WO2025158984A1 - Medical material, method for producing same, and medical tool - Google Patents

Abstract

Provided is a medical material having mechanical properties comparable to those of polytetrafluoroethylene (PTFE) and having slipperiness (slidability) superior to that of PTFE.　A medical material according to the present invention comprises: a block copolymer having a structural unit (A) derived from a reactive monomer having an epoxy group and a structural unit (B) derived from a hydrophilic monomer; and at least one hydrophobic resin selected from the group consisting of polyvinyl chloride resins and polyurethane elastomers. The medical material contains the hydrophobic resin in an amount greater than the amount of the block copolymer, exhibits a sliding resistance of 50 gf or less, and has a tensile strength of 8.0 MPa or more and/or a tensile elongation of more than 80%.

Description

Medical materials, their manufacturing methods, and medical devices

The present invention relates to medical materials, their manufacturing methods, and medical devices.

Medical devices inserted into the body, such as plastic insertion needles, dilators, sheaths (introducers), catheters, and medical tubing, are required to exhibit excellent slip properties (sliding ability, lubricity) to reduce tissue damage to blood vessels and other tissues and improve operability for the surgeon.

Polytetrafluoroethylene (PTFE) is widely used as a material for these medical devices due to its excellent properties, including chemical resistance, non-stickiness, and low friction (see, for example, JP 8-33704 A).

On the other hand, organic fluorine compounds (PFAS, perfluoroalkyl compounds, and polyfluoroalkyl compounds) are known to be highly persistent and accumulate, raising concerns that they place a heavy burden on the environment. For this reason, the EU is currently considering a regulatory proposal that would apply to all PFAS. PFAS includes polytetrafluoroethylene (PTFE).

For this reason, there is a need to develop a resin that has the same level of slipperiness as PTFE and mechanical properties that make it suitable for use in medical devices.

The present invention was made in consideration of the above circumstances, and aims to provide a medical material or medical device that has mechanical properties (particularly tensile strength and tensile elongation) comparable to those of polytetrafluoroethylene (PTFE) and superior slip properties (slidability) to PTFE.

The inventors conducted extensive research to solve the above problems. As a result, they discovered that the above problems could be solved by combining a block copolymer having specific structural units with a specific hydrophobic resin, leading to the completion of the present invention.

The above objective can be achieved by the present invention, which has the following configuration, and includes the following aspects and configurations.

One aspect of the present invention is
1. A medical material comprising a block copolymer having a structural unit (A) derived from a reactive monomer having an epoxy group and a structural unit (B) derived from a hydrophilic monomer, and at least one hydrophobic resin selected from the group consisting of polyvinyl chloride resins and polyurethane elastomers, wherein the content of the hydrophobic resin in the medical material is greater than the content of the block copolymer, the medical material has a sliding resistance of 50 gf or less, and satisfies at least one of a tensile strength of 8.0 MPa or more and a tensile elongation of more than 80%.
2. In the medical material described in 1 above, the hydrophobic resin is preferably contained in a proportion of 125 parts by mass or more and 300 parts by mass or less per 100 parts by mass of the block copolymer.
3. In the medical material described in 1. or 2. above, the reactive monomer having an epoxy group preferably includes at least one selected from the group consisting of glycidyl acrylate, glycidyl methacrylate, 3,4-epoxycyclohexylmethyl acrylate, 3,4-epoxycyclohexylmethyl methacrylate, β-methylglycidyl methacrylate, and allyl glycidyl ether.
4. In the medical material according to any one of 1. to 3. above, it is preferable that the hydrophilic monomer contains at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, 2-hydroxyethyl methacrylate, and N-vinylpyrrolidone.

Another aspect of the present invention is
5. A medical device comprising or consisting of the medical material described in any one of 1. to 4. above.

Yet another aspect of the present invention is
6. A medical device comprising a substrate layer and a coating layer containing or consisting of the medical material described in any one of 1. to 4. above.
7. The medical device described in 5. or 6. above is preferably a plastic insertion needle, a dilator, a sheath (introducer), a catheter, or a medical tube.

Yet another aspect of the present invention is
8. A method for producing the medical material according to any one of 1. to 4. above, comprising: preparing a mixture by mixing the block copolymer, the hydrophobic resin, and an organic solvent so that the content of the hydrophobic resin is greater than the content of the block copolymer; and heat-treating the mixture at a temperature above 100°C and below 150°C for 30 minutes to 3 hours.

One aspect of the present invention relates to a medical material comprising a block copolymer having a structural unit (A) derived from a reactive monomer having an epoxy group and a structural unit (B) derived from a hydrophilic monomer, and at least one hydrophobic resin selected from the group consisting of polyvinyl chloride resin and polyurethane elastomer, wherein the content of the hydrophobic resin in the medical material is greater than the content of the block copolymer, the medical material has a sliding resistance of 50 gf or less, and satisfies at least one of a tensile strength of 8.0 MPa or greater and a tensile elongation of more than 80%. This configuration makes it possible to provide medical materials and medical devices that have mechanical properties (particularly tensile strength and tensile elongation) comparable to those of polytetrafluoroethylene (PTFE) and superior slipperiness (slidability) to PTFE.

In this specification, the structural unit (A) derived from a reactive monomer having an epoxy group is also referred to simply as the "structural unit (A) of the present invention" or "structural unit (A)." In this specification, the structural unit (B) derived from a hydrophilic monomer is also referred to simply as the "structural unit (B) of the present invention" or "structural unit (B)." In this specification, a block copolymer having structural units (A) and (B) is also referred to simply as the "block copolymer of the present invention" or "block copolymer."

In this specification, "at least one hydrophobic resin selected from the group consisting of polyvinyl chloride resin and polyurethane elastomer" is also referred to simply as "the hydrophobic resin of the present invention" or "hydrophobic resin."

In this specification, when a structural unit is defined as being "derived from" a certain monomer, it means that the structural unit is generated by cleavage of one of the polymerizable unsaturated double bonds of the corresponding monomer.

As used herein, the term "(meth)acrylic" includes both acrylic and methacrylic. Thus, for example, the term "(meth)acrylic acid" includes both acrylic acid and methacrylic acid. Similarly, the term "(meth)acryloyl" includes both acryloyl and methacryloyl. Thus, for example, the term "(meth)acryloyl group" includes both acryloyl and methacryloyl groups. Similarly, the term "(meth)acrylate" includes both acrylate and methacrylate. For example, the term "alkoxyalkyl (meth)acrylate" includes both alkoxyalkyl acrylate and alkoxyalkyl methacrylate.

In this specification, the term "X to Y" indicating a range includes both X and Y and means "X or greater and Y or less." Furthermore, "X and/or Y" includes each of X and Y and all combinations of one or more of them, and specifically means at least one of X and Y, and includes X alone, Y alone, and a combination of X and Y.

Unless otherwise specified, operations and measurements of physical properties are performed at room temperature (20-25°C) and relative humidity of 40-50% RH.

