WO2003072156A1 - Biomedical implant surfaces with osteogenic properties - Google Patents

Abstract

The invention relates to a device with a surface covered by a thin film or monolayer of a poly (ethylene glycol)-grafted copolymer exhibiting osteogenic properties, defined in an osteoblastic cell culture test as a surface that increases the expression of growth factor TFG-β1, the expression of hormone PGE2 and/or the expression of mineralization marker osteocalcin by at least a factor of 2 in comparison to the corresponding control surface consisting of either unmodified substrate surface or of tissue-culture polystyrene (TCPS).

Description

Biomedical Implant Surfaces with Osteogenic Properties

Background of the Invention

The use of functional polymers and copolymers is an approach often chosen in the biomaterial area. For example, U.S. Patent Nos. 5,573,934 and 5,626,863 to Hubbell et al. disclose hydrogel materials containing a water-soluble region such as polyethylene glycol and a biodegradable region, including various biodegradable polymers such as polylactide and polyglycolide, terminated with photopolymerizable groups such as acrylates. These materials can be applied to a tissue surface and polymerized, for example, to form tissue coatings. These materials are adhered to tissue surfaces by polymerizing the photopolymerizable groups on the materials after they have been applied to the tissue surface.

U.S. Patent No. 5,462,990 and 5,627,233 to Hubbell et al. discloses multifunctional polymeric materials for use in inhibiting adhesion and immune recognition between cells and tissues. The materials include a tissue-binding component (polycation) and a tissue non-binding component (polyanion). In particular, Hubbell discloses various PEG7PLL copolymers, with molecular weights greater than 300, with structures that include AB copolymers, ABA copolymers, and brush-type copolymers. These polymers are being commercially developed for use as tissue sealants and to prevent surgical adhesions. International Patent WO 98/47948 (29 Oct. 1998) and U. S. Patent Application "Multifunctional Polymeric Tissue Coatings" by Hubbell et al. describes the use of grafted polyionic copolymers that are able to attach to biological and non-biological samples in order to control cell-surface and cell-cell and tissue-surface interactions in biomedical applications. PCT/US98/07590 and USSN 09/403,428 by Textor et al discloses the application of polyionic, PEG-grafted copoylmers in the general area of polymeric coating materials which can be applied to surfaces of substrates used in analytical and sensing devices ('chip')to promote specific recognition of the target analyte and at the same time minimize non-specific adsorption of other molecules in the sampling solution.

BESTATIGUNGSKOPIE More specifically, US 5,567,440 "Methods for Modifying Cell Contact with a Surface",

EP0975691A1 resp. WO 9847948 "Multifunctional Polymeric Tissue Coatings", and EPO 1009451 A2 "Methods and Compositions to Prevent Formation of Adhesions in Biological Tissue" show that cell-cell and cell-surface interactions can be reduced or eliminated by using polymer coatings containing an appropriate amount of grafted "non- interactive" material such as poly(ethylene glycol) (called "PEG") or poly(ethylene oxide) (called " PEO") within the polymer. The preparation of properties of such polyionic, PEG-grafted copolymers for the production of non-interactive surfaces have been published [N. Huang, R. Michel, J. Nόrδs, M. Textor, R. Hofer, A. Rossi, D.L. Elbert, J.A. Hubbell, Ν.D. Spencer, Poly(L-lysine)-g-poly(ethylene glycol) layers on metal oxide surfaces: Surface analytical characterization and resistance to serum and fibrinogen adsorption", Langmuir 17 (2): 489-498 (2001). G.L. Kenausis, J. Vόrδs, D.L. Elbert, Ν.P. Huang, R. Hofer, L.Ruiz, M. Textor, J.A. Hubbell, Ν.D. Spencer, "Poly(L- lysine)-g-poly(ethylene glycol) Layers on Metal Oxide Surfaces: Attachment Mechanism and Effects of Polymer Architecture on Resistance to Protein Adsorption", J. Phys. Chem. B 104: 3298-3309 (2000)].

Artificial (biomaterial) surfaces with such biologically non-interactive properties have their particular applications in the area of biomedical devices where the aim is that the device does not interact (or interacts as little as possible) with the biological environment. The aim can be minimization or prevention of postoperative adhesion, minimization or prevention of thrombosis or infection (EP0975691A1 resp. WO 9847948). Examples of specific applications are stents, vascular or cardiovascular grafts, and catheters in contact with blood or other body fluids. In such cases one usually prefers that proteins or cells do not adhere to the surfaces, blood is not activated and no or little platelet adhesion and thrombus formation occurs, which otherwise would lead to possibly adverse effects in the body of the patient, such as the formation of embolies.

The situation is different in the application of biomedical devices where a strong interaction of the foreign biomaterial and the body of an animal or human patient is usually aimed at. This is particularly true for cases where the device is expected to strongly interact with the surface in order to have the part integrated rapidly and firmly from both a biological and mechanical standpoint. Examples for devices where such requirements are relevant are biosensors placed in vivo, load-carrying implants such as screws in bone applications or cages in spinal surgery. There are three types of approaches commonly discussed or used to induce such strong interactions of cells (and tissue) with biomaterial and biomedical device surfaces: a) The use of special, dedicated topographies that induce particular cellular response [Barbara D. Boyan, David D. Dean, Christoph H. Lohmann, David L. Cochran, Victor L. Sylvia, Zvi Schwartz, The titanium-bone cell interface in vitro: The role of the surface in promoting osteointegration, in: «Titanium in Medicine: Material Science, Surface Science, Engineering, Biological Responses and Medical

