Implant medical device for soft tissue repair and method for manufacturing same

Abstract

The purpose of the present invention is to provide an implant medical device having excellent strength and flexibility. The present invention provides an implant medical device for soft tissue repair, the device having a polyester copolymer comprising a polyhydroxycarboxylic acid in which two types of hydroxycarboxylic acid monomers are main constituent units, and having a thickness of 1.0 mm to 20.0 mm inclusive and a bulk density of 0.40 g / cm3 to 1.50 g / cm3 inclusive.

Description

Implant medical device for soft tissue repair and method of manufacturing same

[0001] The present invention relates to an implantable medical device for soft tissue repair and a method for manufacturing the same.

[0002] Tendons, the rotator cuff, which is a plate-like collection of tendons, and ligaments are tissues that connect muscles to bones or bones to bones, and play an important role in enabling smooth movement of joints such as the shoulder, elbow, knee, wrist, and ankle. Degeneration due to aging, trauma, or overuse in sports can damage or tear tendons, rotator cuffs, and ligaments, interfering with daily activities and causing pain not only during exercise but also at rest and at night.

[0003] Injuries to soft tissues such as tendons, rotator cuffs, and ligaments have traditionally been treated by surgical procedures to physically connect the torn tendons, rotator cuffs, or ligaments to the bone. However, it has been reported that re-rupture occurs in over 40% of cases after rotator cuff repair surgery in the shoulder, and that re-rupture occurs in over 90% of cases after reconstruction of extensive tears.

[0004] Based on the idea that re-rupture occurs because the mechanical strength of surgically connected tendons, rotator cuffs, and ligaments cannot be maintained during the repair process, treatments that aim to maintain mechanical strength and promote biological healing by reinforcing the joints between tendons, rotator cuffs, and ligaments and bones with sheet-like reconstructive materials are becoming increasingly common. In particular, biodegradable reconstructive materials have attracted attention because they do not remain in the body after repair, and biodegradable materials derived from biological or synthetic polymers are being used for this purpose.

[0005] As implant medical devices made of synthetic polymers, there have been disclosed meshes made of polylactic acid fibers with a controlled number of warp and weft threads, and foams made of polylactic acid reinforced with meshes that have been annealed and then rapidly cooled to improve their resistance to dissolution in organic solvents (Patent Documents 1 and 2).

[0006] Furthermore, a biodegradable polyester made of lactic acid and caprolactone has been described as a flexible synthetic polymer (Patent Document 3), and a medical molded article using the biodegradable polyester described in Patent Document 3 has been disclosed, which is intended to be placed in a living body and has a low Young's modulus when there is no hard tissue around the placement site, making it possible to treat the molded article in the same way as soft tissue (Patent Document 4).

[0007] Japanese Patent Publication No. 2008-529749 Japanese Patent No. 4937496 International Publication No. 2019 / 035357 International Publication No. 2020 / 122096

[0008] However, the implant medical devices described in Patent Documents 1 and 2 use polylactic acid, a hard synthetic polymer, and therefore, although they have excellent strength, they lack flexibility and have issues in terms of conformability to soft tissue. Furthermore, even when such a hard synthetic polymer is used, flexibility may be imparted by making the thickness or density extremely small, but because the polymer is easily degraded, it will degrade and be absorbed before the soft tissue repair is complete.

[0009] On the other hand, Patent Document 3 describes a biodegradable polyester composed of lactic acid and caprolactone, but does not specifically describe its application to implant devices for soft tissue repair. Patent Document 4 describes a medical molded body, a medical device, and a nerve regeneration guide tube using the biodegradable polyester described in Patent Document 3, but does not specifically describe its application to implant devices for soft tissue repair. For example, when the shapes (particularly thickness and bulk density) described in Patent Documents 1 and 2 are used, the strength is insufficient. Therefore, even when the soft synthetic polymers described in Patent Documents 3 and 4 are used, the thickness and bulk density must be controlled within an appropriate range in order to be applied to implant devices for soft tissue repair.

[0010] Therefore, the present invention aims to provide an implant medical device for soft tissue repair that has excellent strength and flexibility by using a biodegradable polyester copolymer containing polyhydroxycarboxylic acid, which is a soft synthetic polymer whose main structural unit is two types of hydroxycarboxylic acid monomers, and by controlling the thickness and bulk density within appropriate ranges.

[0011] The present inventors have conducted extensive research to solve the above problems and have discovered the following inventions (1) to (19): (1) A polyester copolymer containing polyhydroxycarboxylic acid, the main structural unit of which is two types of hydroxycarboxylic acid monomers, having a thickness of 1.0 mm or more and 20.0 mm or less and a bulk density of 0.40 g / cm ３ 1.50g / cm or more ３ (2) An implant medical device for soft tissue repair, having a first layer and a second layer, wherein the bulk density of the first layer is 0.20 g / cm or less. ３ 0.95g / cm or more ３ (3) The implant medical device according to (1), which has a first layer and a second layer, the first layer and the second layer having different bulk densities, a thickness of 1.0 mm or more and 20.0 mm or less, and an overall bulk density of 0.40 g / cm or less. ３ (4) An implant medical device for soft tissue repair, having a bulk density of 0.20 g / cm or more and 1.50 g / cm or less. (5) An implant medical device according to (3), comprising a polyester copolymer containing polyhydroxycarboxylic acid having two types of hydroxycarboxylic acid monomers as main structural units. ３ 0.95g / cm or more ３ (6) The implantable medical device according to (3) or (4), wherein the difference between the bulk density of the first layer and the bulk density of the second layer is 0.10 g / cm or less. ３ 0.90g / cm or more ３ (7) The implantable medical device according to any one of (2) to (5), wherein the bulk density of the second layer is 0.95 g / cm or less. ３ Greater than 2.00 g / cm ３(8) The implant medical device according to any one of (2) to (6), wherein the ratio of the thickness of the second layer to the thickness of the implant medical device is 5% or more and 40% or less. (9) The implant medical device according to any one of (2) to (8), wherein the first layer and the second layer are composed of the same components. (10) The implant medical device according to any one of (1) to (9), wherein the maximum tensile load is 20 N or more and 2,000 N or less. (11) The implant medical device according to any one of (1) to (10), wherein the tensile modulus is 0.01 MPa or more and 100 MPa or less. (12) The implant medical device according to any one of (1) to (11), wherein the maximum tensile load is 20 N or more and 2,000 N or less. (13) The implant medical device according to any one of (1), (2), and (4) to (12), wherein the R value of the polyhydroxycarboxylic acid, represented by the following formula 1, is 0.25 or more and 0.99 or less, when the two types of hydroxycarboxylic acid monomers are "monomer A" and "monomer B," respectively. R value = [AB] / (2[A][B]) × 100 ... formula 1 [A]: molar fraction (%) of monomer A residue in the polyhydroxycarboxylic acid [B]: molar fraction (%) of monomer B residue in the polyhydroxycarboxylic acid [AB]: molar fraction (%) of structures (A-B and B-A) in which a monomer A residue and a monomer B residue are adjacent to each other in the polyhydroxycarboxylic acid. (14) The implant medical device according to (13), wherein the monomer A is a monomer selected from the group consisting of lactic acid, glycolic acid, dilactide, and glycolide, and the monomer B is a monomer selected from the group consisting of hydroxyvaleric acid, hydroxycaproic acid, valerolactone, and caprolactone. (15) The implant medical device according to any one of (1) to (14), which consists solely of a polyester copolymer. (16) The implant medical device according to any one of (1) to (15), wherein the soft tissue is a tendon, a rotator cuff, or a ligament. (17) The implant medical device for repairing ruptured soft tissue according to any one of (1) to (16), which is used by being fixed between muscle and bone, or between bone and bone.(18) A method for repairing a ruptured tendon, rotator cuff, or ligament, comprising a fixation step of fixing the implant medical device according to any one of (1) to (17) to muscle and bone or bone and bone. (19) A method for manufacturing the implant medical device according to any one of (1) to (17).