In the present invention, a block copolymer is combined with a hydrophobic resin (at least one of polyvinyl chloride resin and polyurethane elastomer). The block copolymer provides slipperiness (slidability, lubricity) to the material. Furthermore, the block copolymer exhibits superior slipperiness (slidability, lubricity) to PTFE. Therefore, the medical material of the present invention and medical devices made using the medical material exhibit slipperiness (slidability) equal to or greater than PTFE. Furthermore, the hydrophobic resin promotes the ring-opening and crosslinking reaction of the epoxy groups present in the block copolymer, improving mechanical properties (e.g., tensile strength and tensile elongation). Therefore, the medical material of the present invention and medical devices made using the medical material (e.g., plastic insertion needles, dilators, sheaths (introducers), catheters, medical tubing) have excellent slipperiness (slidability) and mechanical properties (e.g., tensile strength and tensile elongation), and these properties are well balanced. Furthermore, the medical material of the present invention contains more hydrophobic resin than block copolymer. Therefore, when a medical device is made using the medical material of the present invention, the ring-opening (crosslinking reaction) of the epoxy groups proceeds at a high density. This increases the strength of the medical device (or the film strength in the case of a coating layer). Therefore, the present invention can provide a medical material that has mechanical properties (particularly tensile strength and tensile elongation) comparable to those of polytetrafluoroethylene (PTFE) and superior slip properties (slidability) to PTFE.

Note that the above mechanism is speculative and does not limit the technical scope of the present invention.

The following describes preferred embodiments of the present invention. Note that the present invention is not limited to the following embodiments, and various modifications are possible within the scope of the claims. Furthermore, the embodiments described in this specification can be combined in any manner to create other embodiments.

Throughout this specification, singular expressions should be understood to include the plural concept, unless otherwise specified. Therefore, singular articles (e.g., "a," "an," "the," etc. in English) should be understood to include the plural concept, unless otherwise specified. Furthermore, terms used in this specification should be understood to have the meaning commonly used in the relevant field, unless otherwise specified. Therefore, unless otherwise defined, all technical and scientific terms used in this specification have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. In the case of conflict, the present specification (including definitions) will take precedence.

(Block copolymer)
The block copolymer according to the present invention has a structural unit (A) derived from a reactive monomer having an epoxy group and a structural unit (B) derived from a hydrophilic monomer.

The reactive monomer having an epoxy group that constitutes the block copolymer has an epoxy group as its reactive group. By introducing the structural unit (A) derived from such a reactive monomer into the block copolymer, the epoxy group opens, causing crosslinking (bonding) between the block copolymers and increasing strength. Furthermore, when a layer of medical material is formed on a substrate layer made of a resin material, the ring-opened epoxy group can also cause crosslinking (bonding) between the block copolymer and the substrate layer.

The reactive monomers that make up the block copolymer are not particularly limited as long as they contain an epoxy group, and known compounds can be used. Among these, because it is easier to control the crosslinking or polymerization of the block copolymer, it is preferable for the reactive monomers containing an epoxy group to include at least one selected from the group consisting of glycidyl acrylate, glycidyl methacrylate (GMA), 3,4-epoxycyclohexylmethyl acrylate, 3,4-epoxycyclohexylmethyl methacrylate, β-methylglycidyl methacrylate, and allyl glycidyl ether. Among these, glycidyl (meth)acrylate is more preferred, with glycidyl methacrylate being particularly preferred, considering its ability to further promote the crosslinking reaction and ease of production.

The above reactive monomers may be used alone or in combination of two or more types. In other words, the structural unit (A) (reactive site) derived from the reactive monomer may be a homopolymer type composed of a single type of reactive monomer, or a copolymer type composed of two or more types of the above reactive monomers. When two or more types of reactive monomers are used, the structural unit (A) may be in the form of a block copolymer or a random copolymer.

In other words, in a preferred embodiment of the present invention, the reactive monomer having an epoxy group includes at least one selected from the group consisting of glycidyl acrylate, glycidyl methacrylate, 3,4-epoxycyclohexylmethyl acrylate, 3,4-epoxycyclohexylmethyl methacrylate, β-methylglycidyl methacrylate, and allyl glycidyl ether. In a more preferred embodiment of the present invention, the reactive monomer having an epoxy group is at least one of glycidyl acrylate and glycidyl methacrylate. In a particularly preferred embodiment of the present invention, the reactive monomer having an epoxy group is glycidyl methacrylate.

The hydrophilic monomers that make up the block copolymer swell when in contact with body fluids (e.g., blood, urine) or aqueous solvents, imparting excellent slipperiness (lubricity). Therefore, by introducing structural units (B) derived from such hydrophilic monomers into the block copolymer, medical devices made from this medical material will have excellent slipperiness (lubricity), reducing friction when the medical device comes into contact with the wall of a lumen, such as a blood vessel wall.

The hydrophilic monomers that make up the block copolymer are not particularly limited as long as they have the above properties, and known compounds can be used. Examples include acrylamide and its derivatives, vinylpyrrolidone, acrylic acid, methacrylic acid and their derivatives, polyethylene glycol acrylate and its derivatives, monomers with sugars or phospholipids in the side chains, and water-soluble monomers such as maleic anhydride. More specifically, acrylic acid, methacrylic acid, N-methylacrylamide, N,N-dimethylacrylamide (DMAA), acrylamide, acryloylmorpholine, N,N-dimethylaminoethyl acrylate, N-vinylpyrrolidone, 2-methacryloyloxyethyl phosphorylcholine, 2-methacryloyloxyethyl-D-glycoside, 2-methacryloyloxyethyl-D-mannoside, vinyl methyl ether, 2-hydroxyethyl (meth)acrylate, 4-hydroxybutyl (meth)acrylate, 2-hydroxypropyl (meth)acrylate, 2-hydroxybutyl (meth)acrylate, 6-hydroxyhexyl (meth)acrylate, 1,4-cyclohexanedimethanol mono(meth)acrylate, 1-chloro-2-hydroxypropyl (meth)acrylate, diethylene glycol mono(meth)acrylate, 1,6-hexanediol mono(meth)acrylate, pentaerythritol tri(meth)acrylate, dipentaerythritol penta(meth)acrylate, neopentyl glycol mono(meth)acrylate, trimethylolpropane di(meth)acrylate, trimethylolethane di(meth)acrylate, 2-hydroxy-3-phenyloxypropyl(meth)acrylate, 4-hydroxycyclohexyl(meth)acrylate, 2-hydroxy-3-phenyloxy(meth)acrylate, 4-hydroxycyclohexyl(meth)acrylate, cyclohexanedimethanol mono(meth)acrylate, poly(ethylene glycol) methyl ether acrylate, and poly(ethylene glycol) methyl ether methacrylate. From the viewpoints of imparting excellent slipperiness (lubricity), ease of synthesis, and operability, the hydrophilic monomer preferably includes at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, 2-hydroxyethyl methacrylate, and N-vinylpyrrolidone, and more preferably at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, and 2-hydroxyethyl methacrylate. Of these, from the viewpoint of excellent lubricity, N,N-dimethylacrylamide is particularly preferred as the hydrophilic monomer.

The above hydrophilic monomers may be used alone or in combination of two or more types. In other words, the structural unit (B) (hydrophilic moiety) derived from the hydrophilic monomer may be a homopolymer type composed of a single hydrophilic monomer, or a copolymer type composed of two or more of the above hydrophilic monomers. When two or more hydrophilic monomers are used, the structural unit (B) may be in the form of a block copolymer or a random copolymer.

In other words, in a preferred embodiment of the present invention, the hydrophilic monomer includes at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, 2-hydroxyethyl methacrylate, and N-vinylpyrrolidone. In a more preferred embodiment of the present invention, the hydrophilic monomer is at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, and 2-hydroxyethyl methacrylate. In a particularly preferred embodiment of the present invention, the hydrophilic monomer is N,N-dimethylacrylamide.

