A ready-to-use, expanded polytetrafluoroethylene vascular graft and method of making same

Abstract

The application provides a ready-to-use expanded polytetrafluoroethylene artificial blood vessel and a preparation method thereof, and the preparation method comprises the following steps: (1) adding a premix containing polytetrafluoroethylene and an organic solvent into an extruder to obtain a green pipe; wherein the extruder adopts a core rod with a tapered gradient and a cylindrical outer sleeve, and the taper of the core rod at the inlet section is larger than that at the outlet section; (2) sequentially performing debinding, longitudinal stretching, heat setting and radial expansion on the green pipe to obtain a base pipe; (3) performing post-treatment on the base pipe to obtain a composite base pipe; (4) performing oxygen plasma pretreatment on the inner surface of the composite base pipe, and then pouring a hydrogel precursor solution into the composite base pipe to obtain the ready-to-use expanded polytetrafluoroethylene artificial blood vessel. The artificial blood vessel provided by the application has an integrated gradient fiber structure, and has excellent overall anti-puncture delamination capacity, stable inner wall anti-thrombosis / endothelialization function and good tissue integration potential.

Description

Technical Field

[0001] This invention relates to the field of implantable medical device technology, particularly to the field of artificial blood vessel technology, and especially to a wearable expanded polytetrafluoroethylene artificial blood vessel and its preparation method. Background Technology

[0002] Expanded polytetrafluoroethylene (ePTFE) possesses excellent chemical stability, biocompatibility, and a tunable porous structure, making it one of the main materials for artificial blood vessels. Chinese patent CN119388815A improves self-sealing after puncture by preparing a ready-to-wear ePTFE artificial blood vessel comprising an inner layer, a middle dense layer, and an outer layer. However, the multi-layer interface still carries the risk of delamination under repeated punctures, and the dense layer may hinder nutrient transport and cell ingrowth. Chinese patent CN120393106A focuses on simple surface modification of the inner wall of ePTFE artificial blood vessels. While this improves blood compatibility, the mechanical interlocking strength between the coating and the porous substrate is limited, potentially leading to peeling under blood flow shear forces and puncture, and it does not contribute to puncture self-healing.

[0003] However, in the application of small-diameter artificial blood vessels with an inner diameter of ≤6mm, three major challenges remain: biocompatibility issues such as thrombosis and intimal hyperplasia, delayed vascular function maturation due to slow endothelialization, and the inherent contradiction between the requirements of wearable design and the structural requirements of antithrombosis / endothelialization promotion.

[0004] Therefore, there is an urgent need to develop a wearable expanded polytetrafluoroethylene artificial blood vessel and its preparation method. Summary of the Invention

[0005] This invention provides a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel and its preparation method. This small-diameter wearable ePTFE artificial blood vessel has excellent overall anti-puncture and delamination ability, stable inner wall anti-thrombotic / endothelialization function, and good tissue integration potential.

[0006] In a first aspect, the present invention provides a method for preparing a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel, comprising: (1) A premix containing polytetrafluoroethylene and organic solvent is added to an extruder to obtain a green tube; wherein the extruder uses a mandrel with a gradient tapered shape and a cylindrical outer die sleeve, and the taper of the mandrel in the inlet section is greater than the taper in the outlet section. (2) The green tube is degreased, longitudinally stretched, heat-set and radially expanded in sequence to obtain the base tube; (3) Post-process the base pipe to obtain a composite base pipe; (4) After oxygen plasma pretreatment of the inner surface of the composite tube, the hydrogel precursor solution is injected into the composite tube to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel.

[0007] Preferably, in step (1): the content of organic solvent in the premix is ​​15wt%~35wt%.

[0008] Preferably, in step (1): the organic solvent is a hydrocarbon solvent.

[0009] More preferably, the organic solvent is n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, or methylcyclohexane.

[0010] Preferably, in step (1): the working section of the mandrel adopts a continuously adjustable taper or a segmented taper so that the mandrel and the outer mold sleeve form a non-equidistant annular flow channel; the working section includes an inlet section, a transition section and an outlet section.

[0011] More preferably, the cone angle of the inlet section ranges from 2° to 8°, and the cone angle of the outlet section ranges from 0.5° to 2°.

[0012] Preferably, in step (1): when the taper of the working section is continuously adjustable, the taper variation rate of the working section is 0.09~1.4° / cm along the length direction of the working section.

[0013] Preferably, in step (1): when the working section adopts a segmented taper, the taper difference between the inlet section and the transition section is 1°~4°; the taper difference between the transition section and the outlet section is 0.5°~2°.

[0014] More preferably, in step (1): when the working section adopts a segmented taper, the taper difference between the inlet section and the working section is 2°~3°; the taper difference between the working section and the outlet section is 1°~1.5°.

[0015] Preferably, in step (2): the degreasing temperature is 150~280℃; the time is 30~180s; and the heating rate is 5~20°C / min.

[0016] More preferably, in step (2): the degreasing temperature is 200~260℃; the time is 60~120s.

[0017] Preferably, in step (2): the longitudinal stretching speed is 2~50m / min; the stretching ratio is 2~10 times; and the stretching temperature is 200~260℃.

[0018] More preferably, in step (2): the longitudinal stretching ratio is 3 to 6 times; the stretching temperature is 220 to 250°C.

[0019] Preferably, in step (2): the heat setting temperature is 250~380℃; the time is 10~120s.

[0020] More preferably, in step (2): the heat setting temperature is 300~360℃; the time is 30~60s.

[0021] Preferably, in step (2): the radial expansion temperature is 200~260℃; the expansion ratio is 1.5~5.0 times; and the expansion speed is 10~100mm / min.

[0022] More preferably, in step (2): the radial expansion temperature is 220~250℃; the expansion ratio is 2.0~3.5 times; and the expansion speed is 20~50mm / min.

[0023] Preferably, in step (3): the post-treatment is at least one of gradient heat treatment or wrapping sintering reinforcement treatment.

[0024] More preferably, the gradient heat treatment temperature is 80~300℃ and the time is 10~30min.

[0025] More preferably, the wrapping and sintering reinforcement treatment involves uniformly wrapping an expanded polytetrafluoroethylene film around the outer surface of the base tube, and then sintering the wrapped base tube at 360°C for 5 to 15 minutes.

[0026] More preferably, the expanded polytetrafluoroethylene film has a thickness of 2~30μm, a width of 3~15mm, and a winding angle of 45°~75°.

[0027] Preferably, in step (4): the gas flow rate used for oxygen plasma pretreatment is 10~30 sccm; the radio frequency power is 20~80W; the working gas pressure is 10~50Pa; and the time is 30~120s.

[0028] More preferably, in step (4): the gas flow rate used for oxygen plasma pretreatment is 15~20 sccm; the radio frequency power is 30~50W; and the working gas pressure is 30~40Pa.

[0029] Preferably, in step (4): the hydrogel precursor solution includes component A and component B in a volume ratio of (4~9):1; component A includes a monomer containing sulfonic acid groups, extracellular matrix-derived peptides and deionized water in a mass ratio of (5~20):(0.01~0.5):100; component B includes an initiator, a crosslinking agent and deionized water in a mass ratio of (0.1~1):(0.01~0.5):100.

[0030] Preferably, the infusion is performed using dynamic pressure circulation infusion at a temperature of 35~40℃, and the circulation is carried out for 10~60 minutes under conditions lower than the ethanol bubble point pressure.

[0031] Preferably, the monomer containing the sulfonic acid group is 2-acrylamido-2-methylpropanesulfonic acid, sodium vinyl sulfonate, sodium styrene sulfonate, sodium salt of 2-acrylamido-2-methylpropanesulfonic acid, potassium salt of 3-sulfopropyl acrylate, or potassium salt of 3-sulfonate methacrylate. The extracellular matrix-derived peptides include one or more of the following: cell adhesion peptides, aspartic-glycyl-glutamyl-alanine tetrapeptide, and elastin-derived valine-alanyl-prolyl-glycine tetrapeptide; the cell adhesion peptides include at least one of the following: arginine-glycine-aspartic acid tripeptide, arginine-glutamic acid-aspartic acid-valine tetrapeptide, tyrosyl-isoleucyl-glycyl-seryl-arginine pentapeptide, and isoleucyl-lysyl-valine-alanyl-valine pentapeptide.

[0032] Preferably, the initiator is ammonium persulfate; The crosslinking agent is N,N'-methylenebisacrylamide.

[0033] Preferably, after injecting the hydrogel precursor solution into the composite base tube to obtain the hydrogel layer, the process further includes: hydrophilic modification treatment.

[0034] More preferably, the hydrophilic modification treatment involves immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 10 seconds to 20 minutes.

[0035] More preferably, the solid content of the perfluorosulfonic acid solution is 10 wt%.

[0036] Secondly, embodiments of the present invention also provide a wearable expanded polytetrafluoroethylene artificial blood vessel prepared by any of the preparation methods described in the first aspect above. The wearable expanded polytetrafluoroethylene artificial blood vessel consists of a hydrogel layer with a thickness of 5-50 μm and an integrated gradient fiber structure from the inside to the outside. The average fiber diameter of the gradient fiber structure increases from the inside to the outside, and the porosity increases from 40%-60% to 70%-85%.

[0037] Compared with the prior art, the present invention has at least the following beneficial effects: (1) This invention employs a mandrel with a gradient tapered shape and a cylindrical outer mold sleeve for mold flow field design. Through extrusion processes such as degreasing, longitudinal stretching, heat setting, and radial expansion, a base tube with a gradient fibrous structure is directly and integrally prepared, replacing the traditional multi-layer compaction composite process. This fundamentally eliminates the interlayer interface and solves the core hidden danger of puncture delamination. Simultaneously, the hydrogel liner constructed in situ on the inner surface forms a deeply rooted composite with this gradient fibrous structure, with a bonding strength far exceeding that of surface physical coatings or adsorption. This makes it less prone to detachment under blood flow and puncture, ensuring long-lasting function. The ready-to-wear expanded polytetrafluoroethylene artificial blood vessel prepared using the method of this invention possesses excellent overall anti-puncture delamination capability, stable inner wall anti-thrombotic / endothelialization function, and good tissue integration potential.

[0038] (2) The wearable expanded polytetrafluoroethylene artificial blood vessel prepared by the present invention has excellent puncture self-sealing properties due to its gradient fibrosis structure and hydrogel composite, and is wearable immediately; the sulfonic acid groups of the hydrogel layer simulate the surface of the vascular endothelium, providing excellent anti-protein adsorption and anti-platelet adhesion capabilities, and is blood compatible; the cell adhesion peptides fixed in the hydrogel layer can actively guide the adhesion and spread of endothelial cells, and accelerate the endothelialization process; at the same time, the macroporous structure of the outer layer of the integrated gradient fibrosis structure is conducive to fibroblast ingrowth and collagen deposition, achieving stable anchoring and promoting tissue integration.

