inorganic particles in the polymer composition can have different types of particle shapes and can be regular or irregular in form. The particle shapes can range from almost spherical to more elongated or flat shapes; such as cigar-shaped, flake-shaped, needle-shaped or fibrous shapes; wherein the cross-section can be circular, oval, triangular, rectangular, square, and the average aspect ratio is 1 to greater than 10. An advantage of substantially round particles can be the isotropic properties of the composition, while an elongated shape can result in a composition with better mechanical properties, but this may depend on the orientation of the particles. In an embodiment, the polymer composition comprises a mixture of particles of different shapes.

[0072] In an embodiment, the size D of the particles 50 The particle size of the polyurethane particles is preferably at least 0.10 μm, 0.15 μm, 0.20 μm, 0.25 μm or 0.30 μm, as the handling properties, such as flow and dosing, generally improve with particle size. The particle size is preferably at most 5.0 μm, 4.5 μm, 4.0 μm, 3.0 μm, 2.5 μm or 2.0 μm, for the purpose of dispersion in the polyurethane and the mechanical properties of the polymer composition. The particles in the composition may also be a mixture of different sizes, such as a mixture of particles with a size at the lower end of the range and particles with a size at the upper end of the range, for example, to optimize density and mechanical properties, such as stiffness or toughness of the composition.

[0073] In an embodiment, the polymer composition comprises a biostable thermoplastic polyurethane and 15-70% by mass of inorganic particles containing transition metal compounds (preferably zirconium oxide) particles. A relatively high amount of such particles will enhance the rigidity (e.g., tensile modulus) and radiopacity of the parts made from the composition, but may degrade the ductility and toughness of the composition. In addition, during compounding to prepare the composition, the polymer degradation caused by shear may be enhanced due to a large amount of particles. In an embodiment, the composition therefore comprises at least 20% by mass, 25% by mass, 30% by mass, or 35% by mass and at most 68% by mass, 66% by mass, 64% by mass, 62% by mass, 60% by mass, 58% by mass, 56% by mass, 54% by mass, 52% by mass, 50% by mass, 48% by mass, or 46% by mass of inorganic particles.

[0074] The polymer composition may also include a combination of different particles, such as zirconium oxide particles in combination with one or more other biocompatible inorganic particles, including other transition metal compound particles as defined above (e.g., titanium dioxide and zinc oxide) and / or natural mineral particles such as clay, mica, talc, etc. The other inorganic particles generally have a similar particle size range and shape as the transition metal compound or oxide particles; but may also have different sizes and different particle shapes to produce, for example, a combination of spherical and needle-shaped or fibrous particles. The other particles may be (bio)inert, such as zirconium oxide, or exhibit bioactivity, such as osteoconductivity, such as bioglass or other silicated bioceramics.

[0075] In an embodiment, at most 25 mass % of the total amount of transition metal compound particles such as zirconium oxide and other inorganic particles present in the polymer composition is formed by other inorganic particles, preferably at most 20 mass %, 15 mass %, 10 mass % or 5 mass % is formed by other inorganic particles.

[0076] In further embodiments, the polymer composition is substantially free of calcium phosphate-based particles, such as hydroxyapatite, in view of potential enhanced degradation as reported in the literature and as observed when preparing such polyurethane compositions (see Experimental). More generally, in embodiments, the composition is substantially free of inorganic particles or other compounds that can initiate or enhance (hydrolytic) degradation of the polyurethane because such particles or compounds contain reactive groups that promote such degradation or have hygroscopicity such that the particles or compounds cannot be properly dried to a sufficiently low moisture content, for example, below 250 ppm, 150 ppm or 100 ppm.

[0077] In other aspects, the present invention provides a method for preparing a polymer composition as described above including all variations and preferred embodiments and combinations thereof, comprising the steps of:

[0078] Providing biostable thermoplastic polyurethanes with a moisture content of up to 300 ppm;

[0079] Providing inorganic particles comprising a biocompatible transition metal compound and having a moisture content of at most 250 ppm;

[0080] optionally subjecting the inorganic particles to mechanical or chemical treatment;

[0081] optionally providing other compounds with a moisture content of at most 250 ppm; and

[0082] Mixing polyurethane, inorganic particles and other compounds.

[0083] The method for preparing the polymer composition can use different mixing equipment and processes, which are known to the skilled person; for example, using a solvent-assisted mixing process or a melt mixing process. Typically, the polyurethane polymer and any transition metal compound, zirconium oxide or other compound to be added, such as inorganic particles and / or additives, are thoroughly dried using conventional methods before mixing. Typically, the polyurethane (including all variants and preferred embodiments as described above) is dried at a temperature below its softening point or melting point for several hours, for example, 4 to 30 hours, to obtain a moisture content of at most 300 ppm, preferably at most 250 ppm, 200 ppm or 150 ppm. The inorganic particles (including all variants and preferred embodiments as described above) can be dried at a higher temperature for a longer time, for example, at 100-200°C for 20-40 hours to obtain a moisture content of at most 250 ppm, preferably at most 150 ppm or 100 ppm.

[0084] In an embodiment, the method comprises mechanically or chemically (preliminary) processing (dried) inorganic particles, to enhance the step of the characteristic of the resulting polymer composition. In an embodiment, mechanically processing comprises grinding or abrasive particles, optionally grinding or abrasive particles in the presence of an auxiliary component. Such auxiliary components can be low viscosity or viscous liquids, dispersion aids and / or polymers, each of which is biocompatible and compatible with polyurethane. Such processing can be promoted with ultrasonic treatment, to decompose the aggregate and / or enhance dispersion of particles. Such mechanical treatment steps can produce powders, dispersions, pastes or solid compositions comprising inorganic particles, and the use of the inorganic particles can result in the raising of the dispersion level of particles in polyurethane in subsequent mixing steps. Treated inorganic particles can be dried to reach the required moisture level of 250ppm at the most.

[0085] In another embodiment, the method includes chemically treating the (dried) inorganic particles to change the type and / or number of functional groups at the surface of the particles; and the steps of dispersibility of the particles in the polyurethane and / or interfacial adhesion of the particles to the polyurethane. The treatment may include corona treatment, plasma treatment and / or wet chemical treatment. Thus, as known to those skilled in the art, corona or plasma treatment can modify the surface of the particles, but it is also possible to combine it with wet chemical modification. In wet chemical treatment, the particles are usually first dispersed in a suitable solvent, and then the reagent is added, usually in an amount of 5-300% by mass relative to the particles. Ultrasonic treatment can be used to promote dispersion to break up particle aggregates. Suitable reagents are, for example, organic compounds with alkyl, amine, carboxyl or peroxide functional groups, or silane compounds. Examples include carbon dioxide, oxygen, unsaturated hydrocarbons, alkylamines, carboxylic acids and various amine and / or alkoxy functional silanes, such as 3-aminopropyltriethoxysilane. Such compounds are also referred to as coupling agents in the art. In view of the targeted medical use of the polymer composition, it is preferred to omit the addition of other components such as catalysts, and to substantially completely remove unreacted reagents and solvents after pretreatment, and to bring the moisture level to at most 250 ppm. Based on common sense and optionally some routine experiments, the skilled person will be able to select the reagents and appropriate treatment conditions.