The block copolymer has the above-mentioned structural unit (A) and structural unit (B). The ratio of the structural unit (A) to the structural unit (B) is not particularly limited as long as the above-mentioned effects are achieved. Considering further improvements in lubricity (slipperiness, slidability), the ratio of the structural unit (A) to the structural unit (B) (molar ratio of structural unit (A):structural unit (B)) is preferably 1:2 to 1:100, more preferably 1:2 to 1:50, even more preferably 1:5 to 1:50, and particularly preferably 1:10 to 1:30. The molar ratio of the structural unit (A):structural unit (B) can be controlled by adjusting the charge ratio (molar ratio) of each monomer during the production stage of the block copolymer. Therefore, the charging ratio (molar ratio) of the reactive monomer to the hydrophilic monomer in the production stage of the block copolymer is preferably 1:2 to 1:100, more preferably 1:2 to 1:50, even more preferably 1:5 to 1:50, and particularly preferably 1:10 to 1:30. The composition (molar ratio) of the structural unit (A):structural unit (B) can be confirmed, for example, by subjecting the copolymer to NMR measurement ( ¹ H-NMR measurement, ¹³ C-NMR measurement, etc.).

The composition of each structural unit (structural units (A) and (B) and other structural units) can be measured by known methods. For example, the composition (molar ratio) of each structural unit can be measured by measuring the integral ratio of the intensity of each signal in ^{the 1H} -NMR spectrum of the block copolymer solution.

In one embodiment of the present invention, the block copolymer of the present invention is essentially composed of, or consists of, structural unit (A) derived from at least one reactive monomer selected from the group consisting of glycidyl acrylate, glycidyl methacrylate, 3,4-epoxycyclohexylmethyl acrylate, 3,4-epoxycyclohexylmethyl methacrylate, β-methylglycidyl methacrylate, and allyl glycidyl ether, and structural unit (B) derived from at least one hydrophilic monomer selected from the group consisting of N,N-dimethylacrylamide, acrylamide, 2-hydroxyethyl methacrylate, and N-vinylpyrrolidone.

In one embodiment of the present invention, the block copolymer according to the present invention is essentially composed of, or consists of, a structural unit (A) derived from at least one reactive monomer of glycidyl acrylate and glycidyl methacrylate, and a structural unit (B) derived from at least one hydrophilic monomer selected from the group consisting of N,N-dimethylacrylamide, acrylamide, and 2-hydroxyethyl methacrylate.

In one embodiment of the present invention, the block copolymer according to the present invention is essentially composed of, or consists only of, a structural unit (A) derived from glycidyl methacrylate (a reactive monomer having an epoxy group) and a structural unit (B) derived from N,N-dimethylacrylamide (a hydrophilic monomer).

The weight-average molecular weight of the block copolymer is preferably 10,000 to 10,000,000 from the viewpoint of solubility. The weight-average molecular weight of the block copolymer is more preferably 100,000 to 5,000,000 from the viewpoint of ease of preparation of the coating liquid. In this specification, "weight-average molecular weight" refers to a value measured by gel permeation chromatography (GPC) using polystyrene as the standard substance.

(hydrophobic resin)
The hydrophobic resin of the present invention is at least one of polyvinyl chloride and polyurethane elastomer. The hydrophobic resin induces ring-opening of epoxy groups present in the block copolymer. Ring-opening of the epoxy groups promotes crosslinking (bonding) between the block copolymers. Furthermore, when a coating layer is formed on a substrate layer made of a resin material using the medical material of the present invention, the ring-opened epoxy groups can also crosslink (bond) the block copolymer to the substrate layer. Therefore, in medical devices obtained using the medical material of the present invention, the coating layer is firmly bonded to the substrate layer. This improves the mechanical properties (e.g., tensile strength and tensile elongation) of medical devices made using the medical material of the present invention. Furthermore, the medical device can maintain its shape well even after sliding. Additionally, hydrophobic resins (polyvinyl chloride, polyurethane elastomers) are already used as medical materials, making them suitable from a safety perspective.

The hydrophobic resin according to the present invention is at least one of polyvinyl chloride and polyurethane elastomer. Of these, from the viewpoint of mechanical properties, particularly tensile strength, it is preferable that the hydrophobic resin contains at least polyvinyl chloride resin, and polyvinyl chloride resin is more preferable. Furthermore, from the viewpoint of mechanical properties, particularly tensile elongation, it is preferable that the hydrophobic resin contains at least polyurethane elastomer, and it is more preferable that it consists solely of polyurethane elastomer.

The weight-average molecular weight (Mw) of the hydrophobic resin is 1,000 or more. From the standpoint of solubility, the weight-average molecular weight of the hydrophobic resin is preferably 500,000 or less. Furthermore, from the standpoint of promoting crosslinking and stability, the weight-average molecular weight of the hydrophobic resin is preferably 10,000 or more. As an example, the weight-average molecular weight (Mw) of the hydrophobic resin is 1,000 to 10,000,000, and preferably 10,000 to 500,000.

In the medical material, the mixing ratio (mass ratio) of block copolymer to hydrophobic resin is such that the content of hydrophobic resin is greater than the content of block copolymer. This allows the production of medical devices (e.g., plastic insertion needles, dilators, sheaths (introducers), catheters, medical tubing) with excellent slipperiness (sliding property, lubricity). While it is possible to apply a separate coating with lubricity, medical devices (which partially include a coating layer made of medical material) formed using the medical material of the present invention have slipperiness (sliding property, lubricity), so there is no need to apply a separate coating with lubricity. From the viewpoints of smoothness (slidability, lubricity), mechanical properties, and a good balance between these, the hydrophobic resin is preferably contained in a ratio of 125 to 300 parts by mass per 100 parts by mass of the block copolymer, more preferably 150 to 230 parts by mass per 100 parts by mass of the block copolymer, and particularly preferably 160 to 180 parts by mass per 100 parts by mass of the block copolymer. When the mixing ratio (mass ratio) of the block copolymer to the hydrophobic resin is within the above range, the medical device (or coating layer) formed using the medical material will have sufficient mechanical properties and lubricity (especially excellent lubricity) and an excellent balance between these.

(Characteristics of medical materials)
The medical material according to the present invention (and therefore the medical device (or coating layer) formed using the medical material) has mechanical properties (particularly tensile strength and tensile elongation) comparable to those of polytetrafluoroethylene (PTFE), while also having excellent slip properties (slidability) comparable to or better than PTFE.

Specifically, the medical material (and therefore the medical device (or coating layer) formed using the medical material) has a sliding resistance of 50 gf or less. The sliding resistance is preferably less than 30 gf, and more preferably less than 20 gf. Since a lower sliding resistance is preferable, there is no particular lower limit and the lower limit is 0 gf, but a value of 3 gf or more is acceptable. Therefore, the sliding resistance of the medical material is, for example, 0 gf to 50 gf, preferably 0 gf to less than 30 gf, and more preferably 0 gf to less than 20 gf. The sliding resistance of the medical material may also be 3 gf to less than 30 gf, or 3 gf to less than 20 gf. In this specification, "sliding resistance" is a value measured according to the method described in the examples.

