Implant for medical care of humans and / or animals and bioresorbable composite material therefor

Abstract

The invention relates to an implant for medical care of humans and / or animals. In order to provide an implant which dissolves in the body during and / or after the healing process and at the same time counteracts possible inflammation after an operation, the implant comprises or consists of: a) a bioresorbable implant main body; b) a bioresorbable fastening body which is connected to the implant main body and by means of which the implant main body can be secured within a human or animal body; c) optionally an antibacterial substance. The invention further relates to a bioresorbable composite material for such an implant and to a method for treating a patient or providing prophylaxis for a patient with such an implant.

Description

[0001] Implant for the medical treatment of humans and / or animals and bioresorbable composite material therefor

[0002] The invention relates to an implant for the medical care of humans and / or animals.

[0003] Furthermore, the invention relates to a bioresorbable composite material for such an implant.

[0004] Modern surgical techniques can utilize a wide variety of implants for the medical treatment of humans and / or animals for applications ranging from cardiovascular diseases to bone fractures. The spectrum of possible implants is very broad and can include, for example, body-like implants in the form of allografts for ruptures or, in particular, re-ruptures of cruciate ligaments. However, implants are often made of exogenous materials such as ceramics, for example, in dental surgery, or, in particular, metal alloys, such as stents for the treatment of heart disease. Titanium alloys are frequently used for bone screws, bone nails, and bone plates, for example, for the healing of a femoral neck fracture.After a healing process, the corresponding implants are generally removed from the body if necessary, medically advisable, and feasible in terms of the risk-benefit ratio. During the healing process, the materials for the implants should be designed in such a way that, in addition to being biocompatible, for example, when exposed to blood, they must also meet a specified mechanical stress profile to prevent further complications due to material failure during the healing process.

[0005] In addition to implants made of corrosion-resistant materials such as titanium alloys, which have become particularly popular for treating fractures, implants have also been developed that dissolve in the human body over time. A separate surgical procedure for implant removal can then be omitted, if at all possible. Magnesium alloys are particularly suitable for this purpose, although for many years magnesium alloys with significant amounts of rare earth metals were the preferred choice. Such magnesium alloys dissolve in the human body over time; however, such alloys are so corrosion-resistant that the implants are not so extensively decomposed by bodily fluids during the healing process that they lose their required functionality.Such implants can also be used, for example, to treat fractures because the contact with body fluid is not so intense that corrosion can occur very quickly.

[0006] Magnesium alloys containing rare earth metals dissolve in the body over time. There are medical concerns regarding the rare earth metals, as these remain in the body after the implant has dissolved. Long-term studies do not appear to exist yet, but the concerns are significant because the concentrations of rare earth metals in the magnesium alloys used can be significant. For example, alloy WE43, a standard alloy for implants, contains approximately 4% yttrium and 3% other rare earth metals (e.g., neodymium and / or gadolinium) by weight.Particularly when such implants are used in areas where the implants come into close contact with body fluids, for example as cardiovascular stents, the rare earth metals from such implants enter the bloodstream in an uncontrolled manner and there are concerns that this could lead to long-term consequences.

[0007] Therefore, the proportion of rare earth metals in biodegradable implants has been reduced. Alloys that are essentially free of rare earth metals have now been developed for the treatment of fractures (P. Holweg et al., A lean bioabsorbable magnesium-zinc-calcium alloy ZX00 used for operative treatment of medial malleolus fractures, Bone Joint Res 2020, 9(8), 477; P. Holweg et al., A lean magnesium-zinc-calcium alloy ZX00 used for bone fracture stabilization in a large growing-animal model, Acta Biomaterialia 113 (2020) 646).

[0008] The development of bioresorbable implants results in very specific requirements for the implant to be inserted surgically: Firstly, the implant must be able to withstand the expected mechanical loads over the duration of the expected healing process. If the implant loses too much strength during the healing process, for example, this can be problematic for the healing process. Furthermore, the implant should ideally only dissolve noticeably after the healing process. Premature dissolution can impair the healing process. Ultimately, this is also related to the fact that the dissolution of the implant is accompanied by a decrease in its mechanical properties. Especially for implants that come into close contact with bodily fluids, especially blood, it is extremely difficult to specify a suitable material.This applies, for example, to implants placed on or in a person's heart, such as stents. Particularly when the implant comes into close contact with body fluids such as blood, there is an increased risk of bacterial infection resulting from the operation. This is even more true if, in addition to the actual implant body, additional components are required to position the implant body, as is necessary, for example, with annuloplasty rings for the treatment of heart valve disease.

[0009] Based on the prior art, the invention aims to further develop an implant of the type mentioned at the outset in such a way that it has sufficient stability even in the case of intensive contact with body fluid, in particular blood, during a healing process, but dissolves essentially completely in the body at the latest after a healing process and optionally at the same time counteracts a bacterial infection, in particular following an operation.

[0010] A further aim of the invention is to provide a composite material suitable for this purpose.

[0011] The object of the invention is achieved if an implant of the type mentioned at the outset comprises or consists of the following components: a) a bioresorbable implant base body; b) a bioresorbable fastening body which is connected to the implant base body and via which the implant base body can be fastened in a human or animal body; c) optionally an antibacterial substance. An implant according to the invention achieves several advantages: Firstly, the implant base body, as well as the fastening body which is connected to the implant base body and via which the implant base body can be fastened in the human or animal body, are bioresorbable. Bioresorbable here means that the implant base body, as well as the fastening body for this purpose, dissolve at least essentially completely over time in the human body.The dissolution in the human body is calibrated for both the implant base body and the attachment body in such a way that the healing process is not jeopardized. An additional antibacterial substance prevents or at least reduces the risk of bacterial infection in the area of ​​the implant, especially immediately after surgery and implantation. The antibacterial substance is a component of the implant and is therefore released as it dissolves. Since the antibacterial substance is generally present not only within one of the components (implant base body and / or attachment body), but also on the respective surface of the component, the antibacterial effect occurs immediately after implantation.As the implant degrades, the antibacterial substance, if present within one of the components, is continuously released. This allows the antibacterial effect to be maintained throughout the entire time the implant is in the body.

[0012] The bioresorbability of the implant body and the fixation body are coordinated depending on the intended use. If the implant base body essentially remains in a fixed position after the implant has been fixed in the human body, the fixation body can be made of a material that dissolves more quickly than the implant base body in the body. However, if fixation is necessary throughout the entire healing process and the fixation body also serves to hold the implant base body in a fixed position, the fixation body should dissolve more slowly than the implant base body. Ideally, the fixation body will then only noticeably dissolve once the healing process is complete, i.e., when the implant base body is no longer required, and the fixation body can fulfill its function as a fixation device for the implant base body until then.In this case, the dissolution of the fixation body can begin before the healing process is complete, but must not progress to such an extent that the implant base body can no longer be or is no longer held in a stable position independently before the healing process is complete.