Applications», Springer Verlag, Heidelberg and Berlin, 2001, pp. 561-586] and/or that favor the mechanical interaction ("interlocking") between neotissue and implant/device [Daniel Buser, Titanium for dental applications: Implants with roughened surfaces, in: «Titanium in Medicine: Material Science, Surface Science, Engineering, Biological Responses and Medical Applications», Springer

Nerlag, Heidelberg and Berlin, 2001, pp. 875- 888]. b) The use of a chemical composition that is interactive with proteins in blood or other body fluid in a non-specific way ("non-specific protein adsorption", meaning the uncontrolled adsorption of usually many different types of proteins). Many of these proteins have peptide sequences that interact with (generally) different types of receptors that are contained in the cell membrane, finally forming focal contact points that are essential in cell adhesion. Surfaces that show such behavior or common: many metallic materials such as titanium, or polymers have surface composition that supports such non-specific adsorption of cell- adhesive proteins. c) A third alternative is the use of surfaces that are resistant to protein- and cell- adhesion as mentioned above, but onto which specifically interactive biological ligands ("bioligands") are grafted. Typical examples of such bioligands are peptides with amino acid sequences that specifically interact with particular types of α-β receptors in the cell membrane or alternatively with heparin-type of moieties in the cell membrane. Again, depending on the density of such bioligands at the surface, focal contacts are formed between cell membrane and device surface, leading to strong attachment and spreading of the cell at the surface. In application where the aim is to induce a strong interaction leading to firm integration in a particular tissue environment, one or a combination of the approaches a), b), c) are typically chosen. This general approach relies on the provision of a high surface concentration of interactive sites (e.g. adhesion proteins or peptides), leading to a high density of strongly attached and spread cells at the interface between the biological environment and the biomaterial or biomedical device.

In explicit contrast to this state-of-the-art approach, we have developed-and claim correspondingly in this patent application-a method to coat biomaterial or biomedical device surfaces for integration in bony tissue, which does NOT rely on the provision of a high density of bioligands at the surface for extensive cell attachment and spreading. The invention is related to the observation that surfaces on titanium or titanium oxide material (with either smooth or rough surface topography) that have been specially treated with a PEG-grafted polymer in order to make them little or not at all interactive with proteins, elicit a biological response of osteoblastic cells in contact with such surfaces, which we describe as osteogenic, strongly stimulating osteoblast differentiation, expression of mineralization (osteocalcin) and growth factor (TGF-βl) factor, and formation of an osteoblast phenotype characterized by a less flattened ("spread") morphology in comparison to conventional, more strongly cell-interactive titanium surfaces.

Summary of the Invention The invention covers the production of biomaterial or biomedical device surfaces with osteogenic properties, mediated in part by the application of a coating on the surfaces based on poly(ethylene glycol) (PEG) or poly(ethylene oxide) (PEO). Such surfaces are shown to favor osteoblastic cell differentiation as measured by alkaline phosphatase activity, stimulation of osteoblast phenotype and expression of factors related to mineralization and general tissue formation (osteocalcin, TGF-βl). This stimulation of an osteogenic response, based on an entirely new approach, not relying on the use of complex growth factors or bone morphogenic proteins, is much stronger than what has so far been reported to our knowledge. The effect can be observed on both smooth/flat as well as on topographically structured/rough surfaces. It is claimed that the combination of dedicated surface topographies as used today in certain dental implants together with the PEG-derived surface chemistry is particularly attractive to the envisaged applications.

A preferred method of producing these osteogenic surfaces is based on the spontaneous assembly of polyionic, PEG-grafted copolymers onto oppositely charged surfaces such as those formed by inorganic oxides or natural oxide films on metallic biomaterials or artificially produced oxide films on metallic biomaterials or charged surfaces on any other material such as on polymers. This way of surface treatment using spontaneous adsorption from solution and attachment through electrostatic and van der Waals interaction is particularly suitable for coating not only smooth but also rough surfaces thus combining in an additive or synergistic way the preferred rough topographies used in bone-related application with the osteogenic PEG-derived surface chemistry.

In detail:

A preferred example is a device with a surface covered by a thin film or monolayer of a poly(ethylene glycol)-grafted copolymer exhibiting osteogenic properties, defined (in an osteoblastic cell culture test) as a surface that increases the rate of differentiation as measured by alkaline phosphatase, the expression of growth factor TGF-βl, the expression of hormone PGE₂ and/or the expression of mineralization marker osteocalcin by at least a factor of 5 in comparison to the corresponding control surface consisting of unmodified substrate surface or of tissue-culture polystyrene (TCPS). Another preferred example is a device, wherein the density of the polyethylene glycol (PEG) chains and their molecular weight is adjusted to optimize the osteogenic activity.

Another preferred example is a device, wherein the copolymer is a PEG-grafted polycationic poly(amino acid) or a PEG-grafted polycationic synthetic polymer or a PEG- grafted polycationic polysaccharide that adsorbs spontaneously from solution onto surfaces that are negatively charged at the pH of adsorption and/or pH of use (i.e. 7.4 in biological media).

