Anti-tumor degradable ureteral stent with bionic structure and preparation method of anti-tumor degradable ureteral stent

Abstract

The invention belongs to the technical field of medical stents, and particularly relates to an anti-tumor degradable ureteral stent with a bionic structure and a preparation method of the anti-tumor degradable ureteral stent. The anti-tumor degradable ureteral stent with the bionic structure comprises a single-layer tubular stent body or a multi-layer tubular stent body, the single-layer tubular stent body is made of aliphatic polyester polymers and drugs, the multi-layer tubular stent body comprises an inner tube and an outer tube which are coaxial, the inner tube is made of polyvinyl alcohol and drugs, and the outer tube is made of polyethylene glycol. And the outer pipe is made of an aliphatic polyester polymer. The anti-tumor degradable ureteral stent with the bionic structure provided by the invention has flexibility required for adapting to dynamic peristalsis of the ureter, is not easy to cause mechanical damage to urinary tract tissues, has dual functions of anti-tumor activity and surface wettability, can realize synchronous matching of a drug release rate and a material degradation rate, and has a good application prospect. Meanwhile, the mechanical anti-fatigue performance of the stent is not sacrificed.

Description

A biomimetic biodegradable ureteral stent for tumor treatment and its preparation method Technical Field

[0001] This invention belongs to the field of medical stent technology, specifically relating to an anti-tumor biodegradable ureteral stent with a biomimetic structure and its preparation method. Background Technology

[0002] Ureteral stents (also known as double-J stents) are widely used auxiliary devices in urological surgery, mainly for relieving ureteral obstruction, draining urine, and supporting ureteral anastomoses. They are particularly useful in the treatment of malignant tumors such as urothelial carcinoma, where they are often used to assist drainage and postoperative recovery.

[0003] Currently, ureteral stents widely used in clinical practice are mainly made of non-degradable polymer materials such as polyurethane (PU) and silicone rubber. Although these materials have good biostability and mechanical properties, they present significant technical problems in clinical application: 1. Requires secondary surgery for removal: Due to the non-degradable nature of the materials, patients must undergo a painful cystoscopic removal surgery after the stent placement period (usually 2 weeks to several months). This not only increases the patient's physical pain and financial burden but may also lead to complications such as urinary tract infections and bleeding.

[0004] 2. Long-term indwelling complications: Long-term indwelling non-degradable stents can easily lead to urinary salt deposits (stone adhesion), bacterial biofilm formation, and urinary tract irritation symptoms, which seriously affect the patient's quality of life.

[0005] To address these issues, biodegradable ureteral stents (such as PLGA-based stents) have become a research hotspot. However, existing biodegradable stent technologies still face a fundamental contradiction: an imbalance between the durability of mechanical support and degradation kinetics. Amorphous PLGA with low lactate content (such as LA:GA=50:50) degrades too quickly and is prone to "burst" acid hydrolysis, causing the stent to collapse prematurely within the ureteral repair window (3-4 weeks) and lose its supporting function. While PLGA with high lactate content has better stability, it often has excessive stiffness and lacks the flexibility required to adapt to the dynamic peristalsis of the ureter, easily causing mechanical damage to urinary tract tissues.

[0006] Furthermore, for the treatment of ureteral tumors, traditional systemic chemotherapy (such as intravenous gemcitabine) suffers from problems such as rapid drug metabolism (short half-life), significant systemic toxicity, and low drug concentration at the local lesion. Although some studies have attempted to use drug-loaded stents for local treatment, how to simultaneously match the drug release rate with the material degradation rate without sacrificing the mechanical fatigue resistance of the stent remains an unsolved technical challenge in this field. Summary of the Invention

[0007] In view of this, the purpose of the present invention is to provide an antitumor biodegradable ureteral stent with a biomimetic structure and its preparation method. The antitumor biodegradable ureteral stent with a biomimetic structure provided by the present invention has the flexibility required to adapt to the dynamic peristalsis of the ureter, is not prone to causing mechanical damage to the urinary tract tissue, has the dual functions of antitumor activity and surface wettability, and can achieve synchronous matching of drug release rate and material degradation rate, without sacrificing the mechanical fatigue resistance of the stent.

[0008] To achieve the above objectives, the present invention provides the following technical solution: The present invention provides an antitumor biodegradable ureteral stent with a biomimetic structure, comprising a single-layer tubular stent or a multi-layer tubular stent; the single-layer tubular stent is made of an aliphatic polyester polymer and a drug; the multi-layer tubular stent comprises a coaxial inner tube and an outer tube; the inner tube is made of polyvinyl alcohol and a drug; the outer tube is made of an aliphatic polyester polymer; the drug comprises an antitumor drug and / or a protein drug; the aliphatic polyester polymer comprises at least one of polylactic acid-glycolic acid copolymer, polycaprolactone, polylactic acid, and polyethylene glycol-polylactic acid-glycolic acid copolymer; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer is 60~85:15~40; the mass of the drug in the single-layer tubular stent is 5~25% of the mass of the aliphatic polyester polymer; the mass of the drug in the multi-layer tubular stent is 5~25% of the mass of the aliphatic polyester polymer.

[0009] Preferably, the antitumor drug includes at least one of gemcitabine, paclitaxel, cisplatin, pirarubicin, and doxorubicin.

[0010] Preferably, the number average molecular weight of the polycaprolactone is not higher than 50,000.

[0011] Preferably, the polylactic acid is a copolymer of L-lactic acid and D-lactic acid, wherein the molar percentage of D-lactic acid is 15-25%; and the intrinsic viscosity of the polylactic acid is 0.8-1.6 dL / g.

[0012] Preferably, the polyethylene glycol segment in the polyethylene glycol-polylactic acid-glycolic acid copolymer has a molecular weight of 1,000 to 5,000 and a mass percentage of 5 to 15% in the polyethylene glycol-polylactic acid-glycolic acid copolymer, and the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer block is 85 to 90: 10 to 15.

[0013] This invention also provides a method for preparing the biomimetic biodegradable ureteral stent with biomimetic structure described in the above technical solution. When the biomimetic biodegradable ureteral stent with biomimetic structure is a single-layer tubular stent, the preparation method includes the following steps: firstly spinning a sacrificial layer spinning solution on a mandrel to form a sacrificial layer film; secondly spinning a drug-loaded spinning solution on the sacrificial layer film to form a composite film; the drug-loaded spinning solution includes an aliphatic polyester polymer, a drug, and a first organic solvent; removing the sacrificial layer film from the composite film to obtain a single-layer tubular stent; when the biomimetic biodegradable ureteral stent with biomimetic structure is a multilayer tubular stent, the preparation method includes the following steps: coaxially electrospinning an inner tube spinning solution and an outer tube spinning solution together to obtain a multilayer tubular stent; the inner tube spinning solution includes polyvinyl alcohol, a drug, and a solvent; the outer tube spinning solution includes an aliphatic polyester polymer and a third organic solvent.

[0014] Preferably, the sacrificial layer spinning solution includes a sacrificial layer material and a second organic solvent; the sacrificial layer material is one or more of polyvinylpyrrolidone, polyethylene oxide, water-soluble gelatin, and polyvinyl alcohol.