[0013] Due to the defined bioresorbability of both the implant base body and the fastening body, an implant body according to the invention can be used for various implants for the treatment of diseases or accidents when it is desirable for the implant to dissolve essentially completely over a predetermined period of time.

[0014] For example, it can be provided that, based on a volume equal to 1 cm 3the material of the implant base body degrades in the body and / or in a simulated body fluid 10%, preferably 15%, in particular 25%, for example 30% or 40% such as 40% to 60% or 100%, faster than the material of the fastening body. Ideally, the materials of the implant base body and the fastening body are matched, taking into account the respective degrading volumes, in such a way that the fastening body sufficiently holds the implant base body during its degradation so that no larger particles from the degradation of the implant base body can enter the bloodstream. In particular, the degrading volumes and the materials can be matched in such a way that no detached particles of the implant base body can enter the bloodstream during its degradation, which could lead to thrombosis.

[0015] The implant base body is advantageously made of a metal, a composite material with predominantly metallic components, or, in particular, a metal alloy. In particular, the implant base body can also be made of a magnesium alloy. The magnesium alloy can contain or consist of the following elements by mass:

[0016] 0.2% to 1.0%, preferably 0.3% to 0.8%, in particular 0.5% to 0.7%, zinc; 0.2% to 1.0%, preferably 0.3% to 0.7%, in particular 0.4% to 0.6%, calcium; optionally further alloying elements such as titanium, boron or silicon in a total amount of less than 0.1%, preferably less than 0.05%;

[0017] The remainder is magnesium and manufacturing-related impurities. Magnesium alloys are particularly suitable for being bioresorbable and thus able to be broken down by the body. At the same time, magnesium alloys have sufficient mechanical properties for many implant applications, particularly for the treatment of heart disease. In this context, it is particularly advantageous if the magnesium alloy contains 0.2% to 1.0%, preferably 0.3% to 0.8%, in particular 0.5% to 0.7%, zinc and at the same time 0.2% to 1.0%, preferably 0.3% to 0.7%, in particular 0.4% to 0.6%, calcium by mass. It is advantageous for the zinc content to be higher than the calcium content in the alloy. Zinc contributes in particular to strength, whereas calcium also contributes in particular to corrosion resistance, so that a certain durability is ensured even in contact with corrosive media such as body fluids, in particular blood.Compared to state-of-the-art alloys, which have lower zinc and calcium contents and a zinc to calcium ratio of 1:1, a coordinated alloy concept, particularly for cardiac implants, can achieve a balanced property profile that combines sufficient durability while maintaining specified strengths on the one hand, and subsequent degradation on the other. Particularly preferred content ranges in mass percent for zinc are in the range of 0.55% to 0.70% and for calcium in the range of 0.50% to 0.68%. In addition, other alloying elements such as titanium, boron, or silicon can optionally be present in total amounts of less than 0.1%, preferably less than 0.05%. The remainder is made up of magnesium and manufacturing-related impurities. The content of manufacturing-related impurities is typically less than 0.01%, preferably less than 0.005%.

[0018] It is also advantageous if the magnesium alloy is free of rare earth metals, with the exception of manufacturing-related impurities. As mentioned, the manufacturing-related impurities are preferably less than 0.005%. Based on the rare earth metals, the corresponding manufacturing-related impurities are particularly preferably less than 0.001% by mass, particularly preferably less than 0.0005%.

[0019] The iron content should be less than 100 ppm, preferably less than 50 ppm, to avoid accelerating corrosion of the magnesium alloy. The shape of the implant base body depends on the specific application. For example, the implant base body can be designed as a tube. It is also possible, particularly for rings, nails, wires, or the like, for the implant base body to be made from solid material, in particular a rod, preferably an extruded rod, which may have been formed once or multiple times using equal channel angular pressing (ECAP). For various diameters, a homogeneous microstructure results from manufacturing a rod, in particular an extruded rod. Such a microstructure can be further improved using ECAP, and also in terms of strength, if desired.Ultimately, this depends on the desired mechanical and corrosion properties, as no steps need to be carried out that make the material even better than absolutely necessary. For small diameters, a drawing process can also be used, whereby a previously extruded rod material can serve as the starting material, for example. Wire drawing can be used to create implants with a diameter of less than 5 mm, and in particular less than 3 mm. Other processes such as hammering can also be used if small implant diameters are to be achieved. Rods or other pre-products can also be formed if the material is suitable, for example to introduce saddle-shaped structures along a rod's longitudinal axis for a mitral annulus replacement, which later enables better coaptation of valves.Magnesium alloys as described above are particularly suitable for this purpose due to their very good formability.

[0020] For many applications, it is sufficient for the implant base body to have a substantially constant cross-section. If the base body is made from a single material, such as a bar, it may also have undergone forming processes, such as bending. It is also possible for the implant base body to be machined from a semi-finished material by machining, such as milling or turning.

[0021] The advantages of the invention become particularly apparent when the implant base body is an annuloplasty ring, for example a mitral ring, in particular an open mitral ring, or a tricuspidal ring. A mitral ring can, for example, be inserted into the heart of a human or, if appropriate, an animal using the fastening body. The implant is fastened accordingly using the provided fastening body. The mitral ring itself is approximately ring-shaped and can be open, i.e., as a non-closed ring. This allows, in particular, the mitral ring to be formed particularly easily from a pre-cut rod, with the pre-cut rod section being bent after cutting to form the mitral ring. The mitral ring is then combined with the fastening body, with the implant additionally comprising the antibacterial substance and, in this configuration, can be surgically inserted.

[0022] To optimize its properties, the implant base body can be heat-treated. Heat treatment can be advantageous, for example, if the implant base body, particularly an annuloplasty ring such as a mitral or tricuspidal ring, exhibits high internal stresses due to the manufacturing process. For example, the implant base body can be formed from a heat-treated magnesium alloy, preferably with a hardness of at least 55 HV2, preferably at least 60 HV2, and / or a tensile strength of at least 200 MPa. Higher HV2 values ​​of at least 65 HV2 and / or tensile strengths of more than 230 MPa or more than 250 MPa are also possible.

[0023] For many applications, it is advantageous if the implant base body is essentially ring-shaped. In particular, the implant base body can be designed as an open ring with free ends. In this case, the distance between the free ends of the open ring can be less than 10%, preferably less than 5%, in particular less than 3.5%, for example 0.5% to 3.0%, of the circumference of the open ring. The circumference includes a gap between the two opposite ends of the open ring. Spacing the free ends can have a positive effect on the overall flexibility of the implant.