Another preferred example is a device, wherein the polymer is a PEG-grafted poly(amino acid) with a polycationic backbone made of lysine, histidine, arginine or ornithine in D-, L- or DL configuration, or the polymer is a PEG-grafted polymer with a cationic backbone of a polysaccharide such as chitosan, partially deacetylated chitin, and amine- containing derivatives of neutral polysaccharides, or the polymer is a PEG-grafted non- peptide polyamine with a polycationic backbone such as poly(aminostyrene), poly(aminoacrylate), poly (N-methyl aminoacrylate), poly (N-ethylaminoacrylate), poly(N,N-dimethyl aminoacrylate), poly(N,N-diethylaminoacrylate), poly(aminomethacrylate), poly (N-methyl amino-methacrylate), poly(N-ethyl aminomethacrylate), poly(N,N-dimethyl aminomethacrylate), poly(N,N-diethyl aminomethacrylate), poly(ethyleneimine), polymers of quaternary amines, such as poly(N,N,N-trimethylaminoacrylate chloride), poly(methacrylamidopropyltrimethyl ammonium chloride), or the polymer is a PEG-grafted charged synthetic polymer with a polycationic backbone such as polyethyleneimine, polyamino(meth)acrylate, polyaminostyrene, polyaminoethylene, poly(aminoethyl)ethylene, polyaminoethylstyrene, and N-alkyl derivatives thereof.

Another preferred example is a device 3 where the substrate material is a metal or metal- oxide-covered surface or a ceramic surface or a polymeric surface that adopts a negative charge at the pH of adsorption and/or the pH of use (i.e. 7.4 in biological media).

Another preferred example is a device, wherein the copolymer is a PEG-grafted polyanionic synthetic polymer or a poly(amino acid) that adsorbs spontaneously from solution onto surfaces that are positively charged at the pH of adsorption and/or pH of use (i.e. 7.4 in biological media).

Another preferred example is a device, wherein the copolymer is a PEG-grafted copolymer with an anionic backbone of a poly(amino acid) grafted with poly(ethylene glycol) where the amino acid contains an additional pendant carboxy group imparting a negative charge to the backbone at pH above 4 and in particular at neutral pH such as polyaspartic acid or polyglutamic acid; or a natural or unnatural polymer with pendant negatively charged groups, particularly carboxylate groups, including alginate, carrageenan, furcellaran, pectin, xanthan, hyaluronic acid, heparin, heparan sulfate, chondroitin sulfate, dermatan sulfate, dextran sulfate, poly(meth)acrylic acid, oxidized cellulose, carboxymethyl cellulose and crosmarmelose, synthetic polymers and copolymers containing pendant carboxyl groups, such as those containing maleic acid or fumaric acid in the backbone.

Another preferred example is a device where the substrate material is a metal or metal- oxide-covered surface or a ceramic surface or a polymeric surface that adopts a positive charge at the pH of adsorption and/or the pH of use (i.e. 7.4 in biological media). Another preferred example is a device, wherein the PEG molecular weight is between 500 and 20,000, preferentially between 2000 and 5000, and the grafting ratio lysine monomersPEDG chains is between 2 and 20, preferentially between 2.5 and 7.

Another preferred example is a device, wherein the material to be coated is a biocompatible metal covered by a natural or artificially strengthened oxide film, such as titanium, niobium, tantalum, zirconium, hafnium, molybdenum, tungsten, cobalt- chromium, cobalt-chromium-molybdenum, stainless steel, as well as further alloys composed of these materials, in particular titanium-aluminum-vanadium and titanium- aluminum-niobium alloys.

Another preferred example is a device, wherein the material to be coated is a biocompatible polymer with a surface that is charged or made charged at the pH of adsorption or pH of use and is used in combination with an oppositely charged PEG- grafted copolymer.

Another preferred example is a device, wherein the polymer is made charged in contact with an aqueous solution through prior application of a surface modification technique, such as a wet-chemical, gas-phase-chemical, plasma or flame exposure treatment with the effect that charged functional groups are introduced in the surface of the polymer.

Another preferred example is a device, wherein a fraction of the PEG chains are linked at the end of the chain to a specific peptide in order to increase the interactivity, as measured by cell attachment and spreading, of the surface with particular types of cells. Another preferred example is a device, wherein the peptide density is so low that the osteogenic activity is still present.

Another preferred example is a device, wherein less than 1% of all PEG chains at the surface are modified with a peptide. Another preferred example is a device, wherein the peptide contains a sequence of an integrin-receptor-type binding type of peptide such as RGD or of a heparin-domain- binding type of peptide such as KRSR or FHRRIKA.

Another preferred example is a device, where the surface of the same device contains a pattern with areas of low interactiveness and high degree of osteogenic activity, where cells strongly express factors that are beneficial to the healing and new bone forming process, and areas of higher interactiveness where stronger cell attachment and spreading takes place inducing stronger proliferation of bone-forming cells.

Another preferred example is a device, where the cell-interactive areas of the pattern are composed of unmodified substrate surface or are made with peptide modified polymers, and a higher peptide density that induces stronger interaction and proliferation.

Another preferred example is a device, wherein the device is an implant for application within bone structure or in contact with bone.

Another preferred example is a device, wherein the application of the implant is in the areas of dental implantology, maxillofacial surgery, osteosynthesis, spinal surgery, or orthopedics.

Another preferred example is a device, wherein the device to be coated is a scaffold for growing boneous tissue ex vivo or in vivo by tissue engineering.

Another preferred example is a device, wherein the material to be coated is a bone substitute such as resorbable or non-resorbable calcium phosphates.