[0015] Preferably, the conditions for the first spinning include: the pump head inner diameter of the injection pump is 0.5~1.0mm, the voltage is 20~25kV, the injection speed is 20~30µL / min, the receiving speed is 1000~2000rpm, and the spinning time is 30~90min.

[0016] Preferably, the conditions for the second spinning include: a voltage of 20~25kV, a feeding speed of 20~30µL / min, a receiving speed of 2000~3000rpm, and a spinning time of 2~3h.

[0017] Preferably, the conditions for coaxial electrospinning are as follows: the columnar collector is made of stainless steel with a diameter of 0.7~1.5mm and a rotation speed of 2000~3000rpm; the flow rate of the spinning solution in the outer tube is 20~30µL / min and the flow rate of the spinning solution in the inner tube is 5~15µL / min; the flow rate ratio of the spinning solution in the outer tube to the spinning solution in the inner tube is 2~5:1; the DC voltage is 20~25kV; the receiving distance is 10~20cm; the temperature is 25±2℃; and the relative humidity is 50±5%.

[0018] This invention provides a biomimetic biodegradable ureteral stent with a biomimetic structure, comprising a single-layer tubular stent or a multi-layer tubular stent; the single-layer tubular stent is made of an aliphatic polyester polymer and a drug; the multi-layer tubular stent comprises a coaxial inner tube and an outer tube; the inner tube is made of polyvinyl alcohol and a drug; the outer tube is made of an aliphatic polyester polymer; the drug comprises an antitumor drug and / or a protein drug; the aliphatic polyester polymer comprises at least one selected from polylactic acid-glycolic acid copolymer, polycaprolactone, polylactic acid, and polyethylene glycol-polylactic acid-glycolic acid copolymer; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer is 60-85:15-40; the mass of the drug in the single-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer; the mass of the drug in the multi-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer.

[0019] Compared with the prior art, the technical solution provided by the present invention has the following beneficial effects: 1. Wide range of adjustable degradation kinetics to meet the needs of different clinical treatment cycles. Advantage description: The ureteral stent prepared by the present invention is not limited to a single degradation cycle, but can achieve precise control of the degradation rate in the range of 3 to 12 weeks according to the clinical needs of different conditions (such as short-term postoperative support or long-term anti-stenosis).

[0020] Explanation: This effect stems from the lactic acid / glycolic acid (LA:GA) molar ratio of the PLGA substrate, which ranges from 60:40 to 85:15. Within this range, by adjusting the monomer ratio, the transformation of the material's microstructure from an amorphous to a semi-crystalline state can be precisely controlled. This change in crystallinity alters the water molecule penetration pathway, allowing the scaffold's degradation mode to switch between "erosion-dominated" and "swelling-dominated," thus enabling on-demand customization of the mechanical support window.

[0021] 2. Description of the dual functional regulation advantages of anti-tumor activity and surface wettability: The stent not only has the ability to kill urothelial carcinoma cells, but can also adjust the hydrophilic and hydrophobic properties of the surface according to the different drug loading, thereby balancing the needs of "drug burst release killing" and "long-term prevention of adhesion".

[0022] Explanation of the cause: This effect is due to the fact that the loading of anti-tumor drugs such as gemcitabine is set at 5-25% in the technical solution.

[0023] Antitumor activity: Within this concentration range, gemcitabine exhibits significant dose-dependent cytotoxicity and can effectively inhibit cancer cell survival.

[0024] Surface modification: The polar groups of gemcitabine can significantly improve the surface energy of PLGA substrates. As the drug loading varies from 5% to 25%, the water contact angle of the scaffold surface can be adjusted significantly from 0° (superhydrophilic) to 115° (hydrophobic). High hydrophilicity can be selected to reduce tissue adhesion, or moderate hydrophobicity can be selected to maintain longer-term structural stability, depending on the needs.

[0025] 3. Description of the high damage tolerance advantage brought by the biomimetic "fiber bridging" mechanism: Regardless of the ratio, the stent exhibits excellent radial flexibility and fracture resistance, can adapt to the dynamic peristaltic environment of the ureter, and avoids brittle disintegration.

[0026] Explanation: This effect stems from the electrospinning process and tubular fiber network structure employed in this invention. By controlling the receiving rotation speed (1500~3500 rpm), the nanofibers form a dense nonwoven network. This structure mimics the "fiber bridging" mechanism of natural lotus stems. Under radial pressure, the fibers undergo synergistic slippage and micro-deformation to dissipate energy, giving the material superior fatigue resistance compared to traditional cast scaffolds.

[0027] 4. Universal and non-destructive demolding and inner wall quality control advantages: The preparation process of this invention solves the industry problem of the difficulty in completely removing electrospun tubular scaffolds from the mandrel, ensuring that scaffolds of various formulations have uniform tube diameter and smooth inner wall without mechanical damage.

[0028] Explanation: This effect stems from the two-step spinning process assisted by the PVP sacrificial layer. Regardless of the thickness or formulation of the outer PLGA drug-loaded scaffold, the inner PVP sacrificial layer can utilize its rapid solubility in water / alcohol to achieve quick and gentle separation of the scaffold from the mold, avoiding microstructural damage caused by forced mechanical demolding. Attached Figure Description

[0029] Figure 1 shows the physical image and SEM image of the single-layer tubular stent prepared according to the present invention. A shows the actual tubular ureter prepared by electrospinning, where (a) is the overall macroscopic morphology of the ureter stent, (b) is the inner diameter measurement of the ureter stent, and (c) is the cross-sectional view of the port of the ureter stent. B and C are biomimetic physical objects. D shows the morphological changes of fibers under different drug loading (0%, 5%, 15%, 20%, 25%), where (a) is PLGA-Gem 0%, (b) is PLGA-Gem 5%, (c) is PLGA-Gem 10%, (d) is PLGA-Gem 15%, (e) is PLGA-Gem 20%, and (f) is PLGA-Gem 25%. Figure 2 shows the XRD diffraction patterns of stents with different LA:GA ratios (50:50, 60:40, 70:30, 80:20). Figure 3 shows the XRD diffraction patterns of stents with different LA:GA ratios (50:50, 60:40, 70:30, 80:20). Figure 4 shows the DSC thermal analysis curves of stents with drug loading concentrations of 0:40, 70:30, and 80:20; Figure 5 shows the water contact angle test results of stents with different drug loading concentrations, where (a) is 0% PLGA-Gem, (b) is 5% PLGA-Gem, (c) is 10% PLGA-Gem, (d) is 15% PLGA-Gem, (e) is 20% PLGA-Gem, and (f) is 25% PLGA-Gem; Figure 6 shows the mass loss curve of the stent during degradation in artificial urine; Figure 7 shows the cell viability (CCK-8) bar chart of the stent co-cultured with 5637 urothelial carcinoma cells; Figure 8 shows the physical image of the ureteral stent in Comparative Example 1; Figure 9 shows the SEM image of the ureteral stent prepared in Comparative Example 3, where a is 1000 rpm, b is 1500 rpm, and c is 3000 rpm. Detailed Implementation

[0030] This invention provides a biomimetic biodegradable ureteral stent with a biomimetic structure, comprising a single-layer tubular stent or a multi-layer tubular stent; the single-layer tubular stent is made of an aliphatic polyester polymer and a drug; the multi-layer tubular stent comprises a coaxial inner tube and an outer tube; the inner tube is made of polyvinyl alcohol and a drug; the outer tube is made of an aliphatic polyester polymer; the drug comprises an antitumor drug and / or a protein drug; the aliphatic polyester polymer comprises at least one selected from polylactic acid-glycolic acid copolymer, polycaprolactone, polylactic acid, and polyethylene glycol-polylactic acid-glycolic acid copolymer; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer is 60-85:15-40; the mass of the drug in the single-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer; the mass of the drug in the multi-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer.