[0012] Furthermore, as a result of intensive research aimed at solving the above problems, the present inventors have discovered the following inventions (20) to (23): (20) A polyester copolymer containing polyhydroxycarboxylic acid, the main structural unit of which is two types of hydroxycarboxylic acid monomers, having a thickness of 1.0 mm or more and 20.0 mm or less and a bulk density of 0.40 g / cm ３ 1.50g / cm or more ３ (21) An implant medical device for soft tissue repair, comprising a first layer and a second layer, wherein the bulk density of the first layer is 0.20 g / cm or less. ３ 0.95g / cm or more ３ (22) The implant medical device according to (20) or (21), wherein the R value of the polyhydroxycarboxylic acid, represented by the following formula 1, is 0.25 or more and 0.99 or less, when the two types of hydroxycarboxylic acid monomers are "monomer A" and "monomer B", respectively. R value = [AB] / (2[A][B]) × 100 ... formula 1 [A]: molar fraction (%) of monomer A residue in the polyhydroxycarboxylic acid [B]: molar fraction (%) of monomer B residue in the polyhydroxycarboxylic acid [AB]: molar fraction (%) of structures in which a monomer A residue and a monomer B residue are adjacent to each other (A-B and B-A) in the polyhydroxycarboxylic acid (23) A method for producing the implant medical device according to (20).

[0013] The present invention can provide an implantable medical device suitable for soft tissue repair applications.

[0014] FIG. 1 is an example of an implant medical device having a rectangular parallelepiped external structure consisting of a first layer and a second layer. FIG. 2 is an example of an attachment surface that is polygonal. FIG. 3 is an example of an attachment surface that is cross-shaped. FIG. 4 is an example of an attachment surface that is gourd-shaped and cloud-shaped. FIG. 5 is an example of an attachment surface that is fan-shaped. FIG. 6 is an example of an attachment surface that is broach-shaped. FIG. 7 is an example of an attachment surface that is lens-shaped. FIG. 8 is a schematic diagram of a rectangular parallelepiped with a curved attachment surface as viewed from the side. FIG. 9 is a schematic diagram of a rectangular parallelepiped with an uneven attachment surface as viewed from the side.

[0015] The implant medical device for soft tissue repair of the present invention has a polyester copolymer containing polyhydroxycarboxylic acid having two types of hydroxycarboxylic acid monomers as main structural units, and has a thickness of 1.0 mm or more and 20.0 mm or less and a bulk density of 0.40 g / cm ３ 1.50g / cm or more ３ The present invention is characterized by the following:

[0016] The implant medical device for soft tissue repair of the present invention has a first layer and a second layer, the first layer and the second layer having different bulk densities, a thickness of 1.0 mm or more and 20.0 mm or less, and an overall bulk density of 0.40 g / cm. ３ 1.50g / cm or more ３ The present invention is characterized by the following:

[0017] In this specification, soft tissue refers to tissues of the human body excluding bones, including cartilage, tendons, rotator cuffs, and ligaments, and particularly refers to tendons, rotator cuffs, and ligaments.

[0018] The polyester copolymer in the present invention refers to a polymer composed of two or more types of monomers, among polymers in which monomers are linked mainly by ester bonds. In this specification, the term polyester also refers to a polymer in which monomers are partially linked by bonds other than ester bonds, such as urethane bonds or amide bonds.

[0019] In order to control biodegradability within a suitable range, the number of bonds other than ester bonds among the bonds used for monomer linkage is preferably 30% or less, more preferably 10% or less, even more preferably 5% or less, and most preferably 0%, i.e., no bonds other than ester bonds, when the number of bonds used for monomer linkage is taken as 100%.

[0020] The polyester copolymer of the present invention contains a polyhydroxycarboxylic acid. Polyhydroxycarboxylic acid refers to a polymer composed of hydroxycarboxylic acid. Hydroxycarboxylic acid refers to a compound having both a hydroxy group and a carboxy group in the same molecule.

[0021] Since polyhydroxycarboxylic acids are biodegradable, polyester copolymers containing polyhydroxycarboxylic acids can be biodegradable. Furthermore, since polyester copolymers are biodegradable, the implantable medical device of the present invention containing a polyester copolymer is also biodegradable.

[0022] For suitable use in soft tissue repair applications, the implant medical device of the present invention is preferably biodegradable. To impart suitable biodegradability, the polyester copolymer of the present invention preferably contains 50% by mass or more of polyhydroxycarboxylic acid relative to 100% by mass of the polyester copolymer, more preferably 70% by mass or more, even more preferably 90% by mass or more, and most preferably 100% by mass, i.e., consists solely of polyhydroxycarboxylic acid.

[0023] From the viewpoint of providing biodegradability to the implant medical device of the present invention, the implant medical device used in the present invention preferably contains 50% by mass or more of polyester copolymer relative to 100% by mass of the implant medical device, more preferably 70% by mass or more, even more preferably 90% by mass or more, and most preferably 100% by mass, i.e., the implant medical device consists solely of polyester copolymer.

[0024] As used herein, biodegradability refers to the property of being decomposed in a living body or in the environment. More specifically, it refers to a degree of biodegradation of 60% or more as determined in Measurement Example 1 described below. Terms that can be used interchangeably with biodegradability include bioabsorbability and biocompatibility, but these terms often refer primarily to the property of being decomposed in a living body.

[0025] Polyhydroxycarboxylic acids can be obtained by polymerization of hydroxycarboxylic acids, polymerization of lactones, which are cyclic compounds formed by intramolecular dehydration condensation of the hydroxyl groups and carboxyl groups of hydroxycarboxylic acids, or polymerization of lactides, which are cyclic compounds formed by dehydration condensation of the hydroxyl groups and carboxyl groups of two molecules of hydroxycarboxylic acids, etc. That is, hydroxycarboxylic acids, lactones, or lactides can be used as hydroxycarboxylic acid monomers.

[0026] Examples of hydroxycarboxylic acids used in the hydroxycarboxylic acid monomer include lactic acid, glycolic acid, hydroxybutyric acid, hydroxyvaleric acid, hydroxycaproic acid, hydroxyheptanoic acid, hydroxyoctanoic acid, hydroxynonanoic acid, hydroxydecanoic acid, hydroxyundecanoic acid, hydroxydodecanoic acid, and (2-hydroxyethoxy)acetic acid. Of these, lactic acid, glycolic acid, hydroxyvaleric acid, and hydroxycaproic acid are preferred, and lactic acid and hydroxycaproic acid are particularly preferred.

[0027] Examples of lactones that can be used in the hydroxycarboxylic acid monomer include butyrolactone, valerolactone, caprolactone (hereinafter, CL), dioxepanone, ethylene oxalate, p-dioxanone, trimethylene carbonate, β-propiolactone, and pivalolactone. In particular, butyrolactone, valerolactone, CL, p-dioxanone, and trimethylene carbonate are preferred, and valerolactone and CL are more preferred.

[0028] As the lactide used for the hydroxycarboxylic acid monomer, dilactide (hereinafter referred to as LA) obtained by dehydration condensation of two molecules of lactic acid, glycolide obtained by dehydration condensation of two molecules of glycolic acid, or tetramethyl glycolide can be used, with LA or glycolide being particularly preferred.

[0029] The polyhydroxycarboxylic acid of the present invention has two types of hydroxycarboxylic acid monomers as its main constituent units. "The polyhydroxycarboxylic acid has two types of hydroxycarboxylic acid monomers as its main constituent units" means that the sum of the number of residues of the two types of monomers is 50 mol % or more, where the sum of the number of all monomer residues contained in the entire polyhydroxycarboxylic acid, including other monomer residues, is taken as 100 mol %, and each residue is 10 mol % or more, where the sum of the number of all monomer residues contained in the entire polyhydroxycarboxylic acid is taken as 100 mol %.

