DE102016122837A1 - Material for bone implants - Google Patents

Abstract

The invention is based on a material for a bone implant, comprising: a support structure having a surface comprising at least one biocompatible material, a matrix covalently bonded to said surface, and calcium phosphate incorporated into said matrix. A medically acceptable, highly compatible and versatile material may be provided when the matrix has at least one polysaccharide.

Description

The present invention relates to a bone implant material comprising a support structure having a surface comprising at least one biocompatible material, a matrix covalently bonded to that surface, and calcium phosphate incorporated into said matrix. The present invention further relates to a method for producing the material according to the invention, to a bone implant to which the material according to the invention is applied, and to its use as a bone implant material.

In recent years, the number of patients suffering from bone damage and / or requiring a bone implant has tended to increase. This trend highlights the need to explore high-quality, stable and functional bone replacement materials.

The organic components in the bone consist of 95% collagen and 5% proteoglycans and other adhesion-promoting glycoproteins. The mineral part of the bone consists almost exclusively of calcium phosphate in the modification of hydroxyapatite. The hard material properties of hydroxyapatite, in combination with the elastic properties of the organic components, make the bone a very versatile composite material.

The demands on high quality and functional bone implants are diverse and it is difficult to meet all requirements with one material. The functionality of an implant material is difficult to predict, since the natural process of bone wound healing and implant healing is very complex and sometimes not fully understood.

Wound healing of hard or soft tissues following surgery, such as bone grafting, involves many cellular and extracellular events. The healing process at the interface between the bone graft and the bone involves four steps that partially overlap each other. These include inflammatory reactions, the formation of a soft callus followed by the formation of a hard callus and finally the remodeling phase.

After surgery, proteins and other molecules of the blood and tissue fluids first adsorb to the surface of the implant. Adding a wound to vascularized tissue not only causes inflammatory responses but also activates a variety of other endogenous protective systems, such as the extrinsic and intrinsic coagulation system, the complement system, the fibrinolytic system, and the kinin system. This is followed by two sequential phases, which can also merge into one another, namely the acute and the chronic inflammatory reaction. The blood coagulates into a blood clot, which is composed of fibrin as the main component. In parallel, cytokines and other growth factors are released to recruit white blood cells to the wound. Neutrophils and mononuclear cells, such as monocytes, are recruited in the acute inflammatory response. Mononuclear cells differentiate into macrophages and attach themselves to the surface of the implant. In normal wound healing, macrophages are responsible for clearing the wound of bacteria, cell debris, and other contaminants via phagocytosis. The implant material is also perceived by the body as a foreign body. However, since the implant is much larger than the macrophages, they can not phagocytose the material. These events eventually lead to the phase of chronic inflammation at the material / tissue border. The macrophages fuse and form multinucleated giant cells in order to enclose the foreign body. The macrophages also provide for the recruitment of other cells, such as fibroblasts, which deposit fibrous tissue on the surface of the implant.

After the inflammatory phase, a soft callus forms. This consists of bone precursor cells and fibroblasts, which are located in a disordered matrix of non-collagenous proteins and collagen. This matrix is gradually formed by said cells as a first reaction and is structurally similar to the woven bone. The soft callus is finally gradually rebuilt by osteoblasts into ordered lamellar bone structure. Osteoblasts secrete type I collagen, calcium phosphate and calcium carbonate with an arbitrary, random orientation. The remodeling phase overlaps with the formation of the hard callus. This is done by resorption of the disordered bone structures by osteoclasts and subsequent formation of ordered bone structures by osteoblasts.

In order for materials to be suitable for use as bone graft materials and to allow optimal healing of the defect site, they must have some special properties. These include, for example, biocompatibility, immunogenicity, osseointegration and osteogenicity. To a good one To achieve biocompatibility, the material or its degradation products must not be toxic, carcinogenic or teratogenic. Neither in the implant environment nor in the rest of the body should any inflammatory, immune or other negative or unfavorable reactions be triggered. If it comes to a defense reaction of the body, the implant is encapsulated by the body by connective tissue from the rest of the body. The isolation in the connective tissue capsule prevents osseointegration of the implant into the adjacent bone tissue, as the connective tissue provides a barrier to the formation of vessels and thereby also to the necessary oxygen and nutrient transport. An implant must under no circumstances trigger an immune response of the body, so it must not show any immunogenicity. Particularly important for the suitability as an implant material is also a good osseointegration, through which a stable attachment of the foreign material in the bone is achieved, which must be able to cope with the everyday stresses of the patient later.

If individual particles detach from the implant, they must not trigger any of the aforementioned reactions and should either be degradable or secretible in the body to avoid permanent accumulation and attachment in the body or aseptic endoprosthesis loosening. Potential bone materials should therefore have reliable strength, high resistance, high abrasion resistance, corrosion resistance, and stiffness similar to bone. The latter plays a major role in the context of so-called "stress shielding". The bone is a dynamic system, which is dismantled or assembled according to mechanical stress. Now, if a bone implant material is used, which has a higher stiffness compared to the bone substance, this takes over a large part of the mechanical stress, whereby the surrounding bone is gradually degraded.

Another criterion is the so-called biocompatibility. A potential implant material should either be as bioinert or bioactive as possible. A bio-inert material does not cause any chemical or biological reactions in the body. The implant is therefore biocompatible and builds with the bone substance at best a positive connection (contact osteogenesis). Thus, only the transfer of pressure between implant and bone is possible, but this is sufficient for many applications, such as the replacement of skull bones, dental implants and fixing pins for bone fractures, completely sufficient. Implants made of titanium, aluminum, cobalt, chromium and polyether ether ketone (PEEK) belong to this type of material.

By contrast, a bioactive material promotes rapid implant ing into the surrounding tissue, ensuring fast and long-lasting fixation of the implant in the body. This effect is called "osseointegration". Bioactive materials therefore often show osteoconductive and osteoinductive properties and usually have a high hydrophilicity. Frequently, such materials are absorbed by the body. Hydroxylapatite, tricalcium phosphate and some bioglasses are among the bioactive materials. Bioactive materials can also cause an immune reaction in the body in the worst case. This ability of a biomaterial to promote cell adhesion and migration is critical to the early stages of wound healing and the later stages of new bone formation, and is highly dependent on the initial contact between the cells and the implant material.

Therefore, prior art bone graft materials have been coated with hydroxyapatite or tricalcium phosphate by various chemical or physical methods. These calcium phosphates are usually prepared by simple chemical methods by precipitation from aqueous solutions. The application to the surface of the implant material takes place either directly from the solution to the surface or by means of physical methods, such as "electrospray deposition". However, in the coating of materials with apatite, for example, both the low adhesion of the calcium phosphates on the implant is disadvantageous, as well as their limited cohesion within the individual calcium phosphate layers. With these procedures, a structure as similar as possible to the bone should be generated on the surface in order to favor the healing of the material into the bone. However, it is disregarded that the bone itself is a highly hierarchical composite material consisting of a matrix and a mineral phase.

Some methods of coating collagen implant materials have been evaluated for their in vivo functionality. Frequently, collagen-coated titanium implants, such as screws or nails, have been analyzed in various in vivo systems. For example, positive effects on the ingrowth of the material into the surrounding bone as well as the formation of new bone have been reported. However, contradictory results with respect to collagen-coated titanium implants are found in the prior art. Thus, for example, no improved osseointegration of collagen-coated porous titanium cylinders implanted in the stem-tibial diaphysis could be detected.

Also, the prior art has reported methods for covalently or noncovalently immobilizing various extracellular matrix proteins such as fibronectin or short peptides on surfaces. These showed some positive effects in in vitro test systems, such as cell adhesion and proliferation.

Often collagen is replaced by cost and handling reasons by its denatured form gelatin. The preparation of the gelatin is usually done by physical and chemical degradation or thermal denaturation of native collagen. Unlike native collagen, gelatin is water-soluble at a physiological pH and melts at a sol-gel transition temperature of 25 to 30 ° C. After cooling, transparent gels are formed. The prior art also reports the non-covalent application of gelatin to surfaces of arterial implant materials. Similarly, gelatin was covalently coupled to PEEK and mineralized with calcium phosphate to form a bone-like layer that resulted in very good proliferation of osteoblasts. The problem with gelatin-based coatings is that gelatine is an animal product that requires Class III certification. However, as gelatin is industrially derived from bovine or porcine skin, there may be religious reservations (pigs) or medical reservations (beef, BSE) that may limit the use of gelatin-based coatings.

On this basis, the present invention has the object to provide a material for bone implants, which is medically safe, highly tolerated and versatile and also has bone-like structures. Furthermore, a corresponding production method is to be provided.

The object is achieved by the features of the independent claims. Favorable embodiments and advantages of the invention will become apparent from the other claims and the description.

