Implantable materials in contact with blood and uses thereof

Abstract

The present invention is directed to blood compatible materials, implantable devices comprising said materials, methods for the preparation of such materials, and medical devices coated therewith, as well as uses thereof for anti-thrombotic and / or cell proliferation aspects.

Description

[Technical field]

[0001] The present invention relates to implantable materials in contact with blood, and their preparation and application, for example in implantable devices such as vascular grafts, shunts, patches, stents, valves, etc. In particular, the present invention relates to biocompatible materials and implantable artificial devices (endoprostheses) useful for the treatment of humans or animals by surgery or therapy, especially of cardiovascular disease, in particular of cardiovascular ischemic disease, degenerative or congenital disease, e.g., coronary artery disease or peripheral vascular disease, or access surgery for dialysis patients, pediatric cardiovascular malformations, trauma, and repair surgery. [Background technology]

[0002] Cardiovascular ischemic disease is the number one cause of death in the Western world, and vascular grafts with small internal diameter (ID) remain a challenge for material scientists. The main failure mechanisms associated with currently commercially available grafts include early thrombosis / occlusion and lack of rapid and confluent endothelialization, as well as late intimal hyperplasia that can lead to stenosis and occlusion.

[0003] This suggests the need for the development of improved artificial vascular conduits that do not cause pathological reactions after implantation. The graft should resemble native blood vessels in terms of biocompatibility, antithrombogenicity, shape, size, and mechanical properties, compliance, regeneration, and growth potential, and resistance to infection. The use of naturally derived materials (e.g., collagen or decellularized tissue matrices) carries the risk of immunological reactions leading to premature degradation causing aneurysms and rupture. The production of fully cell-derived grafts is very time-consuming, taking about 6 months. Thus, in clinical practice, surgery using small-caliber vascular grafts (less than 6 mm) is mainly performed with autologous arteries or veins, which may, however, be pathological or may have been used in previous surgery, and thus there is a strong demand for ready-to-use small-caliber vascular grafts. Synthetic polymers have the advantage that they can be manufactured on a large scale with control over (physical and chemical) properties such as strength, degradation rate, and reproducible microstructure. Commercially available grafts are made of non-degradable polymers such as polyethylene poly(ethylene terephthalate) (PET), Dacron®, or expanded poly(tetrafluoroethylene) (ePTFE) or poly(urethane) (PU), or others, and are widely used as non-degradable synthetic vascular grafts.

[0004] It has been used successfully to replace larger caliber vessels, but has failed as small caliber prostheses (ID<6mm) required for smaller vessels and / or anastomosis in microvascular surgery. Vascular grafts with small internal diameters pose a challenge to material scientists. It may be employed in the treatment of cardiovascular ischemic diseases such as coronary artery disease or peripheral vascular disease, which still ranks first among non-communicable causes of death in the Western world. In addition, it may be necessary to treat end-stage renal disease in access surgery. Whenever possible, bypass surgery is performed with autologous arteries or veins, because the results of synthetic vascular grafts show poor clinical outcomes when used in small caliber grafts (ID<6mm). In contrast to larger caliber grafts, prostheses with internal diameters less than 6mm increase hemodynamic flow disturbances and thrombotic complications. Therefore, endothelialization of small caliber grafts should form an interface compatible with hemodynamic and functional processes. The main failure mechanisms are early thrombosis (cell surface interactions) and lack of rapid and confluent endothelialization. This failure is mainly a result of surface thrombogenicity due to lack of endothelial cells and unfavorable healing process with anastomotic intimal hyperplasia caused by hemodynamic disturbance. Intimal hyperplasia may subsequently lead to stenosis and occlusion, often requiring reoperation or amputation in distal cases. All these limitations also apply to any device (valves, VADs, etc.) that is implanted in contact with blood. As a result of the above, synthetic grafts are not used for coronary surgery. Therefore, the development of an ideal small-caliber prosthesis is a major challenge in vascular research.

[0005] Numerous surface coatings have been developed over the last 30 years to improve the hemocompatibility of biomaterials, and as reviewed in Biran et al., 2016, Adv. Drug Deliv. Rev. 112, 12-23, the anticoagulant heparin has been employed as a covalently immobilized surface coating onto a variety of medical devices to improve hemocompatibility.

[0006] To further improve the hemocompatibility (i.e., heparin conjugation) of expanded poly(tetrafluoroethylene) (ePTFE), so-called non-treated surfaces, and / or polycarbonate-urethane (PCU) surfaces, numerous strategies of surface modification (e.g., chemical immobilization, physical adsorption, and plasma treatment) have been explored. It has been shown that PCU surface functionalization achieved using plasma treatment is significantly more effective in generating functional amine groups on the surface compared to alternative methods such as aminolysis and physical adsorption by polydopamine coating, and that these functional amine groups grafted onto the vascular graft surface can then be used for the subsequent conjugation of heparin by reductive amination, allowing terminal immobilization of heparin molecules on the graft surface (Qiu et al., 2017, Acta Biomater., 51, pp. 138-147). Conjugation of heparin by plasma treatment followed by reductive amination has been demonstrated to advantageously provide not only higher surface density but also better stability and antithrombogenic activity, and improve the performance of vascular grafts with respect to patency and the early stages of endothelialization and graft integration.

[0007] Biodegradable polymers have also been investigated for their potential as scaffolds for the fabrication of small-diameter arterial bypass grafts. Two methods of aminolysis of electrospun poly-L-lactide-co-caprolactone (PLCL) microfibrous vessels were compared: plasma treatment and fmoc-PEG-diamine insertion for terminal heparin conjugation to PLCL scaffolds for optimizing heparin density and bioactivity on biodegradable electrospun PLCL microfibrous vessel grafts (Hsieh et al., 2017, Biomed. Mater., 12, 6). From this study, it was concluded that the superiority of plasma treatment in terms of initial heparin density was achieved, while chemical introduction of amino groups for terminal conjugation of heparin (by fmoc-PEG-diamine insertion) caused partial fiber erosion, dissolution, and structural destruction, while inducing minimal polymer chain degradation. However, plasma activation of surfaces is highly aggressive towards support materials and often leads to undesirable defects in the structure and mechanical properties of implantable materials, and may even hinder further biomedical uses of biodegradable materials such as PCL (polycaprolactone).

[0008] Gao et al., 2018, Regen. Biomater., 5, 105-114, recently prepared small-diameter vascular grafts based on a matrix of electrospun poly(ε-caprolactone) that was subsequently coated with alternating layers of selenium-containing catalyst-organoselenium-modified polyethyleneimine and heparin (layer-by-layer (LbL) assembly) to combine the effects of heparinization and catalytic NO generation to stimulate tissue regeneration, angiogenesis, and vascular remodeling processes. In this assembly, heparin is used as a polyanion for LbL-formation. [Prior art documents] [Patent documents]

[0009] [Patent Document 1] U.S. Patent No. 4,613,665 [Non-patent literature]

[0010]

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[0011] Therefore, there is a need to discover suitable materials and / or coatings that allow for the preparation of small inner diameter vascular grafts that allow for early, transient anti-thrombotic properties that prevent early thrombotic events and allow for early, rapid and confluent endothelialization of the graft. [Means for solving the problem]

[0012] The present invention relates to the unexpected discovery of a novel biocompatible material suitable for in vivo tissue regeneration that allows the preparation of small-diameter grafts (1-6 mm inner diameter) useful in coronary artery bypass grafts (CABG), as described, for example, in Virk et al., 2019, Curr. Cardiol. Rep., 21, 36. The novel material and coating exhibit mechanical properties suitable for matching arterial compliance with strain and stress better than native arteries and for enhancing suture retention and preventing aneurysm formation and rupture, thus allowing its use in vascular grafts, shunts, patches, endoprostheses, valves, and any implantable device in contact with blood. The method developed to prepare such a material shows the main advantage of resulting in materials with highly reproducible structural properties compared to chemical or plasma functionalization, and is more time-efficient than methods of preparing grafts that require in vitro cellular manipulation, such as the formation of an extracellular matrix based on cell type, ingrowth, and duration. (L'Heureux et al., 2006, Nat. Med., 12, pp. 361-365).

