In-vivo indwelling object and method for producing porous structure

Abstract

Provided is an in-vivo indwelling object capable of sufficiently expanding a stent in a radial direction while preventing expansion of a stent cover in the radial direction. The in-vivo indwelling object includes a diameter-expandable cylindrical stent, and a porous structure that is disposed so as to cover the stent and is configured so as to be expandable in accordance with the diameter expansion of the stent, wherein: the porous structure has a skeleton part disposed in a mesh shape, a plurality of void parts partitioned by the skeleton part, and a covering part disposed so as to cover at least a portion of the outer surface of the porous structure and including a polymer material; and the polymer material has a Young's modulus of greater than 0 MPa and less than 200 MPa.

Description

Method for manufacturing an implant and a porous structure

[0001] The present invention relates to an implant and a method for manufacturing a porous structure used for the implant.

[0002] A stent is an implant that is delivered to a lesion in a living body lumen by a stent delivery system and then implanted to treat various diseases caused by stenosis or occlusion of a living body lumen such as a blood vessel, and expands a lesion such as a stenosis or occlusion to secure a lumen.

[0003] For example, Japanese Patent Application Laid-Open No. 2018-161163 discloses a stent provided with a cover layer composed of a knitted fabric that covers the outer periphery of a stent body that can be expanded in order to prevent peripheral embolism (such as restenosis) during stent implantation. In the stent of Japanese Patent Application Laid-Open No. 2018-161163, when the stent body expands, the cover layer provided with meshes expands so as to follow the stent body, and it is possible to prevent the scattering of plaque and thrombus during the expansion of the stent body.

[0004] However, the inventors have found that when using a stent such as that disclosed in Japanese Patent Application Laid-Open No. 2018-161163, a phenomenon occurs in which the stent cover expands in the radial direction at a timing different from the intention of the operator during the procedure. When such a phenomenon occurs, for example, there is a risk that living body tissue will be caught in the expanded portion of the stent cover, and as a result, there may be problems such as damage to the stent cover or the stent itself falling off from a catheter or the like.

[0005] Further, in Japanese Patent Application Laid-Open No. 2018-161163, the stent cover is fixed to the stent over a full circumference in the circumferential direction by using a predetermined synthetic resin, but the inventors have also found that such a stent may not be able to expand sufficiently in the radial direction, that is, there is a risk of poor expansion.

[0006] The present invention has been made in view of the above circumstances, and an object thereof is to provide an implant that can sufficiently expand the stent in the radial direction while preventing the radial expansion of the stent cover.

[0007] The inventors of the present invention conducted diligent research to solve the above problems. As a result, they discovered that the above objectives can be achieved by an in vivo implant comprising a stent and a porous structure that is positioned to cover the stent, is configured to expand as the diameter of the stent expands, and has a covering portion containing a polymer material having predetermined physical properties, and thus completed the present invention.

[0008] The above objective can be achieved by the present invention having the following configuration, and the present invention encompasses the following aspects and forms.

[0009] One aspect of the present invention is an in vivo implant comprising: 1. an expandable cylindrical stent; and a porous structure disposed to cover the stent and configured to expand as the stent expands, wherein the porous structure has a mesh-like skeletal structure, a plurality of voids partitioned by the skeletal structure, and a covering portion disposed to cover at least a part of the outer surface of the porous structure and containing a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

[0010] 2. In the in-vivo implant described in 1. above, it is preferable that the elongation at break of the polymer material is 960% or more and 3500% or less; 3. In the in-vivo implant described in 1. or 2. above, it is preferable that the Young's modulus of the polymer material is greater than 0 MPa and 100 MPa or less; 4. In the in-vivo implant described in any of 1. to 3. above, it is preferable that the polymer material includes a biodegradable polymer; 5. In the in-vivo implant described in 4. above, it is preferable that the biodegradable polymer includes one or more selected from the group consisting of polylactic acid, polyglycolic acid, poly(ε-caprolactone), copolymer of lactic acid and ε-caprolactone, copolymer of lactic acid and glycolic acid, and copolymer of glycolic acid and ε-caprolactone; 6. In the in-vivo implant described in 4. above, it is preferable that the biodegradable polymer includes a polylactic acid-based resin; 7. In the in vivo implant described above, it is preferable that the polylactic acid resin contains a copolymer of lactic acid and a hydroxycarboxylic acid other than lactic acid; 8. In the in vivo implant described above in 7., it is preferable that the lactic acid is L-lactic acid and the hydroxycarboxylic acid other than lactic acid is ε-caprolactone; 9. In the in vivo implant described above in 8., it is preferable that the ratio of L-lactic acid to ε-caprolactone (moles of L-lactic acid:moles of ε-caprolactone) is 0.1 to 10:1; 10. In the in vivo implant described above in 8. or 9., in the polymer material, it is preferable that the ratio (mass%) of the copolymer of L-lactic acid and ε-caprolactone to the total mass of the polymer material is greater than 25% by mass and 100% by mass or less; 11. In the in vivo implant described above in 1. to 10. In the in-vivo implant described in any of the above, it is preferable that the covering portion is arranged to cover at least one of the ends of the porous structure located in the longitudinal direction in the circumferential direction; 12. In the in-vivo implant described in any of the above 1. to 11., it is preferable that the porous structure is fixed to the stent via the covering portion.

[0011] Another aspect of the present invention is a method for manufacturing a porous structure used in an in-vivo implant, comprising: 13. a step of preparing a porous structure to be used in an in-vivo implant, which comprises an expandable cylindrical stent and a porous structure disposed to cover the stent and configured to expand as the diameter of the stent expands; and a step of forming a covering portion on at least a part of the outer surface of the porous structure by applying a covering material containing a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

[0012] This is a schematic plan view showing a stent delivery system equipped with a biological implant according to an embodiment. This is a plan view showing the reduced diameter state of the biological implant according to an embodiment. This is an enlarged view showing a part of the expanded diameter state of the biological implant according to an embodiment. This is an enlarged perspective view showing a part of the biological implant according to an embodiment. This is an enlarged cross-sectional view perpendicular to the axis showing a part of the biological implant according to an embodiment. This is an enlarged cross-sectional view along the axial direction showing a part of the biological implant according to an embodiment. This is a schematic cross-sectional view showing a part of the biological implant according to an embodiment. This is a schematic cross-sectional view showing a part of the biological implant according to an embodiment. This is a diagram for explaining the manufacturing method of the biological implant according to an embodiment. This is a diagram for explaining the manufacturing method of the biological implant according to an embodiment. This is a diagram for explaining the manufacturing method of the biological implant according to an embodiment. This is a perspective view showing an enlarged part of a porous structure according to a modified example. This is an enlarged cross-sectional view perpendicular to the axis showing a part of the biological implant according to a modified example. This is an enlarged cross-sectional view along the axial direction showing a part of the biological implant according to a modified example. This is a diagram for explaining the manufacturing method of the biological implant according to a modified example. This is a diagram for explaining the manufacturing method of the biological implant according to a modified example. This is a diagram of a porous structure showing a simplified example of the arrangement of the covering part. This is a diagram of a porous structure showing a simplified example of the arrangement of the covering part. This is a diagram of a porous structure showing a simplified example of the arrangement of the covering portion. This is a diagram of a porous structure showing a simplified example of the arrangement of the covering portion. This is a diagram showing the results of the embodiment.

[0013] An in vivo implant according to one aspect of the present invention comprises an expandable cylindrical stent and a porous structure disposed to cover the stent and configured to expand in accordance with the expansion of the stent, wherein the porous structure has a mesh-like skeletal portion, a plurality of voids partitioned by the skeletal portion, and a covering portion disposed to cover at least a part of the outer surface of the porous structure and made of a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

[0014] The inventors have found that an in vivo implant having such a configuration allows for sufficient radial expansion of the stent while preventing unintended radial expansion of the stent cover (a porous structure, particularly a mesh-like skeletal structure). The mechanism by which the above effects are achieved by the configuration of the present invention is presumed to be as follows.

[0015] In the in vivo implant according to the present invention, the porous structure covering the stent has at least a portion of its outer surface covered by a covering. The presence of this covering suppresses the radial expansion of the porous structure (particularly the mesh-like skeletal structure). On the other hand, the presence of the covering that suppresses the radial expansion of the porous structure may also hinder the radial expansion of the stent. However, the polymer material contained in the covering according to the present invention has a Young's modulus within a predetermined range. This results in a suitable flexibility of the covering, and in combination with the other components of the in vivo implant, it is presumed that the in vivo implant according to the present invention will prevent the radial expansion of the porous structure (particularly the mesh-like skeletal structure) while allowing the stent to expand sufficiently without being hindered during expansion.

[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," etc. 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 pertains. 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 meaning of terms as defined in the claims. Furthermore, the dimensional ratios in the drawings are exaggerated for illustrative purposes and may differ from actual ratios.

[0019] (Embodiment) An embodiment will be described with reference to Figures 1 to 10.

