In-vivo indwelling object

Abstract

The present invention provides an in-vivo indwelling object that makes it possible to effectively suppress the onset of inflammation which is derived from a porous structure composed of a biodegradable polymer. An in-vivo indwelling object 100 comprises: a cylindrical stent 10 that is expandable; a porous structure 20 that is disposed so as to cover at least a part of the stent 10; and an antithrombotic coating layer 30 that coats at least a part of the surface of the porous structure 20, wherein the porous structure 20 is composed of a first biodegradable polymer, the antithrombotic coating layer 30 is composed of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the degradation time of the porous structure 20 is shorter than the function maintenance time of the antithrombotic coating layer 30.

Description

Implant for in vivo placement

[0001] The present invention relates to an implant for in vivo placement.

[0002] A stent is a medical device that is delivered to a lesion site within a biological lumen by a stent delivery system and then implanted therein to treat various diseases caused by stenosis or occlusion of biological lumens such as blood vessels, and expands the lesion sites such as stenotic or occluded portions to secure the lumen. A common stent has linear struts that form the outer periphery of a cylindrical shape with gaps formed therein. After being delivered to the lesion site, when the stent is expanded (dilated) within the biological lumen, it imparts tensile strength (expansion force) to the lesion site. By maintaining the tensile strength over a predetermined period of time, the stent holds the lesion site in an expanded state and secures the lumen of the biological lumen.

[0003] For example, International Publication No. 2006 / 126182 discloses a mesh-covered stent in which a mesh is disposed to cover the outer periphery of the stent, in order to prevent peripheral embolism during stent implantation. In the mesh-covered stent, the mesh having an expandable knitted structure expands following the expansion of the stent, thereby preventing scattering of plaques and thrombi when the stent is expanded.

[0004] In mesh-covered stents, a polymer mesh is sometimes adopted from the viewpoints of improving the followability of the mesh when the stent expands, improving deliverability to the lesion site, and reducing the influence that may be exerted on the biological lumen after treatment. In general, polymer materials are more likely to cause inflammation in tissues than metal materials, so it is preferable that the polymer material decomposes and disappears after fulfilling its role of suppressing scattering of thrombi and plaques.

[0005] Furthermore, in the case of a mesh-covered stent, the surface area increased by the mesh causes thrombi and plaques to adhere to the stent surface after stent implantation, which may cause poor prognosis in stent implantation procedures. In order to prevent thrombi from adhering to medical devices such as stents, a technique of applying an antithrombotic coating to the surface of a polymer mesh is known.

[0006] On the other hand, if the antithrombotic function of the coating described above is lost faster than the degradation time of the polymer mesh, the polymer mesh surface becomes more susceptible to thrombosis, potentially increasing the risk of thrombosis and other related conditions.

[0007] The present invention has been made in view of the above-mentioned problems, and aims to provide an in vivo implant that can effectively suppress the onset of inflammation originating from a porous structure made of a biodegradable polymer.

[0008] The inventors of this invention conducted intensive research to solve the above problems. As a result, they discovered that the above objectives could be achieved by using an in vivo implant in which the degradation time of the porous structure and the maintenance time of the antithrombotic coating layer were adjusted, and thus completed the present invention.

[0009] One aspect of the present invention is an in vivo implant comprising: 1. an expandable tubular stent; a porous structure disposed to cover at least a portion of the stent; and an antithrombotic coating layer coating at least a portion of the surface of the porous structure, wherein the porous structure is made of a first biodegradable polymer, the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer.

[0010] 2. In the in-vivo implant described in 1. above, it is preferable that the first biodegradable polymer is at least one selected from the group consisting of polyglycolic acid, polyglycolic acid-polylactic acid copolymer, polyglycolic acid-polycaprolactone copolymer, and polylactic acid-polycaprolactone copolymer; 3. In the in-vivo implant described in 1. or 2. above, it is preferable that the number average molecular weight (Mn) of the first biodegradable polymer is 10,000 to 150,000; 4. In the in-vivo implant described in any of 1. to 3. above, it is preferable that the antithrombotic polymer is a hydrophilic polymer and the second biodegradable polymer is a hydrophobic polymer; 5. In the in-vivo implant described in 1. to 4. above. In the in vivo implant described in any of the above, it is preferable that the monomer constituting the antithrombotic polymer is one or more selected from the group consisting of ethylene glycol, methoxyethylene glycol, and propylene glycol, and the monomer constituting the second biodegradable polymer is one or more selected from the group consisting of lactic acid, caprolactone, and glycolic acid; 6. In the in vivo implant described in any of the above 1 to 4, it is preferable that the antithrombotic polymer is one or more selected from the group consisting of polyethylene glycol, methoxypolyethylene glycol, and polypropylene glycol, and the second biodegradable polymer is one or more selected from the group consisting of polylactic acid, polycaprolactone, and polyglycolic acid; 7. In the in vivo implant described in any of the above 1 to 6, it is preferable that the block molecular weight of the antithrombotic polymer in the block copolymer is 1000 to 10000, and the block molecular weight of the second biodegradable polymer is 1000 to 10000; 8. In the in-vivo implant described in any of the above, it is preferable that the ratio (molar ratio) of the antithrombotic polymer to the second biodegradable polymer in the block copolymer constituting the antithrombotic coating layer is 40:60 to 80:20; 9. In the in-vivo implant described in any of the above 1 to 8, it is preferable that the degradation time of the porous structure is 3.6 months or less;10. In the in-vivo implant described in any of 1 to 9 above, it is preferable that the functional maintenance time of the antithrombotic coating layer exceeds 3.6 months; 11. In the in-vivo implant described in any of 1 to 10 above, it is preferable that the first biodegradable polymer is polyglycolic acid, the antithrombotic polymer is polyethylene glycol, and the second biodegradable polymer is polylactic acid.

[0011] This is a schematic plan view showing a stent delivery system equipped with an in-vivo implant according to an embodiment. This is a schematic plan view showing the reduced diameter state of the in-vivo implant according to an embodiment. This is a schematic plan view showing the expanded (widened diameter) state of the in-vivo implant according to an embodiment. This is a partially enlarged view showing an enlarged portion of the in-vivo implant in the state shown in Figure 3. This is a schematic cross-sectional view showing an example of the arrangement of the stent, porous structure, and antithrombotic coating layer. This is a diagram showing the results of Reference Examples 1 to 4. This is a diagram showing the results of Reference Example 5.

[0012] An in-vivo implant according to one aspect of the present invention is an in-vivo implant comprising an expandable tubular stent, a porous structure disposed to cover at least a portion of the stent, and an antithrombotic coating layer coating at least a portion of the surface of the porous structure, wherein the porous structure is made of a first biodegradable polymer, the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer. The in-vivo implant may also be referred to as a biological implant.

[0013] The inventors have found that an in vivo implant having such a configuration can effectively suppress the occurrence of inflammation originating from a porous structure made of a biodegradable polymer. The mechanism by which the above effects are exerted by the configuration of the present invention is presumed to be as follows.

[0014] The in-vivo implant according to the present invention comprises an expandable tubular stent, a porous structure disposed to cover at least a portion of the stent, and an antithrombotic coating layer coating at least a portion of the surface of the porous structure. The porous structure is made of a first biodegradable polymer, and the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and is characterized in that the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer.

[0015] Generally, porous structures made of biodegradable polymers are more susceptible to thrombosis and other adhesion compared to porous structures made of materials such as metals, posing a higher risk of thrombosis. To reduce this risk, it is conceivable to cover the surface of the porous structure with an antithrombotic coating layer. However, for example, if the function of the antithrombotic coating layer is lost faster than the decomposition time of the porous structure after the implant is placed in a blood vessel, the surface of the porous structure becomes more prone to thrombosis, potentially leading to thrombosis. On the other hand, in the implant according to the present invention, the decomposition time of the porous structure is shorter than the time for which the function of the antithrombotic coating layer is maintained, so the function of the antithrombotic coating layer is maintained at least until the porous structure is decomposed. This is presumed to prevent thrombosis and other adhesion to the porous structure, thereby reducing the risk of thrombosis. Furthermore, it is thought that by the time the porous structure is decomposed, the antithrombotic coating layer has been incorporated into the vascular endothelium. Therefore, the risk of the antithrombotic coating layer leaking into the blood vessels after the degradation of the porous structure is low, and it is presumed that it is unlikely to cause peripheral vascular occlusion, for example.