[0039] (3) The preparation method of the present invention combines mold design, material modification and interface polymerization, and the process is controllable and has industrialization potential. Attached Figure Description

[0040] To more clearly illustrate the technical solutions in the embodiments of the present invention or the prior art, the drawings used in the description of the embodiments or the prior art will be briefly introduced below. Obviously, the drawings described below are some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0041] Figure 1 This is a flowchart of a method for preparing a wearable expanded polytetrafluoroethylene artificial blood vessel according to an embodiment of the present invention; Figure 2 This is a schematic diagram of a mandrel structure provided in an embodiment of the present invention; Figure 3 This is a schematic cross-sectional view of a wearable expanded polytetrafluoroethylene artificial blood vessel provided in an embodiment of the present invention. Detailed Implementation

[0042] To make the objectives, technical solutions, and advantages of the embodiments of the present invention clearer, the technical solutions of the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are some embodiments of the present invention, but not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without creative effort are within the scope of protection of the present invention.

[0043] The following describes the specific implementation of the concept in this application.

[0044] Please refer to Figure 1 and Figure 2 This invention provides a method for preparing a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel, comprising: (1) A premix containing polytetrafluoroethylene and organic solvent is added to an extruder to obtain a green tube; wherein the extruder uses a mandrel with a gradient tapered shape and a cylindrical outer die sleeve, and the taper of the mandrel in the inlet section is greater than the taper in the outlet section. (2) The green tube is degreased, longitudinally stretched, heat-set and radially expanded in sequence to obtain the base tube; (3) Post-process the base pipe to obtain a composite base pipe; (4) After oxygen plasma pretreatment of the inner surface of the composite tube, the hydrogel precursor solution is injected into the composite tube to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel.

[0045] In this embodiment of the invention, a mandrel with a gradient tapered shape and a cylindrical outer mold sleeve are used for mold flow field design. Through extrusion processes such as degreasing, longitudinal stretching, heat setting, and radial expansion, a base tube with a gradient fibrous structure is directly and integrally prepared, replacing the traditional multi-layer compaction composite process. This fundamentally eliminates the interlayer interface and solves the core hidden danger of puncture delamination. Simultaneously, a hydrogel liner constructed in situ on the inner surface forms a deeply rooted composite with this gradient fibrous structure, with a bonding strength far exceeding that of surface physical coatings or adsorption. This makes it less prone to detachment under blood flow and puncture, ensuring long-lasting function. The ready-to-wear expanded polytetrafluoroethylene artificial blood vessel prepared using the method of this invention possesses excellent overall anti-puncture delamination capability, stable inner wall anti-thrombotic / endothelialization function, and good tissue integration potential.

[0046] In a preferred embodiment, in step (1): the content of organic solvent in the premix is ​​15wt% to 35wt% (for example, it can be 15 wt%, 20 wt%, 25 wt%, 30 wt% or 35 wt%).

[0047] In a preferred embodiment, in step (1): the organic solvent is a hydrocarbon solvent.

[0048] In a more preferred embodiment, the hydrocarbon solvent is n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, or methylcyclohexane.

[0049] It should be noted that hydrocarbon solvents include, but are not limited to, the following: isoparaffin solvents with low aromatic content such as the Isopar series (e.g., Isopar G, Isopar H), Exxsol series, n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, methylcyclohexane, etc. In the embodiments of the present invention, by limiting the content of organic solvent in the premix to the range of 15wt%~35wt%, good formability, stretching uniformity, and controllable pore structure can be obtained. Specifically, due to the high viscosity of the premix, it cannot be extruded using a screw extruder, but must be extruded using a plunger extruder.

[0050] In one specific embodiment, the extrusion speed of the extruder is 0.5~5.0 m / min (e.g., 0.5 m / min, 1 m / min, 1.5 m / min, 2 m / min, 2.5 m / min, 3 m / min, 3.5 m / min, 4 m / min, 4.5 m / min or 5.0 m / min), preferably 1.0~3.0 m / min; the extrusion temperature is 30~60℃ (e.g., 30℃, 35℃, 40℃, 45℃, 50℃, 55℃ or 60℃), preferably 40~50℃; and the extrusion pressure is 5~30 MPa (e.g., 5 MPa, 10 MPa, 15 MPa, 20 MPa, 25 MPa, 30 MPa, 35 MPa, 40 MPa, 45 MPa or 50 MPa), preferably 10~20 MPa.

[0051] In the implementation of this invention, experiments have confirmed that if the extrusion speed is greater than 5.0 m / min, it easily leads to melt fracture or surface roughness in the premix; however, if the extrusion speed is less than 0.5 m / min, it reduces production efficiency and results in insufficient orientation. Simultaneously, the extrusion temperature is limited to 30~60℃, with no heating or only slight heating, to maintain the lubricity of the organic solvent; the extrusion pressure is limited to 5~30 MPa to avoid insufficient densification due to excessively low pressure, and to avoid excessively high pressure leading to mold wear or overheating and degradation of the premix.

[0052] In a preferred embodiment, in step (1): the working section of the mandrel adopts a continuously adjustable taper or a segmented taper (e.g., Figure 2 , Figure 3 As shown in the figure, so that the mandrel and the outer mold sleeve form a non-equidistant annular flow channel; the working section includes an inlet section, a transition section and an outlet section.

[0053] It should be noted that the working section is divided into 2 to 4 sections (for example, 2, 3, or 4), with 3 sections being preferred. Figure 2 and Figure 3 The middle section consists of inlet sections on the right and inlet sections on the left. The middle section consists of transition sections on the right and outlet sections on the left.

[0054] In this embodiment of the invention, the working section of the mandrel is designed with a continuously adjustable or segmented taper, with a larger taper at the inlet and a gentler taper at the outlet, forming a non-equidistant annular flow channel with the outer mold sleeve. This design allows the inner layer of the premix to withstand higher shear and tensile rates than the outer layer when the premix flows through the mold composed of the mandrel and the outer mold sleeve. During subsequent degreasing, longitudinal stretching, heat setting, and low-temperature radial expansion processes, the inner layer, with its higher degree of molecular chain pre-orientation, will preferentially form a finer, more oriented fiber network, while the outer layer will form a relatively looser, coarser fiber network with larger nodes. This results in an expanded polytetrafluoroethylene (ePTFE) hollow fiber tube with a denser inner layer and a sparser outer layer, exhibiting a continuous gradient change in fiber / node size. Thus, the gradient structure endows the base tube with the following characteristics: a) a dense inner surface that provides a smooth blood contact interface and facilitates the adhesion of the hydrogel layer; b) high mechanical strength in the intermediate transition zone; and c) a loose and porous outer surface that facilitates tissue cell ingrowth and nutrient exchange.

[0055] In a more preferred embodiment, the cone angle of the inlet section ranges from 2° to 8° (e.g., it can be 1°, 2°, 3°, 4°, 5°, 6°, 7° or 8°), and the cone angle of the outlet section ranges from 0.5° to 2° (e.g., it can be 0.5°, 1°, 1.5° or 2°).

[0056] In a preferred embodiment, in step (1): when the taper of the working section is continuously adjustable, the taper variation rate of the working section along the length direction is 0.09~1.4° / cm (for example, it can be 0.09° / cm, 0.1° / cm, 0.12° / cm, 0.15° / cm, 0.18° / cm, 0.2° / cm, 0.25° / cm, 0.3° / cm, 0.4° / cm, 0.5° / cm, 0.6° / cm, 0.7° / cm, 0.8° / cm, 0.9° / cm, 1.0° / cm, 1.1° / cm, 1.2° / cm, 1.3° / cm or 1.4° / cm).

[0057] In a preferred embodiment, in step (1): when the working section adopts a segmented taper, the taper difference between the inlet section and the transition section is 1° to 4° (for example, it can be 1°, 1.5°, 2°, 2.5°, 3°, 3.5° or 4°); the taper difference between the transition section and the outlet section is 0.5° to 2° (for example, it can be 0.5°, 1°, 1.5° or 2°).

[0058] In a more preferred embodiment, in step (1): when the working section adopts a segmented taper, the taper difference between the inlet section and the working section is 2°~3° (for example, it can be 2°, 2.2°, 2.5°, 2.6°, 2.8° or 3°); the taper difference between the working section and the outlet section is 1°~1.5° (for example, it can be 1°, 1.2°, 1.4° or 1.5°).

[0059] More specifically, for segmented tapers, the taper of the inlet section is 2°~8°, the taper of the transition section is 1.5°~4°, and the taper of the outlet section is 0.5°~2°; for continuously adjustable tapers, the taper change rate of the inlet section is 0.35~1.4° / cm, the taper change rate of the transition section is 0.25~0.7° / cm, and the taper change rate of the outlet section is 0.09~0.35° / cm.

[0060] It should be noted that taper refers to the angle formed between the generatrix of the side of the mandrel and the central axis. Specifically, the length ratios of the inlet section, transition section, and outlet section of the mandrel working section are 20%~35%, 30%~50%, and 20%~35%, respectively. The length of the mandrel working section is usually 50~200mm, preferably 80~150mm.

[0061] In this embodiment of the invention, the large taper of the inlet section allows the premix to withstand a high shear rate upon entering the flow channel, promoting the preferential orientation of the inner molecular chains. The gradual taper change in the working section ensures a smooth transition in shear rate, avoiding sudden stress changes. The gentle taper of the outlet section allows the outer melt to achieve final shaping under low shear conditions, forming a differentiated orientation structure between the inner and outer layers. Thus, this segmented taper design or continuously adjustable taper design allows the premix melt to be gradually compressed and released in the flow channel, reducing radial rebound after extrusion and improving the wall thickness uniformity and roundness accuracy of the green tube. When the working section uses continuously adjustable taper, the taper change rate is controlled within the range of 0.09~1.4° / cm along the length of the working section; when the working section uses segmented taper, the tapers of adjacent working sections are different. Experiments have shown that excessive taper variation leads to the following problems: a sharp increase in the shear and tensile rates of the inner premix melt, resulting in localized stress concentration and a tendency to crack or break; excessive difference in flow rates between the inner and outer premix melts, causing rough extruded surfaces and uneven thickness; and excessively steep taper at the exit section, resulting in severe radial springback of the green tube and poor dimensional stability. Insufficient taper variation leads to the following problems: insufficient shear rate difference between the inner and outer premix melts, failing to form the expected highly oriented inner layer structure; and nearly equidistant annular channels, resulting in a small gradient structure difference between the outer and inner layers of the base tube, which reduces the interfacial bonding force between the subsequent hydrogel layer and the inner layer of the base tube; simultaneously, excessively low extrusion pressure and insufficient premix orientation also lead to a decrease in the tensile properties of the base tube.