The medical material (and therefore the medical device (or coating layer) formed using said medical material) satisfies at least one of a tensile strength of 8.0 MPa or more and a tensile elongation of greater than 80%. Preferably, the medical material satisfies both a tensile strength of 8.0 MPa or more and a tensile elongation of greater than 80%. Here, the tensile strength of the medical material is preferably 8.4 MPa or more, more preferably 10 MPa or more, even more preferably 20 MPa or more, and particularly preferably 25 MPa or more. Since a higher tensile strength of the medical material is preferable, there is no particular upper limit, but it is usually 200 MPa or less. Therefore, the tensile strength of the medical material is, for example, 8.0 MPa to 200 MPa, preferably 8.4 MPa to 200 MPa, more preferably 10 MPa to 200 MPa, even more preferably 20 MPa to 200 MPa, and particularly preferably 25 MPa to 200 MPa. In this specification, "tensile strength" is a value measured according to the method described in the examples. The tensile elongation of the medical material is preferably 100% or more, more preferably 120% or more, even more preferably 150% or more, and particularly preferably 400% or more. Since a higher tensile elongation of the medical material is preferable, there is no particular upper limit, but it is usually 500% or less. Therefore, the tensile elongation of the medical material is, for example, more than 80% and 500% or less, preferably 100% or more and 500% or less, more preferably 120% or more and 500% or less, even more preferably 150% or more and 500% or less, and particularly preferably 400% or more and 500% or less. In this specification, "tensile elongation" refers to a value measured according to the method described in the Examples.

(Method for manufacturing medical materials)
The medical material according to the present invention has a sliding resistance of 50 gf or less and at least one of a tensile strength of 8.0 MPa or more and a tensile elongation of more than 80%. A medical material satisfying these properties may be produced by any method, but can be produced particularly by appropriately controlling the heating conditions and the mixing ratio of the block copolymer and the hydrophobic resin.

In other words, the present invention provides a method for producing a medical material according to the present invention, which comprises: preparing a mixture by mixing the block copolymer, the hydrophobic resin, and the organic solvent so that the content of the hydrophobic resin is greater than the content of the block copolymer (mixture preparation step); and heat-treating the mixture at a temperature greater than 100°C and less than 150°C for 30 minutes to 3 hours (mixture heat-treatment step).

A preferred embodiment of the above method is described below. Note that the present invention is not limited to the following embodiment.

(Mixture preparation process)
In this step, a mixture is prepared by mixing a block copolymer, a hydrophobic resin, and an organic solvent. The mixing ratio of the block copolymer to the hydrophobic resin is such that the content of the hydrophobic resin is greater than the content of the block copolymer. Preferably, the hydrophobic resin is mixed with the block copolymer at a mixing ratio of 125 to 300 parts by mass per 100 parts by mass of the block copolymer. It is more preferable to mix the hydrophobic resin with the block copolymer at a mixing ratio of 150 to 230 parts by mass per 100 parts by mass of the block copolymer. It is particularly preferable to mix the hydrophobic resin with the block copolymer at a mixing ratio of 160 to 180 parts by mass per 100 parts by mass of the block copolymer. By setting the mixing ratio of the block copolymer to the hydrophobic resin within the above range, the sliding resistance, tensile strength, and tensile elongation of the medical material (and therefore the medical device (or coating layer) formed using the medical material) can be more appropriately controlled.

The organic solvents that can be used to prepare the mixture are not particularly limited as long as they can dissolve the block copolymer and hydrophobic resin (and other components, if used). They are appropriately selected depending on the type of block copolymer and hydrophobic resin (and other components, if used). From the viewpoint of high solubility, alcoholic solvents such as methanol, ethanol, isopropyl alcohol, and butanol; and organic solvents such as dichloromethane, chloroform, carbon tetrachloride, tetrahydrofuran (THF), dimethyl sulfoxide, N,N-dimethylformamide (DMF), dioxane, and benzene are preferred. These solvents may be used alone or in combination (as a mixed solvent) of two or more. The concentration of the block copolymer in the mixture is 0.1 to 20% by mass, preferably 0.5 to 15% by mass, and more preferably 1 to 10% by mass. The concentration of the hydrophobic resin in the mixture is preferably such that the mixing ratio with the block copolymer falls within the range described above. When the concentrations of the block copolymer and hydrophobic resin are within the above ranges, the sliding resistance, tensile strength, and tensile elongation of the medical material (and therefore the medical device (or coating layer) formed using the medical material) can be more appropriately controlled.

The order in which the block copolymer and hydrophobic resin are mixed is not particularly limited. For example, (1) the block copolymer and hydrophobic resin can be charged all at once to the organic solvent, (2) the block copolymer can be added to the organic solvent and then the hydrophobic resin can be added, or (3) the hydrophobic resin can be added to the organic solvent and then the block copolymer can be added. If necessary, the above additions can be carried out with stirring. Alternatively, the mixture can be stirred after the above additions.

(Heat treatment step of the mixture)
In this step, the mixture obtained in the above (mixture preparation step) is heat-treated at a temperature above 100°C and below 150°C for 30 minutes to 3 hours. This heat treatment allows for more appropriate control of the sliding resistance, tensile strength, and tensile elongation of the medical material, particularly the sliding resistance. Here, if the heat treatment temperature is below 100°C or the heat treatment time is less than 30 minutes, the heat treatment is insufficient, and the desired durability in terms of slipperiness is not achieved. If the heat treatment temperature exceeds 150°C or the heat treatment time exceeds 3 hours, the heat treatment proceeds excessively, and the desired durability in terms of slipperiness is also not achieved. Considering the better lubricity and mechanical properties of the medical material, as well as a better balance between these, the heat treatment temperature is preferably 105°C to 140°C, more preferably above 105°C to 120°C. Considering the better lubricity and mechanical properties of the medical material, as well as a better balance between these, the heat treatment time is preferably 40 minutes to 2 hours, more preferably 50 minutes to 1.5 hours. The heat treatment step may be performed once or may be repeated two or more times. In the latter case, it is preferable that the heat treatment temperature in the entire heat treatment step (all repeated heat treatment steps) is higher than 100° C. and lower than 150° C., and that the heat treatment time in the entire heat treatment step is 30 minutes to 3 hours, both of which are within the above-mentioned ranges.

(Medical devices)
As described above, the medical material according to the present invention has mechanical properties (particularly tensile strength and tensile elongation) comparable to those of polytetrafluoroethylene (PTFE) and has slip properties (slidability) that are equal to or superior to those of PTFE.

The present invention therefore also provides a medical device comprising or consisting of the medical material of the present invention. The sliding resistance of the medical device is 50 gf or less (preferably less than 30 gf, more preferably less than 20 gf), and the medical device satisfies at least one of the following, and preferably both: a tensile strength of 8.0 MPa or more (preferably 8.4 MPa or more, more preferably 10 MPa or more, even more preferably 20 MPa or more, and particularly preferably 25 MPa or more) and a tensile elongation of more than 80% (preferably 100% or more, more preferably 120% or more, even more preferably 150% or more, and particularly preferably 400% or more). The size of the medical device can be appropriately selected depending on the desired application (e.g., a catheter).

Alternatively, the present invention also provides a medical device comprising a substrate layer and a coating layer containing or consisting of the medical material of the present invention. The sliding resistance of the coating layer is 50 gf or less (preferably less than 30 gf, more preferably less than 20 gf), and the coating layer satisfies at least one of the following: a tensile strength of 8.0 MPa or more (preferably 8.4 MPa or more, more preferably 10 MPa or more, even more preferably 20 MPa or more, particularly preferably 25 MPa or more) and a tensile elongation of more than 80% (preferably 100% or more, more preferably 120% or more, even more preferably 150% or more, particularly preferably 400% or more), and preferably satisfies both.

