STENT, METHOD FOR MANUFACTURING STENT, AND METHOD FOR RELIEVING STRAIN IN THE RESPIRATORY TRACT TO ENSURE AIR FLOW

Abstract

A respiratory stent comprising a bioabsorbable polyester copolymer, the bioabsorbable polyester copolymer being a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, and when the two types of ester bond-forming monomers are monomer A and monomer B, respectively, the R value represented by the following formula is 0.25 to 0.99: R=[AB] / (2[A][B])×100, in which [A] is the molar fraction (%) of monomer A residue in the polyester copolymer, [B] is the molar fraction (%) of monomer B residue in the polyester copolymer, and [AB] is the molar fraction (%) of a structure (AB and BA) in which a monomer A residue and a monomer B residue are adjacent to each other in the polyester copolymer.

Description

[Technical field]

[0001] The present disclosure relates to a respiratory stent comprising a bioabsorbable polyester copolymer, a method for making the stent, and a method for using the stent to relieve respiratory tract obstructions and ensure airflow. [Background technology]

[0002] Stents are implanted medical devices that can be placed inside the body, some of which are radially expandable, and are positioned inside various body cavities or vessels (e.g., the vascular system, esophagus, gastrointestinal tract, large and small intestines, bile duct, pancreatic duct, lung duct, ureter, nasal cavity, trachea, etc.) When a body cavity or vessel becomes narrowed, a stent is placed in the narrowed area to secure the lumen.

[0003] Such stents include those that are left in the body cavity or blood vessel for a long period of time, and those that are left in the body cavity or blood vessel for only a specified period of time and are removed from the body after maintaining the patency of the lumen. For example, Non-Patent Document 1 discloses an airway stent that is left in the narrowed site to ensure breathing when the airway or bronchi are narrowed due to lung cancer or the like. In the case of stents that are placed in a body cavity or vessel for a predetermined period of time, there is a need to treat the stent using a biodegradable polymer material.

[0004] As polymeric materials for such bioabsorbable materials, attention has been focused on polylactic acid, polyglycolic acid, polycaprolactone, polydioxane, and bioabsorbable polyesters, which are copolymers of these. For example, Patent Documents 1 and 2 disclose bioabsorbable stents made of polylactic acid or polycaprolactone, but many problems remain to be overcome in the development of bioabsorbable stents. [Prior art documents] [Patent documents]

[0005] [Patent Document 1] International Publication No. 2020-122096 [Patent Document 2] Japanese Patent No. 6505438 [Non-patent literature]

[0006] [Non-Patent Document 1] Yueqi Zhu et al., Materials Today 2017, 20, 516-529 Summary of the Invention [Problem to be solved by the invention]

[0007] Stents are required to function properly in an environment where multiple physical actions such as bending, stretching, and compression occur due to movements in a body cavity or blood vessel, so stents are required to have high compliance with the internal movements of biological organs. In addition, in respiratory stents such as airway stents, the occurrence of complications due to mucus adhesion and low compliance with the movements of biological organs is a major problem. However, the bioabsorbable stents described in Non-Patent Document 1 and Patent Document 2 are hard and therefore poorly compliant with the movements of biological organs, and may move or fall off from the placement site or damage surrounding tissues. Furthermore, the inner diameter of the trocar or catheter for transporting the stent to the placement site is smaller than the airway, and it is difficult to transport a hard stent through the trocar, and it may sometimes be damaged. In addition, even if a hard stent can be successfully deformed, it is impossible to restore it to a shape suitable for the size of the airway.

[0008] In addition, Patent Document 1 discloses a medical molded article that contains a bioabsorbable polyester and has excellent followability to the movement of biological organs, but does not consider mucus adhesion, which may cause complications such as stent stenosis and infection.

[0009] Therefore, an object of the present disclosure is to provide a highly biocompatible respiratory stent that contains a bioabsorbable polyester copolymer and thereby can suppress mucus adhesion and has excellent follow-up properties to the movements of biological organs. [Means for solving the problem]

[0010] In order to solve the above problems, the present disclosure provides the following. [1] 1. A respiratory stent comprising a bioabsorbable polyester copolymer, the bioabsorbable polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, A stent having an R value represented by the following formula, in which the two types of ester bond-forming monomers are monomer A and monomer B, respectively, is 0.25 or more and 0.99 or less. R = [AB] / (2[A][B]) × 100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residue and monomer B residue are adjacent to each other in the polyester copolymer [2] The stent described in [1] has a Young's modulus of 0.1 MPa or more and 50 MPa or less, as measured in accordance with JIS K6251 (2017). [3] The stent according to [1] or [2], having a restoring ability defined by the following formula of 40% or more. Restorability (%) = (L0 x 2 - L1) / L0 x 100 L0: initial length L1: The length after applying a tensile stress in the longest direction of the stent to generate a tensile strain of 100% of the initial length L0, repeated 10 times [4] The stent according to any one of [1] to [3], having a mucus adhesion amount defined by the following formula of 60% or less. Mucus adhesion amount (%) = (As-Asb) x 100 / (Ac-Acb) As: absorbance of sample at 450 nm Asb: absorbance at 450 nm of blank solution of sample (PBS instead of mucin solution, incubated overnight) Ac: absorbance of polylactic acid stent at 450 nm Acb: Absorbance at 450 nm of blank solution of polylactic acid stent (incubated overnight with PBS instead of mucin solution) [5] The stent according to any one of [1] to [4], further comprising a water-soluble polymer. [6] The stent according to [5], wherein the content of the water-soluble polymer defined by the following formula is 0.1% by mass or more and 25% by mass or less. Water-soluble polymer content (mass%) = [M1 / (M1+M2)×100] M1: mass of water-soluble polymer M2: Mass of polyester copolymer [7] The stent according to [5] or [6], wherein the water-soluble polymer is a polyalkylene glycol. [8] The stent according to any one of [1] to [7], wherein the bioabsorbable polyester copolymer is contained in an amount of 50% by mass or more based on 100% by mass of the stent. [9] The monomer A is at least one selected from the group consisting of lactic acid and glycolic acid, The stent according to any one of [1] to [8], wherein the monomer B is at least one selected from the group consisting of caprolactone and δ-valerolactone.

[10] The stent described in any one of [1] to [9] above, having an outer diameter of 4 mm or more and 24 mm or less, and a thickness of 0.2 mm or more and 2 mm or less.

[11] The stent described in any one of [1] to

[10] has a plurality of protrusions or a plurality of irregularities on its outer surface.

[12] The stent described in

[11] , wherein the height of the protrusions or irregularities is 0.1 mm or more and 3.0 mm or less.

[13] A method for producing the stent described in any one of [1] to

[12] , comprising a step of 3D printing using a printing material containing the bioabsorbable polyester copolymer.

[14] A method for relieving respiratory tract narrowing and ensuring airflow using a respiratory stent, comprising: the stent comprises a bioabsorbable polyester copolymer; the bioabsorbable polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, The method, wherein when the two types of ester bond-forming monomers are monomer A and monomer B, respectively, the R value represented by the following formula is 0.25 or more and 0.99 or less. R = [AB] / (2[A][B]) × 100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residue and monomer B residue are adjacent to each other in the polyester copolymer

[0011] INDUSTRIAL APPLICABILITY The present disclosure can provide a respiratory stent that can suppress mucus adhesion, has excellent followability to the movement of biological organs, and is highly biocompatible. [Brief description of the drawings]

[0012] [Figure 1] FIG. 1 is a schematic diagram of a stent according to an embodiment. [Diagram 2] FIG. 2 is a schematic diagram of a stent according to another specific example of an embodiment. [Diagram 3] FIG. 3 is a schematic diagram of a stent according to yet another specific example of an embodiment. [Figure 4]FIG. 4 is a schematic diagram of a stent according to yet another specific example of an embodiment. [Diagram 5] FIG. 5 is a cross-sectional view taken along line AA of the stent shown in FIG. [Figure 6] FIG. 6 is a schematic diagram of a stent according to yet another specific example of an embodiment. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0013] Hereinafter, embodiments will be described. Note that the present disclosure is not limited to the embodiments described below.

[0014] <Stent> The stent according to this embodiment is a respiratory stent comprising a bioabsorbable polyester copolymer.

[0015] In the present disclosure, the "respiratory system" is a general term for organs related to breathing, and examples thereof include the airway, oral cavity, nasal cavity, larynx, trachea, bronchi, bronchioles, and lungs. The stent according to this embodiment is a respiratory stent, and is preferably a stent for the airways, trachea, bronchi or lungs. By placing the stent according to this embodiment in the narrowed respiratory tract, the narrowing of the respiratory tract can be relieved and airflow can be ensured. The stent according to this embodiment can be applied not only to a narrowed respiratory tract but also to an obstructed respiratory tract.

[0016] [Bioabsorbable polyester copolymer] The stent according to this embodiment comprises a bioabsorbable polyester copolymer. The stent according to the present embodiment contains a specific bioabsorbable polyester copolymer, and therefore has the advantage that it is degraded inside and outside the body after a certain period of time, and the resulting product is metabolized or excreted, making it unnecessary to take out the placed stent from the body. In addition, when a part of the stent is degraded inside the body after a certain period of time, the remaining part of the stent becomes compatible with the body tissue, making it less likely to cause complications.

[0017] The stent of this embodiment is not limited in the amount of bioabsorbable polyester copolymer contained therein, as long as it contains a bioabsorbable polyester copolymer; however, it is preferable for the stent to contain 50% or more by mass of bioabsorbable polyester copolymer relative to the entire stent (100% by mass), and more preferably 80% or more by mass. In cases where it is required that the stent completely disappears when applied to a living body, it is particularly preferable that the stent is made only of a bioabsorbable polyester copolymer, that is, that the stent contains 100% by mass of the bioabsorbable polyester copolymer relative to the entire stent (100% by mass). Furthermore, in order to obtain a stent that has excellent followability to the movement of a living organ as required in the present disclosure, it is preferable that the stent contains a bioabsorbable polyester copolymer as described below within the above range.

[0018] Here, "bioabsorbability" of the stent according to the embodiment of the present disclosure refers to the property of the stent being naturally decomposed (i.e., biodegraded) by hydrolysis or enzymatic reaction after being placed inside or outside the body, and the resulting product being metabolized or excreted and disappears.

[0019] In the stent according to this embodiment, the polyester copolymer is bioabsorbable. The polyester copolymer is a copolymer constituted of two or more types of monomers including at least one type of ester bond-forming monomer, and may be a copolymer having an ester bond-forming monomer residue as a main constituent unit. The polyester copolymer according to this embodiment preferably includes a polyester copolymer having two or more types of ester bond-forming monomer residues as main constituent units, and may be a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units.