1. (a) eine Trägerstruktur mit einer Oberfläche, die zumindest ein biokompatibles Material umfasst,
2. (b) eine kovalent an diese Oberfläche gebundene Matrix und
3. (c) in diese Matrix eingelagertes Calciumphosphat.

In particular, an article of the present invention relates to a material for a bone implant, comprising:

1. (a) a support structure having a surface comprising at least one biocompatible material,
2. (b) a matrix covalently bonded to this surface and
3. (c) calcium phosphate incorporated into this matrix.

It is proposed that the matrix has at least one polysaccharide.

The terms "material for bone implants" and "bone implant material" are used synonymously herein. The material for bone implants according to the invention has bioactive properties. As used herein, the term "bioactive" refers to the property of the bone implant material of the present invention to permit rapid ingrowth into the surrounding tissue to provide rapid and prolonged fixation of the implant in the body. This property results from the technical features defined in the above sub-items (a) to (c) in combination with the characterizing part.

Whenever bone implants are discussed in the specialist literature, it is always attempted to make the surface of the implant as bone-like as possible. This can be done in the simplest case by plasma coating with hydroxyapatite. It is known, for example, to use a gelatin / collagen-based covalently bound coating which is mineralized with calcium phosphate. This is the conceivable bone-like coating for implants. This can be used to create the compound that plays a role in mineralization in collagen biomineralization (bone and dentin in teeth) and gelatin as degraded collagen.

In contrast, polysaccharides are not discussed in connection with the biomineralization of calcium phosphate - if at all only as glycoproteins. However, polysaccharides play a role in the biomineralization of calcium carbonate (egg shells - keratan sulfate, coccolith exoskeletons, etc.), but even there they are under-studied compared to proteins, as polysaccharides are notoriously difficult to characterize.

The approach according to the invention can therefore be regarded as a departure from known approaches. It is therefore by no means an obvious idea to think of building a bone-like bioinspired coating for an implant on polysaccharides. The obvious would be, as I said, collagen / gelatin or at least other proteins (here especially acidic proteins) to use. However, it has surprisingly been found that polysaccharides in this application represent a targeted and advantageous alternative to known substances, such as collagens.

In this context, a support structure is to be understood as meaning at least one material layer, which is constructed from the biocompatible material and which can be connected to the matrix or can be covalently bonded. The carrier structure may be, for example, an outermost layer of a basic structure of the implant manufactured entirely from the biocompatible material, for example a shaping element. Or it can be applied to a basic structure made of another material, such as the biocompatible material. Furthermore, in this context biocompatible, the property of a material or substance to be understood to be used in vivo and thereby have little to no negative impact on the patient or the healing process or to be used without consequential damage.

A matrix is to be understood here as meaning a structure which has as its main constituent a polysaccharide in which calcium phosphate is incorporated. In this case, the polysaccharide can be any polysaccharide which can be used by the person skilled in the art. If the polysaccharide is an animal polysaccharide, a substance can be used whose properties are well known and tested. Advantageously, the polysaccharide is a vegetable polysaccharide, whereby a vegan substance can be used, which is medically safe. In general, it is also possible to use "artificial" or non-naturally occurring polysaccharides. Or it is possible to use mixtures of polysaccharides. These can be produced specifically in the laboratory. Such polysaccharides can be selected and used very flexibly by their freely designed coating chemistry.

In a preferred embodiment, the polysaccharide is selected from a group consisting of alginic acid, alginate, hyaluronic acid, hyaluronate, pectin, carrageenan, agarose, amylose and chitosan. As a result, many different substances can be used, which can be selected individually, due to their special properties.

Hyaluronic acid (Hya) is a linear polysaccharide composed of disaccharide repeating units. These are composed of D-glucuronic acid and N-acetyl-glucosamine, which are linked by β-1,4 and β-1,3 glycosidic bonds (see below).

In a physiological environment, the carboxyl groups of hyaluronic acid are to a large extent present as the sodium salt, thus it is negatively charged and immobilizes a large number of water molecules; even at very low concentrations, an aqueous solution is therefore very viscous. Hya is used in many ways in the medical and healthcare industry. Since hyaluronic acid is widely biocompatible and safe, there are many fields of application. This makes them very interesting as a starting material for the surface coating of bone implants.

In the past, hyaluronic acid was extracted mainly from chicken combs that are produced as waste in food production. However, in the meantime, biotechnological production is done by cultivating genetically modified streptococcus bacteria, eliminating the risk of contamination with animal pathogens.

Alginic acid is a copolymer of the two uronic acids D-mannuronic acid (M) and L-guluronic acid (G). Alginic acid can be obtained from plants or algae, which is why their composition and the sequence distribution of the sugar units depend strongly on the place of origin of the plant / alga and their species. The G and M are each linked via β-1,4 glycosidic bonds. The good gelling properties of alginic acid are due to the G blocks that complex the calcium ions. Since alginic acid comes from natural sources, it must be ensured that potential contaminants such as heavy metals, endotoxins, proteins, etc. have been removed, otherwise immune reactions can occur. Provided that it has been sufficiently purified, alginic acid, like hyaluronic acid, does not cause an immune or inflammatory reaction in the body, so it has good biocompatibility.

Unlike hyaluronic acid, alginic acid is not degradable in the human body as there is no alginase. Because of its good biocompatibility, alginic acid is used in numerous biomedical applications, for example in wound dressings and, when RGD-peptide modified, also comes to be used as a bone graft material. If alginic acid is used in combination with hydroxyapatite, the formation of bone tissue can be additionally stimulated.

Furthermore, long-chain polysaccharides, but also branched polysaccharides, such as starch, can be used. Thus, the matrix has a more three-dimensional structure that is similar to the target bone structure, which can facilitate mineralization.

One way to add anti-bacterial properties to the coating / surface finish in addition to bone-identical mineralization (absorbency) is the use of chitosan, a linear polysaccharide-based biopolymer: chitosan may also utilize the material's hemostatic efficacy and antimicrobial properties. Chitosan also has the following material properties: non-allergenic, non-toxic, wound-healing, antibacterial, hemostatic, bacteriostatic, fungicidal, (biodegradable) and anti-microbial.

When the 6-position hydroxyl group present in most polysaccharides is attached to the implant via an ester linkage with linkers such as succinic anhydride or directly to polyacrylic acid-coated PEEK via ester linkage (optionally with activation chemistry via 1-ethyl-3- (3 -dimethylaminopropyl) carbodiimide (EDC) or the like), then the entire repertoire of polysaccharides is available for covalent attachment to the implant. This results in the possibility of using mixtures of polysaccharides or of polysaccharides with synthetic polymers. This will make a completely free composition of the coating available. This can range from negatively charged polysaccharides such as hyaluronic acid, via neutral ones such as amylose, to positively charged ones such as chitosan. Thus, the complete spectrum of charge densities can be covered from neutral to strongly negative or positive. Of course, this also affects the structure of the coating covalently attached to the implant and the mineralization of the calcium phosphate. The mixture of the bound polymers can be varied steplessly and thus also the property profiles.

Furthermore, in such mixtures, the structure of the coating can be adjusted specifically, since, for example, linear can be combined with branched polysaccharides. For example, the branched amylopectin of plant starch may be combined with the linear amylose. Chemically, both molecules are identical but not structural.

A promising pair for chemically different polysaccharides would be, for example, to couple the neutral amylose with the negatively charged alginate, which can additionally be cross-linked by the addition of Ca ²⁺ ions. Subsequent to this crosslinking, even further layers of alginate can be subsequently connected via simple impregnation of the layer with Ca ²⁺ , washing and application of the next alginate layer (layer by layer assembly, LBL). An extreme combination would be the negative alginic acid with the positive chitosan. These ionically stabilized "layer-by-layer assemblies" can be covalently bonded by suitable crosslinking chemistry (by means of EDC or the like). In between, all combinations of polysaccharides are continuously available via a common chemical bond via ester bonds.

In addition, it may be advantageous if the polysaccharide is a chemically modified polysaccharide. Here, chemically modified is intended to mean that on the polysaccharide artificial, chemical modification of a polysaccharide sugar, for example on a free group, e.g. Hydroxyl, aldehyde or acid group were made. This can be used to expand the range of applications. For example, inactive groups can be intentionally "transformed" into active groups, or an undesirable property can be eliminated. In this case, for example, a treatment with hexamethylenediamine (HMDA) or adipic dihydrazide (ADH) or a deacetylation. HMDA and ADH are both diamide linkers. Since ADH shows a lower basicity than HMDA, the coupling is possible even in the acidic pH range of 4.8. Both linkers thus address a coupling chemistry in different pH ranges. Depending on the requirement for the connection of the polysaccharide so a different linker may be necessary. Upon deacetylation, the polysaccharide can be rendered hydrolysis resistant.