When a medical device contains the medical material of the present invention, the medical device may be in a form consisting of the medical material and the other components described above, a form consisting of the medical material and a coiled or braided metal (for example, a structure in which a metal coil or metal braid is embedded in a tubular medical device made using the medical material), or a form coated on a metal wire or metal surface.

When a medical device comprises a substrate layer and a coating layer containing or consisting of the medical material of the present invention, the substrate layer may be made of any material, including, for example, metal materials, polymer materials (resin materials), and ceramics.

The metallic material constituting the substrate layer is not particularly limited, and metallic materials commonly used for medical devices such as catheters, guidewires, and indwelling needles can be used. Specific examples include various stainless steels such as SUS304, SUS314, SUS316, SUS316L, SUS420J2, and SUS630, as well as gold, platinum, silver, copper, nickel, cobalt, titanium, iron, aluminum, tin, and various alloys such as nickel-titanium alloys, nickel-cobalt alloys, cobalt-chromium alloys, and zinc-tungsten alloys. These may be used alone or in combination of two or more. The metallic material best suited for the substrate layer of the intended use, such as a catheter, guidewire, or indwelling needle, can be appropriately selected.

Furthermore, among the materials constituting the substrate layer, the polymer material (resin material or elastomer material) is not particularly limited, and polymer materials commonly used in medical devices such as plastic insertion needles (indwelling needles), dilators, sheaths (introducers), catheters, or medical tubing can be used. Specific examples include polyamide resins, polyolefin resins such as polyethylene resins and polypropylene resins, modified polyolefin resins, cyclic polyolefin resins, epoxy resins, polyurethane resins, diallyl phthalate resins (allyl resins), polycarbonate resins, fluororesins (e.g., polytetrafluoroethylene resins), amino resins (urea resins, melamine resins, benzoguanamine resins), polyester resins such as polyethylene terephthalate resins and polybutylene terephthalate resins, styrene resins, acrylic resins, polyacetal resins, vinyl acetate resins, phenolic resins, vinyl chloride resins, silicone resins (silicon resins), polyether resins, and polyimide resins.

In addition, thermoplastic elastomers such as polyurethane elastomers, polyester elastomers, and polyamide elastomers (nylon elastomers) can also be used as materials for the base layer.

These polymer materials may be used alone, as a mixture of two or more types, or as a copolymer of two or more monomers constituting any of the above resins or elastomers. Among these, polyethylene resins, polyurethane resins, polyethylene terephthalate resins, polyamide resins, and polyamide elastomers are preferred, with polyamide resins and polyamide elastomers being more preferred. The carboxyl and amino groups contained as terminal groups in polyamide resins and polyamide elastomers can undergo crosslinking reactions with epoxy groups in block copolymers. Furthermore, these polymer materials (particularly polyamide resins and polyamide elastomers) are relatively soft and can be easily impregnated with block copolymers and hydrophobic resins. This enhances the bond between the polymer material (particularly polyamide resins and polyamide elastomers) and the block copolymer, allowing for the formation of a more durable coating layer. The polymer material can be appropriately selected to be optimal for the substrate layer of the intended use, such as a plastic insertion needle (indwelling needle), dilator, sheath (introducer), catheter, or medical tubing.

Furthermore, the shape of the substrate layer is not particularly limited and can be selected appropriately depending on the intended use, such as sheet, wire, rod, or tube.

The entire substrate layer may be made of one of the materials listed above. The substrate layer may be a multilayer structure formed by laminating different materials in multiple layers, or a structure in which components made of different materials are joined together for each portion of the medical device. Alternatively, the substrate may have a structure in which the surface of a substrate layer core made of one of the materials listed above is coated with one of the other materials listed above by an appropriate method to form a substrate surface layer. Examples of the latter include a substrate surface layer formed by coating the surface of a substrate layer core made of a resin material or the like with a metal material by an appropriate method (conventionally known methods such as plating, metal vapor deposition, sputtering, etc.); a substrate surface layer formed by coating the surface of a substrate layer core made of a hard reinforcing material such as a metal or ceramic material with a polymer material that is softer than the metal reinforcing material by an appropriate method (conventionally known methods such as dipping, spraying, coating, printing, etc.); or a substrate surface layer formed by combining the reinforcing material that forms the substrate layer core with a polymer material. The core substrate layer may also be a multilayer structure formed by laminating different materials in multiple layers, or a structure in which components formed from different materials are joined together for each portion of the medical device. A separate middle layer may also be formed between the core substrate layer and the substrate surface layer. Furthermore, the substrate surface layer may also be a multilayer structure formed by laminating different materials in multiple layers, or a structure in which components formed from different materials are joined together for each portion of the medical device.

Another layer may be provided between the substrate layer and the coating layer. In this configuration, the other layer may be made of a material similar to the polymer material (resin material or elastomer material) described above.

Alternatively, the coating layer may be made of a medical material and the other components mentioned above, or may be made of a medical material and a coiled or braided metal (for example, a structure in which a metal coil or metal braid is embedded in a coating layer made using a medical material).

(Method for manufacturing medical devices)
When a medical device includes or is composed of the medical material of the present invention, a method for manufacturing the medical device includes molding a mixture containing a block copolymer and a hydrophobic resin, and, if necessary, the other components described above. The mixture can be prepared by mixing the block copolymer and the hydrophobic resin, adding the block copolymer and the hydrophobic resin all at once to a solvent, adding the block copolymer and the hydrophobic resin to a solvent in that order, or adding the hydrophobic resin and the block copolymer to a solvent in that order. When a solvent is used to prepare the mixture, the solvent can be appropriately selected depending on the type of block copolymer and hydrophobic resin used. Specific examples include, but are not limited to, N,N'-dimethylformamide (DMF), chloroform, acetone, tetrahydrofuran (THF), dioxane, benzene, and methanol. These solvents may be used alone or in combination. The concentration of the block copolymer in the mixture is 0.1 to 20% by mass, preferably 0.5 to 15% by mass, and more preferably 1 to 10% by mass. The concentration of the hydrophobic resin in the mixture is preferably such that the mixing ratio with the block copolymer falls within the above-mentioned range.

Known molding methods can be used in the same manner or with appropriate modifications. Specific examples include dipping, which involves applying a medical material to a substrate (e.g., a wire) by immersion and then removing the substrate; melt extrusion molding; paste extrusion molding; and spray coating. Molding conditions are also not particularly limited and can be appropriately selected depending on the type and amount of medical material used and the type and size of the medical device. The molding temperature is, for example, greater than 100°C and less than 150°C, preferably greater than 105°C and less than 140°C, and more preferably greater than 105°C and less than 120°C. The molding time is, for example, 30 minutes to 3 hours, preferably 40 minutes to 2 hours, and more preferably 50 minutes to 1.5 hours. Under these conditions, the resulting medical device exhibits better lubricity and mechanical properties (especially lubricity), while achieving a better balance between these properties. The molding operation may be performed once or repeatedly two or more times. In the latter case, it is preferable that the molding temperature during the entire molding operation (all repeated molding operations) is greater than 100°C and less than 150°C, and that the molding time during the entire molding operation is 30 minutes or more and 3 hours or less, each falling within the above range.