[0030] In this specification, the term "monomer residue" refers, in principle, to a repeating unit of a chemical structure derived from a monomer in the chemical structure of a polyhydroxycarboxylic acid obtained by polymerizing two or more monomers including the monomer.

[0031] For example, "having monomer A residues and monomer B residues as main structural units" means that the sum of the number of monomer A residues and monomer B residues is 50 mol % or more, and the monomer A residues are 10 mol % or more, and the monomer B residues are 10 mol % or more, when the sum of the number of all monomer residues contained in the entire polyhydroxycarboxylic acid is 100 mol %.

[0032] Here, the molar fractions (%) of the monomer A residue, the monomer B residue, and other residues can be determined from the area values ​​of the signals derived from each residue by nuclear magnetic resonance (hereinafter referred to as "NMR") measurement. For example, when the monomer A residue is an LA residue and the monomer B residue is a CL residue, they can be determined by Measurement Example 2 described below.

[0033] The molar fractions (%) of the monomer A residue, the monomer B residue, and other residues are calculated using the molecular weights of the repeating units of the chemical structures derived from the monomers. For example, if the monomer A residue is an LA residue and the monomer B residue is a CL residue, the molecular weight of the repeating unit of the monomer A residue is 144, and the molecular weight of the repeating unit of the monomer B residue is 114.

[0034] The molar fraction (%) of the sum of the monomer A residues and the monomer B residues is, based on the above definition, 50 mol % or more, preferably 75 mol % or more, and more preferably 90 mol % or more, when the sum of the number of all monomer residues contained in the entire polymer, including other monomer residues, is 100 mol %.

[0035] Furthermore, the molar fractions (%) of the monomer A residues and the monomer B residues are each, similarly as defined above, 10 mol % or more, preferably 20 mol % or more, more preferably 30 mol % or more, even more preferably 40 mol % or more, and most preferably 50 mol % or more. A particularly preferred embodiment is a polymer in which the sum of the monomer A residues and the monomer B residues is 100 mol % of the entire polymer, i.e., a polymer consisting only of the monomer A and the monomer B.

[0036] Furthermore, when two types of hydroxycarboxylic acid monomers are designated "monomer A" and "monomer B," respectively, a more preferred embodiment of the present invention is one in which the R value represented by the following formula 1 is 0.25 or more and 0.99 or less (hereinafter, polyhydroxycarboxylic acids having an R value of 0.25 or more and 0.99 or less will be referred to as "random polymers"): R value = [AB] / (2[A][B])×100 Formula 1 [A]: Molar fraction (%) of monomer A residues in the polyhydroxycarboxylic acid [B]: Molar fraction (%) of monomer B residues in the polyhydroxycarboxylic acid [AB]: Molar fraction (%) of structures in which a monomer A residue and a monomer B residue are adjacent to each other (A-B and B-A) in the polyhydroxycarboxylic acid

[0037] It is preferable to use a random polymer as the polyhydroxycarboxylic acid, that is, to set the R value of the polyhydroxycarboxylic acid within the above range, because a flexible implantable medical device can be obtained. From the viewpoint of flexibility, the lower limit of the R value is preferably 0.25 or more, and more preferably 0.45 or more.

[0038] Monomer A and Monomer B used in the present invention are preferably monomers selected from lactic acid, glycolic acid, hydroxybutyric acid, hydroxyvaleric acid, hydroxycaproic acid, hydroxyheptanoic acid, hydroxyoctanoic acid, hydroxynonanoic acid, hydroxydecanoic acid, hydroxyundecanoic acid, hydroxydodecanoic acid, (2-hydroxyethoxy)acetic acid, butyrolactone, valerolactone, CL, dioxepanone, ethylene oxalate, p-dioxanone, trimethylene carbonate, β-propiolactone, pivalolactone, LA, glycolide, and tetramethylglycolide.

[0039] Of the monomers A and B, it is particularly preferred that the monomer A is at least one monomer selected from the group consisting of lactic acid, glycolic acid, LA, and glycolide, and it is particularly preferred that the monomer B is at least one monomer selected from the group consisting of hydroxyvaleric acid, hydroxycaproic acid, valerolactone, and CL.

[0040] The weight-average molecular weight of the polyester copolymer of the present invention can be measured by the method described in Measurement Example 3. From the viewpoint of controlling strength and biodegradability within suitable ranges, the weight-average molecular weight of the polyester copolymer is preferably 50,000 or more, more preferably 100,000 or more, even more preferably 150,000 or more, and most preferably 200,000 or more. As the weight-average molecular weight increases, the strength increases and the number of days required for biodegradation increases. From the viewpoint of moldability, the weight-average molecular weight of the polyester copolymer is preferably 1,000,000 or less, more preferably 800,000 or less, and even more preferably 500,000 or less.

[0041] The polyester copolymer of the present invention may be a multi-polyester copolymer in which polyester copolymers are linked together. In this case, the polyester copolymer before linking is referred to as a "macromer" for convenience. In this specification, the average number of macromer molecules constituting one molecule of the polyester copolymer is referred to as the linkage number. The linkage number refers to the value obtained by dividing the weight average molecular weight of the polyester copolymer by the weight average molecular weight of the macromer.

[0042] The multi-polyester copolymer can be produced by linking polyester copolymers directly or by linking them using a linker molecule. Whether the multi-polyester copolymer has a linear or branched structure is determined by the structure of the polyester copolymer and the structure of the linker molecule.

[0043] When the polyester copolymer has a branched structure, the multi-polyester copolymer will have a branched structure regardless of the linking method. When the polyester copolymer has a linear structure, the multi-polyester copolymer will have a linear structure when linked directly or using a linear linker molecule, and the multi-polyester copolymer will have a branched structure when linked using a branched linker molecule.

[0044] As the linear linker molecule, a compound having two functional groups, such as a carboxyl group, a hydroxyl group, or an amino group, in the molecule can be used. Examples thereof include oxalic acid, malonic acid, succinic acid, glutaric acid, adipic acid, pimelic acid, suberic acid, azelaic acid, sebacic acid, dodecanedioic acid, ethylene glycol, propanediol, butanediol, pentanediol, hexanediol, heptanediol, octanediol, polyethylene glycol, polypropylene glycol, polyethylene glycol-polypropylene glycol copolymer, ethylenediamine, propanediamine, butanediamine, diaminopentane, hexamethylenediamine, diaminoheptane, diaminooctane, lactic acid, glycolic acid, caproic acid, alanine, isoleucine, leucine, methionine, valine, phenylalanine, tryptophan, asparagine, cysteine, glutamine, arginine, histidine, and glycine.

[0045] Furthermore, as a linker molecule having a branched structure, a compound having two or more functional groups, such as a carboxyl group, a hydroxyl group, and an amino group, in the molecule can be used, and examples thereof include malic acid, tartaric acid, citric acid, tyrosine, serine, threonine, aspartic acid, glutamic acid, and lysine.

[0046] In the polyhydroxycarboxylic acid of the present invention, it is preferable that the crystallinity of at least one of the monomer A residue or the monomer B residue is less than 14%. If the crystallinity is less than 14%, an increase in the tensile modulus of the implant medical device is suppressed, and an implant medical device suitable for soft tissue repair applications can be obtained. The crystallinity of the monomer A residue or the monomer B residue is more preferably 10% or less, and even more preferably 5% or less. It is preferable that the crystallinity of all monomer residues is low, but it is also acceptable for only some of the monomer residues to have a low crystallinity.

[0047] As used herein, the crystallization ratio of a monomer residue refers to the ratio of the heat of fusion per unit weight of a monomer residue in the polyhydroxycarboxylic acid of the present invention to the product of the heat of fusion per unit weight of a homopolymer consisting of only a certain monomer residue and the weight fraction of that monomer residue in the polyhydroxycarboxylic acid of the present invention. In other words, the crystallization ratio of a monomer A residue refers to the ratio of the heat of fusion per unit weight of a monomer A residue in the polyhydroxycarboxylic acid to the product of the heat of fusion per unit weight of a homopolymer consisting of only monomer A and the weight fraction of the monomer A residue in the polyhydroxycarboxylic acid of the present invention.