[0016] Throughout this specification, singular expressions should be understood to include the concept of their plural form unless otherwise specified. Therefore, singular articles (e.g., "a," "an," "the" in English) should be understood to include the concept of their plural form unless otherwise specified. Furthermore, terms used in this specification should be understood to have the meaning commonly used in the art unless otherwise specified. Accordingly, unless otherwise defined, all technical and scientific terms used in this specification have the same meaning as commonly understood by those skilled in the art to which this invention belongs. In case of any conflict, this specification (including definitions) shall prevail. This invention is not limited to the embodiments described below and can be modified in various ways within the scope of the claims. Also, in this specification, "X to Y" indicates a range including X and Y, meaning "X or greater and Y or less." "X and / or Y" means at least one of X and Y, encompassing X alone, Y alone, and combinations of X and Y. Furthermore, "%" of concentration refers to mass concentration "mass%" unless otherwise specified.

[0017] Unless otherwise specified, measurements of operation and physical properties shall be taken under room temperature (20-25°C) / relative humidity of 40-50% RH.

[0018] Embodiments of the present invention will be described below with reference to the attached drawings. Note that the following description does not limit the technical scope or the meaning of terms described in the claims. Furthermore, the dimensional ratios in the drawings are exaggerated for illustrative purposes and may differ from actual ratios.

[0019] (Embodiment) The in-vivo implantation 100 according to this embodiment will be described below with reference to Figures 1 to 5.

[0020] Figure 1 is a schematic plan view showing a stent delivery system 300 equipped with an in-vivo implant 100 according to an embodiment of the present invention. Figure 2 is a schematic plan view showing the in-vivo implant 100 in a reduced diameter state according to the embodiment. Figure 3 is a schematic plan view showing the in-vivo implant 100 in an expanded diameter state according to the embodiment. Figure 4 is a partially enlarged view showing the in-vivo implant 100 according to the embodiment. Figure 5 is a schematic cross-sectional view showing an example of the arrangement of the stent 10, the porous structure 20, and the antithrombotic coating layer 30.

[0021] In this specification, the direction in which the in-vivo implant 100 extends is referred to as the "axial direction." The axial direction is the direction from the tip 10A to the proximal end 10B as shown in Figure 2. The side inserted into the body is referred to as the "tip side," and the side opposite the tip side, where the surgeon operates the stent delivery system 300, is referred to as the "proximal end side." The direction perpendicular to the "axial direction" is referred to as the "radial direction" of the in-vivo implant 100.

[0022] As shown in Figure 1, the stent delivery system 300 includes an in-vivo implant 100 and a balloon catheter 200.

[0023] <Balloon Catheter 200> The balloon catheter 200 is used to deliver the catheter to the lesion site in a deflated state and to expand and implant the in vivo implant 100 near the lesion site.

[0024] The balloon catheter 200 comprises a long catheter body 210, a balloon 220 provided at the tip of the catheter body 210, and a hub 230 fixed to the proximal end of the catheter body 210.

[0025] The catheter body 210 comprises an outer tube and an inner tube positioned inside the outer tube.

[0026] An expansion lumen is formed inside the outer tube through which an expansion fluid for expanding the balloon 220 flows. The tip of the outer tube is fixed to the base end of the balloon 220. The base end of the outer tube is fixed to the hub 230.

[0027] A guidewire lumen is formed inside the inner tube into which a guidewire is inserted. The tip of the inner tube penetrates the inside of the balloon 220 and opens on the tip side of the balloon 220. The base end of the inner tube penetrates the side wall of the outer tube on the base side of the balloon 220 and is fixed to the outer tube.

[0028] The constituent material of the catheter body portion 210 is preferably a material having a certain degree of flexibility. Examples include polyolefins such as polyethylene, polypropylene, polybutene, ethylene-propylene copolymer, ethylene-vinyl acetate copolymer, ionomer, or mixtures of two or more of these; thermoplastic resins such as polyvinyl chloride resin, polyamide, polyamide elastomer, polyester, polyester elastomer, polyurethane, and fluororesin; silicone rubber; and latex rubber.

[0029] The balloon 220 is a component that expands inside a narrowed area (lesion) to widen it. The tip of the balloon 220 is fixed to the outer wall surface of the inner tube. The base of the balloon 220 is fixed to the outer wall surface of the tip of the outer tube. Therefore, the inside of the balloon 220 communicates with the expansion lumen formed in the outer tube. The balloon 220 allows expansion fluid to flow in from the base opening 231 via the expansion lumen. The balloon 220 expands when the expansion fluid flows in and contracts and folds up when the fluid that has flowed in is discharged.

[0030] The constituent material of the balloon 220 is preferably a flexible material that expands and contracts with the inflow and outflow of the expansion fluid. Examples include polymer materials such as polyolefins, crosslinked polyolefins, polyesters, polyester elastomers, polyvinyl chloride, polyurethanes, polyurethane elastomers, polyphenylene sulfide, polyamides, polyamide elastomers, and fluororesins, as well as silicone rubber and latex rubber. The constituent material of the balloon 220 is not limited to the use of the above polymer materials alone; a film in which the above polymer materials are appropriately laminated may also be applied. The expansion fluid may be a gas or a liquid, for example, helium gas or CO2. ２ Gas, O２ Examples include gases such as gases, and liquids such as physiological saline and X-ray contrast agents.

[0031] The hub 230 is provided with a base opening 231 that communicates with the expansion lumen of the outer tube. The base opening 231 functions as a port for the inflow and outflow of the expansion fluid.

[0032] The constituent materials of the hub 230 are not particularly limited, but examples include thermoplastic resins such as polyethylene, polyurethane, polyester, polypropylene, polycarbonate, polyamide, polysulfone, polyarylate, and methacrylate-butylene-styrene copolymer.

[0033] <Intra-vivo implant 100> The intra-vivo implant 100 includes an expandable tubular stent 10, a porous structure 20 positioned to cover at least a portion of the stent, and an antithrombotic coating layer 30 that coats at least a portion of the surface of the porous structure.

[0034] In this embodiment, the in-vivo implant 100 is used, for example, to treat narrowing or obstruction in blood vessels, bile ducts, trachea, esophagus, urethra, or other biological tubular lumens. The in-vivo implant 100 is a so-called balloon-expandable type, attached to a balloon 220 in a crimped state, delivered to the lesion, and then expanded and implanted at the lesion.

[0035] The stent 10 has a substantially cylindrical shape with a lumen that extends in the axial direction. The tip 10A is located at the acupoint end of the stent 10, and the proximal end 10B is located at the proximal end of the stent 10.

[0036] As shown in Figures 2 and 3, the stent 10 is formed to allow for radial expansion (as shown in Figure 3) and contraction (as shown in Figure 2). As shown in Figures 3 and 4, the stent 10 has a linear ring 11 that forms the outer circumference of a cylindrical shape with a gap, and a link portion 12 that connects the rings 11 at the gap.

[0037] As shown in Figure 4, the wavy ring 11 has multiple first strut portions 15 made of straight or curved lines, second strut portions 16 made of straight or curved lines, and curved portions 17 formed between the first strut portions 15 and the second strut portions 16. The ring 11 also has multiple third strut portions 14 made of straight or curved lines, which are adjacent to the link portion 12 on one axial side (for example, the left side in Figure 4) and are arranged in pairs in the circumferential direction (up and down in Figure 4).

[0038] The rings 11 are arranged sequentially along the axial direction, and adjacent rings 11 in the axial direction are integrated by the link portion 12. Therefore, by increasing or decreasing the number of rings 11, it is possible to easily obtain a stent 10 of a desired length.

