A micro-channel drug-loaded bioabsorbable polylactic acid stent and a preparation method thereof

Abstract

This invention discloses a microchannel-based bioresorbable polylactic acid (PLA) scaffold and its preparation method. The scaffold comprises a tubular mesh body made of PLA-based bioresorbable material, with at least a portion of the scaffold struts forming a continuous hollow microchannel network, which serves as a structural reservoir for therapeutic drugs. The scaffold preparation method is as follows: a three-dimensional tubular template is constructed using sacrificial template fibers; PLA solution is impregnated / drained into the three-dimensional tubular template and solidified into a composite preform; the composite preform undergoes reinforcement, shaping, and curing treatment; subsequently, the template is dissolved and removed to form the microchannel network; then, a drug solution is perfused into the microchannel network using pressure-driven infusion, followed by drying and curing to complete the loading. Compared to surface coating drug loading methods, this invention can achieve protective drug loading and controlled release while maintaining mechanical support, making it suitable for interventional treatments in cardiovascular, peripheral vascular, and non-vascular lumens.

Description

Technical Field

[0001] This invention relates to the fields of biomedical engineering and medical device technology, and in particular to a microchannel-based bioabsorbable polylactic acid scaffold and its preparation method. Background Technology

[0002] Vascular stent implantation is the mainstream interventional procedure for treating stenosis or occlusion of coronary arteries and peripheral blood vessels. An ideal stent should effectively support the blood vessel, prevent acute recoil and restenosis, and gradually degrade and be absorbed after fulfilling its purpose, thereby reducing long-term foreign body-related risks.

[0003] Currently, most mainstream drug-eluting stents rely on surface coatings for drug loading and release. This approach has inherent drawbacks: the drug is only loaded into a very thin coating on the stent surface. This surface coupling mode results in the drug release kinetics being unrelated to the stent skeleton degradation kinetics. The surface coating is in direct contact with the blood, making it prone to burst release or premature complete release due to coating dissolution or peeling. Furthermore, the polymer skeleton, which serves as the mechanical support, degrades deep within the stent, making it difficult to utilize the degradation behavior of the skeleton material (such as bulk dissolution and pore formation) to synergistically drive or regulate the surface drug release rate. Therefore, existing technologies struggle to achieve precise matching between the drug release cycle and the stent degradation cycle, failing to meet the clinical need for long-term, stable, and controllable release. Moreover, long-term residual coatings or exposed stent skeletons can easily trigger inflammatory responses.

[0004] Therefore, there is an urgent need for a novel bioresorbable drug scaffold that can structurally integrate a drug reservoir into the scaffold framework, achieve synergistic regulation of drug release and framework degradation behavior, and simultaneously take into account both mechanical properties and controlled drug release performance. Summary of the Invention

[0005] To address the aforementioned problems, this invention provides a bioresorbable polylactic acid scaffold based on microchannel drug delivery, thereby resolving the issues raised in the background. This scaffold forms a continuous hollow microchannel network within at least a portion of the scaffold struts, serving as a structural reservoir for the drug, ensuring that the drug is primarily located within the scaffold struts rather than on the surface coating.

[0006] To achieve the above objectives, the present invention provides the following technical solution:

[0007] In a first aspect, the present invention provides a microchannel-based bioabsorbable polylactic acid scaffold, comprising: a tubular mesh scaffold body formed of polylactic acid-based bioabsorbable material, the scaffold body being composed of a plurality of interconnected scaffold rods; wherein at least a portion of the scaffold rods form a continuous hollow microchannel network inside; the microchannel network has an inner diameter of 20-30 micrometers, and the microchannel network is used to contain therapeutic drugs or their formulations.

[0008] Preferably, the inner diameter of the microchannel network is 25 micrometers.

[0009] Preferably, the therapeutic agent is an antiproliferative drug and / or an anti-inflammatory drug, more preferably rapamycin, everolimus, paclitaxel or a pharmaceutically acceptable salt thereof.

[0010] Preferably, the polylactic acid-based bioabsorbable material is selected from poly-L-lactic acid (PLLA), poly-racemic lactic acid (PDLLA), and their copolymers.

