Bioabsorbable fibrous medical materials

Abstract

The present invention addresses the problem of providing a bioabsorbable stretchable fiber-like medical material that requires only a small force to achieve ligation that gives small knots and is less likely to loosen. The present invention provides a fiber-like medical material that comprises a molded article obtained by spinning and stretching a bioabsorbable aliphatic polymer, has an elongation at break of 75% or more, has a lower intermediate tensile elasticity at a strain in the range of 0.25-10% than an initial tensile elasticity at a strain in the range of 0.05-0.25%, the intermediate tensile elasticity being 400 MPa or less, and has a residual strain ratio after 100% deformation of 70% or less.

Description

[Technical Field]

[0001] The present invention relates to a fibrous medical material that uses stretchable and bioabsorbable aliphatic polymer fibers and that allows for ligation in which knots are easily formed, the knots are small, and the knots are difficult to unravel. [Background technology]

[0002] Sutures are classified into monofilament sutures, which are made of a single fiber, and multifilament sutures, which are made of multiple fibers. Sutures are made from either non-absorbable or absorbable polymers. Non-absorbable polymers include polyethylene, polypropylene, nylon, silicone, Teflon, and silk. Absorbable polymers include homopolymers or copolymers made from glycolic acid, lactic acid, ε-caprolactone, or dioxanone. However, it is known that sutures containing glycolic acid or lactic acid tend to induce a strong inflammatory response during the absorption process, which can pose a problem in terms of biocompatibility.

[0003] Conventional bioabsorbable sutures include those made of rigid polymers such as polyglycolic acid (PGA) and poly(glycolide / L-lactide) copolymer, which are used as multifilaments, and those made of copolymers such as poly(glycolide / trimethylene carbonate) copolymer, poly(glycolide / ε-caprolactone) copolymer, poly-p-dioxanone, poly(glycolide / trimethylene carbonate / p-dioxanone) copolymer, poly(glycolide / trimethylene carbonate / ε-caprolactone) copolymer, poly(glycolide / L-lactide / trimethylene carbonate / ε-caprolactone) copolymer, and poly(L-lactide / ε-caprolactone) copolymer, which are used as monofilaments due to their increased flexibility. These sutures are used according to the required strength, duration of tensile strength maintenance, absorption period, application site, tissue reaction, tissue damage potential, required elasticity, knot safety, operability, cost efficiency, infection resistance, and the physician's experience.

[0004] For example, Patent Document 1 describes a surgical suture composed of a monofilament thread made of a copolymer of lactic acid and ε-caprolactone. Patent Document 2 describes a suture obtained by melt-spinning a glycolide / ε-caprolactone copolymer. Patent Document 3 describes a monofilament suture obtained by co-spinning a first polymer and a second polymer synthesized from one or more monomers selected from the group consisting of glycolide, glycolic acid, lactide, lactic acid, caprolactone, dioxanone, trimethylene carbonate, and ethylene glycol, wherein the Young's moduli of the first polymer and the second polymer are 3.0 GPa or less. Patent Document 4 describes a synthetic composite biomaterial containing collagen, at least one organic polymer (such as polyglycolide, polylactide, glycolide-lactide copolymer, polylactone, and polyhydroxyalkanoic acid), and at least one active ingredient.

[0005] Sutures made of polyhydroxyalkanoates (sometimes referred to as PHA) have also been reported. For example, Patent Document 5 describes a polyester molded article containing a biodegradable polyester copolymer consisting of 3-hydroxybutyrate (sometimes referred to as 3HB) units and 4-hydroxybutyrate (sometimes referred to as 4HB) units, with the 4-hydroxybutyrate unit content being greater than 60 mol% and less than 95 mol%. Patent Documents 6 and 7 describe medical devices containing sutures made of biocompatible polyhydroxyalkanoates. Patent Document 8 describes fibers containing poly-4-hydroxybutyrate polymers and having a tensile strength greater than 126 MPa. Patent Documents 9 and 10 describe polymer products obtained by drawing a composition characterized by a biodegradable polyhydroxyalkanoate copolymer containing at least two randomly repeating monomer units. Patent Document 11 describes a polyester copolymer consisting of 97-40 mol% 3-hydroxybutyrate units and 3-60 mol% 4-hydroxybutyrate units, with an [η] measured in chloroform at 30°C in the range of 0.4-10.0 dL / g, and describes that the polyester copolymer has excellent flexibility and good moldability, and that the resulting molded products such as fibers and films are supple and strong. However, none of the above prior art documents describes that fibers with small knots and less tendency to unravel can be obtained by making the fibrous medical material have an elongation at break equal to or greater than a predetermined value and a modulus of elasticity (initial tensile modulus and intermediate tensile modulus, described below) equal to or less than a predetermined value.

[0006] Furthermore, a monofilament suture (MonoMax®) made from a homopolymer of 4-hydroxybutyric acid (also referred to as P(4HB)) has been developed (Patent Document 12, and Non-Patent Documents 1 and 2). This MonoMax suture has been reported to have a modulus of elasticity of 485 MPa (Non-Patent Document 2), which is lower than the modulus of elasticity of PDSII made from poly-p-dioxanone, which is 1370 MPa, and Monocryl made from poly(glycolide / ε-caprolactone) copolymer, which is 725 MPa. Therefore, it is considered a soft suture. However, when actually used by doctors, the suture is stiff and knots tend to loosen, necessitating the need to tie more knots. [Prior art documents] [Patent documents]

[0007] [Patent Document 1] Japanese Patent Application Laid-Open No. 2001-149462 [Patent Document 2] Japanese Patent Application Laid-Open No. 2011-6496 [Patent Document 3] Patent No. 4071661 [Patent Document 4] Special Publication No. 2019-505338 [Patent Document 5] Japanese Patent Application Publication No. 06-336523 [Patent Document 6] U.S. Patent No. 6,867,247 [Patent Document 7] Patent No. 5031144 [Patent Document 8] Special Publication No. 2007-525601 [Patent Document 9] Special Publication No. 2003-513130 [Patent Document 10] Special Publication No. 2003-513131 [Patent Document 11] Japanese Patent Application Publication No. 1-48821 [Patent Document 12] International Publication No. WO2004 / 101002 [Non-patent literature]

[0008] [Non-Patent Document 1] BMC Surgery volume 8, Article number: 12 (2008) [Non-patent document 2] International Journal of Polymer Science Volume 2012, Article ID 216137 Summary of the Invention [Problem to be solved by the invention]

[0009] When using sutures in surgery, tying and tying knots are fundamental techniques. The ease of tying knots and the resistance to untangling are characteristics required of sutures, and these are required not only for non-absorbable sutures but also for bioabsorbable sutures.

[0010] Multifilament sutures are made by weaving together fine fibers and have an uneven surface. This gives them the advantage of knot resistance and flexibility. However, they tend to be highly invasive when passing through tissue and have poor knot lubrication, i.e., a high coefficient of friction makes knots difficult to slip. Furthermore, the formation of tiny capillaries increases the risk of infection compared to monofilament sutures. While monofilament sutures have a smooth surface, which minimizes tissue invasiveness and infection resistance, they also lack flexibility and are prone to loosening. While monofilament sutures can be made to loosen more easily by tying more knots, this can result in larger knots and increased tissue damage. Furthermore, traditional monofilament sutures are rigid, impliable, and incompliant, leading surgeons to often use excessive force to securely tie the knot. Therefore, when the suture is tightened with a strong force to form a knot, excessive force is applied to the tissue at the suture site, raising concerns about unintended tissue damage. Therefore, there is a need for a suture that is highly extensible, pliable, and resistant to knot unraveling, so that an appropriate tension can be uniformly maintained until the tissue self-repairs while minimizing tissue invasion, as well as a suture that can form a knot that is difficult to unravel without tightening with a strong force. Furthermore, there is an even greater need for a bioabsorbable suture that is easy to form a knot, produces a small knot, and is resistant to unraveling, eliminating the need for suture removal or re-incision, and reducing the foreign body sensation felt by the embedded knot in surrounding tissues, such as in subcutaneous sutures and internal sutures. The fact that the knot is small and difficult to untie means that there is less need to leave the cut end of the thread long when cutting off the excess thread from the knot, and if the thread is flexible, it is expected that the cut surface of the thread will be less likely to irritate the tissue due to its rigidity. Furthermore, tissues can swell for a variety of reasons, but existing sutures lack elasticity and are therefore unable to properly adapt to swelling, placing excess tension on the tissue and potentially causing scarring. A suture with a lower elastic modulus than conventional sutures that more closely matches the elasticity of tissue, meaning it can stretch and then shrink, would be able to stretch and distribute tension even when tissue swells, and would shrink when the swelling subsides, allowing it to continue to contribute to wound adhesion. There is also a need for the development of bioabsorbable sutures.

[0011] The present invention aims to provide a bioabsorbable fibrous medical material that can form small knots that are difficult to unravel with little force. Another objective is to provide a bioabsorbable fibrous medical material that is stretchable and can follow the movement of tissue. [Means for solving the problem]

[0012] As a result of extensive research to solve the above problems, the present inventors have found that by using a bioabsorbable aliphatic polymer as a raw material and setting the elongation at break of a molded product obtained by spinning and drawing the bioabsorbable aliphatic polymer to 75% or more, setting the intermediate tensile modulus to a value lower than the initial tensile modulus, setting the intermediate tensile modulus to 400 MPa or less, and setting the residual strain rate after 100% deformation to 70% or less, it is possible to provide a fibrous medical material that is both bioabsorbable and stretchable, which allows for ligation that is easy to form knots, small knots that are difficult to unravel, and which reduces the number of knots that can be tied, and has thus completed the present invention.

[0013] <1> A fibrous medical material comprising a molded product obtained by spinning and drawing a bioabsorbable aliphatic polymer, the fibrous medical material having an elongation at break of 75% or more, an intermediate tensile modulus between 0.25% and 10% strain that is lower than the initial tensile modulus between 0.05% and 0.25%, the intermediate tensile modulus being 400 MPa or less, and a residual strain rate after 100% deformation being 70% or less. <2> The initial tensile modulus is 1000 MPa or less. <1> The fibrous medical material according to claim 1. <3> The initial tensile modulus is 480 MPa or less. <1> or <2> The fibrous medical material according to claim 1. <4> The intermediate tensile modulus is 300 MPa or less. <1> from <3> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <5> The residual strain rate after 100% deformation is 50% or less. <1> from <4> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <6> Porosity is 0% to 55%. <1> from <5> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <7> The diameter of the largest pore (pore, void, gap, or hollow) measured by microscopic observation of a cross section perpendicular to the fiber axis direction is 100 μm or less. <1> from <6> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <8> The ratio of the major axis length to the minor axis length in the width direction cross section (major axis length / minor axis length) is 1.0 or more and 3.0 or less; <1> from <7> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <9> The bioabsorbable aliphatic polymer is an aliphatic polyester. <1> from <8> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <10> The bioabsorbable aliphatic polymer is a polyhydroxyalkanoate. <1> from <9> 10. The fibrous medical material according to claim 9, wherein the fibrous medical material is a fibrous medical material. <11> The polyhydroxyalkanoate is a polyhydroxyalkanoate composed of two or more types of hydroxyalkanoate units. <10> The fibrous medical material according to claim 1. [Effects of the Invention]

[0014] The bioabsorbable fibrous medical material of the present invention is easy to handle, allows knots to be formed with little force, and enables ligation with small knots that are difficult to unravel. [Brief explanation of the drawings]