[0013] According to another aspect, the method of the present invention confers extended anticoagulant properties to the material. Specifically, the material of the present invention, when present on the luminal surface of a vascular patch or graft, provides an unexpected and long-lasting antithrombotic effect that prevents incipient thrombus formation or occlusion. According to certain embodiments, the material of the present invention, when biodegradable, induces a host response with cell growth and de novo ECM (extracellular matrix) formation in vivo.

[0014] Moreover, according to another particular embodiment, when the multilayer polymer layer-by-layer (LbL) coating of the material of the invention comprises at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair of said coating, the material of the invention further exhibits the advantage of imparting favorable cell attachment and growth properties by inducing rapid endothelialization due to the high affinity for endothelial cell growth on the surface. Thus, the material and method of the invention make it possible to achieve the surprising properties of extended anticoagulant properties while maintaining cell-attracting properties.

[0015] Another particularly advantageous and non-obvious aspect is the discovery of compositions and process steps that allow for the combination of the above functions on heterogeneous substrate surfaces without the use of aggressive coupling reactions that would chemically modify the surface or alter the mechanical properties of the surface.

[0016] Thus, the materials and methods of the present invention are expected to have beneficial effects with respect to any biocompatible implantable material (particularly materials for use in implantable devices) that is intended to come into contact with blood upon implantation.

[0017] An aspect of the present invention provides a blood compatible material comprising: (i) a substrate comprising a biocompatible nonwoven fiber; and (ii) a heparinized multi-layer polymeric layer-by-layer (LbL) coating on the biocompatible nonwoven fiber, the multi-layer polymeric coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer layer, the multi-layer polymeric layer-by-layer (LbL) coating being functionalized with heparin on an outer layer pair of the multi-layer polymeric coating by terminal functionalization, the anionic polymer being selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or mixtures thereof, and the cationic polymer being selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or mixtures thereof.

[0018] Another aspect of the invention relates to an implantable device comprising the material according to the invention.

[0019] Another aspect of the invention relates to extracorporeal circulation devices, such as membranes and catheters, oxygenators, and hemodialysis / filtration systems that include the materials of the invention.

[0020] Another aspect of the present invention relates to a method for preparing the blood compatible material according to the present invention.

[0021] Another aspect of the invention relates to a method of preparing the implantable device of the invention.

[0022] Another aspect of the invention relates to a method of treating cardiovascular ischemic disease, e.g., coronary artery disease or peripheral vascular disease, pediatric cardiovascular malformation, trauma, or organ defect, e.g., end-stage renal disease (ESRD) or failure, by reparative surgery in a subject in need thereof, comprising implantation of an implantable device comprising a material of the invention in contact with the blood of the subject.

[0023] Another aspect of the invention relates to the use of the material according to the invention or the method according to the invention for the preparation of an implantable device or an extracorporeal circulation device in contact with blood. [Brief description of the drawings]

[0024] [Figure 1] FIG. 1 shows the chemical structures of (a) poly(allylamine hydrochloride) (PAH), (b) poly(sodium 4-styrenesulfonate) (PSS), (c) fully deacylated chitosan (CHI), (d) dextran sulfate sodium salt (DS), (e) poly(L-lysine) (PLL), and (f) N,N,N-trimethylchitosan (TMC). [Figure 2A] 1 is a schematic diagram showing the structure of a blood compatible material of the present invention, the material comprising (i) a substrate (1) comprising a biocompatible nonwoven fiber; (ii) a heparinized multi-layer polymeric layer-by-layer (LbL) coating of the present invention (2); and an optional pre-coating deposit (3), where heparin (4) is end-functionalized to the multi-layer polymeric coating (2). [Figure 2B] 1 is a schematic diagram showing the structure of a blood compatible material of the present invention, the material comprising: i) a substrate (1) comprising biocompatible nonwoven fibers; (ii) a heparinized multi-layer polymeric layer-by-layer (LbL) coating of the present invention (2); and an optional pre-coating deposit (3), the multi-layer polymeric coating (2) comprising at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair (21) forming a heparinized substrate for terminal functionalization of heparin (4), the lower layer of the multi-layer polymeric coating being formed of a layer pair (22) of another polymer of the present invention. [Figure 2C]1 is a schematic diagram showing the structure of a blood compatible material of the present invention, comprising: i) a support (1) comprising a biocompatible erodible fiber; (ii) a heparinized multi-layer polymer layer-by-layer (LbL) coating (2) of the present invention; and an optional pre-coating deposit (3), wherein the multi-layer polymer coating (2) comprises at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair (21) forming a heparinized substrate for terminal functionalization of heparin (4), and the lower layer of the multi-layer polymer coating (2) is formed of layer pairs of the same polymers. [Diagram 3] FIG. 1 shows the thickness (nm) of (CHI / HEP)n and (CHI / DS)n multilayer polymer coatings versus the LbL layer number (i.e., the number of layer pairs of anionic and cationic polymer layers deposited on a silicon wafer as described in Example 1). Note that a slope is obtained at higher layer numbers since the thickness of the first deposited layer is influenced by the substrate. [Figure 4A] FIG. 1 shows the antithrombotic activity of various materials 1, 2 and 3 of the invention compared to comparative coatings C1-C6, expressed as anti-Xa values ​​(UI / ml) measured after 0 h, 24 h, 7 days and 10 days of incubation. [Figure 4B] FIG. 1 shows the antithrombotic activity of various materials 1, 2 and 3 of the invention compared to comparative coatings C1-C6, expressed as anti-Xa values ​​(UI / ml) measured after 1, 2.5, 5 and 7 days of incubation. [Figure 4C] FIG. 2 shows the thrombin generation inhibition (EPT: nmol / L / min) capacity measured according to the assay of example 2b after 0, 2.5 and 7 days of incubation. [Diagram 5] FIG. 1 shows the antithrombotic activity of materials 4 and 5 of the invention, expressed as anti-Xa values ​​(UI / ml), measured after 1, 2.5, 5 and 7 days of incubation as described in Example 2. [Figure 6]FIG. 4B depicts semi-quantitative cell proliferation scores (0-4) of smooth muscle cells (SMC) (grey) and endothelial cells (EC) (black) from in vitro cell cultures tested on inventive materials 1 and 2 compared to comparative coatings C1 and C6 described in Example 2c after 5 days of incubation for the same samples shown in FIG. 4B and FIG. 4C. The higher the score number of endothelial cells (EC) on the cell-attracting coating, the better the endothelialization. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0025] The term "heparin" refers to a heparin molecule, a fragment of a heparin molecule, or a heparin derivative. Heparin is a polysaccharide belonging to the group of glycosaminoglycans with sulfonic groups distributed along the repeating unit. The structure of heparin contains specific binding sequences that interact with proteins in the blood, in particular, for example, thrombin (T) and antithrombin (AT), which catalyzes the inhibition of factor Xa. In an attempt to circumvent the problems of intravenously administered heparin, this molecule is most commonly immobilized on material surfaces by three different methods: ionic (electrostatic interactions), multipoint attachment (covalent bonds), and end functionalization (covalent bonds). Attractive advantages of using this latter method include the orientation of the heparin molecule in the bulk and the retention of the active sites responsible for the inhibition of thrombin generation. Heparin derivatives can be any functional or structural variation of heparin. Heparin is attached to the last LbL layer of the multilayer polymer layer-by-layer (LbL) coating of the present invention.

[0026] The term "layer-by-layer (LbL) layer" defines a pair of anionic and cationic polymer layers. According to a particular aspect, the number of LbL layers is about 1 to about 8, typically about 1 to about 5 (e.g., 4 or 5). According to a further particular embodiment, for use as a biodegradable implant, it is desirable to keep the number of LbL layers in the material of the invention as low as necessary to reach the intended functionality. As layers are added, the stiffness of the implant increases, as well as the biodegradability decreases. According to another particular embodiment, the multilayer polymer coating comprises a pair of alternating layers of poly(allylamine hydrochloride) (PHA) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair.