[0031] Unless otherwise specified, the present invention does not have special requirements on the source of raw materials used, and commercially available products well known to those skilled in the art can be used.

[0032] In one embodiment, the aliphatic polyester polymer includes at least one of polylactic acid-glycolic acid copolymer (PLGA), polycaprolactone (PCL), polylactic acid (PLA), and polyethylene glycol-polylactic acid-glycolic acid copolymer (PEG-PLGA), with polylactic acid-glycolic acid copolymer (PLGA) being a specific example.

[0033] In one embodiment, the molar ratio of lactic acid (LA) to glycolic acid (GA) in the polylactic-glycolic acid copolymer (PLGA) is 60~85:15~40, and in specific embodiments it is 60:40, 70:30 or 80:20.

[0034] In a specific embodiment of the present invention, the polylactic acid-glycolic acid copolymer was purchased from Shenzhen Polysun Technology Co., Ltd., and its product name is poly(lactic acid-CO-glycolic acid), with the molecular formula (C6H8O4). n The number average molecular weight is 213,000~600,000.

[0035] In one embodiment, the number average molecular weight (Mn) of the polycaprolactone (PCL) is not higher than 50,000, or PCL is blended with PLGA 50 / 50 or PDLLA (polymers with a degradation cycle of less than 12 weeks) at a mass ratio of 60~70:30~40 to form a continuous phase of PCL (rapidly degradable polymer).

[0036] In one embodiment, to obtain an amorphous structure to accelerate degradation, the polylactic acid (PLA) is a copolymer (PDLLA) of L-lactic acid (L-LA) and D-lactic acid (D-LA), wherein the molar percentage of D-lactic acid is 15-25%; and the intrinsic viscosity of the polylactic acid (PLA) is 0.8-1.6 dL / g.

[0037] In one embodiment, the polyethylene glycol (PEG) segment in the polyethylene glycol-polylactic acid-glycolic acid copolymer (PEG-PLGA) has a molecular weight of 1,000 to 5,000 and a mass percentage of 5 to 15% in the polyethylene glycol-polylactic acid-glycolic acid copolymer, and the molar ratio of lactic acid (LA) to glycolic acid (GA) in the polylactic acid-glycolic acid copolymer (PLGA) block is 85 to 90: 10 to 15.

[0038] When the aliphatic polyester polymer is a blend of multiple aliphatic polyester polymers, the aliphatic polyester polymer with the fastest degradation rate in the resulting blend constitutes the continuous phase, and its mass percentage is not less than 60%.

[0039] Materials such as PCL can also be electrospun and possess good flexibility. Although their degradation rate differs from PLGA, this invention regulates the degradation cycle by adjusting the molecular weight or blending ratio of aliphatic polyester polymers. Generally, the higher the molecular weight, the more ester bonds need to be broken, resulting in slower degradation. Dense and ordered crystalline regions act as a barrier to degradation. Reducing crystallinity accelerates degradation. The stronger the hydrophilicity of the material, the easier it is for water molecules to penetrate, leading to faster hydrolysis. Polycaprolactone (PCL) can be copolymerized with PLA or PGA, or blended with fast-degrading polymers to disrupt its highly crystalline structure; polylactic acid (PLA) can be copolymerized, particularly by introducing D-type lactic acid monomers to form an amorphous copolymer (PDLLA); and polyethylene glycol-polylactic acid-glycolic acid copolymer (PEG-PLGA) can be synergistically regulated by controlling the length and content of PEG segments and the LA / GA ratio of PLGA blocks, thereby achieving regulation of the degradation cycle.

[0040] In one embodiment, the drug includes an antitumor drug and / or a protein drug, specifically an antitumor drug; the antitumor drug includes at least one of gemcitabine, paclitaxel, cisplatin, pirarubicin, and doxorubicin, specifically gemcitabine; the protein drug includes one or more of basic fibroblast growth factor (bFGF), vascular endothelial growth factor (VEGF), epidermal growth factor (EGF), transforming growth factor-β3 (TGF-β3), and urokinase, specifically basic fibroblast growth factor (bFGF) or vascular endothelial growth factor (VEGF).

[0041] In this embodiment of the invention, the loaded drug is the antimetabolite gemcitabine. The loaded drug can be replaced with other chemotherapy drugs suitable for the treatment of urinary system tumors, such as paclitaxel, cisplatin, pirarubicin, or doxorubicin. Electrospinning technology is universally applicable; whether the drug is hydrophilic (such as gemcitabine) or hydrophobic (such as paclitaxel), loading and sustained release can be achieved by adjusting the solvent system or using coaxial spinning technology.

[0042] In one embodiment, the mass of the drug in the single-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer, specifically 5, 10, 15, 20 or 25% in the embodiment; the mass of the drug in the multi-layer tubular stent is 5-25% of the mass of the aliphatic polyester polymer, specifically 5, 10, 15, 20 or 25% in the embodiment.

[0043] In one embodiment, the inner diameter of the single-layer tubular stent is 1.5~2.5mm, specifically 2.0mm in this embodiment, and the wall thickness is 0.3~0.7mm, specifically 0.5mm in this embodiment; in the multi-layer tubular stent, the inner diameter of the outer tube is 2.1~3.5mm, specifically 2.5mm in this embodiment, and the wall thickness is 0.1~0.4mm, specifically 0.25mm in this embodiment; the inner diameter of the inner tube is 1.5~2.5mm, specifically 2.0mm in this embodiment, and the wall thickness is 0.1~0.4mm, specifically 0.25mm in this embodiment.

[0044] Biomimetic biodegradable ureteral stent for tumor prevention: 1. Matrix material composition: Polylactic acid-glycolic acid copolymer (PLGA) is used as the matrix.

[0045] Key parameter range: The molar ratio of lactic acid (LA) to glycolic acid (GA) in the PLGA is 60:40~85:15.

[0046] Technical principle: Within this range, the material can form a nanofiber network with a certain degree of orientation through electrospinning, and provide a controllable crystalline structure from amorphous to semi-crystalline state according to different LA contents, thereby meeting the 3 to 12-week degradation window required for ureter repair.