[0011] Secondly, the present invention provides a method for preparing the above-mentioned scaffold: a three-dimensional tubular template is constructed using water-soluble sacrificial template fibers; a polylactic acid solution is infiltrated and solidified using impregnation or dynamic impregnation-dip slurrying to form a composite preform; the composite preform is then subjected to strengthening, shaping, and curing treatment; subsequently, the composite preform is immersed in a solvent to dissolve and remove the template fibers to form a microchannel network; and finally, a drug solution is infused into the microchannel network using pressure-driven perfusion and dried and cured to complete the loading. Wherein:

[0012] The water-soluble sacrificial template fiber is prepared from polyvinyl alcohol, alginate, sucrose fiber or gelatin;

[0013] The polylactic acid solution has a mass fraction of 8-12 wt%, and the solvent is selected from dichloromethane, chloroform, or a mixture thereof;

[0014] The temperature of the strengthening and curing treatment is higher than the glass transition temperature of polylactic acid material but lower than its melting temperature.

[0015] The template fiber removal solvent is deionized water, and the template fiber is completely dissolved and removed by changing the water multiple times; the active pressure driven infusion adopts pulsed pressurization, and the pressure cycle is 0.2 MPa×30 s intermittent and 0.5 MPa×10 s alternating cycle.

[0016] Preferably, the method for preparing the above-mentioned stent includes the following steps:

[0017] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0018] 1. Preparation of polyvinyl alcohol (PVA) solution: Dissolve PVA powder in deionized water at 85℃ to prepare a solution with a concentration of 6-12wt%, and stir magnetically until completely dissolved and clear.

[0019] 2. Electrospinning:

[0020] 1) Fill the syringe with PVA solution and connect the metal needle (specification: 20-25G).

[0021] 2) Set the spinning parameters: feed rate 0.5-2 mL / h, high voltage power supply voltage 15-25 kV, receiving distance 10-20cm.

[0022] 3) Cover the receiving drum with aluminum foil and obtain a PVA fiber membrane by electrospinning. The fiber membrane is composed of PVA fibers ranging from nanometer to micrometer scale.

[0023] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 30-60 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0024] 5) Hot Stretching and Sizing: The above-mentioned water-containing, oriented fiber bundles are guided through a hot air drying channel at a temperature of 80-100℃. During this process, the PVA fiber surface softens and fuses, while a 3:1 draw ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to be bundled and fused into a dense, continuous PVA template fiber with a diameter of 20-30 μm under the combined action of heat and force. Finally, the fiber is wound and collected for later use.

[0025] In this step, the diameter of a single fiber can be controlled by adjusting parameters such as solution concentration, voltage, and receiving distance, thus providing raw materials for subsequent processing.

[0026] Step 2: Prepare a three-dimensional tubular template and perform dynamic impregnation-dip slurry application.

[0027] 1. Dissolve poly-L-lactic acid (PLLA) particles in dichloromethane to prepare a solution with a concentration of 8-12 wt%. Place the solution in a sealed container and stir with a magnetic stirrer at 300-500 rpm for 4-6 hours at room temperature until the PLLA is completely dissolved, forming a homogeneous, transparent, viscous solution without obvious bubbles;

[0028] 2. Weaving and Fixing the PVA Fiber Tube Template: The continuous PVA template fibers, approximately 25 μm in diameter, prepared in the first step, are dried in a vacuum drying oven (40-50℃) for 2-4 hours to completely remove physically adsorbed water, ensuring fiber crystallinity and improving its rigidity in organic solvents. Then, a double-helix cross-winding method is used. The PVA template fibers are continuously wound around the mandrel at a constant helix angle (+45°) to form the first set of helices. Subsequently, the direction is switched without cutting the fibers, or the same / other fiber is used to perform a second set of helical winding at an opposite helix angle (-45°). The two sets of fibers intersect, and at each intersection, a small amount of solvent (such as deionized water mist) is used for local bonding and curing, constructing a three-dimensional tubular mesh structure on a specially made micro mandrel (1-2 mm in diameter), forming a stable tubular PVA fiber template that fits well with the mandrel.