[0039] As shown in Figures 2 to 5, the porous structure 20, which has numerous through-voids, is positioned to cover the outer circumference (outer surface 19) of the stent 10, and an antithrombotic coating layer 30 is provided on at least a portion of its surface.

[0040] As shown in Figures 2 and 3, the porous structure 20 is configured to expand in accordance with the expansion of the stent 10. By positioning the porous structure 20 on the outer circumference of the stent 10, when the stent 10 expands within the blood vessel, the porous structure 20 is biased against the blood vessel by the stent 10, thereby reducing the risk of the porous structure 20 rupturing due to the pressure of the blood flow within the blood vessel.

[0041] The porous structure 20 extends along the axial direction of the stent 10 and has a mesh structure. The gaps in the mesh structure of the porous structure 20 can be configured to be smaller than the gaps between adjacent rings 11. By making the mesh of the porous structure 20 smaller than the gaps between the rings 11 in this way, it is possible to prevent plaque and thrombi from scattering when the stent 10 is expanded at the lesion site. This prevents peripheral embolism caused by plaque and thrombi.

[0042] The porous structure 20 includes, for example, any one of knitted fabrics (knits), woven fabrics (braids), and molded articles. When the porous structure 20 is a knitted fabric, the porous structure 20 can be constituted by, for example, plain knitting. By configuring the porous structure 20 by plain knitting, shortening of the axial direction of the porous structure 20 accompanying expansion of the stent 10 can be suppressed. Note that plain knitting itself is a known knitting method, so a detailed description thereof is omitted in this specification. Further, when the porous structure 20 is a woven fabric, the woven fabric is constituted by a known weaving method. Further, when the porous structure 20 is a molded article, the porous structure 20 is constituted by a cylindrical sheet provided with regularly arranged voids. The size, shape, and number of stitches and weaves of the porous structure 20 are not particularly limited as long as peripheral embolism during expansion of the stent 10 can be prevented.

[0043] The thread diameter of the porous structure 20 is not particularly limited, and may be, for example, 20 μm.

[0044] Further, the size, shape, number, and the like of the stitches of the porous structure 20 are not particularly limited as long as peripheral embolism caused by plaques or thrombi when the stent 10 expands can be prevented. Further, the porous structure 20 can be fixed to any position such as the distal end 10A, the proximal end 10B, and the link portion 12 of the stent 10, for example. There is no particular limitation on the method for fixing the porous structure 20 to the stent 10, and any method can be used, for example, adhesive fixation with a polymer or paraffin, fixation using mechanical means such as a binding band, a clip, a thread, or the like.

[0045] Further, as described above, at least a part of the surface of the porous structure 20 is coated with an antithrombogenic coating layer 30. In one embodiment shown in FIGS. 3, 4 and 5, the entire surface of the porous structure 20 is coated with the antithrombogenic coating layer 30.

[0046] As shown in FIG. 5, in one embodiment, the antithrombotic coating layer 30 is configured to form a layer that entirely covers the outer circumferential surface of the framework of the porous structure 20 in the circumferential direction. Further, in the embodiment shown in FIG. 5, the antithrombotic coating layer 30 is not formed in the gaps of the mesh structure of the coated porous structure. However, in other embodiments, the antithrombotic coating layer 30 may circumferentially cover a part of the outer surface of the framework of the porous structure 20, or may be formed in the gaps of the mesh structure of the porous structure. Note that the rotation direction based on the axial direction is defined as the circumferential direction, and is indicated by arrows R1-R2 in the drawing.

[0047] Hereinafter, a preferred example of materials, physical properties and the like of each part (stent, porous structure, antithrombotic coating layer, etc.) constituting the in-vivo indwelling device 100 will be described.

[0048] <In-vivo Indwelling Device> The in-vivo indwelling device according to the present invention includes an expandable cylindrical stent, a porous structure disposed to cover at least a part of the stent, and an antithrombotic coating layer that coats at least a part of the surface of the porous structure. Further, the porous structure is made of a first biodegradable polymer, the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the decomposition time of the porous structure is shorter than the function maintenance time of the antithrombotic coating layer, which is a feature of the present invention.

[0049] [Stent] The in-vivo indwelling device according to the present invention includes an expandable cylindrical stent. The stent may be made of any material, and may be a non-biodegradable material or a biodegradable material.

[0050] <Non-biodegradable Material> As non-biodegradable materials usable for stents, carbon fibers, metallic materials and the like can be used. From the viewpoint of further enhancing the effect of reducing inflammation, it is particularly preferable that the non-biodegradable material is a metallic material.

[0051] Here, the metal material used when the stent is composed of a metal material is not particularly limited, and metal materials commonly used in the field of stents can be used. Specifically, examples include stainless steels such as SUS304, SUS316, SUS316L, SUS420J2, and SUS630, tantalum, titanium, nickel-titanium alloys, tantalum-titanium alloys, nickel-aluminum alloys, Inconel, gold, platinum, iridium, tungsten, and cobalt-chromium (Co-Cr) alloys. Among stainless steels, SUS316L is preferred because it has the best corrosion resistance. Among cobalt-based alloys, MP35N and L605 are preferred.

[0052] <Biodegradable Materials> There are no particular restrictions on the biodegradable materials that can be used for stents; generally, biodegradable resin materials commonly used in the medical field can be used.

[0053] The biodegradable resin material is not particularly limited, and known biodegradable resin materials such as those described in Japanese Patent Publication No. 2011-528275, Japanese Patent Publication No. 2008-514719, International Publication No. 2008 / 1952, Japanese Patent Publication No. 2004-509205, etc., can be used. Specifically, examples include (1) polymers selected from the group consisting of aliphatic polyesters, polyesters, polyacid anhydrides, polyorthoesters, polycarbonates, polyphosphazenes, polyphosphate esters, polyvinyl alcohols, polypeptides, polysaccharides, proteins, and cellulose; and (2) copolymers composed of two or more monomers constituting the above (1). Here, the aliphatic polyester is not particularly limited and includes, for example, polylactic acid (PLA) such as poly-L-lactic acid (PLLA), poly-D-lactic acid (PDLA), poly-DL-lactic acid (PDLLA), polyglycolic acid (PGA), polyhydroxybutyric acid, polyhydroxyvaleric acid, polyhydroxypentanoic acid, polyhydroxyhexanoic acid, polyhydroxyheptanoic acid, poly(ε-caprolactone) (PCL, also simply called polycaprolactone), polytrimethylene carbonate, poly-2,2-dimethyltrimethylene carbonate, polydioxanone, polybutyrolactone, polyvalerolactone, polymalic acid, polyethylene adipate, polyethylene succinate, polybutylene adipate, and polybutylene succinate. Furthermore, the polycarbonate is not particularly limited and includes, for example, tyrosine-polycarbonate.

[0054] Alternatively, the biodegradable resin material may be a copolymer formed by the arbitrary copolymerization of monomers constituting the polymer. Here, the copolymer is not particularly limited. Specifically, PGA-PLA (polyglycolic acid-polylactic acid copolymer), PGA-PLLA (polyglycolic acid-poly-L-lactic acid copolymer), PGA-PDLA (polyglycolic acid-poly-D-lactic acid copolymer), PGA-PDLLA (polyglycolic acid-poly-DL-lactic acid copolymer), PGA-PCL (polyglycolic acid-polycaprolactone copolymer), PLA-PCL (polylactic acid-polycaprolactone copolymer), PLLA-PCL (poly L-lactic acid-polycaprolactone copolymer), PDLA-PCL (poly-D-lactic acid-polycaprolactone copolymer), PDLLA-PCL (poly-DL-lactic acid-polycaprolactone copolymer), PLLA-PDLLA (poly-L-lactic acid-poly-DL-lactic acid copolymer), PLLA-PDLA (poly-L-lactic acid-poly-D-lactic acid copolymer), PDLA-PDLLA (poly-D-lactic acid-poly-DL-lactic acid copolymer), PLLA-r-PTMC (TMC = trimethylene PAE (Polyanhydride) (poly(lactide-co-glycolide), polyanhydride, polyorthoester, poly(N-(2-hydroxypropyl)methacrylamide), PGA-TMC (poly(glycolide-co-trimethylene carbonate)), and PDO-PGA-TMC (poly(glycolide-co-trimethylene carbonate-co-dioxanone)), PAE (Polyanhydride) in which salicylic acid is chemically introduced into the polymer backbone. Examples include esters)-Salicylate (for example, polymers in which salicylic acid is bonded to both ends of polylactide anhydride or polyadipic acid).