[0015] [Figure 1] FIG. 1 shows a surgeon's knot. [Figure 2]Figure 2 shows a single knot in a surgeon's knot. [Figure 3] Figure 3 shows a surface SEM image of P(3HB-co-4HB) suture 2.5-0, where P(3HB-co-4HB) stands for copolymer of 3-hydroxybutyric acid and 4-hydroxybutyric acid. [Figure 4] Figure 4 shows a SEM photograph of a P(3HB-co-4HB) suture 3-0 surgical knot. [Figure 5] Figure 5 shows a stereomicrograph of a 3-0 size surgical knot of opaque P(3HB-co-4HB) suture. [Figure 6] Figure 6 shows a stereomicrograph of a colorless P(3HB-co-4HB) suture 3-0 size surgical knot. [Figure 7] Figure 7 shows a stereomicrograph of a colorless P(3HB-co-4HB) suture size 1 surgical knot. [Figure 8] Figure 8 shows a stereomicrograph of a white P(3HB-co-4HB) suture 2.5-0 size surgical knot. [Figure 9] FIG. 9 shows a surface SEM photograph of P(4HB) MonoMax® suture 2-0. [Figure 10] FIG. 10 shows a SEM photograph of a P(4HB) MonoMax® suture 2-0 surgeon's knot. [Figure 11] FIG. 11 shows a stereomicrograph of a P(4HB) MonoMax® suture 2-0 surgeon's knot. [Figure 12] FIG. 12 shows a stereomicrograph of a PDSII suture 3-0 surgical knot. [Figure 13] FIG. 13 shows a stereomicrograph of a PDSII suture 4-0 surgical knot. [Figure 14] Figure 14 shows the relationship between the immersion time in buffer solution and the tensile breaking strength and elongation at break of P(3HB-co-4HB) suture. [Figure 15] FIG. 15 shows the decrease in weight average molecular weight Mw of P(3HB-co-4HB) suture with the duration of immersion in buffer solution. [Figure 16]Figure 16 shows the relationship between the implantation period of P(3HB-co-4HB) sutures in rats and the tensile breaking strength and breaking elongation. [Figure 17] FIG. 17 shows the decrease in the weight-average molecular weight Mw of P(3HB-co-4HB) suture over the period of implantation in the rat body. [Figure 18] FIG. 18 shows the state of the sutured area 7 weeks after suturing with bioabsorbable sutures on a microminipig. [Figure 19] FIG. 19 shows HE staining of tissue at the P(3HB-co-4HB) suture. [Figure 20] FIG. 20 shows HE staining of tissue at the site of polyglyconate (PGA) suture. [Figure 21] FIG. 21 shows HE staining of tissue at the P(4HB) suture site. [Figure 22] FIG. 22 shows an example of a stress-strain curve of a tensile test of the fiber of Example 1 up to break (measured with a chuck distance of 1 cm). [Figure 23] FIG. 23 shows an example of a stress-strain curve for a tensile test of the fiber of Example 2 up to break. [Figure 24] FIG. 24 shows an example of a stress-strain curve for a tensile test of the fiber of Example 3 up to break. [Figure 25] FIG. 25 shows an example of a stress-strain curve for a tensile test of the fiber of Example 4 up to break. [Figure 26] FIG. 26 shows the stress-strain curves of 3 cm of the P(3HB-co-4HB) suture of Example 1 cyclically tested to 100% strain. [Figure 27] FIG. 27 shows the stress-strain curves of 3 cm of the P(3HB-co-4HB) suture of Example 2 cyclically tested to 100% strain. [Figure 28] FIG. 28 shows the stress-strain curves of 3 cm of the P(3HB-co-4HB) suture of Example 3 cyclically tested to 100% strain. [Figure 29] FIG. 29 shows the stress-strain curves of 3 cm of the P(3HB-co-4HB) suture of Example 4 cyclically tested to 100% strain. [Figure 30] FIG. 30 shows the stress-strain curves of 12 cm of P(3HB-co-4HB) suture of Example 4 cyclically tested to 50% strain. [Figure 31] FIG. 31 shows an example of a cross section of the yarn of Example 1. [Figure 32] FIG. 32 shows an example of a cross section of the yarn of Example 1. [Figure 33] FIG. 33 shows an example of a cross section of the yarn of Example 2. [Figure 34] FIG. 34 shows an example of a cross section of the yarn of Example 3. [Figure 35] FIG. 35 shows a surface SEM photograph of the P(3HB-co-4HB) suture of Example 4. [Figure 36] FIG. 36 shows a cross-sectional SEM photograph of the P(3HB-co-4HB) suture of Example 4. [Figure 37] FIG. 37 shows a SEM photograph of the surgeon's knot of the P(3HB-co-4HB) suture of Example 4. DETAILED DESCRIPTION OF THE INVENTION

[0016] The present invention will be described in detail below. Note that the following embodiments are merely examples for explaining the present invention, and the present invention is not limited to these embodiments.

[0017] The bioabsorbable fibrous medical material of the present invention comprises a molded article obtained by spinning and stretching a bioabsorbable aliphatic polymer.

[0018] <Bioabsorbable aliphatic polymer> Bioabsorbability refers to the property of being naturally decomposed by hydrolysis or enzymatic reaction after being placed inside or outside the body, and the decomposition products are eliminated by metabolism or excretion. In other words, bioabsorbability means that the substance can be locally eliminated and excreted outside the body.

[0019] The term "locally disappearing" refers to the property of being decomposed and disappearing from the applied site within a predetermined number of days (e.g., 360 days, 240 days, 120 days, 60 days, or 30 days) under physiological conditions. As an example of local disappearance, a sample equivalent to a polymer concentration of at least 1% by mass is placed in physiological saline (pH 4-8) at 37°C, mixed with a rotor mixer, and visually observed. If the sample loses its shape and becomes a transparent aqueous solution within a predetermined number of days, the sample can be determined to have local disappearance. Alternatively, a sample can be implanted in the body and disappears within a predetermined number of days. A sample can also be determined to have local elimination properties if it is decomposed and disappears. "Excretable from the body" means that after a material has disappeared from the application site, it can be excreted from the body without excessive accumulation in organs such as the kidneys or liver. For example, if a material decomposes to a molecular weight of 70,000 or less, or in some cases 40,000 or less, the sample can be determined to have excretable from the body. Alternatively, after the material has disappeared from the application site, it may become partially decomposed or become a low-molecular-weight compound, which may then be metabolized to water or carbon dioxide and utilized in the body or excreted externally.

[0020] Examples of aliphatic polymers include aliphatic polyesters, polyamides such as nylon, polystyrene, polyvinyl alcohol, poly(ethylene-co-vinyl acetate), poly(hydroxyethyl methacrylate), and other polyolefins, acid-modified polyolefins (maleic anhydride-grafted polyethylene, maleic anhydride-grafted polypropylene, and the like), ethylene-vinyl compound copolymers (ethylene-vinyl acetate copolymer, ethylene-vinyl alcohol copolymer, ethylene-vinyl chloride copolymer, ethylene-(meth)acrylic acid copolymer and its ionic cross-linked products (ionomers), ethylene-methyl methacrylate copolymer, and the like), polyvinyl compounds (polymethyl acrylate, polymethyl methacrylate, and the like), polycarbonate, and polyethers (polyethylene oxide, and the like). Of the above, aliphatic polyesters are preferred in terms of bioabsorbability.

[0021] Bioabsorbable aliphatic polyesters include, but are not limited to, polyesters selected from the group consisting of polyglycolic acid, polylactic acid (D, L, DL), polyε-caprolactone, polyhydroxybutyrate, polyhydroxyvalerate, polyorthoester, polyhydroxyhexanoate, polybutylene succinate, polyhydroxyalkanoates other than those mentioned above, poly-p-dioxanone, and copolymers thereof. Monomer units in the copolymers include, but are not limited to, glycolic acid, lactic acid (D, L, DL), ε-caprolactone, hydroxybutyrate, hydroxyvalerate, orthoester, hydroxyhexanoate, butylene succinate, hydroxyalkanoates other than those mentioned above, 1,3-dioxane-2-one (trimethylene carbonate), and p-dioxanone. Polyhydroxyalkanoates are particularly preferred as bioabsorbable aliphatic polymers.

[0022] <Polyhydroxyalkanoate> The polyhydroxyalkanoate used in the present invention is biodegradable and bioabsorbable, and can have repeated stretchability (the property of being repeatedly stretched and contracted). It is preferable that the polymerized units contain 3-hydroxybutyrate units. In the present invention, it is preferable to use a polyhydroxyalkanoate composed of two or more types of hydroxyalkanoate units. That is, the polyhydroxyalkanoate used in the present invention is preferably a copolymer polyester containing two or more types of monomer units. It is preferable that the polymerized units contain 4-hydroxybutyrate (sometimes referred to as 4HB) units in addition to 3-hydroxybutyrate (sometimes referred to as 3HB) units.

[0023] The weight-average molecular weight of the polyhydroxyalkanoate measured by gel permeation chromatography relative to polystyrene standards is preferably 100,000 or more, more preferably 200,000 or more, and may even be 300,000 or more, 400,000 or more, or 500,000 or more. The weight-average molecular weight measured by gel permeation chromatography relative to polystyrene standards may be 600,000 or more, 700,000 or more, 800,000 or more, 900,000 or more, 1,000,000 or more, 1,100,000 or more, 1,200,000 or more, 1,300,000 or more, 1,400,000 or more, 1,500,000 or more, 2,000,000 or more, 3,000,000 or more, or 4,000,000 or more. The upper limit of the weight-average molecular weight measured by gel permeation chromatography relative to polystyrene standards is not particularly limited, but is generally 20,000,000 or less, and may be 10,000,000 or less, 8,000,000 or less, 7,000,000 or less, 6,000,000 or less, 5,000,000 or less, 4,000,000 or less, or 3,000,000 or less. However, when melt molding is performed, taking into consideration the reduction in molecular weight due to thermal decomposition and the viscosity during melting to be not too high, the weight average molecular weight as measured by gel permeation chromatography in terms of polystyrene is desirably 400,000 or more and 2,500,000 or less, more preferably 500,000 or more and 2,200,000 or less, and even more preferably 600,000 or more and 2,000,000 or less.

[0024] The polyhydroxyalkanoate used in the present invention preferably contains 3-hydroxybutyrate (3HB) units as polymerized units, and more preferably contains 3-hydroxybutyrate (3HB) units and 4-hydroxybutyrate (4HB) units. When the polyhydroxyalkanoate contains 3HB units, or when the polyhydroxyalkanoate contains 3HB units and 4HB units, it may contain polymerized units other than 3HB units and 4HB units as polymerized units. Examples of the other polymerized units include lactate (LA), glycolate (GA), 3-hydroxypropionate (3HP), 3-hydroxyvalerate (3HV), 4-hydroxyvalerate (4HV), 5-hydroxyvalerate (5HV), 4-hydroxyhexanoate (4HH), 5-hydroxyhexanoate (5HH), 6-hydroxyhexanoate (6HH), 3-hydroxyhexanoate (3HH), and hydroxyalkanoates having 7 or more carbon atoms. Terpolymers and multipolymers containing the above polymerized units can also be used instead of binary copolymers. The above copolymer compositions can also be mixed (blended) in any ratio and used.

[0025] In the present invention, the 3-hydroxybutyrate unit and the 4-hydroxybutyrate unit are represented by the following formulas, respectively. 3-hydroxybutyrate unit: -OCH(CH3)CH2C(=O)- 4-hydroxybutyrate unit: -OCH2CH2CH2C(=O)-

[0026] When the polyhydroxyalkanoate contains 4-hydroxybutyrate units, the proportion of the 4-hydroxybutyrate units relative to all monomer units is preferably 5 mol% to 40 mol%, more preferably 10 mol% to 40 mol%, and even more preferably 10 mol% to 30 mol%. The proportion of the 4-hydroxybutyrate units relative to all monomer units may be 5 mol% or more, 6 mol% or more, 7 mol% or more, 8 mol% or more, 9 mol% or more, 10 mol% or more, 11 mol% or more, 12 mol% or more, 13 mol% or more, 14 mol% or more, 15 mol% or more, or 16 mol% or more, or even 17 mol% or more, 18 mol% or more, 19 mol% or more, or 20 mol% or more. The ratio of 4-hydroxybutyrate units to all monomer units may be 40 mol% or less, 39 mol% or less, 38 mol% or less, 37 mol% or less, 36 mol% or less, 35 mol% or less, 34 mol% or less, 33 mol% or less, 32 mol% or less, 31 mol% or less, 30 mol% or less, 29 mol% or less, 28 mol% or less, 27 mol% or less, 26 mol% or less, 25 mol% or less, 24 mol% or less, 23 mol% or less, 22 mol% or less, or 21 mol% or less.