[0024] It is also advantageous if the implant base body has a Young's modulus of less than 80 GPa, preferably less than 60 GPa, in particular less than 55 GPa, for example 20 GPa to 50 GPa. Compared to non-degrading implants made of titanium or a titanium alloy, implant base bodies made of magnesium or a magnesium alloy have a significantly lower Young's modulus. This reduces, for example, the maximum and mean mitral valve pressure gradient with such an implant and results in less pericardial effusion. This also leads to improved hemodynamics in the first 12 months after implantation, which is beneficial for patients. Furthermore, implant base bodies with a higher Young's modulus are more prone to dehiscence than flexible devices.

[0025] The term "fastening body" is to be understood broadly within the scope of the present disclosure. The fastening body can, for example, be a thread with which the implant base body is fixed, for example, to surrounding tissue, for example, in the heart, or to a bone portion. However, the fastening body can also be designed, for example, as a flexible or rigid mesh that is fully or partially connected to the implant base body. Preferably, the fastening body is flexible. In particular, the fastening body has greater flexibility than the implant base body.

[0026] The fastening body advantageously accommodates the implant base completely or at least predominantly, especially when it is an annuloplasty ring, such as a mitral ring. With a mitral ring, the fastening body is pushed over the implant base so that the mitral ring can be secured in the appropriate position in the heart via the fastening body. To enable the implant base to be accommodated accordingly, the fastening body can generally be tubular. In particular, for special applications in the area of ​​a human heart, the fastening body can be configured to be attached to a mitral complex above commissures.

[0027] It is preferred that the fastening body essentially completely surrounds the implant base body. This ensures that no detached particles of the implant base body enter the bloodstream and could lead to thrombosis there. The fastening body preferably degrades so slowly, relative to the degradation of the implant base body, that at least no particles with a particle size larger than 5 μm enter the bloodstream, in particular particles with a particle size larger than 5 μm, which would lead to thrombosis. Smaller detached particles can, if necessary, be broken down by the body or do not pose a danger. The fastening body is preferably designed with pores. The term pores means that the fastening body is porous, i.e. not completely densely packed. The pores can be designed with a pore size distribution.The pore volume fraction is preferably 75% to 95%, in particular 85% to 90%. Such a porous structure can also be created in electrospinning by superimposing individual spun threads, resulting in a porous structure with different pore sizes. Such a porous structure is preferred for a balanced property profile with regard to biodegradability on the one hand and a filter for detached particles from the degrading implant base body on the other. Pores of different sizes and / or a predetermined surface roughness also promote implant ingrowth.The following are involved in the ingrowth of an implant in the area of ​​a mitral valve or other heart valves: fibroblasts, which prefer pore sizes of 5 pm to 15 pm, endothelial cells, which prefer pore sizes of 10 pm to 20 pm, muscle and nerve cells, which prefer pore sizes of 20 pm to 100 pm, and macrophages and immune cells, which prefer pore sizes of 0.5 pm to 10 pm. To promote ingrowth of the implant, varying pore sizes in the range of 1 pm to 30 pm are therefore ideal. An average pore size, measured across a distribution of visible pores in a scanning electron micrograph (SEM image), taking the largest diameter of a pore into account, should be less than 50 pm or 25 pm, preferably less than 20 pm, in particular less than 10 pm, for example less than 5 pm.

[0028] The fastening body can be formed with an adjacent flap that allows for attachment of the implant. For example, the fastening body can completely enclose the implant body and be sewn into the body via a flap. The flap can have openings designed to allow suture material to pass through during surgery.

[0029] With regard to the required bioresorbability and flexibility, it is advantageous if the attachment body is made of a polymer. The polymer can be designed so that it dissolves in the body in a manner coordinated with the dissolution of the implant base body. The attachment body can be formed from a polymer deposited by electrospinning. Electrospinning creates micro- and / or nanofibers, i.e. fibers with a specific length and / or a diameter in the range of micrometers down to a few nanometers. The selected electrospinning process creates micro- and nanofibers with a diameter of up to 5 pm, preferably 0.6 pm to 3.5 pm, which are randomly distributed per layer.In addition to the aforementioned preferred pore size between the fibers of 5 pm to 50 pm, the diameters of the micro- and nanofibers in the range of 0.6 pm to 3.5 pm are also responsible for improved cell migration into the material. It has been shown that a correspondingly deposited attachment body can be advantageous, particularly with regard to the uptake of the antibacterial substance and its subsequent release into the body.

[0030] In addition to electrospinning, melt-electro-writing can be particularly advantageous for producing a fastening body, whereby, if necessary, several layers can be deposited on top of each other in order to cover an ideal pore size range.

[0031] It is particularly advantageous if the fastening body is made of or with a polycaprolactone or a derivative thereof. Polycaprolactone has a relatively low melting point of 60 °C with an average molecular weight (Mn) of 80 kDa. This allows polycaprolactone to be processed in various ways, for example, by casting, drop-deposition, or the aforementioned electrospinning or melt electrowriting. In principle, however, any biodegradable polymer can be used.

[0032] The antibacterial substance advantageously comprises or consists of nanoparticles. Nanoparticles are particles that have an average diameter of less than 100 nm in at least one length dimension (x, y, and / or z direction). It has proven particularly useful for the antibacterial substance to contain or consist of silver nanoparticles. Silver nanoparticles can effectively prevent bacterial infections in the area of ​​the implant after surgery or at least reduce the likelihood of them. Another advantage is that the nanoparticles damage bacterial membranes and are therefore effective against both gram-positive and gram-negative bacteria. The nanoparticles used in the invention are preferably produced by laser ablation. Furthermore, the nanoparticles can be ultrapure, i.e., with a purity of more than 99.999%.The nanoparticles are preferably essentially spherical.

[0033] The antibacterial substance can, for example, be applied to the surface of the implant base body, for example, by directly coating it. However, it is preferred if the antibacterial substance is distributed in and / or on the fastening body, in particular homogeneously distributed throughout the entire volume of the fastening body. It is particularly effective if the antibacterial substance is processed together with a base material of the fastening body, for example, a polymer such as polycaprolactone. As a result, the antibacterial substance is not only homogeneous and free of agglomerates in the fastening body, but is also released in appropriate doses over a long period of time as the fastening body dissolves.In other words: With the dissolution of the bioresorbable base body, there is also a continuous, ongoing local release of the antibacterial substance, so that an antibacterial effect is present not only immediately after an operation, but also during the healing process.