When a medical device has a layer (coating layer, covering layer) containing the medical material of the present invention formed on a substrate layer, a manufacturing method for the medical device includes, for example, preparing a coating liquid containing a block copolymer, a hydrophobic resin, and a solvent (preparation step); applying the coating liquid to the substrate layer to form a coating film on the substrate layer (coating step); and heat-treating the coating film at a temperature above 100°C and below 150°C for 30 minutes to 3 hours (heat-treatment step). If necessary, a drying step (drying step) may be performed after the coating step and before the heat-treatment step. Furthermore, a washing step (washing step) may be performed after the heat-treatment step. Furthermore, with the medical material of the present invention, the hydrophobic resin is stably retained in the coating layer (covering layer). Furthermore, the epoxy groups of the block copolymer are ring-opened without the need for the addition of an acid or base. Therefore, with the medical material of the present invention, a separate washing step is not required, which is advantageous for mass production.

Below, a preferred embodiment of the method for manufacturing the above medical device will be described. Note that the present invention is not limited to the following embodiment.

(preparation process)
In this step, a coating liquid containing a block copolymer, a hydrophobic resin, and a solvent is prepared. In this step, a coating liquid containing a block copolymer, a hydrophobic resin, and a solvent may be purchased and used. Alternatively, the coating liquid may be prepared by mixing the block copolymer, the hydrophobic resin, and the solvent.

Hydrophobic resins are stable in the coating solution, making them preferable in terms of safety and ease of operation. Furthermore, if the coating solution is kept at room temperature, the ring-opening of the epoxy groups (crosslinking reaction) will not proceed. This makes them easy to work with.

The concentration of the block copolymer in the coating solution is 0.1 to 20% by mass, preferably 0.5 to 15% by mass, and more preferably 1 to 10% by mass. The concentration of the hydrophobic resin in the mixture is preferably such that the mixing ratio with the block copolymer falls within the range shown below. If the concentrations of the block copolymer and hydrophobic resin are within the above ranges, the sliding resistance, tensile strength, and tensile elongation of the medical material can be more appropriately controlled.

The mixing ratio of the block copolymer and hydrophobic resin when preparing the coating solution is such that the content of the hydrophobic resin is greater than the content of the block copolymer. Preferably, the hydrophobic resin is mixed with the block copolymer at a ratio of 125 to 300 parts by weight per 100 parts by weight of the block copolymer. It is more preferable to mix the hydrophobic resin with the block copolymer at a ratio of 150 to 230 parts by weight per 100 parts by weight of the block copolymer. It is particularly preferable to mix the hydrophobic resin with the block copolymer at a ratio of 160 to 180 parts by weight per 100 parts by weight of the block copolymer. By maintaining the mixing ratio of the block copolymer and hydrophobic resin within the above range, the sliding resistance, tensile strength, and tensile elongation of the medical material can be more appropriately controlled. Furthermore, a uniform coating layer of the desired thickness can be easily obtained with a single coating, and the viscosity of the solution remains within an appropriate range, which is advantageous in terms of operability (e.g., ease of coating) and production efficiency.

The preparation steps other than those mentioned above are the same as those in the (Mixture Preparation Step) in the (Medical Material Manufacturing Method) above.

(Coating process)
In this step, the coating liquid is applied onto a substrate layer to form a coating film on the substrate layer. Here, the substrate layer is the same as that described above (medical device).

The method for applying (coating) the coating liquid to the surface of the substrate layer is not particularly limited, and conventional methods can be applied, such as application/printing, immersion (dipping, dip coating), spraying, spin coating, mixed solution-impregnated sponge coating, bar coating, die coating, reverse coating, comma coating, gravure coating, and doctor knife. Of these, immersion (dipping, dip coating) methods are preferred.

Furthermore, when forming a coating film (coating layer, covering layer) only on a portion of the substrate layer, the coating film (coating layer, covering layer) can be formed on the desired surface area of the substrate layer by immersing only a portion of the substrate layer in the coating liquid and coating the coating liquid onto that portion of the substrate layer.

If it is difficult to immerse only a portion of the substrate layer in the coating liquid, the surface portions of the substrate layer that do not require the formation of a coating film (coating layer, covering layer) can be protected (coated, etc.) with a suitable removable member or material. The substrate layer is then immersed in the coating liquid to coat the substrate with the coating liquid. After that, the protective member (material) covering the surface portions of the substrate layer that do not require the formation of a coating film (coating layer, covering layer) can be removed, and the coating can be reacted by heating or other means to form a coating film (coating layer, covering layer) on the desired surface portion of the substrate layer. However, the present invention is not limited to these formation methods, and conventionally known methods can be used to form a coating film (coating layer, covering layer). For example, if it is difficult to immerse only a portion of the substrate layer in the coating liquid, other coating methods (e.g., applying the coating liquid to a desired surface portion of a medical device using an application device such as a sprayer, bar coater, die coater, reverse coater, comma coater, gravure coater, spray coater, or doctor knife) can be used instead of the immersion method. Furthermore, when the structure of a cylindrical medical device requires that both the outer and inner surfaces of the device have a coating film (coating layer, covering layer), the immersion method is preferably used, as it allows both the outer and inner surfaces to be coated at the same time.

The amount of coating liquid to be applied is preferably such that the thickness (dry film thickness) of the resulting coating layer (coating layer) is 0.1 to 300 μm, more preferably 0.5 to 200 μm, and even more preferably 1 to 100 μm.

(drying process)
In this step, if necessary, the coating film is dried to remove at least a portion of the solvent.

There are no particular restrictions on the drying conditions as long as they allow the solvent to be removed, and they can be selected appropriately depending on the type of solvent. The drying temperature is, for example, 10°C to 50°C, preferably 10°C to 30°C, and more preferably 20°C to 25°C. The drying time is, for example, 10 minutes to 5 hours, preferably 20 minutes to 3 hours, and more preferably 30 minutes to 1.5 hours. There are also no particular restrictions on the pressure conditions during drying, and drying can be carried out under normal pressure (atmospheric pressure).

(Heat treatment process)
In this step, the coating film formed in the above (applying step) or the coating film dried in the above (drying step) is heat-treated at a temperature above 100°C but below 150°C for 30 minutes to 3 hours. This heat treatment allows for more appropriate control of the sliding resistance, tensile strength, and tensile elongation of the medical device, particularly the sliding resistance. The heat treatment temperature is preferably 105°C or higher and 140°C or lower, more preferably above 105°C but below 120°C. The heat treatment time is preferably 40 minutes to 2 hours, more preferably 50 minutes to 1.5 hours. Under these heat treatment conditions, the medical device can exhibit better lubricity and mechanical properties, and these properties can be well balanced. In addition, crosslinking or polymerization in the block copolymer is effectively promoted, forming a strong layer (coat layer, covering layer). Therefore, high lubricity (surface lubricity) can be maintained for a longer period of time. Furthermore, by setting the heat treatment temperature and time below the upper limit values, excessive crosslinking or polymerization can be suppressed. This prevents the layer (coating layer, covering layer) from becoming too hard, thereby maintaining good lubricity (surface lubricity). Another advantage is that even polymer materials that are easily deformed or plasticized by heat can be used as the substrate layer. Therefore, the present invention broadens the range of materials available, enabling the manufacture of medical devices for a variety of applications. Furthermore, since it is possible to form a layer (coating layer, covering layer) with excellent durability at low temperatures, this is also preferable from the perspective of energy costs when manufacturing medical devices. The heat treatment step may be performed once or repeatedly performed two or more times. In the latter case, it is preferable that the heat treatment temperature during the entire heat treatment step (all repeated heat treatment steps) is greater than 100°C and less than 150°C, and that the heat treatment time during the entire heat treatment step is 30 minutes or more and 3 hours or less, each within the above range.