[0027] The polyhydroxyalkanoates of the present invention may be any of random polymers, block polymers, alternating polymers, or graft polymers, but are preferably random polymers.

[0028] [Method for producing polyhydroxyalkanoate] Generally, methods for synthesizing polyhydroxyalkanoates include fermentation synthesis (biosynthesis) and chemical synthesis. The method for producing the polyhydroxyalkanoate used in the present invention may be either fermentation synthesis (biosynthesis) or chemical synthesis, but the fermentation synthesis (biosynthesis) is preferred in order to obtain a polyhydroxyalkanoate with a large molecular weight.

[0029] Chemical synthesis is a method of chemically synthesizing polyhydroxyalkanoates according to conventional organic synthesis techniques. Specific examples of chemical synthesis include P(3HB-co-6HHx) synthesized by catalytic ring-opening polymerization of fatty acid lactones such as (R)-β-butyrolactone and ε-caprolactone (Abe et al., Macromolecules, 28, 7630 (1995)), P(3HB-co-4HB) synthesized by catalytic ring-opening polymerization of fatty acid lactones such as (R)-β-butyrolactone and γ-butyrolactone (Hori et al., Polymer, 36, 4703 (1995)), and P(GA-co-LA) synthesized by catalytic ring-opening polymerization of glycolide and lactide (Gilding et al., Polymer, 20, 1459 (1979)). However, because a catalyst such as stannous octoate is used, careful attention must be paid to the amount of catalyst used in the chemical synthesis of bioabsorbable polyhydroxyalkanoates.

[0030] Fermentation synthesis (biosynthesis) is a method for biosynthesizing PHAs according to conventional culture engineering techniques. Specifically, the 4HB-containing polyhydroxyalkanoates used in the present invention can be produced by fermentation synthesis using a microorganism capable of producing P(3HB) as a carbon source. Examples of such carbon sources include ε-caprolactone (also known as 6-hexanolactone) or its saponification product, 6-hydroxyhexanoate, or a salt thereof; γ-butyrolactone or its saponification product, 4-hydroxybutyrate, or a salt thereof; butyric acid derivatives such as 4-chlorobutyrate and 4-bromobutyrate; and even-chain α,ω-alkanediols having 4 or more carbon atoms, such as 1,4-butanediol, 1,6-hexanediol, 1,8-octanediol, 1,10-decanediol, and 1,12-dodecanediol (Saito et al., Polymer International 39, 169 (1996) and International Publication WO2019 / 044837). By appropriately changing the type of carbon source used and the supply ratio, P(3HB-co-4HB) copolymers with various 4HB ratios can be produced. PHAs obtained by biological enzymatic reactions in this way do not contain metal catalysts such as stannous octoate, as is the case with chemical synthesis, and in this sense, they have the advantage. Fermentation synthesis can be performed using either genetically modified or non-genetically modified bacteria.

[0031] As a method for extracting polyhydroxyalkanoate from bacterial cells, a solvent extraction method in which extraction is performed using a halogenated hydrocarbon solvent such as chloroform followed by precipitation with a poor solvent such as hexane or methanol may be used, as is already known, or an aqueous extraction method may be used as described in JP-B 04-061638, JP-A 07-177894, and WO 2004029266. When producing polyhydroxyalkanoate for use in living organisms, any polyhydroxyalkanoate purification process may be used as long as it is capable of removing impurities such as bacterial cell-derived proteins and endotoxins, and the purified polyhydroxyalkanoate may be depyrogenated using peroxide as described in US Pat. No. 6,245,537.

[0032] <Molecular weight measurement of polyhydroxyalkanoate (gel permeation chromatography (GPC) method)> The molecular weight of the polyhydroxyalkanoate can be measured by gel permeation chromatography as follows. Chloroform was added so that the polyhydroxyalkanoate concentration was approximately 0.5 mg / ml, and the solution was dissolved at 60°C for 2 to 4 hours. The solution was then returned to room temperature and filtered through a PTFE filter with a pore size of 0.2 μm to remove insoluble matter, giving a measurement sample. The GPC conditions were as follows:

[0033] Equipment: Shimadzu HPLC Prominence system Column: Showa Denko Shodex K-806L (two columns in series) Column temperature: 40℃ Mobile phase: chloroform (1 ml / min) Detector: RI (40℃) Standard: Shodex polystyrene molecular weight standard (6.87 million to 1270) Injection volume: 60μl Analysis time: 30 minutes

[0034] <Production of molded articles from bioabsorbable aliphatic polymers such as polyhydroxyalkanoates> The method for producing a molded article (such as an elastic suture) from a bioabsorbable aliphatic polymer such as polyhydroxyalkanoate, which has a slow crystallization rate after melting, is not particularly limited, as long as the polymer is stretched to a state in which the bioabsorbable aliphatic polymer, such as polyhydroxyalkanoate, has solidified or crystallized to the extent that it can be stretched from an amorphous state and fine crystals are present, thereby imparting elasticity. For example, molded articles (such as an elastic suture) can be produced in accordance with the methods described in Japanese Patent Application Nos. 2019-90739, 2020-096144, and 2020-096145.

[0035] Specifically, an elastic suture can be produced by melt-molding a bioabsorbable aliphatic polymer such as polyhydroxyalkanoate, followed by stretching in a state in which microcrystals are formed by allowing an appropriate waiting time at room temperature or a temperature at which crystallization proceeds to promote crystallization, or by melt-molding in a partially molten state and stretching in a state in which microcrystals remain. When melt-molding a bioabsorbable aliphatic polymer such as polyhydroxyalkanoate, further additives may be added as long as they do not impair the effects of the present invention.

[0036] Examples of additives include one or more selected from antioxidants, heat stabilizers (e.g., hindered phenols, hydroquinone, phosphites, and substituted derivatives thereof), ultraviolet absorbers (e.g., resorcinol, salicylate), color inhibitors (phosphites, hypophosphites, etc.), lubricants, mold release agents (montanic acid and its metal salts, its esters, its half esters, stearyl alcohol, stearamide, polyethylene wax, etc.), colorants (dyes or pigments, etc.), conductive agents or carbon black as a colorant, plasticizers, flame retardants (bromine-based flame retardants, phosphorus-based flame retardants, red phosphorus, silicone-based flame retardants, etc.), flame retardant assistants, and antistatic agents.

[0037] The method for incorporating an additive into a bioabsorbable aliphatic polymer such as polyhydroxyalkanoate is not particularly limited, and examples include dry blending, solution blending, and addition during chemical polymerization of the bioabsorbable aliphatic polymer such as polyhydroxyalkanoate.

[0038] Bioabsorbable aliphatic polymers such as polyhydroxyalkanoates can be subjected to known melt molding methods such as injection molding, injection compression molding, compression molding, extrusion molding (melt extrusion molding), blow molding, press molding, and spinning (melt extrusion spinning). Spinning (melt extrusion spinning, partial melt extrusion spinning) is preferred, and it is preferable that a drawing operation is performed to impart stretchability. For melt extrusion spinning, a small-sized plunger-type melt extrusion device at the laboratory level can be used, but a large-sized device such as an industrially used single-screw spinning device or twin-screw spinning device can also be used. The number of melt molding cycles is not particularly limited, but preferably can be performed only once. Furthermore, prior to melt molding, a heat treatment mediated by gas, liquid, or solid may or may not be performed to arbitrarily adjust the primary lamellar thickness of the raw material bioabsorbable aliphatic polymer, such as polyhydroxyalkanoate (Patent Application No. 2020-96145).

[0039] In the present invention, the solidification step after melting can be carried out in a mold, in air, or in liquid (e.g., water). That is, solidification can be achieved by cooling a molten bioabsorbable aliphatic polymer such as polyhydroxyalkanoate in a mold, in air, or in water. Preferably, a molten bioabsorbable aliphatic polymer such as polyhydroxyalkanoate can be cooled in a mold or in air. When cooling in air, the temperature and humidity of the air can be controlled, but cooling at room temperature without special temperature control is also possible. Furthermore, a gas with a modified composition of air (e.g., nitrogen, oxygen, carbon dioxide, water concentration) can also be used, or an environment in which a rare gas (e.g., helium, neon, argon) is added or circulated can also be used. When cooling in liquid, the temperature and components of the liquid (e.g., water, alcohols, glycerol) can be optionally modified. Bioabsorbable aliphatic polymers such as polyhydroxyalkanoates in a molten, amorphous state are in solution at temperatures above the glass transition point, and because the crystallization rate is slow, if they are stretched as is, the molecules will flow and quickly break. However, if they are partially crystallized and stretched in a solidified state, the polymer chains will become oriented and they will be able to be stretched.

[0040] Molded articles of bioabsorbable aliphatic polymers such as polyhydroxyalkanoates produced by the method of the present invention include various fibers such as drawn yarns and ultra-drawn yarns, one example of which is sutures. Sutures may be either monofilament sutures consisting of a single fiber or multifilament sutures consisting of multiple fibers. When producing sutures, the diameter of the suture is not particularly limited, but is generally 1 mm or less, and may be 0.8 mm or less, 0.6 mm or less, 0.5 mm or less, 0.4 mm or less, or 0.3 mm or less, with the lower limit of the diameter generally being 0.001 mm or more. The resulting fibers may be woven into a mesh or a three-dimensional structure. Furthermore, they may be combined with other fibers or materials to produce composite fibers or three-dimensional composites. However, in order for the resulting fiber to exhibit elasticity, it is desirable that the extruded fiber be stretched using methods such as melt spinning or partial melt spinning, so that the α crystals that form the lamellae in the fiber are oriented along the fiber direction, and that the lamellae made of α crystals are stacked perpendicular to the fiber axis. The amorphous layers and tie molecules between the lamellae are deformed by tension, increasing the amount of β crystals, which have an extended planar zigzag structure, and the β crystals decrease or disappear when the load is removed, allowing the fiber to exhibit elastic response (Patent Application No. 2019-90739). The obtained fibers exhibit stretchability without being subsequently heat-treated (annealed) at a temperature equal to or higher than the glass transition point but not melting the fibers, but may be heat-treated.

[0041] <Elongation at break> The elongation at break of the fibrous medical material of the present invention is 75% or more, preferably 100% or more, more preferably 150% or more, even more preferably 180% or more, and particularly preferably 200% or more. There is no particular upper limit to the elongation at break, but it is generally 1000% or less.

[0042] The elongation at break of the fibrous medical material is not particularly limited, but can be measured, for example, by the following method. Using a fibrous medical material with a length of 3 cm and a fiber diameter of approximately 0.1 to 0.4 mm and an AGS-50NX tensile tester (Shimadzu Corporation), a temperature of 23°C, a test speed of 10 mm / min, and an initial length (chuck distance) of 10 mm, the elongation at break can be measured by conducting a tensile test until the fiber breaks. If a sample of sufficient length is available, it is preferable to ensure a chuck distance of 10 cm.

[0043] <Porosity> The fibrous medical material of the present invention may or may not have voids. Preferably, the fibrous medical material of the present invention has voids. When the fibrous medical material of the present invention has voids, the range of porosity is not particularly limited, but is preferably 5 to 55%, more preferably 10 to 50%, and even more preferably 20 to 45%.