[0047] This invention achieves precise control over the scaffold's properties by creatively adjusting the monomer ratio of LA and GA in PLGA. As a crystalline monomer, LA's content directly determines the regularity and crystallinity of the copolymer's molecular chains, thereby controlling the material's hydrophilicity / hydrophobicity and mechanical modulus. Regarding the degradation cycle, the LA content directly dominates the scaffold's hydrolysis rate by influencing the amount of crystalline regions (acting as degradation barriers) and the hydrolysis sensitivity of ester bonds. Based on classical degradation models and previous studies, when the LA / GA ratio is adjusted from 50 / 50 to 90 / 10, the in vitro complete degradation time can be extended from approximately 4 weeks to over 24 weeks. This fully covers the critical time windows for ureteral epithelial repair (approximately 3-6 weeks) and more complete repair of the smooth muscle and fibrous tissue layers (up to 12 weeks), achieving a match between degradability and functional period.

[0048] 2. Active drug loading: Loaded with antimetabolite antitumor drugs, specifically gemcitabine.

[0049] Key parameter range: The mass percentage of the antitumor drug in the stent is 5-25%.

[0050] Technical principle: This concentration range (5~25%) can ensure that the drug is released in the early stage to kill tumor cells, and can also achieve sustained release in the later stage through the degradation of PLGA matrix; at the same time, the addition of drug molecules plays a plasticizing role to a certain extent, and adjusts the glass transition temperature (Tg) and hydrophilicity of the material.

[0051] 3. Microstructure characteristics: The scaffold is composed of interwoven nanofibers with an average fiber diameter ranging from 0.4 to 1.0 µm.

[0052] The fiber network exhibits a "fiber bridging" structure, which can dissipate energy through the cooperative slippage between fibers when subjected to radial pressure, thus providing fatigue resistance.

[0053] The surface water contact angle of the scaffold is 0~115°. The hydrophilic and hydrophobic properties of the surface can be controlled within this range by adjusting the drug content and the LA:GA ratio.

[0054] The LA:GA ratio determines the bulk chemical polarity and crystalline state of the polymer. Increasing the GA ratio enhances hydrophilicity by increasing the density of polar ester bonds and forming an amorphous structure. The addition of hydrophilic drugs further modifies the surface chemical composition at the nanoscale through the "surface segregation effect" during preparation and the direct introduction of polar functional groups. This synergistic effect allows for the modulation of scaffold surface properties over a wide range (e.g., water contact angle from ~50° to ~100°), thereby optimizing initial wettability, protein adsorption spectra, and cellular response after implantation. This creates a smart interface for ureteral repair that supports tissue regeneration while also possessing anti-tumor functions.

[0055] The present invention also provides a method for preparing the biomimetic antitumor biodegradable ureteral stent with biomimetic structure described in the above technical solution.

[0056] When the biomimetic antitumor biodegradable ureteral stent is a single-layer tubular stent, the preparation method includes the following steps: first spinning a sacrificial layer spinning solution on a mandrel to form a sacrificial layer film; second spinning a drug-loaded spinning solution on the sacrificial layer film to form a composite film; the drug-loaded spinning solution includes an aliphatic polyester polymer, a drug, and a first organic solvent; and removing the sacrificial layer film from the composite film to obtain a single-layer tubular stent.

[0057] In one embodiment, the sacrificial layer spinning solution includes a sacrificial layer material and a second organic solvent; the sacrificial layer material is one or more of polyvinylpyrrolidone (PVP), polyethylene oxide (PEO), water-soluble gelatin, and polyvinyl alcohol (PVA), with polyvinylpyrrolidone (PVP) being a specific example; the polyvinylpyrrolidone (PVP) is PVP K30; the number average molecular weight of the polyvinylpyrrolidone is 30,000~58,000, with 40,000~50,000 in a specific example; the second organic solvent is anhydrous ethanol; the mass-to-volume ratio of the sacrificial layer material in the sacrificial layer spinning solution is 5~25% (w / v), with 15.7% (w / v) in a specific example; the method for preparing the sacrificial layer spinning solution is to dissolve the sacrificial layer material in the second organic solvent and stir until completely dissolved.

[0058] In one embodiment, the first organic solvent is one or more of trifluoroacetic acid (TFA), trifluoroethanol (TFE), hexafluoroisopropanol (HFIP), dichloromethane, and dimethylformamide (DCM / DMF), with trifluoroacetic acid (TFA) being a specific example; the mass fraction of the aliphatic polyester polymer in the drug-loaded spinning solution is 15-25%, with 20% being a specific example; the drug-loaded spinning solution is prepared by dissolving the aliphatic polyester polymer in the first organic solvent, adding the drug, and stirring until completely dissolved; the stirring rate is 300-500 r / min, with 400 r / min being a specific example.

[0059] In addition to polyvinylpyrrolidone (PVP), other polymeric materials that are easily soluble in water or specific solvents, such as polyethylene oxide (PEO), water-soluble gelatin, or polyvinyl alcohol (PVA), can also be used as sacrificial layers.

[0060] Spinning solvents: In addition to trifluoroacetic acid (TFA), trifluoroethanol (TFE), hexafluoroisopropanol (HFIP), or dichloromethane / dimethylformamide (DCM / DMF) mixed solvents can also be used to dissolve PLGA.

[0061] The sacrificial layer material can dissolve rapidly to achieve demolding without damaging the outer PLGA structure; the solvent can dissolve PLGA and drugs well and is volatile.

[0062] In one embodiment, the mandrel is made of stainless steel and has a diameter of 1.5~3.0mm, specifically 2mm in this embodiment; the first spinning and the second spinning are two-step continuous electrostatic spinning; during the first spinning and the second spinning processes, the mandrel is in a rotating state; the present invention does not have a special limitation on the rotation rate, and a rotation rate well known in the art can be used.

[0063] In one embodiment, the thickness of the sacrificial layer film is 0.1~0.2mm, and in a specific embodiment it is 0.15~0.2mm.

[0064] In one embodiment, the conditions for the first spinning include: the inner diameter of the injection pump head is 0.5~1.0mm, specifically 0.84mm in this embodiment; the voltage is 20~25kV, specifically 23kV in this embodiment; the injection speed is 20~30µL / min, specifically 24µL / min in this embodiment; the receiving rotation speed is 1000~2000rpm, specifically 2000rpm in this embodiment; and the spinning time is 60~90min, specifically 70min in this embodiment.

[0065] In one implementation, the conditions for the second spinning include: a voltage of 20-25 kV, specifically 23 kV; a feed rate of 20-30 µL / min, specifically 24 µL / min; a receiving rotation speed of 2000-3000 rpm, specifically 2000 rpm; and a spinning time of 2-3 hours, specifically 125 minutes. The spinning time affects the wall thickness and can be controlled according to requirements.

[0066] As one implementation method, the sacrificial layer film in the composite film is removed by immersing the composite film in water; the water is deionized water; the immersion temperature is room temperature, and the time is 20-40 minutes, specifically 30 minutes in this embodiment. This invention utilizes the rapid dissolution of the sacrificial layer to achieve automatic detachment of the PLGA support.