[0086] In embodiments, other compounds as described above for the polymer composition may be provided having a moisture content of up to 250 ppm.Examples of other compounds include bioactive compounds, stabilizers, and compounds that aid in dispersing inorganic particles in the TPU during mixing.

[0087] In an embodiment, a method for preparing a polymer composition comprising a biostable thermoplastic polyurethane and 15-70% by mass of inorganic particles comprises adding dried particles or a dried masterbatch comprising such particles during the synthesis of the polyurethane. For example, such addition can be accomplished by mixing with, for example, a liquid starting chemical before a one-step polymerization reaction, or by mixing the particles with a prepolymer before or during a two-step polymerization step. In other embodiments, a method for preparing a polymer composition comprising a biostable thermoplastic polyurethane and inorganic particles comprises providing a polymerized polyurethane and using a solvent-assisted mixing step; for example, a solution of the dried polyurethane in a good solvent for the polymer (such as THF) is first prepared, followed by mixing with the dried inorganic particles or with inorganic particles pre-dispersed in a liquid, preferably the same good solvent. In an embodiment, a biocompatible dispersant may optionally be added to help evenly disperse the inorganic particles. Depending on the concentration of the polyurethane and the amount of particles in the mixture, a liquid dispersion or a pasty mixture may be obtained. In a subsequent step, the solvent may be removed by known methods such as evaporation, preferably at elevated temperature and / or reduced pressure. The resulting solidified polymer composition can then be brought into a form suitable for use in a shaping step such as compression or injection molding, for example by cutting or grinding. Advantages of this solvent-assisted mixing process include relatively low shear forces and low temperatures, i.e., a lower risk of polymer degradation, and the relatively small scale on which the process can be operated.

[0088] In another embodiment, the method for preparing the polymer composition comprises melt mixing the components, also known as compounding, at a temperature above the softening or melting point of the polyurethane using known equipment, such as a batch mixer or a continuous mixer (e.g., a single-screw extruder or a twin-screw extruder). Optionally, a biocompatible wetting agent or dispersant may be added before or during melt mixing to enhance dispersion of the particles. Prior to melt mixing, the polyurethane and particles are thoroughly dried to minimize hydrolytic degradation during melting and mixing, as described above. For similar reasons, the melt mixing equipment and mixing conditions are selected so as to keep the temperature of the composition as low as possible. In an embodiment, the polymer composition is prepared by mixing the dry components on a twin-screw extruder, applying conditions such as screw configuration, screw speed, throughput rate, and temperature setting that result in sufficient shear or torque to properly disperse the inorganic particles in the polyurethane while minimizing overheating and polymer chain scission or molar mass reduction. In an embodiment, the temperature of the extruder barrel is set to at most 210°C, preferably about 205°C or 200°C. It was observed that melt mixing the polyurethane with the dry particles under such conditions resulted in better dispersion and significantly less polymer degradation than that found for the same thermoplastic polyurethane with other mineral fillers such as hydroxyapatite and bismuth oxide.

[0089] In embodiments, the mass-average molar mass (or weight-average molecular weight), Mw, of the TPU in a polymer composition from which the bone anchoring component of an orthopedic implant can be made is at least 70 kDa. As described in the experimental section, molar mass and molar mass distribution are typically measured using GPC methods. Note that in this disclosure, the ISO terms 'mass' and 'molar mass' are generally used rather than the still commonly used terms 'weight' and 'molecular weight'. It has been found that if the TPU in the composition has such a minimum molar mass, components made from the polymer composition (e.g., made by (melt) mixing the TPU with inorganic particles) exhibit certain minimum desired properties, such as a certain strength or elongation at break. In preferred embodiments, the molar mass, Mw, of the TPU in the composition is at least 75 kDa, 80 kDa, 85 kDa, 90 kDa, 95 kDa, or 100 kDa. The molar mass, Mw, of the TPU used to make the composition also meets these minimum values, but is typically higher; it is also preferably not so high that its melt viscosity would hinder processing and mixing with the inorganic oxide particles; this also limits the molar mass of the TPU in the resulting composition. In an embodiment, the TPU in the prepared polymer composition has a molar mass Mw of at most 400 kDa, 300 kDa, 250 kDa or 200 kDa to produce a balanced combination of processability and mechanical properties.

[0090] In an embodiment, the polymer composition has an E modulus of at least 800 MPa, preferably at least 850 MPa or 900 MPa, and at most 3000 MPa, preferably at most 2500 MPa, 2000 MPa, 1800 MPa, or 1600 MPa, as measured on a dry, molded sample at 20°C. Alternatively, the polymer composition is characterized by an E modulus of at least 200 MPa or at least 225 MPa and at most 700 MPa, preferably at most 600 MPa, 550 MPa, 500 MPa, or 450 MPa, as measured on a wet-conditioned sample (in water at 37°C), which better simulates the biological conditions during targeted use of the polymer composition in an implant. Without wishing to be bound by any theory, the inventors theorize that the bone anchoring component of the implant should mimic the mechanical properties of the surrounding solid tissue, particularly the stiffness or modulus, to achieve more durable results. In this regard, the inventors note that the bone anchoring component of an implant will primarily contact cancellous bone (also known as trabecular bone or spongy bone), rather than subchondral bone or cortical bone, which form the hard layer beneath the cartilage and the hard outer layer of the bone, respectively. For example, if the modulus of the anchoring component is below the stated level, micromotion between the implant and the surrounding bone may induce the formation of a fibrous tissue interface rather than the desired direct bond. If the modulus of the implant portion is too high, stress shielding may occur; in this case, loads on the implant, such as those on a cartilage replacement cap, may be primarily propagated within the implant itself, without loading the surrounding bone tissue. When the bone is unloaded, the bone may reshape and / or be resorbed in unloaded areas, ultimately leading to implant loosening. If the stiffness of the implant is too low, the implant may also deform or be damaged when loaded. Furthermore, compared to less rigid materials such as unfilled polyurethane, the polymer compositions of the present invention exhibit reduced creep and plastic deformation, which contributes to better implant stability.