[0044] The method for measuring the porosity is not particularly limited, but for example, the porosity can be measured by observing the cross section of the fibrous medical material with a scanning electron microscope and performing image analysis of the cross section. Image analysis of the cross section can be performed using software such as ImageJ (an image processing program developed by the National Institutes of Health, USA), but is not limited to this.

[0045] <Initial tensile modulus> If the tensile modulus at a strain of 0.05% to 0.25% is defined as the initial tensile modulus, the initial tensile modulus of the fibrous medical material of the present invention is preferably 1000 MPa or less, more preferably 600 MPa or less, even more preferably 480 MPa or less, still more preferably 400 MPa or less, even more preferably 300 MPa or less, and particularly preferably 200 MPa or less. There is no particular lower limit for the initial tensile modulus, but it is generally 5 MPa or more, and may be 10 MPa or more.

[0046] <Intermediate tensile modulus> If the tensile modulus at strains of 0.25% to 10% is defined as the intermediate tensile modulus, in the present invention, the intermediate tensile modulus is set to a value lower than the initial tensile modulus. The intermediate tensile modulus of the fibrous medical material of the present invention is preferably 400 MPa or less, more preferably 300 MPa or less, even more preferably 250 MPa or less, even more preferably 200 MPa or less, and particularly preferably 150 MPa or less. There is no particular lower limit for the intermediate tensile modulus, but it is generally 5 MPa or more, and may be 10 MPa or more.

[0047] The elastic modulus can be measured, for example, using a tensile tester. The distance between the chucks of the tensile tester is 1 to 10 cm, and the specimen is fixed to a fixture with a distance of 1 cm above and below. The tensile speed is 10 mm / min. The initial tensile modulus can be calculated from the slope of the stress-strain curve between strains of 0.05% and 0.25%, and the intermediate tensile modulus can be calculated from the slope of the stress-strain curve between strains of 0.25% and 10%, for example. The strain range used for the intermediate tensile modulus may be any range as long as it is a range of strain larger than the strain range used to calculate the initial tensile modulus and is close to the strain used to calculate the initial tensile modulus. However, in this specification, the tensile modulus between strains of 0.25% and 10% is defined as the intermediate tensile modulus.

[0048] <Residual strain rate after 100% deformation> The residual strain rate of the fibrous medical material of the present invention after 100% deformation is 70% or less, preferably 60% or less, and more preferably 50% or less. There is no particular lower limit to the residual strain rate after 100% deformation, but it is generally 5% or more, and may be 10% or more, 20% or more, or 30% or more.

[0049] <Residual strain rate after 50% deformation> The residual strain rate of the fibrous medical material of the present invention after 50% deformation is preferably 40% or less, more preferably 30% or less, and even more preferably 20% or less. There is no particular lower limit to the residual strain rate after 50% deformation, but it is generally 5% or more, and may be 10% or more, 20% or more, or 30% or more.

[0050] The tensile elongation recovery rate is measured by using a tensile testing machine to perform a cycle test on a 3cm long fiber at a temperature of 23°C and an initial length of 10mm. The fiber is stretched at a tensile speed of 20mm / min to a strain of 100% (20mm, twice the initial length, i.e., a displacement of 10mm). The gripper is then moved at the same speed to the original length, causing the fiber to shrink. This operation is repeated, and the displacement length at the beginning of the second stretch (i.e., approximately equal to the end of the first shrinkage) is calculated as X. 100 mm, the tensile elongation recovery rate R 100 (%)teeth, R 100 =[20-(X 100 +10)] / 10×100 It is shown as follows. Residual strain rate S 100 (%)teeth S 100 =100-R 100 It is shown as follows.

[0051] Similarly, a 12 cm long fiber is subjected to a cycle test using a tensile tester under the conditions of a temperature of 23°C and an initial length of 100 mm. The fiber is stretched at a tensile speed of 100 mm / min to a strain of 50% (150 mm, which is 1.5 times the initial length, i.e., a displacement length of 50 mm). The gripper is then moved at the same speed to the original length, causing the fiber to shrink. When this operation is repeated, the displacement length at the beginning of the second stretch (i.e., approximately equal to the end of the first shrinkage) is defined as X. 50 mm, the tensile elongation recovery rate R 50 (%)teeth, R 50 = [150-(X 50 +100)] / 50×100 It is shown as follows. Residual strain rate S 50 (%)teeth S 50 =100-R 50 It is shown as follows.

[0052] In general terms, the chuck distance is Y, the initial strain is a%, and the displacement length at the beginning of the second extension (i.e., approximately equal to the end of the first contraction) is X. a mm, the tensile elongation recovery rate R a (%)teeth, R a = [Y+Y×a / 100-(X a +Y)] / (Y×a / 100)×100 It is shown as follows. Residual strain rate S a (%)teeth S a =100-R a It is shown as follows.

[0053] <Maximum pore diameter> In the fibrous medical material of the present invention, the diameter of the largest pore (hole, fine hole, pore, void, gap, or hollow space) measured by microscopic observation of a cross section perpendicular to the fiber axis is preferably 100 μm or less, more preferably 75 μm or less, and even more preferably 50 μm or less. There is no particular lower limit to the diameter of the largest pore (hole, fine hole, pore, void, gap, or hollow space), but it is generally 0.1 μm or more, and may be 0.2 μm or more, or 1 μm or more. There is also no limit to the number of pores per cross section, and there may be one or more, and each pore may be independent or connected.

[0054] The diameter of the largest pore of a fibrous medical material can be measured, for example, by the following method. Images that allowed identification of pores inside the fibers were taken using a scanning electron microscope, and the images were analyzed using ImageJ (an image processing program developed by the National Institutes of Health) to calculate the diameter of the pores.

[0055] <Ratio of major axis length to minor axis length> The cross-sectional shape of the fibrous medical material of the present invention does not necessarily have to be circular, and examples include ellipse, polygon, free curve, and combinations thereof. If the cross-sectional shape is circular, the thread diameter can be determined by measuring the thread width. However, if the cross-sectional shape is elliptical, the ratio of the major axis length to the minor axis length (major axis length / minor axis length) may be 1.0 or more, 1.1 or more, or 1.2 or more. There are no particular restrictions on the upper limit of the ratio of the major axis length to the minor axis length (major axis length / minor axis length), but it is generally 3.0 or less, and may be 2.0 or less.

[0056] The ratio of the minor axis length to the major axis length (major axis length / minor axis length) in the width direction cross section of the fibrous medical material can be measured by the following method. The minor axis length and major axis length were measured using a dial thickness gauge (Teclock Corporation, SM-1201L type, graduation 0.001 mm). Measurements were taken at three locations (1 / 4, 1 / 2, and 3 / 4 of the total length) of the yarn to be measured in the direction of the fiber where the thickness was thinnest, and the average of these measurements was taken as the minor axis length. The major axis length was measured by gently deforming the yarn into a U-shape without folding the fiber, and then measuring the major axis perpendicular to the gauge. The major axis length was measured at three locations (1 / 4, 1 / 2, and 3 / 4 of the total length) of the yarn to be measured, and the average of these measurements was taken as the major axis length.

[0057] The present invention will be explained in more detail with reference to the following examples, but the present invention is not limited to these examples. [Example]

[0058] In the following examples and comparative examples, the side or cross-section of suture knots and threads was observed under an electron microscope. Prior to observation, sutures were cut to an appropriate size with a razor, placed on a specimen stage, and coated with a thin film of osmium using an osmium plasma coater, NL-OPC80NS (Japan Laser & Electronics Co., Ltd.). Observation was performed using a field-emission scanning electron microscope, JSM-7610F (JEOL Ltd.), at an accelerating voltage of 5.0 kV. For the cross-sections of some sutures, the sutures were cut to an appropriate size with a razor, placed on a specimen stage, and platinum (Pt) was deposited using an ion sputtering machine, E1045 (Hitachi High-Technologies Corporation). The cross-sections were then observed using a thermal electron gun-type low-vacuum scanning electron microscope, TM4000plus (Hitachi High-Technologies Corporation), at an accelerating voltage of 5.0 kV.

[0059] Example 1 Using a plunger-type melt extrusion spinning machine IMC-19F8 (Imoto Machinery Co., Ltd.), approximately 5 g of PHA (3HB-co-14.8 mol% 4HB) with a weight-average molecular weight (Mw) of 970,000 was added to the cylinder. A 1 mm die diameter was used, and melt extrusion (extrusion rate: 1 mm / s) was performed at 170 °C for 5 minutes. The yarn was then wound around a 114 mm diameter bobbin at 5 rpm, ensuring no overlapping of the yarn. After winding, the yarn was allowed to solidify for 60 minutes at room temperature (23 °C) to partially promote crystallization. After that, the yarn was manually drawn by hot pin stretching at 60 °C at a draw ratio of approximately 5-10x to produce a slightly flattened and whitened stretchable monofilament yarn. The diameter of the stretchable monofilament yarn produced was measured using a dial thickness gauge (Teclock Corporation, SM-1201L model, 0.001 mm graduation) or a dial gauge (Ozaki Manufacturing Co., Ltd., 5B-HG model, 0.001 mm graduation). Measurements were taken at three locations (1 / 4, 1 / 2, and 3 / 4) of a 10 cm piece of yarn, and the average value was used as the yarn diameter. When the cross section of the P(3HB-co-4HB) yarn was not necessarily circular but elliptical or flattened, the minor and major axis lengths were measured, and the cross-sectional area was calculated as an ellipse and used for evaluation in the tensile test. The cross section of the yarn in Example 1 herein was obtained by winding an amorphous molten polymer and then hot-pin drawing, resulting in a shape that was not circular but was flattened. The yarn diameter in Table 1 is the average of the minor and major axis lengths. Hot-pin drawing is a method of drawing a fiber while pressing it against a heated metal pin.

[0060] The surface of the yarn was also observed using a scanning electron microscope (Figure 3).

[0061] Using this thread, a knot was tied in a soft elastomer artificial skin sheet using a surgeon's knot and tightened with a force gauge (Standard Model Digital Force Gauge: ZTS-100N, IMADA CO., LTD) at approximately 5 N of force. The knot was photographed with a camera (DP26, OLYMPUS CO., LTD) attached to a stereomicroscope (SZX7, OLYMPUS CO., LTD), and the size of the knot (perimeter and area enclosed by the perimeter) was analyzed using image analysis software (cellSens, OLYMPUS CO., LTD) (Figure 5). Because the knot is three-dimensional, it was difficult to accurately estimate its volume. Therefore, the area enclosed by the perimeter and the area enclosed by the perimeter divided by the thread diameter were used as indices of knot size. The area enclosed by the perimeter may vary depending on the viewing angle; however, five knots were randomly analyzed and treated as the average. The results are shown in Table 1 (Example 1).

[0062] The average diameter of the slightly flat suture made of P(3HB-co-14.8mol%4HB) in Example 1 was 0.205mm in the short direction and 0.352mm in the long direction, with the average diameter of the short and long directions being 0.281mm. The size of the surgeon's knot fastened with a force of 5N was 3.46mm in average circumference, and the average area of ​​the region enclosed by that circumference was 0.688mm. 2 The average value of [area surrounded by perimeter / fiber diameter], which is an index of the size of the knot, was 2.45. The weight-average molecular weight Mw of the PHA after spinning was 320,000.

[0063] The results of observing the appearance of the surgeon's knot using a scanning electron microscope are shown in Figure 4. Also, an optical microscope photograph of the surgeon's knot used to calculate the values ​​related to the size of the knot is shown in Figure 5.