[0020] Figure 1 shows a stent delivery system 300 equipped with a biological implant 100 according to an embodiment. Figures 2 to 8 are diagrams illustrating the biological implant 100, stent 10, and porous structure 20 according to an embodiment. Specifically, Figure 2 is a schematic plan view showing the reduced diameter state of the stent 10 and porous structure 20 according to an embodiment, and Figure 3 is a schematic plan view showing the expanded diameter state of the stent 10 according to an embodiment. Figures 4 to 8 are diagrams illustrating characteristic parts of the biological implant 100 according to an embodiment. Figures 9 and 10 are diagrams illustrating the manufacturing method of the porous structure 20 according to an embodiment.

[0021] For the sake of clarity, the following directions are defined in this specification.

[0022] The longitudinal direction in which the stent 10 and the porous structure 20 extend is defined as the "axial direction." The axial direction is the direction from the tip 10A to the base 10B (or from the base 10B to the tip 10A) as shown in Figure 2, and is indicated by arrows X1-X2 in each figure.

[0023] In the stent 10 and the porous structure 20, the side inserted into the living body is referred to as the "proximal side," and the side opposite the proximal side, on which the surgeon operates the stent delivery system 300, is referred to as the "proximal side." The "proximal portion" refers to the part that includes a certain range extending from the proximal (frontmost) end towards the proximal end, and the "proximal end" refers to the part that includes a certain range extending from the proximal (very proximal) end towards the proximal end. Furthermore, the rotational direction relative to the axial direction is defined as the circumferential direction and is indicated by arrows R1-R2 in the figure.

[0024] <Stent Delivery System 300> As shown in Figure 1, the biological implantation 100 according to this embodiment is positioned on the outer circumference of the expandable and deflated balloon 220 provided by the balloon catheter 200.

[0025] 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 base end of the catheter body 210.

[0026] The balloon catheter 200 to which the implanted biological device 100 is attached constitutes a stent delivery system 300. The stent delivery system 300 delivers the implanted biological device 100, to which the porous structure 20 is attached, in a deflated state to the lesion site, and expands the stent 10 and the porous structure 20 as the balloon 220 expands, thereby allowing the stent 10 and the porous structure 20 to be placed in the lesion site.

[0027] The balloon catheter 200 can be configured as a rapid exchange type balloon catheter, for example, by introducing a guidewire W from near the tip of the catheter body 210 and inserting the guidewire W through to the tip of the balloon 220. The balloon 220 can also be configured as a so-called over-the-wire type balloon catheter.

[0028] For example, organic polymer materials can be used as the material constituting the balloon 220. Specifically, polymer materials such as polyolefins (e.g., polyethylene, polypropylene, polybutene, ethylene-propylene copolymer, ethylene-vinyl acetate copolymer, ionomer, or mixtures of two or more of these), polyvinyl chloride, polyamide, polyamide elastomer, polyurethane, polyurethane elastomer, polyimide, fluororesin, or mixtures thereof, or elastic resin materials such as two or more of the above polymer materials can be used, and among these, polyamide resins can be preferably used as the main material.

[0029] <Biological implant 100> The biological implant 100 according to this embodiment is used to treat narrowing or obstruction in blood vessels, bile ducts, trachea, esophagus, urethra, or other biological tubular lumens. The stent 10 used in the biological implant 100 is configured as a so-called balloon-expandable medical device, which is placed in a crimped state on a folded balloon 220, delivered to the lesion, and then expanded and placed in the lesion.

[0030] Furthermore, the stent 10 can also be constructed as a so-called self-expanding medical device, made of a self-expanding material.

[0031] As shown in Figures 2 and 3, the biological implantation device 100 comprises an expandable stent 10 and a porous structure 20 positioned to cover the stent 10 and configured to expand in accordance with the expansion of the stent 10.

[0032] <Stent 10> As shown in Figures 2 and 3, the stent 10 has a cylindrical shape that extends in the axial direction.

[0033] As shown in Figure 3, the stent 10 has linear rings 11 that form the outer circumference of a cylindrical shape with gaps, and link portions 12 that connect the rings 11 in gaps partitioned between adjacent linear rings 11 in the axial direction.

[0034] The ring 11 extends circumferentially around the stent 10 in a wave-like pattern that reciprocates in the axial direction.

[0035] The ring 11 has multiple first strut portions 14, which are straight or curved; second strut portions 15, which are straight or curved; and curved portions 17, which are formed between the first strut portions 14 and the second strut portions 15.

[0036] Furthermore, as shown in Figure 3, the ring 11 has a plurality of third strut portions 16, which are straight or curved, and are adjacent to one side of the link portion 12 in the axial direction and are provided in pairs in the circumferential direction.

[0037] The rings 11 are arranged sequentially along the axial direction. Adjacent rings 11 in the axial direction are connected and integrated by link portions 12.

[0038] The stent 10 can be configured as a drug-eluting stent, for example, in which at least a portion of the stent 10 is coated with a drug. The drug coated on the stent 10 can be supported by a predetermined polymer, for example, to form a drug-carrying portion 60. As the polymer, for example, a biodegradable polymer can be used.

[0039] Figure 7 shows a schematic cross-sectional view of the link portion 12 (or curved portion 17). Figure 8 shows a schematic cross-sectional view of the first strut portion 14 (or third strut portion 16).

[0040] As shown in FIG. 8, for example, the drug-carrying portion 60 can be disposed on the outer surface of the first strut portion 14 of the ring 11 and / or the outer surface 19 of the third strut portion 16. On the other hand, as shown in FIG. 7, it is preferable not to form the drug-carrying portion 60 on the curved portion 17 of the ring 11 and the link portion 12 (the portion where stress concentration and / or distortion occur as the stent 10 expands in diameter). By configuring in this way, when the stent 10 expands, stress concentrates on the drug-carrying portion 60, and accordingly, bending and distortion occur in the drug-carrying portion 60, so that the drug-carrying portion 60 can be prevented from peeling off or falling off from the stent 10.

[0041] The outer surface 19 of the stent 10 (or strut) means the surface of the stent 10 on the side where the porous structure 20 is disposed.

[0042] <Porous structure 20> As shown in FIGS. 2 to 8, the porous structure 20 includes a skeleton portion 30 arranged in a mesh shape, a plurality of void portions 40 partitioned by the skeleton portion 30, and a covering portion 50 including a polymer material arranged to cover at least a part of the outer surface of the porous structure 20.

[0043] As shown in FIG. 2, the porous structure 20 is arranged to cover the outer periphery of the stent 10 and has a cylindrical shape similar to the stent 10.

[0044] The porous structure 20 has a tip portion 20A arranged to cover the vicinity of the tip portion 10A of the stent 10 and a base portion 20B arranged to cover the vicinity of the base portion 10B of the stent 10.

[0045] It is preferable that the area of each of the plurality of void portions 40 included in the porous structure 20 is smaller than the gap between the rings 11 of the stent 10. By adopting such a configuration, the porous structure 20 can prevent the scattering of plaques and thrombi accompanying the expansion of the stent 10. On the other hand, it is preferable that the size of the void portion 40 is larger than the area of a single blood cell contained in the blood. Thereby, the blood cells can pass through the void portion 40. Further, thereby, the void portion 40 enables the porous structure 20 to be imparted with extensibility. Therefore, when the stent 10 expands, the porous structure 20 also elongates in the circumferential direction along with the expansion of the stent 10, and can exhibit good followability (extensibility) with respect to the expansion of the stent 10.

[0046] There are no particular restrictions on the pattern (shape in the development view) etc. of the skeleton portion 30 and the void portion 40 of the porous structure 20.

[0047] The porous structure 20 can be constituted by, for example, a knitted fabric (knit), a woven fabric (braid), a molded product (a member such as a film body in which cuts are formed), etc. When the porous structure 20 is a knitted fabric, the porous structure 20 can be constituted by weft knitting. By constituting the porous structure 20 by weft knitting, it is possible to suppress the axial length of the porous structure 20 from becoming shorter as the stent 10 expands. When the porous structure 20 is a woven fabric, the woven fabric is constituted by a known weaving method. Further, the porous structure 20 may be constituted by a molded product obtained by making holes in a cylindrical object formed by injection molding or the like. The size, shape, and number of the knitting stitches and weaving stitches of the porous structure 20 are not particularly limited as long as it is possible to prevent peripheral embolism during the expansion of the stent 10.

[0048] FIG. 4 is a perspective view showing an enlarged view of the vicinity of the proximal end portion 20B of the porous structure 20, and FIG. 5 is a view showing an enlarged view of a part of the cross section orthogonal to the axis of the proximal end portion 20B of the porous structure 20.

[0049] As shown in FIGS. 4 and 5, the covering portion 50 has concavo-convex portions 53 formed along the surface shape of the porous structure 20.

[0050] The biological implant 100 is protected by a covering portion 50 which covers at least a portion of the porous structure 20. This prevents the porous structure 20 (especially the void portion 40) from getting caught on the biological lumen (e.g., blood vessel) when the biological implant 100 moves within the biological lumen. Furthermore, because the covering portion 50 contains a polymer material, it prevents an excessive decrease in the flexibility of the porous structure 20 where the covering portion 50 is provided. Therefore, it is possible to prevent a decrease in the deliverability of the biological implant 100 or a decrease in the expandability of the porous structure 20 as a result of providing the covering portion 50.