If the biocompatible material consists of only one or a few layers, the material according to the invention for bone implants can be applied to solid materials or bodies (basic structure) which are used as a bone implant. These bodies may have any desired or required three-dimensional shape. Preferably, the entire surface of the material according to the invention for bone implants comprises or consists of the material defined in sub-item (a) above. Suitable materials to which the material of the present invention may be applied may be selected from ceramics known in the art, metals, polymers, composites or combinations thereof. These would be, for example, as metals: titanium / stainless steel, as ceramics: zirconium (dioxide), as polymer: polyether ketone (PEK) and the entire PEK family, but in particular: polyetheretherketone (PEEK), polyetherketone ketone (PEKK), polyetherketone ether ketone ketone (PEKEKK); Carbon Fiber Reinforced PEEK (CFR-PEEK), PEEK COMPOSITE, Glass Fiber Reinforced Polymer, Polyethylene (PE), Ultra High Molecular Weight Polyethylene (UHMWPE), Polyorthoester, Polymethylmethacrylate (PMMA), Polyethylene Terephthalate (PET) or Polyamides ( PA). In a preferred embodiment, the material is PEEK. This material is very similar to mechanically native bone material and thus well suited as a bone graft material.

ist ein sehr weit verbreiteter thermoplastischer High-Performance Kunststoff. Der semikristalline Kunststoff ist wegen seiner hohen Lösungsmittelbeständigkeit, dem hohen Schmelzpunkt von 343 °C und dem hohen Glasübergangspunkt von ca. 143 °C sehr beliebt für Anwendungen bei höheren Temperaturen.Polyetheretherketone (PEEK) with the following structural detail

is a very widespread thermoplastic high-performance plastic. Due to its high solvent resistance, the high melting point of 343 ° C and the high glass transition point of approx. 143 ° C, semicrystalline plastic is very popular for applications at higher temperatures.

Since 1998, PEEK has been available in implantable quality on the market. Since then, the market share of PEEK as implant material has greatly increased. PEEK is widely used in medical technology as a bone substitute material and implant, for example as a fusion cage for spinal fusion in intervertebral disc injuries. Since PEEK is permeable to X-ray radiation and does not interact with magnetic fields, the patient can be easily observed after surgery with imaging techniques to track the healing process in the affected area. The good mechanical properties of PEEK, which are very similar to those of the cortical bone (cortical bone) qualify PEEK as a well-suited material for use as bone substitute material and implants.

PEEK is one of the materials that almost completely bioinert behave, so do not enter into any specific interaction with the body. PEEK is neither rejected by the body nor integrated well into the bone, so ideally the bone is in good contact with the implant. Partial tissue encapsulation occurs, which reduces mechanical stability and may result in loss of the implant. To better integrate PEEK into the bone, some methods have been developed to achieve bioactivity of the material, such as calcium phosphate coating and also the addition of hydroxyapatite particles to the polymer. Other methods of surface modification of PEEK are possible and are for example as described in the following sections.

In order to improve the surface properties of a bone implant so that it is better accepted by the body, it is possible to modify it after the implant has been manufactured. Surface modifications may be more physical, such as coating with hydroxyapatite (HA) and titanium, or chemical nature.

To chemically modify a plastic substrate, such as PEEK, there are basically two options. One method is the direct reaction with small molecules or a plasma treatment to change the surface properties and to introduce, for example, linker molecules, on which one can carry out further reactions. Of course, the concentration of linker molecules per surface is very small here, since the surface of a macroscopic PEEK film is smooth and only the material is directly accessible on the surface. For polymer substrates used in solid phase synthesis, this is different. These have a good swelling behavior in suitable solvents, which can be introduced in the entire volume of the polymer body linker. This behavior is of course not desirable for plastic implants, such as PEEK, for bone implants, since the modification should only be limited to the surface in order not to impair the volume properties, such as hardness or mechanical stability.

However, to introduce many functional groups on the surface, the polymer substrate can be coated with a functional polymer. In this case, a multiple of functional groups is introduced, depending on the molecular weight of the introduced polymer, compared to the modification with small molecules.

Two methods are distinguished. One speaks of the grafting-from method when a polymer chain is initiated from the surface of the substrate and grows increasingly. In this variant, the substrate must be able to act, for example, as the initiator radical of a radical polymerization, or it must be able to be brought to it via an activator reagent. The grafting-from strategy achieves a higher density of functionalization, since the polymer chains grow successively and are not applied as a finished polymer. The reverse approach, grafting-to, is based on a finished polymer grafted to the surface with a suitable mechanism. A disadvantage of this method is that when grafting finished polymers, adjacent potential binding sites are strongly blocked by the macromolecules, this does not happen in the grafting-from approach. In contrast, control over the polymer length and molecular weight distribution of the applied polymers can only be achieved with the grafting-to approach.

The direct modification with small molecules or the reduction of the surface carbonyl groups of PEEK is possible by reduction with sodium borohydride in dimethylsulfoxide (DMSO) at 120 ° C and coupling of primary amines (see below and C. Henneuse; Goret; J. Marchand-Brynaert, Polymer 1998, 39, 835-844 and C. Henneuse-Boxus; E. Duliere; J. Marchand-Brynaert, European Polymer Journal 2001, 37, 9-18) ,

It is known to couple various molecules, including gelatin, to the reduced PEEK surface, thereby increasing hydrophilicity and achieving higher acceptance of bone-forming cells. The contact angle of the coated substrates also decreased significantly, whereby the hydrophilicity of the surface and thus the acceptance of bone-forming cells could be increased ( see J. Knaus, Master Thesis 2013 University of Konstanz and H. Cölfen; Tian Tian; J. Knaus, 2016 ).

As a grafting-from approach, for example, a surface-induced polymerization can be considered. By surface modification by means of small molecules can be introduced in a variety of organic polymers linker, which can be carried out further modifications to the surface. For example, it has been possible to introduce OH groups onto the surface of PEEK by wet-chemical modification, to which ATRP initiators have been coupled by means of acrylic acid derivatives (see US Pat Yameen; M. Alvarez; O. Azzaroni; U. Jonas; W. Knoll, Langmuir 2009, 25, 6214-6220 ).

Furthermore, a UV-indexed polymerization (UV: ultraviolet) into consideration. Instead of applying a radical initiator or other initiators to the surface, depending on the substrate, it is possible to generate radicals directly on the surface. This can be e.g. by auxiliary reagents, such as benzophenone (BP), benzoylbenzoic acid or other photoinitiators, which, upon UV excitation, abstract hydrogen atoms from suitable polymers and thereby generate radicals on the polymer surface which can initiate chain initiation. Polymers, such as PET, form hydroxyl and peroxide groups on the surface after argon plasma treatment in air. These can also serve as free-radical initiators upon excitation with UV light. With polyethylene, this was also done under similar conditions.

The photoinitiator BP is known to undergo a photopinacol reaction upon UV exposure. This results in the formation of a semibenzopinacol radical which can serve as an initiator in polymerizations. Photoinduced cleavage can also generate radicals from the excited molecule that initiate polymerizations. The polymer polyetheretherketone (PEEK) has units in the polymer backbone that behave similarly (see also M. Kyomoto, K. Ishihara, Acs Applied Materials & Interfaces 2009, 1, 537-542). In 2009, Kyomoto demonstrated that the surface of untreated PEEK under UV action is capable of initiating radical polymerizations of various acrylic acid derivatives. This involves a mixed grafting-from and grafting-to mechanism as both growing polymer chains are started on the surface and growing polymer chains terminate in solution at the surface. Building on its work, other groups have self-initiated polymerizations under UV excitation, including acrylic acid. The Kyomoto 2013 direct detection of ketyl radicals was achieved by in situ ESR spectroscopy.

The covalently bonded matrix typically has a layer thickness of 100 to 150 nanometers (nm), but may be thicker or thinner. In particular, the covalently bonded matrix may have a layer thickness of 1 nm to 10 microns (μm), preferably from 10 nm to 1 μm, more preferably from 20 nm to 500 nm, more preferably from 30 nm to 300 nm, more preferably from 50 nm 200 nm, and most preferably from 100 to 150 nm. Further, the covalently bonded matrix preferably covers the entire surface of the bone implant material of the present invention.

An interesting variant of the grafting-to method is the trapping of growing radical polymer chains by immobilized radical scavengers on the surface of the substrate (see also P. Yang, JY Xie, J. Yuan, L. Zhang, WN Liu; WT Yang, Journal of Polymer Science Part a-Polymer Chemistry 2007, 45, 745-755 ). By Yang et al. It has been shown that it is possible to use the polymerization inhibitor hydroquinone to apply polymers on a substrate after the grafting-to approach.

In this case, a hydroquinone derivative was coupled to the surface and carried out a normal free-radical polymerization with a thermal radical initiator. The immobilized hydroquinone quenched the growing polymer chain by homolytic cleavage of the OH bond, it comes to chain termination. The long-lived aryloxyl radical is unable to initiate radical polymerization with the monomers present, but can recombine with a growing polymer radical, capturing the polymer at the surface.