[0062] In a preferred embodiment, in step (2): the degreasing temperature is 150~280℃ (e.g., 150℃, 180℃, 200℃, 250℃ or 280℃); the time is 30~180s (e.g., 30s, 50s, 80s, 100s, 120s, 150s or 180s); and the heating rate is 5~20℃ / min (e.g., 5℃ / min, 10℃ / min, 15℃ / min or 20℃ / min).

[0063] In a more preferred embodiment, in step (2): the degreasing temperature is 200~260°C (e.g., 200°C, 210°C, 220°C, 230°C, 240°C, 250°C or 260°C); the time is 60~120s (e.g., 60s, 70s, 80s, 90s, 100s, 110s or 120s).

[0064] In this embodiment of the invention, experiments have confirmed that if the degreasing temperature is below 150°C, organic solvents are prone to remain, affecting subsequent stretching; however, if the degreasing temperature is above 280°C, the polytetrafluoroethylene will sinter prematurely and lose its stretchability. The degreasing time depends on the wall thickness of the green tube and the organic solvent content, with the specific time determined by ensuring complete evaporation of the organic solvent. If the heating rate is too fast, it can easily lead to rapid evaporation of the organic solvent, resulting in bubbles or cracks. Therefore, under the premise of ensuring rapid evaporation and avoiding the generation of bubbles or cracks, the heating rate is limited to 5~20°C / min.

[0065] In a preferred embodiment, in step (2): the longitudinal stretching speed is 2~50 m / min (e.g., 2 m / min, 5 m / min, 10 m / min, 15 m / min, 20 m / min, 25 m / min, 30 m / min, 35 m / min, 40 m / min, 45 m / min or 50 m / min); the draw ratio is 2~10 times (e.g., 2 times, 3 times, 4 times, 5 times, 6 times, 7 times, 8 times, 9 times or 10 times); the stretching temperature is 200~260℃ (e.g., 200℃, 210℃, 220℃, 230℃, 240℃, 250℃ or 260℃).

[0066] In a more preferred embodiment, in step (2): the longitudinal stretching ratio is 3 to 6 times (for example, it can be 3 times, 3.5 times, 4 times, 4.5 times, 5 times, 5.5 times or 6 times); the stretching temperature is 220 to 250°C (for example, it can be 220°C, 225°C, 230°C, 235°C, 240°C, 245°C or 250°C).

[0067] It should be noted that the feeding speed during the longitudinal stretching process is matched with the extrusion speed of the extruder, and the feeding speed is 0.5~5.0m / min; the stretching speed is the same as the winding speed, and the stretching speed depends on the draw ratio and the target output; the length of the stretching section is 1~10m to ensure sufficient residence time at the stretching temperature.

[0068] In this embodiment of the invention, experiments have confirmed that if the draw ratio is less than 2, the porosity of the resulting base tube is too low; however, if the draw ratio is greater than 10, the fibers are prone to breakage during longitudinal stretching, leading to a decrease in strength. If the stretching temperature is below 200°C, the molecular chain movement is insufficient; however, if the stretching temperature is above 260°C, the polytetrafluoroethylene begins to sinter and densify.

[0069] In a preferred embodiment, in step (2): the heat setting temperature is 250~380℃ (for example, it can be 250℃, 280℃, 300℃, 320℃, 350℃ or 380℃); the time is 10~120s (for example, it can be 10s, 20s, 40s, 50s, 60s, 80s, 100s or 120s).

[0070] In a more preferred embodiment, in step (2): the heat setting temperature is 300~360°C (e.g., 300°C, 310°C, 320°C, 330°C, 340°C, 350°C or 360°C); the time is 30~60s (e.g., 30s, 40s, 50s or 60s).

[0071] It should be noted that heat setting uses a hot air oven or heated rollers, or both, preferably both. Hot air setting ensures more uniform heat setting, while heated rollers improve heat setting efficiency. The heat setting time depends on the film thickness and setting temperature, and it is essential to ensure that the fiber-node structure is fully locked in.

[0072] In this embodiment of the invention, if the heat setting temperature is below 250°C, the heat setting will be insufficient; however, if the heat setting temperature is above 380°C, the polytetrafluoroethylene will be over-sintered, causing the pore structure to collapse.

[0073] In a preferred embodiment, in step (2): the radial expansion temperature is 200~260℃ (e.g., 200℃, 210℃, 220℃, 230℃, 240℃, 250℃ or 260℃); the expansion ratio is 1.5~5.0 times (e.g., 1.5 times, 2.0 times, 2.5 times, 3.0 times, 3.5 times, 4.0 times, 4.5 times or 5.0 times); the expansion speed is 10~100 mm / min (e.g., 10 mm / min, 20 mm / min, 40 mm / min, 50 mm / min, 60 mm / min, 70 mm / min, 80 mm / min, 90 mm / min or 100 mm / min).

[0074] In a more preferred embodiment, in step (2): the radial expansion temperature is 220~250°C (e.g., 220°C, 225°C, 230°C, 235°C, 240°C, 245°C or 250°C); the expansion ratio is 2.0~3.5 times (e.g., 2.0 times, 2.2 times, 2.5 times, 2.8 times, 3.0 times, 3.2 times or 3.5 times); and the expansion speed is 20~50 mm / min (e.g., 20 mm / min, 25 mm / min, 30 mm / min, 35 mm / min, 40 mm / min, 45 mm / min or 50 mm / min).

[0075] It should be noted that the expansion medium is compressed air or nitrogen, and the gas pressure must be kept stable to avoid pressure fluctuations.

[0076] In this embodiment of the invention, the radial expansion temperature is kept consistent with the longitudinal stretching temperature to ensure that the molecular chains have sufficient mobility during radial expansion. During this process, if the expansion ratio is too low, the tube diameter will be insufficient; if the expansion rate is too high, the tube wall will be too thin and the mechanical strength will decrease; if the expansion rate is too fast, it will easily lead to uneven tube wall thickness or rupture. Therefore, the expansion ratio is limited to 1.5 to 5.0 times, and the expansion rate is limited to 10 to 100 mm / min.

[0077] In a preferred embodiment, in step (3): the post-processing is at least one of gradient heat treatment or wrapping.

[0078] In this embodiment of the invention, to further optimize the mechanical properties, structural stability, and interfacial hydrophilicity of the base tube with a gradient fibrous structure, so as to meet the requirements of immediate-penetration artificial blood vessels for puncture self-sealing, tissue integration, and long-term stability, a post-processing control process is introduced after step (2), specifically including the following two processing methods that can be used independently or in combination. For example, the post-processing can only use gradient heat treatment or wrapping sintering reinforcement treatment, or a combination of the two methods. When using the above two methods in combination, the order of post-processing is wrapping sintering reinforcement treatment followed by gradient heat treatment.

[0079] In a more preferred embodiment, the gradient heat treatment temperature is 80~300°C (e.g., 80°C, 100°C, 120°C, 150°C, 200°C, 220°C, 250°C or 300°C), and the time is 10~30 min (e.g., 10 min, 15 min, 20 min, 25 min or 30 min).

[0080] In this embodiment of the invention, gradient heat treatment can be performed in a fixed state (such as with mandrel support) and a non-fixed state (without support), which can achieve microstructural control of the fiber network of the tube wall, thereby controlling the thermal stability, crystallinity and residual stress release of the fiber structure of the tube, optimizing its mechanical properties and dimensional stability, and improving the deformation recovery ability of the tube wall during repeated punctures.

[0081] In a more preferred embodiment, the wrapping and sintering reinforcement treatment involves uniformly wrapping an expanded polytetrafluoroethylene film around the outer surface of the base tube, and then sintering the wrapped base tube at 360°C for 5 to 15 minutes (for example, 5 minutes, 6 minutes, 8 minutes, 10 minutes, 12 minutes or 15 minutes).

[0082] In a more preferred embodiment, the expanded polytetrafluoroethylene film has a thickness of 2-30 μm (e.g., 2 μm, 5 μm, 6 μm, 8 μm, 10 μm, 15 μm, 20 μm, 25 μm or 30 μm), a width of 3-15 mm (e.g., 3 mm, 5 mm, 6 mm, 8 mm, 10 mm, 12 mm or 15 mm), and a winding angle of 45°-75° (e.g., 45°, 50°, 55°, 60°, 65°, 70° or 75°).

[0083] It should be noted that the winding angle is the angle between the winding direction and the central axis of the base tube.

[0084] In this embodiment of the invention, to further enhance the outer structural strength and overall puncture resistance of the base tube, an expanded polytetrafluoroethylene (ePTFE) membrane is wrapped around it and then subjected to high-temperature sintering. Experiments have shown that a thicker ePTFE membrane provides stronger reinforcement, but excessive thickness can lead to uneven adhesion between the wrapping layer and the base tube, and excessive shrinkage stress during sintering causing wrinkles. Conversely, a thinner ePTFE membrane results in insufficient reinforcement, limited improvement in puncture resistance and delamination, decreased radial compliance, and potential obstruction of outer tissue ingrowth. Therefore, the thickness of the ePTFE membrane is limited to 2-30 μm. The width of the ePTFE membrane affects wrapping efficiency and coverage uniformity. An excessively wide ePTFE membrane is prone to wrinkles or uneven overlap on curved surfaces; an excessively narrow ePTFE membrane requires more wrapping turns, increasing the number of joints and potentially creating weak points. Therefore, the width of the ePTFE membrane is limited to 3-15 mm. The winding angle determines the anisotropy of the reinforcing layer. However, if the winding angle is too small, the radial reinforcement is insufficient, resulting in weak burst resistance; if the winding angle is too large, the axial reinforcement is insufficient, leading to low longitudinal stiffness and easy axial deformation of the tube. Therefore, the winding angle is limited to 45°~75°. In this way, this treatment ensures that the outer layer of the composite tube remains hydrophobic while enhancing the bonding between the fibers in the outer layer, improving the peel resistance and structural integrity of the composite tube, while maintaining the gradient structure of the inner layer and enhancing the overall resistance to puncture and delamination.

[0085] In a preferred embodiment, in step (4): the gas flow rate used for oxygen plasma pretreatment is 10~30 sccm (e.g., 10 sccm, 15 sccm, 20 sccm, 25 sccm or 30 sccm); the radio frequency power is 20~80W (e.g., 20 W, 25 W, 30 W, 35 W, 40 W, 45 W, 50 W, 55 W, 60 W, 65 W, 70 W, 75 W or 80 W); the working gas pressure is 10~50 Pa (e.g., 10 Pa, 15 Pa, 20 Pa, 25 Pa, 30 Pa, 35 Pa, 40 Pa, 45 Pa or 50 Pa); and the time is 30~120 s (e.g., 30 s, 40 s, 50 s, 60 s, 80 s, 90 s, 100 s or 120 s).