[0034] If the antibacterial substance is provided distributed in and / or on the fastening body, it is advantageous if it is present in a proportion of 0.10% to 6.0% by mass, preferably 0.20% to 4.5%, in particular 0.30% to 4.0%, for example 0.50% to 1.5%, based on the total mass of antibacterial substance and fastening body. A certain minimum content of antibacterial substance is necessary to achieve the desired effect. Excessively high contents of antibacterial substance are more likely to lead to complications. In this respect, medium contents of, for example, 0.20% to 1.0% (in mass percent) of antibacterial substance are particularly preferred. On the one hand, the desired antibacterial effect is achieved, and on the other hand, complications from the release of the antibacterial substance, for example nanoparticles, are minimized.

[0035] In addition to nanoparticles, antibiotics can also be used alternatively or in addition, especially non- or poorly water-soluble antibiotics such as rifampicin, since such antibiotics are only released during the degradation of the attachment body due to their poor or non-existent water solubility.

[0036] The further object of the invention is achieved by a bioresorbable composite material for an implant which is implanted in the human or animal body, consisting of or comprising a bioresorbable carrier substance and an antibacterial substance, wherein the composite material is electrospun and / or present as a coating.

[0037] It has been recognized that a particularly effective release of an antibacterial substance in a composite material used in connection with implants in the human or animal body can be achieved if the composite material is electrospun and / or present as a coating. Electrospinning can create particularly fine threads in the nanometer range, between which the antibacterial substance is then present. Similarly, preferably particularly thin layers of a few microns on heart valves or other implants, particularly those inserted in a heart, can lead to a rapid release of the antibacterial substance, which is continuous during the degradation of the bioresorbable carrier substance. The coating can also be produced by electrospinning, 3D printing, or foaming.In particular, electrospinning can be used to achieve good mechanical properties for the composite material in terms of strength and sewability combined with long-term stability.

[0038] It is preferred if the carrier substance is meltable at temperatures below 100°C. This allows for relatively simple processing of the composite material, especially since the carrier substance usually represents the predominant component. This is especially true if the carrier substance is a polymer, especially polycaprolactone or a derivative thereof.

[0039] The antibacterial substance can comprise or consist of nanoparticles. The antibacterial substance can contain metallic nanoparticles or consist exclusively of them. The nanoparticles can be selected, for example, from the group consisting of silver, copper, zinc, and magnesium. The nanoparticles are preferably produced by laser ablation in a suitable solvent to achieve high nanoparticle purity. Silver nanoparticles are advantageously used. The nanoparticles can also be produced directly in a dissolved polymer or monomer solution.

[0040] The antibacterial substance can be present in the bioresorbable composite material in a mass percentage of 0.10% to 6.0%, preferably 0.20% to 4.5%, in particular 0.30% to 4.0%, 0.50% to 1.5%, based on the total mass of the antibacterial substance and carrier substance. Preferred proportions are in the range of approximately 0.15% to 1.75%.

[0041] In a further aspect, the invention relates to a method for treating a patient with a disease such as heart valve disease or for prophylaxis thereagainst, wherein an implant is inserted into the patient, preferably for treating a patient with a diseased or defective mitral valve, in particular with endocarditis or a high risk of endocarditis, wherein the method comprises the following steps: a) providing an implant according to the invention; b) surgically implanting the implant in the patient.

[0042] In a method according to the invention, the above-described advantages of an implant according to the invention are fully realized. To protect the patient, implantation is preferably performed minimally invasively via a catheter.

[0043] Further features, advantages, and effects in connection with the present disclosure will become apparent from the following exemplary embodiments. Reference is made to the drawings, which show:

[0044] Fig. 1 a commercially available mitral ring with a surrounding sleeve;

[0045] Fig. 2 a 3D model of an open mitral annulus;

[0046] Fig. 3 non-heat-treated mitral rings made of different alloys and depending on a period of time in contact with simulated body fluid;

[0047] Fig. 4 Hardness values ​​HV2 for two alloys as a function of heat treatment; Fig. 5 shows the dependence of the diameter of heat-treated and non-heat-treated mitral rings as a function of the duration of contact with dilute hydrochloric acid;

[0048] Fig. 6 shows a constant concentration of silver nanoparticles in a carrier substance before and after processing of the carrier substance;

[0049] Fig. 7 an electrospun sleeve on a mitral annulus;

[0050] Fig. 8 and Fig. 9 scanning electron microscopy (SEM) images of an electrospun sleeve;

[0051] Fig. 10 is a diagram showing the tensile strength and elongation of rods made of a magnesium alloy of type ZX00 extruded to different diameters;

[0052] Fig. 11 is a diagram of a fatigue cycling test of a magnesium alloy of type ZX00 with a stress amplitude versus the number of cycles to failure, the test duration and a corresponding implantation duration at an average heart rate;

[0053] Fig. 12 Results of 3-point bending tests on non-implanted pins made of ZX00 and WE43 alloys;

[0054] Fig. 13 Results of 3-point bending tests on pins made of a ZX00 alloy before (AR) and 22 weeks after implantation;

[0055] Fig. 14 Results of 3-point bending tests on pins made of ZX00 and WE43 alloys before (AR) and 22 weeks after implantation;

[0056] Fig. 15 Comparison of force-displacement curves from 3-point bending tests of sterilized and non-sterilized pins of type ZX00;

[0057] Fig. 16 Results of 3-point bending tests on pins made of a ZX00 alloy after different degradation times in simulated body fluid;

[0058] Fig. 17 Maximum force achieved in 3-point bending tests for ZXOO pins and WE43 pins depending on a degradation time of up to 5 weeks in simulated body fluid;

[0059] Fig. 18 Elastic limit achieved in 3-point bending tests for ZXOO pins and WE43 pins depending on a degradation time of up to 5 weeks in simulated body fluid;

[0060] Fig. 19 Diagram of cell viability at different concentrations of silver nanoparticles;

[0061] Fig. 20 Diagram of the OD value at different concentrations of silver nanoparticles; Fig. 21 X-ray images of two rats each with four pins made of magnesium alloys of type WE43 and ZX00 at different times after implantation;

[0062] Fig. 22 Photos of explanted WE43 and ZXOO pins;

[0063] Fig 23 Light micrographs of magnesium alloy pins (WE43 and ZX00) and surrounding skin tissue;

[0064] Fig. 24 an inserted Mg mitral ring after 12 months of implantation in a sheep heart;

[0065] Fig. 25 a histology image of a Mg mitral annulus after 12 months of implantation;

[0066] Fig. 26 to Fig. 28 diagrams of a pericardial effusion and pressure gradients at the mitral valve in comparison of an implant according to the invention with a permanent implant.