[0064] [Table 1]

[0065] <Example 2> A plunger-type melt viscosity measuring device, Flow Tester CFT-500D (Shimadzu Corporation), was used as the melt spinning device, using P(3HB-co-15.3 mol% 4HB) with a weight-average molecular weight Mw of 700,000. The piston diameter was 11.282 mm (piston cross-sectional area 1 cm). 2 Approximately 1 g of PHA was placed in the cylinder, and a die (nozzle) with a hole diameter of 1 mm and a hole length of 1 mm was used. After 120 seconds of preheating at 150°C, partial melt spinning was performed. The weight used was 2.5 kg, and a total load of 3 kg was applied using the weight and fishing gear. Due to the increased force from the pulley, the piston cross-sectional area was 1 cm. 2 The polymer was extruded under an extrusion pressure of 2.942 MPa. Differential scanning calorimetry (DSC) analysis of this PHA revealed a melting peak between 85°C and approximately 155°C, indicating that the melting temperature of 150°C was a partial melt, not a complete melt. The remaining crystals from the partial melt spinning and the molten, fluid extruded polymer were already semi-solidified immediately after extrusion. The extruded polymer was manually stretched approximately 5 to 10 times to produce a nearly transparent, stretchable monofilament yarn. The yarn diameter was measured using a dial thickness gauge (Teclock Corporation, SM-1201L, 0.001 mm graduation) or a dial gauge (Ozaki Manufacturing Co., Ltd., 5B-HG, 0.001 mm graduation). Measurements were taken at three points (1 / 4, 1 / 2, and 3 / 4) of the yarn, cut into 10 cm lengths, and the average was used as the yarn diameter. The cross section of this yarn was nearly circular. Using this 3-0 size thread, a knot was tied using a surgeon's knot in the same manner as in Example 1, and the size of the knot was analyzed using a stereomicroscope (FIG. 6). The results are shown in Table 2 (Example 2).

[0066] The average diameter of the suture made of P(3HB-co-15.3mol%4HB) in Example 2 was 0.207mm (almost circular, major axis length / minor axis length <1.2), which corresponds to a thread diameter of 3-0 according to the USP (United States Pharmacopia) standard. The size of the surgical knot fastened with a force of approximately 5N was 3.61mm in average circumference, and the average area of ​​the region surrounded by that circumference was 0.744mm. 2The average value of [area surrounded by perimeter / fiber diameter], which is an index of the size of the knot, was 3.59, which was slightly larger than that of Example 1. The weight-average molecular weight Mw of the PHA after spinning was 470,000.

[0067] [Table 2]

[0068] Example 3 A plunger-type melt viscosity measuring device, Flow Tester CFT-500D (Shimadzu Corporation), was used as the melt spinning device, using P(3HB-co-15.3 mol% 4HB) with a weight-average molecular weight Mw of 750,000. The piston diameter was 11.282 mm (piston cross-sectional area 1 cm). 2 Approximately 1 g of PHA was placed into the cylinder, and melt-spun using a die (nozzle) with a hole diameter of 1 mm and a hole length of 1 mm at 170°C after 120 seconds of preheating. A 2.5 kg weight was used, and a total load of 3 kg was applied using the weight and fishing gear. The polymer was extruded at an extrusion pressure of 2.942 MPa, as in Example 2. This PHA exhibited a melting peak in differential scanning calorimetry (DSC) analysis from 60°C to approximately 170°C, suggesting that it was almost completely melted at the melting temperature of 170°C. The extruded fiber was not wound on a bobbin but was hung in a straight line and allowed to solidify for 30 minutes at room temperature (23°C), allowing partial crystallization to proceed. It was then manually stretched at a draw ratio of approximately 5x to produce a transparent, stretchable monofilament yarn, and the yarn diameter was measured in the same manner as in Example 2. Using this single-size thread, a knot was tied using a surgeon's knot in the same manner as in Example 1, and the size of the knot was analyzed using a stereomicroscope (FIG. 7). The results are shown in Table 3 (Example 3).

[0069] The average diameter of the suture made of P(3HB-co-15.3mol%4HB) in Example 3 was 0.406mm (almost circular, major axis length / minor axis length <1.2), which corresponds to a diameter of 1 according to the USP standard. The size of the surgeon's knot fastened with a force of approximately 5N was 4.87mm in average circumference, and the average area of ​​the region surrounded by that circumference was 1.35mm. 2The average value of [area surrounded by perimeter / fiber diameter], which is an index of the size of the knot, was 3.33, which was the same as in Example 2. The weight-average molecular weight Mw of the PHA after spinning was 450,000.

[0070] As shown in Figures 5, 6, and 7, the knots of the elastic yarns obtained from P(3HB-co-4HB) shown in Examples 1, 2, and 3 were tightly knotted with no gaps between the yarns. The knot size index (area enclosed by perimeter / yarn diameter) was slightly smaller in Example 1, but was at roughly the same level in Examples 2 and 3.

[0071] [Table 3]

[0072] Example 4 The yarns in Examples 1 to 3 were prepared by manual drawing using a small, laboratory-scale plunger-type melt extrusion device. Using an industrially used single-screw spinning and drawing device with a diameter of 16 mm and a die with a diameter of 1 mm, a P(3HB-co-16.0 mol% 4HB) copolymer with a weight-average molecular weight (Mw) of 560,000 was partially melt-spun at a rate of 0.9 g / min with the extruder temperature set to 145-160°C. The extruded fiber, a mixture of undissolved crystals and fluidized polymer, was passed through water at 50°C and then wound around a multi-stage roller in air at a room temperature of 23°C and drawn (draw ratio of approximately 9 times) to produce a stretchable yarn. The resulting yarn with a circular cross section was used to measure the yarn diameter in the same manner as in Example 2. Using this 2.5-0 standard thread, a knot was tied using a surgeon's knot in the same manner as in Example 1, and the size of the knot was analyzed using a stereomicroscope (FIG. 8). The results are shown in Table 4 (Example 4).

[0073] The average diameter of the suture thread used to evaluate the knot size of the P(3HB-co-16.0 mol% 4HB) suture thread of Example 4 was 0.256 mm (almost circular, major axis length / minor axis length <1.2), which corresponds to a thread diameter of 2.5-0 according to the USP standard. The size of the surgical knot fastened with a force of approximately 5 N was 4.03 mm in average circumference, and the average area of ​​the region surrounded by that circumference was 0.843 mm. 2 The average value of [area surrounded by the perimeter / fiber diameter], which is an index of the size of the knot, was 3.29, which was equivalent to that of Examples 2 and 3. The weight-average molecular weight Mw of the PHA after spinning was 350,000.

[0074] As shown in Figures 5, 6, 7, and 8, the knots of the elastic yarns obtained from P(3HB-co-4HB) shown in Examples 1, 2, 3, and 4 appear to be tightly knotted with no gaps between the yarns. The [area enclosed by the perimeter / yarn diameter], which was used as an index of the size of the knot in Example 1, was slightly smaller than the [area enclosed by the perimeter / yarn diameter] in the other Examples, but was at a similar level in Examples 2, 3, and 4.

[0075] [Table 4]

[0076] <Comparative Example 1> The same procedure as in Example 2 was repeated, except that B.BRAUN's MonoMax suture (size 2-0) made of P(4HB) was used. The thread surface was observed under a scanning electron microscope (Fig. 9), the surgeon's knot was observed under a scanning electron microscope (Fig. 10), and the circumference of the knot and the area enclosed by the circumference were analyzed (Fig. 11), and the ratio of the area enclosed by the circumference to the thread diameter was calculated. The results are shown in Table 5 (Comparative Example 1).

[0077] The average diameter of the MonoMax suture (size 2-0) made of P(4HB) in Comparative Example 1 was 0.346 mm, which was certainly equivalent to the USP standard of 2-0. The average circumference of the knot tightened with a force of 5 N was 6.91 mm, and the average area of ​​the region enclosed by that circumference was 2.60 mm. 2An optical microscope photograph of the surgical knot used in the calculation is shown in FIG. The average value of [area enclosed by perimeter / thread diameter], which was used as an index of the size of the knot, was 7.50, which was clearly larger than those of Examples 1 to 4.

[0078] [Table 5]

[0079] Comparative Example 2: PDSII suture (size 3-0) The same procedure as in Comparative Example 1 was performed using Ethicon PDSII suture (size 3-0) made of polydioxanone, except that scanning electron microscope observation of the suture surface and knot was omitted. The suture diameter, knot circumference, and area enclosed by the circumference were analyzed (Figure 12), and the area enclosed by the circumference / suture diameter ratio was calculated. The results are shown in Table 6.

[0080] The average thread diameter of the PDSII suture (size 3-0) made of polydioxanone in Comparative Example 2 was 0.291 mm. However, the medical device package insert states that PDS sutures are within the USP specifications except for diameter, and that the upper limit of the thread diameter specification is set higher than the USP, with 3-0 being up to 0.056 mm larger than the specification. Therefore, subtracting 0.056 mm from 0.291 mm gives 0.235 mm, which corresponds to USP size 3-0, but the actual thread diameter was on average 0.291 mm. The size of the surgeon's knot fastened with a force of 5 N had an average circumference of 6.48 mm, and the average planar area enclosed by that circumference was 2.28 mm. 2 An optical microscope photograph of the surgical knot used in the calculation is shown in FIG.

[0081] The average value of the [area enclosed by the perimeter / thread diameter], which was used as an index of knot size for the PDSII suture used in Comparative Example 2, was 7.82, which was clearly larger than that of the P(3HB-co-4HB) sutures of Examples 1 to 4, but was almost the same as that of the MonoMax suture, the P(4HB) suture of Comparative Example 1.

[0082] [Table 6]

[0083] Comparative Example 3: PDSII suture (size 4-0) The same procedure as in Example 2 was carried out using Ethicon PDSII suture (size 4-0) made of polydioxanone, except that scanning electron microscope observation of the suture surface and knot was omitted. The suture diameter, knot circumference, and area enclosed by the circumference were analyzed (Figure 13), and the area enclosed by the circumference / suture diameter ratio was calculated. The results are shown in Table 7.

[0084] The average thread diameter of the PDSII suture (size 4-0) made of polydioxanone in Comparative Example 3 was 0.163 mm, but the medical device package insert states that PDS sutures are within the USP specifications except for diameter, and that the upper limit of the thread diameter specification is set larger than the USP, with 4-0 being up to 0.029 mm larger than the specification. The measured diameter of the suture in Comparative Example 3 was 0.163 mm, which was within the USP 4-0 size. The size of the surgeon's knot tightened with a force of 5 N was 4.77 mm on average in circumference, and the average area of ​​the plane enclosed by that circumference was 1.20 mm. 2 An optical microscope photograph of the surgical knot used in the calculation is shown in FIG. The average value of [area enclosed by perimeter / thread diameter], which was used as an index of knot size for the PDSII suture used in Comparative Example 3, was 7.33, which was clearly larger than that of the P(3HB-co-4HB) sutures of Examples 1 to 4, but was approximately the same level as the MonoMax suture (size 2-0), which is the P(4HB) suture of Comparative Example 1, and the PDSII suture (size 3-0) of Comparative Example 2.

[0085] [Table 7]

[0086] Even when comparing the circumferences around the knots of surgical knots, the circumference of the knot of the P(3HB-co-4HB) suture of 1 size (average thread diameter 0.406 mm) in Example 3, which is 4.86 mm, is smaller than the circumference of the knot of the MonoMax suture of 2-0 size (average thread diameter 0.346 mm) in Comparative Example 1, which is 6.91 mm, and the circumference of the knot of the PDSII suture of 3-0 size (average thread diameter 0.291 mm) in Comparative Example 2, which is 6.48 mm. Furthermore, the fact that it is equivalent to the circumference of the knot of the PDSII suture of 4-0 size (average thread diameter 0.163 mm) in Comparative Example 3, which is 4.77 mm, indicates that the knot of the P(3HB-co-4HB) suture, which is an elastic suture, is smaller than those of other absorbent sutures. As described above, from the results of Examples 1 to 4 and Comparative Examples 1 to 3, it was suggested that the knot of the P(3HB-co-4HB) suture is smaller than those of the existing absorbent monofilament sutures, the MonoMax suture and the PDSII suture.