[0086] In a more preferred embodiment, in step (4): the gas flow rate used for oxygen plasma pretreatment is 15-20 sccm (e.g., 15 sccm, 16 sccm, 17 sccm, 18 sccm, 19 sccm or 20 sccm); the radio frequency power is 30-50 W (e.g., 30 W, 32 W, 35 W, 38 W, 40 W, 42 W, 45 W, 48 W or 50 W); and the working gas pressure is 30-40 Pa (e.g., 30 Pa, 32 Pa, 34 Pa, 35 Pa, 36 Pa, 38 Pa or 40 Pa).

[0087] It should be noted that, in order to prevent local overheating of ePTFE from causing junction melting or pore collapse during this process, it is preferable to ensure that the temperature of the composite base tube is below 60°C through water cooling or intermittent treatment. Specifically, since the composite base tube has a certain length, the two ends of the tube cavity are sealed and low-pressure plasma is introduced only into the inner cavity, or the entire tube cavity is evacuated and then diffused with a small amount of gas, or the sample is rotated to ensure uniform exposure of the inner surface.

[0088] In this embodiment of the invention, to achieve shallow (several micrometers) activation of the inner surface of the composite substrate to introduce oxygen-containing groups (including -OH or -COOH), while avoiding damage to the porous ePTFE framework and excessive etching and embrittlement of the outer wall, and maintaining the hydrophobicity of the outer surface, a radio frequency low-pressure oxygen plasma is used to pretreat the inner surface of the composite substrate. The oxygen plasma pretreatment uses pure oxygen or a mixture of oxygen and argon, preferably pure oxygen, to facilitate the introduction of oxygen-containing polar groups. The oxygen flow rate is limited to 10-30 sccm to avoid plasma instability due to excessively low flow rate and excessive etching due to excessively high flow rate. Simultaneously, the working pressure (cavity vacuum) is limited to 10-50 Pa for the inner surface of the composite substrate to ensure uniform glow discharge. If the radio frequency power exceeds 80W, it can easily lead to deep-hole etching, nanopillar structures, or even restore the superhydrophobicity of the composite substrate. Furthermore, to avoid excessively deep CF bond breakage due to excessive processing time, resulting in surface embrittlement or hydrophobicity restoration, the processing time is limited to 30-120 s.

[0089] In this embodiment of the invention, polytetrafluoroethylene has extremely low surface energy (approximately 18~20 mN / m) and high CF bond inertness. If oxygen plasma pretreatment is not performed, direct injection of the hydrogel precursor solution will result in: 1) Inability to form covalent bonds: The inner surface of the composite base tube lacks active sites such as oxygen-containing groups. The hydrogel can only be bonded through van der Waals forces or weak mechanical intercalation, lacking COC or CN covalent bond anchoring, resulting in a bonding strength far lower than that after plasma activation and subsequent polymerization; 2) Poor wettability of the hydrogel precursor solution: The hydrophobic inner surface makes it difficult for component A to completely wet the inner surface, easily leading to problems such as discontinuous coating, bubble entrainment, localized missing coating, and uneven hydrogel layer thickness; 3) Easy peeling under blood flow scouring: Under arterial blood flow shear force (greater than 10~15 dyn / cm), 2 ) and during repeated puncture-hemostasis cycles, the hydrogel layer, which is physically adsorbed or simply interlocked, is very easy to fall off in whole or swell and be lost, thus losing its long-lasting antithrombotic and endothelialization functions; 4) Insufficient mechanical interlocking depth: After slight etching and activation by oxygen plasma, the hydrogel can penetrate into the outermost micropores and covalently crosslink with the activated groups during polymerization; for those without oxygen plasma pretreatment, mechanical interlocking is only achieved through shallow pores with weak bonding force, and the puncture site is prone to delamination along the interface.

[0090] In a preferred embodiment, in step (4): the hydrogel precursor solution includes component A and component B in a volume ratio of (4~9):1 (e.g., 4:1, 4.5:1, 5:1, 5.5:1, 6:1, 6.5:1, 7:1, 7.5:1, 8:1, 8.5:1 or 9:1); component A includes a monomer containing sulfonic acid groups, extracellular matrix-derived peptides and deionized water in a mass ratio of (5~20):(0.01~0.5):100; component B includes an initiator, a crosslinking agent and deionized water in a mass ratio of (0.1~1):(0.01~0.5):100.

[0091] Specifically, in component A, the mass ratio of the monomer containing sulfonic acid groups to deionized water can be any ratio from 5:100 to 20:100, for example, it can be 5:100, 6:100, 7:100, 8:100, 9:100, 10:100, 10.5:100, 11:100, 11.5:100, 12:100, 12.5:100, 13:100, 13.5:100, 14:100, 14.5:100, 15:100, 16:100, 17:100, 18:100, 19:100 or 20:100. The mass ratio of extracellular matrix-derived peptides to deionized water can be any ratio from 0.01:100 to 0.5:100, for example, 0.01:100, 0.02:100, 0.05:100, 0.06:100, 0.08:100, 0.1:100, 0.15:100, 0.2:100, 0.25:100, 0.3:100, 0.35:100, 0.4:100, 0.45:100, or 0.5:100. In component B, the mass ratio of initiator to deionized water can be any ratio from 0.1:100 to 1:100, for example, 0.1:100, 0.2:100, 0.3:100, 0.4:100, 0.5:100, 0.6:100, 0.7:100, 0.8:100, 0.9:100, or 1:100; the mass ratio of crosslinking agent to deionized water can be 0.01: Any ratio from 100 to 0.5:100, for example, 0.01:100, 0.02:100, 0.05:100, 0.06:100, 0.08:100, 0.1:100, 0.15:100, 0.2:100, 0.25:100, 0.3:100, 0.35:100, 0.4:100, 0.45:100, or 0.5:100.

[0092] In this embodiment of the invention, experiments have confirmed that, in component A, under the premise that the amounts of other raw materials remain unchanged, if the amount of monomers containing sulfonic acid groups is too low, the charge density of the hydrogel network will be insufficient, and the ability to resist protein adsorption and platelet adhesion will be significantly reduced. At the same time, the swelling rate of the hydrogel will be too high, the mechanical strength will be weak, and it will be easily deformed or damaged under blood flow. However, if the amount of monomers containing sulfonic acid groups is too high, the viscosity of the system will be too high, which is not conducive to uniform perfusion. It may also lead to excessive hydrophilic swelling of the hydrogel, uncontrolled thickness of the hydrogel layer, and decreased mechanical strength. At the same time, excessive sulfonic acid groups may also inhibit the normal adhesion and spreading of endothelial cells. Assuming other raw material usage remains constant, excessively high levels of extracellular matrix-derived peptides can lead to a significant increase in costs. Furthermore, excessively high ligand density may cause receptor saturation, inhibiting cell migration and proliferation. Some peptide sequences may also cause non-specific protein adsorption at high densities. Conversely, insufficient levels of extracellular matrix-derived peptides will result in inadequate surface peptide density to trigger effective integrin-mediated adhesion signals, leading to a significant decrease in endothelial cell adhesion rate and spreading area, thus failing to achieve the desired endothelialization effect. Therefore, component A is limited to a sulfonic acid-containing monomer, extracellular matrix-derived peptides, and deionized water in a mass ratio of (5~20):(0.01~0.5):100.

[0093] In this embodiment of the invention, for component B, under the premise that the amounts of other raw materials remain unchanged, if the amount of crosslinking agent is too low, it will lead to insufficient crosslinking density, loose hydrogel network, poor mechanical strength, and excessive swelling rate, making it easy to break and be lost under the flushing of blood flow; the thickness of the hydrogel layer will be difficult to control precisely, and local collapse may occur; however, if the amount of crosslinking agent is too high, it will lead to excessively high crosslinking density, the hydrogel network will be too rigid and brittle, the compliance will decrease, and the mechanical matching with the soft blood vessel wall will be poor; the pores will be too small, hindering cell inward growth; the swelling rate will be low, and the swelling compensation capacity of the hydrogel required for self-sealing after puncture will be weakened. Under the premise that the amounts of other raw materials remain unchanged, if the amount of initiator is too low, there will be insufficient free radical generation, incomplete polymerization reaction or extremely slow rate, low conversion rate, poor mechanical properties of hydrogel, and unreacted monomers may be cytotoxic; however, if the amount of initiator is too high, the polymerization rate will be too fast, the exothermic reaction will be violent, and local burst polymerization in the lumen may lead to blockage, and excessive initiator residues will have a toxic effect on cells. Therefore, component B is defined as including an initiator, a crosslinking agent, and deionized water in a mass ratio of (0.1~1):(0.01~0.5):100.

[0094] It should be noted that the amount of deionized water used should only be sufficient to ensure that all solids are dissolved. Specifically, the hydrogel precursor solution should include component A and component B in a volume ratio of (4~9):1 to avoid excessive component B, which would lead to excessively rapid polymerization and cross-linking, resulting in a brittle and hard hydrogel layer with severe shrinkage and poor adhesion to the tube wall. Simultaneously, the gelation time during infusion is relatively short; if component B is too low, the lumen cannot be fully filled during the circulation infusion phase, leading to uneven hydrogel lining thickness or tube blockage. Furthermore, residual monomers containing sulfonic acid groups can increase cytotoxicity.

[0095] In a preferred embodiment, the perfusion is performed using dynamic pressure circulation perfusion at a temperature of 35-40°C (e.g., 35°C, 36°C, 37°C, 38°C, 39°C, or 40°C), and the perfusion is performed for 10-60 minutes (e.g., 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, or 60 minutes) at a pressure lower than the ethanol bubble point.

[0096] Specifically, the dynamic pressure cyclic perfusion includes: sealing one end of the base tube and connecting the other end to a peristaltic pump. First, the inner wall of the composite base tube, pretreated with oxygen plasma, is wetted with ethanol. Then, under pressure lower than the ethanol bubble point, a solution of a mixture of components A and B in a specific ratio is circulated into the tube lumen, and the circulation is maintained at a constant temperature of 35-40°C for 10-60 minutes. During this process, the activated ePTFE inner surface initiates free radical polymerization. The hydrogel precursor solution undergoes polymerization and cross-linking at the inlet of the internal pores of the gradient fiber structure of the tube wall and on the inner surface. After polymerization, the tube is washed and dried to obtain an artificial blood vessel with a hydrogel layer covalently grafted with sulfonic acid groups and cell adhesion peptides on the inner wall. Due to the dense fibers of the gradient structure inner layer, the penetration depth of the hydrogel precursor is controllable, thus forming a composite interface with a large mechanical interlocking depth and strong bonding, rather than a simple surface coating. Preferably, ethanol wettation and hydrogel precursor solution perfusion are completed within 30 minutes after oxygen plasma pretreatment to avoid hydrophobic recovery of surface oxygen-containing groups over time.