[0051] Furthermore, the biological implant 100 has a coating portion 50 with uneven surfaces 53 formed along the surface shape of the porous structure 20. In other words, the coating portion 50 is not formed with a uniform thickness across all parts of the porous structure 20, and some parts of the coating portion 50 are thinner than other parts of the coating portion 50 (for example, the protrusions 51). As a result, the amount of coating portion 50 required (the volume of polymer material contained in the coating portion 50) is reduced compared to the case where the coating portion 50 is provided with a uniform thickness across the entire porous structure 20.

[0052] As shown in Figures 4 and 5, the covering portion 50 has a convex portion 53a that exhibits a convex shape at a position corresponding to the skeletal portion 30, and a recessed portion 53b that is composed of a smaller amount of covering than the convex portion 53a at a position corresponding to the void portion 40.

[0053] In the covering portion 50, the skeletal portion 30 and the void portion 40 are arranged alternately in the circumferential direction of the porous structure 20. Therefore, as shown in Figures 4 and 5, the covering portion 50 has convex portions 53a provided in the areas where the skeletal portion 30 is located and concave portions 53b provided in the areas where the void portion 40 is located, which are arranged alternately along the circumferential direction of the porous structure 20.

[0054] The porous structure 20 has a recess 53b with a small thickness formed in a part of the covering portion 50, thus reducing the amount of covering portion 50 that needs to be installed. Furthermore, because the porous structure 20 has a recess 53b formed in the covering portion 50, when the porous structure 20 expands in conjunction with the expansion of the stent 10, it is possible to prevent the expansion of the porous structure 20 from being hindered at the location where the recess 53b is formed. Therefore, by providing the recess 53b, the porous structure 20 can prevent the smooth expansion of the porous structure 20 from being hindered by the presence of the covering portion 50.

[0055] In particular, in this embodiment, as shown in Figures 4 and 5, the covering portion 50 has convex portions 53a and concave portions 53b arranged alternately in the circumferential direction. Therefore, when the porous structure 20 expands, each of the concave portions 53b located between adjacent convex portions 53a in the circumferential direction begins to expand quickly. Thus, when the porous structure 20 expands, it becomes possible to uniformly expand each part of the porous structure 20 in the circumferential direction.

[0056] As shown in Figures 2, 3, and 6, the covering portion 50 can be positioned at least one of the two ends 20A and 20B located in the axial direction (longitudinal direction) of the porous structure 20. Alternatively, the covering portion 50 can be positioned to cover at least one of the two ends located in the axial direction (longitudinal direction) of the porous structure 20 in the circumferential direction.

[0057] In this embodiment, the covering portion 50 is provided at both ends 20A and 20B of the porous structure 20. However, for example, the covering portion 50 may be provided only at the tip end 20A of the porous structure 20, or only at the base end 20B of the porous structure 20. Furthermore, the covering portion 50 may not be provided in a certain range from each end 20A and 20B of the porous structure 20 to the central part 20C (near the axial center of the porous structure 20). The following description will focus on the covering portion 50 provided at the base end 20B of the porous structure 20, but the covering portion 50 provided at the tip end 20A can be configured in the same way as the covering portion 50 provided at the base end 20B.

[0058] As shown in Figure 6, the amount of coverage of the covering portion 50 gradually decreases from the base end portion 20B located in the axial direction of the porous structure 20 to the central portion 20C. In other words, the covering portion 50 provided at the base end portion 20B of the porous structure 20 gradually decreases in thickness toward the tip portion 20A located on the opposite side in the axial direction. Similarly, the covering portion 50 provided at the tip portion 20A of the porous structure 20 gradually decreases in thickness toward the base end portion 20B located on the opposite side in the axial direction.

[0059] For example, as will be described later, one manufacturing method for the porous structure 20 involves preparing a porous structure longer than the length intended for use in the product and cutting it at a predetermined position. When such a manufacturing method is adopted, both ends 20A and 20B of the porous structure 20 are made up of the cut ends. Since the porous structure 20 has a mesh structure consisting of a skeletal part 30 and a void part 40, fraying and inconsistencies are likely to occur at the cut ends, which can result in a shape that spreads radially outward. When both ends 20A and 20B of the porous structure 20 are formed in this way, the porous structure 20 is more likely to get caught in biological tubular lumens or the like.

[0060] In this embodiment, by arranging the covering portion 50 at both ends 20A and 20B of the porous structure 20, it is possible to prevent the fraying and inconsistencies described above from occurring at both ends 20A and 20B. Furthermore, since the amount of covering portion 50 is gradually reduced from each end 20A and 20B located in the axial direction of the porous structure 20 toward the central part 20C, the amount of covering portion 50 can be reduced even more effectively. In addition, when moving the biological implant 100 within a biological lumen, both ends 20A and 20B in the axial direction of the porous structure 20 are prone to contact with the inner wall of the biological lumen, and are particularly likely to get caught on the porous structure 20. By covering at least one end of both ends 20A and 20B in the axial direction of the porous structure 20 with the covering portion 50, as in this embodiment, it becomes possible to effectively prevent the porous structure 20 from getting caught on the inner wall of the biological lumen, etc.

[0061] As shown in Figure 7, the porous structure 20 can be fixed to the stent 10 via the covering portion 50. By fixing the porous structure 20 to the stent 10 via the covering portion 50, it is possible to prevent fraying or inconsistencies in the porous structure 20 at the cutting position when the porous structure 20 is cut at the position where the covering portion 50 is provided (the application position of the covering material 50a) during the manufacturing of the porous structure 20 (see Figure 10).

[0062] The porous structure 20 can be fixed to the area of ​​the stent 10 where the drug-carrying portion 60 is not located, via the covering portion 50. By fixing the porous structure 20 to the stent 10 only in the area where the drug-carrying portion 60 is not located, the polymer material contained in the covering portion 50 can be used to properly fix the covering portion 50, the porous structure 20, and the stent 10 to each other.

[0063] Examples of areas where the drug-carrying portion 60 is not located include the curved portion 17 and / or the link portion 12. As mentioned above, stress concentration and strain are likely to occur in the curved portion 17 and / or the link portion 12 when the stent 10 is expanded, so it is not preferable to place the drug-carrying portion 60 there (see Figure 7). Therefore, it is preferable to fix the porous structure 20 to the curved portion 17 and the link portion 12, where the drug-carrying portion 60 is not provided, via the covering portion 50. In particular, since the number of curved portions 17 provided in a single stent 10 is greater than the number of link portions 12 (see Figure 3), it is more preferable to fix the porous structure 20 to at least the curved portion 17 via the covering portion 50, from the viewpoint of improving the fixing force of the porous structure 20 to the stent 10.

[0064] As described above, in this embodiment, the covering portion 50 is positioned only near both ends 20A and 20B of the porous structure 20. Therefore, for example, the porous structure 20 can be fixed to the stent 10 via the covering portion 50 at the curved portions 17 and / or link portions 12 located at both ends 20A and 20B.

[0065] The polymer material contained in the covering portion 50 is, for example, a biodegradable polymer. By configuring the covering portion 50 in this way, the covering portion 50 can be decomposed and disappear after a predetermined period of time has elapsed since the stent 10 was implanted. Furthermore, by using a biodegradable polymer as the polymer material, the impact on the human body caused by implanting the covering portion 50 together with the stent 10 and the porous structure 20 can be reduced.

[0066] The polymer material contained in the covering portion 50 has physical properties such as a Young's modulus less than 10 MPa. By configuring the covering portion 50 in this way, it is possible to prevent the smooth expansion of the porous structure 20 from being hindered by the covering portion 50 when the stent 10 expands.

[0067] (Method for manufacturing the porous structure 20 according to the embodiment) Next, a method for manufacturing the porous structure 20 according to the embodiment will be described.

[0068] To begin manufacturing the porous structure 20, a porous structure 20 having a predetermined length in the axial direction is prepared, as shown in Figure 9.

[0069] Next, a predetermined core metal 500A is inserted through the porous structure 20.

[0070] Next, a coating material 50a containing a polymer material is applied to any portion of the outer surface of the porous structure 20, thereby forming a coating portion 50 having irregularities that conform to the surface shape of the porous structure 20 on at least a part of the porous structure 20.

[0071] The above-mentioned coating material 50a is obtained by dissolving the main material of the coating material 50a in a predetermined solvent.

[0072] The coating material 50a can be applied to the porous structure 20 using a predetermined nozzle 400A. The viscosity of the coating material 50a can be adjusted such that, for example, as it is applied to the porous structure 20 by dropping from the nozzle 400A, the coating material 50a flows along the skeletal portion 30 of the porous structure 20 and naturally penetrates into the inside of the voids 40. By adjusting the viscosity of the coating material 50a in this way, it becomes possible to form uneven portions 53 (see Figures 5 and 6) that conform to the surface shape of the porous structure 20 without applying any pressure to the porous structure 20 when dropping the coating material 50a from the nozzle 400A onto the porous structure 20.