[0067] 1. Production of mitral rings

[0068] The alloys Mg-0.6Zn-0.5Ca (hereinafter referred to as ZX00) were defined as the starting material for bioresorbable mitral rings and used as extruded rod material with 6 and 12 mm diameters. The commercially available magnesium-based alloy WE43 was used as a reference material.

[0069] The design considerations for the mitral rings included closed and open rings, with and without closure systems. A commercially available Carpentier-Edwards classic mitral ring (which consists primarily of non-resorbable titanium) was selected as a template for the development of bioabsorbable mitral rings made of magnesium alloys. The different mechanical properties of titanium and magnesium, or alloys with these metals as the base alloy component, were taken into account. A commercially available mitral ring, as shown in Fig. 1, was measured, and a 3D model was created based on these measurements; this model is shown in Fig. 2.

[0070] Based on the 3D model created, a bending tool was manufactured, and mitral rings made of a magnesium alloy were fabricated from rods with a diameter of 2.5 mm and a length of approximately 92 mm at room temperature. References to Mg mitral rings in the following refer to mitral rings made of a magnesium-based alloy, i.e., an alloy with magnesium as the main component. Degradation analyses of the Mg mitral rings in simulated body fluid (SBF for short) showed that after a few weeks, the Mg mitral rings manufactured as described above began to bend as they became thinner due to degradation. This indicates that internal stresses build up in the Mg mitral rings as a result of the bending of the rods.The condition of corresponding Mg mitral rings at the beginning of a degradation analysis in SBF and after 12 weeks is shown in Fig. 3 for a mitral ring made of the frequently used alloy WE43 and the alloy ZX00.

[0071] Internal stresses can be relieved through appropriate heat treatment after bending of the Mg mitral rings. The heat treatment should not impair the mechanical properties of the Mg alloys. To verify this, the hardness of the heat-treated Mg mitral rings was measured at different temperatures, as shown in Fig. 4. The optimal heat treatment temperature for ZX00 was found to be approximately 150°C to 230°C, in particular approximately 200°C, for approximately 1 hour, and for WE43 approximately 250°C for 1 hour. This almost completely eliminates the effect of bending. No significant difference in the degradation behavior of the Mg mitral rings was observed, as can be seen in Fig. 5 for the alloys tested when comparing heat-treated and non-heat-treated Mg mitral rings.

[0072] 2. Development of functional surface coatings

[0073] Pulsed laser ablation in liquids was used to generate ultra-pure nanoparticles from various materials. In the first step, several solvents (dimethylformamide, tetrahydrofuran [THF], dichloromethane, acetone, chloroform, and dimethyl sulfoxide) were analyzed to select the most suitable ones. Acetone and THF were chosen as solvents due to their optimal vapor pressure and lower toxicity, and silver nanoparticles (AgNPs) were selected as a potential antimicrobial additive. The process flow was optimized for nanoparticle stability, starting with laser ablation, through storage of the AgNPs, and viscosity adjustment by solvent evaporation. This also applies to the optimal amount of polycaprolactone (PCL) added to the solvent for nanoparticle production, as well as the optimization of the rate of nanoparticle synthesis and particle concentration in the polymer matrix.Finally, optical evaluation (UV-VIS absorption) was performed by casting AgNP-PCL films to assess particle quality. A doping of 3.6 wt.% AgNP in PCL was achieved with the silver nanoparticles. However, according to the literature, this amount is cytotoxic and was therefore limited to approximately

[0074] 2 wt% AgNP or less. The best results were achieved with acetone and with 1% PCL dissolved in acetone, as AgNP agglomeration was less pronounced in this solvent.

[0075] Furthermore, synthesis protocols for various PCL nanoparticle composites, such as silver, zinc, copper, and magnesium nanoparticles, were developed, and sample batches were prepared. Direct coating of the Mg mitral rings with AgNPs was also tested, but this yielded less satisfactory results. Therefore, the antibacterial properties induced by the AgNPs were incorporated by incorporating the AgNPs into the PCL used for electrospinning. This has the advantage of eliminating the need for an additional coating process, which could affect the consistent quality of the Mg mitral rings.

[0076] Subsequently, tests were conducted to determine whether the particle concentration remained unchanged in the electrospun material or whether there were any changes. As can be seen in Fig. 6, the results of UV-VIS optical spectral analyses show that the particle concentration in the granules and the electrospun material is virtually unchanged.

[0077] To attach the Mg mitral rings to the heart, meshes with different pore sizes were produced using a 3D printer suitable for PCL. However, it was subsequently determined that the 3D-printed sleeves no longer possessed sufficient mechanical quality and stability after just a few weeks, so an alternative process was used to produce the holder for the Mg mitral rings. Electrospinning was shown to be a suitable method for producing sleeves with sufficient mechanical stability over the required time periods. The electrospun sleeves not only offer sufficient mechanical stability and sewability, but also serve as a suitable functional basis for encapsulating antibacterial nanoparticles, such as those made of silver.

[0078] Fig. 7 shows a sleeve produced by electrospinning. Figs. 8 and 9 show scanning electron micrographs illustrating the microstructure of such sleeves. PCL material with up to 3.6 wt.% AgNP also lends itself well to electrospinning. For the production of the final meshes for securing the Mg mitral rings, PCL granules doped with 0.5 wt.% AgNP were produced by electrospinning.

[0079] 3. In vitro tests

[0080] 3.1. Mechanical tests and degradation experiments

[0081] The maximum forces acting on a mitral annulus in the heart were determined from relevant specialist publications. Typical forces acting on the mitral annulus at maximum left ventricular pressure of 100 mm Hg are (4.9 ± 2.0 N vs. 2.1 ± 1.1 N), at 125 mm Hg (5.4 ± 2.3 N vs. 2.3 ± 1.2 N), and at 150 mm Hg (5.7 ± 2.4 N vs. 2.4 ± 1.1 N). Since these forces are very low compared to the material characteristics after extrusion of the Mg alloys or after double ECAP forming, extruded rod material with a 6 mm diameter was subsequently used for the production of the Mg mitral annulus and pins instead of material with a 12 mm diameter before and after double ECAP forming. For the in vitro biological studies, rod material with a diameter of 12 mm was used. The material characteristics of the extruded rod material with a diameter of 6 mm compared to the extruded rod material with a diameter of 12 mm are shown in Fig.10. The 6 mm material has higher tensile strength (290 MPa vs. 230 MPa) and hardness (67 HV2 vs. 58 HV2), but lower ductility and less work-hardening capacity than the 12 mm material. Therefore, the material is more difficult to bend into the shape of the Mg mitral ring at room temperature; fractures of the Mg mitral rings have been observed in some cases. To avoid fractures, the Mg mitral rings can be formed at elevated temperatures and / or created by wire drawing. For wire drawing, a diameter of 2.5 mm, for example, can be selected, which leads to improved bending behavior. Subsequently, results from fatigue tests in simulated body fluid (SBF) were used to evaluate whether the heartbeat could have a negative influence on the mechanical stability of the Mg mitral rings. The results of the previous tests are shown in Fig.11 and demonstrate that a corrosive environment, such as that found in the body, significantly reduces fatigue strength. Under SBF conditions, a significantly lower fatigue limit (30 MPa to 40 MPa) was observed than under room air conditions (145 MPa to 170 MPa). However, this fatigue limit is still significantly higher than the maximum load of the Mg mitral rings in the heart and is therefore sufficient to test the Mg mitral rings on living subjects.

