A composite type expanded polytetrafluoroethylene artificial blood vessel and a method for preparing the same

Abstract

The application provides a composite expanded polytetrafluoroethylene artificial blood vessel and a preparation method thereof, and comprises the following steps: (1) adding inner layer premixing material containing polytetrafluoroethylene and an organic solvent and outer layer premixing material containing polytetrafluoroethylene, an organic solvent, magnetic particles and an elastic polymer precursor into an extruder for co-extrusion to obtain a green pipe; (2) sequentially performing axial stretching, sintering and shaping and secondary axial stretching on the green pipe to obtain a base pipe; (3) using a spinning solution containing polytetrafluoroethylene and a spinning aid to perform electrospinning on the base pipe as a receiving device to obtain a composite base pipe; and (4) sequentially loading a polydopamine layer and an anticoagulant drug layer on the inner surface of the composite base pipe after plasma pretreatment to obtain the composite expanded polytetrafluoroethylene artificial blood vessel. The artificial blood vessel provided by the application has excellent biomimetic mechanical properties, long-acting double anticoagulation function, rapid endothelialization capacity and wireless stenosis monitoring function.

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 composite 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. However, for small-diameter artificial blood vessels with an inner diameter ≤6mm, ePTFE still faces three core challenges: mechanical property mismatch, thrombosis and endothelialization obstacles, and difficulties in postoperative monitoring. Regarding mechanical property mismatch, traditional uniaxially expanded ePTFE is highly rigid, lacking the toe region and high compliance of natural blood vessels, easily leading to anastomotic intimal hyperplasia and restenosis. Existing dual-expansion processes only adjust fiber arrangement through two axial stretches, improving longitudinal compliance but leaving the problem of excessive radial support unresolved. Regarding thrombosis and endothelialization obstacles: the superhydrophobic surface of ePTFE easily induces protein adsorption and platelet activation, leading to thrombosis; simultaneously, its lack of bioactive sites makes it difficult for endothelial cells to adhere and grow, resulting in low long-term patency rates. Postoperative monitoring is difficult: existing imaging examinations (such as Doppler ultrasound and angiography) cannot achieve real-time and continuous monitoring of stenosis, which is often only discovered when the stenosis has become severe, thus missing the best time for intervention.

[0003] Chinese patent CN119388815A aims to improve the self-sealing property after puncture by preparing a wearable expanded polytetrafluoroethylene (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. Clearly, existing technologies still suffer from defects such as interface delamination, coating peeling, single function, or complex structure. In particular, existing anticoagulant modifications often target single targets (such as inhibiting only coagulation factors or only platelets), failing to achieve synergistic dual-mechanism anticoagulation.

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

[0005] This invention provides a composite expanded polytetrafluoroethylene (ePTFE) artificial blood vessel and its preparation method. This small-diameter composite ePTFE artificial blood vessel has excellent biomimetic mechanical properties, long-lasting dual anticoagulation function, rapid endothelialization ability, and wireless stenosis monitoring function.

[0006] In a first aspect, the present invention provides a method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel, comprising: (1) The inner layer premix containing polytetrafluoroethylene and organic solvent, and the outer layer premix containing polytetrafluoroethylene, organic solvent, magnetic particles and elastic polymer precursor are all added to the extruder for co-extrusion to obtain a green tube. (2) The green tube is subjected to a first axial stretching, sintering and shaping, and a second axial stretching in sequence to obtain a base tube; (3) Using the base tube as a receiving device, electrospinning is performed using a spinning solution containing polytetrafluoroethylene and spinning aids, followed by heat treatment to obtain a composite base tube. (4) After plasma pretreatment of the inner surface of the composite base tube, a polydopamine layer and an anticoagulant drug layer are loaded sequentially on the inner surface to obtain the composite expanded polytetrafluoroethylene artificial blood vessel.

[0007] Preferably, in step (1): the mass ratio of polytetrafluoroethylene to organic solvent in the inner layer premix is ​​100:(15~35); The mass ratio of polytetrafluoroethylene, organic solvent, magnetic particles and elastic polymer precursor in the outer premix is ​​100:(15~35):(5~15):(3~10).

[0008] Preferably, in step (1): the organic solvent is a hydrocarbon solvent; The magnetic particles are at least one of terbium-dysprosium-iron alloy particles, amorphous alloys, Fe-Ga alloys, and Co-Fe alloys; The elastic polymer precursor is at least one of the following: polyurethane prepolymer, isocyanate-terminated polyurethane prepolymer, polydimethylsiloxane prepolymer, silicone rubber prepolymer, polyetheretherketone prepolymer, and polylactic acid-glycolic acid copolymer prepolymer.

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

[0010] More preferably, the magnetic particles have a particle size of 1~5μm.

[0011] Preferably, step (1) includes: (11) The inner layer premix is ​​extruded to form an inner layer blank; the cross-section of the inner layer blank is a first ring; (12) The outer premix is ​​extruded to form an outer blank; the cross-section of the outer blank is a second ring; the inner diameter of the second ring matches the outer diameter of the first ring so that the outer blank can coaxially fit the inner blank; (13) The inner layer blank is coaxially inserted into the outer layer blank, and a plunger extruder is used to press it into a mold composed of a mandrel and a die for co-extrusion to obtain the green tube with an inner layer and an outer layer.

[0012] Preferably, the compression ratio used for the inner layer is (30~100):1; and the compression ratio used for the outer layer is (70~200):1.

[0013] Preferably, in step (1): the mandrel is a cylindrical mandrel or a tapered mandrel; the tapered mandrel is a continuously adjustable tapered mandrel or a segmented tapered mandrel, so that a non-equidistant annular flow channel is formed between the mandrel and the die, and the taper of the mandrel in the inlet section is greater than the taper of the outlet section.

[0014] Preferably, in step (2): the stretching speed of the first axial stretching is 50~200m / min, the stretching temperature is 80~150℃, and the draw ratio is 1.5~3 times.

[0015] Preferably, in step (2): the sintering temperature is 330~380℃ and the time is 5~15min.

[0016] Preferably, in step (2): the stretching speed of the second axial stretching is 30~150m / min, the stretching temperature is 200~280℃, and the draw ratio is 2~5 times.

[0017] Preferably, in step (3): the mass ratio of polytetrafluoroethylene resin to spinning aid in the spinning solution is 1:(0.5~2); The spinning aid is at least one of polyvinyl alcohol and polyethylene glycol.

[0018] More preferably, in step (3): the mass ratio of polytetrafluoroethylene resin to spinning aid in the spinning solution is 1:(0.8~1.2).

[0019] More preferably, the degree of alcoholysis of polyvinyl alcohol is ≥88%, and the number average molecular weight of polyethylene glycol is 4000~20000.

[0020] Preferably, the electrospinning voltage is 25~40kV, the receiving distance is 12~18cm, the extrusion speed is 1~2mL / h, and the receiving roller speed is 800~1500rpm.

[0021] More preferably, the heat treatment temperature is 385~395℃ and the time is 15~25min.

[0022] Preferably, in step (4): the plasma pretreatment uses oxygen or a mixture of oxygen and argon, with a gas flow rate of 20~30 sccm, a radio frequency power of 80~120W, a working gas pressure of 20~30Pa, and a time of 60~120s.

[0023] Preferably, in step (4): the loaded polydopamine layer includes: immersing the inner surface of the composite base tube, which has been pretreated by plasma, in a buffer solution containing dopamine for reaction.

[0024] More preferably, the concentration of dopamine in the buffer solution is 2-3 mg / mL, and the reaction time is 3-5 h.

[0025] Preferably, in step (4): the loaded anticoagulant drug layer includes: adding 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide to the anticoagulant drug buffer for activation to obtain an activation solution; and injecting the activation solution into the inner surface of the composite base tube loaded with polydopamine layer for reaction.

[0026] More preferably, the concentration of the anticoagulant in the anticoagulant buffer is 2-3 mg / mL, and the concentration of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide in the activation solution is 8-12 mM and the concentration of N-hydroxysuccinimide is 4-6 mM.

[0027] More preferably, the activation time is 15-30 min and the reaction time is 3-5 h.

[0028] Preferably, the anticoagulant is obtained by the following method: Dissolve sodium heparin in MES buffer, then add 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide and stir at 25-28°C for 1-2 hours to obtain an activated solution; An amino-containing aspirin derivative was added to the activation solution and reacted at 25-28°C in the dark for 12-24 hours to obtain the reaction product. The reaction product was transferred into a dialysis bag with a molecular weight cutoff of 3500-8000 Da and dialyzed in deionized water at 4-8°C for 48-60 hours. After dialysis, the product was freeze-dried to obtain the anticoagulant.

[0029] More preferably, the molar ratio of the carboxyl group of heparin sodium to 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide is 1:(2~5):(1~3).

[0030] More preferably, the molar ratio of the carboxyl group of heparin sodium to the amino group of the aspirin derivative in the activation solution is 1:(0.5~3).

[0031] Secondly, embodiments of the present invention also provide a composite expanded polytetrafluoroethylene artificial blood vessel prepared by any of the preparation methods described in the first aspect above. The composite expanded polytetrafluoroethylene artificial blood vessel comprises, from the inside out, a biofunctional layer including an anticoagulant drug layer and a polydopamine layer, a base tube with a double-layer structure, and an electrospinning layer.

[0032] Compared with the prior art, the present invention has at least the following beneficial effects: (1) The composite expanded polytetrafluoroethylene artificial blood vessel provided in this embodiment of the invention has a four-layer integrated composite structure. From the inside out, it consists of an anticoagulant layer, a polydopamine layer, a base tube with a double-layer structure, and an electrospinning layer. The base tube is made of expanded polytetrafluoroethylene tube prepared by a double expansion process. It has a built-in magnetoelastic sensing layer. Magnetic microparticles and elastic polymers remain stably retained in the outer area of ​​the tube wall during the two stretching and sintering processes, forming a mechanical interlock with the polytetrafluoroethylene fiber network to form an integrated magnetoelastic sensing layer. This layer deforms under the action of hemodynamics, generating a magnetic field change signal that can be captured by an external coil, realizing wireless and real-time stenosis monitoring. Moreover, the base tube exhibits a nonlinear stress-strain curve in the longitudinal direction, and the problem of excessive radial support force is solved by double-bulk co-extrusion and double expansion processes. At the same time, the electrospinning layer forms a mechanical interlock with the base tube through high-temperature heat treatment, providing a high specific surface area nano-topological structure that can promote cell adhesion and tissue infiltration. On the inner surface of the base tube, anticoagulant drugs are covalently fixed through a polydopamine layer to achieve anticoagulation and promote endothelialization. Thus, this artificial blood vessel combines excellent biomimetic mechanical properties, long-lasting dual anticoagulation function, rapid endothelialization capability, and wireless stenosis monitoring function.