[0073] Furthermore, when the coating material 50a, whose viscosity has been adjusted as described above, is applied to the porous structure 20, the coating material 50a flows from the application point toward both ends of the porous structure 20, and the amount of coating material 50a applied decreases as it moves axially away from the application point. As a result, it becomes possible to form a coating portion 50 whose thickness decreases as it moves from both ends 20A and 20B toward the axial center (see Figure 6).

[0074] The viscosity of the coating material 50a can be adjusted, for example, by changing the concentration of the main material contained in the coating material 50a. The specific materials (solvent material and main material) and physical properties of the coating material 50a will be described later.

[0075] After applying the coating material 50a to the porous structure 20, the coating material 50a is dried. Once the coating material 50a has dried and solidified, and the coated portion 50 has been formed, multiple porous structures 20 can be obtained by cutting the porous structure 20 at the position where the coated portion 50 has been formed, as shown in Figure 10. The area near the cutting position of the porous structure 20 constitutes one of the ends 20A, 20B of the porous structure 20.

[0076] A covering portion 50 is formed at the cutting position of the porous structure 20. Therefore, the skeletal portion 30 near the cutting position of the porous structure 20 is covered on the surface by the polymer material contained in the covering portion 50. Thus, when the porous structure 20 is cut, fraying and inconsistencies in the skeletal portion 30 located at the cutting position can be prevented.

[0077] (Modifications) Next, a biological implant relating to a modification of the embodiment described above will be explained. In the explanation of the modification, explanations that overlap with those described in the embodiment described above will be omitted as appropriate. Unless otherwise specified, the biological implant relating to the modification can incorporate the same components as those described in the embodiment described above.

[0078] The porous structure 20 of the biological implant 100 according to the modified example has a different configuration of the covering portion 50A compared to the covering portion 50 according to the above-described embodiment.

[0079] The covering portion 50 of the porous structure 20 according to the above embodiment includes an uneven portion 53 formed along the surface shape of the porous structure 20 (see Figures 4 to 6). On the other hand, as shown in Figures 11 to 13, the covering portion 50A according to this modified example is arranged only along the skeletal portion 30 surrounding the void portion 40, and is not arranged at a position corresponding to the void portion 40.

[0080] As shown in Figures 12 and 13, the covering portion 50A is configured to form a layer that covers the outer surface of the skeletal portion 30.

[0081] In the modified porous structure 20, the covering portion 50A is not formed in the location where the void portion 40 exists. In other words, unlike the embodiment described above, there is no recess 53b in the covering portion 50 that is arranged to fill the void portion 40. Therefore, the amount of covering portion 50A to be mounted is further reduced compared to the covering portion 50 in the embodiment described above. Also, since there is no recess 53b that is arranged to fill the void portion 40 as described above, the flexibility of the porous structure 20 in the location corresponding to the void portion 40 can be improved. Furthermore, the bulkiness caused by providing the covering portion 50 when it is crimped onto the stent 10 (the state in which the biological implant 100 is formed) can be reduced. In addition to these, if the stent 10 is configured to include a drug carrying portion 60, the absence of the recess 53b in the covering portion 50 that is arranged to fill the void portion 40 can improve the drug permeability of the stent 10 (the flow of drug between the outer surface 19 of the stent 10 and the inner wall of the biological lumen).

[0082] As shown in Figure 13, in the modified porous structure 20, the amount of covering portion 50A gradually decreases from the end located in the axial direction of the porous structure 20 (the base end portion 20B in the illustrated example) to the central portion (towards the direction of arrow X1 in the figure). By configuring it in this way, as explained in the above embodiment, it is possible to reduce the amount of covering portion 50A installed while preventing fraying and inconsistencies from occurring at the ends of the porous structure 20.

[0083] In the modified example, an example was described in which a covering portion 50A is formed on the base end portion 20B of the porous structure 20. However, it is also possible to adopt a configuration in which the covering portion 50A is provided on the tip portion 20A and / or the base end portion 20B of the porous structure 20. Furthermore, although not explained here, the covering portion 50A can be configured to fix the porous structure 20 to the stent 10 at a predetermined position (for example, in an area where the drug-carrying portion 60 is not formed), similar to the embodiment described above.

[0084] (Method for manufacturing the porous structure 20 according to a modified example) Next, a method for manufacturing the porous structure 20 according to a modified example will be described.

[0085] To begin manufacturing the porous structure 20, a porous structure 20 having a predetermined length in the axial direction is prepared, as shown in Figure 14.

[0086] Next, the porous structure 20 is set in a predetermined jig 500B. The jig 500B used here can be a stepped core metal in which a small diameter section (for example, an outer diameter of 1.0 mm) is provided at the position where the coating material 50a is applied to the porous structure 20, and large diameter sections (for example, an outer diameter of 2.0 mm) are provided on both sides of the small diameter section.

[0087] Next, a coating material 50a containing a polymer material is applied to any part of the outer surface of the porous structure 20, forming a coating portion 50A on at least a part of the porous structure 20 that is positioned only along the skeletal portion 30 surrounding the void portion 40.

[0088] The coating material 50a can be sprayed onto the porous structure 20 using a predetermined sprayer 400B. When spraying the coating material 50a onto the porous structure 20 from the sprayer 400B, the coating material 50a can be applied along the outer surface of the porous structure 20 by rotating the jig 500B. There are no particular restrictions on the specific type of sprayer 400B, but for example, a known ultrasonic sprayer can be used.

[0089] When applying the coating material with the sprayer 400B, the coating material 50a can be applied to a predetermined position on the porous structure 20 in such a way that it covers only the skeletal portion 30 (i.e., the coating material 50a does not fill the void portion 40) by adjusting the application conditions such as the viscosity of the coating material 50a, the amount sprayed (air flow rate), the application time, and the rotation speed of the jig 500B.

[0090] Furthermore, by adjusting the coating conditions as described above, the coating material 50a flows from the coating position where it is applied from the sprayer 400B toward both ends of the porous structure 20, and the amount of coating material 50a applied decreases as it moves axially away from the coating position. As a result, it becomes possible to form a coating portion 50A whose thickness decreases as it moves toward the axial center from both ends 20A, 20B (see Figure 13).

[0091] After applying the coating material 50a to the porous structure 20, the coating material 50a is dried. Once the coating material 50a has dried and solidified, and the coated portion 50A has been formed, multiple porous structures 20 can be obtained by cutting the porous structure 20 at the position where the coated portion 50A has been formed, as shown in Figure 15. The area near the cutting position of the porous structure 20 constitutes one of the ends 20A, 20B of the porous structure 20.

[0092] A covering portion 50A is formed at the cutting position of the porous structure 20. Therefore, the skeletal portion 30 near the cutting position of the porous structure 20 is covered on the surface by the polymer material contained in the covering portion 50A. Thus, when the porous structure 20 is cut, fraying and inconsistencies in the skeletal portion 30 located at the cutting position can be prevented.

[0093] (Examples of arrangement of the covering portion 50) In the embodiments and modifications described above, the tip portion 20A and / or base portion 20B were exemplified as positions for forming the covering portion 50 (or covering portion 50A) on the porous structure 20. However, there are no particular restrictions on the position where the covering portion 50 is provided in the biological implant according to the present invention.

[0094] For example, as shown in Figure 16, the covering portion 50 (or covering portion 50A) may be formed over the entire axial length of the porous structure 20. Alternatively, as shown in Figure 17, for example, the covering portion 50 (or covering portion 50A) may be arranged to extend linearly in the axial direction, connecting the tip portion 20A, the base portion 20B, and each end portion 20A, 20B of the porous structure 20. Alternatively, as shown in Figure 18, for example, the covering portion 50 (or covering portion 50A) may be arranged to extend spirally in the circumferential direction of the porous structure 20, connecting the tip portion 20A, the base portion 20B, and each end portion 20A, 20B of the porous structure 20. Alternatively, as shown in Figure 19, for example, the covering portion 50 (or covering portion 50A) may be configured to include multiple portions arranged with gaps in the circumferential direction at each end portion 20A, 20B of the porous structure 20.

[0095] The following describes a preferred example of the materials and physical properties of each part (stent, porous structure, covering part, etc.) that constitute the biological implant 100.

[0096] <Intra-biological implant> The intra-biological implant according to the present invention comprises an expandable cylindrical stent and a porous structure disposed to cover the stent and configured to expand in accordance with the expansion of the stent's diameter. The porous structure has a mesh-like skeletal structure, a plurality of voids partitioned by the skeletal structure, and a covering portion made of a polymer material disposed to cover at least a part of the outer surface of the porous structure. The Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

[0097] [Stent] The stent may be made of any of the following materials, and may be made of non-biodegradable or biodegradable materials.

[0098] <Non-biodegradable materials> Non-biodegradable materials that can be used for stents include carbon fibers, metal materials, and non-biodegradable resin materials. Preferably, the non-biodegradable material is a metal material or a non-biodegradable resin material. From the viewpoint of further reducing inflammation, it is particularly preferable that the non-biodegradable material is a metal material.

[0099] 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.