[0097] In a preferred embodiment, the monomer containing a sulfonic acid group is 2-acrylamido-2-methylpropanesulfonic acid, sodium vinyl sulfonate, sodium styrene sulfonate, sodium salt of 2-acrylamido-2-methylpropanesulfonic acid, potassium salt of 3-sulfopropyl acrylate, or potassium salt of 3-sulfonate methacrylate. Extracellular matrix-derived peptides include one or more of the following: cell adhesion peptides, aspartic-glycyl-glutamyl-alanine tetrapeptide, and elastin-derived valine-alanyl-prolyl-glycine tetrapeptide; cell adhesion peptides include at least one of the following: arginine-glycine-aspartic acid tripeptide, arginine-glutamic acid-aspartic acid-valine tetrapeptide, tyrosyl-isoleucyl-glycyl-seryl-arginine pentapeptide, and isoleucyl-lysyl-valine-alanyl-valine pentapeptide.

[0098] In a preferred embodiment, the initiator is ammonium persulfate; The crosslinking agent is N,N'-methylenebisacrylamide.

[0099] In a preferred embodiment, after injecting the hydrogel precursor solution into the composite base tube to obtain the hydrogel layer, the process further includes: hydrophilic modification treatment.

[0100] In a more preferred embodiment, the hydrophilic modification treatment involves immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 10 to 20 minutes (e.g., 10 to 30 minutes, 1 minute, 2 minutes, 5 minutes, 10 minutes, 15 minutes, or 20 minutes), preferably 5 to 10 minutes. Specifically, the solid content of the perfluorosulfonic acid solution is 10 wt%.

[0101] In this embodiment of the invention, the treatment can introduce hydrophilic functional groups on the inner surface without damaging the gradient fibrous structure of the base tube, forming a stable hydrophilic layer, effectively reducing platelet adhesion and improving the blood compatibility of the material.

[0102] This invention also provides a wearable expanded polytetrafluoroethylene artificial blood vessel, which consists of a hydrogel layer with a thickness of 5-50 μm and an integrated gradient fiber structure from the inside out; the average fiber diameter of the gradient fiber structure increases from the inside out and the porosity increases from 40%-60% to 70%-85%.

[0103] The wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel prepared in this embodiment of the invention consists of a hydrogel layer with a thickness of 5-50 μm and an integrated gradient fiber structure from the inside out. The average fiber diameter of this integrated gradient fiber structure gradually increases from 50-200 nm to 300-800 nm from the inner surface to the outer surface, and the porosity gradually increases from 40%-60% to 70%-85%. A 5-50 μm thick biomimetic hydrogel layer is firmly bonded to the inner surface. This hydrogel layer is bound to the ePTFE matrix through both covalent bonds and mechanical interlocking. The sulfonic acid groups provide durable anticoagulant function, and the cell adhesion peptides provide selective adhesion sites for endothelial cells. This blood vessel exhibits good radial compliance, matches natural blood vessels, has a burst pressure >3000 mmHg, and a water permeability <200 mL / min / cm². Because the inner layer of the gradient structure is dense and combined with hydrogel, the fiber-hydrogel complex around the needle hole undergoes elastic deformation rather than brittle tearing during puncture. After needle removal, thanks to the synergistic effect of the elastic memory of ePTFE and the swelling properties of hydrogel, rapid self-sealing is achieved (hemostasis time <3 minutes) without delamination.

[0104] The present invention will be further described below by way of examples, but the scope of protection of the present invention is not limited to these embodiments. The present invention may have many other embodiments, and those skilled in the art can make various corresponding changes and modifications based on the present invention without departing from its spirit and essence. However, all such corresponding changes and modifications should fall within the scope of protection of the appended claims. Unless otherwise specified, the experimental methods used in the following embodiments are conventional methods. Unless otherwise specified, the materials, reagents, etc., used in the following embodiments can be obtained commercially or by existing methods.

[0105] Example 1 A method for preparing a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel, comprising: (1) PTFE suspension resin and Isopar G organic solvent were mixed evenly at a mass ratio of 78:22 and aged for 24 hours to obtain a premix; A three-section mandrel is selected. The total length of the working section is 120mm. The inlet section is 35mm long with a cone angle of 5° and a taper change rate of 0.87° / cm; the transition section is 50mm long with a cone angle of 2.5° and a taper change rate of 0.44° / cm; and the outlet section is 35mm long with a cone angle of 1° and a taper change rate of 0.19° / cm. The outer mold sleeve is cylindrical with an inner diameter of 8.5mm. A non-equidistant annular flow channel is formed between the mandrel and the outer mold sleeve, with an inlet gap of 3.0mm and an outlet gap of 2.5mm, to ensure that the inner diameter of the final artificial blood vessel is ≤6mm. The above premixed material is injected into a plunger extruder equipped with the above mandrel and outer die sleeve. The extrusion speed is 2.0 m / min, the extrusion temperature is 45℃, and the extrusion pressure is 15 MPa to obtain a green tube. (2) The obtained green tube was degreased at 230℃ for 90s, with a heating rate of 12℃ / min; The degreased green tube was then subjected to longitudinal stretching at 240℃, with a feeding speed of 2.0 m / min, a stretching speed of 16 m / min, a draw ratio of 8, and a stretching section length of 2.5 m. Then, the longitudinally stretched green tube is heat-set at 330℃ for 45 seconds using a hot air oven. Subsequently, radial expansion was performed at 235℃ with an expansion ratio of 2.5 times and an expansion rate of 35 mm / min. Compressed nitrogen was used as the expansion medium to obtain a gradient-fiberized ePTFE hollow fiber base tube, i.e., the base tube. At this point, the inner diameter of the base tube was 4.0 mm, and the wall thickness was 350 μm. The average fiber diameter on its inner surface was 120 nm, and the porosity was 48%. The average fiber diameter on its outer surface was 550 nm, and the porosity was 78%. (3) The base tube from step (2) is subjected to gradient heat treatment at 80°C and treated for 10 minutes under a fixed state supported by a mandrel to obtain a composite base tube. (4) Low-pressure plasma was introduced into the inner cavity of the composite base tube to perform oxygen plasma pretreatment on the inner surface. The treatment conditions were as follows: pure oxygen was used, the oxygen flow rate was 20 sccm, the working pressure was 35 Pa, the radio frequency power was 50 W, the treatment time was 60 s, and the substrate temperature was controlled below 50 ℃. Among them, after treatment, the water contact angle of the inner surface decreased from 130° to 42°, the activation depth was 3~5 μm, and the water contact angle of the outer surface was maintained above 120°. Dissolve 12g of 2-acrylamide-2-methylpropanesulfonic acid (AMPS) and 0.15g of acrylamide-arginine-glycine-aspartic acid tripeptide (Acr-PEG6-REDV) in 100mL of deionized water and stir well to obtain component A; dissolve 0.5g of ammonium persulfate (APS) and 0.15g of N,N'-methylenebisacrylamide (MBA) in 100mL of deionized water and stir well to obtain component B; A hydrogel liner was constructed using dynamic pressure circulation perfusion technology, including: sealing one end of a composite base tube pretreated with oxygen plasma and connecting the other end to a peristaltic pump. The inner wall of the tube was first wetted with anhydrous ethanol for 30 seconds, and then the ethanol was drained. Component A and component B were then mixed uniformly at a volume ratio of 7:1 to obtain a hydrogel precursor solution, which was immediately injected into the tube lumen and circulated at a flow rate of 2 mL / min for 20 minutes at a constant temperature of 37°C. After polymerization, the lumen was rinsed with a large amount of deionized water for 30 minutes and then vacuum dried at 40°C for 2 hours. At this point, the hydrogel layer thickness was 15 μm, resulting in a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel with an inner diameter of 4 mm.

[0106] Example 2 Example 2 is basically the same as Example 1, except that the post-processing method used is different.

[0107] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 80°C and treated for 30 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0108] Example 3 Example 3 is basically the same as Example 1, except that the post-processing method used is different.

[0109] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 300°C and treated for 10 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0110] Example 4 Example 4 is basically the same as Example 1, except that the post-processing method used is different.

[0111] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 300°C and treated for 30 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0112] Example 5 A method for preparing a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel, comprising: (1) PTFE suspension resin and Isopar M organic solvent were mixed evenly at a mass ratio of 85:15 and aged for 36 hours to obtain a premix; A three-section mandrel is selected. The total length of the working section is 150mm. The inlet section is 45mm long with a cone angle of 5° and a taper change rate of 0.67° / cm; the transition section is 60mm long with a cone angle of 2.5° and a taper change rate of 0.33° / cm; and the outlet section is 45mm long with a cone angle of 1° and a taper change rate of 0.13° / cm. The outer mold sleeve is cylindrical with an inner diameter of 12.5mm. A non-equidistant annular flow channel is formed between the mandrel and the outer mold sleeve, with an inlet gap of 4.5mm and an outlet gap of 3.5mm, to ensure that the inner diameter of the final artificial blood vessel is ≤6mm. The above premixed material is injected into a plunger extruder equipped with the above mandrel and outer die sleeve. The extrusion speed is 1.5 m / min, the extrusion temperature is 45℃, and the extrusion pressure is 15 MPa to obtain a green tube. (2) Degrease the obtained green tube at 260℃ for 70s, with a heating rate of 5℃ / min; The degreased green tube was then longitudinally stretched at 260°C, with a feeding speed of 1.5 m / min, a stretching speed of 9 m / min, a draw ratio of 6, and a stretching section length of 2.5 m. Then, the longitudinally stretched green tube is heat-set at 250°C for 120 seconds using a hot air oven. Subsequently, radial expansion was performed at 260℃ with an expansion ratio of 2.0 times and an expansion rate of 35 mm / min. Compressed nitrogen was used as the expansion medium to obtain a gradient-fiberized ePTFE hollow fiber base tube, i.e., the base tube. At this point, the inner diameter of the base tube was 6.0 mm, and the wall thickness was 400 μm. The average fiber diameter on its inner surface was 150 nm, and the porosity was 50%. The average fiber diameter on its outer surface was 600 nm, and the porosity was 80%. (3) The base tube from step (2) is subjected to gradient heat treatment at 80°C and treated for 10 minutes under a fixed state supported by a mandrel to obtain a composite base tube. (4) Low-pressure plasma was introduced into the inner cavity of the composite base tube to perform oxygen plasma pretreatment on the inner surface. The treatment conditions were as follows: pure oxygen was used, the oxygen flow rate was 30 sccm, the working pressure was 50 Pa, the radio frequency power was 20 W, the treatment time was 120 s, and the substrate temperature was controlled below 50 ℃. Among them, after treatment, the water contact angle of the inner surface decreased from 130° to 45°, the activation depth was 3~5 μm, and the water contact angle of the outer surface remained above 120°. Dissolve 5g AMPS and 0.05g Acr-PEG6-REDV in 100mL of deionized water and stir well to obtain component A; dissolve 0.1g APS and 0.05g MBA in 100mL of deionized water and stir well to obtain component B; A hydrogel liner was constructed using dynamic pressure circulation perfusion technology, including: sealing one end of a composite base tube pretreated with oxygen plasma and connecting the other end to a peristaltic pump. The inner wall of the tube was first wetted with anhydrous ethanol for 30 seconds, and then the ethanol was drained. Component A and component B were then mixed uniformly at a volume ratio of 7:1, and the mixture was immediately injected into the tube cavity. The solution was circulated at a constant temperature of 37°C and a flow rate of 2 mL / min for 30 minutes. After polymerization, the tube cavity was flushed with a large amount of deionized water for 30 minutes and then vacuum dried at 40°C for 2 hours. At this point, the hydrogel layer thickness was 18 μm, yielding a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel with an inner diameter of 6 mm.