[0027] The term "nonwoven fibers" refers to any material construction that does not involve weaving or knitting of fibers. Nonwoven fibers can be nano or micro fibers or fibrils and materials, including nonwoven fibers obtained by various manufacturing procedures such as extrusion, expansion, or electrospinning / spraying.

[0028] The term "chitosan derivatives" refers to N-sulfofurfuryl chitosan (SFC) and N-[(2-hydroxyl-3-trimethylammonium)propyl]chitosan chloride (HTACC) (Channasanon et al., 2007, J. Colloid Interface Sci., 316, 331-343), N-glycidyltrimethylammonium chloride (Cui et al., 2010, Adv. Funct. Mater., 20, 3303-3312); N-[(2-hydroxyl-3-trimethylammonium)propyl]chitosan chloride (HTACC), N-succinyl chitosan (SCC) and N-sulfofurfuryl chitosan (SFC) (Graisuwan, 2012, J. Colloid Interface Sci., 316, 331-343). Sci., 376, 177-188) or synthetic chitosan analogues such as N-quaternized chitosan like N,N,N-trimethylchitosan (TMC) (Martins, 2014, Int. J. Mol. Sci., 15, 20800-20832).

[0029] The term "endothelial cell-attracting coating" refers to a coating that will allow cell attachment and induce rapid endothelialization due to a high affinity for endothelial cell growth on its surface. The endothelial cell-attracting properties of a coating can be determined by performing the assays described herein or any other suitable assays, such as long-term preclinical in vivo implantation studies. An example of such a suitable endothelial cell-attracting coating according to the present invention includes at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair of the coating.

[0030] The term "artificial device" includes artificial devices as defined in the present invention that are intended to contact blood (e.g., implantable devices or extracorporeal circulation devices, as well as catheters and other devices). Artificial devices include biodegradable and non-biodegradable devices. Implantable devices include artificial blood vessels, grafts, or shunts, cardiovascular patches, cardiovascular valves or leaflets, endoprostheses, such as stents, ventricular assist devices, artificial hearts, intravascular devices, such as filters, occlusion devices, and intravascular leads (pacemakers). Extracorporeal circulation devices include devices used for cardiopulmonary bypass, extracorporeal membrane oxygenation (ECMO), or other blood purification devices, such as dialysis, and components thereof.

[0031] The biodegradable device of the present invention may further comprise a mesh, preferably non-biodegradable, to improve mechanical stability during degradation of the biodegradable device, specifically the mesh improves flexibility and imparts anti-kinking properties to the implantable device, especially when inserted into a vascular graft. When used as a venous replacement, the incorporation of the mesh may prevent external compression of the blood vessel. Furthermore, in the case of a biodegradable vascular graft, late aneurysm formation may be prevented.

[0032] The term "biocompatible nonwoven fiber-containing substrate" includes materials suitable for implants, patches, or leaflets (e.g., but not limited to, degradable and non-degradable polymers described in this application).

[0033] The phrase "hemocompatible" refers to a material that can be used in contact with blood without causing harm (specifically, a material that is antithrombotic). According to certain embodiments, hemocompatibility can be assessed by antithrombotic effect over a period of at least one week or more (typically about one week to about one month) in a standard assay such as those described herein.

[0034] As used herein, "treatment" and "treating" and the like generally mean to obtain a desired pharmacological and / or physical and / or physiological effect. This effect may be prophylactic in that it prevents or partially prevents a disease, condition, or state thereof, and / or it may be therapeutic in that it partially or completely cures a disease, condition, symptom, or adverse effects caused by a disease. The term "treatment" as used herein is directed to any treatment of a disease in a mammal, particularly a human, including: (a) preventing the onset of a disease in a subject who may be predisposed to the disease but has not yet been diagnosed as having the disease; (b) inhibiting the disease, i.e., preventing the onset of the disease; or relieving the disease, i.e., causing regression of the disease and / or its symptoms or conditions (e.g., ameliorating or curing damage).

[0035] The term "subject" as used herein refers to a mammal. For example, mammals contemplated by the present invention include humans, primates, and domestic animals, such as cattle, sheep, pigs, horses, particularly race horses, laboratory rodents, and the like.

[0036] The term "efficacy" of the treatment of the present invention may be measured based on the course of disease changes in response to the use of the present invention. For example, the efficacy of the device of the present invention may be measured by the improvement of laboratory and clinical outcomes with the implanted device / material of the present invention.

[0037] Materials according to the invention Referring to the figures, and particularly initially to FIG. 2A, a blood compatible material may be: - a support (1) comprising biocompatible nonwoven fibers; - a heparinized multi-layered polymeric layer-by-layer (LbL) coating (2) on said biocompatible nonwoven fabric, the multi-layer polymer LbL coating comprises at least one layer pair of an anionic polymer layer and a cationic polymer layer, the multi-layer polymer layer-by-layer (LbL) coating being functionalized with heparin on the outer layer pair of the multi-layer polymer by terminal functionalization, the anionic polymer being selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or a mixture thereof, and the cationic polymer being selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or a mixture thereof; and Includes.

[0038] In another further embodiment of the invention, the material of the invention further comprises an optional pre-coating deposit (3) on a support (1) comprising a biocompatible non-woven fiber, said pre-coating deposit comprising an adhesive for a heparinized multi-layer polymer LbL coating (2). In a particular embodiment, the adhesive can be a cationic polymer, for example, poly(ethyleneimine) (PEI), poly(lysine). Another cationic polymer that can be used as an adhesive can be poly(allylamine hydrochloride) (PAH).

[0039] In another further embodiment of the invention, the material of the invention further comprises an optional endothelial cell attractive coating of a multi-layer polymer layer-by-layer (LbL) coating, on which heparin is functionalized by terminal functionalization. According to certain aspects, the endothelial cell attractive coating is particularly advantageous when used in combination with biodegradable nonwoven fibers.

[0040] In one particular embodiment, the support comprising biocompatible nonwoven fibers comprises or consists of biocompatible nonwoven fibers (eg, microfibers and / or nanofibers, or a combination thereof).

[0041] In another embodiment, the biocompatible nonwoven fibers have a diameter of 0.1-10 μm, for example, 0.5-5 μm or 1-5 μm.

[0042] In one particular embodiment, the nonwoven fibers form a three-dimensional network or matrix with a certain porosity defined by the interstices between the fibers.

[0043] Preferably, the nonwoven fibers are unoriented, that is, they do not have any particular orientation along a particular axis or plane, but can be aligned using various electrospinning parameters.

[0044] In one particular embodiment, said nonwoven fabric comprises or consists of several superimposed layers of fibers.

[0045] In a more particular embodiment, said nonwoven fabric comprises or consists of several superimposed layers with different porosities.

[0046] According to another particular embodiment, the nonwoven fibers according to the invention may be prepared as described in Pfeiffer et al., 2014, J. Biomed. Mater. Res., 102A, pp. 4500-4509.

[0047] In another embodiment, the biocompatible nonwoven fibers are biodegradable. In particular, the biodegradable biocompatible nonwoven fibers can be found in biodegradable polymers selected from the group consisting of degradable polycaprolactone (PCL), polyglycolic acid (PGA), polylactic acid (PLA), poly(lactic-co-glycolic acid) (PLGA), poly-L-lactide (PLLA), poly(glycerol sebacate) (PGS), polydioxanone (PDO, PDS), or poly-p-dioxanone, or PGA-poly-4-hydroxybutyrate (P4HB) copolymers, degradable polyurethane fibers, and combinations thereof, as well as any other polymers used in biomedical applications, for example as described in Miranda et al., 2020, Antibiotics, 9, 174.

[0048] In a more particular embodiment, the biodegradable, biocompatible nonwoven fibers comprise or consist of polycaprolactone (PCL) fibers.

[0049] In a particular embodiment, said fibers are ε-PCL nanofibers, preferably electrospun ε-polycaprolactone nanofibers, which can be obtained, for example, according to the method described in Nottelet et al., 2009, J. Biomed. Mater. Res., 89A, pp. 865-875.

[0050] In a specific embodiment, the fibers are polycaprolactone having a molecular weight in the range of 10 to 200 kDa, preferably in the range of 40 to 120 kDa, preferably in the range of 60 to 100 kDa.