[0100] Furthermore, if the stent is made of a non-biodegradable resin material, there are no particular limitations, and non-biodegradable resin materials commonly used in the stent product field can be used. Specifically, examples include polyolefins such as polyethylene and polypropylene, aromatic polyesters such as polyethylene terephthalate, cellulosic polymers such as cellulose acetate and cellulose nitrate, and fluorine-containing polymers such as polytetrafluoroethylene and tetrafluoroethylene-ethylene copolymers.

[0101] <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.

[0102] 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), 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-derived polycarbonate.

[0103] 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, DL-lactic acid / ε-caprolactone copolymer (PDLLA-r-PCL, PDLLA-b-PCL, PDLLA-alt-PCL), L-lactic acid / ε-caprolactone copolymer (PLLA-r-PCL, PLLA-b-PCL, PLLA-alt-PCL), PLLA-r-PTMC (TMC = trimethylene carbonate), PLLA-r-PDTC (DCT = 2,2-dimethyltrimethylene carbonate), PLLA-b-PTMC, PLLA-b-PDTC, PLGA (poly(lactide-co-glycolide), polyanhydride, polyorthoester, poly(N-(2-hydroxypropyl)methacrylamide), DLPLA-poly(dl-lactide), LPLA-poly(l-lactide), PGA-polyglycolide, PDO-poly(dioxanone), PGA-TMC-poly(glycolide-co-trimethylene) Examples include carbonates, PGA-LPLA-poly(l-lactide-co-glycolide), PGA-DLPLA-poly(dl-lactide-co-glycolide), LPLA-DLPLA-poly(l-lactide-co-dl-lactide), and PDO-PGA-TMC-poly(glycolide-co-trimethylene carbonate-co-dioxanone), and PAE (Polyanhydride esters)-Salicylate (for example, polymers in which salicylic acid is chemically introduced into the polymer main chain, such as polylactide anhydride or polyadipic acid bonded to both ends).

[0104] 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.

[0105] 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.

[0106] 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.

[0107] 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.

[0108] Stents can be suitably formed from materials appropriately selected from the non-biodegradable and biodegradable resin 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 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. Furthermore, if the resin material is biodegradable, the stent will disappear by biodegradation after a predetermined period, thus reducing the potential impact on the body lumen after treatment.

[0109] 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.

[0110] <Drug-carrying section> In a stent, the drug-carrying section may contain a drug in addition to the non-biodegradable resin material and biodegradable resin material described above. The drug contained in the drug-carrying section is not particularly limited and can be appropriately selected according to the desired application. Considering the induction of inflammation by the porous structure, the drug preferably contains an anti-inflammatory agent or an immunosuppressant. The anti-inflammatory agent is not particularly limited and known anti-inflammatory agents can be used. The anti-inflammatory agent is a steroidal anti-inflammatory agent, a non-steroidal anti-inflammatory agent, or a combination thereof. For example, anti-inflammatory drugs include alclofenac, alclometasone dipropionate, algestone acetonide, alpha-amylase, amsinafar, amsinafid, amfenac sodium, ampyrrose hydrochloride, anakinra, anirolac, anitrazafen, apazon, valsalazid disodium, bendazac, benoxaprofen, benzydamine hydrochloride, bromelain, budesonide, carprofen, cycloprofen, syntazone, criprofen, clobetasol propionate, clobetazone butyrate, clopirac, cloticasone propionate, colmetasone acetate, cortodoxone, deflazacort, desonide, desoxymethasone, dexamethasone, dexamethasone dipropionate, diclofenac potassium, diclofenac sodium, diflorazone diacetate, diflumidone sodium, diflunisal, and difluprednate. Diphthalone, Dimethyl sulfoxide, Drocinonide, Endrizone, Enlimomab, Enolicam sodium, Epirizol, Etodrug, Etofenamate, Felbinac, Fenamol, Fenbufen, Fenclofenac, Fenclolac, Fendozal, Fenpiparone, Fenthiazac, Flazaron, Fluazacort, Flufenamic acid, Flumizole, Flunisolid acetate, Flunixin, Flunixin meglumine, Fluocortin butyl, Fluorometholone acetate, Fluquazone, Flurbiprofen, Fluretofen, Fluticasone propionate, Flaprofen, Flobufen, Halcinonide, Halobetasol propionate, Halopredone acetate, Ibufenac, Ibuprofen, Ibuprofen aluminum, Ibuprofen piconol, Ilonidap, Indomethacin, Indomethacin sodium, Indoprofen,Indoxol, Intrazol, Isoflupredone acetate, Isoxepak, Isoxicam, Ketoprofen, Lofemisole hydrochloride, Romoxicam, Loteprednol etavonate, Meclofenamete sodium, Meclofenamic acid, Mechlorizone dibutyrate, Mefenamic acid, Mesalamine, Mesecrazone, Prednisolone, Methylprednisolone sleptanoate, Momiflumate, Nabumeton, Naproxen, Naproxen sodium, Naproxol, Nimazon, Orsalazine sodium, Orgotein, Orpanoxin, Oxaprozin, Oxyfenbutazone, Paraniline hydrochloride, Pentosan sodium polysulfate, Fenbutazone sodium glycerate, Pirfenidone, Piroxicam, Piroxicam cinnamate, Piroxicam olamine, Pirprofen, Prednazate, Prif Examples include feron, prodolic acid, proquazone, proxazole, proxazole citrate, rimexolone, romazalit, sarcorex, sarnasedin, sarsalate, sanguinalium chloride, secrazone, selmethacin, sudoxicam, sulindac, suprofen, talmetacin, talniflumate, talosalate, tebuferon, tenidap, tenidap sodium, tenoxicam, tesicam, tesimide, tetridamine, thiopinac, thixocortol pivalate, tolmetin, tolmetin sodium, triclonide, triflumidate, didomemethacin, zomepirac sodium, aspirin (acetylsalicylic acid), salicylic acid, corticosteroids, glucocorticoids, tacrolimus, pimecorlimus, their prodrugs, and their codrugs. The above anti-inflammatory agents may be used individually or in combination of two or more.

[0111] The immunosuppressants are also not particularly limited, and known immunosuppressants can be used. Examples include sirolimus, everolimus, biolimus A9, pimecrolimus, zotarolimus, sirolimus derivatives such as ABT-578, biolimus (e.g., biolimus A9®), tacrolimus, azathioprine, cyclosporine, cyclophosphamide, mycophenolate mofetil, and gusperimus. The above immunosuppressants may be used individually or in combination of two or more. Furthermore, the above immunosuppressants may be used in combination with anti-inflammatory agents.

[0112] [Porous Structure] In an in-vivo implant according to one embodiment, the material constituting the skeletal part of the porous structure can be the non-biodegradable material and biodegradable material described in the [Stent] section above. However, since the in-vivo implant is implanted in the body, it is preferably a biodegradable resin material, more preferably polyglycolic acid (PGA) or lactic acid-glycolic acid copolymer (PLGA), and even more preferably polyglycolic acid (PGA).

[0113] The weight-average molecular weight of the resin material used in the skeletal portion of the porous structure according to one embodiment is not particularly limited, but is preferably 10,000 or more, more preferably 10,000 to 1,000,000, and even more preferably 20,000 to 500,000. Methods for measuring the weight-average molecular weight include gel permeation chromatography (GPC), light scattering, viscosity measurement, and mass spectrometry (TOFMASS, etc.). In this specification, the weight-average molecular weight is the value measured by GPC using polystyrene as a standard substance.

[0114] <Coating portion> The coating portion according to the present invention is arranged to cover at least a part of the outer surface of the porous structure, and includes a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

[0115] (Polymer material) In the present invention, the polymer material is included in the covering portion, which is arranged to cover at least a part of the outer surface of the porous structure. The Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa. Because the Young's modulus of the polymer material is within this range, the polymer material is easily deformed, which reduces the likelihood of expansion defects in the covering portion and allows the stent to be sufficiently expanded. The Young's modulus can be measured by the method described in the examples.

[0116] The mass ratio of the polymer material to the total mass of the coating ((mass of polymer material / total mass of the coating) × 100 (mass%)) is preferably 80% by mass or more, more preferably 90% by mass or more, even more preferably 95% by mass or more, particularly preferably 99% by mass or more, and most preferably 100% by mass.

[0117] (Young's Modulus) Young's modulus (elastic modulus) is an index used to evaluate the mechanical properties of polymer materials, and it shows the relationship between stress and strain in response to an external force applied to the material. Specifically, Young's modulus is defined as the value obtained by dividing stress by strain, and its unit is expressed in Pa (Pascal). When Young's modulus is high, the material is rigid and tends to be resistant to deformation under external force. Conversely, when Young's modulus is low, the material is flexible and easily deformed by external force.