[0048] The crystallinity ratios of the monomer A residues and the monomer B residues indicate the proportion of the monomer A residues or the monomer B residues in the polyhydroxycarboxylic acid of the present invention that form a crystalline structure.

[0049] In particular, when the monomer A residue is a lactic acid residue derived from LA and the monomer B residue is a caprolactone residue derived from CL, the crystallinity of the lactic acid residue is preferably less than 14%, more preferably 10% or less. The crystallinity is determined specifically by the method described in Measurement Example 4 below.

[0050] An example of the structure of the implant medical device of the present invention is shown in Figure 1. The implant medical device of Figure 1 has a rectangular parallelepiped structure formed by stacking a film-like layer (first layer 2) made of a polyester copolymer and a film-like layer (second layer 3) made of a polyester copolymer, each having a different thickness and bulk density.

[0051] Here, when the implant medical device is a rectangular parallelepiped, the largest surface and the surface facing it, whichever surface is comprised of the layer with the lower bulk density, are referred to as the attachment surface for convenience. Furthermore, when the implant medical device of the present invention is a cylinder, the bottom surface of the layer with the lower bulk density is referred to as the attachment surface for convenience. In the case of the implant medical device of Figure 1, since it has a rectangular parallelepiped structure, the outer surface with the largest surface area of ​​the first layer 2 is the attachment surface 1.

[0052] Although the external structure of the implant medical device of the present invention is shown as a rectangular parallelepiped in Fig. 1, it is not limited to this. Specific examples of the external structure include a rectangular parallelepiped, a pyramid, a sphere, a cylinder, and a cone, but a rectangular parallelepiped or a cylinder is preferable.

[0053] In the present invention, the shape of the rectangular attachment surface is not limited to the quadrilateral shape of the implant device shown in FIG. 1, but also includes polygonal shapes such as a triangle, pentagon, or decagon as shown in FIG. 2, and a cross shape as shown in FIG. 3, etc.

[0054] The attachment surface of the cylinder is not limited to a circle, but also includes shapes such as an ellipse, a gourd shape and a cloud shape as shown in Fig. 4, a fan shape, a half-moon shape and a teardrop shape as shown in Fig. 5, a brooch shape as shown in Fig. 6, and a lens shape as shown in Fig. 7. Furthermore, the surfaces constituting the external structure may have one curved surface as shown in Fig. 8, or one concave or convex surface as shown in Fig. 9.

[0055] In the present invention, the thickness of an implant medical device refers to the distance between the attachment surface and the opposing surface when the implant medical device is a rectangular parallelepiped or cylindrical. When the thickness is not uniform, the thickness is measured at multiple locations, including the maximum and minimum thicknesses, without any deviation within the surface, and the average value is taken as the thickness of the implant medical device. Furthermore, when the implant medical device is a pyramid, the thickness is taken as half the length of the short side of the base. When the implant medical device is a sphere, the thickness is taken as the radius. When the implant medical device is a cone, the thickness is taken as the radius of the base. As a specific example, for a shape with a curved surface as shown in Figure 8, the minimum thickness 4 is the distance between the apex of the curve within the curved surface and the other surface, and the maximum thickness 5 is the distance between the end of the curved surface and the other surface. Furthermore, for a shape with a concave surface as shown in Figure 9, the minimum thickness 4 is the distance between the apex of the concave surface and the other surface, and the maximum thickness 5 is the distance between the smooth surfaces. The thickness of these implant medical devices can be measured using the method described in Measurement Example 5.

[0056] The thickness of the implant medical device of the present invention is preferably 1.0 to 20.0 mm, more preferably 1.2 to 10.0 mm, even more preferably 1.5 to 5.0 mm, and most preferably 2.0 to 3.0 mm, from the viewpoint of controlling the strength and flexibility within suitable ranges.

[0057] Furthermore, in this specification, the direction perpendicular to the thickness direction is referred to as the longitudinal direction, and the direction perpendicular to the thickness and longitudinal directions is referred to as the horizontal direction, with the longitudinal length of the implant medical device of the present invention being referred to as the longitudinal length and the horizontal length being referred to as the horizontal length. The longitudinal length and horizontal length are selected so that their sums are maximized. The ranges of the longitudinal length and horizontal length of the implant medical device of the present invention are not particularly specified, but ranges of 50 cm or less, 30 cm or less, 15 cm or less, or 5 cm or less can be selected as ranges that can be placed in the affected area. From the perspective of feasibility, the lower limit is 0.1 cm or more. The thickness of the implant medical device is always smaller than the longitudinal length and horizontal length of the implant medical device, and the horizontal length is the same as or smaller than the longitudinal length.

[0058] The bulk density of the implant medical device refers to the value obtained by dividing the weight of the implant medical device by the volume including the pores and internal voids, and can be measured by the method described in Measurement Example 6. From the viewpoint of controlling the strength and flexibility within suitable ranges, the bulk density of the implant medical device is set to 0.40 g / cm ３ 1.50g / cm or more ３ Preferably, 0.45 g / cm or less ３ 1.40g / cm or more ３ More preferably, 0.60 g / cm ３ 1.30g / cm or more ３ is more preferably 0.80 g / cm ３ 1.20g / cm or more ３ On the other hand, from the viewpoint of controlling the platelet-rich plasma (hereinafter referred to as PRP) carrying capacity within a suitable range, 1.10 g / cm is most preferable. ３ Preferably, 1.00 g / cm or less ３ More preferably, 0.90 g / cm or less ３ More preferably, 0.80 g / cm ３ The following is most preferred:

[0059] The implant medical device of the present invention preferably comprises at least two layers, a first layer and a second layer. By having a structure including a first layer with excellent PRP-supporting ability and a second layer with high strength, it can exhibit functions suitable for soft tissue repair applications. Note that layers other than the first and second layers may be included, but the thickness of each layer must be smaller than that of the first and second layers.

[0060] The bulk density of the first layer and the bulk density of the second layer can be measured by the method described in Measurement Example 6. From the viewpoint of controlling the PRP loading capacity within a suitable range, the bulk density of the first layer is set to 0.20 g / cm ３ 0.95g / cm or more ３ Preferably, 0.25 g / cm or less ３ 0.85g / cm or more ３ More preferably, 0.30 g / cm or less ３ 0.75g / cm or more ３ More preferably, 0.35 g / cm ３ 0.60g / cm or more ３The following is most preferred:

[0061] From the viewpoint of controlling the strength and flexibility within a suitable range, the bulk density of the second layer is 0.95 g / cm ３ Greater than 2.00 g / cm ３ Preferably, 1.00 g / cm or less ３ 1.80g / cm or more ３ More preferably, 1.05 g / cm or less ３ 1.65g / cm or more ３ More preferably, 1.10 g / cm ３ 1.50g / cm or more ３ The following is most preferred:

[0062] The combination of the bulk densities of the first layer and the second layer may be any combination as long as the bulk density of each layer is within the above range and the bulk density of the implant medical device as a whole is within the above range. However, from the viewpoint of controlling the PRP carrying capacity within a suitable range, it is preferable that the difference between the bulk density of the first layer and the bulk density of the second layer is 0.10 g / cm. ３ 0.90g / cm or more ３ It is preferable that the density is 0.15 g / cm or less. ３ 0.80g / cm or more ３ More preferably, it is 0.20 g / cm or less. ３ 0.75g / cm or more ３ It is even more preferable that:

[0063] Examples of pretreatment methods for measuring the bulk density of the first layer and the second layer include the following: The cross section of the implant medical device is observed with an optical microscope, a scanning electron microscope, or the like to confirm the boundary surface between the first and second layers, and the implant medical device is cut along the boundary surface at a plane sufficiently distant from the boundary surface to obtain all or part of the first and second layers. All or part of the obtained first and second layers are measured using the method described in Measurement Example 6.