[0055] The polymers and copolymers described above may be used individually, in combination of two or more types, or in combination of one or more polymers and one or more copolymers. Furthermore, the polymers and copolymers may be produced by synthesis or commercially available products. The synthesis method is not particularly limited, and known methods can be applied in the same manner or with appropriate modifications. For example, polylactic acid (PLA), polyglycolic acid (PGA), or lactic acid-glycolic acid copolymer (PLGA) can be obtained by selecting the desired structure from L-lactic acid, D-lactic acid, and glycolic acid as raw materials and performing dehydration polycondensation. Alternatively, they can be obtained by selecting the desired structure from lactide, a cyclic dimer of lactic acid, and glycolide, a cyclic dimer of glycolic acid, and performing ring-opening polymerization. Lactides include L-lactide, a cyclic dimer of L-lactic acid; D-lactide, a cyclic dimer of D-lactic acid; meso-lactide, a cyclic dimer of D-lactic acid and L-lactic acid; and DL-lactide, a racemic mixture of D-lactide and L-lactide. Any of these lactides can be used in this disclosure.

[0056] Of these, the biodegradable resin material is preferably a bulk erosion type polymer. More preferably, the biodegradable resin material is selected from the group consisting of a single monomer homopolymer or a copolymer of two or more monomers selected from the group consisting of lactic acid, caprolactone, glycolic acid, dioxanone, butyrolactone, valerolactone, hydroxybutyric acid, and trimethylene carbonate.

[0057] The materials constituting the stent may be used individually, or they may be used in combination as a mixture of two or more materials or as a copolymer of two or more monomers constituting any of the above-mentioned resins.

[0058] In this specification, "biodegradable" is not particularly limited, but refers to materials that satisfy at least one of the following standards: ISO 9408, ISO 9439, ISO 10707, ISO 14855-1, ISO 14855-2, ISO 14851, ISO 14852, ISO 17556, JIS K 6950:2000, JIS K 6951:2000, JIS K 6953-1:2011, JIS K 6953-2:2010, and JIS K 6955:2017. "Non-biodegradable" refers to materials that do not satisfy any of these standards.

[0059] Stents can be suitably formed from materials appropriately selected from the non-biodegradable and biodegradable materials exemplified above, depending on the application site (placement site in the body), etc. For example, if a stent is formed from a metal material, the metal material has excellent strength, making it possible to leave the implant in the body in the lesion site for a desired period of time while maintaining the desired tensile strength. On the other hand, if a stent is formed from a resin material, the polymer material has excellent flexibility, exhibiting excellent effect in the delivery of the implant to the lesion site. Furthermore, if the resin material is biodegradable, the stent will disappear by biodegradation after a predetermined period of time, thus reducing the potential impact on the body lumen after treatment.

[0060] Furthermore, the stent may be a self-expanding stent or a balloon-expanding stent. When the stent is constructed as a self-expanding stent, it is preferable to use a superelastic alloy such as nickel-titanium alloy because it is necessary to restore it to its original shape. Also, when the implanted device is balloon-expanding, it is preferable to use a cobalt-based alloy such as cobalt-chromium (Co-Cr) alloy or stainless steel because it is less likely to return to its original shape after expansion. In addition, when the stent is made of carbon fiber, it exhibits excellent effects in that it is high in strength, has excellent flexibility, and is highly safe in vivo.

[0061] [Porous Structure] The in-vivo implant according to the present invention comprises a porous structure disposed to cover at least a portion of the expandable cylindrical stent described above. The porous structure is made of a first biodegradable polymer.

[0062] <First Biodegradable Polymer> The first biodegradable polymer is not particularly limited as long as it is a biodegradable polymer, and can be appropriately selected from the biodegradable materials described in the [Stent] section above. In particular, the first biodegradable polymer is preferably an aliphatic polyester. Furthermore, from the viewpoint of suitability for medical use, the first biodegradable polymer is more preferably at least one selected from the group consisting of polyglycolic acid, polyglycolic acid-polylactic acid copolymer, polyglycolic acid-polycaprolactone copolymer, and polylactic acid-polycaprolactone copolymer, and even more preferably at least one selected from the group consisting of polyglycolic acid, polyglycolic acid-poly-L-lactic acid copolymer, polyglycolic acid-polycaprolactone copolymer, and poly-L-lactic acid-polycaprolactone copolymer, and is particularly preferably polyglycolic acid. By making the first biodegradable polymer one of the polymers described above, the degradation rate of the porous structure becomes suitable, and the occurrence of inflammation originating from the porous structure can be more effectively suppressed.

[0063] The average molecular weight (number-average molecular weight) of the first biodegradable polymer is preferably 10,000 to 150,000, more preferably 40,000 to 120,000, even more preferably 50,000 to 100,000, particularly preferably 60,000 to 90,000, and most preferably 70,000 to 80,000. By having the average molecular weight of the first biodegradable polymer within the above range, the degradation rate of the porous structure becomes suitable, and the occurrence of inflammation originating from the porous structure can be more effectively suppressed. The average molecular weight (number-average molecular weight) of the first biodegradable polymer can be measured using gel permeation chromatography according to the method described in the reference example.

[0064] The crystallinity of the first biodegradable polymer is preferably 10% to 50%, more preferably 20% to 45%, and even more preferably 30% to 40%. By having the crystallinity of the first biodegradable polymer within the above range, the degradation rate of the porous structure becomes more suitable, and the occurrence of inflammation originating from the porous structure can be more effectively suppressed. The crystallinity of the first biodegradable polymer can be measured by differential scanning calorimetry (DSC).

[0065] [Antithrombotic Coating Layer] The in-vivo implant according to the present invention comprises an antithrombotic coating layer that coats at least a portion of the surface of the porous structure described above. The antithrombotic coating layer is composed of a block copolymer of an antithrombotic polymer and a second biodegradable polymer. The presence of the antithrombotic coating layer on the surface of the porous structure prevents the adhesion of thrombi and the like to the porous structure, thereby reducing the risk of thrombosis and the like.

[0066] As described above, the antithrombotic coating layer consists of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, but it is preferable that the antithrombotic polymer is a hydrophilic polymer and the second biodegradable polymer is a hydrophobic polymer. With the antithrombotic coating layer having such a configuration, the adhesion of thrombi and the like to the porous structure is more effectively prevented, and the risk of thrombosis and the like is further reduced. In this specification, a hydrophilic polymer is a polymer whose solubility in 100 g of water at 20°C is 1 g or more. On the other hand, a hydrophobic polymer is a polymer whose solubility in 100 g of water at 20°C is less than 1 g.

[0067] The average molecular weight (number-average molecular weight) of the block copolymer constituting the antithrombotic coating layer according to one embodiment is preferably 2,000 to 50,000, more preferably 4,000 to 25,000, and even more preferably 8,000 to 15,000. Having the average molecular weight of the block copolymer within the above range ensures sufficient thickness of the antithrombotic coating layer, making it less susceptible to wear and peeling. The average molecular weight (number-average molecular weight) of the block copolymer can be measured using gel permeation chromatography.

[0068] The thickness of the antithrombotic coating layer is not particularly limited, but is preferably 0.001 μm to 1 μm, more preferably 0.005 μm to 0.1 μm, and even more preferably 0.01 μm to 0.05 μm. By having the antithrombotic coating layer within the above thickness range, sufficient antithrombotic properties can be imparted to the antithrombotic coating layer, and the durability of the coating layer against peeling, friction, etc., is improved.