[0051] In one embodiment, said nonwoven comprises or consists of one or more layers with a porosity of 0.8-12 μm, which may be measured by the scaffold porosity (e) or void ratio as explained in de Valence et al., 2012, Acta Biomater., 8, 3914-3920.

[0052] In another embodiment, the biocompatible nonwoven fibers are non-biodegradable. Specifically, non-biodegradable biocompatible nonwoven fibers can be found in polymers such as polyethylene terephthalate (PET), Dacron®, or expanded polytetrafluoroethylene (ePTFE) or polyurethane (PU), or others, which are widely used in the clinic as non-degradable synthetic vascular grafts.

[0053] In more specific embodiments, the non-biodegradable, biocompatible, non-woven fabric comprises or consists of ePTFE, for example, as described in Lian et al., 2003, J. Vasc. Surg., 37, 472-80.

[0054] In one embodiment, the multilayer polymer coating is a layer-by-layer (LbL) deposition multilayer coating. For example, the LbL deposition multilayer coating can be produced as described in Picart et al., 2015, Layer-by-Layer Films for Biomedical Applications, C. Picart, F. Caruso and J.-C. Voegel (eds.), Wiley-VCH: Weinheim, Germany; El-Khouri et al., 2011, Functional Polymeric Ultrathin Films, R. Advincula and W. Knoll (eds.), Wiley-VCH: Weinheim, Germany. Typically, the multilayer polymer coating has a total thickness of about 0.1 nm to about 500 nm, typically about 1 nm to about 500 nm, specifically 1 nm to about 100 nm, as measured, for example, by ellipsometry on a silicon wafer, as shown in FIG. 3, with each layer pair having a thickness of about 0.2 nm to about 50 nm (e.g., about 1 to 25 nm).

[0055] In one embodiment, the anionic or cationic polymer has a molecular weight

[0056]

number

[0057] In particular embodiments, suitable anionic or cationic polymers according to the present invention have a molecular weight of less than 75 kDa.

[0058] In one particular embodiment, the anionic or cationic polymer has a molecular weight

[0059]

number

[0060] is in the range of 7.5 kDa to 75 kDa.

[0061] In one embodiment, the anionic polymer is selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or a mixture thereof.

[0062] In another embodiment, the cationic polymer is selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or a mixture thereof.

[0063] In a further embodiment, the anionic polymer is selected from heparin and poly(sodium 4-styrenesulfonate) (PSS) and the cationic polymer is selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH).

[0064] In one particular embodiment, the anionic polymer is heparin.

[0065] In one particular embodiment, the anionic polymer is poly(sodium 4-styrenesulfonate) (PSS).

[0066] In one particular embodiment, the cationic polymer is chitosan or an analog thereof (eg, N-alkylated chitosan, such as trimethylchitosan).

[0067] In one particular embodiment, the chitosan has a molecular weight of about 7.5 kDa to 75 kDa.

[0068] In another particular embodiment, the chitosan has a degree of deacetylation of at least 80%, preferably at least 85%, typically from about 80 to about 95%.

[0069] In another particular embodiment, trimethylchitosan is prepared as described in Mourya et al., 2009, J. Mater. Sci. Mater. Med., 20, 1057-79, or Malik et al., 2018, Int. J. Nanomedicine, 13, 7959-7970, or Kulkarni et al., 2017, Carbohydr. Polym., 157, 875-902.

[0070] In one particular embodiment, the anionic polymer is poly(allylamine hydrochloride) (PAH).

[0071] In a more specific embodiment, the multilayer polymer LbL coating comprises from about 3 layer pairs to about 8 layer pairs of anionic and cationic polymer layers according to the present invention, said layer pairs being arranged such that the anionic and cationic polymer layers alternate in the multilayer polymer coating.

[0072] In a more specific embodiment, the multilayer polymer LbL coating of the material of the invention comprises at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer pair of layers of said coating which are functionalized with terminally functionalized heparin.

[0073] According to a more specific embodiment, the multilayer polymer LbL coating comprises about 1 layer pair to about 8 layer pairs of anionic and cationic polymer layers, wherein the anionic polymer is heparin, the cationic polymer layer is selected from chitosan or analogs thereof, and the alternating layer pair of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair of the coating is functionalized with terminally functionalized heparin.

[0074] End-functionalization of the membrane with heparin is achieved by reaction of heparin with amine groups in the outer layer of the multi-polymer layer-by-layer coating.

[0075] In one particular embodiment, a substrate comprising biocompatible nonwoven fibers may first be pre-coated with a deposit comprising an adhesive for a multi-layer polymeric coating.

[0076] In certain embodiments, the adhesive is a cationic polymer that contains multiple primary or secondary amino groups (eg, poly(ethyleneimine) (PEI)).

[0077] Referring to the figures (particularly initially to FIG. 2), a heparinized multilayer polymeric LbL coating (2) is obtained by functionalization of heparin (4) on the outer layer of said multilayer polymeric coating (2) by terminal functionalization, independent of the fact that heparin can be used as an anionic polymer in the multilayer polymeric LbL coating (2).

[0078] Various end-functionalization strategies for heparin are known, as described, for example, in Biran et al., 2016, Adv. Drug Deliv. Rev., 112, 12-23.

[0079] According to further specific embodiments, heparin is functionalized on the outer layer pair of the multilayer polymeric LbL coating via end-attachment (EPA), in which the reducing end of heparin (typically first partially depolymerized by nitrous acid deamination) is covalently bonded to the amino functionality of the final layer of the multilayer polymeric LbL coating by reductive amination.

[0080] According to further particular embodiments, heparin is functionalized on the outer layer pair of said multilayer polymeric LbL coating via terminal conjugation (EPC), in particular heparin is bound to the amino functional groups of the final layer of the multilayer polymeric LbL as described, for example, in Hsieh et al., 2017, supra.

[0081] According to more specific embodiments, the heparin is functionalized after diazotization of the heparin to provide highly reactive aldehyde groups on the heparin prior to functionalization onto the polymer coating.

[0082] Moreover, according to certain embodiments, the terminal conjugation of heparin makes it possible to provide sufficient access for endothelial cells to the outer layer of the LbL coating, particularly the cell-attracting coating of a multi-layer polymeric layer-by-layer (LbL) coating in which heparin is functionalized by terminal conjugation.

[0083] Referring to the figures (particularly initially to Figures 2B and 2C), in another particular embodiment, a multilayer polymer LbL coating of a substrate comprising biocompatible nonwoven fibers includes at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair (21) to form a heparinized substrate for end-functionalization of heparin (4). In this case, heparin (4) is end-functionalized to the multilayer polymer coating (2), and the bottom layer of the multilayer polymer coating is formed of a layer pair of either the same polymer (Figure 2C) or preferably another polymer described herein ((22) in Figure 2B).

[0084] According to a more specific embodiment, a multi-layer polymer LbL coating of a substrate comprising a biocompatible nonwoven fiber is provided in which the substrate (optionally pre-coated with a pre-coating agent) is first coated with an anionic polymer of a first LbL layer pair.

[0085] In another particular embodiment, the multilayer polymer LbL coating of the substrate comprises several layer pairs of anionic and cationic polymer layers, where the anionic polymer layers can be formed of the same or different anionic or cationic polymers between the various layer pairs.

[0086] In another particular embodiment, the multilayer polymer LbL coating of the substrate comprises about 3-8 (particularly 3-5) alternating anionic / cationic polymer layer pairs, which polymer layer pairs are selected from heparin / chitosan or an analogue thereof layer pairs, and PAH / PSS layer pairs, or combinations thereof.

[0087] In another particular embodiment, the multilayer polymer coating of the implantable material support comprises about 3-8 (e.g., about 3-7, e.g., about 3 to about 5, particularly 4-5) layer pairs of anionic and cationic polymer layers, where the anionic polymer is heparin and the cationic polymer is chitosan or an analog thereof.