[0064] The ratio of the thickness of the second layer to the overall thickness of the implant medical device can be measured by the method described in Measurement Example 7. From the viewpoint of controlling the bulk density of the implant medical device within a suitable range, the ratio of the thickness of the second layer to the thickness of the implant medical device is preferably 5% or more and 40% or less, more preferably 7% or more and 35% or less, and even more preferably 10% or more and 30% or less.

[0065] The vertical and horizontal lengths of the first and second layers may be different, but it is preferable that the vertical and horizontal lengths of the first and second layers are the same.

[0066] In the implant medical device of the present invention, the first and second layers are preferably composed of the same components. By using the same components, the degradation rates of the first and second layers are equal, and it is expected that the entire implant medical device will be degraded and absorbed uniformly. On the other hand, by using different components for the first and second layers, it is possible to intentionally change the degradation period and maximize the therapeutic effect on soft tissue. Furthermore, when the present invention is used for soft tissue regeneration, it is preferable that the first layer be the application surface.

[0067] The maximum tensile load of the implant medical device of the present invention can be measured by the method described in Measurement Example 8. Specifically, the maximum tensile load refers to the maximum load observed in a single measurement. The maximum tensile load of the implant medical device is preferably 20 N or more, more preferably 50 N or more, and even more preferably 70 N or more. There is no particular upper limit, but as a feasible range, 2,000 N or less, 1,000 N or less, or 500 N or less can be selected.

[0068] The tensile modulus of elasticity of the implant medical device of the present invention can be measured by the method described in Measurement Example 8. The tensile modulus of elasticity of the implant medical device is preferably 100 MPa or less, more preferably 50 MPa or less, even more preferably 20 MPa or less, and most preferably 10 MPa or less. There is no particular lower limit, but as a practicable range, 0.01 MPa or more, 0.05 MPa or more, or 0.10 MPa or more can be selected.

[0069] The retracted length of the implant medical device of the present invention can be measured by the method described in Measurement Example 9. The retracted length of the implant medical device of the present invention is preferably 50 mm or less, more preferably 30 mm or less, even more preferably 15 mm or less, and most preferably 10 mm or less. There is no particular lower limit, but as a feasible range, 0.1 mm or more, 1 mm or more, or 3 mm or more can be selected. The shorter the retracted length, the more easily the device can be accommodated in a tube with a smaller inner diameter, so by setting the retracted length within the above range, it is possible to quickly deliver the device to the affected area during surgery using an endoscope.

[0070] After surgical treatment for soft tissue repair, platelet-rich plasma (PRP) is sometimes injected into the affected area to promote tissue repair. Therefore, the implantable medical device of the present invention preferably has the ability to support PRP.

[0071] The PRP-loading capacity can be evaluated by the amount of PRP loaded per unit weight of the implant medical device, and can be measured by the method described in Measurement Example 10. The injection amount of PRP in a typical treatment is 1 mL, and in order for the implant medical device to load the entire amount of injected PRP, a PRP-loading capacity of 0.20 g / g or more is required. Therefore, the PRP-loading capacity of the implant medical device of the present invention is preferably 0.20 g / g or more, more preferably 0.40 g / g or more, even more preferably 0.60 g / g or more, and most preferably 0.80 g / g or more.

[0072] The implant medical device of the present invention can be produced by various methods such as injection molding, extrusion molding, freeze-drying, electrospinning, or 3D printing. From the viewpoint of precisely controlling the bulk density, however, it is preferable to produce it by freeze-drying, electrospinning, or 3D printing, and it is more preferable to produce it by 3D printing.

[0073] Examples of 3D printers include fused deposition modeling (FDM), stereolithography (SLA or DLP), inkjet printing, powder sintering (SLS or SLM), and powder bonding. The implant medical device of the present invention is preferably produced by the fused deposition modeling (FDM), stereolithography (SLA or DLP), or powder sintering (SLS or SLM), and more preferably by the fused deposition modeling (FDM) or powder sintering (SLS or SLM).

[0074] When fabricating the implant medical device of the present invention by 3D printing, the thickness of each layer to be laminated (layer pitch) can be set. When fabricating the implant medical device of the present invention, any layer pitch can be selected, for example, from 0.01 mm to 1 mm.

[0075] When fabricating the object by 3D printing, the method of filling the interior of the object (infill pattern) can be selected. Examples of infill patterns include line, grid, triangle, trihexagon, cubic subdivision, octet, quarter cubic, zigzag, cross, cross 3D, star, cubic, honeycomb, 3D honeycomb, gyroid, Himbert curve, Archimedes chord, and octagram spiral. However, when fabricating the implant medical device of the present invention, line, grid, triangle, trihexagon, zigzag, star, cubic, concentric circle, honeycomb, 3D honeycomb, or gyroid is preferred.

[0076] The implant medical device of the present invention is used to repair soft tissue, and is particularly suitable for repairing ruptured soft tissue. For use in repairing ruptured soft tissue, it is preferably fixed to muscle and bone or bone and bone. A method for repairing a ruptured tendon, rotator cuff, or ligament preferably includes a fixation step of fixing the implant medical device of the present invention to muscle and bone or bone and bone.

[0077] EXAMPLES The present invention will be described in detail below with reference to examples and comparative examples, but the present invention is not limited to these.

[0078] (Measurement Example 1: Biodegradability Evaluation) Measurement was carried out in accordance with the method specified in JIS K6950 (2000). Specifically, an inorganic salt solution (300 mL), activated sludge (9 mg), and a polyester copolymer or an implant medical device (30 mg) were added to a culture bottle and stirred at 25°C for 90 days. Using a dissolved oxygen meter, the amount of dissolved oxygen before and after stirring was determined, and the difference was taken as the oxygen consumption. The degree of biodegradation was calculated according to the following formula 2 based on the oxygen consumption when each material was added. In this case, a biodegradation degree of 60% or more was considered to be biodegradable. Biodegradability (%) = (BOD - B) / TOD x 100 Formula 2 BOD: Oxygen consumption of the polyester copolymer or implant medical device B: Oxygen consumption of the blank TOD: Theoretical oxygen consumption required for complete oxidation of the polyester copolymer or implant medical device

[0079] (Measurement Example 2: Molar fraction and R value of each residue) The purified polyhydroxycarboxylic acid was dissolved in deuterated chloroform. １ H-NMR was measured to calculate the molar fraction (%) of monomer residues in the polyhydroxycarboxylic acid.

[0080] When the polyhydroxycarboxylic acid has two types of hydroxycarboxylic acid monomers as main structural units, the two types of hydroxycarboxylic acid monomers are designated as "monomer A" and "monomer B," respectively, and the molar fractions (%) of the monomer A residues and the monomer B residues are calculated, respectively. １Using H homospin decoupling, the peaks of monomer A and monomer B were separated by the signal derived from adjacent monomer residues of monomer A or monomer B, and the respective peak areas were quantified. From the respective peak area ratios, [AB] was calculated, and the R value was calculated using the following formula 1: R value = [AB] / (2[A][B]) × 100 ...Formula 1 [A]: Molar fraction (%) of monomer A residues in the polyhydroxycarboxylic acid; [B]: Molar fraction (%) of monomer B residues in the polyhydroxycarboxylic acid; [AB]: Molar fraction (%) of structures in which a monomer A residue and a monomer B residue are adjacent (A-B and B-A) in the polyhydroxycarboxylic acid. Here, [AB] is the molar fraction (%) of structures in which a monomer A residue and a monomer B residue are adjacent (A-B and B-A) in the polyhydroxycarboxylic acid, specifically, the ratio of the number of A-B and B-A to the total number of A-A, A-B, B-A, and B-B.