Finally, the bone implant material of the present invention includes calcium phosphate incorporated into said matrix, preferably calcium orthophosphate in all mineral forms, more preferably selected from the group consisting of amorphous calcium orthophosphate (ACP), dicalcium phosphate dihydrate (DCPD), octacalcium phosphate, and hydroxyapatite with partial fluoride, chloride or carbonate substitution, with ACP, hydroxyapatite and octacalcium phosphate being particularly preferred. Methods for incorporating said calcium phosphates into a corresponding matrix are described below.

In a preferred embodiment, the polysaccharide is attached to the biocompatible material via a linker, wherein the linker is selected from a group consisting of: a diamine linker or diamine linker in combination with a succinic acid linker or a (UV-grafted) polyacrylic acid linker. Corresponding linkers are known in the art. Methods for the covalent attachment of polysaccharides to, for example, PEEK are described below.

In other embodiments, the bone implant material of the present invention comprises or consists of oxide ceramic materials, titanium, polymeric materials, or composite materials, wherein the polysaccharide of the covalently bonded matrix is bound to titanium or ceramic oxide materials via a silane linker. Suitable silane linkers and corresponding methods for binding polysaccharides are known in the art.

1. (a) das biokompatibles Material PEEK ist,
2. (b) das Polysaccharid Alginsäure ist und
3. (c) das in diese Matrix eingelagerte Calciumphosphat Hydroxylapatit, insbesondere kristallines Hydroxylapatit ist.

In a particularly preferred embodiment, the present invention relates to a material for bone implants, comprising:

1. (a) the biocompatible material is PEEK,
2. (b) the polysaccharide is alginic acid and
3. (c) the calcium phosphate embedded in this matrix is hydroxyapatite, in particular crystalline hydroxyapatite.

1. (a) Bereitstellen einer Trägerstruktur mit einer Oberfläche, umfassend ein biokompatibles Material,
2. (b) kovalentes Ankoppeln einer Matrix, umfassend zumindest ein Polysaccharid, an diese Oberfläche, und
3. (c) Mineralisieren der Matrix mit Calciumphosphat.

A further subject matter of the present invention relates to a method for producing a material according to the invention for bone implants, comprising the steps:

1. (a) providing a support structure having a surface comprising a biocompatible material,
2. (b) covalently attaching a matrix comprising at least one polysaccharide to said surface, and
3. (c) mineralizing the matrix with calcium phosphate.

For this object of the present invention, all relevant definitions, advantages and preferred embodiments, which have been listed above for the material according to the invention for a bone implant, apply analogously.

Methods for the covalent coupling of a polysaccharide-comprising matrix to a surface according to step (b) of the method according to the invention are not particularly limited and are known in the prior art.

In a preferred embodiment, in particular if the surface comprises or consists of PEEK, step (b) of the method according to the invention comprises the steps: (b1) covalently coupling a linker molecule selected from the group consisting of a diamine linker or a Diamine linker and a succinic acid-linker or UV-grafted polyacrylic acid (PAA) to this activated surface, and (b2) covalently coupling the polysaccharide having carboxylic acid groups to the diamine linker molecule or the hexamethylenediamine-modified polysaccharide to the succinic acid linker or the unmodified polysaccharide via ester linkages to the polyacryric acid linker.

Methods for coupling linker molecules to a correspondingly activated PEEK surface are also not particularly limited.

Methods for mineralizing a polysaccharide-containing matrix with calcium phosphates according to step (c) of the method according to the invention are not particularly limited. In the case that amorphous calcium phosphate (ACP) is used, for example, they include incubating the surface with a solution comprising calcium chloride, dipotassium hydrogen phosphate and a nucleation inhibitor. This nucleation inhibitor is preferably a non-collagenous protein or protein analog, more preferably poly-aspartic acid and / or fetuin. For example, in the case where hydroxyapatite is used, you include incubating the surface with a solution comprising calcium chloride and dipotassium hydrogen phosphate.

Another object of the present invention relates to a bone implant on which the bone implant material according to the invention is applied.

For this object of the present invention, all relevant definitions, advantages and preferred embodiments, which have been listed above for the material for bone implants according to the invention, apply in an analogous manner.

Another object of the present invention relates to the use of the material according to the invention for a bone implant as a bone implant material. For this object of the present invention, all relevant definitions, advantages and preferred embodiments, which have been listed above for the material for bone implants according to the invention, apply in an analogous manner.

Another object of the present invention relates to the use of the material according to the invention for a bone implant, for example for the treatment of bone damage.

Also for this object of the present invention, all relevant definitions, advantages and preferred embodiments, which were listed above for the material according to the invention for bone implants, apply analogously.

Another object of the present invention relates to the use of the bone implant according to the invention, for example for the treatment of bone damage.

Also for this object of the present invention, all relevant definitions, advantages and preferred embodiments, which were listed above for the material according to the invention for bone implants, apply analogously.

Implants grow better in the human body, and become more stable with the body (among other things by increased accumulation of body cells), the better the implant surface corresponds to the natural bone. This is the goal of the present invention. Furthermore, the coating should be covalently bonded to the surface of the implants. The bone implant materials according to the invention have a higher biocompatibility, better healing in the natural bones and an increased mechanical strength.

The surface modification according to the present invention aims to covalently bond bone-like structures to the surface of bone graft materials that include an organic polysaccharide matrix and the mineral phase of the natural bone. This is to assist the healing of the implant in the bone. These structures include a matrix of a polysaccharide which is eventually mineralized with calcium phosphate. Mineralization occurs with the help of non-collagenous proteins and their analogues, which act as nucleation inhibitors, so that mineralization is controlled and ectopic mineralization is avoided. Such nucleation inhibitors are, for example, poly-aspartic acid or fetuin. Mineralization with octacalcium phosphate or hydroxyapatite is by incubating the polysaccharide matrix in a solution containing calcium ions or phosphate ions. By a slow and controlled addition of a solution of each complementary phosphate ions or calcium ions, octacalcium phosphate and / or hydroxyapatite may be precipitated within the polysaccharide matrix. Due to the relatively disordered structure of the polysaccharide, the resulting surface modification has a braid-like or callus-like structure. The bone cells could thus build up further disordered collagen structures around the material during the healing of the material or link the material directly to the bone. These disordered structures can then eventually be converted into ordered bone structures in the natural remodeling phase of bone wound healing. In this case, however, the cells can not penetrate to the direct surface of the implant material due to the covalent attachment of the polysaccharides during the remodeling and thus always remain in a desired matrix of extracellular proteins. The implant material is thus masked for the cells in order to avoid unwanted reactions during the healing of implants. Since the modifications only affect the surface of the implant materials, material properties are not changed.

The basic chemical reactions can easily be adapted for the modification of different materials. Thus, metal oxide surfaces can be covalently attached via established silane chemistry. This makes the surface coating according to the invention also interesting for oxide ceramic materials. Furthermore, since metals are readily oxidizable on the surface, for example by plasma treatment, the common titanium implant materials of the surface modification according to the invention via silanes are accessible via silane chemistry.

In recent years, many different materials have been developed for use as bone implants. The biological, chemical, as well as mechanical requirements for implant materials must be combined in one material in order to come as close as possible to the properties of the bone. The variety of approved materials reflects the great effort in this field. The plastic polyetheretherketone (PEEK), for example, has very good mechanical properties comparable to natural bone. The disadvantages are obvious with the high hydrophobicity and thus low bioactivity, which is why there are numerous attempts to tackle this problem.

A method was developed in which the gelatin functionalization of the bone implant plastic PEEK was established. In the present invention, a method has now been developed to covalently bond to the surface of PEEK a network of polysaccharides of plant or bacterial origin, in the concrete hyaluronic acid, alginic acid derivatives and fully synthetic polymers, to address the problems of animal-based coating to circumvent, for example, the evidence of sterility and the lengthy approval process due to the potential endotoxin content and the allergen potential. Some patients opt for certain animal products for personal reasons, be it religious or ethical reasons. Vegan or synthetic functionalizations could be interesting alternatives for this group of people. In order to come as close as possible to the chemical structure of natural bone material, finally, the applied coating with calcium phosphate, in particular hydroxyapatite, mineralized.

The description of advantageous embodiments of the invention given so far contains numerous features that are summarized in several dependent claims in several groups. However, these features may conveniently be considered individually and summarized to meaningful further combinations, particularly in terms of claims, so that a single feature of a dependent claim may be combined with a single, several, or all features of another dependent claim. In addition, these features can be combined individually and in any suitable combination both with the method according to the invention and with the device according to the invention according to the independent claims. Thus, process features are also objectively formulated as a property of the corresponding device unit and functional device features also as corresponding process features.