[0069] <Antithrombotic Polymers> Antithrombotic polymers are not particularly limited as long as they possess antithrombotic properties. Here, antithrombotic properties refer to the ability to suppress the formation of blood clots. For example, excellent antithrombotic properties mean that blood clots are less likely to form, or if they do form, they are rapidly broken down.

[0070] In particular, the monomer constituting the antithrombotic polymer is preferably one or more selected from the group consisting of ethylene glycol, methoxyethylene glycol, and propylene glycol, more preferably one or more selected from the group consisting of ethylene glycol and methoxyethylene glycol, and even more preferably ethylene glycol. Furthermore, from the viewpoint of excellent antithrombotic properties and biocompatibility, the antithrombotic polymer is preferably one or more selected from the group consisting of polyethylene glycol, methoxypolyethylene glycol, and polypropylene glycol, more preferably one or more selected from the group consisting of polyethylene glycol and methoxypolyethylene glycol, and even more preferably polyethylene glycol.

[0071] The block molecular weight (number-average molecular weight) of the antithrombotic polymer in the block copolymer is preferably 1,000 to 10,000, more preferably 2,500 to 7,500, and even more preferably 4,000 to 6,000. Having the block molecular weight of the antithrombotic polymer within the above range further improves the antithrombotic properties of the antithrombotic coating layer. The block molecular weight (number-average molecular weight) of the antithrombotic polymer can be measured, for example, by gel permeation chromatography (GPC).

[0072] <Second Biodegradable Polymer> The second biodegradable polymer is not particularly limited as long as it is a biodegradable polymer, and can be appropriately selected from the biodegradable materials described in the [Stent] section above.

[0073] In particular, the monomer constituting the second biodegradable polymer is preferably one or more selected from the group consisting of lactic acid, caprolactone, and glycolic acid, more preferably one or more selected from the group consisting of lactic acid and glycolic acid, and preferably lactic acid, from the viewpoint of suitability for medical applications.

[0074] Furthermore, the second biodegradable polymer is preferably an aliphatic polyester, more preferably one or more selected from the group consisting of polylactic acid, polycaprolactone, and polyglycolic acid, more preferably one or more selected from the group consisting of polylactic acid and polyglycolic acid, and most preferably polylactic acid. By having the second biodegradable polymer be the polymer described above, the duration of the antithrombotic coating layer's function is optimized, and the occurrence of inflammation originating from the porous structure can be more effectively suppressed.

[0075] The block molecular weight (number-average molecular weight) of the second biodegradable polymer in the block copolymer is preferably 1,000 to 10,000, more preferably 2,500 to 7,500, and even more preferably 4,000 to 6,000. By using the above-described polymer as the second biodegradable polymer, the duration of the antithrombotic coating layer's function is optimized, and the occurrence of inflammation originating from the porous structure can be more effectively suppressed. The block molecular weight of the second biodegradable polymer can be measured, for example, by gel permeation chromatography (GPC).

[0076] For example, in the antithrombotic coating layer of an in-vivo implant according to one embodiment, it is preferable that the monomer constituting the antithrombotic polymer is one or more selected from the group consisting of ethylene glycol, methoxyethylene glycol, and propylene glycol, and that the monomer constituting the second biodegradable polymer is one or more selected from the group consisting of lactic acid, caprolactone, and glycolic acid, more preferably that the monomer constituting the antithrombotic polymer is one or more selected from the group consisting of ethylene glycol and methoxyethylene glycol, and that the monomer constituting the second biodegradable polymer is one or more selected from the group consisting of lactic acid and glycolic acid, and even more preferably that the monomer constituting the antithrombotic polymer is ethylene glycol and the monomer constituting the second biodegradable polymer is lactic acid. By having the second biodegradable polymer and the antithrombotic polymer composed of the above-mentioned monomers, the duration of the function maintenance of the antithrombotic coating layer becomes appropriate, and the antithrombotic properties are further improved, so that the occurrence of inflammation originating from the porous structure can be more effectively suppressed.

[0077] For example, in the antithrombotic coating layer of an in-vivo implant according to one embodiment, it is preferable that the antithrombotic polymer is one or more selected from the group consisting of polyethylene glycol, methoxypolyethylene glycol, and polypropylene glycol, and that the second biodegradable polymer is one or more selected from the group consisting of polylactic acid, polycaprolactone, and polyglycolic acid; it is more preferable that the antithrombotic polymer is one or more selected from the group consisting of polyethylene glycol and methoxypolyethylene glycol, and that the second biodegradable polymer is one or more selected from the group consisting of polylactic acid and polyglycolic acid; and it is even more preferable that the antithrombotic polymer is polyethylene glycol and the second biodegradable polymer is polylactic acid. In other words, the block copolymer constituting the antithrombotic coating layer is preferably one or more selected from the group consisting of polyethylene glycol-polylactic acid, methoxypolyethylene glycol-polylactic acid, polypropylene glycol-polylactic acid, polyethylene glycol-polycaprolactone, methoxypolyethylene glycol-polycaprolactone, polypropylene glycol-polycaprolactone, polyethylene glycol-polyglycolic acid, methoxypolyethylene glycol-polyglycolic acid, and polypropylene glycol-polyglycolic acid, more preferably polyethylene glycol-polylactic acid, and even more preferably polyethylene glycol-poly-L-lactic acid. By having a block copolymer constituting the antithrombotic coating layer that is composed of the above-mentioned antithrombotic polymer and a second biodegradable polymer, the duration of function maintenance of the antithrombotic coating layer becomes appropriate, and the antithrombotic properties are further improved, thereby more effectively suppressing the occurrence of inflammation originating from the porous structure.

[0078] In an in-vivo implant according to one embodiment, preferably, the first biodegradable polymer is at least one selected from the group consisting of polyglycolic acid, polyglycolic acid-polylactic acid copolymer, polyglycolic acid-polycaprolactone copolymer, and polylactic acid-polycaprolactone copolymer; the antithrombotic polymer is one or more selected from the group consisting of polyethylene glycol, methoxypolyethylene glycol, and polypropylene glycol; and the second biodegradable polymer is one or more selected from the group consisting of polylactic acid, polycaprolactone, and polyglycolic acid. Furthermore, in an in-vivo implant according to one embodiment, more preferably, the first biodegradable polymer is polyglycolic acid, the antithrombotic polymer is polyethylene glycol, and the second biodegradable polymer is polylactic acid. By having such a configuration, inflammation originating from the porous structure made of biodegradable polymers and its onset can be more effectively suppressed.

[0079] Furthermore, in the block copolymer of a second biodegradable polymer and an antithrombotic polymer constituting the antithrombotic coating layer according to one embodiment, it is preferable that the block molecular weight (number average molecular weight) of the antithrombotic polymer is 1,000 to 10,000, and the block molecular weight (number average molecular weight) of the second biodegradable polymer is 1,000 to 10,000. It is more preferable that the block molecular weight of the antithrombotic polymer is 2,500 to 7,500, and the block molecular weight of the second biodegradable polymer is 2,500 to 7,500, and even more preferable that the block molecular weight of the antithrombotic polymer is 4,000 to 6,000, and the block molecular weight of the second biodegradable polymer is 4,000 to 6,000. By having the block molecular weights of the second biodegradable polymer and the antithrombotic polymer within the above ranges, both a suitable duration of function and sufficient antithrombotic properties can be obtained.

[0080] Furthermore, in the block copolymer constituting the antithrombotic coating layer according to one embodiment, the ratio (molar ratio, moles of antithrombotic polymer:moles of second biodegradable polymer) of the antithrombotic polymer to the second biodegradable polymer is preferably 40:60 to 80:20, more preferably 45:55 to 70:30, and even more preferably 50:50 to 65:35. By having the ratio (molar ratio) of the second biodegradable polymer to the antithrombotic polymer within the above range, both a suitable duration of function and sufficient antithrombotic properties can be obtained.