[0067] In one embodiment, after removing the sacrificial layer film from the composite film, the process further includes: sequentially performing shaping drying and sterilization; the shaping drying involves inserting a support wire of the same diameter into the composite film after removing the sacrificial layer film and then drying it; the support wire is stainless steel wire; the drying equipment is a vacuum drying oven; the drying temperature is room temperature, the time is 12~24h, specifically 24h in this embodiment, and the vacuum degree is -0.095MPa (gauge pressure), i.e., 6kPa absolute pressure. This invention prevents film shrinkage through the support wire and removes residual solvent through drying; the sterilization is performed by ultraviolet lamp irradiation or immersion in medical alcohol; the intensity of the ultraviolet lamp irradiation is 40~100μW / cm². 2 In a specific embodiment, it is 70 μW / cm 2 The soaking time is 12-24 hours, with 24 hours in the specific embodiment; the soaking time in medical alcohol is 20-60 minutes, with 30 minutes in the specific embodiment.

[0068] When the biomimetic antitumor biodegradable ureteral stent is a multilayer tubular stent, the preparation method includes the following steps: coaxial electrospinning the inner tube spinning solution and the outer tube spinning solution together to obtain a multilayer tubular stent; the inner tube spinning solution includes polyvinyl alcohol, a drug and a solvent; the outer tube spinning solution includes an aliphatic polyester polymer and a third organic solvent.

[0069] In one embodiment, the solvent is water, specifically deionized water; the concentration of polyvinyl alcohol in the inner tube spinning solution is 5-12% (w / v), specifically 8% (w / v) in this embodiment, and the concentration of the antitumor drug is 0.5-5% (w / v), specifically 2% (w / v) in this embodiment; the inner tube spinning solution is prepared by dissolving polyvinyl alcohol and the drug in the solvent and magnetically stirring until completely dissolved; the present invention does not have a particular limitation on the rate and time of magnetic stirring, as long as a homogeneous and transparent solution can be obtained by complete dissolution.

[0070] In one embodiment, the third organic solvent is a mixture of dichloromethane and N,N-dimethylformamide; the volume ratio of dichloromethane to N,N-dimethylformamide is 6~9:1~4, and in a specific embodiment it is 8:2; the concentration of aliphatic polyester polymer in the outer tube spinning solution is 10~25% (w / v), and in a specific embodiment it is 18% (w / v); the preparation method of the outer tube spinning solution is to dissolve the aliphatic polyester polymer in the third organic solvent and stir magnetically until completely dissolved; the present invention does not have a special limitation on the rate and time of magnetic stirring, as long as a homogeneous and transparent solution can be obtained by complete dissolution.

[0071] In one embodiment, the coaxial electrospinning is performed by injecting the inner tube spinning solution and the outer tube spinning solution into two independent injection pumps and connecting them to coaxial spinning needles. The needle of the outer tube spinning solution is connected to the outer channel, and the needle of the inner tube spinning solution is connected to the inner channel. The needles are aligned with a rotating columnar collector for spinning, forming a tubular fiber membrane on the columnar collector.

[0072] In one embodiment, the conditions for coaxial electrospinning are as follows: the columnar collector is made of stainless steel with a diameter of 0.7~1.5mm, specifically 1mm in this embodiment; the rotation speed is 1000~3000rpm, specifically 1500rpm in this embodiment; the flow rate of the spinning solution in the outer tube is 20~30µL / min, specifically 25µL / min in this embodiment; the flow rate of the spinning solution in the inner tube is 5~15µL / min, specifically 5µL / min in this embodiment; the flow rate ratio of the spinning solution in the outer tube to the inner tube is 2~5:1, specifically 5:1 in this embodiment; the DC voltage is 15~25kV, specifically 18kV in this embodiment; the receiving distance is 10~20cm, specifically 15cm in this embodiment; the temperature is 25±2℃; and the relative humidity is 50±5%.

[0073] In one embodiment, after coaxial electrospinning, the process further includes: sequential drying, demolding, and sterilization, followed by sealing for later use; the drying equipment is a vacuum drying oven; the drying temperature is 25~35℃, specifically 30℃ in this embodiment, the vacuum degree is -0.08~-0.1MPa, specifically -0.09MPa in this embodiment, and the time is 24~48h, specifically 36h in this embodiment; demolding involves peeling the dried multi-layer tubular support from the collector; sterilization is performed using ethylene oxide gas.

[0074] In the multilayer tubular scaffold prepared by this invention, the drug is encapsulated in the fiber core layer, and the polymer shell provides physical protection and a barrier to control diffusion, thereby enabling more stable and longer-lasting drug release and further optimizing the anti-tumor effect.

[0075] This invention employs a "blended spinning" process (mixing drugs and polymers) to obtain a single-layer tubular scaffold. Coaxial electrospinning technology is used to prepare nanofibers with a core-shell structure, encapsulating the drug in the fiber core and using the polymer as the shell, resulting in a multi-layer tubular scaffold. The core-shell structure can further mitigate the initial burst release effect of the drug, achieving a longer-lasting sustained release, while protecting less bioactive drugs (such as protein drugs) from inactivation. The protein drugs include: basic fibroblast growth factor (bFGF), used to promote angiogenesis and cell proliferation at the repair site; vascular endothelial growth factor (VEGF), used to specifically promote local vascular network reconstruction; epidermal growth factor (EGF), used to directly accelerate the regeneration and coverage of the urothelial layer; transforming growth factor-β3 (TGF-β3), used to regulate extracellular matrix reconstruction, inhibit excessive fibrosis, and promote scarless healing; and urokinase, used to dissolve fibrin deposits and maintain scaffold lumen patency.

[0076] Addressing the three major challenges of existing technologies—namely, the inability to meet the need for tube removal, the difficulty in balancing degradation and mechanical properties, and the lack of precise local chemotherapy—this invention utilizes electrospinning technology to construct a tubular scaffold with a biomimetic nanofiber topology. By simulating a fiber entanglement network at the microscopic level, this invention endows the brittle, biodegradable polymer material with excellent damage tolerance and radial resilience, enabling it to resist the continuous peristaltic pressure of the ureter.

[0077] Based on this biomimetic structure, the present invention further introduces the strategies of "crystallization regulation" and "functionalized drug loading": 1. By precisely controlling the monomer ratio of PLGA (locked at around 70:30), the material is induced to transform from amorphous to semi-crystalline, and the degradation mode is changed from "erosion-dominated" to "swelling-dominated" by using the crystalline region as a physical barrier, thereby precisely matching the approximately 4-week repair window of the ureter.

[0078] 2. The antimetabolite drug gemcitabine was loaded in situ into nanofibers, which not only improved the surface wettability of the scaffold by utilizing its hydrophilic groups to reduce adhesion, but also achieved local long-term inhibition of residual tumor cells.

[0079] In summary, the present invention aims to develop a novel ureteral stent and its manufacturing process that integrates "biomimetic mechanical enhancement", "controllable degradation cycle" and "anti-tumor therapy", so as to completely solve the problems of traditional stents requiring secondary surgery and the poor mechanical properties of existing biodegradable stents.

[0080] The technical solutions of the present invention will be clearly and completely described below with reference to the embodiments of the present invention, but they should not be construed as limiting the scope of protection of the present invention.