[0081] For example, if monomer A is LA and monomer B is CL, under the following condition 1 １ H homospin decoupling was performed to separate the adjacent monomer residues, namely, the methine group of LA (around 5.10 ppm), the α-methylene group of CL (around 2.35 ppm), and the ε-methylene group (around 4.10 ppm), by the signal derived from LA or CL, and the respective peak areas were quantified. Condition 1: Instrument name: JNM-ECZ400R (manufactured by JEOL Ltd.) １ H homospin decoupling Irradiation position: 1.66 ppm Solvent: deuterated chloroform Measurement temperature: room temperature

[0082] (Measurement Example 3: Weight-average molecular weight) A polyester copolymer was dissolved in chloroform and passed through a 0.45 μm syringe filter (DISMIC-13HP; manufactured by ADVANTEC) to remove impurities, and then GPC measurement was carried out under the following condition 2 to calculate the weight-average molecular weight of the polyester copolymer. Condition 2: Instrument name: Prominence (registered trademark; manufactured by Shimadzu Corporation) Mobile phase: Chloroform (for HPLC) (manufactured by Wako Pure Chemical Industries, Ltd.) Flow rate: 1 mL / min Column: TSKgel (registered trademark) GMH ＨＲ－M (φ7.8 mm × 300 mm; manufactured by Tosoh Corporation) Detector: UV (254 nm), RI Column, detector temperature: 35°C Standard material: polystyrene

[0083] (Measurement Example 4: Crystallization rate measured by DSC) Polyhydroxycarboxylic acid was dissolved in chloroform to a concentration of 5% by mass, and the solution was transferred to a polytetrafluoroethylene (hereinafter, referred to as PTFE) petri dish and dried at room temperature under normal pressure for 24 hours. This was then dried under reduced pressure at 50°C for 24 hours to obtain a film having a thickness of approximately 100 μm.

[0084] The obtained film (about 10 mg) was collected on alumina PAN and subjected to DSC measurement under the following condition 3. From the measurement results under temperature conditions (D) to (E), the crystallization rate was calculated using the following equations 3 and 4. Condition 3: Instrument name: EXSTAR 6000 (Seiko Instruments Inc.) Temperature conditions: (A) 25°C → (B) 250°C (10°C / min) → (C) 250°C (5 min) → (D) -70°C (10°C / min) → (E) 250°C (10°C / min) → (F) 250°C (5 min) → (G) 25°C (100°C / min) Standard substance: alumina Crystallization ratio of monomer A = (heat of fusion derived from monomer A residue per unit mass of polyhydroxycarboxylic acid) / {(heat of fusion per unit mass of homopolymer consisting of only monomer A residue) × (mass fraction of monomer A residue in polyhydroxycarboxylic acid)} × 100 ... Equation 3 Crystallization ratio of monomer B = (Heat of fusion derived from monomer B residues per unit mass of polyhydroxycarboxylic acid) / {(Heat of fusion per unit mass of homopolymer consisting of only monomer B residues)×(mass fraction of monomer B residues in polyhydroxycarboxylic acid)}×100 Equation 4

[0085] Here, the heat of fusion per unit mass of the homopolymer was determined as follows. The homopolymer was dissolved in chloroform to a concentration of 5% by mass, and the solution was transferred onto a PTFE petri dish and dried overnight at normal pressure and room temperature. This was then dried under reduced pressure to obtain a film. The obtained film was placed on an alumina PAN and subjected to DSC measurement under the following condition 4. The heat of fusion was read from the melting peak area on the graph obtained from the measurement results under temperature conditions (D) to (E). Condition 4: Instrument name: EXSTAR 6000 (manufactured by Seiko Instruments Inc.) Temperature conditions: (A) 25°C → (B) 250°C (10°C / min) → (C) 250°C (5 min) → (D) -70°C (10°C / min) → (E) 250°C (10°C / min) → (F) 250°C (5 min) → (G) 25°C (100°C / min) Standard material: alumina

[0086] (Measurement Example 5: Thickness of implant medical device) The thickness of 10 locations including the maximum and minimum thicknesses of the implant medical device and ensuring no bias within the surface was measured using a constant pressure thickness measuring device, and the average value was taken as the thickness of the implant medical device.

[0087] (Measurement Example 6: Bulk density of implant medical device, bulk density of first layer, bulk density of second layer) In this specification, the terms are defined as follows. Open pore: A void that exists on the surface of a material and is connected to the outside air. Closed pore: A void that is isolated inside a material and is not connected to the outside air. Dry mass: The mass of the material only. Expressed as W1. Mass in liquid: The mass of a material when immersed in liquid. Expressed as W2. Wet mass: The sum of the mass of the material and the liquid that has filled the open pores. Expressed as W3.

[0088] The bulk density of the implant medical device, first layer, or second layer (hereinafter referred to as the sample) was measured using the following procedure. To dry the sample, vacuum drying at room temperature may be performed, but no other pretreatment is required.

[0089] Step 1: The mass of the sample in a dry state was measured using an electronic balance, and the displayed value was designated as W1. Step 2: A container filled with water was prepared on top of the electronic balance, and the sample was submerged in the water so that the entire sample was immersed but not touching the bottom, and the displayed value was designated as W1-W2. The value of W2 was also determined from the value of W1 determined in Step 1 and the value of W1-W2. Step 3: The sample was removed from the water, and the mass of the sample in a wet state was measured using an electronic balance, and the displayed value was designated as W3.

[0090] The bulk density of the sample was determined using the values ​​of W1, W2, and W3 obtained in steps 1 to 3 and the following equation 5. If the sample dissolves in water, the bulk density can be determined by performing steps 1 to 3 above using a liquid with a known density instead of water. Bulk density = (W1 x ρ) / (W3 - W2) ...Equation 5 ρ: Density of the liquid used in steps 1 to 3. As mentioned above, when water is used, the value of ρ is ρ = 1.

[0091] (Measurement Example 7: Ratio of second layer thickness to overall thickness) A cross section of the implant medical device cut in the thickness direction was observed under an electron microscope, and based on the obtained image, the thickness of the second layer of the implant medical device was measured three times with different fields of view. Furthermore, the above measurement was performed three times with different cut surfaces. The average of the nine obtained values ​​was taken as the thickness of the second layer of the implant medical device, and the ratio of the second layer thickness to the overall thickness was calculated using the following formula 6. Thickness ratio of second layer (%) = thickness of second layer / thickness of implant medical device ... formula 6

[0092] (Measurement Example 8: Maximum tensile load and tensile modulus) A tensile test was performed on the implant medical device in accordance with the method specified in JIS K7161 (2014) under the following condition 5 to measure the maximum tensile load and tensile modulus. Each measurement was performed three times, and the number average was calculated to determine the maximum tensile load and tensile modulus of the implant medical device. Note that the maximum tensile load and tensile modulus were values ​​rounded to one decimal place. Condition 5: Device name: EZ-LX (manufactured by Shimadzu Corporation) Initial length: 10 mm Tensile speed: 25 mm / min Load cell: 1 kN

[0093] (Measurement Example 9: Retracted Length) The implant medical device was rolled up, and the diameter was measured using a digital caliper. The diameter or thickness was measured in a direction perpendicular to the length direction.