The above-described characteristics, features and advantages of this invention, as well as the manner in which they are achieved, will become clearer and more clearly understood in connection with the following description of the embodiments; the examples mentioned in the following description do not limit the invention the combination of features specified therein, not even in terms of functional characteristics. In addition, suitable features of each embodiment may also be explicitly considered isolated, removed from one embodiment, incorporated into another embodiment to complement it, and / or combined with any of the claims.

The methods mentioned and used in the text below (ATR-IR analysis, scanning electron microscopy, confocal laser scanning microscopy, fluorescence spectrometry, NMR measurements, thermogravimetric analysis (TGA), UV investigations, contact angle measurement, nuclear magnetic resonance spectroscopy were used in accordance with well-known principles and procedures and performed with known devices.

It is possible via a wet-chemical modification of the PEEK surface to achieve the reduction of the carbonyl groups in the PEEK skeleton by treating the film with NaBH ₄ in DMSO at 120 ° C.

The reaction can be carried out without exclusion of oxygen with stirring in a 500 milliliter (mL) flask. 10 PEEK plates (1 cm ² (square centimeter) each) were placed in 20 milliliters (mL) of DMSO and stirred. It was heated to 120 ° C and after 20 min 13mmol (490 mg) of NaBH _{4 were} added. The reaction time was 4 h 30 min. The PEEK was washed in 20 mL MeOH for 15 min, in 20 mL H ₂ O for 10 min and in 20 mL 1 M HCl for 35 min. After rinsing in EtOH, the PEEK was dried in a vacuum drying oven for 2 h at 40 ° C and 50 mbar. The reaction was verified by ATR-IR spectrum (ATR-IR: ν 3400 cm ^-1 (m) (cm ^-1 : wavenumber), strong attenuation of 1647 cm ^-1 (w) C = O valence vibration, not shown) ,

10 PEEK-Plättchen (je 1 cm²) wurden in 30 ml Aceton vorgelegt und auf 40°C erhitzt. Es wurde 1 Gramm (g) (10 mmol) Bernsteinsäureanhydrid zugegeben. Die Reaktionszeit betrug 5 h 35 min. Die Plättchen wurden mit je 20 mL Aceton, H₂O und Ethanol gewaschen und im Vakuumtrockenofen für 2 h bei 40°C und 50 mbar getrocknet. Die Reaktion wurde anhand eines ATR-IR-Spektrums verifiziert. ATR-IR: ν 3400 cm^-1 (m), 2924 cm^-1 (w), 2861 cm^-1 (m) sp³ CH₂ Valenzschwingungen, 1705 (w) COOH Valenzschwingung (nicht gezeigt).The reduced PEEK films can be further converted to the succinic acid ester. The esterification can be carried out by means of succinic anhydride in acetone at room temperature:

Ten PEEK flakes (1 cm ² each) were placed in 30 ml of acetone and heated to 40 ° C. 1 gram (g) (10 mmol) of succinic anhydride was added. The reaction time was 5 h 35 min. The platelets were washed with 20 ml of acetone, H ₂ O and ethanol and dried in a vacuum drying oven for 2 h at 40 ° C and 50 mbar. The reaction was verified by ATR-IR spectrum. ATR-IR: ν 3400 cm ^-1 (m), 2924 cm ^-1 (w), 2861 cm ^-1 (m) sp ³ CH ₂ valence oscillations, 1705 (w) COOH valence vibration (not shown).

Purified PEEK films could be reacted in pure diamine (ethylenediamine (EDA) and 1,3-diaminopropane). There is formation of imines (Schiff's base) on the PEEK surface. The reaction may proceed for 3 hours with reflux of the diamine and stirring of the mixture:

5 PEEK platelets (1 cm ² each) were initially charged in 10 ml of ethylenediamine or 1,3-diaminopropane. The reaction mixture was heated with stirring for 3 h under reflux. The reaction mixture was cooled to RT and the PEEK plates washed extensively with acetone. The modified films were dried in a vacuum oven for 2 h at 40 ° C and 50 mbar. The reaction was verified by an ATR-IR spectrum ATR-IR: ν 2925 cm ^-1 (m), 2854 cm ^-1 (m) sp ³ CH ₂ valence oscillations, 1620 cm ^-1 (w) C = N valence vibration (not shown).

It was also possible to introduce a succinic acid linker to the amine-functionalized PEEK samples. The PEEK films were placed in 10 mL dry 10 millimolar (mM) succinic anhydride DMF solution (DMF: dimethylformamide). After 10 hours, the films were washed thoroughly with MilliQ (ultrapure water). And examined by ATR-IR spectroscopy (not shown). ATR-IR: ν 2925 cm ^-1 (m), 2854 cm ^-1 (m) sp ³ CH ₂ valence oscillations, 1706 cm ^-1 (w) C = O valence vibration.

The modified PEEK films were subjected to contact angle measurement. The modified PEEK films were treated with phosphate buffer prior to measurement, rinsed with MilliQ and dried well. The contact angle measurements can be made 5 seconds (s) after application of the drop. Contact angle measurements with ultrapure water (MilliQ) imply a significantly increased hydrophilicity of the ethylenediamine-modified PEEK compared to unmodified PEEK films (84.5 °). In the variant modified with 1,3-diaminopropane, the contact angle at 70 ° has also become smaller since the surface also became more hydrophilic. The increase in hydrophilicity suggests a successful reaction during the reaction of PEEK with the diamines.

Untreated PEEK films can be coated with a polyacrylic acid grafting-from polymerization method (UV light-induced PEEK modification):

The experimental procedure used can be carried out in one step. It can be worked with degassed aqueous solutions of distilled acrylic acid. As UV light source an OSRAM Vitalux 300 can be used without further filters.

4 PEEK plates (1 cm ² each) were placed in a Schlenk flask and degassed by means of 3 vacuum / nitrogen cycles. The appropriate amount of degassed MilliQ water (4 Freeze Pump Thaw cycles) was added. After adding the degassed acrylic acid (passing nitrogen through it for 30 minutes), the flask was irradiated while stirring with the UV lamp from a distance of 15 cm. The reaction time was between 15 minutes and 75 minutes. The concentration of the acrylic acid solution was between 5 wt% and 25 wt%. ATR-IR: ν 1705 cm ^-1 (m) COOH valence vibration.

The respective possible conditions and results are listed in Table 1 Table 1. Table 1: Reactions of UV grafting with acrylic acid. entry Acrylic acid concentration [wt%] Reaction time [min] Gel layer on surface MG10 20 75 Yes MG11 10 70 Yes MG13 7.5 70 Yes MG19 5 30 Yes MG20 5 45 Yes MG12 5 60 Yes

The polyacrylic acid-grafted PEEK films were examined by scanning electron microscope (SEM) (not shown).

In each approach, a substantially homogeneous layer of polyacrylic acid was formed. The higher the polyacrylic acid concentration and the longer the reaction time the more polyacrylic acid was deposited on the surface essentially in spherical form or the higher was the layer thickness. Here, the smaller the polyacrylic acid concentration, the smaller the beads (30 (micron (μm) at MG10 versus 3 μm at MG12).

For further surface functionalization, the samples were polymerized for 30 minutes at 5 wt% acrylic acid moiety. Under these conditions, the coating was still sufficiently thick to speak of a homogeneous coating (not shown), and at the same time, at this layer thickness, one can assume that the mechanical properties of the bulk material are not adversely affected. The polyacrylic acid layers made at higher acrylic acid concentrations are not suitable for targeting as a bone graft material because of the thick gel cushion of up to 5 millimeters (mm) because a gel pad on the surface greatly degrades mechanical contact to the surrounding tissue.

At low magnification, it can be seen that the polyacrylic acid has formed linear structures with the beads. The formation of these lines could be due to the drying method (vacuum furnace) but may also be due to the hydrophobicity of the PEEK surface. The acrylic acid molecules diffusing to the active site have a higher affinity for a growing polyacrylic acid layer than for the hydrophobic PEEK surface. This could explain the line structures consisting of PAA balls. The average bead size under these reaction conditions is 1.7 microns, making it about half as large as at 60 min polymerization. Coating results were verified by ATR-IR spectra (not shown).

The coatings which have been prepared at 5% by weight of acrylic acid and 30 minutes of UV treatment have proven particularly suitable. The PEEK can be thinly coated under these conditions so that too little gel pad is deposited.

The UV-induced grafting polymerization is thus very well suited for the coating of PEEK with polyacrylic acid, since significant amounts of PAA could be detected on the PEEK surface.