[0029] 3. Dynamic Impregnation: Slowly immerse the device, which holds the mandrel and PVA template, into a beaker containing PLLA solution, ensuring the template is completely submerged. Turn on the motor and rotate the mandrel continuously in the solution at a low speed of 10-20 rpm for 3-5 minutes.

[0030] 4. Draining and initial molding: Pull the mandrel assembly out of the PLLA solution at a controlled, uniform speed (approximately 2-5 mm / s). Immediately after pulling it out, start the motor and rotate the mandrel in the air at a medium speed of 50-100 rpm for 30-60 seconds;

[0031] 5. Solvent Evaporation Pre-curing: Place the completed ploughed mandrel with the composite preform horizontally on a support in a fume hood. Let it stand at room temperature for 15-30 minutes to allow most of the dichloromethane solvent to evaporate naturally. At this time, the PLLA coating gradually gels from a liquid state, possessing some initial strength, but it is not yet fully cured.

[0032] In this step, dynamic rotation ensures that the solution fully penetrates the porous, complex three-dimensional template, avoiding dead zones.

[0033] Step 3: Strengthening, Shaping, and Curing

[0034] 1. Place a small, open beaker containing a small amount of dichloromethane (DCM) solvent (about 5-10 mL) at the bottom of a large glass desiccator;

[0035] 2. The composite preform (together with the mandrel) that has completed the "dynamic impregnation-dip slurry" process and undergone preliminary curing at room temperature is suspended in a large glass desiccator to ensure that it does not come into direct contact with the liquid solvent.

[0036] 3. Place in a sealed desiccator and let stand at room temperature (25℃) for 30-60 minutes;

[0037] 4. Thermally induced crystallization: Carefully transfer the composite preform treated with solvent vapor to a preheated forced-air drying oven and treat it at 80-90℃ (this temperature is higher than the glass transition temperature of polylactic acid but lower than its melting temperature) for 20-40 minutes.

[0038] 5. After the procedure is complete, allow it to cool naturally to room temperature, and carefully remove the composite embryo from the mandrel.

[0039] In this step, PLLA is transformed from an amorphous or low-crystallinity gel into a highly crystalline solid, significantly increasing the modulus and strength of the scaffold and improving its dimensional stability, thus providing sufficient structural integrity for subsequent processing.

[0040] Step 4: Dissolve and remove the template to form microchannels

[0041] 1. Immerse the cured composite scaffold preform into a beaker containing a large amount of deionized water;

[0042] 2. Allow to stand or gently agitate at room temperature or with slight heating (37°C, simulating physiological temperature) for 4-12 hours to ensure complete dissolution of the PVA fibers. The water can be changed several times during this period to accelerate the dissolution and elution process.

[0043] 3. Remove the support structure with tweezers, rinse gently with deionized water, and then allow it to dry slightly in air or under low vacuum. At this point, a hollow microchannel network has formed inside the support structure.

[0044] In this step, the network channels and the scaffold body are integrally molded, avoiding the traditional coating interface and achieving physical separation between the drug reservoir and the mechanical support structure. The gentle water dissolution process does not damage the already crystallized and solidified PLLA structure.

[0045] Step 5: Active pressure-driven infusion and curing

[0046] 1. Preparation of drug solution: Dissolve the model drug in an appropriate amount of deionized water (or the corresponding organic solvent of the drug, such as acetone for rapamycin) to prepare a saturated or high-concentration solution;

[0047] 2. Set up a pressure infusion system: Completely immerse the dry support in the drug solution, then place them together in a pressure tank, seal the pressure tank, and connect it to a compressed nitrogen cylinder through a pipeline. Install a precision pressure regulating valve and a pressure gauge in the pipeline.

[0048] 3. Active pressure injection: Slowly fill the pressure tank with inert gas (nitrogen), and set the system pressure cycle to: 0.2MPa×30s intermittent + 0.5MPa×10s cycle, with a total duration of 30-60 minutes;

[0049] 4. Depressurization and Removal: Slowly release the system pressure to atmospheric pressure, open the pressure vessel, and use tweezers to remove the stent from the liquid.

[0050] 5. Drying and curing: Remove the stent from the drug solution, absorb excess solution on the surface with filter paper, and place it in a vacuum drying oven to dry at room temperature or lower temperature for 12-24 hours to completely remove the solvent, allowing the drug to be deposited in the microchannel in crystal or solid form.