[0033] (2) In this embodiment of the invention, a mandrel with continuously adjustable taper or segmented taper is used. This mandrel can induce the orientation degree of the inner PTFE molecular chain to be higher than that of the outer layer, forming a fibrous gradient with denser inner layers and sparser outer layers, which significantly improves radial compliance. Moreover, this invention obtains a base tube with a gradient fibrous structure through integrated preparation, replacing the traditional multilayer compaction composite process, fundamentally eliminating the interlayer interface and solving the core hidden danger of puncture delamination. Attached Figure Description

[0034] 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.

[0035] Figure 1 This is a flowchart of a method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel according to an embodiment of the present invention; Figure 2 This is a schematic cross-sectional view of a mandrel provided in an embodiment of the present invention; Figure 3 This is a schematic cross-sectional view of a composite expanded polytetrafluoroethylene artificial blood vessel provided in an embodiment of the present invention. Detailed Implementation

[0036] 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.

[0037] The following describes the specific implementation of the concept in this application.

[0038] Please refer to Figure 1 This invention provides a method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel, comprising: (1) The inner layer premix containing polytetrafluoroethylene and organic solvent, and the outer layer premix containing polytetrafluoroethylene, organic solvent, magnetic particles and elastic polymer precursor are all added to the extruder for co-extrusion to obtain a green tube. (2) The green tube is subjected to the first axial stretching, sintering and shaping, and the second axial stretching in sequence to obtain the base tube; (3) Using the base tube as a receiving device, electrospinning is performed using a spinning solution containing polytetrafluoroethylene and spinning aids, followed by heat treatment to obtain a composite base tube. (4) After plasma pretreatment of the inner surface of the composite base tube, a polydopamine layer and an anticoagulant drug layer are loaded sequentially on the inner surface to obtain a composite expanded polytetrafluoroethylene artificial blood vessel.

[0039] In this embodiment of the invention, the composite expanded polytetrafluoroethylene (ePTFE) artificial blood vessel is a four-layer integrated composite structure. From the inside out, it consists of an anticoagulant drug layer, a polydopamine layer, a base tube with a double-layer structure, and an electrospun layer. The base tube is composed of ePTFE tubing prepared by a double expansion process, and it has a built-in magnetoelastic sensing layer. Magnetic microparticles and elastic polymers remain stably retained in the outer region of the tube wall during the two stretching and sintering processes, forming a mechanical interlock with the PTFE fiber network to form an integrated magnetoelastic sensing layer. This layer deforms under hemodynamic conditions, generating a magnetic field change signal that can be captured by an external coil, enabling wireless and real-time stenosis monitoring. Moreover, the base tube exhibits a nonlinear stress-strain curve in the longitudinal direction, and the problem of excessive radial support force is solved through double-bulk co-extrusion and double expansion processes. Simultaneously, the electrospun layer forms a mechanical interlock with the base tube through high-temperature heat treatment, providing a high specific surface area nano-topological structure that can promote cell adhesion and tissue infiltration. On the inner surface of the base tube, the anticoagulant drug is covalently fixed through the polydopamine layer to achieve anticoagulation and promote endothelialization. Thus, this artificial blood vessel possesses excellent biomimetic mechanical properties, long-lasting dual anticoagulation function, rapid endothelialization capability, and wireless stenosis monitoring function.

[0040] In a preferred embodiment, in step (1): the mass ratio of polytetrafluoroethylene (PTFE) to organic solvent in the inner layer premix is ​​100:(15~35); The mass ratio of polytetrafluoroethylene, organic solvent, magnetic particles and elastic polymer precursor in the outer premix is ​​100:(15~35):(5~15):(3~10).

[0041] Specifically, in the inner layer premix, the amount of organic solvent used is 15wt% to 35wt% of the amount of polytetrafluoroethylene (PTFE) (e.g., 15 wt%, 20 wt%, 25 wt%, 30 wt%, or 35 wt%). In the outer layer premix, the amount of organic solvent used is 15wt% to 35wt% of the amount of PTFE (e.g., 15 wt%, 20 wt%, 25 wt%, 30 wt%, or 35 wt%); the amount of magnetic particles used is 5wt% to 15wt% of the amount of PTFE (e.g., 5 wt%, 8 wt%, 10 wt%, 12 wt%, or 15 wt%); and the amount of elastic polymer precursor used is 3wt% to 10wt% of the amount of PTFE (e.g., 3 wt%, 4 wt%, 5 wt%, 6 wt%, 7 wt%, 8 wt%, 9 wt%, or 10 wt%).

[0042] In a preferred embodiment, in step (1): the organic solvent is a hydrocarbon solvent; The magnetic particles are at least one of terbium-dysprosium-iron alloy particles (Terfenol-D), amorphous alloy (Metglas), Fe-Ga alloy, and Co-Fe alloy; The elastic polymer precursor is at least one of the following: polyurethane prepolymer, isocyanate-terminated polyurethane prepolymer, polydimethylsiloxane prepolymer, silicone rubber prepolymer, polyetheretherketone prepolymer, and polylactic acid-glycolic acid copolymer prepolymer.

[0043] In a more preferred embodiment, the hydrocarbon solvent is n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, or methylcyclohexane.

[0044] In a more preferred embodiment, the magnetic particles have a particle size of 1 to 5 μm (e.g., 1 μm, 2 μm, 3 μm, 4 μm or 5 μm).

[0045] 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 mass ratio of polytetrafluoroethylene to organic solvent in the inner layer premix to 100:(15~35), good formability, stretching uniformity, and pore structure controllability 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.

[0046] In this embodiment of the invention, for the outer premix, polytetrafluoroethylene (PTFE), organic solvent, magnetic particles, and elastic polymer precursor are mixed uniformly at a mass ratio of 100:(15~35):(5~15):(3~10) to ensure that the magnetic particles and elastic polymer are uniformly dispersed in the PTFE matrix. With other component amounts remaining constant, if the amount of elastic polymer precursor is less than 3 wt% of the PTFE amount, a continuous network cannot be formed; however, if it is higher than 10 wt%, the mechanical properties of the matrix will be excessively weakened. With other component amounts remaining constant, if the amount of magnetic particles is less than 5 wt% of the PTFE amount, the signal sensitivity is insufficient, making real-time monitoring difficult; however, if it is higher than 15 wt%, extrusion will be difficult and mechanical properties will be impaired.

[0047] In a preferred embodiment, step (1) includes: (11) The inner layer premix is ​​extruded to form an inner layer blank; the cross-section of the inner layer blank is a first ring; (12) The outer premix is ​​extruded to form an outer blank; the cross-section of the outer blank is a second ring; the inner diameter of the second ring matches the outer diameter of the first ring so that the outer blank can be coaxially fitted with the inner blank; (13) The inner layer blank is coaxially inserted into the outer layer blank, and a plunger extruder is used to press it into a mold composed of a mandrel and a die for co-extrusion to obtain a green tube with an inner layer and an outer layer.

[0048] In a preferred embodiment, the compression ratio used for the inner layer is (30~100):1 (e.g., 30:1, 35:1, 40:1, 50:1, 60:1, 70:1, 80:1, 90:1 or 100:1); and the compression ratio used for the outer layer is (70~200):1 (e.g., 70:1, 75:1, 80:1, 90:1, 100:1, 120:1, 140:1, 150:1, 160:1, 180:1 or 200:1).

[0049] It should be noted that both the inner and outer blanks are in the shape of tubes, and the wall thicknesses of the inner and outer blanks can be designed according to the design requirements. No specific limitations are made here, but the sum of the wall thicknesses of the inner and outer blanks must be compatible with the mold used. The compression ratio refers to the ratio of the inlet area to the outlet area when the inner or outer blank is extruded into a tube.

[0050] In one specific implementation, to ensure the wall thickness ratio and avoid excessive packing of the outer layer or overstretching of the inner pure PTFE layer, the extrusion parameters of the inner and outer layers are usually different. The extrusion speed of the extruder is 0.5~2.0 m / min (e.g., 0.5 m / min, 0.6 m / min, 0.8 m / min, 1 m / min, 1.5 m / min, or 2 m / min). The extrusion speed of the inner layer can be slightly faster than that of the outer layer, while the outer layer needs to be decelerated to prevent uneven orientation of the magnetic particles. This avoids the problem of premix melt fracture or surface roughness caused by excessively high extrusion speed, while avoiding the problem of reduced production efficiency and insufficient orientation caused by excessively low extrusion speed. Since the inner layer is a premix of pure PTFE, it has good flowability; the outer layer contains magnetic solid particles and elastic polymer precursors, resulting in a higher apparent viscosity. It often requires a higher feed pressure or a slightly lower local compression ratio to prevent delamination. Therefore, the compression ratio of the inner layer is limited to (30~100):1, and the compression ratio of the outer layer is limited to (70~200):1. The temperature of the extruder's feeding section is 50~80℃ to prevent premature sintering and agglomeration, while the die zone can be maintained at room temperature (e.g., 25℃)~80℃. Meanwhile, for concentricity control, a mandrel fine-tuning mechanism ensures that the inner and outer preforms are coaxial, keeping the deviation within 15%.

[0051] In this invention, the mandrel can be a cylindrical rod or a tapered mandrel. In a preferred embodiment, the tapered mandrel is a continuously adjustable tapered mandrel or a segmented tapered mandrel (e.g., Figure 2 As shown in the figure, a non-equidistant annular flow channel is formed between the mandrel and the die, and the taper of the mandrel in the inlet section is greater than the taper of the outlet section.

[0052] It should be noted that the working section of the tapered mandrel is divided into 2 to 4 sections (for example, 2, 3, or 4), preferably 3 sections. Figure 2 In the middle, the working section includes the inlet section, the transition section and the outlet section. The right end is the inlet section, the middle is the transition section and the left end is the outlet section.

[0053] 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 die. This design allows the inner and outer preforms, which are coaxially nested, to experience higher shear and tensile rates when flowing through the non-equidistant annular flow channel. This induces a higher degree of pre-orientation of the inner molecular chains, preferentially forming a finer, more oriented fiber network, while the outer layer forms a relatively looser fiber network with coarser fibers and larger nodes. After two axial stretching and sintering processes, a base tube with a dense inner surface and a sparse outer surface, exhibiting a continuous gradient in fiber / node size, is obtained. Thus, the dense inner surface of the base tube provides a smooth blood contact interface and facilitates the adhesion of the polydopamine layer. Furthermore, the high mechanical strength of its intermediate transition zone and the fibrous gradient of dense inner surface and sparse outer surface significantly improve radial compliance.