It is also possible to carry out polymerizations in pure acrylic acid and, for comparison, in pure methyl acrylate, ie a polymerization in the pure monomer. The results were verified by ATR-IR spectra (not shown). Coupling of 1,4-diaminobutane to the PAA-coated PEEK surface:

The polyacrylic acid layer can be modified by the coupling of diamine linkers, so that later can be formed with organic acids amide bonds. The coupling of the diamine species to the carboxyl groups can be carried out using the modern coupling reagent 4- (4,6-dimethoxy-1,3,5-triazin-2-yl) -4-methylmorpholinium chloride (DMT-MM). Activation and coupling can be performed with DMT-MM at a buffered pH of 9:

nach dem Fachmann bekannter Weise durchgeführt werden: (Stern = 7-Hydroxicumarin-Fluorophor, Sub = Substrat und siehe S. Shiota; S. Yamamoto; A. Shimomura; A. Ojida; T. Nishino; T. Maruyama, Langmuir 2015, 31, 8824-8829) :

Quantification of the amino groups on the surface of the left-functionalized PEEK films (substrate) can be carried out by means of the fluorescent dye C-coumarin:

be carried out in a manner known to the person skilled in the art: (star = 7-Hydroxicumarin fluorophore, Sub = substrate and see S. Shiota; S. Yamamoto; A. Shimomura; A. Ojida; T. Nishino; T. Maruyama, Langmuir 2015, 31, 8824-8829) :

The fluorescence intensity can be used to determine the concentration of the dye in solution and thus to draw conclusions about the amount of surface amino groups. For the functionalized PEEK samples (with ethylenediamine, 1,3-diaminopropane 4 and tetramethylenediamine (TMDA)), a higher NH ₂ density than an unfunctionalized PEEK sample could be determined (data not shown).

Syntheses of Modified Polysaccharides:

To modify the PEEK films with polysaccharides, two strategies can be pursued: the introduction of the linker molecule (diamine) on the carboxylated PEEK surface and the introduction of the linker on the polymer. In the first case, in the polymer coupling step, unmodified polysaccharide would be coupled to free amino groups on the substrate, whereas in the second case, amine functionalized polysaccharides would be anchored to carboxyl groups on the substrate.

oder Alginsäure (gezeigt als Strukturausschnitte von Alginsäure mit den verschiedenen Poly-G, Poly-M und alternierenden Blöcken. Je nach Herkunft der Alginsäure ist das Verhältnis aus G und M unterschiedlich):

mit einem Amin funktionalisiert werden, um danach an Carboxylgruppen auf der PEEK-Oberfläche gekuppelt zu werden. Hierbei kann es nötig sein, vor der Funktionalisierung eine Deacetylierung durchzuführen.In one approach, unmodified hyaluronic acid:

or alginic acid (shown as structural sections of alginic acid with the various poly-G, poly-M and alternating blocks, depending on the source of alginic acid, the ratio of G and M is different):

be functionalized with an amine to be subsequently coupled to carboxyl groups on the PEEK surface. It may be necessary to perform a deacetylation prior to functionalization.

Deacetylation of hyaluronic acid:

Unmodified hyaluronic acid consists of a D-glucuronic acid and an N-acetyl-glucosamine unit. It therefore has one free carboxyl function and one acetylated amine function per disaccharide monomer. Two common methods for introducing amine functionalities, in addition to substitution reactions on the OH groups, the functionalization of the carboxyl groups with diamine linkers, or the deprotection of the N-acetyl group to the free amine.

The deacetylation can be carried out in aqueous hydrazine solution under Hydrazinsulfatkatalyse:

To 1 g of sodium hyaluronate was added 50 mL of hydrazine monohydrate and 0.5 g of hydrazine sulfate to give a solution, based on the polymer, of 2 percent by weight (wt%). After stirring for 72 h at 55 ° C, the polymer product was precipitated in cold ethanol, filtered and dried in vacuo (24 h). The residue was taken up in 20 mL of 5% acetic acid. To this solution was added aqueous iodic acid solution (10 mL, 0.5 M) keeping the temperature at 4 ° C for 1 h in an ice bath. Aqueous solution of hydrogen iodide (57%, 3 mL) was added. After 15 minutes, the violet solution in the separating funnel was extracted five times with 25 mL diethyl ether each time until the aqueous phase was colorless. The pH of the solution was adjusted to 7-7.5 with 0.2 M NaOH solution. The polymer was precipitated in 1 volume equivalent of ethanol, dissolved in H ₂ O and dialyzed against deionized water. The dialysis water was changed twice a day. After three days of dialysis, the solution was freeze-dried and the deacetylated hyaluronic acid was obtained as a product.

The free amino groups could be used as anchor groups for coupling to the carboxyl groups of the PEEK-PAA surface. The polymer was examined by NMR spectroscopy to determine the degree of deacetylation (not shown).

Modification of hyaluronic acid by means of hexamethylenediamine and adipic dihydrazide:

The free carboxyl groups of hyaluronic acid may be suitable for various modification possibilities of the polymer. In the literature, e.g. reported by amidation, ester formation or Ugi condensation.

For the synthesis of the amidated hyaluronic acid hexamethylenediamine (HMDA) can be used. The coupling of the amine can proceed via the classic EDC / NHS coupling (1-ethyl-3- (3-dimethylaminopropyl) carbodiimide / N-hydroxysuccinimide coupling) in an aqueous medium. The reaction is split into an activation phase in slightly acidic and a coupling phase in slightly basic. The pH value of the reaction had to be continuously monitored and adjusted:

wurde hergestellt (500 Milligramm (mg) in 167 mL MilliQ). Bezogen auf die Menge der Carboxylgruppen im Polymer wurden 30 äq. Hexamethylendiamin (HMDA, 30äq, 39,6 mmol, 4,6 g) zugegeben. Der pH Wert der Lösung wurde auf 7,5 eingestellt (0,1 M NaOH, bzw. 0,1 M HCl). EDC (4 äq., 5,28 mmol, 0,9 g) und NHS (4 äq., 5,28 mmol, 0,607 g) wurden in 10 mL Wasser gelöst und dann zu der Reaktionslösung gegeben. Der pH Wert der Mischung wurde mittels Zugabe von 0,1 M NaOH bei 7,5 gehalten. Die Reaktion wurde über Nacht gerührt. Der pH Wert wurde auf 7 eingestellt und das Polymer in Ethanol (3 Volumenäquivalente) ausgefällt. Das Polymer wurde in MilliQ gelöst $\left(\text{5 }\frac{mg}{mL}\right)$

und 6 Tage gegen VE-Wasser (Vollentsalztes Wasser) dialysiert. Das VE-Wasser wurde täglich 2 Mal gewechselt. Das gereinigte Produkt wurde 4 Tage gefriergetrocknet. Der Grad der HMDA-Funktionalisierung wurde mittels ¹H-NMR Spektroskopie bestimmt. 46 % HMDA Funktionalisierung.An aqueous sodium hyaluronate solution with $3\frac{mG}{mL}$

was prepared (500 milligrams (mg) in 167 mL MilliQ). Based on the amount of carboxyl groups in the polymer were 30 eq. Hexamethylenediamine (HMDA, 30 eq, 39.6 mmol, 4.6 g) was added. The pH of the solution was adjusted to 7.5 (0.1 M NaOH, or 0.1 M HCl). EDC (4 eq., 5.28 mmol, 0.9 g) and NHS (4 eq., 5.28 mmol, 0.607 g) were dissolved in 10 mL of water and then added to the reaction solution. The pH of the mixture was maintained at 7.5 by the addition of 0.1 M NaOH. The reaction was stirred overnight. The pH was adjusted to 7 and the polymer precipitated in ethanol (3 volume equivalents). The polymer was dissolved in MilliQ $\left(\text{5}\frac{mG}{mL}\right)$

and dialyzed against demineralised water for 6 days. The demineralised water was changed twice a day. The purified product was freeze-dried for 4 days. The degree of HMDA functionalization was determined by ¹ H NMR spectroscopy. 46% HMDA functionalization.

The reaction was carried out with a very large excess of diamine (30 times, based on the amount of carboxyl groups in the polymer) to prevent the cross-linking of hyaluronic acid. By ¹ H NMR spectra, the successful coupling of the HMDA linker and ATR-IR spectra revealed the successful functionalization of the carboxyl groups through formation of amide bonds with the HMDA linkers (not shown).