[0051] In this step, pulsed pressure is used to drive the drug solution to overcome capillary resistance, allowing it to penetrate and fill deep into the entire microchannel network. The subsequent drying process, through solvent evaporation, causes the drug to become supersaturated and precipitate, depositing as crystals or a solid on the inner wall of the channel.

[0052] Thirdly, the present invention provides the use of the above-mentioned stent as or in the preparation of materials for cardiovascular, peripheral vascular and / or non-vascular lumen interventional therapy.

[0053] Compared with the prior art, the present invention has at least the following beneficial effects:

[0054] 1) Microchannel networks provide internal drug reservoirs, which helps reduce the risks of burst release and peeling associated with surface coatings;

[0055] 2) The drug load is relatively independent of the main stent structure, reducing the negative impact on mechanical support;

[0056] 3) The drug loading and release kinetics can be controlled by adjusting the channel size, network structure and perfusion process parameters to improve the controllable release capability. Attached Figure Description

[0057] Figure 1 This is a mechanical diagram for testing the bending stiffness of the support.

[0058] Figure 2 A bar chart to test the bending stiffness of the support.

[0059] Figure 3 This is the test data for the radial performance of the test bracket.

[0060] Figure 4 The graph shows the cumulative release rate of rapamycin in the stents obtained in Examples 1-3 and Comparative Examples 1-5. Detailed Implementation

[0061] The present invention will be further described in detail below with reference to embodiments. It should be understood that the specific embodiments described herein are merely illustrative of the invention and are not intended to limit the invention.

[0062] In this embodiment of the invention, some of the raw materials are as follows:

[0063] Polyvinyl alcohol (PVA), product number: S30196-500g, purchased from Shanghai Yuanye Biotechnology Co., Ltd.

[0064] Poly-L-lactic acid (PLLA), catalog number: 33135-50-1, purchased from Wuhan Haishan Biotechnology Co., Ltd.;

[0065] Rapamycin, product number: 53123-88-9, was purchased from Wuhan Dongkangyuan Biotechnology Co., Ltd.

[0066] Example 1

[0067] This embodiment provides a bioabsorbable polylactic acid scaffold based on microchannel drug delivery. The scaffold struts form a microchannel network internally, with an inner diameter of approximately 25 μm and a strut thickness of approximately 75 μm. Its preparation method includes:

[0068] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0069] 1. Preparation of polyvinyl alcohol (PVA) solution: Dissolve PVA powder in hot deionized water at 85℃ to prepare a 9wt% solution. Stir magnetically until completely dissolved and clear.

[0070] 2. Electrospinning:

[0071] 1) Fill the syringe with PVA solution and connect the metal needle (specification: 25G).

[0072] 2) Set the spinning parameters: feed rate 1 mL / h, high voltage power supply voltage 20 kV, receiving distance 15 cm.

[0073] 3) Cover the receiving drum with aluminum foil and obtain a PVA fiber membrane by electrospinning. The fiber membrane is composed of PVA fibers ranging from nanometer to micrometer scale.

[0074] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 45 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0075] 5) Hot Stretching and Sizing: The aforementioned water-containing, oriented fiber bundles are guided through a hot air drying channel at 90°C. During this process, the PVA fiber surface softens and fuses, while a 3:1 stretch ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to bundle and fuse under the combined action of heat and force into a dense, continuous PVA template fiber with a diameter of 25 μm, which is then wound and collected for later use.

[0076] Step 2: Prepare a three-dimensional tubular template and perform dynamic impregnation-dip slurry application.

[0077] 1. Dissolve poly-L-lactic acid (PLLA) particles in dichloromethane to prepare a 10 wt% solution. Place the solution in a sealed container and stir at 400 rpm for 5 hours at room temperature using a magnetic stirrer until the PLLA is completely dissolved, forming a homogeneous, transparent, viscous solution without obvious bubbles.