[0082] 3.2. 3-point bending test on pins (initial condition and after 22 weeks of implantation)

[0083] To assess the stability of pins after 22 weeks of implantation in a subcutaneous rat model, a 3-point bending test was set up.

[0084] Fig. 12 shows the results of the 3-point bending tests with non-implanted ZX00 and WE43 pins. It is clearly visible that WE43 pins are stronger due to a significantly higher alloy content, but two pins fractured significantly earlier than ZX00 pins. The mean maximum force for ZX00 is 89 ± 1 N and for WE43 98 ± 4 N. The elastic range is approximately the same for both alloys, ranging from 45 N to 50 N.

[0085] Fig. 13 shows the results of 3-point bending tests with pins made of the ZX00 alloy, where the pins had previously been implanted for 22 weeks. It is clearly visible that these pins have a lower bending strength than the original pins (-37%), which was also expected, but the ductility has remained the same or even improved slightly. The mean maximum force for pins made of the ZX00 AR alloy (AR corresponds to "as received," i.e., in the extruded state) is 89 N ± 1 N and for ZX00 in vivo it is 56 N ± 3 N. The elastic range for ZX00 in vivo is lower or has decreased, but is still approximately 30 N, which is still well above the maximum stress in the heart.

[0086] Fig. 14 shows results from 3-point bending tests on pins made of WE43 alloy, where the pins were implanted for 22 weeks. It is clearly visible that the in vivo tested WE43 pins exhibit lower bending strength (-31%), which was also expected. However, in contrast to the in vivo pins made of ZX00 alloy, ductility also decreased significantly, which can lead to premature mechanical failure. The mean maximum force for WE43 AR pins is 98 N ± 4 N and for WE43 in vivo pins is 68 ± 5 N. The elastic range for WE43 in vivo is only slightly lower in comparison, at approximately 40 N.

[0087] To investigate any potential influence of gamma sterilization on mechanical properties, 3-point bending tests were also conducted on non-implanted pins prior to sterilization. ZXOO pins had the same force-displacement curves before and after sterilization, albeit with slightly different shapes. There was no difference in the force-displacement curves for WE43 pins, but the non-sterile WE43 pins were more likely to fail earlier. This is evident in Fig. 15.

[0088] 3.3. Influence of cleaning and sterilization on ZXOO pins

[0089] Some pins that were not required for implantation and were sterile packaged and sterilized were found to show signs of corrosion on the surface after some time. This effect was systematically investigated to determine the cause.

[0090] For this purpose, ZX00 alloy plates were immersed in various cleaning agents such as distilled water, isopropanol, ethanol, or acetone and examined under a light microscope after 1 hour and 4 hours. Except for distilled water, as expected, no changes to the surface were observed with the different cleaning agents. Therefore, the observed traces of corrosion could not have been caused by a faulty cleaning process of the pins.

[0091] It was therefore assumed that environmental influences could have led to these corrosion traces. To investigate this, a ZXOO platelet was exposed to an environment with very high humidity for several days. This sample was placed in a sealed plastic bag along with a small bowl of water. It was observed that at higher humidity, corrosion traces similar to those previously observed form on the surface. In implants subject to mechanical stress, this can trigger cracking and lead to premature failure. When cleaning and packaging Mg implants, care must be taken to ensure that the humidity is as low as possible and that the implant is packaged with as little ambient air as possible.

[0092] 3.4. Degradation and 3-point bending tests on in vitro degraded pins

[0093] A comprehensive in vitro degradation test was conducted for 5 weeks in SBF with Tris / HCl buffer on sterilized pins made of the alloys ZX00 and WE43. Each bottle filled with 250 ml of SBF contained three to four pins suspended without contact with the bottle wall. All bottles were placed in a water bath heated to body temperature (36.5 °C) and sealed with Parafilm. The initial pH of the SBF was between 7.35 and 7.45. The SBF was changed every seven days.

[0094] The degradation results at different time points (2, 4, and 5 weeks) were visually assessed. Visual inspection showed that the degradation proceeded as observed in preliminary tests, i.e., a white surface was obtained for the ZXOO pins and a black surface with falling black particles was obtained for the WE43 pins. However, the degradation occurred more rapidly and with greater variance from pin to pin than expected, which is why the test was terminated after only five weeks.

[0095] Three-point bending tests were also performed on the in vitro degraded pins to evaluate their mechanical behavior at different degradation times. The results for ZXOO pins are shown in Fig. 16 and can be summarized as follows: Strength decreased with increasing degradation time, while ductility remained almost unchanged until week 4. None of the pins fractured during the test, but there was a larger scatter in the force-displacement curves after 5 weeks.

[0096] The evaluation and summary of the maximum forces and elastic limit from the force-displacement curves is shown in Fig. 17 and Fig. 18. The ZX00 pins have lower strength than the WE43 pins, both in the undegraded state and after various stages of degradation. However, the WE43 pins show a higher variation in strength values ​​after just 2 weeks, but especially after 4 and 5 weeks. High scatter is undesirable for implants, so the ZX00 alloy is preferable to the WE43 alloy for this reason as well. The elastic deformation is analogous to the maximum force or strength for both alloys and is shown in Fig. 18.

[0097] 3.5. Biological Tests - Biocompatibility

[0098] Samples were prepared for biocompatibility testing of the PCL (80 kDa) sleeve material with varying levels of silver nanoparticles (AgNP). A methodological approach was planned based on an adapted test protocol for determining antibacterial activity on plastic and other non-porous surfaces based on the ISO 22196:2011 standard (Plastic Materials). This protocol relies on surface rather than mass testing. Escherichia coli and Staphylococcus aureus were selected as the bacterial strains. For cytotoxicity testing, electrospun meshes were mounted on so-called CellCrowns (available at www.scaffdex.com).