[0113] Example 6 Example 6 is basically the same as Example 5, except that the post-processing method used is different.

[0114] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 80°C and treated for 30 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0115] Example 7 Example 7 is basically the same as Example 5, except that the post-processing method used is different.

[0116] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 300°C and treated for 10 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0117] Example 8 Example 8 is basically the same as Example 5, except that the post-processing method used is different.

[0118] Specifically, (3) the base tube from step (2) is subjected to gradient heat treatment at 300°C and treated for 30 minutes in a fixed state supported by a mandrel to obtain a composite base tube.

[0119] Example 9 Example 9 is basically the same as Example 1, except that the post-processing method used is different.

[0120] Specifically, (3) the base tube is fixed on a rotating fixture and the expanded polytetrafluoroethylene (ePTFE) membrane is evenly wrapped around the outer surface of the base tube at a winding angle of 60°. After the wrapping is completed, the wrapped base tube is placed at 360°C for sintering for 5 minutes. After sintering is completed, it is naturally cooled to room temperature (for example, 25°C) to obtain a composite base tube. The single layer thickness of the ePTFE membrane is 25μm, the width is 8mm, and the total wrapping thickness is 100μm, that is, 4 layers are wrapped.

[0121] Example 10 Example 10 is basically the same as Example 9, except that the sintering time of the wrapping sintering reinforcement treatment is different.

[0122] Specifically, (3) the base tube is fixed on a rotating fixture and the expanded polytetrafluoroethylene (ePTFE) membrane is evenly wrapped around the outer surface of the base tube at a winding angle of 60°. After the wrapping is completed, the wrapped base tube is placed at 360°C for sintering for 10 minutes. After the sintering is completed, it is naturally cooled to room temperature (for example, 25°C) to obtain a composite base tube. The single layer thickness of the ePTFE membrane is 25μm, the width is 8mm, and the total wrapping thickness is 100μm, that is, 4 layers are wrapped.

[0123] Example 11 Example 11 is basically the same as Example 9, except that the sintering time of the wrapping sintering reinforcement treatment is different.

[0124] Specifically, (3) the base tube is fixed on a rotating fixture and the expanded polytetrafluoroethylene (ePTFE) membrane is evenly wrapped around the outer surface of the base tube at a winding angle of 60°. After the wrapping is completed, the wrapped base tube is placed at 360°C for sintering for 15 minutes. After sintering is completed, it is naturally cooled to room temperature (for example, 25°C) to obtain a composite base tube. The single layer thickness of the ePTFE membrane is 25μm, the width is 8mm, and the total wrapping thickness is 100μm, that is, 4 layers are wrapped.

[0125] Example 12 Example 12 is basically the same as Example 1, except that step (4) also includes hydrophilic modification treatment.

[0126] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 10 seconds, removing it after immersion, rinsing the surface with deionized water to remove any residual solution, and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0127] Example 13 Example 13 is basically the same as Example 12, except that the hydrophilic modification treatment time in step (4) is different.

[0128] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 1 minute, removing it after immersion, rinsing the surface with deionized water to remove any residual solution, and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0129] Example 14 Example 14 is basically the same as Example 12, except that the hydrophilic modification treatment time in step (4) is different.

[0130] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 5 minutes; removing it after immersion; rinsing the surface with deionized water to remove any residual solution; and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0131] Example 15 Example 15 is basically the same as Example 12, except that the hydrophilic modification treatment time in step (4) is different.

[0132] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 10 minutes; removing it after immersion; rinsing the surface with deionized water to remove any residual solution; and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0133] Example 16 Example 16 is basically the same as Example 12, except that the hydrophilic modification treatment time in step (4) is different.

[0134] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 15 minutes; removing it after immersion; rinsing the surface with deionized water to remove any residual solution; and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0135] Example 17 Example 17 is basically the same as Example 12, except that the hydrophilic modification treatment time in step (4) is different.

[0136] Specifically, step (4) further includes: immersing the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 20 minutes; removing it after immersion; rinsing the surface with deionized water to remove any residual solution; and finally vacuum drying at 60°C for 1 hour to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The perfluorosulfonic acid solution has a solid content of 10 wt%, and the solvent consists of isopropanol and deionized water in a volume ratio of 1:1.

[0137] Example 18 Example 18 is basically the same as Example 10, except that the post-processing method is different.

[0138] Specifically, (3) the base tube is fixed on a rotating clamp and the expanded polytetrafluoroethylene (ePTFE) membrane is evenly wrapped around the outer surface of the base tube at a winding angle of 60°. After the wrapping is completed, the wrapped base tube is placed at 360°C for sintering for 10 minutes. After sintering is completed, it is naturally cooled to room temperature (for example, 25°C). The single layer thickness of the ePTFE membrane is 25μm, the width is 8mm, and the total wrapping thickness is 100μm, that is, 4 layers are wrapped. Then, the base tube that has been reinforced by wrapping and sintering is placed in a fixed state (supported by a mandrel) and heat-treated at 250°C for 20 minutes. After completion, it is allowed to cool naturally to room temperature (e.g., 25°C). (4) Using the same oxygen plasma pretreatment and dynamic pressure cyclic perfusion of the hydrogel precursor solution as in Example 10, the inner surface of the composite base tube containing the hydrogel layer was immersed in perfluorosulfonic acid solution for 10 min. After immersion, it was removed, and the surface residual solution was rinsed with deionized water. Finally, it was vacuum dried at 60°C for 1 h to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel. The solid content of the perfluorosulfonic acid solution was 10 wt%, and the solvent was composed of isopropanol and deionized water in a volume ratio of 1:1.

[0139] The wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessels prepared in the above embodiments were used as samples for the following performance tests, and the test results are shown in Table 1. All performance tests mentioned below were performed in accordance with relevant international and domestic standards: Burst pressure: Tested in physiological saline at 37°C, according to ASTM F2399-07(2019); Water permeability: The permeability per unit time was measured at a pressure of 16 kPa, in accordance with ASTM F2399-07 (2019). Radial compliance: determined using a pulsating flow circulation system and laser diffraction method, in accordance with ISO 7198:2016; In vitro platelet adhesion rate: According to GB / T 16886.4-2022, platelet-rich plasma incubation and LDH method were used for quantitative analysis; Endothelial cell proliferation rate: determined using HUVECs cell line and CCK-8 assay according to ISO 10993-5:2009; Puncture hemostasis time: determined in a simulated blood circulation circuit according to ASTM F3036-13 (2019); Layering after puncture: According to ASTM F2477-07 (2019), the layering was determined by macroscopic and SEM microscopic observation after 100 punctures.

[0140] Table 1 As shown in Examples 1 to 4, the data indicates that with increasing heat treatment temperature, the burst pressure of the artificial blood vessel increases, water permeability decreases, radial compliance slightly decreases, platelet adhesion rate decreases, and puncture hemostasis time shortens, without any stratification. The sample treated at 80℃ exhibits better radial compliance and endothelial cell proliferation rate, while the sample treated at 300℃ shows superior burst pressure and antithrombotic properties. For Examples 1 to 4 and Examples 5 to 8, compared to the artificial blood vessel with an inner diameter of 4mm, the 6mm inner diameter artificial blood vessel shows a slightly lower burst pressure, slightly higher water permeability, and improved radial compliance. This is mainly due to the reduced circumferential stress on the vessel wall after the increase in diameter. For Examples 5 to 8, the trends of other properties are consistent with increasing heat treatment temperature.

[0141] As shown in Examples 9 to 11, after the wrapping and sintering reinforcement treatment, the burst pressure of the artificial blood vessel was significantly improved compared to Example 1. Furthermore, with the extension of sintering time, the burst pressure further increased, water permeability decreased, antithrombotic performance improved, and puncture hemostasis time shortened. Specifically, the sample sintered for 15 minutes achieved a burst pressure of 3900 mmHg and a puncture hemostasis time shortened to 1.5 minutes, exhibiting optimal comprehensive mechanical properties and puncture self-sealing performance. However, excessively long sintering time (15 minutes) can lead to a decrease in radial compliance and a reduction in outer layer porosity, potentially affecting tissue ingrowth. Considering all factors, sintering for 10 minutes is the preferred process condition, which can achieve significantly enhanced anti-puncture delamination ability while maintaining good compliance and tissue integration potential.

[0142] For Examples 12 to 17, the water contact angles of the inner surfaces of the ready-to-wear expanded polytetrafluoroethylene (ePTFE) artificial blood vessels after hydrophilic modification were 68°, 52°, 38°, 32°, 28°, and 26°, respectively. From the above data, it can be seen that, except for Examples 12 and 13, the hydrophilicity of the inner surface of the base tube was significantly improved after hydrophilic modification with perfluorosulfonic acid solution, and the water contact angle further decreased from 42° (hydrogel layer only) in Example 1 to 26°. Furthermore, with prolonged immersion time, the hydrophilicity of the inner surface continued to increase, the platelet adhesion rate decreased accordingly, and the antithrombotic performance improved. However, after immersion time exceeded 10 minutes, the improvement in the hydrophilicity of the inner surface tended to plateau, while the burst pressure slightly decreased and the water permeability slightly increased. This may be due to the permeation effect of the perfluorosulfonic acid solution having a certain influence on the interfacial bonding between the hydrogel layer and the ePTFE matrix. Taking into account antithrombotic properties, mechanical strength, and process efficiency, a soaking time of 5-10 minutes is the optimal process window to obtain excellent antiplatelet adhesion ability without significantly sacrificing mechanical properties.

[0143] Comparing Examples 1, 10, 15, and 18, it is evident that the overall performance of the artificial blood vessel prepared in Example 18 was further optimized after combined treatment with gradient heat treatment, wrapping sintering reinforcement treatment, and hydrophilic modification treatment. Specifically, compared to Example 1, the burst pressure was significantly increased (from 3250 mmHg to 3880 mmHg), the platelet adhesion rate was significantly reduced (from 8.5% to 3.0%), and the puncture hemostasis time was shortened to 1.5 minutes. Compared to a single post-treatment method, the combined treatment achieved the lowest platelet adhesion rate while maintaining a high burst pressure and low water permeability, exhibiting optimal antithrombotic performance and puncture self-sealing performance. This indicates a synergistic effect among the three treatment methods: wrapping sintering reinforcement treatment improved the strength of the outer structure and the resistance to puncture delamination; gradient heat treatment regulated and optimized the dimensional stability of the vessel wall fiber network; and perfluorosulfonic acid hydrophilic modification treatment further improved the blood compatibility of the inner surface.