[0078] 2. Weaving and Fixing the PVA Fiber Tube Template: The continuous PVA template fibers, approximately 25 μm in diameter, prepared in the first step, are dried in a vacuum drying oven (50°C) for 2 hours to completely remove physically adsorbed water, ensuring fiber crystallinity and improving its rigidity in organic solvents. Then, a double-helix cross-winding method is used. The PVA template fibers are continuously wound around a mandrel at a constant helix angle (+45°) to form the first set of helices. Subsequently, without cutting the fibers, the guide is switched, or the same / other fiber is used to perform a second set of helical winding at an opposite helix angle (-45°). The two sets of fibers intersect, and at each intersection, a small amount of solvent (such as deionized water mist) is used for local bonding and curing, constructing a three-dimensional tubular mesh structure on a specially made micro-mandrel (1-2 mm in diameter), forming a stable tubular PVA fiber template that fits well with the mandrel.

[0079] 3. Dynamic Immersion: Slowly immerse the device, which holds the mandrel and PVA template, into a beaker containing PLLA solution, ensuring the template is completely submerged. Turn on the motor to rotate the mandrel continuously in the solution at a low speed of 15 rpm for 4 minutes.

[0080] 4. Draining and initial molding: Pull the mandrel assembly out of the PLLA solution at a controlled, uniform speed (approximately 3 mm / s). Immediately after pulling it out, start the motor and rotate the mandrel in the air at a medium speed of 75 rpm for 45 seconds;

[0081] 5. Solvent Evaporation Pre-curing: Place the completed ploughed mandrel with the composite preform horizontally on a support in a fume hood. Let it stand at room temperature for 20 minutes to allow most of the dichloromethane solvent to evaporate naturally. At this point, the PLLA coating gradually gels from a liquid state, possessing some initial strength, but is not yet fully cured.

[0082] Step 3: Strengthening, Shaping, and Curing

[0083] 1. Place an open beaker containing a small amount of dichloromethane (DCM) solvent (about 10 mL) at the bottom of a large glass desiccator;

[0084] 2. The composite preform (together with the mandrel) that has completed the "dynamic impregnation-dip slurry" process and undergone preliminary curing at room temperature is suspended in a large glass desiccator to ensure that it does not come into direct contact with the liquid solvent.

[0085] 3. Place in a sealed desiccator and let stand at room temperature (25℃) for 45 minutes;

[0086] 4. Thermally induced crystallization: Carefully transfer the composite preform treated with solvent vapor to a preheated forced-air drying oven and treat it at 85°C for 30 minutes;

[0087] 5. After the procedure is complete, allow it to cool naturally to room temperature, and carefully remove the composite embryo from the mandrel.

[0088] Step 4: Dissolve and remove the template to form microchannels

[0089] 1. Immerse the cured composite scaffold preform into a beaker containing a large amount of deionized water;

[0090] 2. Allow to stand or gently agitate for 8 hours at room temperature or with slight heating (37°C, simulating physiological temperature) to ensure complete dissolution of the PVA fibers. The water can be changed several times during this period to accelerate the dissolution and elution process.

[0091] 3. Remove the support structure with tweezers, rinse gently with deionized water, and then allow it to dry slightly in air or under low vacuum. At this point, a hollow microchannel network has formed inside the support structure.

[0092] Step 5: Active pressure-driven infusion and curing

[0093] 1. Preparation of drug solution: Dissolve rapamycin in acetone to prepare a nearly saturated solution of 30 mg / mL;

[0094] 2. Set up a pressure infusion system: Completely immerse the dry support in the drug solution, then place them together in a pressure tank, seal the pressure tank, and connect it to a compressed nitrogen cylinder through a pipeline. Install a precision pressure regulating valve and a pressure gauge in the pipeline.

[0095] 3. Active pressure injection: Slowly fill the pressure tank with inert gas (nitrogen), and set the system pressure cycle to: 0.2MPa×30s intermittent + 0.5MPa×10s cycle, for 45 minutes;

[0096] 4. Depressurization and Removal: Slowly release the system pressure to atmospheric pressure, open the pressure vessel, and use tweezers to remove the stent from the liquid.

[0097] 5. Drying and curing: Remove the stent from the drug solution, absorb excess solution on the surface with filter paper, and place it in a vacuum drying oven to dry at room temperature or lower temperature for 20 hours to completely remove the solvent, allowing the drug to be deposited in the microchannel in crystal or solid form.