[0099] A cytotoxicity test of the electrospun PCL-Ag nanocomposites using XTT tests revealed that a concentration of 4 wt% AgN P doping resulted in a significant decrease in cell viability (see Fig. 19 and Fig. 20). No significant differences in cell viability were observed at the lower concentrations. Therefore, the focus was subsequently on the concentrations of 0.1 wt%, 0.5 wt%, and 1 wt%.

[0100] To determine the final concentration of silver nanoparticle doping in PCL for the meshes of the Mg mitral rings, the following AgNP-doped PCL starting materials were used for electrospinning: 0 wt% AgNP, 0.1 wt% AgNP, 0.5 wt% AgNP, and 1 wt% AgNP. The results of the determination of antibacterial activity in a certified testing laboratory are presented in the following table for PCL doped with different concentrations and for both bacterial strains (E. coli and S. aureus). It was shown that PCL doped with 0.5 wt% AgNP achieved a log reduction factor R of > 2 for sufficient antibacterial effectiveness for both bacterial strains (see Table 1 below). This concentration was therefore used to produce the final meshes for the Mg mitral rings. Table 1: Antibacterial activity R of tested samples 3.6. In vivo tests

[0101] An ethics application to conduct a small animal study was submitted and approved. These are tolerability studies in which the individual components (Mg pin, sleeve, and the entire Mg mitral annulus) are implanted subcutaneously in sequential phases. These preclinical tolerability studies were initially conducted in rats. Mg pins were implanted subcutaneously on the back of the rats. The two Mg alloys ZX00 and WE43 were used for the pins, and four pins of the same Mg material were implanted in a total of eight rats. During the first two weeks, X-ray images were taken several times to monitor degradation, possible blistering, and unilateral fixation of the implants. Thereafter, X-ray images were taken once a week, as shown in Fig. 21.No gas bubble formation was observed, and all pins, except for one WE43 pin, remained stable at the implantation site until week 17. This one pin was apparently removed by another rat, as sutures had torn. No measurable degradation was observed on X-ray during the first 17 weeks. All animals were in good health. The implantation period was extended. The pins were explanted at week 22 after implantation.

[0102] Images of the pins during explantation of WE43 pins in rat 1 and ZX00 pins in rat 2 after 22 weeks of implantation showed the following features: In situ, there were no gross signs of degradation of the pins. This confirms the radiographic findings. There were also no visible signs of inflammation. The implants were surrounded by a thin tissue capsule. Only in one ZX00 pin were very small gas bubbles visible. With the WE43 pins, it was observed during explantation that the tissue adhered less than with the ZX00 pins.

[0103] Figure 22 shows stereomicroscopic images of explanted WE43 and ZXOO pins. The images were taken after 24 h of fixation with 2.5% glutaraldehyde to preserve the biological tissue samples on the surface.

[0104] The explanted pins were then stored in a dry place to prevent further degradation. There was no change in the length of the pins. The WE43 pins had a very uniform black surface compared to a heterogeneous, predominantly white surface of the ZXOO pins. Small, localized corrosion hot spots were present on the ZXOO pins. A comparison of the weights of the WE43 and ZXOO pins before implantation and after 22 weeks of implantation yielded the following values: WE43 before implantation: 43.0 mg, and after 22 weeks: 42.2 mg to 42.8 mg; ZX00 before implantation: 40.0 mg, and after five weeks: 36.7 mg to 39.1 mg. The difference in weight before implantation can be explained by the fact that the WE43 alloy contains heavy rare earth elements. It is also evident that the weight loss is greater for the ZXOO pins than for the WE43 pins. This was already the case in the in vitro experiments.

[0105] Light microscopy images of the explanted ZXOO pins after 22 weeks showed that the pins, including the degradation layer, remained almost the same in diameter when viewed purely visually. The hole for attaching the pins became larger in some pins. The degradation layer on the ZXOO pins looks significantly different than on the WE43 pins and is also significantly thicker. The degradation layer on the ZXOO pins partially resembles an organic surface and ranges in color from gray to white to yellow.

[0106] After 22 weeks of implantation, WE43 pins display a black surface, just like before insertion, with some processing marks still visible. The diameter appears to have barely changed, and in some places, pock-like structures are visible, which could be degradation products. At higher resolution, many small cracks are visible in the degradation layer, and in some places, this layer has crumbled away, creating dents approximately 20 to 25 pm deep.

[0107] The pins, which had already been examined under light microscopy, were then tested for their mechanical stability in 3-point bending tests. The results are presented above.

[0108] The rats' venous blood was examined. No significant differences were observed between the two implant materials. The liver enzymes alkaline phosphatase (AP), aspartate aminotransferase (AST), and alanine aminotransferase (ALT) were within the normal range compared to untreated healthy animals, according to the reference values ​​in the publication "Reference values ​​for selected hematological, biochemical, and physiological parameters of Sprague-Dawley rats at the Animal House, Faculty of Medicine, University of Colombo, Sri Lanka, Shehani L. Delwatta et al., Animal Model Exp Med. 2018." Only the AP values ​​for WE43 and ZX00 were below the limit, with ZX00 being slightly closer to the limit than WE43. The kidney enzymes creatinine and blood urea (BUN) were also within the normal range, with the BUN value being in the upper limit range for both alloys.No significant differences were found between the two Mg alloys in the blood count and in the magnesium levels in the blood. No significant differences were found between the two Mg alloys in the blood count and in the magnesium levels in the blood. The histological images confirm adequate integration into the adjacent tissue with a low inflammatory reaction. Furthermore, the degradation was very homogeneous, and gas production was very low and constant for most implants. Histopathological analyses of the subcutaneous tissue surrounding the implant were performed, and the excised tissue was stained as required. No inflammation was detected in the skin tissue (Fig. 23).

[0109] The microscopic findings observed during extensive histopathological analyses of the brain, heart, lungs, liver, kidneys, lymph nodes, and spleen were mostly minor and are among the spontaneous background findings that occur in untreated rats of this age. The histopathological analyses therefore revealed no abnormalities compared to healthy rats, so good tolerability can be assumed for both Mg alloys over a period of 22 weeks or more.

[0110] 3.7. Acute trials in pigs and implantation in sheep

[0111] The acute test in pigs was conducted to confirm the surgical suturability of the electrospun meshes in vivo, and to test the mechanical stability of the ZX00 Mg mitral ring with a PCL mesh during and shortly after surgery. Analysis of the explanted Mg mitral rings after the acute test with the electrospun meshes showed that both requirements—fixation of the Mg mitral rings in the heart and the operability of the Mg mitral rings with PCL mesh—were met very well.