Brief Description of the Figures

As an example of a PEG-grafted polymers that can show osteogenic activities, Figure 1 shows the chemical structure of the graft copolymer poly-L-lysine-grafted-poly(ethylene glycol with a polycationic backbone of poly-L-lysine and a fraction of the lysine sie chains grafted with PEG.

Caption of Figure 1: Molecular structure of a typical representative of a polycationic PEG-grafted copolymer (poly-L-lysine-grafted-poly(ethylene glycol, PLL-g-PEG). Schematic representation of a smooth and rough negatively-charged surface coated with such a copolymer.

Figure 2 shows the expression of PGE₂ in cell culture tests using an osteoblastic cell line (MG 63) for surfaces smooth and rough (SLA) surfaces covered with PLL-g-PEG ("PEG"), PLL-g-PEG-RGD ("PEG-RGD") and PLL-g-PEG-RDG ("PEG-RDG") respectively to demonstrate the exceptionally large increase in factor expression for the PEG and the PEG-RDG surfaces.

Caption of Figure 2: PGE₂ production on the smooth and rough Ti surfaces coated with PEG, PEG-RDG, and PEG-RGD. MG63 cells were cultured to confluence on the smooth surfaces (Panel A) or the rough SLA surfaces (Panel B) and PGE₂ release into the media measured at harvest . *P<0.05, vs. plastic; #P<0.05, vs. glass; ∑P<0.05, vs. Ti; +P<0.05, vs. Ti-PEG or Ti-PEG-RDG. Detailed Description of the Invention

The present invention is related to the generation of surfaces with osteogenic properties and its applications to materials and devices for use in the area of bone-related implantology and tissue engineering.

The invention originates from the observation of osteoblast behavior at surface coated with particular types of PEG-grafted copolymers (Figure 1) that can be described as osteogenic. This claim is based on the fact that osteoblasts on such surfaces express factors of a kind and at a release rate (more than twice and up to 100-fold concentration or more if compared to standard surfaces; see Figure 2 as an example) that can be described as being typical for materials or surfaces with osteogenic properties, such as biocompatible substrates loaded with growth factors (e.g. collagens/TGF or hydroxyapatite/BMP combinations). Details are given in the section "Example".

The materials of choice to impart osteogenic properties to implants and scaffolds are copolymers characterized by two components: one is poly(ethylene glycol) (PEG) that is uncharged, hydrophilic and exposed to the biological environment, and a second polymer that is charged to interact with oppositely charged surfaces of an artificial material or device, thus binding strongly to the surface (Figure 1) [N. Huang, R. Michel, J. Vδros, M. Textor, R. Hofer, A. Rossi, D.L. Elbert, J.A. Hubbell, N.D. Spencer, Poly(L-Lysine)-g- poly(ethylene glycol) Layers on Metal Oxide Surfaces: Surface Analytical Characterization and Resistance to Serum and Fibrinogen Adsorption, Langmuir 17 (2): 489-498 (2001)]. For surfaces that are negatively charged at neutral pH, a positively charged polymer backbone would be used such as poly-L-lysine, and vice versa. An example is given in Figure 1 and in the Example section. Other suitable polymers are described in the earlier patents and patent applications compiled in the section "Background of the Invention".

The invention relies on the observation as demonstrated in the Example, that the osteogenic properties are related to surfaces that are less interactive than standard surfaces such as on a titanium implant. This reduction of interactiveness is achieved through a reduction of the extent of protein adsorption caused by the grafting of "protein- repellent" PEG. The reduced protein adsorption causes a reduction in cell adhesion, spreading and proliferation of osteoblasts while strongly upregulating processes (by at least a factor of five compared to control surfaces) that are essential to bone formation: differentiation and increased phenotype expression of osteblastic cells, and strongly increased expression of growth factors and factors related to mineralization.

Therefore, the invention makes it feasible to steer and tailor the balance between proliferation, differentiation and maturation of bone-related cells via the composition and architecture of the PEG-grafted copolymer. On the one hand, the degree of protein adsorption (or resistance) depends on the density of the PEG chains at the surface [G.L. Kenausis, J. Vδrδs, D.L. Elbert, N.P. Huang, R. Hofer, L.Ruiz, M. Textor, J.A. Hubbell, N.D. Spencer, Poly(L-lysine)-g-poly(ethylene glycol) layers on metal oxide surfaces: Attachment mechanism and effects of polymer architecture on resistance to protein adsorption, J. Phys. Chem. B 104: 3298-3309 (2000)]. The lower the PEG density at the surface is, the higher is the amount of proteins that adsorb from serum-containing cell culture medium (in vitro) or from blood or other body fluid (in vivo), making the surface becoming increasingly similar to standard surfaces used in bone application such as titanium or titanium alloys. On the other hand, even at the high PEG surface density, the surface can be made increasingly more interactive with cells by introducing at the end of the PEG-chains cell-adhesive ligands such as peptides. In the Example, we demonstrate that by introducing RGD-containing peptide ligands at 5% of the PEG chains, the surface has again properties that are no longer osteogenic and close to the non-functionalized standard titanium surfaces (i.e. control surfaces). The invention therefore covers procedures of producing surfaces that possess osteogenic properties, whereas the degree of osteogeneity and the balance between cell differentiation and phenotype expression on the one hand and cell attachment, spreading and proliferation on the other hand can be tailored through the polymer composition and architecture. In relation to the osteogenic properties, important design aspects of the PEG-grafted copolymer are the molecular weight and grafting density of PEG and possibly, where applicable, the concentration of peptide functions. Regarding the latter aspect, we only claim very low densities of cell- adhesive peptides at the surface (meaning a very small fraction of the PEG chains are modified by peptides), since larger densities lead to properties, which can no longer be described as osteogenic in the sense of the invention.