[0081] The test drugs, specifications, and manufacturers are shown in Table 1.

[0082] Table 1 Test Drugs

[0083] The raw materials and manufacturers used in cell experiments are listed in Table 2.

[0084] Table 2 Materials and Reagents for Cell Experiments

[0085] Example 1: Ureteral stent with low LA content and low drug loading (focusing on short-term support and long-term sustained release): 1. Experimental method: Raw material ratio: PLGA (poly(lactic acid-CO-glycolic acid) purchased from Shenzhen Polystyrene Technology Co., Ltd.) with a lactic acid to glycolic acid molar ratio (LA:GA) of 60:40 was selected as the matrix; the amount of gemcitabine, an antitumor drug, added was 5% of the total mass of the spinning solute.

[0086] Spinning solution preparation: Dissolve PLGA in trifluoroacetic acid (TFA) to prepare a 20wt% solution, add gemcitabine at 5% of the mass of PLGA in proportion, and stir magnetically at 400r / min until completely dissolved.

[0087] Sacrificial layer preparation: 0.5 g of PVP K30 (number average molecular weight of 40,000) was dissolved in 3.17 mL of anhydrous ethanol and electrospun on a 2 mm diameter stainless steel mandrel rotating at 1500 r / min. The voltage was 23 kV, the injection speed was 24 µL / min, the receiving speed was 1500 rpm, and the spinning time was 70 min to form a PVP sacrificial layer with a thickness of 0.2 mm.

[0088] Scaffold formation: PLGA drug-loaded spinning was performed on the PVP sacrificial layer (voltage 23kV, injection rate 24µL / min, receiving speed 2000rpm, spinning time 125min) to form a tubular structure.

[0089] Post-treatment: Immerse in deionized water for 30 min to dissolve the PVP layer and demold. Vacuum dry for 24 h under vacuum conditions of -0.095 MPa and 25℃, then sterilize with ultraviolet light (70 μW / cm²). 2 (Irradiated with ultraviolet light for 24 hours).

[0090] 2. Test results (product structure and effect data): Structural identification (SEM, D in Figure 1): Scanning electron microscopy shows that the fiber surface is smooth, with no obvious adhesion, and the average fiber diameter is relatively fine, distributed in the range of 0.4~0.5µm.

[0091] Wettability (performance data): The contact angle test result was approximately 111.5°, exhibiting a certain degree of hydrophobicity. This indicates that low-concentration drugs have minimal impact on the surface energy of the material, making it suitable for short-term implantation scenarios where hydrophilicity requirements are not high.

[0092] Mechanical properties and degradation (performance data): Due to the low LA content (60:40), the material has low crystallinity and is mainly in an amorphous state. It degrades rapidly in artificial urine, with significant mass loss and mechanical degradation beginning in week 3, matching the short-term healing requirements.

[0093] Antitumor activity (efficacy data): It has basic cytotoxicity, and the survival rate of 5637 cells after co-culturing for 1 day is less than 50%, and the drug release cycle is relatively long.

[0094] Example 2: Ureteral stent with moderate LA content and moderate drug loading (emphasizing performance balance and standard treatment): 1. Experimental method: Raw material ratio: PLGA with an LA:GA molar ratio of 70:30 was selected as the matrix; gemcitabine was added at 15% of the mass of PLGA.

[0095] Preparation process: The solvent and concentration of the spinning solution are the same as in Example 1, the electrospinning receiving speed is adjusted to 2500 rpm, and the remaining steps are consistent with Example 1.

[0096] 2. Test results (product structure and effect data): Structural identification (SEM, D in Figure 1): The uniformity of fiber morphology was significantly improved, with an average diameter of about 0.5~0.6 µm. The fiber network structure was dense and the orientation was moderate, with no obvious beading or defects.

[0097] Crystal structure (Figure 2 XRD, Figure 3 DSC): The XRD pattern shows that the material has semi-crystalline characteristic peaks, and the DSC shows obvious endothermic melting peaks, indicating that this ratio induces the ordering of the microstructure.

[0098] Wettability (Effectiveness Data): The water contact angle is significantly reduced to approximately 35.0°, demonstrating excellent hydrophilicity. This facilitates the formation of a hydration layer on the surface, reduces protein adsorption and tissue adhesion, and alleviates patient discomfort.

[0099] Mechanical properties and degradation (performance data): It exhibits a biomimetic toughening effect of "fiber bridging" and excellent radial resilience. The degradation mode is dominated by swelling, maintaining structural integrity within 4 weeks, and the tensile strength remains above 5N, perfectly matching the standard repair window of 3-4 weeks for the ureter.

[0100] Antitumor activity (efficacy data): The drug release exhibited a benign combination of burst release and sustained release. After 5 days of co-culture, the survival rate of cancer cells (human urothelial carcinoma cell line (5637)) was significantly reduced (<30%), demonstrating a significant antitumor effect.

[0101] Example 3: Ureteral stent with high LA content and high drug loading (focusing on long-term support and high-intensity treatment): 1. Experimental method: Raw material ratio: PLGA with an LA:GA molar ratio of 80:20 was selected as the matrix; gemcitabine was added at 25% of the mass of PLGA.

[0102] Preparation process: The solvent and concentration of the spinning solution are the same as in Example 1, the electrospinning receiving speed is adjusted to 3000 rpm, and the remaining steps are consistent with Example 1.

[0103] 2. Test results (product structure and effect data): Structural identification (SEM, D in Figure 1): The fiber diameter increased, and fiber adhesion and drug aggregation occurred in some areas, but the overall tubular structure remained intact and the fiber network became denser.

[0104] Wettability (Effectiveness Data): The water contact angle drops to 0°, exhibiting extreme hydrophilicity (superhydrophilicity). This greatly promotes the penetration of body fluids, and although it accelerates the initial drug release, it is also balanced by the highly crystalline matrix.

[0105] Mechanical properties and degradation (effect data): Due to the high crystallinity resulting from the high LA content (80:20), the material has a relatively high initial stiffness. Although the high drug loading (25%) leads to a rapid decrease in mechanical properties (<5N) in the later stages of degradation (after 4 weeks), its structural stability is extremely high in the early stages of degradation, making it suitable for initial treatments requiring strong anti-tumor intervention.

[0106] Antitumor activity (efficacy data): The drug release concentration is high, and the killing efficiency against cancer cells (human urothelial carcinoma cell line (5637)) is extremely high, making it suitable for the early stage of treatment when the tumor burden is large.

[0107] Example 4 Preparation method of antitumor biodegradable ureteral stent with core-shell structure that can realize drug sustained release, the specific steps are as follows: (1) Solution preparation: a) Preparation of shell solution: Dissolve polylactic acid-glycolic acid copolymer (PLGA, LA:GA=75:25) in a mixed solvent of dichloromethane and N,N-dimethylformamide (volume ratio 8:2), stir magnetically until completely dissolved, and obtain a homogeneous transparent solution with a concentration of 18% (w / v).