From the perspective of particularly effectively (efficiently) promoting the crosslinking or polymerization of the block copolymer by the hydrophobic resin, the heat treatment may be carried out after the drying treatment. In this way, by undergoing the drying and heat treatment, the solvent is distilled away and further heat treatment is carried out in a state where the block copolymer and hydrophobic resin are easily brought into contact with each other, thereby further improving the effect of promoting the crosslinking or polymerization of the block copolymer by the hydrophobic resin. Furthermore, because the heat treatment can be shortened, even polymer materials that are easily deformed or plasticized by heat can be used as the base layer.

The conditions (temperature, time, etc.) for the drying and heating treatments when these are performed are not particularly limited, but from the perspective of efficiently manufacturing medical devices, it is preferable to perform a drying treatment at a temperature of 10°C to 50°C, maintaining it for 10 minutes to 5 hours, followed by a heating treatment at a temperature of 105°C to 140°C, maintaining it for 40 minutes to 2 hours. From the same perspective, it is even more preferable to perform a drying treatment at a temperature of 10°C to 30°C, maintaining it for 20 minutes to 3 hours, followed by a heating treatment at a temperature above 105°C to 120°C, maintaining it for 50 minutes to 1.5 hours. After the above heating treatments, a further drying treatment may be performed.

Under the above conditions (temperature, time, etc.), a strong coating layer (covering layer) can be formed on the surface of the substrate layer. Furthermore, depending on the type of substrate layer, a crosslinking reaction can occur via the epoxy groups in the block copolymer in the layer (coating layer, covering layer), forming a high-strength coating layer (covering layer) that does not easily peel off from the substrate layer. Therefore, the drying/heating process described above can effectively suppress or prevent peeling of the coating layer (covering layer) from the substrate layer. Furthermore, there are no particular restrictions on the pressure conditions during the heating process, and the process can be carried out under normal pressure (atmospheric pressure).

As a heating means (device), for example, an oven can be used.

(Use of medical devices)
Medical devices are preferably used in devices that come into contact with body fluids, blood, etc., and have a surface that is lubricious in body fluids, physiological saline, and other aqueous liquids, enabling improved operability and reduced damage to tissues and mucous membranes. Specific examples include plastic insertion needles, dilators, sheaths (introducers), catheters, medical tubing, etc. used in blood vessels, but other examples include the following medical devices. That is, in one embodiment of the present invention, the medical device is a plastic insertion needle, dilator, sheath (introducer), catheter, or medical tubing.

(a) Catheters that are inserted or left in the digestive tract via the mouth or nose, such as gastric catheters, nutritional catheters, and enteral feeding tubes;
(b) Catheters that are inserted or placed in the airway or trachea via the mouth or nose, such as oxygen catheters, oxygen cannulas, endotracheal tubes and cuffs, tracheostomy tubes and cuffs, and endotracheal suction catheters;
(c) Catheters that are inserted or placed in the urethra or ureter, such as urethral catheters, urinary catheters, and urethral balloon catheters;
(d) Catheters inserted or left in various body cavities, organs, or tissues, such as suction catheters, drainage catheters, and rectal catheters;
(e) Catheters that are inserted or placed in blood vessels, such as indwelling needles (e.g., plastic indwelling needles), IVH catheters, thermodilution catheters, angiography catheters, vasodilator catheters, and dilators or introducers, or guide wires and stylets for these catheters;
(f) Artificial trachea, artificial bronchi, etc.

The effects of the present invention will be explained using the following examples and comparative examples. However, the technical scope of the present invention should not be interpreted as being limited to the following examples, and examples obtained by appropriately combining the technical means disclosed in each example are also included within the scope of the present invention. In the following examples, unless otherwise specified, operations were performed at room temperature (25°C). Furthermore, unless otherwise specified, "%" and "parts" mean "% by mass" and "parts by mass," respectively.

Synthesis Example 1
The following reaction was carried out to produce a block copolymer (1).

29.7 g of triethylene glycol was added dropwise to 72.3 g of adipic acid dichloride at 50°C, and then the hydrochloric acid was removed under reduced pressure at 50°C for 3 hours to obtain an oligoester. Next, 4.5 g of methyl ethyl ketone was added to 22.5 g of the obtained oligoester, and this was added dropwise to a solution consisting of 5 g of sodium hydroxide, 6.93 g of 31% hydrogen peroxide, 0.44 g of dioctyl phosphate as a surfactant, and 120 g of water, and the mixture was allowed to react at -5°C for 20 minutes. The resulting product was repeatedly washed with water and methanol, and then dried to obtain a polyperoxide (PPO) with multiple peroxide groups within the molecule.

Next, 0.5 g of this PPO, 9.5 g of glycidyl methacrylate (GMA), and 30 g of benzene were used as a solvent and polymerized at 80°C for 2 hours while stirring under reduced pressure. The reaction product obtained after polymerization was reprecipitated with diethyl ether to obtain polyglycidyl methacrylate (PPO-GMA) with multiple peroxide groups in the molecule.

Subsequently, 1.35 g of the obtained PPO-GMA (corresponding to 9.5 mmol of GMA) was dissolved in chlorobenzene together with 11.2 g (113 mmol) of N,N-dimethylacrylamide (DMAA) to give concentrations of 1.35 mass % (PPO-GMA concentration) and 11.2 mass % (DMAA concentration), respectively, and the solution was heated to 80°C under a nitrogen atmosphere for 7 hours to polymerize. The reaction product was reprecipitated with cyclohexane and recovered to obtain block copolymer (1) (structural unit (A):structural unit (B) = GMA:DMAA = 1:12 (molar ratio)) containing epoxy groups in the molecule and exhibiting lubricity when wet. Analysis of the block copolymer (1) thus obtained by ^1H -NMR and ATR-IR confirmed the presence of epoxy groups in the molecule.

Example 1
Polyvinyl chloride resin (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd., trade name: Polyvinyl chloride, average degree of polymerization (n) = 1,050) (PVC) was dissolved in N,N-dimethylformamide (DMF) so that the final concentration in the coating solution was 9.0 mass% (solution (1)). The block copolymer (1) synthesized in Synthesis Example 1 above was added to and dissolved in the solution (1) so that the final concentration in the coating solution was 5.0 mass% to prepare coating solution (1).

A copper wire (diameter: 1.775 mm) was dip-coated in the coating solution (1) prepared above at a speed of 10 mm/sec, and then heated at 110°C for 1 hour to carry out a cross-linking reaction. After heating, the temperature was returned to room temperature (25°C), and the copper wire was removed to obtain a tube (1) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, thickness: 0.085 mm, cross-sectional area: 0.512 ^mm2 ).

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (1) obtained above were measured according to the methods below. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (1) were 11.6 gf, 26.5 MPa, and 120%, respectively.

[Evaluation of sliding resistance]
The tube was cut to a length of 200 mm to prepare test specimens (inner diameter: 1.775 mm, outer diameter: 1.950 mm). Each test specimen was immersed in tap water and set in a pinch tester (OAKRIVER TECHNOLOGY, DL1000), and slid 100 times at a grip force of 500 gf, a test speed of 8.3 mm/s, and a test stroke of 25 mm (grip pad material: silicone, grip pad height: 12.35 mm). The sliding resistance (gf) after 100 slides was measured to evaluate the sliding properties. The lower the sliding resistance, the better the sliding properties were judged to be.