[0118] The Young's modulus of the polymer material according to the present invention is greater than 0 MPa and less than 200 MPa, but in one embodiment, the Young's modulus may be greater than 0 MPa and 100 MPa or less, greater than 0 MPa and 50 MPa or less, greater than 0 MPa and 25 MPa or less, greater than 0 MPa and 10 MPa or less, greater than 0 MPa and 5 MPa or less, 0.1 MPa or more and 200 MPa or less, 0.1 MPa or more and 100 MPa or less, 0.1 MPa or more and 50 MPa or less, 0.1 MPa or more and 25 MPa or less, 0.1 MPa or more and 10 MPa or less, 0.1 MPa or more and 5 MPa or less, 0.5 MPa or more and 200 MPa or less, 0.5 MPa or more and 100 MPa or less, 0.5 The Young's modulus of the polymer material may be between MPa and 50 MPa, 0.5 MPa and 25 MPa, 0.5 MPa and 10 MPa, 0.5 MPa and 5 MPa, 1 MPa and 200 MPa, 1 MPa and 100 MPa, 1 MPa and 50 MPa, 1 MPa and 25 MPa, 1 MPa and 10 MPa, 1 MPa and 5 MPa, 2 MPa and 200 MPa, 2 MPa and 100 MPa, 2 MPa and 50 MPa, 2 MPa and 25 MPa, 2 MPa and 10 MPa, 2 MPa and 5 MPa, 5 MPa and 100 MPa, 5 MPa and 50 MPa, or 5 MPa and 10 MPa. By having the Young's modulus of the polymer material within the above range, the flexibility of the polymer material becomes more suitable, expansion defects of the covered portion become less likely, and the stent can be expanded more sufficiently.

[0119] (Elongation at Break) The elongation at break of the polymer material according to one embodiment of the present invention may be 960% or more and 3500% or less, 960% or more and 3250% or less, 960% or more and 3000% or less, 960% or more and 2500% or less, 960% or more and 2000% or less, 1000% or more and 3500% or less, 1000% or more and 3250% or less, 1000% or more and 3000% or less, 1000% or more and 2500% or less, 1000% or more and 2000% or less, 1500% or more and 3500% or less, 1500% or more and 3250% or less, 1500% or more and 3000% or less, 1500% or more and 2500% or less, 1500% or more and 2000% or less, 1000% or more and 1500% or less, or 1000% or more and 1250% or less. Because the elongation at break of the polymer material falls within the above range, expansion defects of the coating are less likely to occur, and the ductility and deformation resistance of the polymer material are also improved, thus suppressing the rupture of the coating that occurs during stent expansion.

[0120] Here, "elongation at break" is one of the mechanical properties that indicates how much a polymer material can be stretched before it breaks, and it can be used to evaluate the ductility and deformation resistance of a polymer material. Elongation at break is usually expressed as the percentage of the final elongation relative to the initial length when a tensile force is applied to the material, and can be measured specifically by the method shown in the examples below.

[0121] (Biodegradable Polymer) The polymer material according to one embodiment of the present invention preferably contains a biodegradable polymer. A biodegradable polymer is a resin that is biodegradable. Since the in vivo implant according to the present invention is implanted in a living body, it is preferable that the polymer material contains a biodegradable polymer so that the covering portion can be decomposed in the living body. The biodegradable polymer content in the polymer material ((mass of biodegradable polymer / total mass of polymer material) × 100 (mass%)) is preferably 80% by mass or more, more preferably 90% by mass or more, even more preferably 95% by mass or more, particularly preferably 99% by mass or more, and most preferably 100% by mass, from the viewpoint of the biodegradability of the polymer material.

[0122] The biodegradable polymer is not particularly limited, but any of the compounds listed in the section on biodegradable resin materials above can be used as appropriate. The biodegradable polymer may be, for example, at least one selected from the group consisting of polybutylene succinate (PBS), polybutylene adipate terephthalate (PBAT), polylactic acid (PLA), polyvinyl alcohol (PVA), polyglycolic acid (PGA), poly(ε-caprolactone) (PCL), polybutylene succinate-co-adipate (PBSA), polybutylene adipate terephthalate (PBAT), polyethylene terephthalate succinate (PETS), PBAT / PLA compound, starch polyester resin, cellulose acetate, cellulose, polyhydroxyalkanoic acid (PHA), and 3-hydroxybutyrate-co-3-hydroxyhexanoate polymer (PHBH).

[0123] Furthermore, from the viewpoint of having excellent biodegradability and flexibility, it is preferable that the monomer constituting the biodegradable polymer includes one or more selected from the group consisting of lactic acid, glycolic acid, and ε-caprolactone. In other words, from the viewpoint of having excellent biodegradability and flexibility, it is preferable that the biodegradable polymer includes one or more selected from the group consisting of polylactic acid, polyglycolic acid, poly(ε-caprolactone), copolymer of lactic acid and ε-caprolactone, copolymer of lactic acid and glycolic acid, and copolymer of glycolic acid and ε-caprolactone.

[0124] In the polymer material according to one embodiment of the present invention, the biodegradable polymer preferably contains a polylactic acid resin, from the viewpoint of having biodegradability while also possessing desirable flexibility. In this specification, a polylactic acid resin refers to a resin in which 40 mol% or more (up to 100 mol%) of lactic acid monomer is present in the total monomers constituting the resin, preferably 45 mol% or more.

[0125] The polylactic acid resin may be polylactic acid, but from the viewpoint of superior biodegradability and flexibility, it may also be a copolymer of a lactic acid monomer and a monomer of a hydroxycarboxylic acid other than lactic acid. Furthermore, the polylactic acid resin may contain a small amount of chain extender residues. The content of the polylactic acid resin in the biodegradable polymer ((mass of polylactic acid resin / total mass of biodegradable polymer) × 100 (mass%)) is preferably 80% by mass or more, more preferably 90% by mass or more, even more preferably 95% by mass or more, particularly preferably 99% by mass or more, and most preferably 100% by mass.

[0126] Examples of lactic acid include L-lactic acid, D-lactic acid, and DL-lactic acid. Examples of hydroxycarboxylic acids other than lactic acid include difunctional aliphatic hydroxycarboxylic acids such as glycolic acid, 3-hydroxybutyric acid, 4-hydroxybutyric acid, 2-hydroxybutyric acid, 2-hydroxy-3,3-dimethylbutyric acid, 2-hydroxy-3-methylbutyric acid, 2-methyllactic acid, and 2-hydroxycaproic acid, as well as lactones such as caprolactone, butyrolactone, and valerolactone. These may be used individually or in combination of two or more.

[0127] When the polylactic acid resin is a copolymer of lactic acid and a hydroxycarboxylic acid other than lactic acid, the ratio (molar ratio, moles of lactic acid : moles of hydroxycarboxylic acid other than lactic acid) in the polylactic acid resin is preferably 0.1 to 100:1, more preferably 0.5 to 50:1, even more preferably 0.5 to 20:1, and particularly preferably 0.5 to 8:1.

[0128] In one embodiment, the polylactic acid-based resin contained in the polymer material is a copolymer of lactic acid and a hydroxycarboxylic acid other than lactic acid. In this case, it is preferable that the lactic acid is L-lactic acid and the hydroxycarboxylic acid other than lactic acid is ε-caprolactone. By including a polylactic acid-based resin having such a configuration in the polymer material, the flexibility of the polymer material becomes more suitable, and the stent can be expanded more sufficiently in the radial direction.

[0129] When the polylactic acid resin is a copolymer of L-lactic acid and ε-caprolactone, the ratio (molar ratio, number of moles of L-lactic acid : number of moles of ε-caprolactone) of L-lactic acid to ε-caprolactone in the polylactic acid resin is preferably 0.1 to 10:1, more preferably 0.5 to 5:1, and even more preferably 0.5 to 2:1.

[0130] In one embodiment of the polymer material, the ratio (mass%) of the copolymer of L-lactic acid and ε-caprolactone to the total mass of the polymer material may be greater than 25% by mass and 100% by mass or less, 30% by mass or more and 100% by mass or less, 40% by mass or more and 100% by mass or less, 50% by mass or more and 100% by mass or less, greater than 50% by mass and 100% by mass or less, 60% by mass or more and 100% by mass or less, 70% by mass or more and 100% by mass or less, 30% by mass or more and 90% by mass or less, 40% by mass The mass percentage of the copolymer may be % or more and 90% or less by mass, 50% or more and 90% or less by mass, more than 50% and 90% or less by mass, 60% or more and 90% or less by mass, 70% or more and 90% or less by mass, 30% or more and 80% or less by mass, 40% or more and 80% or less by mass, 50% or more and 80% or less by mass, more than 50% and 80% or less by mass, 60% or more and 80% or less by mass, 70% or more and 80% or less by mass, 40% or more and 70% or less by mass, or 40% or more and 60% or less by mass. By having the mass percentage of the copolymer within the above range, the flexibility and ductility of the polymer material are improved, the occurrence of breakage of the coating is more sufficiently suppressed, and stent expansion defects are also more sufficiently suppressed.

[0131] A polymer material according to one embodiment may contain two or more types of polylactic acid resins. Hereinafter, when two types of polylactic acid resins are included, one of the polylactic acid resins will be referred to as the first polylactic acid resin, and the other as the second polylactic acid resin.