[0098] Example 2

[0099] This embodiment provides a bioabsorbable polylactic acid scaffold based on microchannel drug delivery. The scaffold struts form a microchannel network with an inner diameter of approximately 20 μm and a strut thickness of approximately 75 μm. Its preparation method includes:

[0100] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0101] 1. Preparation of polyvinyl alcohol (PVA) solution: Dissolve PVA powder in hot deionized water at 85℃ to prepare a 6wt% solution. Stir magnetically until completely dissolved and clear.

[0102] 2. Electrospinning:

[0103] 1) Fill the syringe with PVA solution and connect the metal needle (specification: 25G).

[0104] 2) Set the spinning parameters: feed rate 0.5 mL / h, high voltage power supply voltage 25 kV, receiving distance 12 cm.

[0105] 3) Cover the receiving drum with aluminum foil and obtain a PVA fiber membrane by electrospinning. The fiber membrane is composed of PVA fibers ranging from nanometer to micrometer scale.

[0106] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 45 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0107] 5) Hot Stretching and Sizing: The aforementioned water-containing, oriented fiber bundles are guided through a hot air drying channel at 90°C. During this process, the PVA fiber surface softens and fuses, while a 3:1 stretch ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to bundle and fuse under the combined action of heat and force into a dense, continuous PVA template fiber with a diameter of 20 μm, which is then wound and collected for later use.

[0108] The remaining steps are the same as in Example 1.

[0109] Example 3

[0110] This embodiment provides a bioresorbable polylactic acid scaffold based on microchannel drug delivery. The scaffold struts form a microchannel network with an inner diameter of approximately 30 μm and a strut thickness of approximately 75 μm. Its preparation method includes:

[0111] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0112] 1. Preparation of polyvinyl alcohol (PVA) solution: Dissolve PVA powder in hot deionized water at 85℃ to prepare a 12wt% solution. Stir magnetically until completely dissolved and clear.

[0113] 2. Electrospinning:

[0114] 1) Fill the syringe with PVA solution and connect the metal needle (specification: 25G).

[0115] 2) Set the spinning parameters: feed rate 2 mL / h, high voltage power supply voltage 18 kV, receiving distance 20 cm.

[0116] 3) Cover the receiving drum with aluminum foil and obtain a PVA fiber membrane by electrospinning. The fiber membrane is composed of PVA fibers ranging from nanometer to micrometer scale.

[0117] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 45 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0118] 5) Hot Stretching and Sizing: The aforementioned water-containing, oriented fiber bundles are guided through a hot air drying channel at 90°C. During this process, the PVA fiber surface softens and fuses, while a 3:1 stretch ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to bundle and fuse under the combined action of heat and force into a dense, continuous PVA template fiber with a diameter of 30 μm, which is then wound and collected for later use.

[0119] The remaining steps are the same as in Example 1.

[0120] Comparative Example 1

[0121] Compared to Example 1, the electrospinning template fiber is eliminated, and a solid PLLA mesh is laser-cut, which has the same macroscopic structure but no internal microchannels and a surface coating of drug.

[0122] The preparation method is as follows:

[0123] Step 1: Laser cutting of PLLA solid mesh

[0124] PLLA membrane preparation: PLLA particles were dissolved in dichloromethane (10 wt%), the solution was transferred to a centrifuge tube, centrifuged at 3000 rpm for 10 minutes to remove micro-bubbles, the degassed PLLA solution was poured onto a horizontal glass plate, and coated unidirectionally with a spatula at a speed of 5 cm / s, covered with a dust cover, and allowed to evaporate at room temperature for 2 hours, and then transferred to a vacuum drying oven (25℃, -0.1 MPa) for 12 hours.

[0125] Laser cutting of the grid: The PLLA film is fixed on the cutting platform for cutting. The power is 15W, the speed is 20 mm / s, the frequency is 5 kHz, and the blowing angle is 45°.

[0126] Step 2: Surface loading of drugs

[0127] Drug coating solution: Dissolve the model drug in ethanol / water (7:3 v / v) to prepare a 5 mg / mL solution, and add 0.1% HPMC (thickening agent).