[0108] b) Preparation of core layer solution: The antitumor drug gemcitabine and polyvinyl alcohol (PVA) were dissolved together in deionized water and magnetically stirred until completely dissolved to obtain a drug-loaded aqueous solution with a PVA concentration of 8% (w / v) and a gemcitabine concentration of 2% (w / v).

[0109] (2) Coaxial electrospinning: The shell solution and core solution were injected into two separate injection pumps and connected to coaxial spinning needles (shell solution connected to the outer channel, core solution connected to the inner channel). The needles were aligned with a high-speed rotating cylindrical stainless steel collector (1 mm in diameter, 1500 rpm). The process parameters were set as follows: shell solution flow rate of 25 µL / min, core solution flow rate of 5 µL / min (flow rate ratio 5:1), DC voltage of 18 kV, and receiving distance of 15 cm. Spinning was carried out at a temperature of 25 ± 2 °C and a relative humidity of 50 ± 5% until a tubular fiber membrane of a predetermined thickness was formed on the collector.

[0110] (3) Post-processing: The collector containing the fiber membrane was transferred to a vacuum drying oven and dried at 30°C and -0.09 MPa for 36 h. Subsequently, the dried tubular scaffold was carefully peeled off from the collector to obtain a core-shell fiber scaffold with PLGA as the shell and PVA / gemcitabine as the core. Finally, the scaffold was sterilized with ethylene oxide gas and sealed for later use.

[0111] The scaffold prepared in this embodiment has its drug encapsulated in a fiber core layer, with a polymer shell providing physical protection and a barrier to control diffusion, thereby enabling more stable and longer-lasting drug release and further optimizing the anti-tumor effect.

[0112] Comparative Example 1: Comparison of Existing Commercial Technologies (Non-degradable Polyurethane Stents) 1. Preparation / Acquisition Method: Commercially available polyurethane (PU) ureteral stents commonly used in clinical urology were selected as control products.

[0113] The scaffold is manufactured using a traditional melt extrusion process, non-electrostatic spinning, and contains no anti-tumor drugs.

[0114] Cut to the same dimensions as the example (inner diameter 2.0mm, length 20mm).

[0115] 2. Test Results and Performance Comparison: Microstructure (Figure 8): SEM observation shows that the surface of Comparative Example 1 is flat and dense, without nanofiber structure and lacking biomimetic topological features. Although this smooth surface has low friction, it lacks the ability of cells to "contact guide," which is not conducive to epithelialization crawling during tissue repair.

[0116] Degradation performance: After immersion in artificial urine for 4 weeks, the mass loss rate is <1%, indicating no degradation. This means that patients must undergo a second cystoscopic catheter removal surgery after treatment, increasing pain and medical costs, and failing to address the clinical pain point of "catheter-free" treatment.

[0117] Antitumor activity: After co-culturing with 5637 urothelial carcinoma cells for 5 days, the cell survival rate was close to 100%, with no effect on inhibiting tumor growth. This means that the stent can only play a physical drainage role and cannot provide local treatment for ureteral cancer. It not only cannot prevent tumor recurrence, but may even aggravate inflammation due to foreign body stimulation.

[0118] Conclusion: Compared with Comparative Example 1, the present invention (Examples 1-3) successfully achieved the dual functions of "controllable degradation" and "anti-tumor", solving the defects of traditional commercial stents that require a second surgery and cannot treat tumors.

[0119] Comparative Example 2: Comparison of material components outside the protection scope (LA:GA=50:50) 1. Preparation method: Except for replacing the matrix material with PLGA (LA:GA=50:50), the drug concentration and preparation process (rotation speed 2500rpm) are the same as in Example 2.

[0120] Note: The LA:GA ratio (50:50) in this comparative example is outside the scope of protection of this invention (60:40~85:15).

[0121] 2. Test Results and Performance Comparison: Crystallinity and Mechanical Properties: Due to the excessively high glycolic acid (GA) content, the material exhibits a completely amorphous state, lacking crystalline support. Radial compression tests show that its initial modulus is low and its compressive resilience is poor.

[0122] Degradation kinetics: The degradation rate in artificial urine is extremely rapid. Significant mass loss occurs in the first week, the scaffold structure collapses and breaks in the second week, and the mass loss exceeds 50% in the fourth week.

[0123] Failure analysis: Due to its rapid degradation, the stent cannot cover the 3-4 week recovery period usually required after ureteral surgery, which can easily cause the stent to break prematurely, block the ureter, and lead to hydronephrosis or infection risk.

[0124] Conclusion: This demonstrates the necessity of limiting the PLGA molar ratio to the range of 60:40 to 80:20 as per the present invention. Below this range (e.g., 50:50), degradation becomes uncontrolled, failing to meet the clinical support window.

[0125] Comparative Example 3: Comparison of preparation process parameters outside the protection range (low speed 1000 rpm) 1. Preparation method: Except for setting the receiving speed of electrospinning to 1000 rpm, the other material formulations are consistent with those in Example 2.

[0126] Note: The rotational speed (1000 rpm) in this comparative example is outside the scope of protection of this invention (1500~3500 rpm).

[0127] 2. Test results and performance comparison: Macroscopic forming: Due to the low rotation speed and insufficient centrifugal force, the fiber deposition layer is loose, the tube wall thickness is uneven, the tube diameter increases very slowly and is difficult to control precisely at the standard size of 2.0mm.

[0128] Microstructure (Figure 9): SEM shows that the fibers are arranged randomly and the porosity is too high.

[0129] Degradation performance: The loose structure allows moisture to easily penetrate into the scaffold, inducing an autocatalytic hydrolysis effect. Tests showed that its mass loss rate at week 3 was significantly higher than that of Example 2 (dense structure), and its mechanical strength decayed too quickly, making it unable to maintain a stable supporting function.

[0130] Conclusion: This demonstrates the necessity of limiting the electrospinning receiving speed to the range of 1500~3500 rpm in this invention. Too low a speed cannot obtain a dense orientation structure, thus affecting dimensional accuracy and degradation stability.

[0131] Comparative Example 4: Comparison of drug concentrations outside the protection range (drug loading 30%) 1. Preparation method: Except for increasing the drug loading of gemcitabine to 30%, the other conditions were the same as in Example 2.

[0132] Note: The drug loading (30%) in this comparative example is outside the scope of protection of this invention (5%~25%).

[0133] 2. Test Results and Performance Comparison: Molding Quality: High drug concentrations caused changes in the viscosity of the spinning solution, resulting in unstable jetting during electrospinning. SEM showed a large number of drug crystals agglomerated on the fiber surface, causing severe adhesion and fusion of the fibers, resulting in the loss of the independent structure of the nanofibers.

[0134] Mechanical properties: Due to drug aggregation disrupting the continuity of the polymer chains, the tensile strength of this scaffold is significantly reduced even before degradation. Four weeks after degradation, its tensile strength is far below the critical value of 5N, posing a risk of breakage after implantation.

[0135] Surface properties: Excessive hydrophilic drugs cause the material to absorb water extremely well, making it prone to premature swelling and deformation within the body.

[0136] Conclusion: This demonstrates the necessity of limiting the drug loading to the range of 5-25% in this invention. Exceeding this range leads to structural defects and loss of mechanical properties in the material.