PEG-grafted copolymers can be synthesized according to procedures described in the literature or in existing patents (see section "Background of the Invention"). The backbone is chosen according to the type of implant or scaffold material surface to which the polymer has to be applied. If the surface of the material to be coated is positively charged at the pH of the coating solution or the pH of later use in application, a polyanionic backbone polymer or poly(amino acid) will be used such as polyacrylic acid or polyglutamic acid. Conversely, if the surface of the device is negatively charged at the relevant pH, then a polycationic backbone polymer or poly(amino acid)will be used such as polyamino(meth)acrylate or polylysine [PCT/USOO/11708 and USSN 09/560,472, WO00065352A1]. The electrostatic interaction between the polyionic backbone of the polymer (with a large number of charges) and the charged material surface ensures a strong binding between copolymer and device to be coated. The molecular weight (MW) of the backbone of the polymer can be changed. Increasing MW leads in general to higher attachment strength of the polymer at the material surface. Alternatively, instead of using electrostatic interaction for the attachment of the copolymer, van der Waals interactions, in particular interactions between hydrophobic parts of the copolymer backbone and the hydrophobic surface can be used to immobilize the polymer. This is a well known mechanism for surface attachment and is used for example in the immobilization of Pluronics-type of molecules such as PPO-PEG to surfaces.

For the synthesis of the copolymer, the amount of backbone and reactive PEG are chosen such that the required grafting ratio g between PEG side chains and backbone monomer results. Preferred grafting ratios are those where there results a high density of PEG chains at the surface of the treated material or device, leading to a non- or little- interactive surface, since this is the regime where osteogenic activities are observed. The range of grafting ratios where osteogenic properties are given, depends on the MW (or size) of the PEG used and on the amount of total copolymer that adsorbs to the surface.

Furthermore, there remains the possibility to use a fraction of PEG chains linked to cell- interactive peptides. For osteoblasts, integrin-interactive peptides such as those containing the active sequence RGD, or heparin-interactive sequences such as FHRRIKA or KRSR are preferred [M.E. Hasenbein et al, Abstract of the Society for Biomaterials Meeting, St. Paul, 2001; Rezania A, Healy KE. Integrin subunits responsible for adhesion of human osteoblast-like cells to biomimetic peptide surfaces. J Orthop Res. 1999 Jul;17(4):615-23]. This may be advantageous in the sense of increasing the number of bone-forming cells at the interface between implant and the body. However, its concentration must be kept much lower than in conventional applications, where the aim

^• is to induce strong cell attachment and spreading. Typically, the peptide density has to be kept below 1% of the PEG chains being peptide modified. Low peptide surface densities can be easily achieved through exposing the surface to a solution of unmodified and peptide-modified copolymer. Through the mixing of the two polymers at any ratio, the peptide concentration at the surface can be easily tailored to the needs of the application. Finally, in application where there is a need to have a combination of cell-surface interactions that lead to strong expression of beneficial factors such as growth factors mineralization marker osteocalcin, high differentiation rate and fast osteoblast phenotype expression, and cell-surface interactions that induce strong cell attachment, spreading and proliferation, patterned surfaces can be introduced that consist of two differently composed areas: a) areas according to this invention having osteogenic properties, and areas with high densities of ligands inducing strong interactions and proliferation. The latter areas may be standard surfaces (e.g. uncoated, unmodified substrates such as metals or metal oxides or polymers), or they may be designed to have a higher density of peptide or protein ligands at the surface fabricated by any type of immobilization strategies.

The copolymer is applied to the surface by exposing the material, implant or device to a solution of the polymer at a concentration and volume-to-surface-ratio that guarantees full coverage of the surface. Typically, polymer concentrations in the range of 0.01 to 10 mg/mL are used. The polymer spontaneously assembles at the surface. Typical assembly times are 1 to 30 min. The substrate has first to be thoroughly cleaned. A number of techniques are suitable: solvent cleaning, alkaline or acid etching of metals, plasma cleaning and oxidation, etc.

Materials to which the invention can be applied are numereous: metals, in particular those metals that are covered by a natural oxide film (e.g. titanium, aluminum, niobium, tantalum, zirconium, hafnium, molybdenum, tungsten, and its alloys), metals with an artificially prepared oxide film (e.g. by anodization), ceramic materials (titanium oxide, zirconium oxide, aluminum oxide, etc.) and polymers. If the copolymer should attach through electrostatic interaction, most metal oxides and metal-oxide-covered surfaces have a sufficient (depending on the isoelectric point: negative or positive charge at the pH of use, i.e. generally 7). However, polymers often have to be treated to increase the number of charged species at the surface, Depending on the polymer type, this can be done by a number of techniques such as chemical or plasma etching or oxidation (e.g. introducing negatively charged carboxy groups), by flame oxidation, etc. Alternatively, the polymer (or any other material in fact) can be coated by a thin film of a suitable metal oxide with a preferred isoelectric point. If the interaction should be of the hydrophobic-hydrophobic type, then the surface may have to be hydrophobized by applications of a thin film of a hydrophobic polymer or by self-assembly processes, such as alkane phosphate monolayers that assemble spontaneously at metal oxide surfaces. Another application can be found in the context of tissue engineering scaffolds, which are used to produce bone or bone-like tissues in vivo or in vitro. Here again, the provision of surfaces within the three-dimensional scaffold, which have osteogenic properties, can improve the speed of bone formation and/or the quality of the newly formed bone. Such scaffolds may be made from polymers or inorganic material, either as resorbable or stable materials. Moreover, bone substitutes such as resorbable or non-resorbable calcium phosphates are another class of materials to which the polymer can be applied in order to impart onto such materials osteogenic properties or increase their genuine osteogenic properties.