[0112] For the sheep study protocol, a control group of two sheep was defined with a standard mitral ring, commercially available from Edwards, and a group of five sheep was selected with the final Mg mitral ring with a PCL coating doped with 0.5 wt% AgNP. The heat-treated Mg mitral rings were first cleaned with isopropanol in an ultrasonic bath and then coated with an electrospun PCL material doped with 0.5 wt% AgNP. These Mg mitral rings were packaged in a sterile package and gamma sterilized.

[0113] The implantation of the Mg mitral rings made of the ZX00 alloy in five sheep proceeded without major problems, although the surgical technique presented some challenges when operating on the beating heart. The five sheep recovered quickly after surgery, and no clinical abnormalities were detected during the follow-up period. Continuous ultrasound examinations were performed to ensure the adequate stability of the implants according to the invention and their degradation. Furthermore, the postoperative condition of the animals was monitored using blood count and blood chemistry analyses, with very satisfactory results achieved in all animals.

[0114] 3.8. Long-term trials in sheep

[0115] The images of the implanted Mg mitral rings with an AgNP-doped PCL shell in the sheep model show excellent ingrowth after 12 months, immediately prior to explantation. All Mg mitral rings with an AgNP-doped PCL shell are surrounded by fibrous tissue, which stabilizes the annulus. No signs of perforation or thrombosis are visible, as documented in Figure 24.

[0116] Histological images of the implanted Mg mitral rings after HE staining show excellent ingrowth (Fig. 25). Endothelialization is visible in all Mg mitral rings. No signs of thrombi, perforation, or air inclusions are visible.

[0117] The Mg mitral rings had degraded by approximately half after 12 months. No critical magnesium particles (larger than 5 μm) were released, as the degraded magnesium rings are surrounded by the AgNP-doped PCL shell on the one hand and by fibrous tissue on the other, thus preventing release. No signs of thrombosis were observed in any of the sheep, indicating that the entire implant is sufficiently stable for direct cardiac implantation.

[0118] Pericardial effusion with mitral rings according to the invention is significantly less after implantation than with permanent annuloplasty rings, which means less stress on the heart. Both the maximum and mean mitral valve pressure gradients are significantly lower than with permanent annuloplasty rings, meaning no stenosis occurs, whereas with permanent annuloplasty rings, signs of mild stenosis are evident. This is demonstrated by a comparison of the pericardial effusion shown in Fig. 26, as well as the maximum pressure gradient at the mitral valve (Fig. 27), and the mean pressure gradient at the mitral valve (Fig. 28). The comparison data ("crtl") refer to a permanent implant, the other data ("Mg") to an implant according to the invention.

Claims

Patent claims 1. An implant for the medical treatment of humans and / or animals, comprising or consisting of: a) a bioresorbable implant base body; b) a bioresorbable fastening body connected to the implant base body and by means of which the implant base body can be fastened in a human or animal body; c) optionally an antibacterial substance.

2. Implant according to claim 1, wherein the implant base body is formed from a metal, a composite material with predominantly metallic components or in particular a metal alloy.

3. Implant according to claim 1 or 2, wherein the implant base body is formed from a magnesium alloy.

4. Implant according to claim 3, wherein the magnesium alloy comprises or consists of the following elements in mass percent: 0.2% to 1.0%, preferably 0.3% to 0.8%, in particular 0.5% to 0.7%, zinc; 0.2% to 1.0%, preferably 0.3% to 0.7%, in particular 0.4% to 0.6%, calcium; optionally further alloying elements such as titanium, boron or silicon in a total amount of less than 0.1%, preferably less than 0.05%; Rest magnesium and manufacturing impurities.

5. Implant according to claim 3 or 4, wherein the magnesium alloy is free of rare earth metals except for impurities caused by manufacturing.

6. Implant according to one of claims 1 to 5, wherein the implant base body is an annuloplasty ring, for example a mitral ring, in particular an open mitral ring.

7. Implant according to one of claims 1 to 6, characterized in that the implant base body is substantially annular.

8. Implant according to claim 7, characterized in that the implant base body is designed as an open ring with free ends.

9. Implant according to claim 8, characterized in that a distance between the free ends of the open ring is less than 10%, preferably less than 5%, in particular less than 3.5%, for example 0.5% to 3.0%, of a circumference of the open ring.

10. Implant according to one of claims 1 to 9, characterized in that the implant base body has an E-modulus of less than 80 GPa, preferably less than 60 GPa, in particular less than 55 GPa, for example 20 GPa to 50 GPa.

11. Implant according to one of claims 1 to 10, wherein the fastening body is formed from a polymer.

12. Implant according to one of claims 1 to 11, wherein the fastening body is formed from a polymer deposited by electrospinning.

13. Implant according to one of claims 1 to 12, characterized in that the fastening body substantially completely surrounds the implant base body.

14. Implant according to one of claims 1 to 13, characterized in that the fastening body is formed with pores.

15. Implant according to one of claims 1 to 14, wherein the antibacterial substance comprises or consists of nanoparticles.

16. Implant according to one of claims 1 to 15, wherein the antibacterial substance is distributed in and / or on the fastening body, in particular is distributed homogeneously.

17. Implant according to claim 16, wherein the antibacterial substance in mass percentage amounts to 0.10% to 6.0%, preferably 0.20% to 4.5%, in particular 0.30% to 4.0%, for example 0.50% to 1.5%, based on the total mass of antibacterial substance and fixing body.

18. Bioresorbable composite material for an implant to be implanted in the human or animal body, consisting of or comprising a bioresorbable carrier substance and an antibacterial substance, wherein the composite material is electrospun and / or in the form of a coating.

19. Bioresorbable composite material according to claim 18, wherein the carrier substance is meltable at temperatures of less than 100°C.

20. Bioresorbable composite material according to claim 18 or 19, wherein the antibacterial substance is present in a mass percentage of 0.10% to 6.0%, preferably 0.20% to 4.5%, in particular 0.30% to 4.0%, for example 0.50% to 1.5%, based on the total mass of antibacterial substance and carrier substance.

21. Bioresorbable composite material according to one of claims 18 to 20, wherein the antibacterial substance comprises or consists of nanoparticles.

22. A method for treating a patient with a disease such as heart valve disease or for the prophylaxis thereof, wherein an implant is inserted into the patient, preferably for treating a patient with a diseased or defective mitral valve, in particular with endocarditis or at high risk for endocarditis, characterized in that the method comprises the following steps: a) providing an implant according to one of claims 1 to 17; b) surgically implanting the implant in the patient.

23. Method according to claim 22, characterized in that the implantation is carried out via a catheter.

Images