[0137] Performance test (1) Figure 1 shows the physical picture and SEM image of the single-layer tubular stent prepared by the present invention. A is the actual object of the tubular ureter prepared by electrospinning, where (a) is the overall macroscopic morphology of the ureter stent, (b) is the inner diameter measurement of the ureter stent, (c) is the port cross-section of the ureter stent, B and C are the actual objects based on the biomimetic principle, and D is the morphological change of the fiber under different drug loading (0%, 5%, 15%, 20%, 25%), where (a) is PLGA-Gem 0%, (b) is PLGA-Gem 5%, (c) is PLGA-Gem 10%, (d) is PLGA-Gem 15%, (e) is PLGA-Gem 20%, and (f) is PLGA-Gem 25%.

[0138] As can be seen from Figure 1, the present invention can form fibers in a wide range.

[0139] (2) Figure 2 shows the XRD diffraction patterns of scaffolds with different LA:GA ratios (50:50, 60:40, 70:30, 80:20), and Figure 3 shows the DSC thermal analysis curves of scaffolds with different LA:GA ratios (50:50, 60:40, 70:30, 80:20).

[0140] As can be seen from Figures 2 and 3, the present invention can achieve structural control from amorphous to semi-crystalline by adjusting the ratio.

[0141] (3) Figure 4 shows the water contact angle test diagram of stents with different drug loading concentrations, where (a) is PLGA-Gem 0%, (b) is PLGA-Gem 5%, (c) is PLGA-Gem 10%, (d) is PLGA-Gem 15%, (e) is PLGA-Gem 20%, and (f) is PLGA-Gem 25%.

[0142] Figure 4 shows the change in contact angle from 111.5° (hydrophobic) to 0° (superhydrophilic), demonstrating the tunability of surface properties.

[0143] (4) Figure 5 is a curve of mass loss of the stent during the degradation process in artificial urine, and Figure 6 is a curve of mechanical property change of the stent during the degradation process in artificial urine.

[0144] Figures 5 and 6 demonstrate that different formulations correspond to different degradation kinetics.

[0145] (5) Figure 7 is a bar chart of cell survival rate (CCK-8) of scaffold co-cultured with 5637 urothelial carcinoma cells.

[0146] As can be seen from Figure 7, all embodiments demonstrate significant antitumor activity.

[0147] Although the above embodiments have provided a detailed description of the present invention, they are only some embodiments of the present invention and not all embodiments. People can obtain other embodiments based on these embodiments without creative effort, and these embodiments all fall within the protection scope of the present invention.

Claims

1. A biomimetic biodegradable ureteral stent with a biomimetic structure, characterized in that, The stent may include a single-layer tubular stent or a multi-layer tubular stent; the single-layer tubular stent is made of an aliphatic polyester polymer and a drug; the multi-layer tubular stent includes a coaxial inner tube and an outer tube; the inner tube is made of polyvinyl alcohol and a drug; the outer tube is made of an aliphatic polyester polymer; the drug includes antitumor drugs and / or protein drugs; the aliphatic polyester polymer includes at least one of polylactic acid-glycolic acid copolymer, polycaprolactone, polylactic acid, and polyethylene glycol-polylactic acid-glycolic acid copolymer; the molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer is 60~85:15~40; the mass of the drug in the single-layer tubular stent is 5~25% of the mass of the aliphatic polyester polymer; the mass of the drug in the multi-layer tubular stent is 5~25% of the mass of the aliphatic polyester polymer.

2. The biomimetic biodegradable ureteral stent with a biomimetic structure according to claim 1, characterized in that, The antitumor drugs include at least one of gemcitabine, paclitaxel, cisplatin, pirarubicin, and doxorubicin.

3. The biomimetic biodegradable ureteral stent with a biomimetic structure according to claim 1, characterized in that, The number average molecular weight of the polycaprolactone is not higher than 50,000.

4. The biomimetic biodegradable ureteral stent with a biomimetic structure according to claim 1, characterized in that, The polylactic acid is a copolymer of L-lactic acid and D-lactic acid, wherein the molar percentage of D-lactic acid is 15-25%; the intrinsic viscosity of the polylactic acid is 0.8-1.6 dL / g.

5. The biomimetic biodegradable ureteral stent with a biomimetic structure according to claim 1, characterized in that, The polyethylene glycol-polylactic acid-glycolic acid copolymer has a molecular weight of 1,000 to 5,000 and a mass percentage of 5 to 15% in the polyethylene glycol-polylactic acid-glycolic acid copolymer. The molar ratio of lactic acid to glycolic acid in the polylactic acid-glycolic acid copolymer block is 85 to 90: 10 to 15.

6. The method for preparing the biomimetic biodegradable ureteral stent with an antitumor structure according to any one of claims 1 to 5, characterized in that, When the biomimetic antitumor biodegradable ureteral stent is a single-layer tubular stent, the preparation method includes the following steps: firstly spinning a sacrificial layer spinning solution on a mandrel to form a sacrificial layer film; secondly spinning a drug-loaded spinning solution on the sacrificial layer film to form a composite film; the drug-loaded spinning solution includes an aliphatic polyester polymer, a drug, and a first organic solvent; removing the sacrificial layer film from the composite film to obtain a single-layer tubular stent; when the biomimetic antitumor biodegradable ureteral stent is a multilayer tubular stent, the preparation method includes the following steps: coaxially electrospinning an inner tube spinning solution and an outer tube spinning solution together to obtain a multilayer tubular stent; the inner tube spinning solution includes polyvinyl alcohol, a drug, and a solvent; the outer tube spinning solution includes an aliphatic polyester polymer and a third organic solvent.

7. The preparation method according to claim 6, characterized in that, The sacrificial layer spinning solution includes a sacrificial layer material and a second organic solvent; the sacrificial layer material is one or more of polyvinylpyrrolidone, polyethylene oxide, water-soluble gelatin, and polyvinyl alcohol.

8. The preparation method according to claim 6, characterized in that, The conditions for the first spinning process include: the pump head inner diameter of the injection pump is 0.5~1.0mm, the voltage is 20~25kV, the injection speed is 20~30µL / min, the receiving speed is 1000~2000rpm, and the spinning time is 30~90min.

9. The preparation method according to claim 6, characterized in that, The conditions for the second spinning process include: a voltage of 20-25 kV, a feed rate of 20-30 µL / min, a receiving speed of 2000-3000 rpm, and a spinning time of 2-3 h.

10. The preparation method according to claim 6, characterized in that, The conditions for coaxial electrospinning are as follows: the columnar collector is made of stainless steel with a diameter of 0.7~1.5mm and a rotation speed of 2000~3000rpm; the flow rate of the spinning solution in the outer tube is 20~30µL / min, and the flow rate of the spinning solution in the inner tube is 5~15µL / min; the flow rate ratio of the spinning solution in the outer tube to the inner tube is 2~5:1; the DC voltage is 20~25kV; the receiving distance is 10~20cm; the temperature is 25±2℃; and the relative humidity is 50±5%.

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