There is no limitation in terms of the dimension and form of a part to be coated by the technique. Since it is a spontaneous adsorption process, even very complex implants such as screws or cages for spinal surgery can be easily coated.

Methods of Use

The type of surfaces described here have a variety of applications in the area of metallic, metal-coated, ceramic or polymeric devices for biomedical implant applications in bone, where the stimulation of an osteogenic response and the regulation of corresponding pathways are desired. In particular, the invention relates to the use in medical situations where a fast integration of the device in bone (osseointegration) through rapid differentiation of osteogenic cells, increased expression of factors that stimulate healing and fast formation of neobone tissue is aimed at. The envisaged profit is twofold: Firstly, earlier stabilization of the implant in bone through rapid formation of functional tissue around the implant allows the patient to take up functions (mechanical load on joints, use of dental root implants, etc.) sooner than would be the case otherwise. Secondly, osteogenic surfaces produce a higher amount of direct bone-implant contact and therefore better guarantee long-term stability and permanent medical success in the clinical application. This is particularly important in the area of permanent implants, e.g. artificial joints or dental root implants.

Typical clinical applications of such osteogenic surfaces based on PEG-grafted copolymers, are for example: dental root implants, maxillofacial implants (plates, screws, wires), spinal surgery devices (screws, plates, wires, cages), osteosynthesis implants and devices to stabilize fractures (screws, plates; internal and external fixation devices ["fixateur interne", "fixateur externe"]), artificial hip and knee joints. Furthermore, the technique of making surfaces according to the invention could be applied to scaffolds for supporting boneous tissue growth ex vivo. The surface treatment according to the invention can be applied to any implant or scaffold surface that interacts strongly with the backbone of the PEG-grafted polymer, either through electrostatic interactions between oppositely charged polymer backbone and implant surface, and/or by van der Waals interactions (e.g. hydrophobic-hydrophobic interaction) between the polymer backbone and the surface to be coated. Charged surfaces are preferentially oxide-passive metal device, meaning metals that have a naturally formed, protective oxide layer which carries a charge at the pH of use (e.g. pH of 7.4 for physiological conditions). Examples include titanium and its alloys, iron-based alloys, steel, tantalum and its alloys, niobium and its alloys, hafnium and its alloys, cobalt-chromium, cobalt-chromium-molybdenum. Furthermore, the surface treatment according to the invention can be applied to stable or bioresorbable polymers such as polyethylene, polystyrene, polytetrafluoroethylene (PTFE), polymethylmetacrylate, polyurethanes, poly(lactic acid), poly(glycolic acid or corresponding copolymers. The surfaces of the devices, in particular of polymers, may be charged naturally or they can be made charged positively or negatively by introducing charged surface species. The application of the PEG-grafted copolymer to the device surface is in general through a simple dipping process in a solution of the PEG-polymer in aqueous or non-aqueous solution. This ensures a complete coating of the surface of three-dimensional devices. However, it is also feasible to apply the PEG-polymer selectively to a particular area of the implant, while leaving other adjacent areas of the same device uncoated. This can be accomplished by e.g. spray or brush coating. In order to ensure sterility of the medical implant, either sterile PEG-grafted copolymer solutions (e.g. sterile-filtrated) are applied to a previously (e.g. gamma-ray-) sterilized device and packed under sterile conditions, or the final, coated device is sterilized at the end, using a technique and conditions that do not impair the quality of the osteogenic surface.

Example

Modification of titanium surfaces with PEG-grafted copolymer with and without peptide-grafting and osteoblast response to these surfaces Introduction

Previous studies have shown that osteoblasts are sensitive to surface topography. On rougher surfaces, cells exhibit morphology more typical of differentiated osteoblasts. This is supported by increased alkaline phosphatase specific activity (ALP) and osteocalcin production, as well as increased production of local factors, and suggests that differences in cell attachment may contribute to phenotypic expression. To test this hypothesis, we examined the response of MG63 osteoblast-like cells to smooth and rough Ti surfaces that were modified using poly-L-lysine-g-poly(ethylene oxide) (PEG) to have increased attachment sites.

Methods The rough Ti surface was prepared by grit blasting and acid etching, resulting in macropits (ca. 60 μm in diameter) and micropits (ca. 1 μm in diameter) across the surface (SLA, Institut Straumann). These disks were cut to fit into the well of a 24-well plastic tissue culture plate. Smooth surfaces were 8 x 8 mm silicon wafers coated with 100 nm thick metallic Ti. The Ti-coated wafers were placed in 8-well glass tissue culture slides. Both surfaces had a natural, amorphous 5 nm thick layer of titanium oxide (TiO₂). Each surface was further modified by treatment with PEG brush copolymers with either 5% RGD peptide, which binds integrin receptors and thereby increases cell attachment, or 5% RDG peptide, which has no effect on integrin binding. Plastic and glass culture dishes, and Ti wafers and SLA disks with and without PEG were used as controls. Surfaces were rinsed in distilled water and sterilized under oxygen plasma for 3 min. prior to use in culture. Analyses were performed 24 h after MG63 osteoblast-like cell cultures had achieved confluence on the plastic surface. Cell proliferation was assessed by measuring cell number; changes in differentiation were assessed by measuring alkaline phosphatase specific activity (ALP) and osteocalcin production; and changes in local factor production were determined by measuring TGF-β and PGE2 production. Differences between groups were assessed by performing ANONA; when differences were detected, post-hoc testing was performed using Bonferroni's modification of Student's t-test. Code of polymers/surfaces: PLL-g-PEG is called "PEG", PLL-g-PEG-RGD is called "PEG-RGD", and PLL-g-PEG-RDG is called "PEG-RDG".