[0020] In the stent according to this embodiment, a more preferred embodiment of the bioabsorbable polyester copolymer is one in which the bioabsorbable polyester copolymer comprises a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units (hereinafter, such a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units contained in such a bioabsorbable polyester copolymer may be simply referred to as the "polyester copolymer of the present disclosure").

[0021] The content of the polyester copolymer in the bioabsorbable polyester copolymer (100% by mass) in the stent is preferably 50% by mass or more, more preferably 70% by mass or more, even more preferably 95% by mass or more, and particularly preferably 100% by mass. By setting the content at such a level, even if the stent is deformed by multiple physical actions such as bending, stretching, and compression due to movement in a body cavity or a blood vessel, it can return to its original shape and can suppress mucus adhesion.

[0022] The term "ester bond forming monomer" refers to a monomer which, after polymerization, forms a polymer in which monomer units are linked by ester bonds, i.e., a monomer which, after polymerization, gives rise to a polyester.

[0023] The ester bond-forming monomer is preferably at least one selected from the group consisting of hydroxycarboxylic acids, cyclic esters of hydroxycarboxylic acids, and dimeric cyclic esters of hydroxycarboxylic acids, and it is more preferable to use hydroxycarboxylic acids, etc. That is, it is preferable that the polyester copolymer includes a polyester copolymer having two or more types of ester bond-forming monomer residues as main constituent units, and the ester bond-forming monomer is at least one selected from the group consisting of hydroxycarboxylic acids, cyclic esters of hydroxycarboxylic acids, and dimeric cyclic esters of hydroxycarboxylic acids.

[0024] As the hydroxycarboxylic acid, it is particularly preferable to use an aliphatic hydroxycarboxylic acid, such as lactic acid, glycolic acid, hydroxybutyric acid, hydroxyvaleric acid, hydroxypentanoic acid, hydroxycaproic acid, and hydroxyheptanoic acid, with lactic acid, glycolic acid, hydroxypentanoic acid, and hydroxycaproic acid being particularly preferred.

[0025] As the lactic acid, L-lactic acid, D-lactic acid, and a mixture thereof can be used, but in terms of the physical properties and biocompatibility of the resulting polymer, it is preferable to use L-lactic acid. When a mixture of L-lactic acid and D-lactic acid is used as a monomer, the content of L-isomer is preferably 85% or more, and more preferably 95% or more.

[0026] The cyclic ester of hydroxycarboxylic acid is preferably a lactone, which is a cyclic compound formed by intramolecular dehydration condensation of a hydroxy group and a carboxyl group of a hydroxycarboxylic acid, and the dimeric cyclic ester of hydroxycarboxylic acid is preferably lactide, which is formed by dehydration condensation of the hydroxy group and the carboxyl group of two molecules of hydroxycarboxylic acid.

[0027] Examples of lactones that can be used include caprolactone, dioxepanone, ethylene oxalate, dioxanone, 1,4-dioxane-2,3-dione, trimethylene carbonate, δ-valerolactone, β-propiolactone, β-butyrolactone, γ-butyrolactone, and pivalolactone. Caprolactone and δ-valerolactone are particularly preferred.

[0028] As the lactide, dilactide obtained by dehydration condensation of two molecules of lactic acid, glycolide obtained by dehydration condensation of two molecules of glycolic acid, and tetramethyl glycolide can be used.

[0029] As the ester bond-forming monomer, derivatives of the monomers exemplified above can also be used.

[0030] Among these, in the present embodiment, it is more preferable that the ester bond-forming monomer is at least one selected from the group consisting of lactic acid, glycolic acid, hydroxybutyric acid, hydroxyvaleric acid, hydroxypentanoic acid, hydroxycaproic acid, hydroxyheptanoic acid, caprolactone, dioxepanone, ethylene oxalate, dioxanone, 1,4-dioxane-2,3-dione, trimethylene carbonate, δ-valerolactone, β-propiolactone, butyrolactone (β-butyrolactone, γ-butyrolactone), pivalolactone, dilactide, glycolide, and tetramethylglycolide.

[0031] It is preferable that Monomer A (described later) is at least one selected from the group consisting of lactic acid, glycolic acid, dilactide, and glycolide, and Monomer B (described later) is at least one selected from the group consisting of hydroxyvaleric acid, hydroxycaproic acid, δ-valerolactone, and caprolactone. Moreover, monomer A is more preferably lactic acid or glycolic acid, and monomer B is more preferably caprolactone or δ-valerolactone.

[0032] In this specification, when the bioabsorbable polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, the monomer that forms a homopolymer consisting only of the two types of ester bond-forming monomers and has high crystallinity is referred to as monomer A, and the other monomer that forms a homopolymer consisting only of the monomer residue and has low crystallinity is referred to as monomer B. The crystallinity of the homopolymer can be measured using a differential scanning calorimeter (DSC) as follows.

[0033] The homopolymer is sampled on an aluminum pan and measured by the DSC method under the following condition A using a differential scanning calorimeter (EXTAR 600; manufactured by Seiko Instruments Inc.) to calculate the heat of fusion. The higher the heat of fusion per unit mass, the higher the crystallinity. For example, the heat of fusion per unit mass of polylactic acid is 93 J / g when determined by the above method.

[0034] (Condition A) Device name: EXTAR 600 (Seiko Instruments Inc.) Temperature conditions: 25℃→250℃ (heating rate: 10℃ / min) Standard material: α-alumina

[0035] In the present disclosure, it is preferable that the crystallinity of both the monomer A residue and the monomer B residue is less than 14%. If the crystallinity is less than 14%, an increase in Young's modulus is suppressed, and a polyester copolymer suitable for a stent can be obtained. The crystallinity of the monomer A residue and the monomer B residue is preferably 10% or less, and more preferably 5% or less. Here, the crystallization rate of a monomer residue is the ratio of the heat of fusion per unit mass of the monomer residue in the polyester copolymer of the present disclosure to the product of the heat of fusion per unit mass of a homopolymer consisting of only a certain monomer residue and the mass fraction of the monomer residue in the polyester copolymer of the present disclosure. That is, the crystallization rate of the monomer A residue is the ratio of the heat of fusion per unit mass of the monomer A residue in the polyester copolymer of the present disclosure to the product of the heat of fusion per unit mass of a homopolymer consisting of only monomer A and the mass fraction of the monomer A residue in the polyester copolymer of the present disclosure. The crystallization rates of the monomer A residue and the monomer B residue indicate the proportion of the monomer A residue or monomer B residue in the polyester copolymer of the present disclosure that form a crystal structure, respectively.

[0036] In particular, when the monomer A residue is a lactic acid residue and the monomer B residue is a caprolactone residue, the crystallization rate of the lactic acid residue and the caprolactone residue is preferably less than 14%, more preferably 10% or less. The crystallization rate is specifically determined by the following method.

[0037] The polyester copolymer is dissolved in chloroform to a concentration of 5% by mass, and the solution is transferred onto a Teflon (registered trademark) petri dish and dried for 20 to 24 hours at normal pressure and room temperature (20°C to 25°C). The resultant is dried under reduced pressure to obtain a polyester copolymer film. The obtained polyester copolymer film is collected on an alumina PAN and measured under the following conditions by the DSC method using a differential scanning calorimeter, and the heat of fusion is calculated from the measurement results under temperature conditions (D) and (E). The crystallization ratio is calculated from the following formula.

[0038] Crystallization rate of lactic acid residue (%)=[(heat of fusion per unit mass of lactic acid residue of polyester copolymer) / (heat of fusion per unit mass of homopolymer consisting of only lactic acid residue)×(mass fraction of lactic acid residue in polyester copolymer)]×100

[0039] Crystallization rate of caprolactone residue (%)=[(heat of fusion per unit mass of caprolactone residue in polyester copolymer) / (heat of fusion per unit mass of homopolymer consisting of only caprolactone residue)×(mass fraction of caprolactone residue in polyester copolymer)]×100

[0040] Device name: EXTAR 600 (Seiko Instruments Inc.) Temperature conditions: (A) 25℃ → (B) 250℃ (heating rate: 10℃ / min) → (C) 250℃ (holding time: 5min) → (D) -70℃ (cooling rate: 10℃ / min) → (E) 250℃ (heating rate: 10℃ / min) → (F) 250℃ (holding time: 5min) → (G) 25℃ (cooling rate: 100℃ / min) Standard material: Alumina

[0041] In this specification, the term "monomer residue" refers, in principle, to a repeating unit of a chemical structure derived from a monomer in the chemical structure of a copolymer obtained by polymerizing two or more monomers including the monomer. For example, when lactic acid (CH3CH(OH)COOH) and caprolactone (ε-caprolactone: formula below) are polymerized to obtain a copolymer of lactic acid and caprolactone,

[0042] [ka]

[0043] The lactic acid monomer residue is represented by the following formula:

[0044] [ka]

[0045] Units expressed by the following formula

[0046] [ka]

[0047] is a caprolactone monomer residue.

[0048] However, as an exception, when a dimer such as lactide is used as a monomer, the term "monomer residue" refers to one of the two repeating structures derived from the dimer. For example, dilactide (L-(-)-lactide: formula below)

[0049] [ka]

[0050] When and caprolactone are polymerized, the chemical structure of the resulting copolymer has a structure in which the structure shown in the above formula (R1) is repeated twice as a dilactide residue. In this case, one of the lactic acid units is regarded as a "monomer residue," and two "monomer residues," i.e., two lactic acid residues, are considered to have been formed derived from the dilactide.

[0051] The bioabsorbable polyester copolymer of the present embodiment is a polyester copolymer having two types of ester bond-forming monomer residues as main structural units. Here, the term "main structural units" means that, when the sum of all monomer residues contained in the entire polymer, including other monomer residues, is 100%, the sum of the two types of monomer residues is 50 mol% or more, and the number of each of the two monomer residues is 20 mol% or more. For example, the term "main structural units" means that, when the sum of all monomer residues contained in the entire polymer, including other monomer residues, is 100%, the sum of the residues of monomer A residue and monomer B residue is 50 mol% or more, the number of monomer A residue is 20 mol% or more, and the number of monomer B residue is 20 mol% or more. Here, the molar fractions of the monomer A residue, the monomer B residue, and other residues can be determined from the area values ​​of the signals derived from each residue by nuclear magnetic resonance (NMR) measurement. For example, when the monomer A residue is a lactic acid residue and the monomer B residue is a caprolactone residue, their molar fractions can be measured by the method described in Measurement Example 1 below.