[0091] In an embodiment, and similar to the above paragraph, the polymer composition exhibits an elongation at break of at least 5%, preferably at least 10%, 20%, 30%, 40% or 50% during a tensile test (Eab; dry / 20°C), or an Eab (wet / 37°C) of at least 10%, 20%, 30%, 40% or 50%. In an embodiment, the polymer composition has a tensile strength at break (TS; dry / 20°C) of at least 30 MPa, 35 MPa or 40 MPa or at least 15 MPa, 20 MPa or 25 MPa after conditioning (wet / 37°C). This stiffness and strength properties of the composition in the conditioned state, which are of the same order of magnitude as cancellous bone, are found to enhance the compatibility and bonding of implants made of the composition with such living tissue over time. This also simplifies or improves the insertion of the implant into the bone hole, with a lower risk of damaging the tissue, relative to more rigid implants made of metal. In addition, it was found that implants made of the polymer composition are strong enough to withstand forces after and during implantation. For example, a typical implant procedure performed by an orthopedic surgeon may involve inserting and pressing the implant into a hole drilled at the site of defective cartilage in the bone, such as manually using a hammer and directional guide, which may also limit the depth of penetration and is often in response to their sensory perception (such as noticing changes in sound during hammering).

[0092] In an embodiment, the polymer composition has a Shore hardness of 76-85 ShD, typically 78-82 ShD (dry / 20°C).

[0093] In another aspect, the present invention relates to the use of a polymer composition comprising the above-described biostable thermoplastic polyurethane and 15-70% by weight of inorganic particles (including all features, embodiments, and combinations thereof) in orthopedic implants and in the manufacture of orthopedic implants, in particular in bone anchoring components thereof. Examples of such uses include implants formed by an injection molding process, in particular, such processes comprising the step of molding the bone anchoring component using the composition.

[0094] In another aspect, the present invention relates to a method for producing an orthopedic implant or a bone anchoring component thereof, the method comprising forming the implant by an injection molding process, the process comprising the step of molding the orthopedic implant or the bone anchoring component thereof from the polymer composition of the present invention as defined above. Such molding processes are generally known to those skilled in the art.

[0095] Methods of using the polymer compositions of the present invention and / or making orthopedic implants or bone anchoring components of implants can involve bolts, screws, or other components that can be used to connect or secure further implant components to the bone, such as sutures, artificial ligaments or tendons, menisci, labral replacement devices, or cartilage replacement devices. In embodiments, such further components can form an integral part of the implant, such as a bone anchoring component and a cartilage replacement component, or can be detachably connected, such as sutures.

[0096] In an embodiment, the use and / or the method comprises manufacturing an implant comprising at least two parts, such as a bone anchoring component and a cartilage replacement component. The cartilage replacement component is typically made of a resilient and wear-resistant biocompatible material, such as a segmented block polymer having a hard segment based on polyester, polyamide or polyurethane. In an embodiment, the cartilage replacement component is made of a biostable, resilient thermoplastic polyurethane (TPU), preferably comprising a hard block and a soft block that are chemically similar to the TPU of the polymer composition. In an embodiment, the use and / or the method comprises manufacturing an implant, wherein the cartilage replacement component is made of a biostable, resilient thermoplastic polyurethane having a hardness of 55-100 ShA. The advantage of using TPU material for the bone anchoring component and the other component of the implant is that the two components can be formed by injection molding to form an article with adhered, integrated parts; for example by applying an insert molding method or by a multi-component or two-component molding method.

[0097] In an embodiment, the use and / or the method comprises making a bone anchoring component of an implant, the component consisting essentially of the polymer composition. The bone anchoring component may have a smooth outer surface or a textured surface and may optionally be provided with a surface coating; for example, to influence interaction with body tissue.

[0098] In an embodiment, the use and / or the method comprises making an orthopedic implant comprising a bone anchoring component and a cartilage replacement component, the method comprising forming the implant using a multi-component injection molding process, the multi-component injection molding process comprising the steps of injecting the polymer composition of the present invention into a mold comprising an insert in the form of a cartilage replacement component to form the bone anchoring component, and the subsequent steps of removing the insert from the mold and injecting a resilient and wear-resistant biocompatible TPU material into the mold partially filled with the polymer composition to form the cartilage replacement component on the bone anchoring component. The manufacture of implants from different polymer grades by multi-component molding techniques is known in the art. For example, WO 2011 / 098473 A1 describes the manufacture of orthopedic implants, such as meniscus or intervertebral disc implants, having two or more different segments, each segment comprising different but chemically related polymeric materials; such as different TPU grades.

[0099] In a further embodiment, the use and / or method comprises manufacturing a bone anchoring component having a textured surface, for example a surface having a surface roughness Ra of at least 1, 3 or 5 μm, in order to enhance interaction with surrounding bone tissue after implantation. This surface roughness may be introduced during or after manufacturing the component; for example, by applying a mold having a certain surface texture; for example, as defined by the VDI3400 scale commonly used in the industry. For examples of polymer implants having such a textured surface and methods for their manufacture, reference is hereby made to WO 2019 / 068903 A1. In a further embodiment, the surface roughness Ra of the bone anchoring component of the implant is at least 6 μm, 8 μm, 10 μm, 12 μm, 14 μm, 16 μm, 18 μm or 20 μm and at most 25 μm.

[0100] In other embodiments, the use and / or the method comprises the further step of providing the (smooth or textured) surface of the bone anchoring component of the implant with a surface coating; for example, a bioactive coating, preferably an osteoconductive coating, that further promotes interaction with tissue. Various bioactive or osteoconductive coatings based on organic and inorganic bioactive materials, as described in the art, can be applied to the bone anchoring component of the implant.

[0101] In another embodiment, the use and / or the method comprises providing bioactive ceramic particles on at least part of the surface of a bone anchoring component made of a polymer composition in order to induce osteogenic properties in the component. A suitable method for providing a surface of a bone anchoring component made of a polymer composition with a bioactive coating comprises treating at least part of the surface with a dispersion of bioactive ceramic particles in a solvent for the polyurethane contained in the polymer composition; as described in WO 2019 / 068903 A1.

[0102] Suitable bioactive ceramic particles include all inorganic materials that show the ability to bind directly to living bone, for example by forming bioactive bone-like apatite via interaction or chemical reaction of the particles with surrounding body fluids. Examples of suitable materials include various calcium phosphates, so-called bioglasses and other silica-based ceramics (silicated ceramics). Various types of calcium phosphates have been described for such applications, such as calcium hydrogen phosphate dehydrate (CaHPO4; DCPA), calcium hydrogen phosphate dihydrate (CaHPO4.2H2O; DCPD), octacalcium phosphate (Ca8(HPO4)2.5H2O; OCP), tricalcium phosphate (Ca3(PO4)2; TCP) and hydroxyapatite (Ca 10(PO4)6(OH)2; HA). Different types of blends can also be used or even show advantages, such as mixtures of HA with TCP or HA with bioglass. In addition to their main components, ceramic particles may also contain small or trace amounts of other (inorganic) elements or ions, such as Si, Na, Mg, Fe, Zn, Ti, Ag, Cu, or -SO4 or -CO3, which can improve specific properties of the particles.