[0132] The first polylactic acid resin and the second polylactic acid resin can be appropriately selected from the polylactic acid resins described above, but it is preferable that the first polylactic acid resin contains L-lactic acid as a monomer and the second polylactic acid resin contains DL-lactic acid. More preferably, the first polylactic acid resin is a copolymer of L-lactic acid and a hydroxycarboxylic acid other than lactic acid, and the second polylactic acid resin is a copolymer of DL-lactic acid and a hydroxycarboxylic acid other than lactic acid. Even more preferably, the first polylactic acid resin is a copolymer of L-lactic acid and ε-caprolactone, and the second polylactic acid resin is a copolymer of DL-lactic acid and ε-caprolactone.

[0133] In the first polylactic acid resin, the ratio of L-lactic acid to ε-caprolactone in the polylactic acid resin can be the same as the ratio used when the polylactic acid resin is a copolymer of L-lactic acid and ε-caprolactone. Furthermore, the ratio (molar ratio, moles of DL-lactic acid : moles of ε-caprolactone) of DL-lactic acid to ε-caprolactone in the second polylactic acid resin is preferably 1 to 20:1, more preferably 5 to 10:1, and even more preferably 8 to 10:1.

[0134] In the polymer material according to one embodiment, the mass ratio (mass of the first polylactic acid resin:mass of the second polylactic acid resin) of the first polylactic acid resin to the second polylactic acid resin is more preferably 0.5 to 10:1, more preferably 0.5 to 8:1, even more preferably 0.5 to 5:1, particularly preferably 1 to 5:1, and most preferably 2 to 4:1. By having this mass ratio within the above range, the flexibility and ductility of the polymer material are improved, the occurrence of breakage of the coating portion is more sufficiently suppressed, and stent expansion defects are also more sufficiently suppressed.

[0135] Furthermore, the total content of the first polylactic acid resin and the second polylactic acid resin in the polymer material ((mass of the first polylactic acid resin and the second polylactic acid resin / total mass of the polymer material) × 100 (mass%) is preferably 80% by mass or more, more preferably 90% by mass or more, even more preferably 95% by mass or more, particularly preferably 99% by mass or more, and most preferably 100% by mass.

[0136] (Other Components) The coating portion may contain other components in addition to the polymer material described above, as long as the effects of the present invention are not impaired. Other components are not particularly limited, and examples include, when the medical device is intended for insertion into a body cavity or lumen, drugs (bioactive substances) such as 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 substances. The amount of other components added is not particularly limited, and the amount normally used is applied in the same manner. 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.

[0137] [Method for Manufacturing a Porous Structure] One embodiment of the present invention is a method for manufacturing a porous structure used in an in-vivo implant, comprising: a step of preparing a porous structure (porous structure preparation step) for a stent that is expanded in diameter and a porous structure that is arranged to cover the stent and is configured to expand as the diameter of the stent expands; and a step of applying a coating material containing a polymer material to an arbitrary portion of the outer surface of the porous structure to form a coating portion on at least a part of the outer surface of the porous structure (coating portion formation step), wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa. An in-vivo implant covered with a stent using a porous structure manufactured by this method will be able to sufficiently expand the stent in the radial direction while preventing the expansion of the porous structure in the radial direction.

[0138] (I) Porous structure preparation process The porous structure preparation process is a process for preparing a porous structure to be used in an in-vivo implant, comprising an expandable cylindrical stent and a porous structure that is positioned to cover the stent and is configured to expand in accordance with the expansion of the stent's diameter.

[0139] In this process, a commercially available porous structure may be prepared, or a manufactured porous structure may be prepared. The method for manufacturing the porous structure is not particularly limited, as long as a porous structure having the configuration described in the [Porous Structure] section above can be manufactured. As a manufacturing method, for example, the method described in International Publication No. 2008 / 062414 can be applied in a similar manner or with appropriate modifications.

[0140] (II) Coating Formation Process The coating formation process is a process of forming a coating on at least a part of the outer surface of a porous structure by applying a coating material containing a polymer material to any part of the outer surface of the skeletal part of the porous structure. Here, the polymer material can be any of those listed in the (polymer material) section of the above-mentioned <coating>.

[0141] The method for applying (coating) a coating material containing a polymer material to any part of the outer surface of a porous structure 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 spraying method is preferred as the application method because it is easy to form a coating even on fine structures such as the voids of the porous structure and is easy to fill the voids of the porous structure with the coating material.

[0142] The conditions for the coating material application method using a nozzle are not particularly limited. For example, the discharge pressure is preferably 1 to 10 kPa, and more preferably 2 to 5 kPa. The discharge time is not particularly limited, but is preferably 0.5 to 10 seconds, and more preferably 1 to 8 seconds.

[0143] Furthermore, it is preferable to rotate the porous structure circumferentially while the coating material containing the polymer material is being sprayed onto the porous structure. The rotation speed in the circumferential direction is preferably, for example, 5 to 30 rpm, and more preferably 10 to 20 rpm.

[0144] Coating materials containing polymer materials are preferably applied (coated) in a solution containing the coating material (hereinafter referred to as the coating solution). The solvent for the coating solution containing the coating material is not particularly limited as long as it can sufficiently dissolve or disperse the polymer material. Specifically, examples include 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, propionnitrile, and benzonitrile; aromatics such as benzene and toluene; amides such as N,N-dimethylformamide (DMF); sulfoxides such as dimethyl sulfoxide; and are not limited to these. These may be used individually or in combination of two or more.

[0145] In particular, from the viewpoint of uniformly dissolving polymer materials, solvents such as ketones like acetone, ethers like tetrahydrofuran, and nitriles like acetonitrile are preferred. Furthermore, from the same viewpoint, a preferred embodiment includes at least one solvent selected from the group consisting of acetone, tetrahydrofuran, and acetonitrile, and a more preferred embodiment includes acetone.

[0146] The coating solution is prepared by mixing polymer material, a solvent, and other components added as needed to create a copolymer solution. The order and method of adding the components are not particularly limited. The components can be added to a mixing container all at once or separately, in stages or continuously. The mixing method is also not particularly limited, and known methods can be used. A preferred method for preparing the coating solution involves adding the copolymer to a good solvent and stirring in the good solvent. The stirring method is not particularly limited as long as it can uniformly mix the copolymer solution.

[0147] The concentration of the polymer material in the coating solution is not particularly limited. From the viewpoint of being able to sufficiently dissolve or disperse the polymer material in the solvent, the concentration of the copolymer in the solution is preferably 0.1 to 1,000 mg / mL, more preferably 1 to 500 mg / mL, and particularly preferably 100 to 300 mg / mL. The amount of coating solution to be applied is not particularly limited, but it may be an amount sufficient to fill all the voids in the area to be coated with the coating solution.

[0148] 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.

[0149] Furthermore, by drying the applied coating solution, a coating is formed on at least a portion of the outer 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, and more preferably 80°C to 150°C. The drying time is preferably 30 minutes to 24 hours, more preferably 1 hour to 12 hours, and even more preferably 1 hour to 6 hours. For example, if the heating temperature is 80°C to 150°C, the heating time may be 1 hour to 6 hours.

[0150] [Method for Manufacturing Intraviviparous Devices] An intraviviparous device according to one embodiment of the present invention can be manufactured by attaching (crimping) a porous structure, manufactured by the above-described method for manufacturing porous structures, to an expandable cylindrical stent. That is, the method for manufacturing an intraviviparous device according to one embodiment of the present invention includes an attachment step of attaching a porous structure, manufactured by the above-described method for manufacturing porous structures, to an expandable cylindrical stent. With this manufacturing method, the intraviviparous device manufactured by this method can sufficiently expand the stent in the radial direction while preventing the expansion of the porous structure in the radial direction.

[0151] 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.

[0152] 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 the covering portion with laser light or the like to fix the porous structure to the stent.

[0153] 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.

[0154] The effects of the present invention will be explained using the following examples and comparative examples. However, the technical scope of the present invention is not limited to the following examples. In the following examples, unless otherwise specified, the operations were carried out at room temperature (25°C). Unless otherwise specified, "%" and "parts" mean "mass%" and "parts by mass," respectively.

[0155] <Preparation of test films> Test films for Examples 1 and 2 and Comparative Examples 1 and 2 were prepared according to the following procedure.

[0156] First, an L-lactic acid / ε-caprolactone copolymer (LCL5050, manufactured by BMG, trade name BioDegmer®) with a molar ratio of L-lactic acid:ε-caprolactone = 50:50, and a DL-lactic acid / ε-caprolactone copolymer (DLCL9010, manufactured by EVONIK) with a molar ratio of DL-lactic acid:ε-caprolactone = 90:10 were prepared. Next, LCL5050 and DLCL9010 were weighed to a total mass of 1 g at the predetermined blend ratio (mass%) shown in Table 1, and dissolved in 20 ml of acetone to prepare each polymer solution. Each of the obtained polymer solutions was poured into a φ100 mm PFA petri dish, taking care not to mix in air bubbles, air-dried at room temperature, and then dried under reduced pressure in a vacuum oven at 120°C for 2 hours. Each formed film (approximately 0.1 mm thick) was peeled off the PFA petri dish to obtain test films for Examples 1 and 2 and Comparative Examples 1 and 2.