[0052] The sum of the number of monomer A residues and the number of monomer B residues is 50 mol% or more, preferably 75 mol% or more, and more preferably 90 mol% or more, when the sum of all monomer residues contained in the entire polymer including other monomer residues is 100 mol%, as defined above. Also, the number of monomer A residues and the number of monomer B residues are each 20 mol% or more, preferably 30 mol% or more, and more preferably 40 mol% or more, as defined above. A polymer in which the sum of the number of monomer A residues and the number of monomer B residues is 100 mol% (entire polymer), i.e., a polymer consisting only of monomer A and monomer B, is a particularly preferred embodiment.

[0053] In addition, as long as the effect of the present disclosure is not impaired, another monomer that can be copolymerized with the two types of ester bond-forming monomers that constitute the main structural unit can also be copolymerized. As such a monomer, another one of the above-mentioned ester bond-forming monomers can be used.

[0054] In addition, it is also a preferred embodiment to copolymerize a monomer that functions as a linker. Examples of the monomer that functions as a linker include hydroxycarboxylic acids other than the two types of ester bond-forming monomers that constitute the main structural units, dialcohols, dicarboxylic acids, amino acids, diamines, diisocyanates, diepoxides, and the like.

[0055] In this specification, the term "polyester copolymer" also refers to copolymers that contain monomers other than ester bond-forming monomers as constituent units and thus contain constituent units linked by bonds other than ester bonds.

[0056] The polyester copolymer according to the present embodiment must be bioabsorbable. A person skilled in the art would be able to synthesize a copolymer that exhibits appropriate bioabsorbability depending on the application by appropriately combining the above-exemplified monomers and adjusting the ratio of the monomers within the range specified in the present disclosure.

[0057] In the polyester copolymer of the present disclosure, when either monomer A or monomer B is present in excess, the polyester copolymer approaches homopolymer-like properties, and therefore the molar ratio of monomer A residues to the total of monomer A residues and monomer B residues (total mole number 100%) is preferably 20% to 80%, more preferably 30% to 70%, and even more preferably 40% to 60%.

[0058] In the stent according to this embodiment, the polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, and when, of the two types of ester bond-forming monomers, a monomer having high crystallinity in a homopolymer composed only of that monomer residue is designated as monomer A, and the other monomer having low crystallinity in a homopolymer composed only of that monomer residue is designated as monomer B, it is preferable that the R value represented by the following formula is 0.25 or more and 0.99 or less.

[0059] R = [AB] / (2[A][B]) × 100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residue and monomer B residue are adjacent to each other in the polyester copolymer

[0060] The R value is used as an index showing the randomness of the arrangement of monomer residues in a copolymer whose main building blocks are two types of ester bond-forming monomer residues, i.e., monomer A residue and monomer B residue. For example, the R value is 1 for a random copolymer in which the monomer arrangement is completely random. The R value can be determined by quantifying the ratio of the number of AA, BB, AB, and BA among combinations of two adjacent monomer residues (hereinafter sometimes referred to as dyads) by nuclear magnetic resonance (NMR) measurement, and specifically, it is measured by the method described in Measurement Example 1 below. For example, when a polyester copolymer is composed of only monomer A and monomer B, [AB] refers to the ratio of the total number of AB dyads and the number of BA dyads to the total number of all dyads (AA, BB, AB, BA) in the polyester copolymer. In addition, when three or more types of ester bond-forming monomers are present, the two ester bond-forming monomers with the highest content are selected, and of the two, the monomer that forms a homopolymer composed only of that monomer residue is highly crystalline is designated as monomer A, and the other monomer that forms a homopolymer composed only of that monomer residue is poorly crystalline is designated as monomer B. For example, if a polyester copolymer is composed of three types of monomers, monomer X, monomer Y, and monomer Z, the top two monomers with the highest content are selected, and among the two monomers, the monomer with the highest homopolymer crystallinity is designated as monomer A, the other monomer with the lowest homopolymer crystallinity is designated as monomer B, and the monomer with the lowest content is designated as monomer C. In this case, [AB] refers to the ratio of the total number of AB dyads and the total number of BA dyads to the total number of all dyads (AA, BB, AB, BA, AC, CA, BC, CB, CC) in the polyester copolymer. The same applies when the polyester copolymer is composed of four or more types of monomers. If the R value is less than 0.25, the polyester copolymer has high crystallinity, and the resulting stent may become hard, increasing the Young's modulus and decreasing the restorability. On the other hand, if the R value exceeds 0.99, the resulting stent may become too soft and sticky, decreasing the ease of handling and the restorability. Therefore, in order to exhibit high conformity to the movement of biological organs and high biocompatibility by suppressing mucus adhesion, the R value is preferably 0.25 or more and 0.99 or less, more preferably 0.45 or more and 0.99 or less, more preferably 0.50 or more and 0.85 or less, and even more preferably 0.50 or more and 0.80 or less.

[0061] In order to control the crystallization rate within a suitable range, the weight average molecular weight of the polyester copolymer according to this embodiment is preferably within the range of 80,000 to 1,000,000, more preferably within the range of 100,000 to 1,000,000, more preferably within the range of 120,000 to 750,000, more preferably within the range of 150,000 to 750,000, even more preferably within the range of 150,000 to 500,000, and particularly preferably within the range of 200,000 to 500,000. The weight average molecular weight of the polyester copolymer can be measured, for example, by the method described in Measurement Example 2.

[0062] [Method for producing polyester copolymer] The polyester copolymer according to the present embodiment is, for example, a macromer synthesis step of blending and polymerizing two types of ester bond-forming monomers, namely, monomer A and monomer B, so that, at the completion of polymerization, the sum of the number of monomer A residues and the number of monomer B residues is 50 mol % or more of the total number of residues, the number of monomer A residues is 20 mol % or more of the total number of residues, and the number of monomer B residues is 20 mol % or more of the total number of residues; a multiplication step of linking the macromers obtained in the macromer synthesis step to each other, or of multiplying the macromers by additionally adding the monomer A and the monomer B to the macromer solution obtained in the macromer synthesis step; The polyester copolymer having the formula:

[0063] [Macromer synthesis process] In the macromer synthesis step, monomer A and monomer B are mixed and polymerized so that, theoretically, at the completion of polymerization, the sum of the number of monomer A residues and the number of monomer B residues is 50 mol % or more of the total number of residues, the number of monomer A residues is 20 mol % or more of the total number of residues, and the number of monomer B residues is 20 mol % or more of the total number of residues. This results in a polyester copolymer having monomer A residues and monomer B residues as main building blocks. Since this production method further includes a multi-production step (described below), in this specification, the polyester copolymer obtained by this step is referred to as a "macromer."

[0064] As the ester bond-forming monomer, the same ones as those described above can be used, and the preferred combinations of the ester bond-forming monomers are also as described above.

[0065] The randomness of the distribution of monomer residues constituting a polyester copolymer having two types of ester bond-forming monomer residues as main building blocks varies depending on the difference in the reactivity of the monomers during polymerization. That is, if one of the two types of monomers is bound to the other monomer with equal probability during polymerization, a copolymer in which the monomer residues are distributed completely randomly is obtained. However, if there is a tendency for one monomer to be bound to the other monomer, a gradient copolymer in which the distribution of monomer residues is biased is obtained. The composition of the monomer residues in the obtained gradient copolymer changes continuously from the polymerization initiation end to the polymerization termination end along the molecular chain.

[0066] Here, if the initial polymerization rate of monomer A is higher than that of monomer B, when monomer A and monomer B are copolymerized in the macromer synthesis step, monomer A is likely to bond after monomer A. Therefore, in the synthesized macromer, a gradient structure is formed that forms a composition gradient in which the proportion of monomer A units gradually decreases from the polymerization initiation end to the polymerization termination end. That is, the macromer obtained in this step has a gradient structure in which monomer A residues and monomer B residues are arranged to form a composition gradient in the skeleton. That is, by using monomer A and monomer B with different initial polymerization rates in this step, a macromer having a gradient structure that forms a composition gradient in the skeleton can be obtained. Such a macromer may be referred to as a "gradient macromer" in this specification.

[0067] In order to realize such a gradient structure in the macromer synthesis step, it is desirable to synthesize the macromer by a polymerization reaction that occurs in one direction from the initiation terminal. Preferred examples of such a synthesis reaction include reactions that utilize ring-opening polymerization and living polymerization.

[0068] In order to facilitate the final production of a polyester copolymer that satisfies the aforementioned R value range, it is preferable that the macromer obtained in this process has an R' value similar to the R value desired for the polyester copolymer, that is, an R' value represented by the following formula of 0.25 or more and 0.99 or less. R' = [AB'] / (2 [A'] [B']) × 100 [A']: Molar fraction (%) of monomer A residues in the macromer [B']: Molar fraction (%) of monomer B residues in the macromer [AB']: Molar fraction (%) of structures in which monomer A residue and monomer B residue are adjacent to each other (A'-B' and B'-A') in the macromer

[0069] The weight average molecular weight of the macromer synthesized in the macromer synthesis step is preferably at least 10,000, more preferably at least 20,000. In order to suppress crystallinity and maintain the flexibility of the macromer, the weight average molecular weight of the macromer synthesized in the macromer synthesis step is preferably at most 150,000, more preferably at most 100,000.

[0070] [Multi-layering process] In the multi-formation step, the macromers obtained in the macromer synthesis step are linked together, or monomer A and monomer B are added to the macromer solution obtained in the macromer synthesis step to form a multi-formation. In this step, macromers obtained in one macromer synthesis step may be linked together, or multiple macromers obtained in two or more macromer synthesis steps may be linked. Note that "multi-formation" refers to the formation of a structure in which multiple molecular chains having a gradient structure in which monomer A residues and monomer B residues are arranged so as to have a composition gradient in the skeleton are repeated. The number of multimeric macromer units may be 2 or more, but since a larger number of connections results in improved tensile strength due to entanglement of molecular chains, the number of multimeric macromer units is preferably 3 or more, more preferably 4 or more, and even more preferably 6 or more. On the other hand, if the molecular weight of the polyester copolymer produced increases excessively, there is a concern that the viscosity will increase and this will have a negative effect on moldability, so the number of macromer units is preferably 80 or less, more preferably 40 or less, and even more preferably 20 or less. The number of linked macromer units can be adjusted by the catalyst and reaction time used in the multi-polymerization step. When linking macromers together to form a multi-polymer, the number of macromer units can be determined by dividing the weight-average molecular weight of the final polyester copolymer by the weight-average molecular weight of the macromer.

[0071] As a catalyst for the multi-production step, for example, 4-dimethylaminopyridinium p-toluenesulfonate, 4-dimethylaminopyridine, etc. can be used.

[0072] The polyester copolymer of the present disclosure may be a linear polymer in which macromer units are linked together in a linear fashion, or a branched polymer in which macromer units are linked in a branched fashion.

[0073] A linear polyester copolymer can be synthesized, for example, by linking a gradient macromer to a similar gradient macromer, one molecule at a time, end-to-end.