[0087] <Comparison of Knot security factor (knot stability)> The evaluation of the difficulty of untying the ligated part (knot stability) was evaluated in vitro using the Knot Security factor (KSF), which is also described by Odermatt et al. (International Journal of Polymer Science, Vol. 2012, Article ID 216137).

[0088] The suture was wound around a plastic tube with a diameter of 2.9 cm, tightly knotted with a surgical knot (surgical knot: Figure 1), and one thread was created by cutting the side opposite to the knot. Both sides of them were attached to a tensile testing machine and tensile testing was performed at a speed of 100 mm / min. Multiple sets were prepared with 1 set of 10 samples. If even one of the 10 samples was untied at the Knot part, a single knot was added above the surgical knot (Figure 2), and single knots were added until none of them were untied at the Knot part. The number of single knots added until all 10 samples were no longer untied at the Knot part was defined as the Knot Security factor (KSF).

[0089] Therefore, KSF can be exemplified as follows: KSF=0: Surgeon's knot only, the knot will not come undone under tension, and there are no additional single knots. KSF=1: A surgeon's knot with only one additional single knot will not come undone under tension. KSF=2 The surgeon's knot cannot be untied by tension with only two additional single knots. KSF=3 A surgeon's knot with only three additional single knots will not come undone under tension. KSF=4 A surgeon's knot with only four additional single knots will not come undone under tension. KSF=n A surgeon's knot cannot be untied under tension with only n additional single knots.

[0090] In the case of monofilament yarn, generally, when 2≦KSF≦4, it can be said that the knot stability is excellent, and when KSF≦1, it can be said that the knot stability is absolutely excellent.

[0091] The KSF of the P(3HB-co-14.8mol%4HB) elastic monofilament yarn, including the P(3HB-co-14.8mol%4HB) yarn shown in Example 1, was evaluated as KSF = 1, meaning that all elastic yarns with average diameters equivalent to 3-0, 2-0, and 0 could not be untied by tension with just one additional single knot in a surgeon's knot, and were judged to have absolutely excellent (even better) knot stability.

[0092] The KSF of the 3-0 size P(3HB-co-15.3mol%4HB) elastic monofilament yarn shown in Example 2 and the 3-0 size P(3HB-co-15.3mol%4HB) elastic monofilament yarn shown in Example 3 both had two additional single knots compared to a surgeon's knot and could not be untied under tension, and was evaluated as KSF = 2, which is considered to be excellent knot stability.

[0093] The KSF of the 2.5-0 size P(3HB-co-16.0 mol% 4HB) elastic monofilament yarn shown in Example 4 was evaluated as KSF = 2, which is excellent knot stability, as in Examples 2 and 3. The knot did not come undone under tension with only two additional single knots compared to the surgeon's knot.

[0094] The KSFs of the MonoMax sutures made of P(4HB) shown in Comparative Example 1, rated 3-0, 2-0, and 0, were evaluated as 2, 2, and 3, respectively (International Journal of Polymer Science, Vol. 2012, Article ID 216137).

[0095] The KSF of the PDSII sutures (3-0, 4-0) made of polydioxanone shown in Comparative Examples 2 and 3, as well as the 2-0 PDSII suture, was evaluated as KSF = 3, which means that the knot would not come undone under tension if three additional single knots were added to the surgeon's knot, and this can be judged to be excellent knot stability.

[0096] <Comparative Examples 4 and 5> Furthermore, when the data from Silver E et al. (J. Oral. Maxillofac. Surg. 2016 Jul;74(7):1304-1312.) are substituted for KSF, the KSF of both the 3-0 and 4-0 nylon sutures (Comparative Example 4) is evaluated as 3, and the KSF of the 3-0 and 4-0 multifilament Vicryl sutures (Comparative Example 5) is evaluated as 4 and 3, respectively. These KSF values ​​are shown in Table 8.

[0097] [Table 8]

[0098] As described above, the KSF of P(3HB-co-4HB) stretchable monofilament suture was superior to that of other sutures, including other bioabsorbable and nonabsorbable monofilament sutures and bioabsorbable multifilament (braided) sutures. With P(3HB-co-4HB) stretchable monofilament suture, fewer single knots can be added to the surgeon's knot to prevent postoperative unraveling. This reduces the number of knots required and the flexibility of the suture itself allows for smaller knots to be formed during surgery. This reduces the volume of the knots themselves, potentially contributing to a reduced foreign body sensation in the surrounding tissue. Furthermore, despite being a monofilament, P(3HB-co-4HB) stretchable suture has a low initial tensile modulus, and its intermediate tensile modulus is even lower than the initial tensile modulus. It is pliable, and when forming a knot, it easily and securely tightens without the need for excessive force, preventing immediate loosening, providing excellent operability. Furthermore, since the previously tied knot does not loosen, adding a single knot is easy. Other monofilament sutures (such as MonoMax, PDSII, and nylon, which are conventional monofilament sutures) are rigid and impliable, with a high intermediate tensile modulus exceeding 400 MPa. They are less flexible than P(3HB-co-4HB) sutures, and as a result, excessive force is often applied to tighten the knot to prevent loosening. This situation can lead to excessive force being applied to the tissue at the suture site when tightening the knot, potentially causing unintended tissue damage and raising concerns about its impact on the tissue. Surgery does not necessarily require ligation in areas with a wide, open space; there are also many situations requiring ligation in a limited range of motion or within a narrow surgical field. In such cases, a suture that allows for easy tying and secure knots with light force is required. Sutures that meet the knot tensile strength requirements of the current USP standard also have a high modulus of elasticity, and are not necessarily satisfactory in terms of ease of knot formation and ease of loosening of the knot.Therefore, there is a need in the medical field for a suture thread that has a lower elastic modulus (initial tensile elastic modulus or intermediate tensile elastic modulus) than existing suture threads, has a high followability to tissue deformation, and is easy to operate. Also, the fact that the knot is small and difficult to untie means that when cutting the excess thread from the knot, it is not necessary to leave a long cutting end of the thread. Moreover, if the thread is flexible, the cutting surface of the thread is less likely to stimulate the tissue due to its rigidity, and in this regard, it is expected to contribute to reducing the foreign body sensation given to the surrounding tissue.

[0099] The Young's modulus of each tissue of the living tissue is summarized by Funai et al. (Research Report of Shizuoka Prefectural Industrial Technology Center, 2007, No. 52, p. 33 - 37, "Creation of a database of physical property values of living tissues for biomechanical simulation and its application examples"). While the elastic modulus of teeth and cortical bone exceeds 10000 MPa, the elastic modulus of ligaments is 248 MPa, the elastic modulus of cartilage is 23 MPa, the elastic modulus of the cornea is 20 MPa, and the elastic modulus of other soft tissues such as various internal organs, muscles, and skin is 10 MPa or less. Even the lowest elastic modulus of existing absorbable suture threads is 485 MPa of MonoMax, and there is no existing absorbable suture thread having an elastic modulus close to that of soft tissues.

[0100] <Ethylene Oxide Gas (EOG) Sterilization of the Stretchable Suture Thread> The P(3HB - co - 4HB) suture thread of Example 1 was sterilized with EO gas. First, the P(3HB - co - 4HB) suture thread cut to an appropriate length was packaged with a sterilization packaging material having ethylene oxide gas permeability (Hybrid Plating Bag HM - 1304: manufactured by Hogimedical Co., Ltd.), and the opening was heated and sealed using a heat sealer. The P(3HB - co - 4HB) stretchable suture thread wrapped in the sterilization packaging material was sterilized with ethylene oxide gas at 40°C for 5 hours using 15 g of 95% ethylene oxide (Exitech 95, manufactured by Nippon Yuki Carbon Co., Ltd.) in a fully automatic ethylene oxide gas sterilizer (Eogelk, SA - N160, manufactured by Erk). After 2 hours of air replacement, further aeration treatment was performed for 14 hours. No obvious macroscopic or microscopic structural or physical property changes were observed in the P(3HB-co-4HB) suture thread.

[0101] <In vitro degradation of P(3HB-co-4HB) stretchable suture thread> To evaluate the degradability of the P(3HB-co-4HB) stretchable suture thread, its degradation behavior in buffer solution was evaluated in vitro. The method is shown below. The 3-0 size PHA suture thread of Example 1 sterilized by EOG was immersed in Dulbecco's phosphate buffer (pH 7.4; 37 °C), taken out at 1, 2, 3, 4, 6, 8, 12, and 16 weeks, gently washed with water, then vacuum dried, and subjected to tensile test and molecular weight measurement. Those not immersed in the acid buffer were taken as 0 week (initial).

[0102] For the tensile test of the P(3HB-co-4HB) stretchable suture thread, a stretched PHA fiber with a length of 3 cm and a fiber diameter of about 0.1 - 0.3 mm was used with a tensile testing machine AGS-50NX (manufactured by Shimadzu Corporation) under the conditions of a temperature of 23 °C, a test speed of 10 mm / min, and an initial length (distance between chucks) of 10 mm to conduct a tensile test until the fiber broke.

[0103] The results of the tensile test are shown in Figure 14, and taking the weight average molecular weight Mw of 320,000 of the sample before immersion as a relative value of 100%, the degree of molecular weight decrease during the immersion period is shown in Figure 15. As shown in Figure 14, the initial breaking elongation of the P(3HB-co-4HB) stretchable suture thread exceeded 180%.

[0104] When the PHA stretchable suture thread was immersed in the buffer solution, the period during which it could maintain half of the initial linear tensile breaking strength was approximately 16 weeks (Figure 14), and the weight average molecular weight also decreased by half at approximately 16 weeks (Figure 15). Regarding the breaking elongation, it also maintained an average of 150% after 16 weeks (Figure 14), indicating that the characteristic of high stretchability of this thread was maintained.

[0105] According to Williams et al. (Biomed Tech (Berl). 2013 Oct;58(5):439-452, Poly-4-hydroxybutyrate (P4HB): a new generation of resorbable medical devices for tissue repair and regeneration), the breaking elongation of oriented P(4HB) fibers is approximately 25 to 90% or less, and it has been reported that the breaking elongation of Monomax sutures with a diameter of 0.154 mm made of oriented P(4HB) is 35%. Also, according to Albertsmeier et al. (Langenbecks Arch Surg (2012) 397:363-371, Evaluation of the safety and efficacy of MonoMax R suture material for abdominal wall closure after primary midline laparotomy-a controlled prospective multicentre trial: ISSAAC [NCT005725079]), the breaking elongation of Monomax sutures is 90%, and for PDS sutures and MonoPlus sutures made of polydioxanone, the breaking elongation is described as 45 to 50%. Compared with these sutures, the breaking elongation of P(3HB-co-4HB) stretchable sutures is clearly about 2 to 5 times greater, and it is expected to have the property of being able to follow the deformation of tissues.

[0106] <In vivo degradation of P(3HB-co-4HB) stretchable sutures> To evaluate the degradability of P(3HB-co-4HB) stretchable sutures, their in vivo degradation behavior was evaluated in rats. The method is shown below.

[0107] An 8 cm incision was made in the dorsal skin of a rat (F344 / NSlc male, 20 weeks old) along the spinal column, and the PHA suture of Example 1 sterilized with EOG was implanted into the subcutaneous tissue. Samples were collected after 4, 8, 12, 16, and 26 weeks, lightly washed with water, vacuum dried, and subjected to tensile testing and molecular weight measurement. Samples without implantation were designated as week 0 (initial period).