[0054] 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°).

[0055] 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).

[0056] 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°).

[0057] 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°).

[0058] 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.

[0059] 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.

[0060] In this embodiment of the invention, the large taper of the inlet section allows the inner layer preform to withstand a high shear rate upon entering the flow channel, promoting the preferential orientation of the inner layer 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 layer preform to complete its final shaping under low shear conditions, forming a differentiated orientation structure between the inner and outer layers. Experiments have shown that excessive taper variation leads to the following problems: a sharp increase in the shear and tensile rates experienced by the inner layer, resulting in localized stress concentration and a tendency for cracks or breakage; excessively large differences in flow rates between the inner and outer layers, causing rough surfaces and uneven thickness in the extruded material; and an excessively steep taper in the outlet section, resulting in severe radial springback of the green tube and poor dimensional stability.

[0061] In a preferred embodiment, in step (2): the stretching speed of the first axial stretching is 50~200m / min (e.g., 50m / min, 55m / min, 60m / min, 70m / min, 80m / min, 100m / min, 120m / min, 140m / min, 150m / min, 160m / min, 180m / min or 200m / min), the stretching temperature is 80~150℃ (e.g., 80℃, 90℃, 100℃, 110℃, 120℃, 130℃, 140℃ or 150℃), and the draw ratio is 1.5~3 times (e.g., 1.5 times, 2 times, 2.5 times or 3 times).

[0062] In a preferred embodiment, in step (2): the sintering temperature is 330~380℃ (for example, it can be 330℃, 335℃, 340℃, 345℃, 350℃, 355℃, 360℃, 365℃, 370℃, 375℃ or 380℃), and the time is 5~15min (for example, it can be 5min, 8min, 10min, 12min or 15min).

[0063] In a preferred embodiment, in step (2): the stretching speed of the second axial stretching is 30~150m / min, for example, it can be 30m / min, 40m / min, 50m / min, 60m / min, 70m / min, 80m / min, 90m / min, 100m / min, 120m / min, 140m / min or 150m / min), the stretching temperature is 200~280℃, for example, it can be 200℃, 210℃, 220℃, 230℃, 240℃, 250℃, 260℃, 270℃ or 280℃), and the draw ratio is 2~5 times (for example, it can be 2 times, 2.5 times, 3 times, 3.5 times, 4 times, 4.5 times or 5 times).

[0064] Specifically, a concentric die, along with a coaxially fitted mandrel and die, is used for co-extrusion to produce a double-layer composite PTFE green tube. Subsequently, a first axial stretching is performed to initially construct the PTFE node-fiber skeleton structure and fix the dispersion of the outer magnetic microparticles and the elastic polymer precursor. During the first axial stretching, the inventors discovered that if the draw ratio is less than 1.5 times, the degree of fiberization is insufficient, the nodes are not fully stretched, leading to insufficiently fine fibrils during the subsequent second axial stretching, resulting in low porosity and excessively high radial stiffness. However, if the draw ratio is greater than 3 times, due to the high brittleness of PTFE at low temperatures, excessive stretching can cause fibril breakage, forming numerous microcracks. Furthermore, the outer elastic polymer precursor is not yet solidified, and excessive stretching will disrupt its continuity, leading to structural collapse during subsequent sintering. Sintering then melts the PTFE nodes. Finally, during the second axial stretching, the crystallinity of PTFE decreases at 200-280℃, and molecular chain segments begin to move, but the nodes are not yet completely melted. Therefore, during the second axial stretching, the outer elastic polymer precursor has softened or pre-crosslinked, allowing it to deform synchronously with the PTFE matrix without breaking. Meanwhile, the PTFE nodes are further stretched, significantly increasing porosity and drastically reducing radial stiffness. Simultaneously, magnetic particles displace synchronously with the elastic mesh and are permanently locked within the stretched pore structure, forming a stable magnetoelastic sensing layer. This is a crucial step in achieving the inner dense and outer sparse gradient structure and high radial compliance. Experiments have shown that if the stretching ratio is less than 2 times during the second axial stretching process, pore expansion is insufficient, and the decrease in radial stiffness is not significant, failing to reach the compliance level of natural blood vessels. However, if the stretching ratio is greater than 5 times, excessive stretching leads to PTFE fibril breakage, node structure destruction, and a significant decrease in burst pressure. Furthermore, the outer elastic mesh becomes thinner due to excessive stretching, potentially exposing or detaching magnetic particles, resulting in poor signal stability.

[0065] In this embodiment of the invention, a synergistic process of two axial stretching steps is employed: the first low-temperature stretching initially constructs the node-fiber skeleton structure of PTFE, fixing the dispersion state of the outer magnetic microparticles and elastic polymer precursor; the second high-temperature stretching utilizes the high ductility of PTFE near its melting point to further expand the porosity and reduce radial stiffness, while permanently locking the elastic mesh and magnetic microparticles within the gradient structure. Thus, the base tube exhibits a continuous gradient structure from the inside out, with the inner layer being high-purity ePTFE and the outer layer being an ePTFE composite region where magnetic microparticles / elastic polymer are uniformly embedded. Experiments have confirmed that if only one stretching is performed, insufficient porosity will lead to poor radial compliance, or an unstable skeleton will cause filler migration and structural inhomogeneity, neither of which can achieve the mechanical gradient design of a dense inner layer and a sparse outer layer.

[0066] It should be noted that the feeding speed during the longitudinal stretching process is matched with the extrusion speed of the extruder, ranging from 0.5 to 2.0 m / min. The stretching speed is the same as the winding speed, and it depends on the draw ratio and target output. The stretching section length is 1 to 10 m to ensure sufficient residence time at the stretching temperature. The stretching speed of the second axial stretching must be lower than that of the first axial stretching. This is because after the first axial stretching and sintering, the PTFE matrix has formed a stable node-fiber network, increasing material strength but decreasing ductility; simultaneously, the outer elastic polymer precursor has partially cross-linked, enhancing viscoelasticity. Therefore, reducing the stretching speed during the second axial stretching ensures that the elastic network and the PTFE matrix deform synchronously and in coordination, avoiding fiber breakage, elastomer interface peeling, or localized necking due to excessively high strain rates, thus ensuring the integrity of the gradient structure and the uniformity of performance.

[0067] In this embodiment of the invention, heat setting can be performed using a hot air oven or heated rollers, or both. Taking isocyanate-terminated polyurethane prepolymer as an example, during the first axial stretching and heating stage, the viscosity of the prepolymer decreases and fills the pores between PTFE particles. During the heating stage of 150~250℃, the isocyanate groups in the prepolymer undergo chain extension and cross-linking reactions with water and its own hydroxyl groups, forming a three-dimensional network structure, which significantly improves thermal stability. Although the subsequent sintering and setting is a short-time high-temperature treatment, the prepolymer has already completed preliminary cross-linking and is physically wrapped by the PTFE fiber network, so only a small amount of thermal oxidative decomposition occurs (mass loss is less than 10wt%), and the main structure is still completely preserved. At the same time, the cross-linked polyurethane and PTFE fiber nodes form a tight mechanical interlock, further preventing its migration or volatilization.

[0068] In a preferred embodiment, in step (3): the mass ratio of polytetrafluoroethylene resin to spinning aid in the spinning solution is 1:(0.5~2) (for example, it can be 1:0.5, 1:0.6, 1:0.8, 1:1.0, 1:1.5, 1:1.8 or 1:2). The spinning aid is at least one of polyvinyl alcohol (PVA) and polyethylene glycol (PEG).

[0069] In a more preferred embodiment, in step (3): the mass ratio of polytetrafluoroethylene resin to spinning aid in the spinning solution is 1:(0.8~1.2) (for example, it can be 1:0.8, 1:0.9, 1:1, 1:1.1 or 1:1.2).

[0070] In this embodiment of the invention, for the spinning solution, if the mass ratio of polytetrafluoroethylene resin to spinning aid is greater than 1:0.5, the resin dosage is too high, the viscosity of the spinning solution is too high, the jet is not easy to split, resulting in a coarser fiber diameter (>1μm), a decrease in porosity, and is not conducive to cell adhesion. However, if the mass ratio of polytetrafluoroethylene resin to spinning aid is less than 1:2, the resin dosage is too low, the fiber continuity is poor, and it is easy to break and bead, resulting in poor mechanical properties and easy shedding after high-temperature treatment. Specifically, in the spinning process, the spinning solution is prepared by the following method: polytetrafluoroethylene dispersion resin and spinning aid are added to deionized water at a mass ratio of 1:(0.5~2), and stirred evenly at 60°C to obtain a spinning solution with a solid content of 9wt%~12wt%.

[0071] In a more preferred embodiment, the degree of alcoholysis of polyvinyl alcohol is ≥88%, and the number average molecular weight of polyethylene glycol is 4000~20000.

[0072] Specifically, PVA with a degree of hydrolysis ≥88% and PEG with a number average molecular weight of 4000~20000 are preferred. These spinning aids can be completely decomposed and volatilized during heat treatment at 385~395℃, leaving no inorganic carbonaceous impurities.

[0073] In a preferred embodiment, the electrospinning voltage is 25-40 kV (e.g., 25 kV, 30 kV, 35 kV, or 40 kV), the receiving distance is 12-18 cm (e.g., 12 cm, 14 cm, 15 cm, 16 cm, or 18 cm), the extrusion speed is 1-2 mL / h (e.g., 1 mL / h, 1.2 mL / h, 1.5 mL / h, 1.6 mL / h, 1.8 mL / h, or 2 mL / h), and the receiving roller speed is 800-1500 rpm (e.g., 800 rpm, 900 rpm, 1000 rpm, 1200 rpm, 1400 rpm, or 1500 rpm).

[0074] Specifically, during the electrospinning process, the relative humidity is controlled at 45-55%RH, and the spinning time is 3-5 hours. Based on these electrospinning parameters, a nanofiber coating layer (i.e., the electrospinned layer) with a thickness of 50-100 μm can be deposited on the outer surface of the base tube. The ratio of the thickness of this nanofiber layer to the wall thickness of the base tube is 1:(4-10) (for example, it can be 1:4, 1:5, 1:6, 1:7, 1:8, 1:9, or 1:10). This ratio ensures that the nanofiber layer provides sufficient cell attachment area and biological functional sites without excessively increasing the overall wall thickness and affecting surgical suturing performance. However, if the ratio of the nanofiber layer thickness to the base tube wall thickness is less than 1:10, the fiber layer is too thin, and the biofunctionalization effect is limited; if the ratio is greater than 1:3, the overall wall thickness is too large, and it is easy to cause interlayer delamination and poor tissue infiltration.