Modification of hyaluronic acid with adipic acid diazide:

For the synthesis of the adipic dihydrazide (ADH) derivative, the method of synthesis was modified. Since ADH shows a lower basicity than HMDA, the coupling is possible even in the acidic pH range of 4.8. Therefore, the addition of NHS could be omitted:

wurde hergestellt durch lösen von 500 mg Natriumhyaluronat in 170 mL H₂O. Bezogen auf die Menge der Carboxylgruppen im Polymer wurde ein 40-fach molarer Überschuss an Adipinsäuredihydrazid (ADH, 52,8 mmol, 9,2 g) zugegeben. Es wurde gewartet bis das ADH vollständig gelöst war (15 min). Der pH Wert der Reaktionsmischung wurde mittels 1 M HCl Lösung auf 4 eingestellt. Es wurde Ethanol (50 mL, 50 Volumenprozent (Vol%)) zugegeben und 30 Minuten gerührt. Es wurden 4 äq. EDC-HCl (5,3 mmol, 0,9 g) zugegeben.A sodium hyaluronate solution with $3\frac{mG}{mL}$

was prepared by dissolving 500 mg of sodium hyaluronate in 170 mL of H ₂ O. Based on the amount of carboxyl groups in the polymer, a 40-fold molar excess of adipic dihydrazide (ADH, 52.8 mmol, 9.2 g) was added. It was waited until the ADH was completely dissolved (15 min). The pH of the reaction mixture was adjusted to 4 using 1 M HCl solution. Ethanol (50 mL, 50% by volume (vol%)) was added and stirred for 30 minutes. There were 4 eq. EDC-HCl (5.3 mmol, 0.9 g) was added.

und für 9 Tage dialysiert. 1 Tag wurde gegen 100 mM NaCl Lösung dialysiert und danach abwechselnd einen Tag gegen 25 Vol% Ethanollösung und einen Tag gegen VE Wasser. Der Ethanol/VE Wasser Zyklus wurde 4 Mal wiederholt. Die Polymerlösung wurde schließlich 3 Tage gefriergetrocknet. Der Grad der ADH Funktionalisierung wurde mittels ¹H-NMR Spektroskopie bestimmt. 53 % ADH Funktionalisierung.The pH was maintained at 4.8 for 2 h using 1 M HCl. After 2 h, the reaction was stopped neutralization of the solution using 1 M NaOH (pH = 7). The reaction solution was placed in prewashed dialysis membrane tubing $\left(\text{MWCO}=3500\frac{G}{mOl}\right)$

and dialysed for 9 days. 1 day was dialyzed against 100 mM NaCl solution and then alternately one day against 25 vol% ethanol solution and one day against deionized water. The ethanol / VE water cycle was repeated 4 times. The polymer solution was finally freeze-dried for 3 days. The degree of ADH functionalization was determined by ¹ H NMR spectroscopy. 53% ADH functionalization.

The synthesis can be carried out with a large excess of ADH to prevent the cross-linking of hyaluronic acid here as well. In both syntheses, the product had to be extensively dialyzed to remove the large excess of HMDA or ADH. By ¹ H NMR spectra, the successful synthesis of ADH-modified hyaluronic acid and ATR-IR spectra demonstrated the successful functionalization of the carboxyl groups by forming amide bonds with the ADH (not shown).

HMDA modified alginic acid:

Similar to hyaluronic acid, alginic acid can also be functionalized with HMDA. The coupling can be carried out according to a specification for the octylamine functionalization of alginic acid:

im Polymer. Nach 5 min Reaktionszeit wurden 10 äq. Hexamethylendiamin (7, 05 g) zugegeben. Die Lösung wurde 24 h bei Raumtemperatur gerührt. Das Polymer wurde in Aceton ausgefällt, nach Trocknung in H₂O gelöst und gegen H₂O dialysiert. Das Wasser wurde 2 Mal täglich gewechselt. Nach 9-tägiger Dialyse wurde die Polymerlösung 4 Tage gefriergetrocknet. Es wurde 938 mg Produkt erhalten. Der Grad der HMDA Funktionalisierung wurde mittels ¹H-NMR Spektroskopie bestimmt. 31,5 % HMDA Funktionalisierung.30 mL aqueous sodium alginate with 3 wt% (1 g sodium alginate) was placed in a flask and the pH was adjusted to 3.4 by means of 0.1 M HCl. The polymer solution was thereby diluted to 50 mL (2 wt%). In 4 mL H ₂ O, 797.5 mg (4.16 mmol) of EDC-HCl was dissolved and added to the polymer solution. The ratio of EDC amount to carboxyl functionalities was thus 0.7. The concentration of EDC was determined by the molar frequency of sodium alginate monomers $\left(\text{M}=168.11\frac{G}{mOl}\right)$

in the polymer. After a reaction time of 5 minutes, 10 eq. Hexamethylenediamine (7.05 g) was added. The solution was stirred at room temperature for 24 h. The polymer was precipitated in acetone, dissolved after drying in H ₂ O and against H ₂ O dialysed. The water was changed twice a day. After 9 days of dialysis, the polymer solution was freeze-dried for 4 days. There was obtained 938 mg of product. The degree of HMDA functionalization was determined by ¹ H NMR spectroscopy. 31.5% HMDA functionalization.

The synthesis can be done with a large excess of HMDA to prevent crosslinking of alginic acid. The product can be extensively dialyzed to rid it of excess HMDA. By ¹ H NMR spectra, the successful coupling of the HMDA linker and ATR-IR spectra revealed the successful functionalization of the carboxyl groups through formation of amide bonds with the HMDA linkers (not shown).

Functionalization of the PEEK surface with polysaccharides: Different strategies can be pursued for the further functionalization of the PEEK surface. On the one hand, differently modified polysaccharides can be provided with amine functionalities and, in the next step, coupled to the carboxyl groups on the PEEK surface. The other approach is based on the modification of the carboxyl groups on the surface by diamines in order to be able to couple unmodified polysaccharides in the next step.

Coupling of amine-functionalized polysaccharides:

In order to be able to apply the polysaccharides to the polymer substrate, a peptide bond should be built up. On the one hand, the classic EDC / NHS coupling chemistry can be used. On the other hand, it is also possible to work with the modern coupling reagent DMT-MM. The classic EDC / NHS coupling can be carried out in two stages. After activation of the PEEK film in the slightly acidic MES buffer for 20 minutes, the coupling with the modified polysaccharides in the slightly basic phosphate buffer can be carried out overnight:

As an alternative coupling reaction, the one-step coupling with the modern coupling reagent 4- (4,6-dimethoxy-1,3,5-triazin-2-yl) -4-methylmorpholinium chloride (DMT-MM) can be chosen. The activation mechanism and the subsequent formation of the peptide bond can be carried out as follows:

It has the advantage that the reaction can be carried out at a constant pH of 9 and the reaction vessel between activation and coupling must not be changed. The coupling at pH 9 is significantly faster compared to the coupling to the NHS ester at pH 7.3, since the amines at pH 9 are largely free amine, which is important for nucleophilic attack on the activated carbonyl center. The NHS coupling can not be carried out at such high pH values as otherwise the NHS ester is hydrolyzed too quickly in the aqueous solution:

Modification of the imino-functionalized PEEK surface:

The imine-functionalized PEEK surface can be reacted under identical reaction conditions with native hyaluronic acid and alginic acid.

Direct modification of the imitated PEEK surface with native hyaluronic acid. The coupling was carried out by means of DMT-MM in an aqueous phosphate buffer solution at pH = 8:

Direct modification of the imitated PEEK surface with native alginic acid. The coupling was carried out by means of DMT-MM in an aqueous phosphate buffer solution at pH = 8 (aminated PEEK film was shaken with polysaccharide in PBS buffer (PBS: phosphate buffer saline) at pH = 8 (67 mM) in 1 mM DMT-MM solution overnight. Concentration of hyaluronic acid: 0.1 mg / mL, alginic acid: 0.05 mg / mL, washed three times with MilliQ of water):

Modification of the polyacrylic acid coated PEEK surface:

The coating of the PEEK substrates with polyacrylic acid was very successful as described above. The very large amount of carboxyl groups that could thereby be introduced on the PEEK surface should now serve as a target for further functionalization with various polysaccharide derivatives. Thus, coupling reactions were carried out with ADH-hyaluronic acid, HMDA-hyaluronic acid, deacetylated hyaluronic acid and HMDA-alginic acid.

Modification of the polyacrylic acid coated PEEK surface with ADH-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

The reaction was carried out under the same conditions as the previous couplings.

Modification of the polyacrylic acid coated PEEK surface with HMDA-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with HMDA-alginic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with deacetylated hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Overall, significantly more coupling experiments were carried out on the PEEK substrate. The overview of the reactions is shown in Table 2, Table 2. The couplings marked with an X were carried out. All reactions were carried out under the same basic conditions using DMT-MM as coupling reagent. Table 2: Overview of the polysaccharide couplings to PEEK substrates. Hya Alg Deac.-Hya ADH Hya HMDA Hya HMDA Alg PEEK Imin X X PEEK-imine COOH X X X X PEEK PAA X X X X PEEK-PAA-amine X X

The reaction schemes associated with each reaction are as follows: ATR-IR spectra demonstrated successful functionalization (not shown).