[0074] When the gradient macromer has a hydroxyl group and a carboxyl group at each end, the ends are condensed with a condensing agent to obtain a multi-polyester copolymer. Examples of the condensing agent include 4-dimethylaminopyridinium p-toluenesulfonate, 1-[3-(dimethylamino)propyl]-3-ethylcarbodiimide, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride, N,N'-dicyclohexylcarbodiimide, N,N'-diisopropylcarbodiimide, N,N'-carbonyldiimidazole, 1,1'-carbonyldi(1,2,4-triazole), 4-(4,6-dimethoxy-1,3,5-triazin-2-yl)-4-methylmorpholinium. ammonium chloride hydrate, (4,6-dimethoxy-1,3,5-triazin-2-yl)-(2-octoxy-2-oxoethyl)dimethylammonium trifluoromethanesulfonate, 1H-benzotriazol-1-yloxytris(dimethylamino)phosphonium hexafluorophosphate, 1H-benzotriazol-1-yloxytripyrrolidinophosphonium hexafluorophosphate, (7-azabenzotriazol-1-yloxy)tripyrrolidinophosphonium hexafluorophosphate, chlorine Lotripyrrolidinophosphonium hexafluorophosphate, Bromotris(dimethylamino)phosphonium hexafluorophosphate, 3-(diethoxyphosphoryloxy)-1,2,3-benzotriazin-4(3H)-one, O-(benzotriazol-1-yl)-N,N,N',N'-tetramethyluronium hexafluorophosphate, O-(7-azabenzotriazol-1-yl)-N,N,N',N'-tetramethyluronium hexafluorophosphate, O-(N-succinimidyl)-N,N, N',N'-Tetramethyluronium tetrafluoroborate, O-(N-Succinimidyl)-N,N,N',N'-Tetramethyluronium hexafluorophosphate, O-(3,4-Dihydro-4-oxo-1,2,3-benzotriazin-3-yl)-N,N,N',N'-Tetramethyluronium tetrafluoroborate, S-(1-Oxido-2-pyridyl)-N,N,N',N'-Tetramethylthiuronium tetrafluoroborate, O-[2-Oxo-1(2H)-pyridyl]-N,N,N',N'-tetramethyluronium tetrafluoroborate, {{[(1-cyano-2-ethoxy-2-oxoethylidene)amino]oxy}-4-morpholinomethylene}dimethylammonium hexafluorophosphate, 2-chloro-1,3-dimethylimidazolinium hexafluorophosphate, 1-(chloro-1-pyrrolidinylmethylene)pyrrolidinium hexafluorophosphate, 2-fluoro-1,3-dimethylimidazolinium hexafluorophosphate, fluoro-N,N,N',N'-tetramethylformamidinium hexafluorophosphate, and the like can be used.

[0075] Furthermore, when the polymerization reaction has a living nature, i.e., when the polymerization reaction can be initiated continuously from the end of the polymer, multimerization can be achieved by repeatedly adding monomer A and monomer B to the gradient macromer solution after the polymerization reaction is completed.

[0076] Alternatively, gradient macromers may be multi-linked via a linker as long as the mechanical properties of the polymer are not affected. In particular, by using a linker having multiple carboxyl groups or multiple hydroxyl groups, such as 2,2-bis(hydroxymethyl)propionic acid, a branched polyester copolymer can be synthesized in which the linker is the branch point.

[0077] The polyester copolymer obtained by the above-mentioned production method is a copolymer having a structure in which two or more macromer units in which monomer A residues and monomer B residues are arranged so as to have a composition gradient in the skeleton are linked together, which is a preferred embodiment of the polyester copolymer of the present disclosure. In this specification, for the sake of convenience, such a structure may be referred to as a "multi-gradient" structure, and a copolymer having a multi-gradient structure may be referred to as a "multi-gradient copolymer".

[0078] In other words, the polyester copolymer of the present disclosure is preferably a multi-gradient copolymer, and the multi-gradient copolymer preferably has a structure in which two or more macromer units having a gradient structure in which monomer A residues and the monomer B residues form a composition gradient in the skeleton are linked, and more preferably has a structure in which three or more such macromer units are linked. Furthermore, the upper limit for the number of linked macromer units having a gradient structure in which monomer A residues and the monomer B residues form a composition gradient in the skeleton is preferably 80 or less, more preferably 40 or less, and even more preferably 20 or less.

[0079] As mentioned above, polyester copolymers in which the monomer A residue is lactic acid and the monomer B residue is caprolactone residue are particularly preferred in the present disclosure. Such polyester copolymers are preferably produced by the following production method.

[0080] First, in the macromer synthesis process, dilactide and ε-caprolactone are polymerized in the presence of a catalyst. The dilactide and ε-caprolactone monomers are preferably purified to remove impurities before use. Dilactide can be purified, for example, by recrystallization from toluene dried with sodium. ε-caprolactone can be purified, for example, by vacuum distillation from CaH2 under a N2 atmosphere.

[0081] As a catalyst for the synthesis step of a macromer having a lactic acid residue and a caprolactone residue, a polyester polymerization catalyst such as a normal germanium-based, titanium-based, antimony-based, or tin-based catalyst can be used. Specific examples of such polyester polymerization catalysts include tin octylate, antimony trifluoride, zinc powder, dibutyltin oxide, and tin oxalate. The method of adding the catalyst to the reaction system is not particularly limited, but is preferably a method in which the catalyst is added in a state dispersed in the raw materials when the raw materials are charged, or a method in which the catalyst is added in a state dispersed at the start of pressure reduction. The amount of the catalyst used is 0.01 to 3 mass% (metal atom equivalent) based on the total amount of the monomers used, more preferably 0.05 to 1.5 mass%.

[0082] A macromer having lactic acid residues and caprolactone residues can be obtained by placing dilactide, caprolactone, and a catalyst in a reaction vessel equipped with a stirrer and reacting the dilactide and caprolactone at 120°C to 250°C under a nitrogen stream.

[0083] The reaction may be carried out using an initiator such as hydroxypivalic acid, alcohol, etc. When water is used as a co-initiator, it is preferable to carry out a co-catalyst reaction at about 90° C. prior to the polymerization reaction.

[0084] The reaction time is 2 hours or more, preferably 4 hours or more, and in order to increase the degree of polymerization, the reaction time is preferably longer, for example, 8 hours or more. However, if the reaction is carried out for too long, the polymer may become discolored, so the reaction time is preferably 3 to 30 hours.

[0085] Next, in the multi-polymerization step, the ends of the gradient macromers having lactic acid residues and caprolactone residues are linked to each other by a condensation reaction to form a multi-polymer. The reaction temperature for the condensation reaction is preferably 10°C to 100°C, more preferably 20°C to 50°C. The reaction time is preferably 1 day or more, more preferably 2 days or more. However, if the reaction is carried out for too long, a problem of coloring of the polymer occurs, so the reaction time is preferably 2 to 4 days.

[0086] The polyester copolymer according to the present embodiment has a structure in which two or more macromer units are linked together, and the rate of the faster initial polymerization rate of the monomer A or the monomer B is V X , the slower one is V Y In this case, the macromer unit is 1.1≦V X / V Y It is preferable that the polyester copolymer has as main structural units monomer A residues and monomer B residues satisfying the above condition of ≦40. 1.1≦V X / V Y By forming a polyester copolymer having a structure in which two or more macromer units, which are polyester copolymers having monomer A residues and monomer B residues as main constituent units satisfying the relationship ≦40, are linked together, it is possible to obtain macromer units with a gradient structure, and as a result, the polyester copolymer of this embodiment is preferred because it has a multi-gradient structure.

[0087] In this specification, the term "macromer" refers to the polyester copolymer obtained in the above-mentioned macromer synthesis step, and since this polyester copolymer is to be used in the above-mentioned multi-polymerization step after the macromer synthesis step, this polyester copolymer is referred to as a "macromer" to avoid confusion. The term "macromer unit" refers to a portion consisting of one macromonomer in the molecular chain of a polyester copolymer. For example, when two macromers are linked to form a polyester copolymer, the polyester copolymer is a polyester copolymer having a structure in which two macromer units are linked.

[0088] In addition, "two types of monomer residues in a macromer unit are 'main structural units'" means that the sum of the number of the two types of monomer residues is 50 mol % or more of the total number of all monomer residues contained in the entire macromer unit including other monomer residues (100%), and the number of each residue is 20 mol % or more of the total number of all monomer residues contained in the entire macromer unit (100%). For example, "monomer A residue and monomer B residue are main structural units" means that the sum of the number of these two types of monomer residues is 50 mol % or more of the total number of all monomer residues contained in the entire macromer unit (100%), the number of monomer A residues is 20 mol % or more, and the number of monomer B residues is 20 mol % or more. Here, the molar fractions of monomer A residue, monomer B residue, and other residues can be determined from the area value of the signal derived from each residue by nuclear magnetic resonance (NMR) measurement. For example, when the monomer A residue is a lactic acid residue and the monomer B residue is a caprolactone residue, their molar fractions can be measured by the method described in Measurement Example 1 below.

[0089] Here, V, which is the faster or slower initial polymerization rate for monomer A or monomer B, is X and V Y can be obtained by the following method. Equimolar amounts of monomer A and monomer B are mixed, a solvent and catalyst are added as necessary, and the polymerization reaction is started by adjusting the temperature and other conditions so that the R value is the same as the above-mentioned R value of the polyester copolymer finally synthesized or to be synthesized within an error range of 10%. Sampling is taken periodically from the sample during polymerization, and the remaining amounts of monomer A and monomer B are measured. The remaining amounts are measured, for example, by chromatography or nuclear magnetic resonance (NMR) measurement. The amount of monomer used in the polymerization reaction can be obtained by subtracting the remaining amounts from the charged amounts. When the amount of monomer used in the polymerization reaction is plotted against the sampling time, the initial slope of the curve is V X , V Y It is.

[0090] If monomer A has a higher initial polymerization rate than monomer B, then when such monomer A and monomer B are reacted, there is a high probability that monomer A will bond to the end of the polymer during polymerization in the early stages of polymerization. On the other hand, in the later stages of polymerization when monomer A is consumed and its concentration in the reaction solution decreases, there is a high probability that monomer B will bond to the end of the polymer during polymerization. As a result, a gradient polymer is obtained in which the proportion of monomer A residues gradually decreases from one end. Such a gradient polymer has low crystallinity and suppresses the increase in Young's modulus. To facilitate the formation of such a gradient structure, V X / V Y is more preferably 1.3 or more, and even more preferably 1.5 or more. On the other hand, if the difference in the polymerization rate between monomer A and monomer B is too large, the gradient polymer will have a structure similar to a block polymer in which only monomer A is polymerized and then monomer B is polymerized, and the crystallinity will become high, which may lead to an increase in Young's modulus. X / V Y is more preferably 30 or less, even more preferably 20 or less, and even more preferably 10 or less.