[0075] In a more preferred embodiment, the heat treatment temperature is 385~395°C (e.g., 385°C, 390°C or 395°C) and the time is 15~25 min (e.g., 15 min, 20 min or 25 min).

[0076] In this embodiment of the invention, high-temperature heat treatment completely decomposes and volatilizes the spinning aids. Simultaneously, PTFE nanofibers and the D-ePTFE fibers of the substrate undergo micro-melting bonding at the contact points, forming a stable mechanically interlocking interface without disrupting the overall porous topology of the fibers. The resulting composite tube has an electrospun layer fiber diameter of 100-800 nm and a porosity ≥70%, which can simulate the nanostructure of the natural vascular extracellular matrix, significantly promoting endothelial cell adhesion and proliferation. It should be noted that the heat treatment must be carried out under nitrogen protection.

[0077] In a preferred embodiment, in step (4): the oxygen plasma pretreatment uses oxygen or a mixture of oxygen and argon, with a gas flow rate of 20-30 sccm (e.g., 20 sccm, 22 sccm, 25 sccm, 28 sccm or 30 sccm); the radio frequency power is 80-120 W (e.g., 80 W, 85 W, 90 W, 95 W, 100 W, 105 W, 110 W, 115 W or 120 W); the working pressure is 20-30 Pa (e.g., 20 Pa, 25 Pa or 30 Pa); and the time is 60-120 s (e.g., 60 s, 80 s, 90 s, 100 s or 120 s).

[0078] 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.

[0079] 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, excessive etching and embrittlement of the outer wall, or disruption of the composite phase interface stability of the bilayer substrate, 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 20-30 sccm to avoid plasma instability due to excessively low flow rate and accelerated etching due to excessively high flow rate. Simultaneously, the working pressure (cavity vacuum) is limited to 20-30 Pa for the inner surface of the composite substrate to ensure uniform glow discharge. If the radio frequency power exceeds 120W, it can easily lead to deep-hole etching, nanopillar structures, or even damage to the PTFE node structure or disruption of the composite phase interface stability on the inner surface of the composite substrate; however, if the radio frequency power is below 80W, insufficient activation will result. In addition, to avoid excessively deep CF bond breakage due to excessive processing time, which could lead to surface embrittlement or hydrophobic recovery, the processing time is limited to 60~120s.

[0080] In this embodiment of the invention, the surface energy of polytetrafluoroethylene is extremely low (approximately 18~20 mN / m), and the CF bond has high inertness. If the polydopamine layer is directly loaded without oxygen plasma pretreatment, it 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. Polydopamine can only be linked by van der Waals forces or weak mechanical intercalation, lacking COC or CN covalent bond anchoring. The bonding strength is far lower than that after plasma activation and repolymerization, and the mechanical interlocking depth is insufficient; under arterial blood flow shear force (greater than 10~15 dyn / cm) 2 During repeated puncture-hemostasis cycles, the polydopamine layer, which is physically adsorbed or simply intercalated, is easily detached or swollen and lost, thus losing its endothelialization-promoting function.

[0081] In a preferred embodiment, in step (4): the loaded polydopamine layer includes: immersing the inner surface of the plasma-pretreated composite base tube in a buffer solution containing dopamine for reaction.

[0082] In a more preferred embodiment, the concentration of dopamine in the buffer solution is 2-3 mg / mL (e.g., 2 mg / mL, 2.5 mg / mL or 3 mg / mL), and the reaction time is 3-5 h (e.g., 3 h, 3.5 h, 4 h, 4.5 h or 5 h).

[0083] It should be noted that the dopamine-containing buffer solution is a Tris-HCl buffer solution with a pH of 8.3~8.5, and the reaction is allowed to proceed at room temperature in the dark for 3~5 hours, so that dopamine can polymerize in situ on the inner surface of the tube to form a uniform polydopamine (PDA) layer; the PDA layer is rich in catechol and free amino groups, which provide reaction sites for subsequent coupling.

[0084] In this embodiment of the invention, the dopamine concentration is controlled at 2-3 mg / mL and the reaction time at 3-5 h to precisely regulate the thickness and chemical activity of the polydopamine (PDA) layer. If the dopamine concentration is too low or the reaction time is too short, the PDA coating will be too thin and discontinuous, resulting in insufficient surface-active amino sites, low graft density of anticoagulant drugs, and easy detachment. If the dopamine concentration is too high or the reaction time is too long, the PDA coating will be too thick (>100 nm) and accompanied by the deposition of a large number of self-aggregating particles, which will not only block the microporous structure of the ePTFE matrix, hindering substance transport and cell infiltration, but may also lead to a decrease in coupling activity due to the oxidation of amino groups on the surface. By controlling the above parameters, this invention can construct a continuous, dense PDA transition layer with a thickness of 20-40 nm on the inner surface of ePTFE. This thickness can completely cover the substrate to provide sufficient biofunctionalization sites without destroying the original porous topology of the matrix, ensuring the efficient and stable fixation of subsequent anticoagulant drugs.

[0085] In a preferred embodiment, in step (4): the loading of the anticoagulant drug layer includes: adding 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide to the anticoagulant drug buffer to activate it, thereby obtaining an activation solution; and injecting the activation solution into the inner surface of the composite base tube loaded with the polydopamine layer for reaction.

[0086] In a more preferred embodiment, the concentration of the anticoagulant in the anticoagulant buffer is 2-3 mg / mL (e.g., 2 mg / mL, 2.5 mg / mL, or 3 mg / mL), the concentration of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide in the activation solution is 8-12 mM (e.g., 8 mM, 9 mM, 10 mM, 11 mM, or 12 mM), and the concentration of N-hydroxysuccinimide is 4-6 mM (e.g., 4 mM, 5 mM, or 6 mM).

[0087] In a more preferred embodiment, the activation time is 15-30 min (e.g., 15 min, 20 min, 25 min or 30 min), and the reaction time is 3-5 h (e.g., 3 h, 4 h or 5 h).

[0088] In this embodiment of the invention, the anticoagulant is preferably a heparin-aspirin complex (Hep-ASA). The Hep-ASA complex is prepared as a 2-3 mg / mL solution, with a concentration range consistent with that of the dopamine buffer solution. This ensures that the supply of Hep-ASA matches the amino density on the surface of the PDA layer (formed by in-situ polymerization of a 2-3 mg / mL dopamine-containing buffer solution). This achieves high-density, stable covalent grafting of Hep-ASA without overloading or underloading of functional molecules, while simplifying process parameter control and improving batch repeatability.

[0089] In a preferred embodiment, the anticoagulant is obtained by the following method: Dissolve sodium heparin in MES buffer, then add 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide and stir at 25-28°C for 1-2 hours to obtain an activated solution; An amino-containing aspirin derivative was added to the activation solution and reacted at 25-28°C in the dark for 12-24 hours to obtain the reaction product. The reaction product was transferred into a dialysis bag with a molecular weight cutoff of 3500~8000 Da and dialyzed in deionized water at 4~8℃ for 48~60h. After dialysis was completed, the product was freeze-dried to obtain the anticoagulant.

[0090] In a more preferred embodiment, the molar ratio of the carboxyl group of heparin sodium to 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide is 1:(2~5):(1~3) (for example, it can be 1:2:1, 1:3:1, 1:4:1, 1:5:1, 1:2:2, 1:3:22, 1:4:2, 1:5:2, 1:2:3, 1:3:3, 1:4:3 or 1:5:3).

[0091] In a more preferred embodiment, the molar ratio of the carboxyl group of heparin sodium to the amino group of the aspirin derivative in the activation solution is 1:(0.5~3) (for example, it can be 1:0.5, 1:1, 1:1.5, 1:2, 1:2.5 or 1:3).

[0092] In this invention, the molar ratio of the free carboxyl group on the heparin sodium molecule to 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) and N-hydroxysuccinimide (NHS) is specified to be 1:(2~5):(1~3), which achieves efficient activation of the carboxyl group while avoiding excessive cross-linking or degradation of the heparin molecule. By adjusting the molar ratio of the carboxyl group of heparin sodium to the amino group of the aspirin derivative to be 1:(0.5~3), the grafting density of aspirin can be precisely controlled. While retaining the anticoagulation factor activity of heparin, the material is endowed with excellent antiplatelet aggregation function, achieving dual synergistic anticoagulation and promoting endothelialization. Furthermore, this invention chemically couples heparin (which inhibits coagulation factors) and aspirin (which inhibits platelet aggregation) to form a single bifunctional molecule (Hep-ASA) and covalently fixes it, solving the problems of easy explosive release and single function of small molecule drugs.

[0093] This invention also provides a composite expanded polytetrafluoroethylene artificial blood vessel, such as... Figure 3 As shown, from the inside out, there are a biofunctional layer 10 including an anticoagulant drug layer and a polydopamine layer, a base tube with a double-layer structure consisting of an inner layer 20 and an outer layer 30, and an electrospinning layer 40.

[0094] The composite expanded polytetrafluoroethylene artificial blood vessel provided by this invention can not only replace diseased blood vessels, but also actively monitor postoperative stenosis. It can also significantly reduce the risk of thrombosis through long-lasting and powerful dual anticoagulation function, realize closed-loop management of "early warning-intervention-protection", significantly improve the safety of patients after surgery, and can be widely used in the fields of coronary artery bypass grafting, peripheral vascular reconstruction, and hemodialysis access, with great economic benefits and value.

[0095] 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.