[0103] The term bioglass, including commercial Products are mixed inorganic oxide or silicate ceramics with a surface reactive glass film that is compatible with tissues and can be used as surface coatings in medical and dental implants. For example, Grade 45S5 indicates a glass composed of 45% SiO2, 24.5% CaO, 24.5% Na2O, and 6.0% PO5 by mass. The high calcium-phosphorus ratio in this material promotes the formation of apatite crystals; the calcium and silicon ions act as crystallization nuclei. Glass is a non-crystalline, amorphous solid typically composed of a silica-based material and small amounts of other inorganic elements.

[0104] In one embodiment, the particle size of the bioactive ceramic particles is in the range of 0.1-10 μm. Particle size and particle size distribution can be measured using SEM or optical microscopy or using (laser) light diffraction techniques. In the present disclosure, D is measured using light diffraction according to ISO 13320:2009, for example using a Malvern Mastersizer 2000. 50 The value is defined as the particle size of the bioceramic particles. The particle size does not appear to be particularly critical, but larger particles may interact more effectively with body fluids and cells. In view of processability, ceramic particles with a particle size of at least 200 nm, or at least 300 nm, 400 nm or 500 nm are preferred. In other embodiments, the implant has ceramic particles with a particle size of at most 10 μm, 8 μm, 6 μm, 5 μm, 4 μm, 3 μm, 2 μm or at most 1 μm at the surface of the bone anchoring component.

[0105] In embodiments, the use of the polymer composition and / or the method of manufacturing an orthopedic implant comprising a bone anchoring component and an optional cartilage replacement component may involve implants of various forms or shapes proposed in the art, such as implants having substantial axial symmetry.

[0106] In one example, as represented by the oblique top view Figure 1As schematically shown in the simplified figure in FIG, implant 1 is cylindrical in shape with a substantially constant diameter, wherein the larger section 3 of the cylinder represents the bone anchoring component, and the cartilage replacement component is a section (or layer) 2 at one end of the cylinder. The diameter of the cylinder may be 5-20 mm, 10-18 mm, 12-17 mm, or typically approximately 15 mm. Alternatively, the substantially cylindrical form may exhibit some taper, wherein the cartilage replacement component or layer has a diameter that is up to 10% larger than the smallest diameter of the bone anchoring component. A slightly tapered bone anchoring component has been found to facilitate release from the mold used to manufacture the component, as well as placement in a bone hole and ensuring contact with the bone during implantation. The outer side may exhibit a taper of 1-5°, preferably at least 1.5° or 2.0°, and up to 4.5°, 4.0°, 3.5°, or 3.0°, relative to the longitudinal orientation of the anchoring component. The cartilage replacement component, or at least its top surface, can be substantially flat or can be curved or contoured to simulate the curvature of the articulation joint in which it is to be implanted. In embodiments, the cartilage replacement component of an orthopedic implant for cartilage replacement has a substantially constant or uniform thickness, preferably within the same range as the cartilage it is to be replaced; for example, a layer of material over a bone anchoring member that also serves as a support for the elastic layer.

[0107] In other embodiments, the implant 1 can have substantial axial symmetry, having a mushroom-like shape with a cap and stem of varying diameters. The simplified diagrams shown in Figures 2A and 2B schematically illustrate an oblique top view and a cross-sectional side view of such a mushroom-shaped implant. The stem 3b is cylindrical with a substantially constant diameter or a slight taper and is formed together with the lower section 3a of the cap of the bone anchoring member 3. As described above, the outer side of the stem may exhibit a taper of 1-5° relative to its longitudinal orientation, with the stem diameter decreasing slightly from the cap to the bottom. Both sections 3a and 3b forming the bone anchoring member 3 are made of the same polymer composition (such as the compositions described in the present disclosure). The top section (or layer) 2 of the cap represents the cartilage replacement component, for example, made of an elastic TPU with a hardness of 55-100 ShA. The stem diameter may be approximately 5-15 mm, typically approximately 6-10 mm; and the cap diameter may be approximately 5-25 mm, 10-20 mm, 12-18 mm, or approximately 15 mm. In practice, typically a range of implants of different sizes, in particular implants with different cap sizes, can be prepared for use as part of, for example, a kit; thereby allowing the selection of a suitable implant, depending on the patient to be treated. The cartilage replacement component 2, or at least its top surface, can be curved or contoured in one or more directions to simulate the curvature of an articulated joint in which the implant can be implanted. In an embodiment, the cartilage replacement component of such an orthopedic implant for cartilage replacement is a layer of substantially constant thickness, which thickness is preferably in the same range as the cartilage layer it is to replace. The bottom surface of the cap 3a, forming part of the bone anchoring component, is substantially flat; preferably having rounded edges and a rounded transition to the stem. Similarly, the bottom edge of the stem is also rounded (not shown in the simplified view of Figure 2). In general, sharp edges and transitions are omitted to reduce stress concentrations. During implantation, a bone hole is created with a diameter equal to or slightly smaller than that of the plug's stem 3b. The damaged cartilage is then removed to a diameter and depth that exposes the bone tissue, preferably such that the side and bottom surfaces of segment 3a contact the patient's bone tissue during insertion; similarly, the surface of stem 3b contacts the cancellous bone. Cartilage replacement component 2 then preferably contacts only the natural cartilage on its sides (as well as the cartilage surface in the joint opposite the top surface of component 2).

[0108] In other embodiments, the use of the polymer composition and / or the method of making an orthopedic implant relates to an orthopedic implant that is not axisymmetric as a whole, but can, for example, be considered to include two similar axisymmetric mushroom-shaped implant components, similar to the components discussed above, that are partially fused together; such as conjoined twins. An example of such an implant is schematically shown in Figure 3In the embodiment of the present invention, the implant comprises a substantially oval or elliptical cap having a contoured top surface and two substantially identical cylindrical or conical stems. In this embodiment, the domed top layer 2 of the cap having a substantially uniform thickness forms the cartilage replacement component, and the underside of the cap 3a together with the two stems 3b' and 3b" form the bone anchoring component 3.

[0109] In other embodiments, the use and / or method may further include the step of including an orientation marker within the implant, such as at the interface between the bone anchoring component and the cartilage replacement component. For example, the marker may include a recessed or raised line or indicator on the top surface of the bone anchoring component during molding, or by placing a small, elongated object in the mold between the steps of injecting the polymer composition to form the bone anchoring component and the biocompatible material to form the cartilage replacement component. Such a marker may be visible through a transparent or translucent cartilage replacement component, as may be the case if it is made of unfilled TPU; this allows implants that are not completely axially symmetrical to be placed in the desired orientation at the implant site. The orientation marker may be opaque so that the orientation of the plug can be visualized during and after implantation. For example, if the surface of the cartilage replacement component of the implant has more than one contour radius to better match the curvature of the bone, the orthopedic implant may further include a relatively small, radiopaque, elongated marker, such as a short piece of thin metal wire or strip, or an array of small marker particles. For example, such a radiopaque orientation marker may be approximately 2-6 mm in length, such as a piece of tantalum wire. In an embodiment, such an orientation marker is present in the region between the bone anchoring component and the cartilage replacement component. Figure 2b 4. Thus, the surgeon can place the plug in the desired direction using his eyes and / or medical imaging techniques; medical imaging techniques can also be used to check its position and any potential changes therein after surgery.