[0157] <Measurement of Young's Modulus of Polymer Materials> The Young's modulus of the test films prepared in Examples 1 and 2 and Comparative Examples 1 and 2, as described above, was measured according to the following procedure. The results are shown in Table 1.

[0158] Using a die, the test film was cut into the shape of a Type 5B dumbbell test specimen as shown in ISO 527-2:2012. Then, a tensile test was performed using a tensile testing machine with a constant temperature chamber (Autograph AG-1kNIS, manufactured by Shimadzu Corporation) at 37°C with a chuck distance of 20 mm and a test speed of 1 mm / min. Young's modulus was determined from the initial slope of the stress-strain curve within the elastic deformation region.

[0159] <Measurement of Elongation at Break> The elongation at break was measured for the test films of Examples 1 and 2 and Comparative Examples 1 and 2 prepared above, according to the following procedure. The results are shown in Table 1.

[0160] Using a die-cutting mold, the test film was cut into the shape of a Type 5B dumbbell test specimen as specified in ISO 527-2:2012. Then, a tensile test was performed using a tensile testing machine with a constant temperature chamber (Autograph AG-1kNIS, manufactured by Shimadzu Corporation) at 37°C with a chuck distance of 20 mm and a test speed of 10 mm / min. The elongation at break was determined from the amount of strain at the time the sample broke.

[0161] <Preparation of Samples (Intravivo Implants)> Samples (intravivo implants) for Examples 1-2 and Comparative Examples 1-2 were prepared according to the following procedure.

[0162] LCL5050 and DLCL9010 were prepared in the same manner as those used in the above-mentioned <Preparation of Test Film>. Subsequently, LCL5050 and DLCL9010 were mixed at the predetermined blend ratio (mass%) shown in Table 1, dissolved with acetone, and a coating solution (polymer solution) with a concentration of 200 mg / mL was obtained.

[0163] Next, a φ2.5 mm core was inserted into the lumen of a tubular knitted micromesh (coarse size: 150 μm, wale size: 300 μm) made of 20 μm diameter PGA (polyglycolic acid) yarn. The coating solution was then applied to the micromesh by dispensing it from an air-operated dispenser (discharge pressure: 3 kPa, discharge time: 4 seconds) connected to a 25 G non-bevel needle while rotating the micromesh at a rotation speed of 18 rpm. As a result, it was confirmed that the coating solution filled the voids in the micromesh. The coating was applied to two locations on the mesh, with a distance of 18 mm between the coated areas (hereinafter referred to as coated areas).

[0164] Subsequently, the coated mesh was dried under reduced pressure at 120°C for 2 hours. Next, the coated mesh was cut at both ends with a razor blade so that the width of the coated area was 230 μm to 520 μm and the length was equivalent to the stent length described later, thereby obtaining a stent attachment mesh (porous structure) having a coated area of ​​230 μm to 520 μm at both ends.

[0165] Next, the stent attachment mesh was placed over the stent (balloon-expandable type, Ultimaster Nagomi, manufactured by Terumo Corporation, stent length 18 mm), and the stent attachment mesh was pressed against the curve at the outermost end of the stent with a pressing body made of silicone rubber. In this state, the covering area was irradiated with laser light, thereby welding and fixing the stent attachment mesh to the stent, and a sample of the implanted device was prepared.

[0166] Next, a sample of the implanted device was crimped onto a balloon catheter using a crimping machine. The final outer diameter of the stent portion at this time was 1.40 mm. The appearance of this state was photographed with a digital microscope (VHX-2000, KEYENCE). The resulting image is shown in Figure 20 as "before expansion".

[0167] <Expansion Test> Expansion tests were performed on samples of the in vivo implants prepared as described above for Examples 1-2 and Comparative Examples 1-2, according to the following procedure. The results are shown in Table 1 and Figure 20.

[0168] An indeflater (20 / 30, Abbott) was connected to the hub of a balloon catheter into which a sample of an implanted in vivo was crimped. After applying a pressure of 4 atm for 30 seconds, the appearance was photographed and observed using a digital microscope (VHX-2000, KEYENCE) and evaluated according to the following criteria. The captured image is shown in Figure 20 as "After 4 atm expansion".

[0169] Here, the expansion ratio was calculated by measuring the diameters of both ends of the stent (for convenience, one end will be referred to as the base side and the other as the tip side) and the diameter of the central part of the stent in the external images taken above, and using the following formula (1) for both the base side and the tip side.

[0170]

[0171] Furthermore, the presence or absence of fracture was confirmed by visual inspection by two evaluators. If the two evaluators' assessments differed, they consulted with each other, re-examined the presence or absence of fracture visually, and reached a conclusion on the evaluation.

[0172] Good: Expansion rate is 95% or higher, and no breakage of the coating is observed visually; Poor: Expansion rate is less than 95%, and / or breakage of the coating is observed visually.

[0173]

[0174] First, as can be seen from Figure 20, in the in vivo implants of Examples 1 and 2 and Comparative Examples 1 and 2 "before expansion," the radial expansion of the porous structure (particularly the mesh-like skeletal structure) was suppressed by the covering portions on both the tip and base sides. On the other hand, as shown in Table 1 and Figure 20 "after 4 atm expansion," in the samples of in vivo implants of Examples 1 and 2, the stent expansion rate was 95% or more, indicating that the stent was sufficiently expanded radially. It should be noted that the expansion rate is preferably 95% or more, more preferably 96% or more, and even more preferably 96.5% or more.

[0175] Furthermore, in the in-vivo implantation samples of Examples 1 and 2, the covering portion did not rupture after expansion. On the other hand, in the in-vivo implantation sample of Comparative Example 1, the expansion rate was low, and the stent was not sufficiently expanded radially, resulting in a poor result. In addition, although the in-vivo implantation sample of Comparative Example 2 appears to show a good expansion rate, this is because the covering portion ruptured, and therefore Comparative Example 2 also resulted in a poor result.

[0176] This application is based on Japanese Patent Application No. 2024-170444, filed on 30 September 2024, the disclosures of which are referenced and incorporated in whole.

[0177] 10 Stent 10A Stent tip 10B Stent base 11 Ring 12 Link 14 First strut 15 Second strut 16 Third strut 17 Curved section 19 Outer surface of stent 20 Porous structure 20A Tip of porous structure 20B Base of porous structure 20C Central part of porous structure 30 Skeleton 40 Gap 50 Covering section 50A Covering section 50a Covering material 53 Uneven section 53a Convex section 53b Recess 60 Drug carrying section 100 Implant 200 Balloon catheter 220 Balloon 300 Stent delivery system 400A Nozzle 400B Sprayer 500A Core 500B Jig

Claims

1. An in vivo implant comprising: an expandable cylindrical stent; and a porous structure disposed to cover the stent and configured to expand as the stent expands, wherein the porous structure has a mesh-like skeletal structure; a plurality of voids partitioned by the skeletal structure; and a covering portion disposed to cover at least a portion of the outer surface of the porous structure and containing a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

2. The in-vivo implant according to claim 1, wherein the elongation at break of the polymer material is 960% or more and 3500% or less.

3. The in-vivo implant according to claim 1, wherein the Young's modulus of the polymer material is greater than 0 MPa and 100 MPa or less.

4. The in-vivo implant according to claim 1, wherein the polymer material includes a biodegradable polymer.

5. The in vivo implant according to claim 4, wherein the biodegradable polymer comprises one or more selected from the group consisting of polylactic acid, polyglycolic acid, poly(ε-caprolactone), copolymer of lactic acid and ε-caprolactone, copolymer of lactic acid and glycolic acid, and copolymer of glycolic acid and ε-caprolactone.

6. The in vivo implant according to claim 4, wherein the biodegradable polymer includes a polylactic acid resin.

7. The in vivo implant according to claim 6, wherein the polylactic acid resin comprises a copolymer of lactic acid and a hydroxycarboxylic acid other than lactic acid.

8. The in vivo implant according to claim 7, wherein the lactic acid is L-lactic acid and the hydroxycarboxylic acid other than lactic acid is ε-caprolactone.

9. The in vivo implant according to claim 8, wherein the ratio of L-lactic acid to ε-caprolactone in the copolymer (moles of L-lactic acid : moles of ε-caprolactone) is 0.1 to 10:

1.

10. The intracellular implant according to claim 8, wherein the ratio (mass%) of the mass of the copolymer of L-lactic acid and ε-caprolactone to the total mass of the polymer material is greater than 25% by mass and less than or equal to 100% by mass.

11. The in vivo implant according to claim 1, wherein the covering portion is arranged to cover at least one of the ends of the porous structure located in the longitudinal direction in the circumferential direction.

12. The in vivo implant according to claim 1, wherein the porous structure is fixed to the stent via the covering portion.

13. A method for manufacturing a porous structure to be used in an in-vivo implant, comprising: a step of preparing a porous structure to be used in an in-vivo implant, which comprises an expandable cylindrical stent and a porous structure disposed to cover the stent and configured to expand in accordance with the expansion of the stent's diameter; and a step of forming a covering portion on at least a part of the outer surface of the porous structure by applying a covering material containing a polymer material, wherein the Young's modulus of the polymer material is greater than 0 MPa and less than 200 MPa.

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