[0128] Spraying process: The spray gun nozzle diameter is 0.3 mm, the air pressure is 0.2 MPa, the distance is 15 cm, and the spraying is repeated 3 times (with a 10-minute drying interval). The drug load is controlled by weighing (target: same as in Example 1).

[0129] Comparative Example 2

[0130] Compared to Example 1, in step five, active pressure perfusion, the pulsed pressurization will be replaced with constant pressurization, with a system pressure of 0.2 MPa and a duration of 45 minutes. The rest is the same as in Example 1.

[0131] Comparative Example 3

[0132] Compared to Example 1, in step five, during active pressure infusion, the pulsed pressurization will be replaced with constant pressurization, with a system pressure of 0.5 MPa and a duration of 45 minutes. The rest is the same as in Example 1.

[0133] Comparative Example 4

[0134] This embodiment provides a bioabsorbable polylactic acid scaffold based on microchannel drug delivery. The scaffold struts form a microchannel network with an inner diameter of approximately 10 μm. Its preparation method includes:

[0135] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0136] 1. Prepare PVA solution: Dissolve PVA powder in hot deionized water to prepare a 4 wt% solution, and stir magnetically until completely dissolved and clear.

[0137] 2. Electrospinning:

[0138] 1) Fill the syringe with PVA solution and connect the metal needle.

[0139] 2) Set the spinning parameters: feed rate 0.3 mL / h, high voltage power supply voltage 28 kV, receiving distance 10 cm.

[0140] 3) Cover the receiving drum with aluminum foil and obtain PVA fibers by electrospinning.

[0141] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 45 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0142] 5) Hot Stretching and Sizing: The aforementioned water-containing, oriented fiber bundles are guided through a hot air drying channel at 90°C. During this process, the PVA fiber surface softens and fuses, while a 3:1 stretch ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to bundle and fuse under the combined action of heat and force into a dense, continuous PVA template fiber with a diameter of 10 μm, which is then wound and collected for later use.

[0143] The remaining steps are the same as in Example 1.

[0144] Comparative Example 5

[0145] This embodiment provides a bioabsorbable polylactic acid scaffold based on microchannel drug delivery. The scaffold struts form a microchannel network with an inner diameter of approximately 40 μm. Its preparation method includes:

[0146] Step 1: Preparation of PVA sacrificial template fibers (electrospinning method)

[0147] 1. Prepare PVA solution: Dissolve PVA powder in hot deionized water to prepare a 15wt% solution, and stir magnetically until completely dissolved and clear.

[0148] 2. Electrospinning:

[0149] 1) Fill the syringe with PVA solution and connect the metal needle.

[0150] 2) Set the spinning parameters: feed rate 2.5 mL / h, high voltage power supply voltage 16 kV, receiving distance 22 cm.

[0151] 3) Cover the receiving drum with aluminum foil and obtain PVA fibers by electrospinning.

[0152] 4) Fiber membrane dispersion and orientation: The collected PVA fiber membrane was immersed in deionized water and dispersed into a short fiber suspension by low-frequency ultrasonic treatment (100W power, 45 seconds). Subsequently, the suspension was introduced into a microfluidic device, and laminar shear force was used to align the short fibers along the flow direction to form a pre-oriented fiber bundle.

[0153] 5) Hot Stretching and Sizing: The aforementioned water-containing, oriented fiber bundles are guided through a hot air drying channel at 90°C. During this process, the PVA fiber surface softens and fuses, while a 3:1 stretch ratio is applied along the fiber axis for hot stretching. This step causes the short fibers to bundle and fuse under the combined action of heat and force into a dense, continuous PVA template fiber with a diameter of 40 μm, which is then wound and collected for later use.

[0154] The remaining steps are the same as in Example 1.

[0155] Stent performance testing

[0156] 1. Mechanical support

[0157] Bending stiffness test

[0158] This invention uses a bending stiffness measuring device for measurement, and the specific operation is as follows:

[0159] (1) Nest one end of the bracket into a plug gauge with an outer diameter of about 2.5 mm, with a nesting length of about 9 mm, and leave the other end free;

[0160] (2) Apply a vertically downward force to the braided bracket through the sensor extension rod, with a loading speed of 10 mm / min and a loading distance of 4 mm;

[0161] (3) Rotate the braided bracket and test at three different ends of the bracket in three circumferential directions. Take the average value of the test as the bending stiffness value of the bracket to eliminate the influence of the loading position on the result.