[0144] Comparative Example 1 Comparative Example 1 is basically the same as Example 1, except that the mold used is different.

[0145] Specifically, (1) PTFE suspension resin and Isopar G organic solvent are mixed evenly at a mass ratio of 78:22 and aged for 24 hours to obtain a premix; A traditional uniform mold is selected, with a cylindrical mandrel with a diameter of 4.0 mm and a cylindrical outer mold sleeve with an inner diameter of 8.5 mm. Equidistant annular flow channels are formed between the mandrel and the outer mold sleeve, with a gap of 2.5 mm. The premixed material is injected into a plunger extruder equipped with the mandrel and outer die sleeve. The extrusion speed is 2.0 m / min, the extrusion temperature is 45℃, and the extrusion pressure is 15 MPa to obtain a green tube.

[0146] Thus, after the same steps (2) as in Example 1, the resulting base tube has an inner diameter of 4.0 mm and a wall thickness of 350 μm. Due to the use of a conventional uniform mold, there is no significant difference in fiber morphology between the inner and outer surfaces of the tube wall. The average fiber diameter is 350–450 nm, and the porosity is 60–65%, which does not have the gradient structure characteristics of dense inner and sparse outer.

[0147] Comparative Example 2 According to the method described in the embodiment of Chinese Patent CN119388815A, a dense ePTFE intermediate layer (porosity of about 30%~35%), a porous ePTFE inner layer, and a porous ePTFE outer layer (porosity of about 70%~80%) are prepared respectively. The inner layer, intermediate layer, and outer layer are sequentially assembled or stacked to obtain a composite cylinder. After pressing and unidirectional stretching, a three-layer composite ePTFE artificial blood vessel base tube with an inner diameter of 4.0 mm and a total wall thickness of about 360 μm is obtained. The three-layer composite ePTFE artificial blood vessel base tube was pretreated with oxygen plasma under the same conditions as in step (4) of Example 1, and then a hydrogel layer was formed on the inner surface using the same hydrogel precursor solution as in step (4) of Example 1 to obtain the final composite artificial blood vessel.

[0148] Comparative Example 3 Comparative Example 3 is basically the same as Example 1, except that the gradient heat treatment in step (3) was not performed. Specifically, the base tube from step (2) of Example 1 is directly applied to step (4), that is, after the base tube is pretreated with oxygen plasma under the same conditions, the same hydrogel precursor solution is injected into the base tube to obtain a wearable expanded polytetrafluoroethylene artificial blood vessel.

[0149] Comparative Example 4 Comparative Example 4 is basically the same as Example 1, except that the composition of the premix is ​​different.

[0150] Specifically, in step (1), PTFE suspension resin and Isopar G organic solvent are mixed evenly at a mass ratio of 90:10 and aged for 24 hours to obtain a premix.

[0151] Comparative Example 5 Comparative Example 5 is basically the same as Example 1, except that the composition of the premix is ​​different.

[0152] Specifically, in step (1), PTFE suspension resin and Isopar G organic solvent are mixed evenly at a mass ratio of 60:40 and aged for 24 hours to obtain a premix.

[0153] Comparative Example 6 Comparative Example 6 is basically the same as Example 1, except that the extrusion speed is different.

[0154] Specifically, in step (1), the extrusion speed is 0.3 m / min, and in step (2), the feeding speed of the green tube is also 0.3 m / min.

[0155] Comparative Example 7 Comparative Example 7 is basically the same as Example 1, except that the extrusion temperature is different.

[0156] Specifically, in step (1), the extrusion temperature is 70°C.

[0157] Comparative Example 8 Comparative Example 8 is basically the same as Example 1, except that the extrusion pressure is different.

[0158] Specifically, in step (1), the extrusion pressure is 35 MPa.

[0159] Comparative Example 9 Comparative Example 9 is basically the same as Example 1, except that the composition of the hydrogel precursor solution is different.

[0160] Specifically, in step (4), 2g of 2-acrylamide-2-methylpropanesulfonic acid (AMPS) and 0.15g of acrylamide-arginine-glycine-aspartic acid tripeptide (Acr-PEG6-REDV) are dissolved in 100mL of deionized water and stirred evenly to obtain component A.

[0161] Comparative Example 10 Comparative Example 10 is basically the same as Example 1, except that the composition of the hydrogel precursor solution is different.

[0162] Specifically, in step (4), 25g of 2-acrylamide-2-methylpropanesulfonic acid (AMPS) and 0.15g of acrylamide-arginine-glycine-aspartic acid tripeptide (Acr-PEG6-REDV) are dissolved in 100mL of deionized water and stirred evenly to obtain component A.

[0163] Comparative Example 11 Comparative Example 11 is basically the same as Example 1, except that the composition of the hydrogel precursor solution is different.

[0164] Specifically, in step (4), 12g of 2-acrylamide-2-methylpropanesulfonic acid (AMPS) and 0.001g of acrylamide-arginine-glycine-aspartic acid tripeptide (Acr-PEG6-REDV) are dissolved in 100mL of deionized water and stirred evenly to obtain component A.

[0165] Comparative Example 12 Comparative Example 12 is basically the same as Example 1, except that the composition of the hydrogel precursor solution is different.

[0166] Specifically, in step (4), 12g of 2-acrylamide-2-methylpropanesulfonic acid (AMPS) and 1g of acrylamide-arginine-glycine-aspartic acid tripeptide (Acr-PEG6-REDV) are dissolved in 100mL of deionized water and stirred evenly to obtain component A.

[0167] The artificial blood vessels prepared in the above comparative model were used as samples for the above performance tests, and the test results are shown in Table 2.

[0168] Table 2 Compared with Example 1, Comparative Example 1 exhibits poorer mechanical properties: the burst pressure is 2980 mmHg, lower than the 3250 mmHg of Example 1, and the water permeability is higher (210 mL / min·cm²), indicating insufficient tube wall density; poor blood compatibility: the platelet adhesion rate is 12.5%, significantly higher than the 8.5% of Example 1, because the tube wall does not have a gradient structure with a dense inner surface and a sparse outer surface, and the inner surface is not dense and smooth enough, which is not conducive to the uniform coverage and stable adhesion of the hydrogel liner; weak endothelialization promotion ability: the endothelial cell proliferation rate is 78%, lower than the 92% of Example 1, indicating that the uniform tube wall has poor support and nutrient transport capabilities for the hydrogel liner; poor puncture self-sealing performance: the puncture hemostasis time is 4.5 min, much longer than the 2.8 min of Example 1, and after 100 repeated punctures, local peeling of the hydrogel layer and microcracks in the tube wall appear, indicating that the traditional uniform structure cannot effectively resist the interfacial stress caused by repeated punctures.

[0169] The bonding ability of the hydrogel layer of the artificial blood vessel was evaluated using a combination of visual observation and ultrasound. It was observed that the hydrogel layer bonding of Comparative Example 2 was generally poor, with visible micro-gaps at the interface, while the hydrogel layer of Example 1 was firmly bonded with no interface gaps. Compared to Example 1, Comparative Example 2 exhibited a potential interlayer interface problem: the three-layer composite structure relied on hot-pressing sintering for interlayer bonding. Under 100 repeated punctures and cyclic filling-depressurization, micro-cracks appeared at the interlayer interface, eventually developing into local delamination. This demonstrates that multi-layer physical composites cannot completely eliminate the risk of interlayer delamination. In contrast, Example 1, through mold flow field design, directly formed a gradient structure with continuously varying fiber density and orientation on a single tube wall, fundamentally eliminating the interlayer interface. The puncture self-sealing performance was also poor: the puncture hemostasis time of Comparative Example 2 was 4.2 minutes, longer than the 2.8 minutes of Example 1. This was due to the difference between the intermediate layer and the... The inner and outer layers have mismatched moduli, resulting in inconsistent deformation around the needle hole and inconsistent rebound after needle removal, which reduces the self-sealing efficiency. The hydrogel layer has weak bonding strength: the inner surface of the three-layer composite tube is a uniform porous ePTFE, which is not as dense and smooth as the inner layer of the gradient structure in Example 1. In Example 1, the hydrogel layer and the matrix mainly rely on physical interlocking after surface plasma activation, and the mechanical interlocking depth is shallow. After 100 punctures, the edge of the hydrogel layer can be seen to lift up. Blood compatibility and endothelialization promotion are slightly inferior: the platelet adhesion rate of Comparative Example 2 is higher, and the endothelial cell proliferation rate is lower, which is related to the insufficient density of the inner surface and the insufficient bonding strength of the hydrogel.

[0170] Meanwhile, the dry-wet diameter change rate of the artificial blood vessels in Example 1 and Comparative Example 3 was tested. The dry-wet diameter change rate of Example 1 was only 2.1%, while that of Comparative Example 3 was 4.8%. Based on Table 2, it can be seen that Comparative Example 3 has insufficient mechanical properties compared to Example 1: the burst pressure of Comparative Example 3 is lower than that of Example 1; the water permeability is higher than that of Example 1, indicating that the untreated ePTFE fiber network is not yet sufficiently stable, the wall density is poor, and the mechanical strength is low; the dimensional stability is poor: the diameter change rate of Comparative Example 3 during the dry-wet transition reaches 4.8%, much higher than that of Example 1. This indicates that gradient heat treatment can effectively lock the fiber-node structure, reduce the diameter change caused by hydrogel swelling and blood flow impact, and ensure the dimensional stability of the artificial blood vessel in vivo; the puncture self-sealing performance is moderate: the puncture hemostasis time of Comparative Example 3 is 3.5 min, slightly higher than that of Example 1, indicating the swelling compensation ability of the hydrogel before heat treatment. The results were acceptable, but due to the insufficient density of the tube wall matrix structure and the inadequate elastic recoil force around the needle hole, the hemostasis time was still longer than that of the heat-treated sample. The endothelialization promotion ability was excellent: the endothelial cell proliferation rate of Comparative Example 3 was slightly higher than that of Example 1. This is because the tube wall porosity was higher and the pore size was larger when not heat-treated, which facilitated the transport of nutrients to the hydrogel liner and the discharge of metabolic waste, thereby promoting the growth of endothelial cells. However, this high porosity came at the cost of sacrificing mechanical strength and dimensional stability. The antithrombotic performance was moderate: the platelet adhesion rate of Comparative Example 3 was higher than that of Example 1, indicating that the inner surface was not dense enough when not heat-treated, and the ePTFE fiber structure under the hydrogel layer did not provide smooth enough support for the blood contact interface, resulting in increased platelet adhesion.