[0096] Example 1 A method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel, comprising: (1) PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:25 to obtain inner layer premix; PTFE dispersion resin, Isopar G organic solvent, high magnetostrictive magnetic microparticles Terfenol-D with a particle size of 3μm and isocyanate-terminated polyurethane prepolymer are mixed evenly at a mass ratio of 100:25:10:6 to obtain outer layer premix; The extrusion die is a plunger type, consisting of a mandrel and a die; the inner barrel has an inner diameter of 12mm, the outer barrel has an inner diameter of 17mm, and the die has an inner diameter of 6.0mm; the mandrel is a three-section tapered mandrel, with a total working section length of 120mm. Inlet section: length 30mm, inlet end mandrel diameter 4.0mm (inlet gap 1.0mm), outlet end mandrel diameter 4.8mm, cone angle 5°, taper change rate 0.267° / cm; Transition section: 40mm in length, 5.0mm in diameter of the mandrel at the outlet end, 3° in cone angle, and a taper variation rate of 0.25° / cm; Outlet section: 50mm in length, 5.2mm diameter of mandrel at the outlet end, 1.5° cone angle, and a taper change rate of 0.15° / cm; a non-equidistant annular flow channel is formed between the mandrel and the die, with an inlet gap of 1.0mm and an outlet gap of 0.4mm; The inner premix is ​​extruded to form an inner preform; the cross-section of the inner preform is a first ring. The outer premix is ​​extruded to form an outer preform; the cross-section of the outer preform is a second ring. The inner diameter of the second ring matches the outer diameter of the first ring so that the outer preform can coaxially fit the inner preform. The inner preform is coaxially fitted into the outer preform, and the fitted inner and outer preforms are pressed into the die using a plunger extruder for co-extrusion to obtain a double-layer composite PTFE green tube. The extrusion speed is 0.8 m / min, the extrusion temperature is 50℃, and the extrusion pressure is 12 MPa. The compression ratio used for the inner layer is 60:1, and the compression ratio used for the outer layer is 120:1. (2) The obtained green tube was subjected to a first axial stretch at 120℃ with a stretching speed of 100 mm / min and a draw ratio of 2.5 times; then it was kept at 365℃ for 10 min to allow the PTFE nodes to melt and bond, while the polyurethane prepolymer underwent cross-linking and curing; finally, a second axial stretch was performed at 250℃ with a stretching speed of 80 mm / min and a draw ratio of 3.5 times, resulting in a D-ePTFE base tube with a wall thickness of approximately 0.45 mm and an inner diameter of 4.0 m. The inner wall of the base tube was made of high-purity ePTFE, and the outer layer was embedded with a uniformly distributed magnetoelastic sensing layer; the radial compliance of the base tube was tested to be 4.8% / 100 mmHg. (3) Polytetrafluoroethylene dispersion resin and polyvinyl alcohol (PVA) with a degree of alcoholysis of 88% are added to deionized water at a mass ratio of 1:1 and stirred evenly at 60°C to obtain a spinning solution with a solid content of about 12wt%. Using the D-ePTFE base tube from step (2) as a rotating receiving device, electrospinning was performed for 4 hours at an extrusion speed of 1.5 mL / h using a spinning solution under conditions of 30 kV and 50% RH to deposit a nanofiber layer with a thickness of approximately 70 μm on the outer surface of the base tube; wherein, the distance between the spinneret and the rotating receiving device was 15 cm; the receiving roller speed was 1000 rpm; the ratio of this thickness to the wall thickness of the D-ePTFE base tube was approximately 1:6.4; The aforementioned substrate tube with deposited nanofiber layer was then placed under nitrogen protection and subjected to high-temperature heat treatment at 390℃ for 20 minutes to obtain a composite substrate tube; wherein, the fiber diameter of the electrospun layer on the outer surface of the composite substrate tube is 200~500nm, and the porosity reaches 78%; (4) Preparation method of anticoagulant drugs: a. Dissolve heparin sodium in MES (2-(N-morpholino)ethanesulfonic acid) buffer at pH 6.0 to prepare a 20 mg / mL solution; add EDC and NHS at a molar ratio of free carboxyl group on the heparin sodium molecule to EDC and NHS of 1:3:2, and stir at room temperature (25℃) for 1.5 h to form a stable active ester intermediate, thus obtaining the activated solution; b. Add an amino-containing aspirin derivative (para-aminoacetylsalicylic acid) to the activation solution and react at room temperature (25°C) in the dark for 18 hours to obtain the reaction product; wherein, the molar ratio of the carboxyl group of heparin sodium to the amino group of aspirin derivative in the activation solution is 1:1.5. c. The reaction product was transferred into a dialysis bag with a molecular weight cutoff of 3500 Da and dialyzed in deionized water at 4°C for 48 h, with the deionized water being changed every 8 h. After dialysis, the product was pre-frozen at -35°C for 4 h, and then placed in a freeze dryer (cold trap temperature ≤ -50°C, vacuum degree 5 Pa) for sublimation drying for 30 h to obtain a white, loose, spongy purified Hep-ASA complex, i.e., an anticoagulant. The inner surface of the composite tube obtained in step (3) was subjected to plasma pretreatment. The treatment conditions were: pure oxygen was used, the oxygen flow rate was 25 sccm, the working pressure was 25 Pa, the radio frequency power was 100 W, the treatment time was 90 s, and the substrate temperature was controlled below 50 ℃. Then, the inner surface of the plasma-pretreated composite tube was immersed in a dopamine-Tris buffer solution (pH=8.5) with a dopamine concentration of 2.5 mg / mL and allowed to stand in the dark at room temperature (25 ℃) for 4 h to allow dopamine to polymerize in situ on the inner surface of the tube, forming a uniform polydopamine layer with a thickness of about 20~40 nm and rich in catechol and free amino groups. Then, the purified Hep-ASA complex was dissolved in MES buffer solution with pH=6.0 to prepare a solution of 2.5 mg / mL and EDC was added. The PDA-loaded composite tube was activated with NHS for 20 min to obtain an activation solution. The concentration of EDC in the activation solution was 10 mM and the concentration of NHS was 5 mM. The activation solution was filled into the inner cavity of the composite tube loaded with PDA layer and sealed. The tube was slowly rotated at 30 rpm at room temperature (25℃) for 4 h to allow the amino groups on the PDA layer to be covalently linked to the carboxyl groups on the Hep-ASA complex through amide bonds. After the reaction was completed, the tube was thoroughly rinsed with deionized water and PBS and vacuum dried at 30℃ for 4 h to obtain the biofunctionalized composite tube. The biofunctionalized composite base tube was placed in an ethylene oxide (EO) sterilizer and sterilized for 3 hours at 35°C, 50% RH relative humidity, and an EO concentration of 600 mg / L. Subsequently, it was desorbed at 50°C for 10 hours under ventilated conditions to remove residual EO monomers, and finally, a composite expanded polytetrafluoroethylene artificial blood vessel was obtained.

[0097] Example 2 A method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel, comprising: (1) PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:20 to obtain inner layer premix; PTFE dispersion resin, Isopar G organic solvent, high magnetostrictive magnetic microparticles Terfenol-D with a particle size of 3μm and isocyanate-terminated polyurethane prepolymer are mixed evenly at a mass ratio of 100:20:8:4 to obtain outer layer premix. The extrusion die is a plunger type, with an inner barrel diameter of 11mm, an outer barrel diameter of 16mm, and a die inner diameter of 6.0mm. The mandrel is a three-section tapered mandrel with a total working section length of 120mm. Inlet section: length 40mm, inlet end mandrel diameter 4.2mm (inlet gap 0.9mm), outlet end mandrel diameter 4.9mm, cone angle 4°, taper change rate 0.2° / cm; Transition section: 30mm in length, 5.05mm in diameter of the mandrel at the outlet end, 2.5° in cone angle, and a taper variation rate of 0.25° / cm; Outlet section: 50mm in length, 5.2mm diameter of the mandrel at the outlet end (0.4mm outlet gap), 1.5° cone angle, and a taper change rate of 0.15° / cm; the inlet gap is 0.9mm and the outlet gap is 0.4mm. The inner premix is ​​extruded to form an inner preform; the cross-section of the inner preform is a first ring. The outer premix is ​​extruded to form an outer preform; the cross-section of the outer preform is a second ring. The inner diameter of the second ring matches the outer diameter of the first ring so that the outer preform can coaxially fit the inner preform. The inner preform is coaxially fitted into the outer preform, and the fitted inner and outer preforms are pressed into the die using a plunger extruder for co-extrusion to obtain a double-layer composite PTFE green tube. The extrusion speed is 1.0 m / min, the extrusion temperature is 50℃, and the extrusion pressure is 10 MPa. The compression ratio of the inner layer is 50:1, and the compression ratio of the outer layer is 100:1. (2) The obtained green tube was subjected to a first axial stretch at 100℃ with a stretching speed of 120 mm / min and a draw ratio of 2.0 times; then it was kept at 350℃ for 15 min to allow the PTFE joints to melt and bond, while the polyurethane prepolymer underwent cross-linking and curing; finally, a second axial stretch was performed at 230℃ with a stretching speed of 100 mm / min and a draw ratio of 3.0 times, resulting in a D-ePTFE base tube with a wall thickness of approximately 0.50 mm and an inner diameter of 4.0 m. The inner wall of the base tube was made of high-purity ePTFE, and the outer layer was embedded with a uniformly distributed magnetoelastic sensing layer; the radial compliance of the base tube was tested to be 3.9% / 100 mmHg. (3) Polytetrafluoroethylene dispersion resin and polyvinyl alcohol (PVA) with a degree of alcoholysis of 88% are added to deionized water at a mass ratio of 1:0.8 and stirred evenly at 60°C to obtain a spinning solution with a solid content of about 10wt%. Using the D-ePTFE base tube from step (2) as a rotating receiving device, electrospinning was performed for 3 hours at an extrusion speed of 2.0 mL / h using a spinning solution under conditions of 35 kV and 55% RH, to deposit a nanofiber layer with a thickness of approximately 50 μm on the outer surface of the base tube; wherein, the distance between the spinneret and the rotating receiving device was 12 cm; the receiving roller speed was 1500 rpm; the ratio of this thickness to the wall thickness of the D-ePTFE base tube was approximately 1:10; The nanofiber-coated substrate was then placed under nitrogen protection and subjected to high-temperature heat treatment at 385°C for 25 minutes to obtain a composite substrate. The electrospun fibers on the outer surface of the composite substrate had a diameter of 100-400 nm and a porosity of 75%. (4) Preparation method of anticoagulant drugs: a. Dissolve heparin sodium in MES (2-(N-morpholino)ethanesulfonic acid) buffer at pH 6.0 to prepare a 15 mg / mL solution; add EDC and NHS at a molar ratio of free carboxyl group on the heparin sodium molecule to EDC and NHS of 1:2:1, and stir at room temperature (25℃) for 1.0 h to form a stable active ester intermediate, thus obtaining the activated solution; b. Add an amino-containing aspirin derivative (para-aminoacetylsalicylic acid) to the activation solution and react at room temperature (25°C) in the dark for 12 hours to obtain the reaction product; wherein, the molar ratio of the carboxyl group of heparin sodium to the amino group of aspirin derivative in the activation solution is 1:0.8. c. The reaction product was transferred into a dialysis bag with a molecular weight cutoff of 3500 Da and dialyzed in deionized water at 4°C for 48 h, with the deionized water being changed every 8 h. After dialysis, the product was pre-frozen at -35°C for 4 h, and then placed in a freeze dryer (cold trap temperature ≤ -50°C, vacuum degree 5 Pa) for sublimation drying for 30 h to obtain a white, loose, spongy purified Hep-ASA complex, i.e., an anticoagulant. The inner surface of the composite base tube obtained in step (3) is subjected to plasma pretreatment. The treatment conditions are: pure oxygen is used, the oxygen flow rate is 30 sccm, the working pressure is 30 Pa, the radio frequency power is 80 W, the treatment time is 60 s, and the substrate temperature is controlled below 50 °C. The inner surface of the plasma-pretreated composite tube was then immersed in a dopamine-Tris buffer solution (pH=8.5) with a dopamine concentration of 2.0 mg / mL and allowed to stand in the dark at room temperature (25℃) for 3 hours to allow dopamine to polymerize in situ on the inner surface of the tube, forming a polydopamine layer. The purified Hep-ASA complex was then dissolved in MES buffer at pH=6.0 to prepare a 2.0 mg / mL solution. EDC and NHS were added, and the solution was activated for 15 minutes to obtain an activation solution. The concentration of EDC in the activation solution was 8 mM, and the concentration of NHS was 4 mM. The activation solution was filled into the inner cavity of the composite tube loaded with the PDA layer and sealed. The tube was then slowly rotated at 30 rpm at room temperature (25℃) for 3 hours to allow the amino groups on the PDA layer to be covalently linked to the carboxyl groups on the Hep-ASA complex via amide bonds. After the reaction, the tube was thoroughly rinsed with deionized water and PBS and then vacuum dried at 30℃ for 4 hours to obtain the biofunctionalized composite tube. The biofunctionalized composite base tube was placed in an ethylene oxide (EO) sterilizer and sterilized for 3 hours at 35°C, 50% RH relative humidity, and an EO concentration of 600 mg / L. Subsequently, it was desorbed at 50°C for 10 hours under ventilated conditions to remove residual EO monomers, and finally, a composite expanded polytetrafluoroethylene artificial blood vessel was obtained.