[0110] In another aspect, the present invention relates to an orthopedic implant comprising a bone anchoring component comprising the polymer composition of the invention as described above, including all features, embodiments and combinations thereof, and which has been obtained by the method of the invention as described above.

[0111] Unless otherwise indicated herein or obviously contradictory with context, otherwise in the context of describing the present invention (especially in the context of the following claims) the use of term " one " and " an " and " the " and similar indicators should be interpreted as covering both singular and plural. Unless otherwise indicated, otherwise the terms " comprise ", " have ", " include " and " contain " should be interpreted as open terms (that is, meaning " including but not limited to "). The description of value ranges herein is only intended to be used as a shorthand method for quoting each individual value falling into the scope, and each individual value is incorporated into this specification as if it were individually quoted in this article. Unless otherwise indicated, the use of any and all examples or illustrative language (for example, " for example ") provided herein is only intended to better illustrate the present invention, rather than to limit the scope of the present invention claimed. Any language in the description should not be interpreted as representing that any element not claimed for protection is essential for practicing the present invention.

[0112] The preferred embodiments of the present invention are described herein, including the best mode known to the inventor for implementing the present invention. After reading the foregoing description, variations of those preferred embodiments will be apparent to those of ordinary skill in the art. The inventors hope that the skilled person will appropriately adopt such variations, and the inventors hope to practice the present invention in a manner different from that specifically described herein. Therefore, the present invention includes all modifications and equivalents of the subject matter described in the appended claims herein as allowed by applicable law. Although some optional features are described as embodiments of the present invention, this description is intended to include and specifically disclose all combinations of these embodiments, unless otherwise expressly noted or physically impossible.

[0113] The following experiments and samples further illustrate embodiments of the present invention, but should of course not be construed as limiting the scope of the claims in any way.

[0114] Experimental part

[0115] method

[0116] Molar mass

[0117] As described in ASTM D5296-11, the TMThe molar mass and molar mass distribution of the samples were measured by GPC (gel permeation chromatography, also known as size exclusion chromatography or SEC) on a Viscotek GPCMax VE2001 system equipped with 10 μm 10E6A, 10 μm 10E4A, and 10 μm 100A columns. The detector and column were operated at 80°C. The polymer samples were dissolved at a concentration of 1.0 mg / ml in DMF containing 0.05% by mass of LiBr and 300 mg / l of DHT at 70°C for up to 4 hours and then filtered through a 0.2 μm PTFE membrane. This solvent composition also served as the eluent. Molar mass calculations were based on a calibration curve obtained using EasyCal polystyrene standards and adjusted based on the results of polyurethane samples of known molar mass.

[0118] hardness

[0119] The hardness of the molded samples was measured using a Zwick Shore Durometer 3131 according to ISO 868 (average of 5 tests, measuring time 15 s, at 20.9° C. / 51.1% RH).

[0120] Tensile properties

[0121] The tensile modulus, (ultimate) tensile strength and elongation at break were measured on a Zwick Z010 universal tensile testing device equipped with a 2.5 kN pneumatic clamp and a temperature chamber using a method based on ISO 527. Injection molded test bars (ISO 5271 BA type) were tested at 20 ° C for dry molding and conditioned in 37 ° C water at 37 ° C. The samples for dry molding measurements were dried overnight using an 80 ° C vacuum oven with a small N2 flow rate and stored in a closed box with silica gel. The samples were conditioned by immersing the samples in water maintained at 37 ° C and measuring the mass increase every 24 h until the change was less than 0.1% by mass (usually at least 360 h). Each sample was stored under the conditions until shortly before being placed in the temperature chamber of the tensile testing machine. Just before the test, the width and thickness of the sample were measured at the center of the sample. The sample was placed in a clamp with a spacing of 54 mm between the clamps at the starting position. A preload of 0.5 N was applied before the tensile test began. The modulus of the sample was measured at a speed of 1 mm / min within the first 0.05% and 0.25% strains. Thereafter, stress and strain measurements were performed at a speed of 50 mm / min until the sample broke, and the elongation at 60% strain was measured using an extensometer.

[0122] Particle size

[0123] The particle size distribution and particle size (D10, D 50 and D 90 ).

[0124] Crystallization behavior

[0125] Differential scanning calorimetry (DSC) was performed using a standard heat flux DSC from Mettler Toledo. Samples of approximately 5 mg mass were cut from the pellets, weighed using a precision balance, and packaged in a (crimped) aluminum pan of known mass. An identical empty pan was used as a reference. Nitrogen was purged at a rate of 50 ml / min. A heating-cooling-heating cycle was applied to determine the parameters that numerically characterize the thermal behavior of the material under investigation. The temperature program applied was: [1] 0.0-70.0°C, 10.00 K / min; [2] 70.0°C, 60.00 min; [3] 70.0-90.0°C, -10.00 K / min; [4] -90.0°C, 10.00 min; [5] -90.0-240.0°C, 10.00 K / min; [6] 240.0°C, 2.00 min; [7] 240.0-90.0°C, -10.00 K / min; [8] -90.0°C, 10.00 min; [9] -90.0-240.0°C, 10.00 K / min.

[0126] polymer composition

[0127] Experiments 1 to 3.

[0128] In the experiments, a polycarbonate urethane based on MDI, butanediol, and poly(hexamethylene carbonate) glycol with a hardness of 75 ShD and a mass-average molar mass (Mw) of 388 kDa was used after drying at 80°C for 24 hours to a moisture content of 113 ppm. Hydroxyapatite (Merck, Hydroxyapatite for Bioceramics) with a particle size of 5 μm was dried at 120°C for 24 hours to a moisture level of approximately 2650 ppm.

[0129] The polyurethane and filler components are mixed in Coperion ZSK Mc 18Melt mixing was carried out on a twin-screw extruder equipped with two Colortronic LabLine feeders and a die plate with a 3 mm opening. Polyurethane granules were metered on the feed barrel in a TPU / HA mass ratio of 80 / 20, and HA powder was metered into barrel 2 on a side feeder. The temperature of all zones was set to 190°C. The screws rotated at a speed of 150-200 rpm, with a production capacity of approximately 2.5 kg / h and torque levels varying slightly in the range of 50-70%. Extrusion conditions were selected to produce a melt temperature of up to approximately 200°C; to achieve stable compounding, resulting in smooth and regular strands. The extruded strands were cooled in a water bath with a total cooling length of 2 m and then cut into pellets using an electric strand picker and a Scheer 50E pelletizer (operating at low speed).

[0130] The extruder was flushed before and after the experiments with mixing with polyurethane for 5-10 min at a throughput of 2 kg / h.