[0171] As shown in Table 2, due to the organic solvent content being less than 15 wt% in Comparative Example 4, the excessively high viscosity of the premix resulted in poor extrusion flowability, significant melt fracture, and numerous longitudinal microcracks on the surface of the green tube. The molding yield was only 40%, and the extrusion pressure exceeded the equipment's tolerance range, significantly increasing the risk of mold wear. After degreasing, the radial shrinkage rate of the green tube reached 8.5%, exhibiting extremely poor dimensional stability. During subsequent stretching, uneven molecular chain orientation prevented the formation of a continuous and uniform gradient fiber structure, leading to frequent breakage and directly resulting in a decrease in burst pressure and an increase in water permeability in the finished product. The rough inner surface also increased the probability of platelet adhesion, and the hydrogel was difficult to adhere uniformly to the defective surface, resulting in a significant decrease in bonding strength. After puncture, hydrogel peeling and tube wall cracks were prone to occur, and the hemostasis time was significantly prolonged. In Comparative Example 5, due to the organic solvent content exceeding 35 wt%, the premix was too thin, resulting in poor moldability and excessive fluidity. During extrusion, the green tube was prone to edge collapse and uneven wall thickness, with a molding yield of only 35%. During degreasing, excessive solvent evaporated rapidly, causing numerous honeycomb-like micropores to appear on the surface of the green tube. The subsequent tensile breakage rate reached 65%, and the yield was less than 20%, with severely deteriorated mechanical properties of the finished product. At the same time, the loose matrix structure led to excessive swelling of the hydrogel and extremely poor bonding. Numerous micropore defects on the inner surface became platelet retention sites, significantly reducing blood compatibility. After puncture, the tube wall may even perforate, completely losing its usability.

[0172] As shown in Table 2, in Comparative Example 6, the premix remained in the mold cavity for too long during extrusion, leading to premature evaporation of local organic solvents. This resulted in rough textures on the surface of the green tube, with a wall thickness deviation of ±12% (±3% in Example 1). During subsequent stretching, insufficient molecular chain orientation and uneven longitudinal strength resulted in a breakage rate of 45%, failing to form the expected gradient fibrous structure. The rough surface and insufficient inner layer density of the green tube directly led to a decrease in the mechanical properties of the finished product and an increase in water permeability. The rough inner surface also increased the probability of platelet adhesion, and the hydrogel was difficult to adhere evenly to the defective surface, resulting in a significant decrease in bonding strength. After puncture, hydrogel peeling and tube wall cracks were prone to occur, and the hemostasis time was significantly prolonged. In Comparative Example 7, the extrusion temperature was too high, causing partial vaporization of the organic solvent, increasing the range of pressure fluctuations in the mold cavity, and resulting in a large number of bubbles with a diameter of 5~20μm inside the green tube. At the same time, the PTFE particles underwent slight pre-sintering due to overheating, increasing the fiber knot breakage rate by 60% during subsequent stretching, and causing hidden microcracks in the finished tube wall. Furthermore, the pores formed in the tube wall after the bubbles burst become stress concentration points and retention sites for bacteria / platelets, which not only significantly reduce the burst pressure but also cause the hydrogel layer to lose its bonding force at the edge of the pores. The increased brittleness of the pre-sintered area prevents the tube wall from elastically deforming to close the pinhole after puncture, prolonging the hemostasis time and greatly increasing the risk of delamination after repeated punctures. In Comparative Example 8, the excessive extrusion pressure led to aggravated wear on the inner wall of the mold cavity, and the coaxiality deviation between the mandrel and the outer mold sleeve increased from 0.02 mm to 0.08 mm, resulting in increased ellipticity of the green tube. At the same time, the premixed material underwent local degradation due to high-pressure heat generation, and stress concentration occurred in the degradation area during subsequent stretching, resulting in a 28% decrease in the burst pressure of the finished product.

[0173] Tests showed that the thicknesses of the hydrogel layers in Comparative Examples 9 to 12 were 12.5 μm, 55.3 μm, 14.8 μm, and 16.2 μm, respectively. SEM was used to observe the morphology of the hydrogel layers in Comparative Examples 9 to 12 after 100 punctures. Comparative Example 9 showed a smooth surface and complete coverage; Comparative Example 10 showed a cracked surface with numerous macroscopic pores; Comparative Example 11 showed a discontinuous island-like distribution with a coverage of approximately 40%; Comparative Example 12 showed a smooth surface, but a large number of apoptotic cells were observed after cell culture. Furthermore, as shown in Table 2, this invention strictly limits component A to include monomers containing sulfonic acid groups, cell adhesion peptides, and deionized water in a mass ratio of (5~20):(0.01~0.5):100. This is to precisely balance the relationship between antithrombotic properties (sufficient sulfonic acid groups), endothelialization promotion (appropriate amount of adhesion peptides), and mechanical stability (suitable gel network structure).

[0174] The parts of this invention not described in detail are techniques known to those skilled in the art.

[0175] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention, and not to limit them; although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art should understand that modifications can still be made to the technical solutions described in the foregoing embodiments, or equivalent substitutions can be made to some of the technical features; and these modifications or substitutions do not cause the essence of the corresponding technical solutions to deviate from the spirit and scope of the technical solutions of the embodiments of the present invention.

Claims

1. A method for preparing a wearable expanded polytetrafluoroethylene (ePTFE) artificial blood vessel, characterized in that, include: (1) A premix containing polytetrafluoroethylene and an organic solvent is added to an extruder to obtain a green tube; wherein the extruder uses a mandrel with a gradient tapered shape and a cylindrical outer die sleeve, and the taper of the inlet section of the working section of the mandrel is greater than the taper of the outlet section. (2) The green tube is subjected to degreasing, longitudinal stretching, heat setting and radial expansion in sequence to obtain the base tube; (3) The base tube is post-processed to obtain a composite base tube; (4) After oxygen plasma pretreatment of the inner surface of the composite base tube, the hydrogel precursor solution is injected into the composite base tube to obtain the wearable expanded polytetrafluoroethylene artificial blood vessel.

2. The preparation method according to claim 1, characterized in that, In step (1): The organic solvent content in the premix is ​​15wt%~35wt%; and / or, The organic solvent is a hydrocarbon solvent, preferably n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, or methylcyclohexane.

3. The preparation method according to claim 1, characterized in that, In step (1): The working section of the mandrel adopts a continuously adjustable taper or a segmented taper to form a non-equidistant annular flow channel between the mandrel and the outer mold sleeve; the working section includes an inlet section, a transition section and an outlet section; Preferably, the cone angle of the inlet section ranges from 2° to 8°, and the cone angle of the outlet section ranges from 0.5° to 2°.

4. The preparation method according to claim 3, characterized in that, In step (1): When the taper of the working section is continuously adjustable, the taper variation rate of the working section along its length is 0.09~1.4° / cm; or, When the working section adopts a segmented taper, the taper difference between the inlet section and the transition section is 1°~4°, preferably 2°~3°; the taper difference between the transition section and the outlet section is 0.5°~2°, preferably 1°~1.5°.

5. The preparation method according to claim 1, characterized in that, In step (2): The degreasing temperature is 150~280℃, preferably 200~260℃; the time is 30~180s, preferably 60~120s; the heating rate is 5~20°C / min; and / or, The longitudinal stretching speed is 2~50m / min; the draw ratio is 2~10 times, preferably 3~6 times; the stretching temperature is 200~260℃, preferably 220~250℃.

6. The preparation method according to claim 1, characterized in that, In step (2): The heat setting temperature is 250~380℃, preferably 300~360℃; the time is 10~120s, preferably 30~60s; and / or, The radial expansion temperature is 200~260℃, preferably 220~250℃; the expansion ratio is 1.5~5.0 times, preferably 2.0~3.5 times; and the expansion speed is 10~100mm / min, preferably 20~50mm / min.

7. The preparation method according to claim 1, characterized in that, In step (3): The post-treatment is at least one of gradient heat treatment or wrapping sintering reinforcement treatment; Preferably, the gradient heat treatment is performed at a temperature of 80~300℃ for a time of 10~30 min; Preferably, the wrapping and sintering reinforcement treatment involves uniformly wrapping an expanded polytetrafluoroethylene (ePTFE) film around the outer surface of the base tube, and then sintering the wrapped base tube at 360°C for 5-15 minutes; more preferably, the ePTFE film has a thickness of 2-30 μm, a width of 3-15 mm, and a wrapping angle of 45°-75°.

8. The preparation method according to any one of claims 1 to 7, characterized in that, In step (4): The oxygen plasma pretreatment uses a gas flow rate of 10-30 sccm, preferably 15-20 sccm; a radio frequency power of 20-80 W, preferably 30-50 W; a working gas pressure of 10-50 Pa, preferably 30-40 Pa; and a time of 30-120 s; and / or, The hydrogel precursor solution comprises component A and component B in a volume ratio of (4~9):1; component A comprises monomers containing sulfonic acid groups, extracellular matrix-derived peptides, and deionized water in a mass ratio of (5~20):(0.01~0.5):100; component B comprises an initiator, a crosslinking agent, and deionized water in a mass ratio of (0.1~1):(0.01~0.5):

100. Preferably, the infusion is performed using dynamic pressure circulation infusion at a temperature of 35~40℃, and the circulation time is 10~60 minutes under conditions lower than the ethanol bubble point pressure.

9. The preparation method according to claim 8, characterized in that, The monomer containing the sulfonic acid group is 2-acrylamido-2-methylpropanesulfonic acid, sodium vinyl sulfonate, sodium styrene sulfonate, sodium salt of 2-acrylamido-2-methylpropanesulfonic acid, potassium salt of 3-sulfopropyl acrylate, or potassium salt of 3-sulfonate methacrylate. The extracellular matrix-derived peptides include one or more of the following: cell adhesion peptides, aspartic-glycyl-glutamyl-alanine tetrapeptide, and elastin-derived valine-alanyl-prolyl-glycine tetrapeptide; the cell adhesion peptides include at least one of the following: arginine-glycine-aspartic acid tripeptide, arginine-glutamic acid-aspartic acid-valine tetrapeptide, tyrosyl-isoleucyl-glycyl-seryl-arginine pentapeptide, and isoleucyl-lysyl-valine-alanyl-valine pentapeptide; And / or, The initiator is ammonium persulfate; The crosslinking agent is N,N'-methylenebisacrylamide; Preferably, after injecting the hydrogel precursor solution into the composite base tube to obtain the hydrogel layer, the method further includes: hydrophilic modification treatment; more preferably, the hydrophilic modification treatment is to immerse the inner surface of the composite base tube containing the hydrogel layer in a perfluorosulfonic acid solution for 10s to 20min.

10. A wearable expanded polytetrafluoroethylene artificial blood vessel prepared by the preparation method according to any one of claims 1 to 9, characterized in that, The wearable expanded polytetrafluoroethylene artificial blood vessel consists of a hydrogel layer with a thickness of 5~50μm from the inside out and an integrated gradient fiber structure; the average fiber diameter of the gradient fiber structure increases from the inside out and the porosity increases from 40%~60% to 70%~85%.

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