[0098] Example 3 Example 3 is basically the same as Example 1, except that a cylindrical core rod is used and the content of organic solvent in the outer premix is ​​higher than that in the inner premix.

[0099] Specifically, (1) PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:18 to obtain an inner layer premix; PTFE dispersion resin, Isopar G organic solvent, high magnetostrictive magnetic microparticles Terfenol-D with a particle size of 3μm and isocyanate-terminated polyurethane prepolymer are mixed evenly at a mass ratio of 100:30:12:7 to obtain an outer layer premix; The extrusion die is a plunger type. The inner barrel has an inner diameter of 13mm, the outer barrel has an inner diameter of 19mm, and the die has an inner diameter of 6.0mm. The mandrel is a cylindrical mandrel with a constant diameter of 120mm and a constant diameter of 5.2mm. An equidistant annular flow channel is formed between the mandrel and the die, with a constant gap of 0.4mm. The taper variation rate of the entire section is 0° / cm. The inner premix is ​​extruded to form an inner preform; the cross-section of the inner preform is a first ring. The outer premix is ​​extruded to form an outer preform; the cross-section of the outer preform is a second ring. The inner diameter of the second ring matches the outer diameter of the first ring so that the outer preform can coaxially fit the inner preform. The inner preform is coaxially fitted into the outer preform, and the fitted inner and outer preforms are pressed into the die using a plunger extruder for co-extrusion to obtain a double-layer composite PTFE green tube. The extrusion speed is 0.6 m / min, the extrusion temperature is 50℃, and the extrusion pressure is 15 MPa. The compression ratio of the inner layer is 70:1, and the compression ratio of the outer layer is 150:1. (2) The obtained green tube was subjected to a first axial stretch at 130℃ with a stretching speed of 80 mm / min and a draw ratio of 2.8 times; then it was kept at 370℃ for 8 min to allow the PTFE nodes to melt and bond, while the polyurethane prepolymer underwent cross-linking and curing; finally, a second axial stretch was performed at 260℃ with a stretching speed of 70 mm / min and a draw ratio of 4.0 times, resulting in a D-ePTFE base tube with a wall thickness of approximately 0.42 mm and an inner diameter of 4.0 m. The inner wall of the base tube is made of high-purity ePTFE, and the outer layer is embedded with a uniformly distributed magnetoelastic sensing layer. The inner wall is dense and smooth, while the outer wall is loose and porous. The radial compliance of the base tube was tested to be 5.2% / 100 mmHg.

[0100] Example 4 Example 4 is basically the same as Example 1, except that the ratio of the outer layer premix is ​​different.

[0101] Specifically, in step (1), for the outer layer premix, PTFE dispersion resin, Isopar G organic solvent, Co-Fe alloy with a particle size of 3 μm and polydimethylsiloxane prepolymer are mixed in a mass ratio of 100:35:15:10 to obtain the outer layer premix.

[0102] Example 5 Example 5 is basically the same as Example 1, except that the ratio of the inner layer premix is ​​different.

[0103] Specifically, in step (1), for the inner layer premix, PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:35 to obtain the inner layer premix.

[0104] Example 6 Example 6 is basically the same as Example 1, except that the stretching speed used in the two axial stretching operations is different.

[0105] Specifically, (2) the obtained green tube was subjected to a first axial stretch at 120°C with a stretching speed of 100 mm / min and a draw ratio of 3; then it was kept at 365°C for 10 min to allow the PTFE nodes to melt and bond, while the polyurethane prepolymer underwent cross-linking and curing; finally, a second axial stretch was performed at 250°C with a stretching speed of 80 mm / min and a draw ratio of 5, ultimately producing a D-ePTFE base tube with a wall thickness of approximately 0.45 mm and an inner diameter of 4.0 m. The inner wall of the base tube is made of high-purity ePTFE, and the outer layer is embedded with a uniformly distributed magnetoelastic sensing layer; the radial compliance of the base tube was tested to be 4.8% / 100 mmHg.

[0106] Example 7 Example 7 is basically the same as Example 1, except that the spinning solution used is different.

[0107] Specifically, in step (3), polytetrafluoroethylene dispersion resin and polyethylene glycol with a number average molecular weight of 4000 are added to deionized water at a mass ratio of 1:2 and stirred evenly at 60°C to obtain a spinning solution with a solid content of about 12wt%. Using the D-ePTFE base tube from step (2) as a rotating receiving device, electrospinning was performed for 4 hours at an extrusion speed of 1.2 mL / h using a spinning solution under conditions of 30 kV and 50% RH, to deposit a nanofiber layer with a thickness of approximately 65 μm on the outer surface of the base tube; wherein, the distance between the spinneret and the rotating receiving device was 15 cm; the receiving roller speed was 1000 rpm; the ratio of this thickness to the wall thickness of the D-ePTFE base tube was approximately 1:7; Then, the substrate tube with the deposited nanofiber layer was placed under nitrogen protection and subjected to high-temperature heat treatment at 390°C for 20 minutes to obtain the composite substrate tube. The fiber diameter of the electrospun layer on the outer surface of the composite substrate tube is 150~450nm, and the porosity is 76%. Compared with the PVA system, the fiber has better flexibility and is more conducive to the extension of cell pseudopodia.

[0108] Example 8 Example 8 is basically the same as Example 1, except that the electrospinning parameters used are different.

[0109] Specifically, in step (3), the D-ePTFE base tube from step (2) is used as a rotating receiving device. Electrospinning is performed for 4 hours at an extrusion speed of 1 mL / h using a spinning solution under conditions of 25 kV and 50% RH to deposit a nanofiber layer with a thickness of about 70 μm on the outer surface of the base tube. The distance between the spinneret and the rotating receiving device is 18 cm, and the receiving roller speed is 800 rpm. Then, the substrate tube with the deposited nanofiber layer was placed under nitrogen protection and subjected to high-temperature heat treatment at 390°C for 20 minutes to obtain the composite substrate tube.

[0110] Example 9 Example 9 is basically the same as Example 1, except that the plasma pretreatment parameters are different.

[0111] Specifically, in step (4), the inner surface of the composite base tube obtained in step (3) is subjected to plasma pretreatment; the treatment conditions are: pure oxygen is used, the oxygen flow rate is 20 sccm, the working pressure is 20 Pa, the radio frequency power is 120 W, the treatment time is 120 s, and the substrate temperature is controlled below 50 °C.

[0112] Example 10 Example 10 is basically the same as Example 1, except that it uses a method for loading polydopamine.

[0113] The inner surface of the plasma-pretreated composite tube was immersed in a dopamine-Tris buffer solution (pH=8.5) with a dopamine concentration of 3.0 mg / mL and allowed to stand in the dark at room temperature (25℃) for 3 h to allow dopamine to polymerize in situ on the inner surface of the tube, forming a polydopamine layer. Then, the purified Hep-ASA complex was dissolved in MES buffer (pH=6.0) to prepare a 3.0 mg / mL solution. EDC and NHS were added, and the solution was activated for 15 min to obtain an activation solution. The concentration of EDC in the activation solution was 8 mM, and the concentration of NHS was 4 mM. The activation solution was filled into the inner cavity of the composite tube loaded with the PDA layer and sealed. The tube was then slowly rotated at 30 rpm at room temperature (25℃) for 3 h to allow the amino groups on the PDA layer to be covalently linked to the carboxyl groups on the Hep-ASA complex via amide bonds. After the reaction, the tube was thoroughly rinsed with deionized water and PBS and then vacuum dried at 30℃ for 4 h to obtain the biofunctionalized composite tube.

[0114] Comparative Example 1 Comparative Example 1 is basically the same as Example 1, except that only one axial stretching was used in step (2).

[0115] Specifically, (2) the obtained green tube is axially stretched at 200°C with a stretching speed of 60 mm / min and a stretching ratio of 8.75 times; then it is kept at 365°C for 10 min to melt and bond the PTFE nodes, while the polyurethane prepolymer undergoes cross-linking and curing; finally, a uniaxially stretched D-ePTFE base tube with a wall thickness of about 0.44 mm and an inner diameter of 4.0 m is obtained.