[0108] Tensile tests of P(3HB-co-4HB) elastic sutures were conducted using a tensile testing machine AGS-50NX (Shimadzu Corporation) on stretched PHA fibers with a length of 3 cm and a fiber diameter of approximately 0.1 to 0.3 mm, at a temperature of 23°C, a test speed of 10 mm / min, and an initial length (distance between chucks) of 10 mm, until the fibers broke.

[0109] The results of the tensile test are shown in FIG. 16, and the degree of molecular weight reduction during the implantation period is shown in FIG. 17, with the weight average molecular weight of the sample before implantation set as a relative value of 100%.

[0110] When PHA elastic sutures were implanted subcutaneously in rats, half of the initial straight tensile breaking strength was maintained for approximately 16 to 24 weeks (Figure 16), and the decrease in weight-average molecular weight was also halved at approximately 16 weeks (Figure 17). The breaking elongation was also maintained at 150% even after 26 weeks (Figure 16).

[0111] It has been shown that the in vivo tensile breaking strength of 3-0 MonoMax sutures made of P(4HB) can maintain half of its initial breaking strength for 12 weeks, while the in vivo tensile breaking strength of 3-0 PDSII sutures made of polydioxanone was reduced to half in 6 weeks (International Journal of Polymer Science, Vol. 2012, Article ID 216137). Based on these findings, it is believed that P(3HB-co-4HB) elastic suture can be used in areas where tensile strength must be maintained for a longer period of time than MonoMax suture or PDSII suture.

[0112] <Microminipig abdominal wall suture test> Suturing tests were conducted using the P(3HB-co-4HB) elastic suture of Example 1 and other absorbable sutures in 32-week-old female microminipigs (Fuji Micra Co., Ltd.) to evaluate macroscopically and microscopically the presence or absence of complications and the extent of inflammatory reactions. After the fetus was removed by cesarean section, the abdominal wall was closed with three cranial P(3HB-co-4HB) sutures, two central poly(glycolide-co-trimethylene carbonate) (PGA-TMC copolymer) sutures (Maxon sutures), and three caudal P(4HB) sutures (MonoMax sutures). Seven weeks after suturing, the sutured areas were visually examined for any complications (signs of infection, wound dehiscence, abdominal wall incisional hernia, adhesions), but no significant complications were observed at the sutured areas using any of the three types of thread (Figure 18).

[0113] Next, the abdominal wall including the suture site was collected, and the suture site of each thread was cut into pieces. After each was fixed in paraffin, it was stained with hematoxylin and eosin (HE) using standard methods, and the degree of inflammation was observed under an optical microscope. The tissue observation was scored by a single pathologist in a blinded manner, without revealing which sample belonged to which suture, on a 4-point scale of 0, 1, 2, or 3 for inflammation, necrosis, and fibrous thickening. 0: No inflammatory reaction, no necrosis, no hyperplasia 1: Very little inflammatory reaction, very little necrosis, very little hyperplasia 2: Inflammatory reaction, necrosis, and hyperplasia 3: Strong inflammatory reaction, strong necrosis, strong hyperplasia

[0114] The results were reevaluated two weeks after the initial observation, and if the evaluations differed between the first and second observations, a third evaluation was conducted two weeks later to determine the final evaluation. The results are shown in Table 9. Photographs of the results of HE staining of the tissues sutured with P(3HB-co-4HB) suture, poly(glycolide-co-trimethylene carbonate) (PGA-TMC) suture, and P(4HB) suture are shown in Figures 19, 20, and 21, respectively. As a result of histological observation, it was suggested that P(3HB-co-4HB) threads had less inflammation than the other two types of sutures. Also, they might be non-inferior with regard to necrosis and fibrous thickening.

[0115]

Table 9

[0116] The characteristics of having less inflammation and potentially being non-inferior with regard to necrosis and fibrous thickening contribute to indicating the usefulness of P(3HB-co-4HB) sutures. Considering also the characteristic of the suture knots becoming smaller, the P(3HB-co-4HB) absorbable elastic suture, which is a long-term absorbable and elastic thread and the knot itself also becomes smaller, is a medical device with new application possibilities and charm, as it can be applied to sites where it could not be used in previous medicine because the degradation absorption was too fast, and also in cases where the tissue side is damaged in soft tissues because the tension of the thread is too strong or it is difficult to stretch compared to the tissue.

[0117] <Tensile fracture evaluation of P(3HB-co-4HB) elastic suture> The same P(3HB-co-4HB) elastic suture used in Example 1 was re-evaluated by a tensile test that performed tensile loading until fracture. A P(3HB-co-4HB) suture with a length of 3 cm, an average fiber diameter of about 0.24 mm for the short axis, and an average fiber diameter of about 0.40 mm for the long axis (long axis length / short axis length = 1.7) was subjected to a tensile test until fiber fracture under the conditions of a temperature of 23°C, a test speed of 10 mm / min, and an initial length (distance between chucks) of 10 mm using a tensile testing machine AGS-50NX (manufactured by Shimadzu Corporation). An example of the stress-strain curve results is shown in Fig. 22. Similarly, examples of the stress-strain curve results of the tensile tests until fracture for Examples 2 to 4 are shown in Figs. 23 to 25.

[0118] The tensile fracture strength of the fiber used in Example 1, also shown in Fig. 22 as an example, was 161 MPa on average for 5 points, and the fracture elongation was 240% (the variation was 180 - 282%). Also, the tensile fracture strength of the fiber used in Example 2 shown in Fig. 23 was 120 MPa on average for 5 samples, and the fracture elongation was 183% (the variation was 157 - 209%). The tensile fracture strength of the fiber used in Example 3 shown in Fig. 24 was 69 MPa on average for 5 samples, and the fracture elongation was 250% (the variation was 178 - 338%). The tensile fracture strength of the fiber used in Example 4 shown in Fig. 25 was 110 MPa on average for 5 samples, and the fracture elongation was 232% (the variation was 192 - 272%).

[0119] <Evaluation of the Stretch Recovery and Residual Strain of P(3HB - co - 4HB) Elastic Suture Threads> The P(3HB - co - 4HB) elastic suture thread used in Example 1 was evaluated by a cycle test of repeatedly stretching and contracting it. A P(3HB - co - 4HB) suture thread with a length of 3 cm and a fiber diameter of about 0.2 mm in the long - axis thickness was subjected to a cycle test using a tensile testing machine AGS - 50NX (manufactured by Shimadzu Corporation) under the conditions of a temperature of 23°C and an initial length of 10 mm. At a tensile speed of 20 mm / min, it was stretched to a strain of 100% (twice the length), and then the gripper was moved back to the original length at the same speed to contract the PHA fiber. This was repeated 5 times. The stress - strain curves during the contraction from the 1st to the 5th times are shown in Fig. 26.

[0120] For the P(3HB - co - 4HB) elastic suture thread used in this Example 1, when a strain of 100% was applied during elongation, at the first point of the second elongation (which is approximately equal to the end point of the first contraction), the tensile elongation recovery rate (%) was about 60%, and the residual strain rate was about 40%. At the first points of the third to fifth elongations, the tensile elongation recovery rate (%) was about 60% - about 55%, and the residual strain rate was about 40% - about 45% (Fig. 26).

[0121] Here, the tensile elongation recovery rate R 100(%) is a tensile strength test performed on a 3cm long fiber using a tensile testing machine at a temperature of 23°C and an initial length of 10mm. The fiber is stretched at a tensile speed of 20mm / min to a strain of 100% (20mm, twice the initial length, i.e., a displacement of 10mm), and then the gripper is moved at the same speed to the original length, causing the fiber to shrink. When this operation is repeated, the displacement length at the beginning of the second stretch (i.e., considered to be approximately equal to the end of the first shrinkage) is X. 100 mm, the tensile elongation recovery rate R 100 (%)teeth, R 100 =[(20-(X 100 +10)) / 10]×100 It is shown as follows. Residual strain rate S 100 (%)teeth S 100 =100-R 100 It is shown as follows.

[0122] The P(3HB-co-4HB) stretchable suture used in Example 2 was evaluated by a cyclic test in which it was repeatedly stretched and contracted. A P(3HB-co-4HB) suture measuring 3 cm in length and 0.207 mm in fiber diameter was subjected to a cyclic test using a tensile testing machine, AGS-50NX (Shimadzu Corporation), at a temperature of 23°C and an initial length of 10 mm. The suture was stretched to a strain of 100% (double its original length) at a tensile speed of 20 mm / min, and then the clamps were moved at the same speed to the original length, causing the PHA fiber to contract. This cycle was repeated five times. The stress-strain curves for the first through fifth contractions are shown in Figure 27.

[0123] When the P(3HB-co-4HB) elastic suture used in Example 2 was stretched to 100% strain, the tensile elongation recovery (%) was approximately 67% and the residual strain was approximately 33% at the beginning of the second stretch (i.e., considered to be approximately equal to the end of the first contraction).At the beginning of the third to fifth stretches, the tensile elongation recovery (%) was approximately 63% to approximately 60%, and the residual strain was approximately 37% to approximately 40% (Figure 27).

[0124] The P(3HB-co-4HB) stretchable suture used in Example 3 was evaluated in a cyclic test in which it was repeatedly stretched and contracted. A 3 cm long P(3HB-co-4HB) suture with a fiber diameter of 0.410 mm was subjected to a cyclic test using a tensile testing machine AGS-50NX (Shimadzu Corporation) at a temperature of 23°C and an initial length of 10 mm. The suture was stretched to a strain of 100% (double its original length) at a tensile speed of 20 mm / min, and then the clamps were moved at the same speed to the original length, causing the PHA fiber to contract. This cycle was repeated five times. The stress-strain curves for the first to fifth contractions are shown in Figure 28.

[0125] When the P(3HB-co-4HB) elastic suture used in Example 3 was stretched to 100% strain, at the beginning of the second stretch (i.e., considered to be approximately equal to the end of the first contraction), the tensile elongation recovery (%) was approximately 70% and the residual strain was approximately 30%. At the beginning of the third to fifth stretches, the tensile elongation recovery (%) was approximately 63% to approximately 68%, and the residual strain was approximately 32% to approximately 37% (Figure 28).

[0126] The P(3HB-co-4HB) stretchable suture used in Example 4 was evaluated in a cyclic test in which it was repeatedly stretched and contracted. A P(3HB-co-4HB) suture measuring 3 cm in length and approximately 277 mm in fiber diameter was subjected to a cyclic test using a tensile testing machine, AGS-50NX (Shimadzu Corporation), at a temperature of 23°C and an initial length of 10 mm. The suture was stretched to 100% strain (double its original length) at a tensile speed of 20 mm / min, and then the clamps were moved at the same speed to return to the original length, causing the PHA fiber to contract. This cycle was repeated five times. The stress-strain curves for the first through fifth contractions are shown in Figure 29. When the P(3HB-co-4HB) elastic suture used in Example 4 was stretched to 100% strain, at the beginning of the second stretch (i.e., considered to be approximately equal to the end of the first contraction), the tensile elongation recovery (%) was approximately 74% and the residual strain was approximately 26%. At the beginning of the third to fifth stretches, the tensile elongation recovery (%) was approximately 72% to approximately 66%, and the residual strain was approximately 28% to approximately 34% (Figure 29).

[0127] Furthermore, the P(3HB-co-4HB) suture used in Example 4, 12 cm long and approximately 0.283 mm in fiber diameter, was subjected to a cyclic test using a tensile testing machine AGS-50NX (Shimadzu Corporation) at 23°C and an initial length of 100 mm. The suture was stretched to a strain of 50% (1.5 times its original length) at a tensile speed of 100 mm / min, and then the clamps were moved at the same speed to the original length, causing the PHA fiber to shrink. This was repeated five times. The stress-strain curves for the first through fifth shrinkages are shown in Figure 30.