[0088] In another particular embodiment, the multilayer polymer LbL coating of the substrate comprises about 3 to 8 (e.g., about 3 to 7, e.g., about 3 to about 4, particularly 4 to 5) layer pairs of anionic polymer layers and cationic polymer layers, where the anionic polymer is PSS and the cationic polymer is PAH.

[0089] In another specific embodiment, the multilayer polymer LbL coating of the substrate comprises about 3 to about 7 (e.g., about 3 to about 5) layer pairs (e.g., 3, 4, or 5) of anionic polymer layers and cationic polymer layers, where the anionic polymer is heparin (HEP) and the cationic polymer is chitosan (CHI), the outer polymer layer pair consists of about 1 to about 2 layer pairs of anionic polymer layers and cationic polymer layers, where the anionic polymer is PSS and the cationic polymer is PAH, and the outer polymer layer pair PSS / PAH is end-functionalized with heparin.

[0090] In a particular embodiment, in a multilayer polymer coating comprising or consisting of a (HEP / CH) layer pair and a multilayer polymer coating comprising or consisting of a (PSS / PAH) layer pair, the thickness increment per layer pair from the multilayer polymer LbL is in the range of 1 to 20 nm.

[0091] According to an advantageous embodiment of the present invention, the thickness of the multi-polymer LbL coating can be adjusted by selection of specific polymer layer pairs.

[0092] In another more particular embodiment, the material according to the invention is selected from the group consisting of: - (optional adhesive)-(Hep / Chi)5-Hep, where the last LbL layer of (Hep / Chi)5 is functionalized with heparin by terminal conjugation; and - (Optional Adhesive) -(PSS / PAH)4-Hep, where the last LbL layer of -(PSS / PAH)4 is functionalized with heparin by terminal conjugation. The present invention relates to a method for producing a biocompatible nonwoven fiber coated with a system comprising a multilayer (LbL) polymer assembly from

[0093] According to a particular embodiment, an adhesive is present.

[0094] According to another particular embodiment, the adhesive is PEI.

[0095] According to another particular embodiment, the adhesive is a PAH.

[0096] Preparation of the material according to the invention According to a particular embodiment, - providing a biocompatible nonwoven textile support; - coating said support with a multi-layered polymeric LbL coating, coating the multi-layer polymer LbL coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer, the anionic polymer being selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or a mixture thereof, and the cationic polymer being selected from chitosan or a mixture thereof, and poly(allylamine hydrochloride) (PAH), or a mixture thereof; - functionalizing the outer layer pair of said multi-polymer layer-by-layer (LbL) coating with heparin by terminal functionalization. A method for preparing a blood compatible material is provided, comprising:

[0097] According to certain embodiments, the implantable material is in the form of a biocompatible non-woven fiber that is provided as a support by electrospinning.

[0098] According to another particular embodiment, biocompatible nonwoven fibers are spun to form a support in the form of a tubular structure forming a lumen, the wall of which forms the luminal surface.

[0099] According to certain embodiments, the nonwoven fiber substrate is pretreated by immersion in an alcohol solution for about 10 minutes, followed by immersion in ultrapure water for about 10 minutes, and then drying prior to coating (e.g., in a solution containing ethanol, such as about 100% to about 95% ethanol).

[0100] According to another particular embodiment, the method of preparing a blood compatible material further comprises an optional pre-coating step of providing a pre-coating deposit of adhesive on the surface of the implantable material prior to the multi-layered polymer coating.

[0101] According to a particular embodiment, an adhesive pre-coating step for a multi-layered polymeric LbL coating is advantageously used when the first polymer layer of the first layer pair is a cationic polymer.

[0102] According to another particular embodiment, a pre-coating step with an adhesive for the multi-layered polymer coating is advantageous when the potting material is PCL.

[0103] According to another particular embodiment, there is provided a method for preparing a blood compatible material according to the present invention, in which a multi-layered polymeric LbL coating is applied to a biocompatible nonwoven fiber (either directly or on a pre-coated biocompatible nonwoven fiber) by layer-by-layer (LbL) deposition.

[0104] According to another particular embodiment, a method for preparing a blood compatible material according to the present invention is provided, wherein heparin is functionalized to a multi-layered polymer coating by terminal functionalization.

[0105] According to another particular embodiment, there is provided a method for preparing a blood compatible material according to the present invention, wherein the multi-layer polymeric LbL coating of a biocompatible nonwoven fiber comprises at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as an outer layer pair and forms a heparinized substrate for end-functionalization of heparin.

[0106] According to another particular embodiment, the heparinized multilayer polymer coating of the present invention has antithrombotic and / or cell proliferation properties and can therefore be used on any substrate intended to come into contact with blood for any in vitro, animal or human application.

[0107] The artificial device of the present invention According to another particular embodiment, the substrate of the material of the invention may be an implantable or non-implantable medical device in contact with blood, in particular a stable or biodegradable implantable prosthetic device in contact with blood, such as those defined herein.

[0108] According to certain embodiments, the implantable medical devices of the present invention have anti-thrombotic properties.

[0109] According to certain embodiments, the implantable medical devices of the present invention have anticoagulant properties.

[0110] According to certain embodiments, the implantable medical devices of the present invention do not require a pre-endothelialization step prior to implantation.

[0111] According to certain embodiments, there is provided an implantable medical device suitable for contact with blood, said device comprising a surface (e.g., a wall) comprising a biocompatible nonwoven fabric, said surface being coated with a multi-layer polymeric coating according to the present invention.

[0112] According to certain embodiments, an implantable medical device suitable for contact with blood is provided, the surface comprising biodegradable, biocompatible nonwoven fibers being formed from electrospun, biocompatible nonwoven fibers.

[0113] In one embodiment, a biodegradable implantable medical device, particularly a vascular graft, is provided having a tubular structure that defines a lumen, with the wall of the device forming the luminal surface.

[0114] According to certain embodiments, an implantable medical device suitable for contact with blood is provided that is a vascular graft.

[0115] In one embodiment, the device is a biodegradable vascular graft having a tubular structure that forms a lumen and the wall of the device forms the luminal surface, the biodegradable vascular graft being of any size, diameter, or configuration (straight or branched), but in particular having an inner diameter of less than 6 mm (ID<6 mm), typically from about 1 to about 6 mm.

[0116] In one specific embodiment, the biodegradable vascular graft is a coronary vascular graft.

[0117] Advantageously, an implantable prosthetic device, such as a biodegradable vascular graft according to the present invention, is a permanent or temporary intracorporeal medical device.

[0118] In one embodiment, the implantable prosthetic device according to the present invention is selected from the group consisting of stents, e.g. bifurcated stents, balloon expandable stents, self-expanding stents, stent grafts, e.g. bifurcated stent grafts, grafts, e.g. vascular grafts, bifurcated grafts, embolic filters, artificial blood vessels, drug delivery devices / balloons, patches, endovascular occlusion devices, CNS shunts (e.g. ventriculopleural shunts, VA shunts, or VP shunts), ventriculoperitoneal shunts, ventriculoatrial shunts, portovenous shunts, and shunts for ascites, heart valve leaflets, shunts for pediatric cardiac surgery, vascular grafts, and shunts for access surgery for dialysis.Furthermore, the biodegradable implantable medical device according to the present invention is also an artificial heart, an extracorporeal cardiac assist device such as a left ventricular assist device (LVAD), or a non-implantable extracorporeal assist device such as cardiopulmonary bypass (CPB), extracorporeal membrane oxygenation (ECMO), or a blood purification device, and their catheters.

[0119] Further, according to a particular embodiment, there is provided an extracorporeal circulation device, such as a cardiopulmonary bypass / ECMO, including an oxygenator, or a hemodialysis / filtration system, or a blood purification system, as well as all catheters and tubes, comprising the material according to the present invention (in particular the coating material 1 or 3 according to the present invention) or the material obtainable according to the method of the present invention.