[0091] Such preferred combinations of monomer A and monomer B include dilactide and ε-caprolactone, glycolide and ε-caprolactone, glycolide and dilactide, dilactide and dioxepanone, ethylene oxalate and dilactide, dilactide and δ-valerolactone, and glycolide and δ-valerolactone.

[0092] As described above, the stent according to the present embodiment preferably contains a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units as the bioabsorbable polyester copolymer. The stent according to the present embodiment also preferably contains a water-soluble polymer different from the bioabsorbable polyester copolymer. The content of the water-soluble polymer different from the bioabsorbable polyester copolymer is not particularly limited, but the stent according to the present embodiment more preferably contains 0.1 to 50 mass %, and even more preferably contains 0.1 to 10 mass %, of the water-soluble polymer different from the bioabsorbable polyester copolymer.

[0093] In addition, from the viewpoint of suppressing mucus adhesion, the stent of this embodiment preferably has a water-soluble polymer content defined by the following formula of 0.1% by mass or more and 25% by mass or less, more preferably 1.0% by mass or more and 10% by mass or less, and even more preferably 2.0% by mass or more and 5.0% by mass or less.

[0094] Water-soluble polymer content (mass%) = [M1 / (M1+M2)×100] M1: mass of water-soluble polymer M2: Mass of polyester copolymer

[0095] The water-soluble polymer other than the bioabsorbable polyester copolymer refers to a polymer other than the bioabsorbable polyester copolymer that dissolves in water. More specifically, whether a certain polymer belongs to this category can be determined by evaluating whether it dissolves in water according to the measurement method described below.

[0096] The water-soluble polymer is composed of a material having water solubility. However, the water-soluble polymer may contain additives other than the above materials as long as they do not impair the expression of water solubility. Here, 1 g of the polymer is added to 100 mL of water heated to 37°C, stirred for 3 hours, and visually confirmed. If the polymer is dissolved, the material (polymer) is judged to be "water-soluble". The water-soluble material is preferably a material that is soluble in 100 parts by mass of water at 37°C at a rate of 1 part by mass or more.

[0097] Examples of water-soluble polymers other than bioabsorbable polyester copolymers include polyalkylene glycols (e.g., polyethylene glycol, polypropylene glycol, etc.), polyacrylic acid, polymethacrylic acid, polyacrylamide, polyvinylpyrrolidone, polyvinylacetamide, polymaleic acid, polysulfonic acid, polyvinyl alcohol, polyethyleneimine, carboxymethylcellulose, alginic acid, polyphosphoric acid, starch, agar, gelatin, pullulan, dextrin, xanthan gum, and salts, copolymers, or copolymer salts thereof. Of these, polyalkylene glycols, polyacrylic acid, polymethacrylic acid, polyacrylamide, polyvinylpyrrolidone, polyvinyl alcohol, and salts, copolymers, or copolymer salts thereof are particularly preferred, polyalkylene glycols, polyvinyl alcohol, and salts, copolymers, or copolymer salts thereof are more preferred, and polyalkylene glycols are most preferred.

[0098] As described above, the stent according to this embodiment preferably contains, as the bioabsorbable polyester copolymer, a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units. The stent also preferably contains a bioabsorbable homopolymer. In other words, a stent containing both a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units and a homopolymer as the bioabsorbable polyester copolymer is a preferred embodiment of the stent according to this embodiment.

[0099] That is, in the stent according to this embodiment, the bioabsorbable polyester copolymer may further contain a polyester other than the bioabsorbable polyester copolymer. Examples of such polyesters other than bioabsorbable polyester copolymers include polyesters selected from the group consisting of polyglycolic acid, polylactic acid (D, L, DL form), polyε-caprolactone, polyhydroxybutyrate, polyhydroxybutyrate valerate, polyorthoester, polyhydroxyvaleric acid, polyhydroxyhexanoic acid, polybutylene succinate, polytrimethylene terephthalate, polyhydroxyalkanoate, and polydioxanone.

[0100] In particular, the stent according to the present embodiment preferably contains either a copolymer of lactic acid and glycolic acid or a copolymer of lactic acid and ε-caprolactone, and the stent according to the present embodiment may further contain polylactic acid.

[0101] In addition, in the case where the stent of this embodiment contains both a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units and a homopolymer as the bioabsorbable polyester copolymer, in order to maintain the ability to follow the movement of biological organs, the content of the homopolymer in the stent is preferably 50 mass % or less in the bioabsorbable polyester copolymer (100 mass %), more preferably 30 mass % or less, and while there is no particular lower limit, the content of the homopolymer in the stent is preferably 1 mass % or more.

[0102] The stent according to this embodiment preferably has a Young's modulus of 0.1 MPa or more and 50 MPa or less, as measured according to JIS K6251 (2017). This measurement method is as described in Measurement Example 3 described later. Since the stent is placed in various body cavities or blood vessels, if the Young's modulus of the stent is too high, when an external force is applied to the stent due to deformation such as bending or curvature, the stent may compress, abrade, puncture, or otherwise damage the tissue around the placement site. Therefore, the Young's modulus of the stent is preferably 50 MPa or less, more preferably 20 MPa or less, and even more preferably 10 MPa or less. On the other hand, if the Young's modulus of the stent is too low, the stent will not be able to maintain its shape when an external force is applied to the stent due to deformation such as bending or curvature, so the Young's modulus of the stent is preferably 0.1 MPa or more, more preferably 0.5 MPa or more, and even more preferably 1.0 MPa or more.

[0103] The stent according to this embodiment preferably has a tensile strength of 4 MPa or more as measured according to JIS K6251 (2017). The measurement method is as described in Measurement Example 3 described later. Tensile strength is a factor directly related to the breaking strength of the stent. Therefore, the tensile strength of the stent is preferably 4 MPa or more, and more preferably 5 MPa or more. For a stent used in a site where more severe deformation such as bending or curvature occurs, the tensile strength is preferably 10 MPa or more, and more preferably 15 MPa or more. The tensile strength of the stent is preferably as high as possible, and there is no particular upper limit to the tensile strength, but the practical upper limit of the tensile strength is thought to be about 500 MPa.

[0104] The stent according to this embodiment preferably has a breaking elongation of 200% or more as measured according to JIS K6251 (2017). The measurement method is as described in Measurement Example 3 described later. The breaking elongation is a factor indicating the breaking strength of the stent. In the case of a stent that is installed in an environment where multiple physical actions such as bending, stretching, and compression are applied due to movement in a body cavity or a blood vessel, the breaking elongation of the stent is preferably 200% or more. In the case of a stent used in a site where more severe deformation occurs, the breaking elongation of the stent is preferably 500% or more, more preferably 800% or more, and even more preferably 1000% or more. The greater the breaking elongation of the stent, the more preferable it is. There is no particular upper limit to the breaking elongation, but the practical upper limit of the breaking elongation is thought to be about 2500%.

[0105] In addition, since the stent according to the present embodiment is used by being placed in various body cavities or blood vessels, it is necessary for the stent to have restoring ability (i.e., the ability to return to its original shape) even if it is deformed by multiple physical actions such as bending, stretching, and compression due to the movement of the inner surface of the body cavity or blood vessel. In addition, restoring ability is a characteristic that is also required for inserting and transporting a deformed stent into a trocar or catheter used for transporting the stent, and for the stent to restore to its original shape and be firmly fixed to the airway wall when it exits the outlet of the trocar at the destination. Therefore, it is important that the stent according to the present embodiment has a restoring ability defined by the following formula of 40% or more. Note that the restoring ability can be quantitatively evaluated from the following formula as in Measurement Example 3 described later.

[0106] Restorability (%) = ((L0 x 2 - L1) / L0) x 100 L0: initial length L1: The length after applying a tensile stress in the longest direction of the stent to generate a tensile strain of 100% of the initial length L0, repeated 10 times

[0107] The closer the restoring ability of the stent is to 100%, the less likely it is to lose its intended function due to deformation. Since the stent is subjected to multiple physical actions such as bending, stretching, and compression due to the movement of the inner surface of the body cavity or blood vessel, the restoring ability of the stent according to this embodiment is preferably 40% or more, more preferably 50% or more, even more preferably 70% or more, even more preferably 75% or more, even more preferably 80% or more, and particularly preferably 85% or more. The greater the restoring ability of the stent, the more preferable it is, with the upper limit of the restoring ability being 100%.

[0108] Furthermore, since the stent according to this embodiment is used by being placed in various body cavities or blood vessels, it is preferable that mucus adhesion can be suppressed. Therefore, the stent according to this embodiment preferably has a mucus adhesion amount (%) defined by the following formula of 60% or less. The mucus adhesion amount can be quantitatively evaluated from the following formula as in Measurement Example 5 described later.

[0109] Mucus adhesion amount (%) = (As-Asb) x 100 / (Ac-Acb) As: absorbance of sample at 450 nm Asb: absorbance at 450 nm of blank solution of sample (PBS instead of mucin solution, incubated overnight) Ac: absorbance of polylactic acid stent at 450 nm Acb: Absorbance at 450 nm of blank solution of polylactic acid stent (incubated overnight with PBS instead of mucin solution)

[0110] The closer to 0% the amount of mucus adhesion on the stent is, the more the stent can suppress mucus adhesion. Since mucus adhesion can cause complications such as stent stenosis and infection, the stent according to this embodiment preferably has a mucus adhesion amount of 60% or less, and more preferably 50% or less.

[0111] The stent of this embodiment can be produced (molded or formed) by molding such a bioabsorbable polyester copolymer using melt molding, solvent molding, electrospinning, or a molding method using a 3D printer.

[0112] The melt molding method is a method in which a polymer is heated and melted, and molded using a mold, an extruder, a press, etc., and a stent can be produced by molding it into a fiber, film, tube, etc. For example, the copolymer described in this specification can be heated to 200°C in an extruder equipped with a 1 mm diameter die and extruded to form the polymer into a thread, and the fiber can be woven or knitted to produce a stent.

[0113] Solvent molding is a method in which a polymer is dissolved in a solvent, the resulting solution is poured into a mold or coagulation bath, and the solvent and solute are separated to form the polymer, and a stent can be produced by molding the polymer into a fiber, film, tube, etc. A specific example of the solvent molding method is to immerse a rod with a diameter of 0.5 to 20 mm in a 20% polymer solution in chloroform, then pull it out, evaporate the solvent, and then immerse the rod again, repeating this process 5 to 50 times, and finally pulling out the core rod to form a tube-shaped stent.