[0128] At the beginning of the second extension (i.e., considered to be approximately equal to the end of the first contraction) after a 50% strain load, the P(3HB-co-4HB) elastic suture used in Example 4 had a tensile elongation recovery rate (%) of approximately 94% and a residual strain rate of approximately 6%. At the beginning of the third to fifth extensions, the tensile elongation recovery rate (%) was approximately 93% to approximately 90%, and the residual strain rate was approximately 7% to approximately 10% (Figure 30). As the proportion of initial strain decreases, the residual strain also decreases, indicating that the material is more susceptible to elastic recovery.

[0129] Here, a 12 cm long fiber is subjected to a cycle test using a tensile testing machine under the conditions of a temperature of 23°C and an initial length of 100 mm, and is stretched at a tensile speed of 100 mm / min to a strain of 50% (150 mm, which is 1.5 times the initial length, i.e., a displacement length of 50 mm). When this operation is repeated, the gripper is then moved at the same speed to the original length, causing the fiber to shrink. The displacement length at the start of the second stretch (i.e., approximately equal to the end of the first shrinkage) is defined as X.50 If it is set to mm, the tensile elongation recovery rate R 50 (%) is R 50 = [(150 - (X + 100)) / 50] × 100 as shown by. The residual strain rate S 50 (%) is S 50 = 100 - R 50 as shown by.

[0130] <Cross - sectional Analysis of P(3HB - co - 4HB) Elastic Suture> To explore the mechanism by which the knots of P(3HB - co - 4HB) elastic sutures become smaller, the cross - section of the P(3HB - co - 4HB) suture of Example 1 was observed with a scanning electron microscope. The results are shown in Figs. 31 and 32. As shown in Fig. 3, the surface of the fiber is smooth and there are no pores. However, as shown in Figs. 31 and 32, there were pores in the fibers of the P(3HB - co - 4HB) elastic suture used in Example 1. Through image analysis, it was found that there were 43.5% voids in the cross - sectional view of Fig. 31 and 24.9% voids in the cross - sectional view of Fig. 32. The cross - sections of 10 threads were measured, and the average porosity was 40 ± 15%.

[0131] Also, the scanning electron microscope observations of the P(3HB - co - 4HB) sutures used in Example 2 and Example 3 are shown in Figs. 33 and 34. Different from Figs. 31 and 32, no pores were observed in the cross - sections of Figs. 33 and 34, and it was observed that they were densely packed. This result supports the fact that the P(3HB - co - 4HB) sutures of Example 2 and Example 3 were colorless and transparent, while the P(3HB - co - 4HB) suture of Example 1 was turbid.

[0132] Regarding changes in polymer structure during stretching, as described in Patent Application No. 2019-90739, unstretched P(3HB-co-4HB) copolymer molded products retain a randomly oriented α structure (α crystals), and the α crystals are not uniformly oriented. However, when the molded product is stretched after a certain period of crystallization treatment, the degree of α crystal orientation increases in the stretched direction, while the amorphous molecular chains between the α crystals are stretched, resulting in the appearance of a β structure (planar zigzag structure). Upon unloading, the β structure decreases or disappears while maintaining the α crystal orientation, suggesting an elastic response. Here, the α structure is a folded lamellar structure, and the β structure is a planar zigzag extended chain structure.

[0133] The P(3HB-co-4HB) copolymer is not only a soft material with a low modulus of elasticity, but also can be spun and drawn into a stretchable structure. Furthermore, the presence of voids within the fibers is also thought to be one of the factors that contribute to the smaller knots.

[0134] There are various methods for creating gaps, voids, pores, and porous structures within non-bioabsorbable polymers, including phase separation, extraction, electron beam irradiation and etching, polymer particle fusion, foaming agent mixing, gas mixing, and stretching. The microcrystalline nucleus stretching method is known to form voids in PHA fibers, a bioabsorbable polymer, and attempts have also been made to impregnate PHA fibers with drug solutions. However, this is the first time that the presence of such voids has been found to contribute to the small size of knots in suture fibers, making it a noteworthy feature. There are no particular limitations on the method for introducing voids into fibers, as long as a strong fiber is obtained.

[0135] <Study on elastic modulus> A tensile tester was used to measure the initial and intermediate tensile moduli of the P(3HB-co-4HB) elastic suture threads of Examples 1 to 4. Herein, the initial tensile modulus is defined as the modulus calculated from the slope of the stress-strain curve corresponding to the two points between strains of 0.05% and 0.25%, and the intermediate tensile modulus is defined as the modulus calculated from the slope of the stress-strain curve corresponding to the two points between strains of 0.25% and 10%.

[0136] The distance between the chucks of the tensile tester was 1 cm, and the sutures were fixed using fixtures 1 cm above and below. The tensile speed was 10 mm / min. The initial tensile modulus of the suture of Example 1 was 520-645 MPa, averaging 589 MPa for five samples. The intermediate tensile modulus was 175-296 MPa, averaging 245 MPa for five samples. The initial tensile modulus of the suture of Example 2 was 328-599 MPa, averaging 492 MPa for five samples. The intermediate tensile modulus was 105-166 MPa, averaging 144 MPa for five samples. The initial tensile modulus of the suture of Example 3 was 222-467 MPa, averaging 373 MPa. The intermediate tensile modulus was 99-134 MPa, averaging 116 MPa for five samples. The initial tensile modulus of the suture of Example 4 was 354-484 MPa, averaging 391 MPa for five samples, and the intermediate tensile modulus was 139-184 MPa, averaging 167 MPa for five samples. The modulus of the MonoMax suture of Comparative Example 1 has been reported to be 485 MPa (literature value, International Journal of Polymer Science, Vol. 2012, Article ID 216137). The measured initial tensile modulus was 576-626 MPa, averaging 600 MPa for three samples, and the intermediate tensile modulus was 457-578 MPa, averaging 531 MPa for three samples. The modulus of elasticity of the PDSII of Comparative Examples 2 and 3 has been reported to be 1370 MPa (literature value, International Journal of Polymer Science, Vol. 2012, Article ID 216137), and the values ​​actually measured for the yarn of Comparative Example 2 were: initial tensile modulus of elasticity 1480 to 1660 MPa, averaging 1560 MPa at three sample points, intermediate tensile modulus of elasticity 1140 to 1210 MPa, averaging 1180 MPa at three sample points. Regarding the modulus of elasticity actually measured for the yarn of Comparative Example 3, the initial tensile modulus of elasticity was 1680 to 1710 MPa, averaging 1710 MPa at three sample points, and intermediate tensile modulus of elasticity 1050 to 1080 MPa, averaging 1070 MPa at three sample points.The measured initial tensile modulus of the nylon suture of Comparative Example 4 (Nesco Suture nylon suture 4-0 was used for modulus evaluation) was 1250-1450 MPa, averaging 1350 MPa for three samples, and the intermediate tensile modulus was 1020-1090 MPa, averaging 1040 MPa for three samples. Furthermore, Vicryl, described in Comparative Example 5, is a braided yarn. If we assume that the cross section of the 3-0 yarn is a circular monofilament, the initial tensile modulus was calculated to be 10,000 MPa (measured value) for three samples, and the intermediate tensile modulus was 4460 MPa for three samples. The moduli are summarized in Table 10.

[0137] As shown above, when comparing the initial tensile moduli of the P(3HB-co-4HB) elastic suture threads of Examples 1 to 4, the initial tensile moduli are 589 MPa in Example 1, 492 MPa in Example 2, 373 MPa in Example 3, and 391 MPa in Example 4, while the PDSII of Comparative Examples 2 and 3 have an initial tensile modulus of 1370 MPa to 1710 MPa, the nylon of Comparative Example 4 has an initial tensile modulus of 1350 MPa, and the Vicryl of Comparative Example 5, although a braided thread, is 10,000 MPa if it is assumed to be a monofilament. Comparing these, it can be said that the initial tensile moduli of the P(3HB-co-4HB) of Examples 1 to 4 are significantly lower than that of Comparative Examples 2 to 5, but the initial tensile moduli of the MonoMax suture are approximately 485 MPa to 600 MPa, meaning that the initial tensile moduli of the MonoMax suture of Examples 1 and 2 are roughly equivalent to those of Comparative Example 1. However, the intermediate tensile moduli of Example 1 are 245 MPa, and that of Example 2 is 144 MPa, while the intermediate tensile moduli of the MonoMax suture are 531 MPa, meaning that the intermediate tensile moduli of P(3HB-co-4HB) are considerably lower than that of the MonoMax suture. It can be seen that as strain increases to 0.25% to 10%, the stretchable properties of the P(3HB-co-4HB) elastic suture become more pronounced. At the same time, P(3HB-co-4HB) elastic suture has the property of returning to its original shape even when stretched, so when a knot is formed, it is stretched moderately, and the knot is formed from the thin part of the stretched thread, while the part of the thread other than the knot shrinks and returns to its original thickness, which is thought to enable the formation of a knot that is small and difficult to unravel. The intermediate tensile moduli of the sutures of Examples 3 and 4 were 116 MPa and 167 Pa, respectively, which are also low values ​​similar to those of Examples 1 and 2.

[0138] Using the raw polymer used in Example 4, yarns with an initial tensile modulus of elasticity of approximately 180 MPa to 500 MPa could be obtained by changing the spinning conditions (screw temperature, spinneret temperature, discharge rate, crystallization temperature, crystallization time, draw ratio) and heat treatment temperature (annealing step). It is expected that further changes in the molecular weight and composition of the polymer used may enable spinning that covers an even wider range of elastic modulus. An example of the fiber shown in Example 4, which was industrially spun, is shown in Figures 35 and 36. Also, the state of a surgeon's knot tied is shown in Figure 37. There were no voids inside the fiber, but the yarn had an initial tensile modulus of 391 MPa and an intermediate tensile modulus of 167 Pa. The fiber was soft and tolerant of stretching, and had the property of shrinking. It was observed that there were no gaps between the threads in the knot, and the knot was tight.

[0139] As described above, the stretchable bioabsorbable fibrous medical material of the present invention is easy to tie knots in, allows for smaller knots, and can reduce the number of knots, thereby reducing the burden on doctors during surgery and reducing the physical irritation to patients' tissues, thereby making a useful contribution to medical care.

[0140] <Summary of Examples and Comparative Examples> A summary of the above examples and comparative examples is provided in Table 10 below.

[0141] [Table 10] JPEG0007785293000011.jpg137170

Claims

1. A fibrous medical material made of spun and drawn monofilaments of a bioabsorbable aliphatic polymer, the fibrous medical material having an elongation at break of 75% or more, an intermediate tensile modulus between 0.25% and 10% strain that is lower than the initial tensile modulus between 0.05% and 0.25% strain, the intermediate tensile modulus being 400 MPa or less, and a residual strain rate after 100% deformation being 70% or less, the bioabsorbable aliphatic polymer comprising 3-hydroxybutyrate units and 4-hydroxybutyrate units, and the ratio of the 4-hydroxybutyrate units to all monomer units being 10 mol% or more and 30 mol% or less.

2. The fibrous medical material according to claim 1 , wherein the initial tensile modulus is 1000 MPa or less.

3. 3. The fibrous medical material according to claim 1, wherein the initial tensile modulus is 480 MPa or less.

4. The fibrous medical material according to claim 1 , wherein the intermediate tensile modulus is 300 MPa or less.

5. 5. The fibrous medical material according to claim 1, wherein the residual strain rate after 100% deformation is 50% or less.

6. 6. The fibrous medical material according to claim 1, which has a porosity of 0% to 55%.

7. A fibrous medical material according to any one of claims 1 to 6, wherein the diameter of the largest pore (pore, void, gap, or hollow) measured by microscopic observation of a cross section perpendicular to the fiber axis direction is 100 μm or less.

8. 8. The fibrous medical material according to claim 1, wherein the ratio of the major axis length to the minor axis length (major axis length / minor axis length) in the width direction cross section is 1.0 or more and 3.0 or less.

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