[0120] Preparation of the Prosthetic Device of the Invention According to a particular embodiment, - providing an implantable device or a part thereof (e.g. a wall of the implantable device) comprising a biocompatible nonwoven fabric as a support; - coating said support with a multi-layered polymeric LbL coating according to the invention A method for preparing an implantable device is provided, comprising:

[0121] According to a particular embodiment, - forming the wall by either electrospinning or extrusion / expansion of biocompatible non-woven fibers; - coating said wall with a multi-layer polymer coating by layer-by-layer (LbL) deposition, the multi-layer polymer coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer layer according to the invention; - functionalizing said multi-layered polymeric LbL coating with heparin by terminal functionalization on its outer layer A method for preparing a blood compatible implantable device is provided, comprising:

[0122] According to certain embodiments, the implantable device or a portion thereof is provided as a support by forming a wall by electrospinning of biocompatible non-woven fibers.

[0123] According to a particular embodiment, a method for preparing the implantable device of the present invention is provided, in which a non-degradable or degradable mesh is introduced into a non-biodegradable or fully biodegradable implantable device (e.g., artificial) that is in contact with blood, in particular a fully biodegradable vascular graft. According to this particular embodiment, when the fully biodegradable vascular graft is implanted in the arterial or venous circulation or any other tubular structure, the non-degradable or degradable mesh will prevent the graft wall from bursting or expanding or kinking in case of insufficient revascularization after complete polymer degradation. This mesh improves the mechanical stability of the graft before, during and after degradation. Thus, the present invention relates to an artificial device comprising a fully biodegradable implantable artificial device as defined in the present invention and a non-degradable or degradable mesh.

[0124] patient In one embodiment, a patient of the present invention is a subject suffering from or at risk of suffering from ischemic, degenerative, or congenital cardiovascular disease, trauma (e.g., neurological, peripheral, cardiac, vascular, or orthopedic trauma), or organ loss or failure, such as end-stage renal disease (ESRD).

[0125] In a further embodiment, a patient according to the present invention is a subject undergoing reconstructive surgery, such as revascularization surgery.

[0126] Use according to the invention According to certain embodiments, the materials, processes, and methods of the present invention are useful for the treatment of cardiovascular disease by surgery or therapy, in particular cardiovascular ischemic disease, e.g., coronary artery disease, peripheral vascular disease, or access surgery for dialysis, or for the application of grafts or shunts in pediatric patients, pediatric cardiovascular malformations, trauma, or repair surgery.

[0127] According to a particular embodiment, the biocompatible material according to the invention may be useful for the preparation of non-degradable or degradable implantable prosthetic devices (e.g. vascular grafts (endoprostheses) useful for the treatment of humans or animals by surgery or therapy, in particular for cardiovascular diseases, in particular cardiovascular ischemic diseases, e.g. coronary artery disease, or peripheral vascular disease, or access surgery for dialysis patients, pediatric cardiovascular malformations, trauma, and repair surgery).

[0128] Some of the problems with the material properties of non-degradable polymers such as ePTFE, which is extremely hydrophobic (contact angle >100°), may be solved by the proposed material of the present invention, since it renders the surface more hydrophilic, thus promoting cell adhesion / growth and less thrombogenicity, and therefore more biocompatible, and may also reduce inflammation / foreign body reaction, which may lead to less late complications such as intimal hyperplasia and calcification.

[0129] An illustrative example of the invention will now be described in more detail and with reference to embodiments shown in the drawings. EXAMPLES

[0130] Example 1: Preparation of vascular grafts using materials of the present invention The material of the present invention was applied onto a substrate useful as an implantable vascular graft.

[0131] All reagents were used as received without any additional purification steps.

[0132] Poly(allylamine hydrochloride) (PAH, ≈50,000 g / mol), poly(sodium 4-styrenesulfonate) (PSS, ≈70,000 g / mol), and dextran sulfate sodium salt (DS, ≥500,000 g / mol) from Leuconostoc ssp. were purchased from Sigma-Aldrich (St. Quentin Fallavier, France), poly(L-lysine) hydrobromide (PLL, ≈15,000–30,000 g / mol) and chitosan (CHI, low molecular weight, 75–85% deacetylated) were purchased from Sigma-Aldrich (Schnelldorf, Germany). Branched poly(ethyleneimine) (PEI, Lupasol WF, ≈25,000 g / mol) was purchased from BASF (Ludwigshafen, Germany). Heparin sodium salt from porcine intestinal mucosa (>180 USP units / mg, reference H4784 from SIGMA) was used. N,N,N-trimethylchitosan (TMC,

[0133]

number

[0134] ≈90,000 g / mol, degree of quaternization 0.25) was obtained by chemical modification of chitosan (Golden-Shell Biochemical Co., China) with methyl iodide (CH3I) as described elsewhere (Verheul et al. 2008 Biomaterials, 29, 3642-3649; Martins et al. 2013 Carbohydr. Res., 381, 153-160).

[0135] A solution of PEI was prepared at a concentration of 2.5 mg / mL in ultrapure water. The polyelectrolyte solutions used for the assembly of LbL films can be classified into three different groups based on the polyanion used: HEP, DS, or PSS.

[0136] For HEP-containing LbL films, polymer solutions were prepared in 0.15 M sodium chloride solution at a concentration of 1.0 mg / mL, except for CHI. Chitosan was dissolved in a solution of acetic acid (0.5 M) and 0.15 M NaCl to a final concentration of 2.0 mg / mL. HEP solution at 1 mg / mL was prepared using acetate buffer (pH 4).

[0137] For DS-based LbL films, an aqueous solution of DS (pH ≈ 6.5) was prepared in 1 M NaCl at a polymer concentration of 1 mg / mL. Chitosan was dissolved in a solution of acetic acid (0.5 M) and 1 M NaCl to a final concentration of 2.0 mg / mL. A PLL solution was prepared at a concentration of 1 mg / mL and 1 M NaCl.

[0138] For PSS-containing LbL films, polymer solutions were prepared in 0.15 M sodium chloride solution at a concentration of 1.0 mg / mL.

[0139] No final adjustment of the pH value was made to any of the polyelectrolyte solutions.

[0140] A mixture DS:HEP 1:10, taking into account the molar ratio (number of chains) of each component per milliliter, was prepared in acetate buffer at 0.15 M NaCl to maintain a final polymer concentration of 1 mg / mL (77.3 mg DS and 23.2 mg HEP dissolved in 100 mL of buffer).

[0141] a) providing a support comprising a biocompatible nonwoven fiber; Biocompatible nonwoven fibers were used as support for the coating of the present invention. For this purpose, ε-polycaprolactone (PCL) micro- and nanofibrous scaffolds were prepared by electrospinning. Briefly, a solution of 15% PCL (80,000 Da, Sigma) in CHCl3 / EtOH (70% v / v) was charged at 20 kV (Spellman High Voltage Electronics Ltd, Pulborough, West Sussex, UK) and pumped through a stainless steel needle at a continuous rate of 12 mL / h by a syringe pump (Fresenius Vial SA, Brezins, France). The polymer fiber was then transferred to a grounded, rotating (4,500 rpm) and translating (200 travel s) tube mounted on a home-built stand. -1 The fibers were harvested by a stainless steel mandrel (2 mm diameter, or less than 2 m, or more than 2 m) with a needle-to-recovery distance of 20 cm. After electrospinning, the grafts were placed in vacuum overnight to remove any residual solvent. The average size and distribution of such PCL fibers was measured in a previous study, which corresponds to an average fiber diameter of 2.2 ± 0.6 μm.

[0142] The resulting fiber meshes (herein referred to as "grafts" and "patches") were used as bioartificial surfaces mimicking vascular grafts used in cardiovascular surgical implants. Brief pre-treatment of the nonwoven fiber supports (e.g., grafts / patches) was performed by immersion in ethanol for 10 min and rinsing with ultrapure water for 10 min and drying. Sterilization prior to implantation was performed using gamma radiation with a minimum dose of 25 kGy and a maximum dose of 32 kGy.

[0143] b) Optionally, coating the substrate with a pre-coating deposit comprising an adhesive. When using a pre-coating deposit of adhesive, the PEI solution is coated onto the non-woven fibrous substrate obtained above by immersion for 15 minutes and rinsing with ultrapure water three times for 2 minutes, followed by drying.