[0114] The electrospinning method is a technique that can produce a fiber structure made of nanofibers with a diameter of several nanometers by applying a high voltage to a polymer solution in a spinning nozzle, and the thickness of the fiber structure can be adjusted to a desired range by adjusting the spinning time. For example, a tube-shaped fiber structure can be produced by accumulating polymer fibers while rotating a cylindrical collector and then removing the collector from the fiber structure. In addition, a custom-made stent can be produced by using the above-mentioned bioabsorbable polyester copolymer as an ink material for a 3D printer. The present disclosure also relates to a method for producing a stent, which includes a step of 3D printing using a printing material containing the above-mentioned bioabsorbable polyester copolymer.

[0115] The shape of the stent according to this embodiment will be described below. Figures 1 to 3 are schematic diagrams showing a stent according to a specific example of this embodiment. The shape of the stent according to this embodiment is not particularly limited, but the stent may include a tubular structure portion as shown in Fig. 1. In the stent 10 shown in Fig. 1, the inside of the tubular structure portion is the inner surface 11, and the surface other than the inner surface 11 is the outer surface 12. The tubular structure portion may be radially expandable. Fig. 3 shows another stent according to this embodiment, which has a branch. The stent according to this embodiment preferably has a shape that matches the shape of the respiratory tract to which the stent is applied.

[0116] Furthermore, the size of the stent according to this embodiment is not particularly limited, but in order to function as a respiratory stent, it is preferable that the outer diameter of the tubular structure portion is 4 mm or more and 24 mm or less, and the thickness is 0.2 mm or more and 2 mm or less, more preferably the outer diameter is 6 mm or more and 20 mm or less, and the thickness is 0.25 mm or more and 1.5 mm or less, and even more preferably the outer diameter is 6 mm or more and 20 mm or less, and the thickness is 0.3 mm or more and 1.2 mm or less. By controlling the thickness in this way, a safer stent with less risk of causing complications can be provided. The thickness of the stent can be calculated by the formula "(outer diameter of the stent - inner diameter of the stent) / 2".

[0117] In addition, the "outer diameter" here is defined to include protrusions or irregularities on the outer peripheral surface if they exist. If there are no protrusions or irregularities on the outer peripheral surface, the "outer diameter" is defined not to include the protrusions or irregularities, and it is sufficient that there is a portion of the stent with an outer diameter within the above range.

[0118] Although the shape of the stent according to the present embodiment is not limited to the above, it is preferable that the outer surface of the stent has projections or irregularities to prevent the stent from moving after placement. It is preferable that the number of projections or irregularities is multiple. The multiple protrusions or irregularities may be arranged regularly or randomly. When protrusions are arranged on the outer surface of the stent, for example, the multiple protrusions 40A may be arranged regularly or randomly as shown in Figures 1, 2 and 3.

[0119] Furthermore, the protrusions or irregularities may be arranged locally on the outer surface, may be arranged over the entire outer surface, or may be scattered over the outer surface.

[0120] The shape of the protrusion 40A is not particularly limited, and may be, for example, a hemisphere, a cylinder, a cone, a column, a polygonal pyramid, a hook, etc. More specifically, the shape of the protrusion 40A may be, for example, a hemisphere as shown in Fig. 1, or a cylinder as shown in Fig. 2. The shape of the unevenness is not particularly limited, and may be pleated, embossed, patterned (e.g., line-like, wavy), etc. More specifically, the shape of the unevenness may be pleated as shown in Fig. 6, or may be a line-like pattern like the protrusions 40B in the stent shown in Fig. 4.

[0121] In addition, when protrusions or irregularities are arranged on the outer surface of the stent, the size of the protrusions is not particularly limited, and the size (height) of the protrusions or irregularities is preferably 4.0 mm or less, more preferably 3.0 mm or less, and even more preferably 2.0 mm or less from the viewpoint of suppressing irritation to tissues. In addition, in order to exert the effect of preventing the stent from moving after placement, the size (height) of the protrusions or irregularities is preferably 0.1 mm or more, and more preferably 0.2 mm or more.

[0122] In order to prevent the stent from moving after placement, for example, highly accurate anatomical data based on 3DCT can be created, and a stent of a size and shape suitable for the shape of the respiratory system of the patient to which the stent is to be applied can be made using 3D printing technology based on anatomical analysis. In addition, a stent in which the difference between the inner diameter of the respiratory system in which the stent is placed and the outer diameter of the stent is 10% or less is preferable, 8% or less is more preferable, 6% or less is even more preferable, and 5% or less is particularly preferable.

[0123] The composition of the stent according to this embodiment is not particularly limited other than that it contains the bioabsorbable polyester copolymer described above, and the entire stent or only a portion of the stent may be made of the bioabsorbable polyester copolymer. Furthermore, the stent according to this embodiment may have a base material, and may have a resin layer on the surface of the base material.

[0124] [Base material] The material of the base material (stent base material) of the stent is not particularly limited, and may include metal or resin. Examples of metals include stainless steel, cobalt alloys, titanium alloys, nickel-titanium alloys (nitinol), and the like. Examples of resins include polyurethane, polyester, bioabsorbable polyester copolymers, PTFE (polytetrafluoroethylene), and silicone resins. From the viewpoints of biocompatibility, mechanical properties, processability, and the like, bioabsorbable polyester copolymers or silicone resins are preferred.

[0125] When the material of the substrate contains a bioabsorbable polyester copolymer, the substrate may be formed entirely from the bioabsorbable polyester copolymer, or may contain only a portion of the bioabsorbable polyester copolymer. Examples of the bioabsorbable polyester copolymer contained in the substrate include those mentioned above, and preferred examples of the bioabsorbable polyester copolymer contained in the substrate are also the same. The substrate may be formed from one material or from two or more materials.

[0126] [Resin Layer] The stent of this embodiment may have a resin layer as a surface layer. The resin layer in the surface layer of the stent is formed on the surface of a substrate and is made of a resin.

[0127] When the stent according to the present embodiment has a resin layer, it preferably contains a mixed layer between the substrate and the resin layer, in which the components of the substrate and the components of the resin layer are mixed. Examples of resins that can be used as the material for the resin layer include polyurethane, polyester, bioabsorbable polyester copolymers, PTFE (polytetrafluoroethylene), and water-soluble polymers, with bioabsorbable polyester copolymers and water-soluble polymers being preferred. When the material of the resin layer contains a bioabsorbable polyester copolymer, the resin layer may be formed entirely from the bioabsorbable polyester copolymer, or may contain only a portion of the bioabsorbable polyester copolymer. The stent according to the present embodiment has a resin layer on the surface of the stent, and the resin layer (100% by mass) preferably contains 50% by mass or more of the bioabsorbable polyester copolymer. In 100% by mass of the resin layer, the content of the bioabsorbable polyester copolymer is more preferably 70% by mass or more, even more preferably 95% by mass or more, and particularly preferably 100% by mass.

[0128] Fig. 5 is a cross-sectional view taken along the line AA of a stent 10 according to one specific example of the present embodiment shown in Fig. 1. As shown in Fig. 5, the stent 10 preferably includes a mixed layer 22 between the substrate 21 and the resin layer 23, in which a component of the substrate 21 and a component of the resin layer 23 are mixed.

[0129] The mixed layer formed between the substrate and the resin layer may be one in which a part of the resin constituting the resin layer penetrates into the substrate. The mixed layer may also be one in which a part of the material constituting the substrate penetrates into the resin layer. When the stent according to the present embodiment includes a mixed layer, the layer containing the resin has a laminated structure of two or more layers including a resin layer and a mixed layer.

[0130] The stent according to the present embodiment preferably has a resin layer on at least a part of the inner surface 11, and may have a resin layer formed on the entire inner surface 11. The stent according to the present embodiment may further have a resin layer on a part or the entire outer surface 12. From the viewpoint of biocompatibility, the stent according to this embodiment preferably has a resin layer on the entire inner surface 11 and the entire outer surface 12, that is, on the entire surface of the stent.

[0131] For the bioabsorbable polyester copolymer and water-soluble polymer forming the resin layer, the above explanations regarding the bioabsorbable polyester copolymer and water-soluble polymer can be applied as is. The resin layer may be formed from one type of material, or two or more types of materials.

[0132] Moreover, the stent according to this embodiment can also be used as a drug-eluting stent by carrying or adsorbing a drug. The stent according to the present embodiment can be applied to applications in which the stent is implanted into the narrowed site of various body cavities or vessels (e.g., the vascular system, esophagus, gastrointestinal tract, large and small intestines, bile duct, pancreatic duct, lung duct, ureter, nasal cavity, trachea, etc.) to secure the lumen, but is not limited thereto. Since the stent according to the present embodiment contains a bioabsorbable polyester copolymer, it has high restoring property and can be used as a drug-eluting stent, so the stent according to the present embodiment is suitable for use as a stent placed in the vascular system, trachea, nasal cavity, etc., and since it can suppress mucus adhesion, the stent according to the present embodiment is particularly suitable for use as a stent placed in the trachea, nasal cavity, etc.

[0133] <How to relieve respiratory constriction and ensure airflow> The method for relieving the constriction of the respiratory system and ensuring airflow according to the present embodiment is a method for relieving the constriction of the respiratory system and ensuring airflow using a stent for the respiratory system, The stent comprises a bioabsorbable polyester copolymer according to the above embodiments.

[0134] The method of this embodiment for relieving respiratory stenosis and ensuring airflow uses the above-mentioned stent, which can suppress mucus adhesion, is a method with excellent biocompatibility, and can suppress the occurrence of complications. In the method for relieving respiratory tract constriction and ensuring airflow according to this embodiment, the above description can be applied as is to the stent. [Industrial Applicability]

[0135] The stent according to the present embodiment is a stent that can be applied to respiratory organs such as the airway, oral cavity, nasal cavity, larynx, trachea, bronchi, bronchioles, lungs, etc. The stent according to the present embodiment can also be applied to applications in which the stent is transplanted into narrowed sites of various body cavities or vessels other than the respiratory organs (e.g., the vascular system, esophagus, gastrointestinal tract, large intestine and small intestine, bile duct, pancreatic duct, pulmonary duct, ureter, etc.) to secure the lumen. EXAMPLES

[0136] The present disclosure will be described in detail below with reference to examples and comparative examples, but the present disclosure is not limited thereto.