[0116] The artificial blood vessels prepared in Example 1 and Comparative Example 1 were subjected to performance tests, and the results are shown in Table 1 below.

[0117] Table 1 As shown in Table 1, although the dimensions of the base tube prepared in Comparative Example 1 are basically the same as those in Example 1, the single stretching in Comparative Example 1 was performed at a higher temperature in a single molding process. The PTFE fiber nodes were uniformly stretched apart, lacking the node-fiber gradient relaxation structure formed by step stretching. This resulted in a higher overall rigidity of the tube wall, failing to effectively simulate the elasticity of natural blood vessels, and thus significantly reducing radial compliance. Furthermore, the single stretching in Comparative Example 1 caused localized stress concentration in the fiber network, and the outer elastic polymer mesh, lacking the fixation of low-temperature pre-stretching, was prone to plastic flow or fracture, weakening the overall pressure-bearing capacity, and consequently reducing the burst pressure. Moreover, the single stretching disrupted the inner dense and outer sparse gradient structure, leading to asynchronous deformation of different layers of the tube wall. When the blood vessel was compressed, the correlation between the deformation of the outer magnetic microparticle area and the overall tube diameter change decreased, resulting in signal hysteresis and distortion. The uneven surface energy distribution of the base tube after a single stretch in Comparative Example 1 affected the uniform anchoring of the subsequent PDA coating and the mechanical interlocking effect of the nanofibers. Therefore, the electrospun layer and the substrate were more prone to micro-delamination during the bending test. Due to the radial compliance mismatch, Comparative Example 1 exhibited severe hemodynamic disturbances and a "stress shielding" effect at the vascular anastomosis, stimulating excessive proliferation of smooth muscle cells, thus verifying the importance of double stretching in resolving mechanical mismatch.

[0118] Comparative Example 2 Comparative Example 2 is basically the same as Example 1, except that it has no magnetic elastic sensing layer.

[0119] Specifically, (1) PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:25 to obtain an inner layer premix; PTFE dispersion resin and Isopar G organic solvent are mixed evenly at a mass ratio of 100:25 to obtain an outer layer premix.

[0120] The artificial blood vessels prepared in Example 1 and Comparative Example 2 were subjected to performance tests, and the results are shown in Table 2 below.

[0121] Table 2

[0122] As shown in Table 2, due to the lack of elasticity from the elastic polymer in Comparative Example 2, the radial compliance of its base tube is lower than that of Example 1. In the in vitro simulated stenosis experiment, the blood vessel of Example 1 generated a significant magnetic field frequency shift upon deformation; Comparative Example 2 showed no magnetic response and could not be used for wireless monitoring. Comparative Example 2 can only detect stenosis through microscopy or subsequent angiography, failing to provide early warning, while this invention solves the problem of difficult postoperative monitoring.

[0123] Comparative Example 3 Comparative Example 3 is basically the same as Example 1, except that it has no electrospinning layer.

[0124] Specifically, instead of performing electrospinning in step (3), the D-ePTFE base tube from step (2) was subjected to plasma pretreatment in step (4) using the same parameters, and a polydopamine layer and an anticoagulant layer were loaded onto its inner surface in sequence.

[0125] The artificial blood vessels prepared in Example 1 and Comparative Example 3 were subjected to performance tests, and the results are shown in Table 3 below.

[0126] Table 3

[0127] As shown in Table 3, since Comparative Example 3 lacks an electrospun layer, the anchoring density of the PDA layer on the pure ePTFE nodes is slightly lower. Therefore, the Hep-ASA grafting amount in Comparative Example 3 is slightly lower than that in Example 1. The platelet adhesion rate in Comparative Example 3 is slightly higher, but still significantly lower than that in pure ePTFE (platelet adhesion rate >20%), further indicating that the dual anticoagulation mechanism of the inner surface of Comparative Example 3 is independently effective. The outer electrospun layer mainly contributes to improving radial blood flow compliance and adventitia tissue integration, indirectly assisting in antithrombosis rather than directly determining the inner surface load. Meanwhile, since the nanofiber layer in Example 1 simulates the extracellular matrix (ECM) topology of natural blood vessels, it guides the extension and spread of endothelial cell pseudopodia; while the smooth PTFE surface of Comparative Example 3 makes cell adhesion difficult, proliferation slow, and difficult to form a functional endothelial layer, resulting in a low endothelial cell proliferation rate.

[0128] Comparative Example 4 Comparative Example 4 is basically the same as Example 1, except that the anticoagulant used is different.

[0129] Specifically, in step (4), the inner surface of the composite base tube obtained in step (3) is subjected to plasma pretreatment. The treatment conditions are: pure oxygen is used, the oxygen flow rate is 25 sccm, the working pressure is 25 Pa, the radio frequency power is 100 W, the treatment time is 90 s, and the substrate temperature is controlled below 50 °C. The inner surface of the plasma-pretreated composite tube was then immersed in a dopamine-Tris buffer solution (pH=8.5) with a dopamine concentration of 2.5 mg / mL and allowed to stand in the dark at room temperature (25℃) for 4 hours to allow dopamine to polymerize in situ on the inner surface of the tube, forming a uniform polydopamine layer with a thickness of about 20-40 nm, rich in catechol and free amino groups. Heparin sodium and para-aminoacetylsalicylic acid were then directly dissolved in MES buffer at pH 6.0 to prepare a mixed solution (wherein the concentration of heparin sodium and para-aminoacetylsalicylic acid in the mixed solution was 2.5 mg / mL, and the molar ratio of the carboxyl group of heparin sodium to the amino group of para-aminoacetylsalicylic acid was 1:1.5). This mixed solution was filled into the inner cavity of the composite tube loaded with the PDA layer and sealed. The tube was then slowly rotated at 30 rpm at room temperature (25℃) for 4 hours. After the reaction, the tube was thoroughly rinsed with deionized water and PBS and vacuum dried at 30℃ for 4 hours to obtain the biofunctionalized composite tube. The biofunctionalized composite base tube was placed in an ethylene oxide (EO) sterilizer and sterilized for 3 hours at 35°C, 50% RH relative humidity, and an EO concentration of 600 mg / L. Subsequently, it was desorbed at 50°C for 10 hours under ventilated conditions to remove residual EO monomers, and finally, a composite expanded polytetrafluoroethylene artificial blood vessel was obtained.

[0130] The artificial blood vessels prepared in Example 1 and Comparative Example 4 were subjected to performance tests, and the results are shown in Table 4 below.

[0131] Table 4

[0132] Drug release behavior was measured in PBS buffer at 37°C. As shown in Table 4, the initial drug loading of both samples was similar, but the drug distribution in Comparative Example 4 was uneven, and it contained a large amount of non-covalently bound drug. The physically mixed drug in Comparative Example 4, relying solely on van der Waals forces or electrostatic adsorption, was prone to significant detachment (burst release) upon initial contact with bodily fluids. In contrast, the chemically coupled Hep-ASA in Example 1, linked by stable amide bonds, released extremely slowly, providing long-term anticoagulant protection. Furthermore, Comparative Example 4 initially showed a prolonged APTT due to burst release; however, as the drug rapidly depleted, its anticoagulant activity was almost completely lost after 7 days, returning to near-bare tube levels. Example 1, on the other hand, maintained a high APTT value for a long period. After burst release, the aspirin concentration in Comparative Example 4 rapidly decreased, failing to sustain platelet aggregation inhibition; simultaneously, the detached drug molecules might cause excessively high local concentrations, inducing thrombocytopenia. In Example 1, the heparin fragment of Hep-ASA stably binds to endothelial growth factor, promoting endothelialization. In Comparative Example 4, the drug was physically mixed, and due to rapid loss and high concentrations of free aspirin, cell proliferation may be inhibited, leading to delayed endothelialization. Moreover, in the early postoperative period, Comparative Example 4 experienced systemic side effects or local ineffective concentrations due to drug burst release, and in the later period, the anticoagulant protection was lost due to drug depletion, ultimately resulting in a high thromboembolism rate.

[0133] Comparative Example 5 Comparative Example 5 is basically the same as Example 1, except that the ratio of the outer layer premix is ​​different.

[0134] Specifically, in step (1), for the outer layer premix, PTFE dispersion resin, Isopar G organic solvent, Terfenol-D high magnetostrictive magnetic microparticles with a particle size of 3 μm and isocyanate-terminated polyurethane prepolymer are mixed at a mass ratio of 100:25:10:18 to obtain the outer layer premix.

[0135] Comparative Example 6 Comparative Example 6 is basically the same as Example 1, except that the stretching ratio of the second axial stretch is different.

[0136] Specifically, in step (2), for the second axial stretching, the second axial stretching is performed at 250°C, the stretching speed is 80 mm / min, and the stretching ratio is 6.5 times.

[0137] Comparative Example 7 Comparative Example 7 is basically the same as Example 1, except that the thickness of the polydopamine layer is different.

[0138] Specifically, in step (4), for the loaded polydopamine layer, the inner surface of the plasma-pretreated composite base tube is immersed in a dopamine-Tris buffer solution (pH=8.5) with a dopamine concentration of 2.5 mg / mL, and allowed to stand in the dark at room temperature (25°C) for 10 h to allow dopamine to polymerize in situ on the inner surface of the tube to form a polydopamine layer.

[0139] Comparative Example 8 Comparative Example 8 is basically the same as Example 1, except that the concentration of anticoagulant used in the loaded anticoagulant layer is different.

[0140] Specifically, in step (4), for the anticoagulant-loaded layer, the purified Hep-ASA complex was dissolved in MES buffer at pH=6.0 to prepare a 6 mg / mL solution. EDC and NHS were added, and the solution was activated for 20 min to obtain an activation solution. The concentration of EDC in the activation solution was 10 mM, and the concentration of NHS was 5 mM. The activation solution was filled into the inner cavity of the composite tube loaded with the PDA layer and sealed. The tube was slowly rotated at 30 rpm at room temperature (25 °C) for 4 h to allow the amino groups on the PDA layer to be covalently linked to the carboxyl groups on the Hep-ASA complex through amide bonds. After the reaction was completed, the tube was thoroughly rinsed with deionized water and PBS and vacuum dried at 30 °C for 4 h to obtain the biofunctionalized composite tube.

[0141] The artificial blood vessels prepared in Example 1 and Comparative Examples 5 to 8 were subjected to performance tests, and the results are shown in Table 5 below.