[0094] (Measurement Example 10: PRP holding capacity) PRP was prepared according to the following procedure. 25 mL of bovine whole blood was placed in a 50 mL centrifuge tube and centrifuged under the following condition (1), and the supernatant was collected. The remaining liquid was then centrifuged under the following condition (2), and the supernatant was collected and mixed with the previous supernatant to prepare PRP. Condition (1): 140×g, 10 minutes, 4°C → 200×g, 10 minutes, 4°C → 300×g, 10 minutes, 4°C Condition (2): 3000 rpm, 10 minutes, 4°C

[0095] The weight of the implant medical device in a dry state was measured and designated as Wi. Subsequently, 15 mL of PRP was placed in a polystyrene container, and the implant medical device was immersed and left to stand for 5 minutes. The implant medical device was then removed and its weight was measured. The weight of PRP supported on the implant medical device was calculated from the difference in weight of the implant medical device before and after immersion in PRP, and designated as Wp. The PRP supporting capacity was calculated using the following formula 7. PRP supporting capacity = Wp / Wi... formula 7

[0096] Synthesis Example 1 50.0 g of L,L-dilactide (PURASORB® L; manufactured by PURAC Corporation) and 39.6 g of ε-caprolactone (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) as monomers and 0.45 g of hydroxypivalic acid as an initiator were placed in a separable flask. Under an argon atmosphere, 37.4 mL of xylene (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) was added, followed by the addition of 0.28 g of tin(II) octoate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) as a catalyst dissolved in 2.0 mL of xylene (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.). The mixture was reacted at 130°C for 15 hours to obtain a crude copolymer. The obtained crude copolymer was dissolved in 260 mL of chloroform and added dropwise to 3,000 mL of hexane in a stirred state to obtain a precipitate. The precipitate was dried under reduced pressure at 50°C to obtain Macromer 1.

[0097] 50 g of macromer 1, 1.3 g of the catalyst 4,4-dimethylaminopyridinium p-toluenesulfonate (synthetic product), and 0.52 g of 4,4-dimethylaminopyridine (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) were collected and dissolved in 184 mL of dichloromethane (dehydrated) (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) under an argon atmosphere, and 0.56 mL of the condensing agent diisopropylcarbodiimide (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) was added, followed by condensation polymerization at room temperature overnight.

[0098] The reaction mixture was diluted with 240 mL of chloroform, and 10.0 g of D,L-lactic acid (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) was added and stirred for 30 minutes. Next, 500 mL of 0.5 M hydrochloric acid was added, followed by stirring for 20 minutes, and the aqueous layer was removed by decantation. Thereafter, 500 mL of ion-exchanged water was added, and the mixture was stirred for 10 minutes, followed by the process of removing the aqueous layer by decantation. This process was repeated until the pH of the removed aqueous layer reached 7. The remaining organic layer was added dropwise to 3,000 mL of methanol while stirring, to obtain a precipitate. This precipitate was dried under reduced pressure at 50°C, yielding polyester copolymer 1.

[0099] Synthesis Example 2 50.0 g of L-lactide (PURASORB® L; manufactured by PURAC Corporation) and 38.5 mL of ε-caprolactone (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) were collected as monomers in a separable flask. Under an argon atmosphere, 0.81 g of tin(II) octoate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) as a catalyst dissolved in 14.5 mL of toluene (ultra-dehydrated) (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) and ion-exchanged water as a co-initiator were added so that the monomer / co-initiator ratio was 285.7. A co-catalyst reaction was carried out at 90°C for 1 hour, followed by a copolymerization reaction at 150°C for 12 hours to obtain a crude copolymer.

[0100] The resulting crude copolymer was dissolved in 100 mL of chloroform and added dropwise to 1,400 mL of stirred methanol to obtain a precipitate. This procedure was repeated three times, and the precipitate was dried under reduced pressure at 70° C. to obtain Macromer 2.

[0101] 30 g of macromer 2, 1.12 g of 4,4-dimethylaminopyridinium p-toluenesulfonate (synthetic product) as a catalyst, and 0.40 g of 4,4-dimethylaminopyridine (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) were collected. These were dissolved in dichloromethane (dehydrated) (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) to a concentration of 20% under an argon atmosphere, and 5.52 g of dicyclohexylcarbodiimide (manufactured by Sigma-Aldrich Co.) as a condensation agent dissolved in 20 mL of dichloromethane was added, followed by condensation polymerization at room temperature for 2 days.

[0102] 120 mL of chloroform was added to the reaction mixture, and the mixture was added dropwise to 2000 mL of stirred methanol to obtain a precipitate. This precipitate was dissolved in 200 mL of chloroform, and the resulting solution was added dropwise to 2000 mL of stirred methanol to obtain a precipitate. This procedure was repeated twice, and the obtained precipitate was dried under reduced pressure at 50°C to obtain polyester copolymer 2.

[0103] Synthesis Example 3 50.0 g of L-lactide (PURASORB® L; manufactured by PURAC Corporation) and 38.5 mL of ε-caprolactone (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) were collected as monomers in a separable flask. Under an argon atmosphere, 0.81 g of tin(II) octoate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) as a catalyst dissolved in 14.5 mL of toluene (ultra-dehydrated) (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) was added, and a copolymerization reaction was carried out at 150° C. for 6 hours to obtain a crude copolymer.

[0104] The obtained crude copolymer was dissolved in 100 mL of chloroform and added dropwise to 1,400 mL of methanol under stirring to obtain a precipitate. This procedure was repeated three times, and the obtained precipitate was dried under reduced pressure at 50° C. to obtain polyester copolymer 3.

[0105] Synthesis Example 4 50.0 g of L-lactide (PURASORB® L; manufactured by PURAC Corporation) was collected as a monomer in a separable flask. Under an argon atmosphere, 0.81 g of tin(II) octoate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) as a catalyst dissolved in 14.5 mL of toluene (ultra-dehydrated) (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) and ion-exchanged water as a co-initiator were added so that the monomer / co-initiator ratio was 142.9. A co-catalyst reaction was carried out at 90°C for 1 hour, followed by a polymerization reaction at 150°C for 3 hours.

[0106] To this was added 38.5 mL of ε-caprolactone (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.), and the mixture was further polymerized at 150° C. for 6 hours to obtain a crude copolymer.

[0107] The resulting crude copolymer was dissolved in 100 mL of chloroform and added dropwise to 1,400 mL of stirred methanol to obtain a precipitate. This procedure was repeated three times, and the precipitate was dried under reduced pressure at 70° C. to obtain Macromer 3.

[0108] 7.5 g of Macromer 3, 0.28 g of 4,4-dimethylaminopyridinium p-toluenesulfonate (synthetic product) as a catalyst, and 0.10 g of 4,4-dimethylaminopyridine (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) were collected. These were dissolved in dichloromethane (dehydrated) (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) to a concentration of 27% under an argon atmosphere, and 1.38 g of dicyclohexylcarbodiimide (manufactured by Sigma-Aldrich Co.) as a condensation agent dissolved in 5 mL of dichloromethane was added, followed by condensation polymerization at room temperature for 2 days.

[0109] 30 mL of chloroform was added to the reaction mixture, and the mixture was added dropwise to 500 mL of stirred methanol to obtain a precipitate. This precipitate was dissolved in 50 mL of chloroform, and the resulting solution was added dropwise to 500 mL of stirred methanol to obtain a precipitate. This procedure was repeated twice, and the obtained precipitate was dried under reduced pressure at 50°C to obtain polyester copolymer 4.

[0110] The obtained polyester copolymers 1 to 4 were subjected to the measurements described in Measurement Examples 1 to 4. The results are shown in Table 1.