Results

Cell number was reduced on smooth Ti wafers and SLA, and further decreased on surfaces coated with PEG or PEG-RDG, but was comparable to glass in cultures grown on PEG-RGD. Differentiation was also affected and the response was both dependent on surface topography and surface treatment. Cellular ALP was reduced on all smooth Ti surfaces coated with PEG, whether they were functionalized or not. In contrast, enzyme activity was stimulated on the SLA surfaces coated with PEG and PEG-RDG, but not with PEG-RGD. A similar response was evident when ALP was assayed using cell layer homogenates, which also contain ALP-enriched matrix vesicles. Cell layer ALP was reduced on all smooth Ti surfaces and there was a further reduction when the surface was coated with PEG alone or PEG-RDG, but not with PEG-RGD. The reverse effect was seen in cultures grown on SLA. Cell layer ALP was increased in all SLA cultures and the effect was greatest in cultures grown on PEG and PEG-RDG, whereas PEG-RGD was comparable to SLA alone. Osteocalcin levels were increased on all Ti surfaces, whether they had smooth or rough topographies, although absolute levels were greater in cultures grown on SLA. On smooth Ti, PEG alone and PEG-RDG supported a 2-fold increase in osteocalcin. On SLA, there was > 6-fold increase in cultures grown on PEG or PEG- RDG. Cells grown on PEG-RGD produced less osteocalcin than those on Ti or SLA alone. Levels of TGF-βl in the conditioned media of cells cultured on smooth Ti treated with either PEG or PEG-RDG was increased more than 3-fold but in cells cultured on SLA coated with the same materials, TGF-βl levels were increased 20 fold; levels on PEG-RGD were comparable to Ti or SLA alone. PGE2 levels were increased on all Ti surfaces as well. PEG and PEG-RDG supported a 20-fold increase on smooth Ti, but PEG-RGD had no effect compared with smooth Ti alone (Figure 2A). On SLA, however, PEG caused a 100-fold increase over plastic and a 50-fold increase over SLA alone (Figure 2B). This stimulatory effect was decreased in cultures grown on PEG- RDG by 30% and by 75% in cultures grown on PEG-RGD.

Conclusions These results were unexpected. PEG alone stimulated differentiation, as measured by osteocalcin production, of the cells and this effect was similar to the effect of PEG complexed with the nonsense peptide RDG. In contrast, the integrin attachment peptide RGD had no effect over that of surface topography alone, or reduced the surface roughness effect, or reduced the PEG effect. These observations suggest that PEG may have osteogenic properties, mediated in part by PGE2. Moreover, decreased attachment of the cell to the surface may also favor expression of the osteoblast phenotype. This is consistent with the observation that osteoblasts assume a less flattened morphology when cultured on rough surfaces.

Claims

Claims

1. Device with a surface covered by a thin film or monolayer of a poly(ethylene glycol)- grafted copolymer exhibiting osteogenic properties, defined in an osteoblastic cell culture test [References: Martin JY, Schwartz Z, Hummert TW, Schraub DM, Simpson J, Lankford J, Dean DD, Cochran DL, Boyan BD (1995) Effect of titanium surface roughness on proliferation, differentiation, and protein synthesis of human osteoblast-like cells (MG63), J Biomed Mater Res 29:389^101; Kieswetter K, Schwartz Z, Hummert TW, Cochran DL, Simpson J, Dean DD, Boyan BD (1996) Surface roughness modulates the local production of growth factors and cytokines by osteoblast-like MG63 cells, J Biomed Mater Res 32:55-63] as a surface that increases the expression of growth factor TGF-βl, the expression of hormone PGE and/or the expression of mineralization marker osteocalcin by at least a factor of 2 in comparison to the corresponding control surface consisting of either unmodified substrate surface or of tissue-culture polystyrene (TCPS).

2. Device according to claim 1 wherein the density of the polyethylene glycol (PEG) chains (grafting ratio) and their molecular weight is adjusted to optimize the osteogenic activity.

3. Device according to claim 1, wherein the copolymer is a PEG-grafted polycationic poly(amino acid) or a PEG-grafted polycationic synthetic polymer or a PEG-grated polycationic polysaccharide that adsorbs spontaneously from solution onto surfaces that are negatively charged at the pH of adsorption and/or pH of use (i.e. 7.4 in biological media).