[0142] Table 5

[0143] As shown in Table 5, in Comparative Example 5, the addition of excessive elastic polymer resulted in the formation of an excessive rubbery continuous phase after sintering, severely weakening the load-bearing structure of the PTFE node-fiber, leading to a significant decrease in burst pressure. Furthermore, the excessive elastomer caused signal hysteresis. In Comparative Example 6, the excessively high secondary stretching ratio led to overstretching and even breakage of the fibrils, resulting in insufficient macroscopic strength of the vessel wall and a decrease in burst pressure. Moreover, the overstretching caused uneven vessel wall structure; when the vessel deformed, the signal and diameter changes were no longer linearly related, making it impossible to quantitatively analyze the degree of stenosis. In Comparative Example 7, the in-situ polydopamine generation reaction time was too long, causing a large amount of dopamine to self-polymerize in the solution, forming large particles deposited on the surface instead of a uniform film formed by in-situ polymerization. This not only blocked the micropores of the ePTFE but also resulted in extremely rough surfaces. Simultaneously, the excessively thick PDA layer altered surface hardness and inhibited the pseudopodia extension of endothelial cells. In Comparative Example 8, the excessively high Hep-ASA concentration caused the surface molecular density to exceed the critical value, resulting in a "stereohindrance" effect that hindered the binding of cell adhesion receptors. At the same time, considering cytotoxicity, this excessive anticoagulation was accompanied by a deterioration in biocompatibility, which could easily lead to bleeding tendency after implantation.

[0144] Comparative Example 9 Comparative Example 9 is basically the same as Example 1, except that the concentration of anticoagulant used in the loaded anticoagulant layer is different.

[0145] 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 4mm and a total wall thickness that is the same as that of the D-ePTFE base tube in Example 1 is obtained. The three-layer composite ePTFE artificial blood vessel base tube was electrospun and heat-treated under the same conditions as step (3) of Example 1. Then, the plasma pretreatment, loading of polydopamine layer and anticoagulant drug layer were carried out under the same conditions as step (4) of Example 1 to obtain the final composite artificial blood vessel.

[0146] The artificial blood vessels prepared in Example 1 and Comparative Example 6 were subjected to performance tests, and the results are shown in Table 6 below.

[0147] Table 6 As shown in Table 6, Comparative Example 9, using uniaxial stretching and preform pressing, has a relatively brittle intermediate dense layer (porosity 30%~35%), and its node bonding is weaker than that of Example 1, which uses biaxial stretching to form an inner dense and outer sparse fibrous interlocking structure. Furthermore, Example 1 has an outer layer with an elastic polymer mesh to enhance node toughness. Comparative Example 9 only uses uniaxial stretching, resulting in high radial stiffness. Example 1 uses a first low-temperature stretching to establish the skeleton and a second high-temperature stretching to expand the outer layer porosity; the elastic polymer reduces stiffness, and its radial compliance is close to that of natural blood vessels. Comparative Example 9's intermediate dense layer provides rebound hemostasis, but the outer layer lacks elastic polymer cushioning, leading to slightly higher puncture stress concentration. Example 1, on the other hand, can synergistically heal itself, resulting in slightly faster hemostasis. Comparative Example 9 uses integrated preform pressing and preform pressing, but the interfaces of different compression ratios rely solely on physical compression, lacking the co-extrusion in-situ interlocking and elastic polymer bridging of Example 1. Moreover, Comparative Example 9 relies on pure PTFE node rebound, which is relatively slow. The outer layer of elastic polymer in Example 1 provides rebound recovery force, resulting in slightly faster rebound.

[0148] It should be noted that the anastomotic endometrial hyperplasia rate was obtained from a 4-week animal experiment. The long-term patency rate was also obtained from a 4-week animal experiment. Drug release behavior was measured in PBS buffer at 37°C.

[0149] The composite 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 7. 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); 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.

[0150] Table 7

[0151] As shown in Table 7, the composite expanded polytetrafluoroethylene artificial blood vessel prepared by this invention has excellent biomimetic mechanical properties, long-lasting dual anticoagulation function, rapid endothelialization ability, and wireless stenosis monitoring function.

[0152] 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. The parts of the present invention not described in detail are techniques known to those skilled in the art.

Claims

1. A method for preparing a composite expanded polytetrafluoroethylene artificial blood vessel, characterized in that, include: (1) The inner layer premix containing polytetrafluoroethylene and organic solvent, and the outer layer premix containing polytetrafluoroethylene, organic solvent, magnetic particles and elastic polymer precursor are all added to the extruder for co-extrusion to obtain a green tube. (2) The green tube is subjected to a first axial stretching, sintering and shaping, and a second axial stretching in sequence to obtain a base tube; (3) Using the base tube as a receiving device, electrospinning is performed using a spinning solution containing polytetrafluoroethylene and spinning aids, followed by heat treatment to obtain a composite base tube. (4) After plasma pretreatment of the inner surface of the composite base tube, a polydopamine layer and an anticoagulant drug layer are loaded sequentially on the inner surface to obtain the composite expanded polytetrafluoroethylene artificial blood vessel.

2. The preparation method according to claim 1, characterized in that, In step (1): The mass ratio of polytetrafluoroethylene to organic solvent in the inner layer premix is ​​100:(15~35); The mass ratio of polytetrafluoroethylene, organic solvent, magnetic particles, and elastic polymer precursor in the outer premix is ​​100:(15~35):(5~15):(3~10); preferably, the content of organic solvent in the outer premix is ​​higher than the content of organic solvent in the inner premix; and / or, The organic solvent is a hydrocarbon solvent, preferably n-hexane, n-heptane, n-octane, petroleum ether, cyclohexane, or methylcyclohexane; The magnetic particles are at least one of terbium-dysprosium-iron alloy particles, amorphous alloys, Fe-Ga alloys, and Co-Fe alloys; preferably, the particle size is 1~5μm. The elastic polymer precursor is at least one of the following: polyurethane prepolymer, isocyanate-terminated polyurethane prepolymer, polydimethylsiloxane prepolymer, silicone rubber prepolymer, polyetheretherketone prepolymer, and polylactic acid-glycolic acid copolymer prepolymer.

3. The preparation method according to claim 1, characterized in that, Step (1) includes: (11) The inner layer premix is ​​extruded to form an inner layer blank; the cross-section of the inner layer blank is a first ring; (12) The outer premix is ​​extruded to form an outer blank; the cross-section of the outer blank is a second ring; the inner diameter of the second ring matches the outer diameter of the first ring so that the outer blank can coaxially fit the inner blank; (13) The inner layer blank is coaxially inserted into the outer layer blank, and a plunger extruder is used to press it into a mold composed of a mandrel and a die for co-extrusion to obtain the green tube with an inner layer and an outer layer. Preferably, the compression ratio used for the inner layer is (30~100):1; and the compression ratio used for the outer layer is (70~200):

1.

4. The preparation method according to claim 3, characterized in that, The mandrel is a cylindrical mandrel or a tapered mandrel; the tapered mandrel is a continuously adjustable tapered mandrel or a segmented tapered mandrel, so that a non-equidistant annular flow channel is formed between the mandrel and the die, and the taper of the mandrel in the inlet section is greater than the taper of the outlet section.

5. The preparation method according to claim 1, characterized in that, In step (2): The first axial stretching is performed at a stretching speed of 50-200 m / min, a stretching temperature of 80-150℃, and a draw ratio of 1.5-3 times; and / or, The sintering and shaping temperature is 330~380℃, and the time is 5~15min; Preferably, the stretching speed of the second axial stretching is 30~150m / min, the stretching temperature is 200~280℃, and the draw ratio is 2~5 times.

6. The preparation method according to claim 1, characterized in that, In step (3): The mass ratio of polytetrafluoroethylene resin to spinning aid in the spinning solution is 1:(0.5~2), preferably 1:(0.8~1.2); The spinning aid is at least one of polyvinyl alcohol and polyethylene glycol; preferably, the degree of alcoholysis of polyvinyl alcohol is ≥88%, and the number average molecular weight of polyethylene glycol is 4000~20000; and / or, The electrospinning voltage is 25~40kV, the receiving distance is 12~18cm, the extrusion speed is 1~2mL / h, and the receiving roller speed is 800~1500rpm; Preferably, the heat treatment temperature is 385~395℃ and the time is 15~25min.

7. The preparation method according to claim 1, characterized in that, In step (4): The plasma pretreatment uses oxygen or a mixture of oxygen and argon, with a gas flow rate of 20-30 sccm, a radio frequency power of 80-120 W, a working pressure of 20-30 Pa, and a time of 60-120 s; and / or, The loaded polydopamine layer comprises: immersing the inner surface of the composite base tube, which has undergone plasma pretreatment, in a buffer solution containing dopamine for reaction; preferably, the concentration of dopamine in the buffer solution is 2~3 mg / mL, and the reaction time is 3~5 h.

8. The preparation method according to any one of claims 1 to 7, characterized in that, In step (4): The anticoagulant-loaded layer comprises: activating an anticoagulant buffer solution by adding 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide to the anticoagulant buffer solution to obtain an activation solution; injecting the activation solution into the inner surface of the composite base tube loaded with the polydopamine layer for reaction; preferably, the concentration of the anticoagulant in the anticoagulant buffer solution is 2~3 mg / mL, the concentration of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide in the activation solution is 8~12 mM, and the concentration of N-hydroxysuccinimide is 4~6 mM; more preferably, the activation time is 15~30 min, and the reaction time is 3~5 h.

9. The preparation method according to claim 8, characterized in that, The anticoagulant drug is obtained by the following method: Sodium heparin is dissolved in MES buffer, and then 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide are added and stirred at 25-28°C for 1-2 h to obtain an activated solution; preferably, the molar ratio of the carboxyl group of sodium heparin to 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide is 1:(2-5):(1-3); An amino-containing aspirin derivative is added to an activation solution and reacted at 25-28°C in the dark for 12-24 hours to obtain the reaction product; preferably, the molar ratio of the carboxyl group of heparin sodium to the amino group of the aspirin derivative in the activation solution is 1:(0.5-3). The reaction product was transferred into a dialysis bag with a molecular weight cutoff of 3500-8000 Da and dialyzed in deionized water at 4-8°C for 48-60 hours. After dialysis, the product was freeze-dried to obtain the anticoagulant.

10. A composite expanded polytetrafluoroethylene artificial blood vessel prepared by any one of the preparation methods described in claims 1 to 9, characterized in that, The composite expanded polytetrafluoroethylene artificial blood vessel consists of, from the inside out, a biofunctional layer including an anticoagulant drug layer and a polydopamine layer, a base tube with a double-layer structure, and an electrospinning layer.

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