[Evaluation of tensile strength and tensile elongation]
Tensile strength and tensile elongation were measured in accordance with ASTM D412 using a tensile tester (Shimadzu Corporation, Autograph AGX-X). Specifically, each tube was cut to a length of 20 mm to prepare a test specimen. Next, the chuck distance of the tensile tester was set to 10 mm, and the test specimen was clamped in the tensile tester. Measurements were performed at room temperature (25°C), with a chuck pressure of 0.5 MPa and a tensile speed of 50 mm/min. The tensile strength (MPa) was determined by dividing the strength at break of each test specimen by the cross-sectional area of the sample before the test. The elongation at break of each test specimen was defined as the tensile elongation (%). Measurements were performed on five test specimens, and the average value was calculated.

Example 2
A polyurethane elastomer (manufactured by Lubrizol, trade name: Pellethane 2363-80AE) (TPU) was dissolved in N,N-dimethylformamide (DMF) so that the final concentration in the coating solution was 8.0% by mass (solution (2)). The block copolymer (1) synthesized in Synthesis Example 1 above was added to and dissolved in solution (2) so that the final concentration in the coating solution was 5.0% by mass, thereby preparing coating solution (2).

A tetrafluoroethylene-hexafluoropropylene copolymer (FEP) wire (diameter: 1.775 mm) was dip-coated with the coating liquid (2) prepared above at a speed of 10 mm/sec, and then heated at 110°C for 1 hour to carry out a crosslinking reaction. After heating, the temperature was returned to room temperature (25°C), and the FEP wire was removed to obtain a tube (2) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, cross-sectional area: 0.512 ^mm2 ).

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (2) obtained above were measured using the same methods as in Example 1. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (2) were 18.3 gf, 8.4 MPa, and 430%, respectively.

Example 3
A polyurethane elastomer (manufactured by Lubrizol, trade name: Pellethane 2363-80AE) (TPU) was dissolved in N,N-dimethylformamide (DMF) so that the final concentration in the coating solution was 10.0% by mass (solution (3)). The block copolymer (1) synthesized in Synthesis Example 1 above was added to and dissolved in the solution (3) so that the final concentration in the coating solution was 4.5% by mass, thereby preparing coating solution (3).

A tetrafluoroethylene-hexafluoropropylene copolymer (FEP) wire (diameter: 1.775 mm) was dip-coated with the coating solution (3) prepared above at a speed of 10 mm/sec, and then heated at 110°C for 1 hour to carry out a crosslinking reaction. After heating, the temperature was returned to room temperature (25°C), and the FEP wire was removed to obtain a tube (3) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, cross-sectional area: 0.512 ^mm2 ).

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (3) obtained above were measured using the same methods as in Example 1. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (3) were 26.1 gf, 8.0 MPa or more, and 80% or more, respectively.

Comparative Example 1
A tube (4) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, cross-sectional area: 0.512 mm 2 ) made of polyvinyl chloride resin (manufactured by Fujifilm Wako Pure Chemical Industries, Ltd., trade name: polyvinyl chloride, average degree of polymerization (n) = 1,050) ( ^PVC ) was prepared.

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (4) obtained above were measured using the same methods as in Example 1. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (4) were 657.4 gf, 46.5 MPa, and 60%, respectively.

Comparative Example 2
A polytetrafluoroethylene (PTFE) tube (manufactured by Chukoh Chemical Industry Co., Ltd., model number: TUF-100, outer diameter: 3 mm, inner diameter: 2 mm, cross-sectional area: 3.93 mm ² ) (5) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, cross-sectional area: 0.512 mm ² ) was prepared.

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (5) obtained above were measured using the same methods as in Example 1. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (5) were 353.7 gf, 27.5 MPa, and 300%, respectively.

Comparative Example 3
A polyurethane elastomer (trade name: Pellethane 2363-80AE, manufactured by Lubrizol) (TPU) was dissolved in N,N-dimethylformamide (DMF) so that the final concentration in the coating solution was 8.0% by mass (solution (4)).

A tetrafluoroethylene-hexafluoropropylene copolymer (FEP) wire (diameter: 1.775 mm) was dip-coated with the coating liquid (4) prepared above at a speed of 10 mm/sec, and then heated at 110°C for 1 hour. After heating, the temperature was returned to room temperature (25°C), and the FEP wire was removed to obtain a polyurethane elastomer (TPU) tube (6) (inner diameter: 1.775 mm, outer diameter: 1.950 mm, cross-sectional area: 0.512 ^mm2 ).

The sliding resistance (gf), tensile strength (MPa), and tensile elongation (%) of the tube (6) obtained above were measured using the same methods as in Example 1. As a result, the sliding resistance, tensile strength, and tensile elongation of the tube (6) were 750.0 gf, 12.5 MPa, and 470%, respectively.

The results are summarized in Table 1 below. In Table 1 below, "mixing ratio (parts by mass)" indicates the mixing ratio of hydrophobic resin to 100 parts by mass of block copolymer (hydrophobic resin (parts by mass) / 100 parts by mass of block copolymer).

The above results show that all of the tubes in the examples exhibit excellent slip properties (low sliding resistance). Furthermore, tube (1) in Example 1 has a tensile strength similar to that of the PTFE tube (tube (5)). Furthermore, tube (2) in Example 2 has a higher tensile elongation than the PTFE tube (tube (5)).

This application is based on Japanese Patent Application No. 2024-007808, filed on January 23, 2024, the disclosure of which is incorporated by reference in its entirety.

Claims (9)

A medical material comprising: a block copolymer having a structural unit (A) derived from a reactive monomer having an epoxy group and a structural unit (B) derived from a hydrophilic monomer; and at least one hydrophobic resin selected from the group consisting of polyvinyl chloride resins and polyurethane elastomers,
In the medical material, the content of the hydrophobic resin is greater than the content of the block copolymer,
A medical material having a sliding resistance of 50 gf or less and satisfying at least one of a tensile strength of 8.0 MPa or more and a tensile elongation of more than 80%. The medical material according to claim 1, wherein the hydrophobic resin is contained in an amount of 125 parts by mass or more and 300 parts by mass or less per 100 parts by mass of the block copolymer. The medical material according to claim 1, wherein the reactive monomer having an epoxy group includes at least one selected from the group consisting of glycidyl acrylate, glycidyl methacrylate, 3,4-epoxycyclohexylmethyl acrylate, 3,4-epoxycyclohexylmethyl methacrylate, β-methylglycidyl methacrylate, and allyl glycidyl ether. The medical material according to claim 1, wherein the hydrophilic monomer includes at least one selected from the group consisting of N,N-dimethylacrylamide, acrylamide, 2-hydroxyethyl methacrylate, and N-vinylpyrrolidone. A medical device comprising or consisting of the medical material described in any one of claims 1 to 4. The medical device according to claim 5, which is a plastic insertion needle, dilator, sheath (introducer), catheter, or medical tube. A medical device comprising a substrate layer and a coating layer containing or consisting of the medical material described in any one of claims 1 to 4. The medical device according to claim 7, which is a plastic insertion needle, dilator, sheath (introducer), catheter, or medical tube. A method for producing a medical material according to any one of claims 1 to 4,
preparing a mixture by mixing the block copolymer, the hydrophobic resin, and an organic solvent such that the content of the hydrophobic resin is greater than the content of the block copolymer;
The method comprises heat treating the mixture at a temperature of more than 100°C and less than 150°C for 30 minutes to 3 hours.