[0144] c) Coating the support (precoated with an adhesive or not) with a multi-layer polymer LbL coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer layer of the present invention. Approximately 3 cm x 6 cm samples of the PCL-patches obtained above were coated by immersing the substrate in the solution. The duration of the polymer deposition and rinsing steps was approximately 1 min in the polyelectrolyte solution, followed by 3 x 8 s rinses in ultrapure water.

[0145] The following materials of the present invention, including LbL film combinations of the present invention (LbL Systems 1-3), and comparative coatings C1-C6, were prepared on biocompatible nonwoven fiber surfaces of various devices, as listed in Table 1 below.

[0146] [Table 1]

[0147] d) Heparinization of the multilayered polymer coating on the final LbL layer A heparinized multi-polymeric LbL coating is obtained by functionalization of heparin on the outer layer of said multi-polymeric coating, independent of the fact that heparin is incorporated as an anionic polymer of said multi-polymeric LbL coating or not.

[0148] Heparinization can be carried out by various known techniques. In certain embodiments, heparin molecules are attached to the surface of the multilayered polymer coating using terminal conjugation (EPC), as described in Hsieh et al., 2017, supra.

[0149] For heparin terminal conjugation, the material obtained in the previous step was incubated for 24 hours in Hep EPC solution (0.02 M monobasic sodium phosphate, 0.2 M sodium chloride, and 3 mg mL -1Immerse in sodium cyanoborohydride (NaC12O4) and agitate using an orbital shaker (200 rpm), then rinse 3 x 2 min with ultrapure water and dry.

[0150] According to certain embodiments, a heparinized substrate favorable for heparin terminal conjugation is formed with at least one outer layer pair comprising poly(allylamine hydrochloride) (PAH) and poly(allylamine hydrochloride) (PSS).

[0151] The amount of heparin deposited on the multilayered polymer coating was determined using the carbazole assay described in US Pat. No. 4,613,665 to be approximately 9.6 μg / cm 2 .

[0152] These results suggest that the multilayered polymer LbL coating presents a sufficient amount of primary amino groups in the bulk direction to react efficiently with heparin (aldehyde-terminated) molecules.

[0153] Given the current clinical use of vascular grafts and patches made of microporous non-degradable polymers such as ePTFE, the coatings of the present invention have also been tested on these materials. To this end, surfaces prepared with or without traditional surface activation procedures (e.g., plasma treatment, etc.) of ePTFE and layer-by-layer coating methods were adjusted as follows compared to those used with PCL described above.

[0154] a) providing a support comprising a biocompatible nonwoven fiber; Prior to deposition, the e-PTFE patches were soaked in EtOH for 1 hour and then in Milli-Q water for 30 minutes.

[0155] b) Optionally, coating the substrate with a pre-coating deposit comprising an adhesive. When adhesive pre-coating deposits were used, the patches were immersed in the PEI or PAH solution for 20 minutes, followed by 3 rinses for 2 minutes with Milli-Q water.

[0156] c) Coating the support (which may or may not be precoated with an adhesive) with a multi-layer polymer LbL coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer layer of the present invention. The ePTFE-patches obtained above were coated by deposition of a (PSS / PAH)4 multilayer film by immersing the patch in the corresponding polyelectrolyte (PSS or PAH) for 1 min, followed by rinsing three times with Milli-Q water for 8 s.

[0157] The following materials of the present invention (LbL Systems 4-5) comprising LbL membrane combinations of the present invention are prepared on biocompatible nonwoven fiber surfaces of various devices, as listed in Table 2 below.

[0158] [Table 2]

[0159] d) Heparinization of the multilayered polymer coating on the final LbL layer A heparinized multi-layer polymer LbL coating is obtained by functionalization of the outer layer of said multi-layer polymer coating by terminal attachment of heparin on the LbL-modified e-PTFE patch obtained above using the EPC procedure described above for the PCL sample. The e-PTFE patch was dried only at the end of the deposition process, not after each deposition step as in the case of the PCL sample.

[0160] Example 2: Characterization of implants coated with materials of the present invention The implants obtained in Example 1 were tested as follows.

[0161] a) Graft stability Blood media is simulated using immersion in 0.15 M NaCl aqueous solution at room temperature or 37° C. for at least 24 hours up to 7 and / or 10 days, taking into account additional events that would include primarily protein adsorption on top of the coating, to evaluate the stability of such membranes under conditions similar to those of the antithrombotic experiments. Total membrane thickness (nm) over time (h) was evaluated by ellipsometry on silicon wafers, and the multilayered polymer coating thickness maintained more than 90% of its initial value, suggesting that membrane integrity was maintained for more than 7 days under these conditions.

[0162] This stability suggests that it may help the host body to undergo the endothelialization process safely from thrombus formation and complications arising therefrom.

[0163] Additionally, they observed that gamma sterilization (25 KGy) and storage (refrigerated for up to 9 months) did not alter the stability results for the vascular grafts.

[0164] b) Antithrombotic properties The antithrombotic performance of the vascular grafts was evaluated in vitro using two complementary methods.

[0165] First, the amount of anticoagulant released from the grafts when incubated with human plasma was evaluated. For this purpose, the grafts were placed in tubes and incubated with 700 μL of pooled human plasma (CRYOcheck, PrecisionBiologic) under static conditions for 24 h, 60 h, 5, 7, and 10 days at 37° C. in triplicate for each condition. Anti-Xa activity in the plasma after each incubation period was measured with a chromogenic anti-Xa assay using an Atellica instrument (Siemens, Germany). Next, the efficacy of each graft to inhibit the formation of thrombin in pooled human plasma (CRYOcheck, PrecisionBiologic) using a Calibrated Automated Thrombogram® (Stago, Asnieres sur Seine, France) was evaluated. For this purpose, the grafts were incubated as described above and then placed in wells with 300 μL of fresh pooled plasma (CRYOcheck). Thrombin generation was then measured with the PPP reagent (Stago, reference 86193) at 37° C. as recommended by the manufacturer.

[0166] The results of the grafts coated with the material of the invention compared to the comparative systems C1-C6 are shown in Figures 4A-B and 4C, respectively. The control patch without heparin shows no anti-Xa increase. The longest anti-Xa effect was observed with PEI-(Hep / Chi)5-Hep terminal (1). The variability of the results of the invention (coating 1) between Figures 4A and 4B is due to different time points and different series of experiments.

[0167] FIG. 4A shows that the coating using the chitosan analog TMC (Inventive Coating 3) had similar anticoagulant properties as the one with chitosan (Inventive Coating 1).

[0168] The (PSS / PAH)-based coatings of the present invention (Inventive Coatings 4 and 5) were also applied to ePTFE patches and analyzed as described above. Figure 5 shows that the (PSS / PAH)-based coatings also had anticoagulant properties on ePTFE, and this property was not dependent on the presence of a PEI sublayer as an adhesive.

[0169] From these data it can be clearly seen that the material according to the invention shows sustained inhibition of thrombin in human plasma up to 5 or 7 days compared to the control where thrombin is normally generated at all time points. Notably, the inhibition of thrombin generation was effective in all samples according to the invention, despite an initial release of anticoagulant (heparin) during the first 24 hours of incubation, as assessed by anti-Xa assay in the supernatant plasma.

[0170] c) Cell adhesion The performance of the vascular grafts in recruiting endothelial and smooth muscle cells was assessed in an in vitro cell proliferation assay described below.

[0171] Cell numbers were semi-quantitatively counted after 5 days of incubation with smooth muscle cells (SMC) and endothelial cells (EC), as shown in Figure 6. SMC proliferation is important in degradable materials to repopulate the vessel wall and trigger ECM formation. Heparin is known to reduce this activity, as shown in Figure 6 for the LbL system 1.

[0172] These results confirm that endothelial cells can rapidly attach to a multilayer polymeric LbL coating when the multilayer polymeric LbL coating comprises a heparinized substrate comprising at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair, and heparin is functionalized by terminal conjugation to the heparinized substrate.

[0173] This finding confirms that rapid and complete endothelialization can be obtained with the material of the present invention, which is particularly useful in the context of biodegradable vascular grafts comprising the material of the present invention whose surface is in contact with blood, which can induce in vivo conversion to neovessels.