[0137] (Measurement Example 1: Measurement of the Molar Fraction and R Value of Each Residue by Nuclear Magnetic Resonance (NMR)) The purified polyester copolymer was dissolved in deuterated chloroform. 1 The ratios of lactic acid monomer residues and caprolactone monomer residues in the polyester copolymer were calculated by H-NMR. 1By H homospin decoupling, the combination of two adjacent monomer residues was separated by the signal indicating whether lactic acid or caprolactone is adjacent to a certain monomer residue for the methine group of lactic acid (around 5.10 ppm), the α-methylene group of caprolactone (around 2.35 ppm), and the ε-methylene group (around 4.10 ppm), and the peak area of ​​each was quantified. When δ-valerolactone was used instead of ε-caprolactone, the combination of two adjacent monomer residues was separated by the signal indicating whether lactic acid or valerolactone is adjacent to a certain monomer residue for the methine group of lactic acid (around 5.10 ppm), the α-methylene group of valerolactone (around 2.35 ppm), and the ε-methylene group (around 4.10 ppm), and the peak area of ​​each was quantified. From the peak area ratio, [AB] and R values ​​were calculated. Here, [AB] is the molar fraction of the structure in which lactic acid residues and caprolactone residues or valerolactone residues are adjacent to each other in the polyester copolymer. Specifically, [AB] is the ratio (%) of the total number of AB dyads and the number of BA dyads to the total number of AA dyads, AB dyads, BA dyads, and BB dyads. The results are shown in Table 1. Device name: JNM-ECZ400R (manufactured by JEOL Ltd.) 1 H homospin decoupling irradiation position: 1.66 ppm Solvent: deuterated chloroform Measurement temperature: room temperature (20℃~25℃)

[0138] R = [AB] / (2[A][B]) × 100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residue and monomer B residue are adjacent to each other in the polyester copolymer

[0139] (Measurement Example 2: Measurement of weight average molecular weight by gel permeation chromatography (GPC)) The weight average molecular weight of the polyester copolymer used was measured under the conditions shown below.

[0140] (GPC measurement conditions) Device name: Prominence (manufactured by Shimadzu Corporation) Mobile phase: Chloroform (for HPLC) (manufactured by Wako Pure Chemical Industries, Ltd.) Flow rate: 1mL / min Column: TSKgel GMHHR-M (φ7.8 mm × 300 mm; Tosoh Corporation) Detector: UV (254 nm), RI Column and detector temperature: 35℃ Standard material: polystyrene The purified polyester copolymer was dissolved in chloroform and passed through a 0.45 μm syringe (DISMIC-13HP; manufactured by ADVANTEC) to remove impurities, and then the weight average molecular weight of the polyester copolymer was calculated by measuring the molecular weight by GPC. The results are shown in Table 1.

[0141] (Measurement example 3: Tensile test) A 40 mm x 5 mm test piece was cut out from the stent (thickness 1.0 mm) produced in each Example and Comparative Example, and a tensile test was performed under the following conditions according to JIS K6251 (2017) using a Tensilon universal testing machine RTM-100 (manufactured by Orientec Co., Ltd.), and the breaking elongation and tensile strength were calculated. In addition, in the stress-strain graph, the slope of the linear approximation equation obtained from the data of five points from the start of stress generation was calculated as Young's modulus. If necessary, two marks were made on the specimen using a suitable marker, the marks being accurately and clearly marked perpendicular to the parallel portion of the specimen and equidistant from the centre of the specimen while the specimen was in the unstrained position. Device name: EZ-1kNLX (Shimadzu Access) Gauge distance before test: 10mm Distance between grips: 10 mm (gripped at the marked line) Tensile speed: 500mm / min Load cell: 1kN Number of tests: 5

[0142] Furthermore, to evaluate the resilience, a tensile strain of 100% was generated relative to the distance between the grips before the test at a tensile speed of 500 mm / min (operation 1). Then, immediately after operation 1 (i.e., the shape retention time was set to 0 seconds), the tensile strain was relaxed at a speed of 500 mm / min, and the distance between the grips was returned to 5 mm (operation 2). Immediately after operation 2 (i.e., the shape retention time was set to 0 seconds), the above-mentioned operations 1 and 2 were performed again. This was repeated, and operations 1 and 2 were each performed 10 times, after which the resilience was calculated from the following formula using the obtained value of L1. (Resiliency (%)) = ((L0 x 2 - L1) / L0) x 100 L0: Initial length (gauge length before test) L1: The length after applying tensile stress in the longest direction of the stent to produce a tensile strain of 100% of the initial length 10 times (gauge line distance after test)

[0143] (Measurement Example 4: Evaluation of Water Solubility of Polymers) 1 g of the polymer was added to 100 mL of water at 37° C., and the resulting solution was stirred for 3 hours. After that, the solution was visually confirmed, and if the polymer was dissolved, the polymer was determined to be water-soluble.

[0144] (Measurement Example 5: In vitro mucus adhesion test) Mucin was purified from saliva and prepared to a concentration of 100 μg / mL to prepare a mucin solution. Stents were punched into disk shapes with a diameter of 4 mm and placed in 48 wells of a microtiter plate. 600 μL of mucin solution with a concentration of 100 μg / mL was added to each well and incubated at 37 °C for 20 to 24 hours. As a control, PBS was added instead of the mucin solution and incubated at 37 °C for 20 to 24 hours. After washing three times with PBS, blocking buffer (ThermoFisher Scientific 37570) was added and incubated at room temperature (20 °C to 25 °C) for 1 hour. After washing three times with PBS, WGA (Biotinylated Wheat Germ Agglutinin (WGA), Vector Laboratories B-1025-5, diluted 500 times with PBS) was added and incubated at room temperature for 1 hour. After washing three times with PBS, horseradish peroxidase (HRP)-labeled streptavidin (HRP-Streptavidin, Sigma-Aldrich RABHRP3-600UL) was added and incubated at room temperature for 1 hour. After washing three times with PBS, 250 μL of TMB (3,3′,5,5′-Tetramethylbenzidine (TMB) substrate, Thermo Scientific PI34028) solution was added and incubated at room temperature for 15 to 30 minutes. After that, the sample was removed, 250 μL of 2 mol / L sulfuric acid was added, and the absorbance at 450 nm was measured with a microplate reader. The amount of mucus (mucin) attached was calculated as shown in the following formula (1).

[0145] [Mucus adhesion amount] (Mucus adhesion amount (%)) = (As-Asb) x 100 / (Ac-Acb) Formula (1) As: absorbance of sample at 450 nm Asb: absorbance at 450 nm of blank solution of sample (PBS instead of mucin solution, incubated overnight) Ac: absorbance of polylactic acid stent at 450 nm Acb: Absorbance at 450 nm of blank solution of polylactic acid stent (incubated overnight with PBS instead of mucin solution)

[0146] [Phosphate buffer] The composition of the phosphate buffer solution is as follows: KCl 0.2g / L KH2PO40.2g / L NaCl 8.0g / L Na2HPO4(anhydrous) 1.15g / L EDTA 0.25g / L

[0147] <Example 1> 50.0 g of L-lactide (PURASORB L; manufactured by PURAC Co., Ltd.) and 39.6 g of ε-caprolactone (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) were collected as monomers and 0.46 g of hydroxypivalic acid was collected as an initiator in a separable flask. These were placed under an argon atmosphere, and 0.27 g of tin(II) octylate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) was added as a catalyst dissolved in 5.8 mL of toluene (ultra-dehydrated, manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) and reacted at 140°C for 9.5 hours to obtain a crude copolymer. The resulting crude copolymer was dissolved in 200 mL of chloroform and added dropwise to 3000 mL of hexane under stirring to obtain a precipitate, which was then dried under reduced pressure at 50° C. to obtain a macromer.

[0148] 50 g of the macromer, 2.9 g of 4-dimethylaminopyridinium p-toluenesulfonate (synthetic product) as a catalyst, and 1.2 g of 4,4-dimethylaminopyridine (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) were collected. These were dissolved in 200 mL of dichloromethane (dehydrated, manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) under an argon atmosphere, and 2.4 mL of diisopropylcarbodiimide (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) as a condensation agent was added, followed by condensation polymerization at room temperature (20°C to 25°C) for 20 hours. The reaction mixture was diluted with 220 mL of chloroform, 10 g of lactic acid was added, and the mixture was stirred for 3 hours. Then, 330 mL of ion-exchanged water was added, the mixture was stirred for 15 minutes, and the aqueous phase was removed by decantation. This process was repeated until the pH of the removed aqueous phase reached 7. The remaining organic phase was dropped into 2200 mL of methanol under stirring to obtain a precipitate. The precipitate was dried under reduced pressure at 50° C. to obtain the purified polyester copolymer of Example 1. The obtained polyester copolymer was cut into approximately 5 mm squares, then fed into an extruder, set at 100°C to 200°C, and extruded so that the diameter of the obtained filament was 1.75 mm. Using this filament and a fused deposition model 3D printer, a tubular molded body (inner diameter: 10 mm, thickness: 1.0 mm, length: 40 mm) was obtained as shown in Figure 6.

[0149] <Example 2> 60.0 g of L-lactide (PURASORB L; manufactured by PURAC Co., Ltd.) and 31.7 g of ε-caprolactone (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) were collected as monomers and 0.46 g of hydroxypivalic acid was collected as an initiator in a separable flask. These were placed under an argon atmosphere, and 0.27 g of tin(II) octylate (manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) was added as a catalyst dissolved in 5.8 mL of toluene (ultra-dehydrated, manufactured by FUJIFILM Wako Pure Chemical Industries, Ltd.) and reacted at 140°C for 9.5 hours to obtain a crude copolymer. The resulting crude copolymer was dissolved in 200 mL of chloroform and added dropwise to 3000 mL of hexane under stirring to obtain a precipitate, which was then dried under reduced pressure at 50° C. to obtain a macromer.

[0150] 50 g of the macromer, 2.1 g of 4-dimethylaminopyridinium p-toluenesulfonate (synthetic product) as a catalyst, and 0.87 g of 4,4-dimethylaminopyridine (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) were collected. These were dissolved in 200 mL of dichloromethane (dehydrated, manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) under an argon atmosphere, and 1.7 mL of diisopropylcarbodiimide (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd.) as a condensation agent was added, followed by condensation polymerization at room temperature (20°C to 25°C) for 20 hours.

[0151] The reaction mixture was diluted with 220 mL of chloroform, 10 g of lactic acid was added, and the mixture was stirred for 3 hours. Then, 330 mL of ion-exchanged water was added, the mixture was stirred for 15 minutes, and the aqueous phase was removed by decantation. This process was repeated until the pH of the removed aqueous phase reached 7. The remaining organic phase was dropped into 2200 mL of methanol under stirring to obtain a precipitate. The precipitate was dried under reduced pressure at 50° C. to obtain the purified polyester copolymer of Example 2. In addition, in the same manner as in Example 1, a tubular molded product, a stent, was obtained.