[0081] [Decomposition Time and Functional Maintenance Time] In the in vivo implant of the present invention, the decomposition time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer. With this configuration, the function of the antithrombotic coating layer is lost after the porous structure has decomposed and disappeared. Therefore, as long as the porous structure is present, the function of the antithrombotic coating layer is maintained, thereby reducing the risk of thrombosis and other complications caused by the porous structure.

[0082] In an in-vivo implant according to one embodiment, the decomposition time of the porous structure is preferably 3.6 months or less (where "or less" is also referred to as "within"), more preferably 3.5 months or less, even more preferably 3 months or less, particularly preferably 2.5 months or less, and most preferably 2 months or less. Furthermore, the decomposition time of the porous structure is preferably 0.5 months or more, more preferably 1 month or more, and even more preferably 1.5 months or more. For example, the preferred decomposition time for the porous structure may be 0.5 months to 3.6 months, 0.5 months to 3.5 months, 0.5 months to 3 months, 0.5 months to 2.5 months, 0.5 months to 2 months, 1 month to 3.6 months, 1 month to 3.5 months, 1 month to 3 months, 1 month to 2.5 months, 1 month to 2 months, 1.5 months to 3.6 months, 1.5 months to 3.5 months, 1.5 months to 3 months, 1.5 months to 2.5 months, or 1.5 months to 2 months. If the decomposition time of the porous structure is within the above range, the porous structure can sufficiently suppress the scattering of thrombi and plaque, and will decompose and disappear at an appropriate time after fulfilling its role in preventing the scattering of thrombi and plaque. Here, the decomposition time of a porous structure as used herein refers to the period (time) during which the porous structure decomposes, and is the time measured according to the method described in the reference example [Measurement of Decomposition Time of Porous Structure (First Measurement Method)]. In this specification, one month is defined as 30.4 days.

[0083] In an in-vivo implant according to one embodiment, the duration of function maintenance of the antithrombotic coating layer is preferably more than 3.6 months, more preferably 4 months or more, even more preferably 4.2 months or more, and particularly preferably 4.5 months or more. Furthermore, the duration of function maintenance of the antithrombotic coating layer is preferably 18 months or less, more preferably 15 months or less, even more preferably 12 months or less, and particularly preferably 9 months or less. For example, the preferred duration of function of the antithrombotic coating layer may be more than 3.6 months and up to 18 months, 4 months and up to 18 months, 4.2 months and up to 18 months, 4.5 months and up to 18 months, more than 3.6 months and up to 15 months, 4 months and up to 15 months, 4.2 months and up to 15 months, 4.5 months and up to 15 months, more than 3.6 months and up to 12 months, 4 months and up to 12 months, 4.2 months and up to 12 months, 4.5 months and up to 12 months, more than 3.6 months and up to 9 months, 4 months and up to 9 months, 4.2 months and up to 9 months, or 4.5 months and up to 9 months. By having the duration of function of the antithrombotic coating layer within the above range, the antithrombotic function can be maintained until the porous structure is degraded, in relation to the degradation time of the porous structure. Here, the duration of function of the antithrombotic coating layer as used herein refers to the period (time) during which the antithrombotic function of the layer is maintained, and is the period measured according to the method described in the reference example [Measurement of duration of function of the antithrombotic coating layer (second measurement method)].

[0084] In an in-vivo implant according to one embodiment, there are no particular limitations as long as the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer. For example, it is preferable that the degradation time of the porous structure is 3.6 months or less and the functional maintenance time of the antithrombotic coating layer exceeds 3.6 months. It is more preferable that the degradation time of the porous structure is 3.5 months or less and the functional maintenance time of the antithrombotic coating layer is 4 months or more. It is even more preferable that the degradation time of the porous structure is 3 months or less and the functional maintenance time of the antithrombotic coating layer is 4 months or more. It is even more preferable that the degradation time of the porous structure is 2.5 months or less and the functional maintenance time of the antithrombotic coating layer is 4 months or more. It is particularly preferable that the degradation time of the porous structure is 2.5 months or less and the functional maintenance time of the antithrombotic coating layer is 4.2 months or more. It is most preferable that the degradation time of the porous structure is 2.5 months or less and the functional maintenance time of the antithrombotic coating layer is 4.5 months or more. If the degradation time of the porous structure and the duration of the antithrombotic coating layer's function are within the above range, the function of the antithrombotic coating layer can be more reliably maintained until the porous structure degrades, and the risk of inflammation caused by the porous structure can be sufficiently reduced.

[0085] Furthermore, the difference between the degradation time of the porous structure and the functional maintenance time of the antithrombotic coating layer is preferably 30 days or more, more preferably 40 days or more, even more preferably 50 days or more, particularly preferably 60 days or more, and most preferably 70 days or more. In addition, the difference between the degradation time of the porous structure and the functional maintenance time of the antithrombotic coating layer is not particularly limited, but is preferably 200 days or less, more preferably 150 days or less, and even more preferably 100 days or less. With such a difference, the function of the antithrombotic coating layer can be maintained more reliably until the porous structure degrades, and thus the risk of inflammation caused by the porous structure can be sufficiently reduced.

[0086] [Other Components] The antithrombotic coating layer may contain other components as long as the effects of the present invention are not impaired. Other components are not particularly limited and include, for example, anticancer agents, immunosuppressants, antibiotics, antirheumatic drugs, antithrombotic drugs, HMG-CoA reductase inhibitors, ACE inhibitors, calcium channel blockers, antihyperlipidemic drugs, integrin inhibitors, antiallergic agents, antioxidants, GPIIb / IIIa antagonists, retinoids, flavonoids, carotenoids, lipid-improving agents, DNA synthesis inhibitors, tyrosine kinase inhibitors, antiplatelet agents, vascular smooth muscle proliferation inhibitors, anti-inflammatory drugs, bio-derived materials, interferon, and NO production-promoting agents (bioactive substances). The amount of other components added is not particularly limited and is applied in the same manner as commonly used amounts. Ultimately, the amount of other components added is appropriately selected considering the severity of the disease to which it is applied, the patient's weight, etc.

[0087] <Method for Manufacturing Intravivo Implants> One embodiment of the present invention is a method for manufacturing an intravivo implant comprising an expandable tubular stent, a porous structure disposed to cover at least a portion of the stent, and an antithrombotic coating layer coating at least a portion of the surface of the porous structure, wherein the porous structure is made of a first biodegradable polymer, the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer.

[0088] A method for manufacturing an in-vivo implant according to one embodiment includes the steps of (i) preparing a coating solution, (ii) applying the coating solution to at least a portion of the surface of a porous structure to form an antithrombotic coating layer, and (iii) attaching the porous structure to an expandable cylindrical stent.

[0089] (i) Step of preparing the coating solution: In this step, a coating solution is prepared which includes a block copolymer of an antithrombotic polymer and a second biodegradable polymer, used to form an antithrombotic coating layer that coats at least a portion of the surface of the porous structure.

[0090] The coating solution can be prepared by dissolving and dispersing the block copolymer of the antithrombotic polymer and the second biodegradable polymer described above in a predetermined solvent. The solvent for the coating solution is not particularly limited as long as it can sufficiently dissolve or disperse the block copolymer, but a mixed solvent of a good solvent and a poor solvent is preferred.

[0091] Examples of good solvents include, but are not limited to, ketones such as acetone, methyl ethyl ketone, and cyclohexanone; esters such as ethyl acetate; halides such as chloroform; olefins such as hexane; ethers such as tetrahydrofuran (THF) and butyl ether; nitriles such as acetonitrile, propionite, and benzonitrile; aromatics such as benzene and toluene; amides such as N,N-dimethylformamide (DMF); and sulfoxides such as dimethyl sulfoxide. One of these good solvents may be used alone, or two or more may be used in combination.

[0092] Examples of poor solvents include water; and lower alcohols having 1 to 6 carbon atoms, such as methanol, ethanol, n-propanol, iso-propanol, n-butanol, sec-butanol, and tert-butanol, but the invention is not limited to these. These poor solvents may be used individually or in combination of two or more.