4. Device according to claim 3, wherein the polymer is a PEG-grafted poly(amino acid) with a polycationic backbone made of lysine, histidine, arginine or ornithine in D-, L- or DL configuration, or the polymer is a PEG-grafted polymer with a cationic backbone of a polysaccharide such as chitosan, partially deacetylated chitin, and amine-containing derivatives of neutral polysaccharides, or the polymer is a PEG- grafted non-peptide polyamine with a polycationic backbone such as poly(aminostyrene), poly (aminoacrylate), poly (N-methyl aminoacrylate), poly (N- ethylaminoacrylate), poly(N,N-dimethyl aminoacrylate), poly(N,N- diethylaminoacrylate), poly (aminomethacrylate), poly (N-methyl aminomethacrylate), poly(N-ethyl aminomethacrylate), poly(N,N-dimethyl aminomethacrylate), poly(N,N-diethyl aminomethacrylate), poly(ethyleneimine), polymers of quaternary amines, such as ρoly(N,N,N-trimethylaminoacrylate chloride), poly(methacrylamidopropyltrimethyl ammonium chloride), or the polymer is a PEG-grafted charged synthetic polymer with a polycationic backbone such as polyethyleneimine, polyamino(meth)acrylate, polyaminostyrene, polyaminoethylene, poly(aminoethyl)ethylene, polyaminoethylstyrene, and N-alkyl derivatives thereof. Device according to claim 3 where the substrate material is a metal or metal-oxide- covered surface or a ceramic surface or a polymeric surface that adopts a negative charge at the pH of adsorption and/or the pH of use (i.e. 7.4 in biological media). Device according to claim 1, wherein the copolymer is a PEG-grafted polyanionic synthetic polymer or a PEG-grafted poly(amino acid) that adsorbs spontaneously from solution onto surfaces that are positively charged at the pH of adsorption and/or pH of use (i.e. 7.4 in biological media). Device according to claim 6, wherein the copolymer is a PEG-grafted copolymer with an anionic backbone of a poly(amino acid) where the amino acid contains an additional pendant carboxy group imparting a negative charge to the backbone at pH above 4 and in particular at neutral pH such as polyaspartic acid or polyglutamic acid; or a natural or unnatural polymer with pendant negatively charged groups, particularly carboxylate groups, including alginate, carrageenan, furcellaran, pectin, xanthan, hyaluronic acid, heparin, heparan sulfate, chondroitin sulfate, dermatan sulfate, dextran sulfate, poly(meth)acrylic acid, oxidized cellulose, carboxymethyl cellulose and crosmarmelose, synthetic polymers and copolymers containing pendant carboxyl groups, such as those containing maleic acid or fumaric acid in the backbone. Device according to claim 6 where the substrate material is a metal or metal-oxide- covered surface or a ceramic surface or a polymeric surface that adopts a positive charge at the pH of adsorption and/or the pH of use (i.e. 7.4 in biological media). Device according to claim 3, 4, 6 or 7, wherein the PEG molecular weight is between 500 and 20,000, preferentially between 2000 and 5000, and the grafting ratio lysine monomers/PEDG chains is between 2 and 20, preferentially between 2.5 and 7. Device according to claim 3, 5, 6 or 8, wherein the material to be coated is a biocompatible metal covered by a natural or artificially strengthened oxide film, such as titanium, niobium, tantalum, zirconium, hafnium, molybdenum, tungsten, cobalt- chromium, cobalt-chromium-molybdenum, stainless steel, as well as further alloys composed of these materials, in particular titanium-aluminum-vanadium and titanium-aluminum-niobium alloys. Device according to claim 3, 5, 6, 8, wherein the material to be coated is a biocompatible polymer with a surface that is charged or made charged at the pH of adsorption or pH of use and is used in combination with an oppositely charged PEG- grafted copolymer. Device according to claim 11, wherein the polymer is made charged in contact with an aqueous solution through prior application of a surface modification technique, such as a wet-chemical, gas-phase-chemical, plasma or flame exposure treatment with the effect that charged functional groups are introduced in the surface of the polymer. Device according to claim 3 or 6, wherein a fraction of the PEG chains are linked at the end of the chain to a specific peptide in order to increase the interactivity, as measured by cell attachment and spreading, of the surface with particular types of cells. Device according to claim 13, wherein the peptide density is so low that the osteogenic activity, as defined in claim 1, is still present. Device according to claim 14, wherein less than 1% of all PEG chains at the surface are modified with a peptide. Device according to claim 13, wherein the peptide contains a sequence of an integrin- receptor-type binding type of peptide such as RGD or of a heparin-domain-binding type of peptide such as KRSR or FHRRIKA or another typical cell binding motif. Device according to claim 1 where the surface of the same device contains a pattern with areas of low interactiveness and high degree of osteogenic activity, where cells strongly express factors that are beneficial to the healing and new bone forming process, and areas of higher interactiveness where stronger cell attachment and spreading takes place inducing stronger proliferation of bone-forming cells. Device according to claim 17, wherein the pattern area that fosters osteogenic activity is made according to any combination of claims 1 to 14, and where the cell- interactive areas are composed of unmodified substrate surface or are made according to claims 13 and 14, but with a higher peptide density that induces stronger interaction and proliferation. Device according to any combination of claims 1 to 18, wherein the device is an implant for application within bone structure or in contact with bone. Device according to claim 19, wherein the application of the device is in the areas of dental implantology, maxillofacial surgery, osteosynthesis, spinal surgery, or orthopedics. Device according to any of the claims 1 to 18, wherein the device to be coated is a scaffold for growing boneous tissue ex vivo or in vivo by tissue engineering. Device according to claim 19, 20 or 21, wherein the material to be coated is a bone substitute such as resorbable or non-resorbable calcium phosphates. Device according to claim 21 wherein the scaffold is made from a non-resorbable biocompatible material or from a resorbable biocompatible material.