[0152] <Example 3> The purified polyester copolymer of Example 3 was obtained in the same manner as in Example 1, except that the amount of hydroxypivalic acid was changed to 0.45 g, the reaction temperature for obtaining a crude copolymer was changed to 150° C., the amount of 4-dimethylaminopyridinium p-toluenesulfonate was changed to 2.1 g, the amount of 4,4-dimethylaminopyridine was changed to 0.87 g, and the amount of diisopropylcarbodiimide was changed to 1.7 mL. In addition, in the same manner as in Example 1, a tubular molded product, a stent, was obtained.

[0153] <Example 4> 50.0g of L-lactide (PURASORB L; manufactured by PURAC) and 38.5mL of ε-caprolactone (manufactured by Wako Pure Chemical Industries, Ltd.) were collected as monomers in a separable flask. These were placed under an argon atmosphere, and 0.81g of tin(II) octylate (manufactured by Wako Pure Chemical Industries, Ltd.) dissolved in 14.5mL of toluene (ultra-dehydrated, manufactured by Wako Pure Chemical Industries, Ltd.) as a catalyst and ion-exchanged water as a co-initiator were added so that the monomer / co-initiator ratio was 142.9. After carrying out a co-catalyst reaction at 90°C for 1 hour, a copolymerization reaction was carried out at 150°C for 6 hours to obtain a crude copolymer.

[0154] The resulting crude copolymer was dissolved in 100 mL of chloroform and added dropwise to 1400 mL of stirred methanol to obtain a precipitate. This procedure was repeated three times, and the precipitate was dried under reduced pressure at 70° C. to obtain a macromer.

[0155] 30 g of the macromer, 0.28 g of 4-dimethylaminopyridinium p-toluenesulfonate (synthetic product) as a catalyst, and 0.10 g of 4,4-dimethylaminopyridine (manufactured by Wako Pure Chemical Industries, Ltd.) were collected. These were dissolved in dichloromethane (dehydrated, manufactured by Wako Pure Chemical Industries, Ltd.) to a concentration of 30% under an argon atmosphere, and 0.47 g of Amylene (manufactured by Tokyo Chemical Industry Co., Ltd.) as a condensation agent dissolved in 5 mL of dichloromethane was added, followed by condensation polymerization at room temperature (20°C to 25°C) for 2 days.

[0156] 30 mL of chloroform was added to the reaction mixture, and the resulting solution was dropped into 500 mL of methanol under stirring to obtain a precipitate. The precipitate was dissolved in 50 mL of chloroform, and the resulting solution was dropped into 500 mL of methanol under stirring to obtain a precipitate. This operation was repeated twice, and the precipitate was dried under reduced pressure at 50°C to obtain a purified polyester copolymer of Example 4. In addition, a stent, which is a tubular molded body, was obtained in the same manner as in Example 1.

[0157] <Example 5> 45 g of the polyester copolymer obtained in Example 3 and 5 g of polylactic acid (Nature3D) were added to 500 mL of chloroform (Fujifilm Wako Pure Chemical Industries, Ltd.) and dissolved, and the resulting solution was added to 3000 mL of hexane in a stirred state, and the resulting precipitate was dried at room temperature (20°C to 25°C) for 24 hours under normal pressure. The precipitate was further dried under reduced pressure at 50°C to obtain a polymer composition of Example 5. A stent, which is a tubular molded product, was also obtained in the same manner as in Example 1.

[0158] <Example 6> A polymer composition of Example 6 was obtained in the same manner as in Example 5, except that the amount of polyester copolymer was changed to 35 g and the amount of polylactic acid was changed to 15 g. A stent, which was a tubular molded product, was also obtained in the same manner as in Example 1.

[0159] <Example 7> 48.5 g of the polyester copolymer obtained in Example 1 and 1.5 g of polyethylene glycol (Sigma-Aldrich) were dissolved in 500 mL of chloroform (Fujifilm Wako Pure Chemical Industries, Ltd.), and the resulting solution was added to 3000 mL of hexane in a stirred state to obtain a precipitate. The resulting precipitate was dried at room temperature (20°C to 25°C) under normal pressure for 24 hours. The precipitate was further dried under reduced pressure at 50°C to obtain a polymer composition of Example 7. A stent, which is a tubular molded product, was also obtained in the same manner as in Example 1.

[0160] <Example 8> The same procedure as in Example 7 was carried out, except that the amount of the polyester copolymer obtained in Example 1 was 48 g and the amount of polyethylene glycol was 2.0 g, to obtain a polymer composition of Example 8. Also, a stent, which is a tubular molded product, was obtained in the same manner as in Example 1.

[0161] <Example 9> The polymer composition of Example 9 was obtained by the same procedure as in Example 7, except that the amount of the polyester copolymer obtained in Example 1 was 47.5 g and the amount of polyethylene glycol was 2.5 g. In addition, in the same manner as in Example 1, a tubular molded product, a stent, was obtained.

[0162] <Example 10> The purified polyester copolymer of Example 10 was obtained by the same synthesis method as in Example 1, except that the amount of 4-dimethylaminopyridinium p-toluenesulfonate was changed to 1.5 g and the amount of diisopropylcarbodiimide was changed to 1.2 mL. In addition, in the same manner as in Example 1, a tubular molded product, a stent, was obtained.

[0163] <Example 11> 50.0 g of L-lactide (PURASORB L; manufactured by PURAC) was collected as a monomer in a separable flask. This was placed under an argon atmosphere, and 0.81 g of tin(II) octoate (manufactured by Wako Pure Chemical Industries, Ltd.) was added as a catalyst dissolved in 14.5 mL of toluene (ultra-dehydrated, manufactured by Wako Pure Chemical Industries, Ltd.) and polymerized at 150°C for 3 hours. 38.5 mL of ε-caprolactone (manufactured by Wako Pure Chemical Industries, Ltd.) was added to this and polymerized at 150°C for 6 hours to obtain a crude copolymer. The crude copolymer was dissolved in 100 mL of chloroform, and the resulting solution was added dropwise to 1400 mL of methanol under stirring to obtain a precipitate. This operation was repeated three times, and the precipitate was dried under reduced pressure at 70° C. to obtain a purified polyester copolymer of Example 11. In addition, in the same manner as in Example 1, a tubular molded product, a stent, was obtained.

[0164] <Comparative Example 1> Polylactic acid (Nature3D) was cut into approximately 5 mm squares and then placed in an extruder, the temperature was set at 100°C to 200°C, and a polyester copolymer was extruded so that the diameter of the resulting filament was 1.75 mm. Using this filament, a tubular molded product, a stent, was obtained in the same manner as in Example 1.

[0165] The results of various measurements of the polymers and stents obtained in Examples 1 to 11 and Comparative Example 1 are shown in Table 1.

[0166] [Table 1]

[0167] [Table 2]

[0168] In Table 1, "monomer A residue ratio" refers to the molar ratio (mol%) of monomer A residue to the total number of moles of monomer A residue and monomer B residue (100%). In Examples 1 to 11, monomer A is L-lactide, and monomer B is ε-caprolactone. X is the initial polymerization rate of L-lactide, V Y represents the initial polymerization rate of ε-caprolactone. [Explanation of symbols]

[0169] 10 Stents 11 Inner surface 12 Outer surface 21 Base material 22 Mixed layer 23 Resin layer 40A protrusion 40B Convex

Claims

1. 1. A respiratory stent comprising a bioabsorbable polyester copolymer, The bioabsorbable polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, A stent, in which the R value represented by the following formula, in which the two types of ester bond-forming monomers are monomer A and monomer B, respectively, is 0.25 or more and 0.99 or less. R=[AB] / (2[A][B])×100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residues and monomer B residues are adjacent to each other in the polyester copolymer

2. The stent of claim 1, having a Young's modulus measured according to JIS K6251 (2017) of 0.1 MPa or more and 50 MPa or less.

3. 3. The stent according to claim 1, wherein the restoring ability defined by the following formula is 40% or more. Recovery (%) = (L ０ × 2 - L １ ) / L ０ × 100 L ０ : Initial length L １ : A tensile stress is applied in the direction of the longest length of the stent to obtain the initial length L ０ The length after repeating the operation of causing 100% tensile strain 10 times

4. The stent according to any one of claims 1 to 3, wherein the amount of mucus adhered thereto is 60% or less, as defined by the following formula: Mucus adhesion amount (%) = (As-Asb) x 100 / (Ac-Acb) As: absorbance of the sample at 450 nm Asb: absorbance at 450 nm of blank solution of sample (PBS instead of mucin solution, incubated overnight) Ac: absorbance of polylactic acid stent at 450 nm Acb: absorbance at 450 nm of blank solution of polylactic acid stent (incubated overnight using PBS instead of mucin solution)

5. The stent of any one of claims 1 to 4, further comprising a water soluble polymer.

6. The stent according to claim 5, wherein the content of the water-soluble polymer defined by the following formula is 0.1% by mass or more and 25% by mass or less. Water-soluble polymer content (mass%) = [M1 / (M1+M2) x 100] M1: mass of water-soluble polymer M2: mass of polyester copolymer

7. The stent of claim 5 or 6, wherein the water soluble polymer is a polyalkylene glycol.

8. The stent according to any one of claims 1 to 7, wherein the bioabsorbable polyester copolymer is contained in an amount of 50 mass% or more based on 100 mass% of the stent.

9. The monomer A is at least one selected from the group consisting of lactic acid and glycolic acid, The stent according to any one of claims 1 to 8, wherein the monomer B is at least one selected from the group consisting of caprolactone and δ-valerolactone.

10. The stent according to any one of claims 1 to 9, having an outer diameter of 4 mm or more and 24 mm or less, and a thickness of 0.2 mm or more and 2 mm or less.

11. The stent according to any one of claims 1 to 10, having a plurality of protrusions or a plurality of irregularities on its outer surface.

12. The stent according to claim 11, wherein the height of the projections or irregularities is 0.1 mm or more and 3.0 mm or less.

13. A method for producing a stent according to any one of claims 1 to 12, comprising the step of 3D printing using a printing material comprising the bioabsorbable polyester copolymer.

14. A method for relieving respiratory tract narrowing and ensuring airflow using a respiratory stent, comprising: the stent comprises a bioabsorbable polyester copolymer; The bioabsorbable polyester copolymer is a polyester copolymer having two types of ester bond-forming monomer residues as main constituent units, The method, wherein when the two types of ester bond-forming monomers are monomer A and monomer B, respectively, the R value represented by the following formula is 0.25 or more and 0.99 or less. R=[AB] / (2[A][B])×100 [A]: Molar fraction (%) of monomer A residues in the polyester copolymer [B]: Molar fraction (%) of monomer B residues in the polyester copolymer [AB]: mole fraction (%) of structures (AB and BA) in which monomer A residues and monomer B residues are adjacent to each other in the polyester copolymer

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