[0131] After drying at 120°C for 24 hours under vacuum / N2, the resulting pellets were injection molded into test bars on an Xplore IM12 microinjection molding machine. The barrel temperature was set to 200°C and the mold temperature was set to 100°C. After a melt time of at least 5 minutes, the material was injected into the vented mold at an injection pressure of 10 bar within 2.2 seconds and a filling pressure of 10 bar within 10 seconds.

[0132] The tensile properties of the molded specimens were measured at room temperature (20°C) in a dry state and at 37°C in a wet state to simulate the conditions when used as implant materials.

[0133] In Table 1 the results are summarized as (Comparative) Experiment 2. Also provided in this table are the molar mass data (Mw) determined on the pellets and the molded test specimens.

[0134] Experiment 1 relates to the results of a reference test carried out on unfilled polyurethane pellets which were injection moulded similarly to Experiment 2 above, but with an ageing temperature of 235° C. and a pressure of 15 bar.

[0135] Experiment 3 was carried out similarly to Experiment 2, but the amount of HA dosed to the extruder was 40% by mass.

[0136] The tensile measurement results summarized in Table 1 show that the hydroxyapatite particles act as a non-reinforcing filler in the polyurethane; the dry modulus does not change significantly, the tensile strength and elongation decrease significantly, but the water absorption decreases and has little effect on the properties (20°C / dry vs. 37°C / wet). The Shore hardness of the molded bars increases from 75.7 ShD (Experiment 1) to 80.2 ShD (for Experiment 2). The GPC results show a significant reduction in molar mass when compared to the unfilled material, which is likely due to the residual moisture contained in the crystallized HA particles inducing polymer degradation during the melt processing step; mainly during the compounding step. The observation that the melt temperature should be kept below 200°C during compounding, as otherwise stable strand extrusion cannot be achieved, supports this interpretation.

[0137] Experiment 4.

[0138] Similar to the compounding procedure of Experiments 2 and 3, a particle size of 4 μm (D 50 Bismuth oxide particles (Helos / Rodos; 5N Plus Product Data Sheet (Helos / Rodos)) were used in place of HA. Various variations in processing conditions were attempted, but experiments were discontinued when stable processing could not be achieved and / or when blistering and discoloration of the extrudate were observed. This was likely caused by excessive degradation of the polyurethane. No further experiments were conducted with this material.

[0139] Experiment 5 to Experiment 6.

[0140] In these experiments, zirconium oxide TZ-0 (Tosoh Europe BV), a free-flowing powder with a (primary) particle size of 0.04 μm (according to the Tosoh brochure), was added to the polyurethane as a filler during compounding. The sample was dried at 120°C for 24 h and compounded with the same 75ShD polyurethane applying the procedures of Experiments 2 and 3. By setting the barrel temperature to 200°C and increasing the screw speed to 300 rpm, strand breakage during compounding was prevented and particle dispersion appeared to be improved. The torque level was limited to 60-70% and the melting temperature was at most 212°C. The metering behavior of the powder was very good, but the coarseness variations observed in the extruded strands indicated that the particles were not well dispersed.

[0141] Adding these zirconium oxide particles at loadings of 20 and 40 mass% resulted in reduced strength; at a loading of 40 mass%, the composition was even quite brittle. This is likely due to insufficient dispersion of the agglomerated primary particles within the polyurethane matrix. Pretreatment of the particles and / or addition of dispersants can be used to enhance dispersibility and improve mechanical properties.

[0142] Experiment 7 to Experiment 8.

[0143] Using the same conditions and procedures as in Experiments 5 and 6, a composition of 75ShD polyurethane with zirconium oxide TZ-Y3-E (Tosoh) was prepared. This zirconium oxide grade TZ-Y3 contains 3 mol% yttrium and is also sold as a free-flowing powder with a (primary) particle size of 0.04 μm. Processing was similar to that observed for TZ-0 powder. As with the pure zirconium oxide grades, the polyurethane composition tended to show brittle failure during tensile testing, especially at a load of 60% by mass. As determined and as Figure 6 The particle size distributions shown for the TZ-0 and TZ-Y3 grades confirm that the actual particles are much larger than their primary particles; the maximum values ​​are approximately 60 μm and 50 μm, respectively, with a smaller maximum value of approximately 1 μm. This again suggests that the observed processing and brittleness issues can be overcome if these particles are pretreated and / or (partially) deagglomerated before mixing with the polymer, and / or if a dispersing aid is added during mixing. Furthermore, using mixing equipment that allows higher forces than the small laboratory-scale equipment used can also enhance dispersion.

[0144] Experiments 9 to 11.

[0145] In these examples, medical grade zirconium oxide (referred to herein as ZrO / SA) obtained from Sigma Aldrich after drying at 120°C for 24 hours was used. The particle size D of these particles was found to be 50 is about 1.8 μm; this is much smaller than the particle size of the particles used in Experiments 5 to 8 above, as Figure 6 As shown. Compounding was carried out in a similar manner, but some additional mixing elements were used in the feeder to ensure the correct metering of the dry, somewhat sticky powder. The polymer strands and granules were made in a stable process. The prepared compositions showed an increase in tensile modulus with increasing zirconia loading (20 mass %, 40 mass % and 60 mass %), but a decrease in tensile strength and elongation. All samples showed a tensile property profile (also in the conditioned state) that would be suitable for use in the manufacture of bone anchors. The addition of zirconia increased the hardness from 75.7 ShD (Experiment 1) to 81.1 ShD and 80.2 ShD (Examples 10 and 11)

[0146] GPC measurements show that the molar mass decreases during melt processing, especially during compounding; however, these decreases appear to be significantly smaller than those observed for the compositions containing HA particles. Unfilled polyurethanes exhibit higher molar masses after injection molding; however, this material has not undergone a previous compounding step. The crystallization behavior of selected samples was investigated by conventional DSC; the relevant results (temperature and enthalpy; (J / g sample)) are summarized in Table 2, and the DSC curves for Experiment 1 and Experiments 9-10-11 are shown in Figure 7 For the composition TPU / HA 80 / 20 (Experiment 2), crystallization appears to initiate during cooling at a higher temperature than for the reference polyurethane, which may be related to the low molar mass of the degraded polyurethane and / or a nucleation effect of the hydroxyapatite particles. On the other hand, the broadening of the crystallization and melting peaks, especially that seen for higher amounts of filler particles, indicates that polymer crystallization is hindered (lower rate). The composition containing zirconium oxide particles shows a nucleation effect during cooling and a higher melting temperature in the reheating scan.

[0147]

[0148] Table 2

[0149]

[0150] Cartilage plug

[0151] A prototype cartilage plug having a mushroom-like shape as shown in FIG. 2A to FIG. 2B was produced on a desktop Xplore IM12 microinjection molding machine using a mold holder component comprising a 2-part mold and applying a set of inserts to achieve two-component molding.