[0111]

[0112] (Example 1) Using polyester copolymer 1, a vertical length and horizontal length were 35 mm, a thickness was 2 mm, and the bulk density of the first layer was 0.55 g / cm ３ , the bulk density of the second layer is 1.10 g / cm ３ The implant medical device was fabricated by 3D printing so that the thickness ratio of the second layer was 30%. The device was fabricated using a fused deposition modeling (FDM) 3D printer with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0113] (Example 2) Using polyester copolymer 1, a vertical length and a horizontal length were 35 mm, a thickness was 2 mm, and the bulk density of the first layer was 0.35 g / cm ３ , the bulk density of the second layer is 1.10 g / cm ３ The implant medical device was fabricated by 3D printing so that the thickness ratio of the second layer was 10%. The device was fabricated using a fused deposition modeling (FDM) 3D printer with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0114] (Example 3) Using polyester copolymer 1, a vertical length and horizontal length were 35 mm, a thickness was 2 mm, and the bulk density of the first layer was 0.85 g / cm ３ , the bulk density of the second layer is 1.00 g / cm ３ The implant medical device was fabricated by 3D printing so that the thickness ratio of the second layer was 10%. The device was fabricated using a fused deposition modeling (FDM) 3D printer with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0115] (Example 4) Using polyester copolymer 1, a vertical length and horizontal length of 35 mm, a thickness of 2 mm, and a bulk density of 1.10 g / cm ３ The implant medical device was fabricated by 3D printing so that it would look like this. A fused deposition modeling (FDM) 3D printer was used, with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0116] Example 5 An implantable medical device was produced using polyester copolymer 2 in the same manner as in Example 1.

[0117] Example 6 An implantable medical device was produced using polyester copolymer 3 in the same manner as in Example 1.

[0118] Example 7 An implantable medical device was produced using polyester copolymer 4 in the same manner as in Example 1.

[0119] (Comparative Example 1) Using polyester copolymer 1, a vertical length and a horizontal length were 35 mm, a thickness was 0.4 mm, and a first layer bulk density was 0.55 g / cm ３ , the second layer bulk density is 1.10 g / cm ３ The implant medical device was fabricated by 3D printing so that the thickness ratio of the second layer was 50%. The device was fabricated using a fused deposition modeling (FDM) 3D printer with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0120] (Comparative Example 2) Using polyester copolymer 1, a vertical length and horizontal length of 35 mm, a thickness of 2 mm, and a bulk density of 0.35 g / cm ３ An implant medical device was fabricated by 3D printing so that the thickness of the implant medical device was 35 mm. A fused deposition modeling (FDM) 3D printer was used, with a layer pitch of 0.2 mm and an infill pattern set to a grid. (Example 8) Using polyester copolymer 1, an implant medical device was fabricated with a length and width of 35 mm, a thickness of 20 mm, and a bulk density of the first layer of 0.55 g / cm. ３ , the bulk density of the second layer is 1.10 g / cm ３ An implant medical device was fabricated by 3D printing so that the thickness ratio of the first layer was 30%. A fused deposition modeling (FDM) 3D printer was used, with a layer pitch of 0.2 mm and an infill pattern set to a grid. (Example 9) Using polyester copolymer 1, an implant medical device was fabricated with a length and width of 35 mm, a thickness of 2 mm, and a bulk density of the first layer of 0.20 g / cm. ３ , the bulk density of the second layer is 1.10 g / cm ３An implant medical device was fabricated by 3D printing so that the thickness ratio of the first layer was 30%. A fused deposition modeling (FDM) 3D printer was used, with a layer pitch of 0.2 mm and an infill pattern set to a grid. (Example 10) Using polyester copolymer 1, an implant medical device with a length and width of 35 mm, a thickness of 2 mm, and a bulk density of the first layer of 0.15 g / cm was fabricated. ３ , the bulk density of the second layer is 1.10 g / cm ３ The implant medical device was fabricated by 3D printing so that the thickness ratio of the second layer was 10%. The device was fabricated using a fused deposition modeling (FDM) 3D printer with a layer pitch of 0.2 mm and an infill pattern set to a grid.

[0121] The implantable medical devices obtained in Examples 1 to 10 and Comparative Examples 1 and 2 were subjected to the measurements described in Measurement Examples 5 to 10. The results are shown in Table 2.

[0122]

[0123] The results of Example 1 and Comparative Example 1 revealed that by making the thickness of the implant medical device 1.0 mm or more, a maximum tensile load suitable for soft tissue repair applications can be obtained. Furthermore, the results of Example 1 and Comparative Example 2 also revealed that by making the bulk density of the implant medical device 0.40 g / cm ３ It has become clear that by doing the above, a maximum tensile load suitable for use in repairing soft tissue can be obtained.

[0124] The implant medical device of the present invention can be used for repairing soft tissues, particularly as a repair material for cartilage, tendons, rotator cuffs, and ligaments.

[0125] 1... Application surface, 2... First layer, 3... Second layer, 4... Maximum thickness, 5... Minimum thickness

Claims

1. A polyester copolymer containing polyhydroxycarboxylic acid having two types of hydroxycarboxylic acid monomers as main structural units, having a thickness of 1.0 mm or more and 20.0 mm or less, and a bulk density of 0.40 g / cm ３ 1.50g / cm or more ３ An implantable medical device for soft tissue repair, which is:

2. A composite material having a first layer and a second layer, the bulk density of the first layer being 0.20 g / cm ３ 0.95g / cm or more ３ 2. The implantable medical device of claim 1, wherein:

3. A composite material having a first layer and a second layer, the bulk density of which is different from that of the first layer and the second layer, a thickness of 1.0 mm or more and 20.0 mm or less, and an overall bulk density of 0.40 g / cm ３ 1.50g / cm or more ３ An implantable medical device for soft tissue repair, which is:

4. The implant medical device according to claim 3, which has a polyester copolymer containing polyhydroxycarboxylic acid having two types of hydroxycarboxylic acid monomers as main building blocks.

5. The bulk density of the first layer is 0.20 g / cm ３ 0.95g / cm or more ３ 5. The implant medical device according to claim 3 or 4, wherein:

6. The difference between the bulk density of the first layer and the bulk density of the second layer is 0.10 g / cm ３ 0.90g / cm or more ３ The implant medical device according to any one of claims 2 to 5, wherein:

7. The bulk density of the second layer is 0.95 g / cm ３ Greater than 2.00 g / cm ３ The implant medical device according to any one of claims 2 to 6, which is as follows:

8. An implant medical device according to any one of claims 2 to 7, wherein the ratio of the thickness of said second layer to the thickness of said implant medical device is 5% or more and 40% or less.

9. The implantable medical device according to any one of claims 2 to 8, wherein the first layer and the second layer are composed of the same component.

10. The implant medical device according to any one of claims 1 to 9, having a maximum tensile load of 20 N or more and 2,000 N or less.

11. The implant medical device according to any one of claims 1 to 10, having a tensile modulus of elasticity of 0.01 MPa or more and 100 MPa or less.

12. The implantable medical device according to any one of claims 1 to 11, which is biodegradable.

13. The implant medical device according to any one of claims 1, 2, and 4 to 12, wherein the R value of the polyhydroxycarboxylic acid, represented by the following formula 1, is 0.25 or more and 0.99 or less, when the two types of hydroxycarboxylic acid monomers are "monomer A" and "monomer B", respectively. R value = [AB] / (2[A][B]) x 100 ... formula 1 [A]: Molar fraction (%) of monomer A residues in polyhydroxycarboxylic acid [B]: Molar fraction (%) of monomer B residues in polyhydroxycarboxylic acid [AB]: Molar fraction (%) of structures (A-B and B-A) in which monomer A residues and monomer B residues are adjacent to each other in polyhydroxycarboxylic acid 14. The implant medical device according to claim 13, wherein said monomer A is a monomer selected from the group consisting of lactic acid, glycolic acid, dilactide and glycolide, and said monomer B is a monomer selected from the group consisting of hydroxyvaleric acid, hydroxycaproic acid, valerolactone and caprolactone.

15. The implantable medical device according to any one of claims 1 to 14, which consists solely of polyester copolymer.

16. An implantable medical device according to any one of claims 1 to 15, wherein the soft tissue is a tendon, a rotator cuff or a ligament.

17. An implant medical device for repairing ruptured soft tissue according to any one of claims 1 to 16, which is fixed to muscle and bone or bone and bone.

18. A method for repairing a ruptured tendon, rotator cuff or ligament, comprising a fixing step of fixing the implant medical device according to any one of claims 1 to 17 to muscle and bone or bone and bone.

19. A method for producing an implant medical device according to any one of claims 1 to 17.