[0093] For each solvent combination, it is preferable that the good solvent includes at least one selected from the group consisting of acetone, tetrahydrofuran, and acetonitrile, and the poor solvent includes at least one selected from the group consisting of water, methanol, ethanol, n-propanol, iso-propanol, n-butanol, sec-butanol, and tert-butanol. A more preferred embodiment is that the good solvent includes acetone, and the poor solvent includes at least one selected from the group consisting of methanol, ethanol, and n-propanol. A particularly preferred embodiment is that the good solvent includes acetone, and the poor solvent includes ethanol.

[0094] In one embodiment, the volume ratio of the good solvent to the poor solvent (good solvent:poor solvent) is more preferably 0.5:9.5 to 5:5, more preferably 1:9 to 4:6, and even more preferably 1:9 to 3:7.

[0095] The coating solution is prepared by mixing the block copolymer of the antithrombotic polymer and the second biodegradable polymer described above, a solvent, and other components added as needed. The order and method of adding these components are not particularly limited. Each of the above components can be added to a mixing container all at once or separately, in stages or sequentially. The mixing method is also not particularly limited, and known methods can be used.

[0096] The concentration of the block copolymer in the coating solution according to one embodiment is not particularly limited, but is preferably 0.1 mg / mL or more and 15 mg / mL or less, more preferably 1 mg / mL or more and 10 mg / mL or less, and even more preferably 2 mg / mL or more and 5 mg / mL or less. When the concentration of the block copolymer in the coating solution is within the above range, the viscosity of the coating solution is within an appropriate range, which is advantageous in terms of operability (e.g., ease of coating, i.e., applicability) and production efficiency.

[0097] The temperature (liquid temperature) when mixing the above components is not particularly limited, but it is preferably 0 to 60°C, and more preferably 10 to 30°C.

[0098] (ii) A step of applying the coating solution to at least a portion of the surface of a porous structure to form an antithrombotic coating layer.

[0099] In this process, the porous structure may be a commercially available one or one that has been manufactured. As for the manufacturing method, for example, the method described in International Publication No. 2008 / 062414 can be applied in the same manner or with appropriate modifications.

[0100] The method for applying a coating solution to at least a portion of the surface of a porous structure (porous structure surface) is not particularly limited, and conventionally known methods such as coating / printing, immersion (dipping method, dip coating method), spraying method, and spin coating method can be applied. Of these, the immersion method (dipping method, dip coating method) is preferred as the coating method because it is easy to form a coating on porous structures that have a fine structure.

[0101] The conditions for the immersion method are not particularly limited. For example, the temperature during immersion (temperature of the coating solution) is preferably 0 to 50°C, and more preferably 10 to 30°C. Also, for example, the immersion time is preferably 8 hours or less, and more preferably 5 hours or less. By setting the immersion time within the above range, the volume exclusion effect by the copolymer is increased, and the antithrombotic properties of the coating are further improved. This is also preferable from the viewpoint of productivity. On the other hand, the immersion time is preferably 5 seconds or more, more preferably 10 seconds or more, and even more preferably 1 minute or more. By setting the immersion time within the above range, the copolymer (coating) can be applied almost uniformly to the surface of the porous structure. In addition, the amount of copolymer adsorbed onto the porous structure increases, which improves the antithrombotic properties.

[0102] Furthermore, when forming a coating (coating with a copolymer) only on a portion of a porous structure, the coating can be formed on the desired surface area of ​​the porous structure by immersing only a portion of the porous structure in a coating solution and coating that portion with the solution.

[0103] Furthermore, by drying the applied coating solution, an antithrombotic coating layer is formed on at least a portion of the surface of the porous structure. The coating solution may be dried at room temperature or by applying heat. When heat is applied, the heating temperature is preferably 50°C to 200°C. The drying time is preferably 30 minutes to 24 hours. For example, if the heating temperature is 50°C to 100°C, the heating time may be 1 hour to 6 hours.

[0104] (iii) Step of attaching a porous structure to an expandable cylindrical stent This step is a step of manufacturing an in vivo implant by attaching (crimping) a porous structure, which has an antithrombotic coating layer formed on at least a part of its surface, to an expandable cylindrical stent in the step (ii) described above.

[0105] Here, the stent may be a commercially available product or a manufactured product, as long as it has the configuration described in the [Stent] section. The method of manufacturing the stent is not particularly limited and can be appropriately selected from general manufacturing methods used depending on the structure and material of the stent. For example, a manufacturing method using etching techniques such as laser etching and chemical etching, and laser cutting techniques can be selected. Furthermore, the materials constituting the stent can be appropriately selected from the resin materials, metal materials, and ceramic materials described in the [Stent] section above.

[0106] The method for attaching the porous structure to the stent is not particularly limited, and known methods can be used. For example, one method involves placing the porous structure over the stent, pressing it with a pressing body such as silicone rubber, and then irradiating it with laser light or the like to fix the porous structure to the stent.

[0107] The in-vivo implant produced by the above manufacturing method can effectively suppress the onset of inflammation originating from the porous structure made of biodegradable polymer when implanted in a blood vessel.

[0108] While embodiments of the present invention have been described in detail, these are descriptive and illustrative, and not limiting, and it is clear that the scope of the present invention should be interpreted by the appended claims.

[0109] This application is based on Japanese Patent Application No. 2025-056468, filed on 28 March 2025, the disclosures of which are referenced and incorporated in whole.

[0110] 10: Stent, 10A: Tip, 10B: Proximal end, 11: Ring, 12: Link section, 14: Third strut section, 15: First strut section, 16: Second strut section, 17: Curved section, 19: Outer surface, 20: Porous structure, 30: Antithrombotic coating layer, 100: Intravivo implant, 200: Balloon catheter, 210: Catheter body, 220: Balloon, 230: Hub, 231: Proximal opening, 300: Stent delivery system.

[0111] (Reference Examples) The effects of the present invention will be explained using the following reference examples. However, the technical scope of the present invention is not limited to the following reference examples. In the reference examples below, unless otherwise specified, the operations were performed at room temperature (25°C). Also, unless otherwise specified, "%" and "parts" mean "mass%" and "parts by mass," respectively.

[0112] [Measurement of Decomposition Time of Porous Structures (First Measurement Method)] The decomposition time of each biodegradable polymer of Reference Examples 1 to 4 was measured according to the following procedure. Reference Examples 1 to 3 correspond to the first biodegradable polymer constituting the porous structure. Reference Example 1: Polyglycolic acid (PGA, number average molecular weight 71010, crystallinity 38%) Reference Example 2: Polyglycolic acid-poly L-lactic acid block copolymer (PGA-PLLA, glycolic acid: L-lactic acid = 90:10 (molar ratio), number average molecular weight 86947, crystallinity 23%) Reference Example 3: Polyglycolic acid-polycaprolactone block copolymer (PGA-PCL, glycolic acid: caprolactone = 75:25 (molar ratio), number average molecular weight 64239, crystallinity 22%) Reference Example 4: Poly L-lactic acid (PLLA, number average molecular weight 82943, crystallinity 45%).

[0113] 0.3 mg each of the biodegradable polymers from Reference Examples 1-4 was measured and placed in screw-cap test tubes. Next, pH 7.4 phosphate-buffered saline (PBS) was poured into the test tubes up to the 15 ml line, and the biodegradable polymers from Reference Examples 1-4 were immersed. In this process, phosphate-buffered saline (PBS) was used as a simulated body fluid. The test tubes were then stored in a 37°C constant temperature bath for a predetermined period (0, 1, 3, 7, and 14 days). After the predetermined period, the screw-cap test tubes were removed from the constant temperature bath, the PBS was aspirated from the test tubes using a Pasteur pipette, and the water in the test tubes was completely evaporated by vacuum drying. Then, a mobile phase solvent (hexafluoroisopropanol containing 5 mM sodium trifluoroacetate) was added to the test tubes so that the concentration of the biodegradable polymers from Reference Examples 1-4 was 1 mg / ml, and after the samples were completely dissolved, the weight-average molecular weight (Mw) and number-average molecular weight (Mn) were measured according to the method described below. Figure 6 shows the measurement results for weight-average molecular weight (Mw) and number-average molecular weight (Mn) on days 0, 1, 3, 7, and 14. In Figure 6, PGA represents the results for Reference Example 1, PGA-PLLA for Reference Example 2, PGA-PCL for Reference Example 3, and PLLA for Reference Example 4.