[0152] The materials used were a TPU composition containing 60% by mass of zirconium oxide (Experiment 11) and an unfilled end-group modified polycarbonate urethane with a hardness of 80 ShA ( II 80A; DSM Biomedical BV, Sittard-Geleen, NL). The materials were dried at 120°C for 24 h or at 80°C for 72 h under vacuum / nitrogen flushing before use.

[0153] Different mould parts together define the implant component with following basic dimensions: length is 6.5mm, diameter is 6.1mm and the stem section with 45 ° of rounding at the bottom; Diameter is the cap section of 10.1mm, the top surface of the contoured shape with two radii of 11mm and 18mm at right angles to each other, wherein the total cap height is 3.5mm, wherein lower 1.0mm is connected to the stem made of rigid polymer composition. All other edges are all with 0.5mm radius rounding. Etching limits the surface of the mould part of stem section to produce the texture of VDI3400 36 with roughness.

[0154] In the first step, a polyurethane / zirconia composition was injected into a mold containing an insert in the cap section, applying a barrel temperature of 210°C, a mold temperature of 80°C, an injection pressure of 10 bar for 2.2 seconds, and a packing pressure of 10 bar for 10 seconds. The insert was then removed from the mold, and a 4 mm long tantalum wire, preheated on a hot plate set at 250°C, was placed on the injection surface in the mold, oriented with an 18 mm radius cap. Subsequently, an elastic polyurethane material was injected into the partially filled mold, applying a barrel temperature of 235°C, a mold temperature of 80°C, an injection pressure of 12 bar for 2.2 seconds, and a packing pressure of 12 bar for 10 seconds to form the cap's top section.

[0155] The molded plugs were functionalized with bioceramic particles by coating the underside of the stem and cap with a dispersion of BCP particles (biphasic calcium phosphate; CamBioCeramics) in THF; the samples were air-dried, rinsed multiple times with ethanol, and dried (50° C. under reduced pressure). These plugs according to the invention were designated as Plug I).

[0156] Similar to the above process, the plug is molded from the corresponding unfilled material, namely polycarbonate urethane ( 75D; DSM Biomedical BV, Sittard-Geleen NL) and stem segments made of II 80A material cap section. These unfilled plugs are referred to as plugs U hereinafter.

[0157] Furthermore, a metal plug with geometry and dimensions corresponding to the polymer plug was produced, wherein the stem portion consisted of titanium, the cap portion was made of cobalt chromium, and the stem was post-treated by corundum blasting (referred to as plug M).

[0158] In vivo testing

[0159] The three types of plugs were evaluated in an animal study in which the devices were implanted in the knees of 32 Dutch goats. The study was approved by the local and national Animal Ethical Committees under project permit PV2015-018-003.

[0160] The animals' knees (32 x 2) were divided into 4 groups, 3 of which were implanted with 16 plugs each of I, U, and M; and the remaining group underwent sham surgery (placebo surgery without plug placement) as a reference for natural cartilage degradation.

[0161] The surgical procedure involved a medial parapatellar skin incision, opening the joint capsule to expose the medial femoral condyle, and locating the center of its weight-bearing portion. To insert the plug, an osteochondral defect was created using a cannulated drill guided by a Kirschner wire. The drilling depth was controlled so that the implant was flush with or slightly recessed from the adjacent cartilage. The plug was then inserted using a press-fit fixation, while the cap's double curvature was aligned to the knee's morphology using orientation markers. No intraoperative or postoperative complications occurred.

[0162] After 26 weeks (6 months) at the laboratory animal testing facility of Maastricht University, 4 animals in each group were euthanized with an overdose of pentobarbital (pentobarbital) (200 mg / kg body weight). The knee was excised and subsequently dissected. The medial femoral condyle and tibial plateau were separated and fixed in neutral buffered formalin. The fixed specimen was dehydrated by incubation in an ethanol aqueous solution (up to 100% ethanol) with increasing concentration. The medial femoral condyle was embedded in a resin based on hydroxyethyl methacrylate (Technovit 8100, Hereaus Kulzer, Hanua, DE) under vacuum. Subsequently, a polymethyl methacrylate (PMMA, Technovit 3040, Hereaus Kulzer, Hanua, DE) coat was created for each block to prevent expansion. The plastic block was then cut using a band saw to horizontally orient the implant and mounted on a diamond saw (SP1600, Leica Biosystems, Nussloch, DE) using ultra-low viscosity cyanoacrylate glue. An incision was made in the middle of the implant. Safranin-O / Fast-Green (Carl Roth, Karlsruhe, DE) was used for staining. The tissue was gently wiped dry and allowed to air dry for five to ten minutes. A glass coverslip was glued to the tissue using cyanoacrylate glue. 50-70 μm sections were cut and glued to a slide using cyanoacrylate glue. The sections were scanned using a bright light microscope (M8 Microscope, Precipoint, Freising, DE) with a magnification of 200x. Bone-implant contact (BIC) was determined using a custom-written MATLAB script (MathWorks, Natick MA, US), which is defined as the percentage of direct contact between the implant surface and the bone.

[0163] The remaining animals were euthanized after 12 months, and the knees and implants were evaluated using the procedures described above.

[0164] Figure 6The results of the BIC scores of the implants after 6 and 12 months, expressed as mean numerical scores (%), are shown in . It can be observed that plug I, having a stem made of a zirconium oxide-TPU composite and having BCP particles on the surface, showed the highest mean BIC score; indicating that this plug had better bone-to-implant contact or in vivo osseointegration than observed for plugs M and U. It is noted that one plug M and one plug I showed hardly any bone contact after 6 months; this seems to be caused by misalignment or tilting of the metal plug, as well as cracking caused by air entrapment in the plugs based on filled polymer compositions, for example. Of the 8 implants made entirely of unfilled polyurethane, 4 had very low BIC scores, which may be due to deformation of the relatively soft plugs.

[0165] After 12 months of implantation, one of eight M plugs, four of eight I plugs, and eight of eight U plugs had a BIC score of (almost) 0%. Again, the unfilled plugs appear to lack sufficient stiffness, and in the case of the metal and composite plugs, this appears to stem from defects caused by the minor compounding and / or molding processes employed, or by the method used for plug insertion during surgery, rather than from the actual performance of a properly manufactured and placed plug. The other M and I plugs all performed well, with photographs showing increased osseointegration compared to the 6-month results. This is not surprising for metal plugs, as they are known to initially exhibit relatively slow bone integration, but after several years, the so-called stress shielding effect often leads to debonding. Ignoring these erroneous results, the BIC scores for the M and I plugs evaluated after 12 months were 40% and 46%, respectively. It is anticipated that if the production of the polymer composition and the molding of the composite plug I are scaled up and improved, and if the implantation of the plugs is better controlled, the observed defects and negative outcomes could be reduced or prevented.