Direct modification of the imitated PEEK surface with native alginic acid. The coupling was carried out by means of DMT-MM in an aqueous phosphate buffer solution at pH = 8:

Modification of the iminated and carboxylated PEEK surface with ADH-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the iminated and carboxylated PEEK surface with HMDA-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the iminated and carboxylated (succinic acid) PEEK surface with HMDA-alginic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the iminated and carboxylated PEEK surface with deacetylated hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with ADH-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with HMDA-hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with HMDA-alginic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid coated PEEK surface with deacetylated hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid-coated and then tetramethylenediamine-treated PEEK surface with native hyaluronic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the polyacrylic acid-coated and then treated with tetramethylenediamine PEEK surface with native alginic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

Modification of the imitated PEEK surface with coupled succinic acid with native alginic acid. The coupling was carried out by means of DMT-MM in aqueous phosphate buffer solution at pH = 8:

• Zudem können drei unterschiedliche Ansätze zur Mineralisierung von Hydroxylapatit in per PEEK-PAA Probe durchgeführt werden.
• Zum einen können zwei Versuche mit Calciumprästrukturierung unternommen werden und einer mit Phosphat-Prästrukturierung.

Mineralization experiments with hydroxyapatite:

• In addition, three different approaches to the mineralization of hydroxyapatite in PEEK-PAA sample can be performed.
• On the one hand, two attempts can be made with calcium pre-structuring and one with phosphate pre-structuring.

Here the following ca-prestructuring would be possible:

zugegeben. Nach vollständiger Zugabe wurde die Mischung über Nacht gerührt.An established hydroxyapatite synthesis route can be modified and used for the mineralization of PEEK-PA films. The PEEK-PA film can be placed in 0.3 M calcium chloride solution at a buffered pH of 9 and stirred for 30 minutes. Now, a disodium hydrogen phosphate solution also buffered at pH = 9 was added at a rate of $3\frac{mL}{min}$

added. After complete addition, the mixture was stirred overnight.

In addition, a Ca-prestructuring and phosphate pre-structuring could be carried out:

Simpler variants of mineralization can also be performed. PEEK-PAA samples were incubated for 72 h in diammonium hydrogen phosphate (phosphate pre-structuring, 6 mL vials with 1 M aqueous (NH ₄ ) ₂ HPO ₄ solution) or calcium nitrate Cal (6 mL vials with 1 M aqueous Ca (NO ₃ ) ₂ solution) given. After this rest period, the samples were placed in the other solution (0.6 M (NH ₄ ) ₂ HPO ₄ or 0, respectively). 6M Ca (NO ₃ ) ₂ solution) and left for a week to allow the counterions to diffuse into the gel as well.

Summary

Work has been done on the surface functionalization of the bone graft plastic polyetheretherketone to allow for better incorporation into the treated bone area. Polyetheretherketone surfaces have been successfully chemically modified by various methods. The surface properties were changed with the help of small molecules and hydroxyl, carboxyl and imine functionalities were obtained on the surface. The modified surface was analyzed and characterized by ATR-IR spectroscopy (not shown). Furthermore, functional polymers, such as polyacrylic acid, but also polymethyl acrylate (PMA) were deposited on the polyetheretherketone surface by means of UV-induced grafting polymerization. The polyacrylic acid layer was examined by various surface analysis methods, such as ATR infrared spectroscopy, scanning electron microscopy, and confocal laser scanning microscopy, to obtain spectroscopic information about the surface and to obtain an accurate picture of its topography (not shown). The polyacrylic acid layer was modified by the coupling of diamine linkers, so that later can be formed with organic acids amide bonds. In order to be able to make quantitative statements about the degree of surface functionalization with amino groups, the cleavable fluorescent dye C-coumarin was synthesized with which the accessible amino groups on the surface could be indirectly quantified. The quantification was successful for the samples which were directly imin-functionalized with diamines.

Hyaluronic acid with adipic dihydrazide and hexamethylenediamine and alginic acid was modified only with the diamine to couple amine linker for subsequent anchoring to the various polyetheretherketone substrates. Likewise, hyaluronic acid was deacetylated to introduce amine functionality on the polysaccharide. The modified polysaccharides were characterized by NMR and ATR infrared spectroscopic methods (not shown).

The numerous modified and unmodified polysaccharides were coupled to the complementary PEEK substrates. The coupled samples were analyzed by ATR infrared spectroscopy, scanning electron microscopy, and partly by thermogravimetry (not shown).

outlook

Since the surface-induced radical polymerization has been very successful, there are a number of approaches in this area where further work could be done. Since the introduction of polysaccharide structures on the polyetheretherketone surface did not prove to be trivial, but the polymerization with acrylic acid worked very well, it would be a promising approach to coat the PEEK surface directly with polymers of modified acrylic acid derivatives. Suitable acrylic acid-based monomer units would be, for example, acrylic acid derivatives modified with sugar molecules, which are known to play an important role in the cell adhesion of osteoblasts. It would also be interesting if the monomer units would carry oligosaccharides of hyaluronic acid, or else short adhesion-promoting RGD peptide sequences, etc.

A very sensitive surface analysis method is X-ray photoelectron spectroscopy. In this way, polysaccharides could possibly be detected on the surfaces of the substrates produced.

The investigation of the swelling behavior of the polyacrylic acid layers on the polyether ether ketone substrate could be an approach to optimize the coupling conditions so that polysaccharide detection would be possible even with simple analytical methods. Couplings in non-aqueous media would also be conceivable, but are certainly also problematic because of the poor solubility of the polysaccharides.

Further attempts to deposit hydroxyapatite in the polyacrylic acid layer should also be made, as the hydroxyapatite coating has been shown to have beneficial effects on organism acceptance.

QUOTES INCLUDE IN THE DESCRIPTION

This list of the documents listed by the applicant has been generated automatically and is included solely for the better information of the reader. The list is not part of the German patent or utility model application. The DPMA assumes no liability for any errors or omissions.

Cited non-patent literature

• C. Henneuse; Goret; J. Marchand-Brynaert, Polymer 1998, 39, 835-844 and C. Henneuse-Boxus; E. Duliere; J. Marchand-Brynaert, European Polymer Journal 2001, 37, 9-18). [0046]
• see J. Knaus, Master Thesis 2013 University of Konstanz and H. Cölfen; L.F. Tian; J. Knaus, 2016 [0047]
• Yameen; M. Alvarez; O. Azzaroni; U. Jonas; W. Knoll, Langmuir 2009, 25, 6214-6220 [0048]
• W.T. Yang, Journal of Polymer Science Part a-Polymer Chemistry 2007, 45, 745-755. [0052]
• 7-Hydroxicumarin fluorophore, Sub = substrate and see S. Shiota; S. Yamamoto; A. Shimomura; A. Ojida; T. Nishino; T. Maruyama, Langmuir 2015, 31, 8824-8829) [0098]

Claims (14)

A bone implant material comprising: (a) a support structure having a surface comprising at least one biocompatible material, (b) a matrix covalently bonded to said surface, and (c) calcium phosphate incorporated into said matrix, characterized in that said matrix is at least having a polysaccharide. Material after Claim 1 , characterized in that the polysaccharide is a vegetable polysaccharide or animal polysaccharide. Material after Claim 1 or Claim 2 , characterized in that the polysaccharide is selected from a group consisting of alginic acid, alginate, hyaluronic acid, pectin, carrageenan, agarose, amylose and chitosan. Material according to one of the preceding claims, characterized in that the polysaccharide is a chemically modified polysaccharide. Material according to one of the preceding claims, characterized in that the biocompatible material is selected from a group consisting of an oxide ceramic material, a polymer material, a composite material and titanium. Material according to one of the preceding claims, characterized in that the biocompatible material is polyetheretherketone (PEEK). The material of any one of the preceding claims, characterized in that the polysaccharide is linked to the biocompatible material via a linker, wherein the linker is selected from a group consisting of a diamine linker, a diamine and succinic acid linker, and a polyacrylic acid linker. Material according to any one of the preceding claims, characterized in that (a) the biocompatible material is PEEK, (b) the polysaccharide is alginic acid and (c) the calcium phosphate embedded in this matrix is hydroxyapatite, in particular crystalline hydroxyapatite. Material according to one of the preceding claims, characterized in that the covalently bonded matrix has a layer thickness of 100 to 150 nm. Material according to one of the preceding claims, characterized in that the matrix covers the entire surface of the support structure. Method for producing a material for a bone implant, in particular according to one of the Claims 1 to 10 comprising the steps of: (a) providing a support structure having a surface comprising a biocompatible material, (b) covalently attaching a matrix comprising at least one polysaccharide to said surface, and (c) mineralizing the matrix with calcium phosphate. Method according to Claim 11 characterized in that step (b) comprises the steps of: (b1) covalently coupling a linker molecule selected from the group consisting of a diamine linker or a diamine linker and a succinic acid linker or UV-grafted polyacrylic acid to which has been activated Surface, and (b2) covalently coupling the polysaccharide having carboxylic acid groups to the diamine linker molecule or the hexamethylenediamine modified polysaccharide to the succinic acid linker or the unmodified polysaccharide via ester linkages to the polyacryric acid linker. Bone implant on the bone implant material after at least one of Claims 1 to 10 is applied. Use of the material according to at least one of Claims 1 to 10 as a bone implant material.