[0174] Taken together, these data support, but are not limited to, the ability of the materials of the present invention to inhibit clotting and thrombus formation following implantation of biocompatible devices in contact with blood, particularly vascular grafts.

[0175] Moreover, the method of preparation of the material of the present invention advantageously provides a way to obtain biocompatible nonwoven fibers that are uniformly coated with a multi-layered polymer coating by LbL techniques, the uniformity being not only demonstrated by optical and fluorescent microscopy, but also primarily supported by the fact that there is a strong anticoagulant reaction that confirms that essentially all of the fibers are coated, thereby preventing clotting events at the small defects.

[0176] LbL coating of the material (2 of the present invention) furthermore allowed the adhesion and proliferation of cells, particularly endothelial cells (Figure 6). Finally, end-functionalization, particularly the conjugation of heparin, considerably improved the performance of the device (patch) of the present invention in terms of delaying prethrombin formation and, consequently, thrombus formation.

Claims

**Claim 1** - A support comprising biocompatible non-woven fibers; - A heparinized multilayer polymer layer-by-layer (LbL) coating of the biocompatible non-woven fibers, the multilayer polymer coating comprising at least one layer pair of an anionic polymer layer and a cationic polymer layer, the multilayer polymer layer-by-layer (LbL) coating being functionalized with heparin on the outer layer pair of the multilayer polymer coating by terminal functionalization, the anionic polymer being selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or mixtures thereof, the cationic polymer being selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or mixtures thereof, the analog of chitosan being N-alkylated chitosan, particularly trimethyl chitosan, a heparinized multilayer polymer layer-by-layer (LbL) coating A blood-compatible material comprising. **Claim 2** The biocompatible non-woven fibers are biodegradable, the material according to claim 1. **Claim 3** The biocompatible non-woven fibers are composed of a biodegradable polymer selected from the group consisting of degradable polycaprolactone (PCL), polyglycolic acid (PGA), polylactic acid (PLA), poly(lactic-co-glycolic acid) (PLGA), poly-L-lactide (PLLA), poly(glycerol sebacate) (PGS), polydioxanone (PDO, PDS), or poly-p-dioxanone, or a PGA-poly-4-hydroxybutyrate (P4HB) copolymer, degradable polyurethane fibers, and combinations thereof, the material according to claim 1 or 2. **Claim 4** The biodegradable biocompatible non-woven fibers comprise polycaprolactone (PCL) fibers, particularly electrospun ε-PCL nano / microfibers, or consist only thereof, the material according to any one of claims 1 to 3. **Claim 5** The biocompatible non-woven fibers are non-biodegradable, the material according to claim 1. **Claim 6** The biocompatible non-woven fibers are composed of a non-biodegradable polymer selected from polyethylene terephthalate (PET), Dacron®, and expanded polytetrafluoroethylene (ePTFE) or polyurethane (PU), the material according to claim 1 or 5. **Claim 7** The biocompatible non-woven fiber has a diameter of about 0.1 to 10 μm, particularly 0.5 to about 5 μm, for example about 1 to about 5 μm, and is the material according to any one of claims 1 to 6.

8. The anionic polymer is heparin, and is the material according to any one of claims 1 to 7.

9. The cationic polymer is chitosan or an analog thereof, and the analog of chitosan is N-alkylated chitosan, particularly trimethyl chitosan, and is the material according to any one of claims 1 to 7.

10. The material further includes an endothelial cell-attracting coating of the multilayer polymer layer-by-layer (LbL) coating, and heparin is functionalized on the endothelial cell-attracting coating by terminal functionalization, particularly terminal conjugation, and is the material according to any one of claims 1 to 9.

11. The biocompatible non-woven fiber is first coated with a precoating deposit containing an adhesive for the multilayer polymer coating, particularly a cationic polymer such as poly(ethyleneimine) (PEI) or poly(allylamine hydrochloride) (PAH), and is the material according to any one of claims 1 to 10.

12. The multilayer polymer LbL coating includes 3 to 8 layer pairs, particularly 3 to 5 layer pairs, of the anionic polymer layer and the cationic polymer layer defined in any one of claims 1 to 11, and the layer pairs are arranged such that the anionic polymer and the cationic polymer layers alternate in the multilayer polymer coating, and is the material according to any one of claims 1 to 11.

13. The multilayer polymer LbL coating includes 3 to 8, particularly 3 to 5, alternating anionic / cationic polymer layer pairs, and the polymer layer pairs are selected from heparin / chitosan or an analog layer pair, and a PAH / PSS layer pair, or a combination thereof, and is the material according to any one of claims 1 to 12.

14. The multilayer polymer coating of the biocompatible non-woven fiber includes at least one pair of alternating layers of poly(allylamine hydrochloride) (PAH) and poly(sodium 4-styrenesulfonate) (PSS) as the outer layer pair, forming a substrate for the terminal functionalization of heparin, and is the material according to any one of claims 1 to 13.

15. The following group: - (Optional adhesive)-(Hep / Chi) 5 -Hep (wherein the last LbL layer of (Hep / Chi) 5 is functionalized with heparin by terminal conjugation); and - (Optional adhesive)-(PSS / PAH) 4 -Hep (wherein the last LbL layer of (PSS / PAH) 4 is functionalized with heparin by terminal conjugation) The material according to any one of claims 1 to 14, comprising a biocompatible nonwoven fiber coated with a multilayer LbL polymer assembly from

16. - A step of preparing a biocompatible nonwoven fiber as a support; - Optionally, a step of coating the support with a precoating deposit containing an adhesive for a multilayer polymer LbL coating; - A step of coating the support, which is precoated or not precoated, with a multilayer polymer LbL coating, wherein the multilayer polymer LbL coating comprises at least one layer pair of an anionic polymer layer and a cationic polymer layer, the anionic polymer is selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or a mixture thereof, the cationic polymer is selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or a mixture thereof, and the analog of chitosan is N-alkylated chitosan, particularly trimethylchitosan, the step of coating; - A step of functionalizing the outer layer pair of the multilayer polymer layer-by-layer (LbL) coating with heparin by end-functionalization A method for preparing a blood-compatible material.

17. An implantable device comprising the material according to any one of claims 1 to 14 or the material obtained according to the method according to claim 16, such as an artificial blood vessel.

18. The device is selected from a stent, a stent graft, an internal graft, a vascular graft, an embolization filter, an artificial blood vessel, a drug delivery device / balloon, a patch, an intravascular occlusion device, a CNS shunt, a ventriculoperitoneal shunt, a ventriculoatrial shunt, a portal systemic shunt and a shunt for ascites, a heart valve and a heart leaflet, a shunt for pediatric cardiac surgery, a vascular graft and a shunt for access surgery for artificial dialysis, an artificial heart, and an LVAD, the implantable device according to claim 17.

19. A cardiopulmonary bypass / ECMO including an oxygenator, or a hemodialysis / filtration system, or a blood purification system, and all extracorporeal circulation devices such as all catheters and tubes, comprising the material according to any one of claims 1 to 14, or the material obtained according to the method according to claim 16.

20. - providing an implantable device or a part thereof (e.g., the wall of the implantable device) comprising a biocompatible non-woven fiber as a support; - coating the support with a heparinized multilayer polymer layer-by-layer (LbL) of the biocompatible non-woven fiber, wherein the multilayer polymer coating comprises at least one layer pair of an anionic polymer layer and a cationic polymer layer, and the multilayer polymer layer-by-layer (LbL) coating is functionalized with heparin on the outer layer pair of the multilayer polymer coating by end-functionalization, the anionic polymer is selected from heparin, and poly(sodium 4-styrenesulfonate) (PSS), or a mixture thereof, the cationic polymer is selected from chitosan or an analog thereof, and poly(allylamine hydrochloride) (PAH), or a mixture thereof, and the analog of chitosan is N-alkylated chitosan, particularly trimethylchitosan, the coating step; A method for preparing an implantable device, comprising.

21. Use of the material according to any one of claims 1 to 15, or use of the material obtained by the method according to claim 16 or 20, for the preparation of an implantable device or an extracorporeal circulation device in contact with blood.

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