[0166] The above situation is solved by Figure 7 Representative photographs of histological sections of each plug type are shown for further illustration (after 6 months and using transmitted light). It should be noted that these photographs are grayscale versions of the original color photographs. However, it can be unequivocally concluded that Plug 1 exhibits a tight fit with the surrounding tissue, with no voids or other irregularities at the interface or in the tissue. Also note the clearly visible distinction between the top cartilage replacement component and the bone anchoring component of the TPU-based plug, as well as the presence of a tantalum orientation marker (shown as a black dot, oriented perpendicular to the slice) at the interface of the two components.

[0167] Furthermore, MRI images of the knee with the implants taken 6 m after implantation showed no significant differences between Plug I and Plug U, but images of Plug I showed better contrast between the plug and tissue. Neither plug appeared to interfere with MRI assessment of the cartilage in the joint, whereas MRI imaging of metallic Plug M and adjacent cartilage was not feasible due to artifacts caused by the metal portion. The stem portion of Plug I, like Plug M, was clearly visible on the radiographs, but Plug U was only partially visible and faintly discernible.

[0168] After 6 months and 12 months, after the tibial plateau was fixed and dehydrated, a 3-4 mm thick coronal osteochondral plate was cut from the tibial plateau with a band saw. The individual plates were then decalcified in formic acid at room temperature for at least six weeks. The specimens were then embedded in paraffin and 5 μm thick slices were cut using a standard microscope microtome (Leica RM2245, Leica Biosystems, Nussloch, DE). The slices were stained with Safranin-O / Fast-Green (Carl Roth, Karlsruhe, DE) and subsequently digitized using a bright light microscope slide scanner (M8 Microscope, Precipoint, Freising, DE) with a magnification of 200 times. Finally, the quality of tibial cartilage was scored using the modified Mankin score (MMS) system, as described by Little et al. (DOI: 10.1016 / j.joca.2010.04.016). In Figure 11, the results of numerically scoring cartilage damage or degradation on the opposing bone surfaces are summarized. In these cases, the metal implants were found to have much more severe cartilage damage than the cartilage damage caused by the TPU-based plugs and the sham-operated group. Plug I, which had a zirconia-filled polyurethane stem and a cartilage contact component that was not filled with polyurethane, was found to perform similarly to the sham-operated goats (i.e., goats with only natural cartilage in their knee joints). Plug U was scored similarly to Plug I, with the soft polyurethane causing no damage; however, it did not. Figure 8 Shown in.

[0169] In summary, a cartilage plug (Plug 1) having a stem made of a zirconium oxide / TPU composition and a cartilage replacement cap made of unreinforced TPU exhibited the best overall performance; it had appropriate handleability and implantability, showed good osseointegration, caused little damage to the opposing cartilage surface, and was imageable as an implant using micro-CT, X-ray, and MRI techniques. Furthermore, several options are described in this disclosure to further improve the properties of the filled TPU composition, and therefore the performance of this implant.

Claims

1. A method for making an orthopedic implant having a bone anchoring component connected to a cartilage replacement component, comprising: a. Providing a first polymer composition comprising a biocompatible, biostable thermoplastic polyurethane and 15-70% by mass of biocompatible inorganic particles containing zirconium dioxide, wherein the particle size D of the inorganic particles 50 In the range of 0.1-5 µm; b. providing a second polymer composition comprising an elastic and wear-resistant biocompatible thermoplastic material; and c. Forming the orthopedic implant using an insert molding or multi-component molding method, wherein the bone anchoring component is injection molded from a first polymer composition and the cartilage replacement component is injection molded from a second polymer composition.

2. The method according to claim 1, wherein the thermoplastic polyurethane has a Shore hardness of 40-90 ShD.

3. The method according to claim 1 or 2, wherein the thermoplastic polyurethane comprises an aliphatic polyether, an aliphatic polyester, an aliphatic polycarbonate or a combination thereof in the soft block.

4. The method according to claim 1 or 2, wherein the thermoplastic polyurethane is a blend of a thermoplastic polyurethane having a hardness of 50-80 ShA and a thermoplastic polyurethane having a hardness of 70-85 ShD. 5 . The method according to claim 1 , wherein the thermoplastic polyurethane further comprises 0.01 to 5 mass % of one or more biocompatible additives based on the amount of the polyurethane. The method according to claim 1 , wherein the inorganic particles further comprise other transition metal oxides.

7. The method according to claim 6, wherein the other transition metal oxide is selected from oxides of at least one of Ti, Zn, Y, La, Yb, Hf and Ta.

8. The method according to claim 1 or 2, wherein the inorganic particles consist of zirconium dioxide.

9. The method according to claim 1 or 2, wherein the first polymer composition comprises 20-70% by mass of inorganic particles.

10. The method of claim 9, wherein the first polymer composition comprises up to 50% by mass of inorganic particles.

11. The method according to claim 1 or 2, wherein the first polymer composition comprises a) 30-65 mass % of the biostable thermoplastic polyurethane, b) 35-70 mass % of the inorganic particles, and c) 0-10 mass % of other compounds, the sum of a) to c) being 100 mass %.

12. The process according to claim 1 or 2, wherein the thermoplastic polyurethane has a mass average molar mass Mw of 70 to 400 kDa, measured by GPC according to ASTM D5296-11.

13. The method according to claim 1 or 2, wherein the bone anchoring member has an E-modulus of 800-2000 MPa, measured at 20°C on a dry molded sample according to ISO 527.

14. The method according to claim 1 or 2, wherein the bone anchoring member has an E-modulus of 200-700 MPa, measured on a sample conditioned in water at 37°C, based on ISO 527.

15. The method according to claim 1 or 2, wherein the bone anchoring member has an elongation at break of at least 5%, measured on a dry, shaped sample at 20°C according to ISO 527.

16. The method according to claim 1 or 2, wherein the thermoplastic polyurethane has a Shore hardness of 60-85 ShD.

17. The method of claim 1 or 2, wherein the thermoplastic polyurethane comprises an aliphatic polycarbonate.

18. The method of claim 1 or 2, wherein the first polymer composition is free of calcium phosphate-based particles.

19. An orthopedic implant comprising a bone anchoring component and a cartilage replacement component, wherein the bone anchoring component is formed by the method according to any one of claims 1 to 18.

20. A method for making an orthopedic implant having a bone anchoring component connected to a cartilage replacement component, comprising: a. Providing a first polymer composition comprising a biocompatible, biostable thermoplastic polyurethane and 15-70% by mass of biocompatible inorganic particles containing zirconium dioxide, wherein the particle size D of the inorganic particles 50 In the range of 0.1-5 µm; b. providing a second polymer composition comprising an elastic and wear-resistant biocompatible thermoplastic material; and c. injecting the first polymer composition into a mold comprising an insert in the form of a cartilage replacement component to form a bone anchoring member; d. removing the insert from the mold; and e. injecting the second polymer composition into the mold to form a cartilage replacement component.

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