[0114] (Method for measuring weight-average molecular weight (Mw) and number-average molecular weight (Mn)) The weight-average molecular weight (Mw) and number-average molecular weight (Mn) were measured using gel permeation chromatography (GPC) with polymethyl methacrylate as the standard substance under the following conditions: (Measurement conditions) Apparatus: Semi-micro GPC system LC-VP (Shimadzu Corporation) Detector: Shodex® RI-104 (Showa Denko K.K.) Column: Shodex® GPC LF-404 (Showa Denko K.K.) Column temperature: 40°C Mobile phase solvent: 5 mM sodium trifluoroacetate-containing hexafluoroisopropanol (HFIP) Sample concentration: 1 mg / ml Flow rate: 0.30 mL / min Injection volume: 10 μL Sample preparation: 1 mg of the sample to be measured was dissolved in 1 mL of mobile phase solvent, and then filtered through a 0.45 μm PTFE membrane filter.

[0115] Next, a logarithmic plot (semi-logarithmic plot) of the obtained number-average molecular weight (Mn) was created, and the time at which Mn was between 50 and 800 (preferably 50 to 150, more preferably 100) was identified from the linear regression equation (ln(y) = ax + b) of this plot. This time was defined as the period during which the porous structure decomposed and disappeared, i.e., the decomposition time. The results are shown in Table 1.

[0116]

[0117] [Measurement of the duration of function of the antithrombotic coating layer (second measurement method)] The duration of function of each biodegradable polymer of Reference Example 5 was measured according to the following procedure. Reference Example 5 corresponds to the block copolymer that constitutes the antithrombotic coating layer. Reference Example 5: Polyethylene glycol-poly L-lactic acid block copolymer (PEG-PLLA, manufactured by NOF Corporation, trade name SUNBRIGHT ME-050LA050, ethylene glycol: L-lactic acid = 59:41 (molar ratio), number average molecular weight of block copolymer 10,000 (PEG: number average molecular weight 5,000, PLLA: number average molecular weight 5,000)) (Preparation of coating solution) 90 mg of the above PEG-PLLA copolymer was weighed and dissolved in 30 ml of solvent (acetone:ethanol = 2:8 (volume ratio)) to prepare the coating solution.

[0118] (Coating process) Two polyethylene terephthalate (PET) films, each approximately 2 cm x 1 cm, were immersed in the coating solution at room temperature for about 10 seconds, and then the PET films were slowly pulled out by hand.

[0119] (Drying process) The PET film was left to stand in a vacuum oven maintained at 60°C for 2 hours to dry and solidify. After that, the PET film was washed with RO water at 40°C for about 1 hour, and then vacuum dried at room temperature to obtain a PEG-PLLA coated PET film (hereinafter referred to as the sample).

[0120] (Evaluation of Zeta Potential) The zeta potential of the sample surface was evaluated over time according to the following procedure. First, the sample was placed in a container with 10 mL of pure water added, and left standing at 37°C while immersed in the pure water.

[0121] Next, after 0, 34, 48, and 89 days of standing, samples were removed from the pure water and attached to the surface of the movable block in the clamp cell installed in the measuring device using special adhesive tape. The zeta potential was then measured under conditions of approximately pH 7, 300 mbar (30,000 Pa), and 1 mM KCl. The zeta potential of the PET film before the PEG-PLLA coating was applied was also measured using the same procedure. The results are shown in the left graph of Figure 7. In the graph, "Sample 0 Days" represents the zeta potential of the sample after 0 days of standing, "Sample 34 Days" represents the zeta potential of the sample after 34 days of standing, "Sample 48 Days" represents the zeta potential of the sample after 48 days of standing, and "Sample 89 Days" represents the zeta potential of the sample after 89 days of standing. Furthermore, the "PEG-PLLA uncoated PET film" is the result of measuring the zeta potential of the PET film before the PEG-PLLA coating was applied.

[0122] - Device: Solid surface ζ potential measurement device (SurPASS manufactured by Anton Paar).

[0123] Next, an approximation formula (linear function) was obtained using the least squares method for the four data points consisting of zeta potentials obtained at 0, 34, 48, and 89 days (see graph on the right in Figure 7). Subsequently, the period (number of days) during which the absolute value of the zeta potential was 20.0 mV or higher was calculated from this approximation formula, and this period was defined as the period during which the function of the antithrombotic coating layer is lost (function loss time). In other words, the period during which the absolute value of the zeta potential was less than 20.0 mV was defined as the period during which the function of the antithrombotic coating layer is maintained (function maintenance time). This is because if the absolute value of the zeta potential is less than 20.0 mV, it is considered that the function of the antithrombotic coating layer, which suppresses the adhesion of proteins and cells (i.e., suppresses the adhesion of thrombi), is maintained.

[0124] The evaluation results are shown in Table 2.

[0125]

[0126] As can be seen from Tables 1 and 2, the decomposition time of the biodegradable polymers in Reference Examples 1 to 3, which correspond to the first biodegradable polymer of the present invention, was shorter than the functional maintenance time of Reference Example 5, which corresponds to the block copolymer of the present invention.

Claims

1. An in vivo implant comprising: an expandable tubular stent; a porous structure disposed to cover at least a portion of the stent; and an antithrombotic coating layer coating at least a portion of the surface of the porous structure, wherein the porous structure is made of a first biodegradable polymer, the antithrombotic coating layer is made of a block copolymer of an antithrombotic polymer and a second biodegradable polymer, and the degradation time of the porous structure is shorter than the functional maintenance time of the antithrombotic coating layer.

2. The in vivo implant according to claim 1, wherein the first biodegradable polymer is at least one selected from the group consisting of polyglycolic acid, polyglycolic acid-polylactic acid copolymer, polyglycolic acid-polycaprolactone copolymer, and polylactic acid-polycaprolactone copolymer.

3. The in-vivo implant according to claim 1 or 2, wherein the number-average molecular weight (Mn) of the first biodegradable polymer is 10,000 to 150,000.

4. The in vivo implant according to claim 1 or 2, wherein the antithrombotic polymer is a hydrophilic polymer and the second biodegradable polymer is a hydrophobic polymer.

5. The in vivo implant according to claim 1 or 2, wherein the monomer constituting the antithrombotic polymer is one or more selected from the group consisting of ethylene glycol, methoxyethylene glycol, and propylene glycol, and the monomer constituting the second biodegradable polymer is one or more selected from the group consisting of lactic acid, caprolactone, and glycolic acid.

6. The in vivo implant according to claim 1 or 2, wherein the antithrombotic polymer is one or more selected from the group consisting of polyethylene glycol, methoxypolyethylene glycol, and polypropylene glycol, and the second biodegradable polymer is one or more selected from the group consisting of polylactic acid, polycaprolactone, and polyglycolic acid.

7. The in vivo implant according to claim 1 or 2, wherein the block copolymer has a block molecular weight of 1,000 to 10,000 of the antithrombotic polymer and a block molecular weight of 1,000 to 10,000 of the second biodegradable polymer.

8. The in-vivo implant according to claim 1 or 2, wherein the ratio (molar ratio) of the antithrombotic polymer to the second biodegradable polymer in the block copolymer constituting the antithrombotic coating layer is 40:60 to 80:

20.

9. The in-vivo implant according to claim 1 or 2, wherein the decomposition time of the porous structure is within 3.6 months.

10. The in vivo implant according to claim 9, wherein the functional maintenance time of the antithrombotic coating layer exceeds 3.6 months.

11. The in vivo implant according to claim 1, wherein the first biodegradable polymer is polyglycolic acid, the antithrombotic polymer is polyethylene glycol, and the second biodegradable polymer is polylactic acid.

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