[0162] Bending stiffness is an important indicator for evaluating the ability of a support to resist axial deformation. The cantilever length of the support is L, and the deflection W is generated when the end of the sensor extension rod presses down on the support. B The testing software records the bending force (F) and deflection (W) in real time. B The change in bending stiffness (E) is plotted on a curve. I The bending stiffness can be calculated from the bending force-deflection curve data with a slope of k. A mechanical schematic diagram of bending stiffness is shown below. Figure 1 As shown, according to the cantilever beam theory, the calculation formula is as follows: Let the applied force at the free end be F, the distance from the clamping point to the loading point be L, and the deflection at the free end be W. B Then we have:

[0163] WB =FL 3 / 3E I

[0164] k=F / W B

[0165] From the above formula, we can obtain:

[0166] E I =kL 3 / 3

[0167] Where k is the slope of the force-displacement curve.

[0168] In this experiment, the blank stent, drug-loaded stent, and post-drug-release stent prepared in Example 1 were tested for bending stiffness. The test data are shown in Table 1.

[0169] Blank stent: A stent that possesses only a complete microchannel network structure but does not load any drugs within the channels, i.e., the stent obtained in step 4. It is an intermediate product in the preparation process of this invention and is used as a control benchmark to evaluate the impact of the drug loading step alone on the mechanical properties of the stent.

[0170] Drug-loaded stents: Drug-loaded stents were immersed in a specific release medium (phosphate-buffered saline PBS, pH=7.4, temperature 37℃), and samples were taken at 7, 14 and 30 days. After removal, the stents were rinsed and dried to obtain the drug-release stents at the corresponding time points, which were used to evaluate the effect of the degradation-drug release process on mechanical properties.

[0171] Table 1 Bending stiffness test data (N·mm) 2 )

[0172]

[0173] From Table 1 and Figure 2 Data shows that compared to the blank stent, the tortuosity of the drug-loaded stent decreased by only 1.28%, indicating that the drug loading process did not significantly weaken the stent's tortuosity. This suggests that the drug is mainly located in the microchannel / pore network inside the stent strut, and the drug reservoir is relatively decoupled from the main supporting structure. The tortuosity gradually decreased over time after drug release; the average daily decrease rate was 1.7% in the first 7 days, 1.9% in the first 14 days, and 1.5% in the first 30 days; the retention rate was greater than 55% in the first 30 days, indicating the potential to meet the short-term support requirements of coronary stents. PBS solution provides an accelerated degradation environment and involves static immersion; in the human vascular environment, due to blood flow carrying away acidic degradation products and different stent stress environments, the actual in vivo degradation rate is usually slower than the in vitro PBS immersion data.

[0174] Radial force test

[0175] Radial force is an important indicator for evaluating a stent's resistance to radial load deformation, and it is mainly divided into three stages: compression, retention, and recovery. According to the radial force testing standard ASTM F3067-14, the slope of the compression stage curve is defined as the stent's radial stiffness; the maximum radial force during the compression stage is defined as the stent's peak radial force. Radial stiffness and peak radial force are commonly used to characterize the stent's ability to resist compressive loads; the radial force at the point of use during the recovery stage is defined as the stent's chronic outward force, characterizing the stent's ability to dilate blood vessels.

[0176] This invention uses a Blockwise TTR2 radial force tester (Blockwise, TTR2, USA) to measure the radial force of the support. Taking the support in Example 1 as an example, the specific operation is as follows:

[0177] (1) Test the lumen temperature preheated to 37°C to simulate human body temperature, and place the stent in the center of the lumen;

[0178] (2) Set the gripping cavity to grip from the initial diameter of 3.0 mm to the predetermined diameter of 1.2 mm at a loading speed of 0.02 mm / s. The predetermined diameter is between the minimum diameter of the support and the working diameter to simulate the process of the support being loaded into the conveying system.

[0179] (3) Hold at the predetermined diameter for 60 seconds to simulate the delivery process of the stent in the blood vessel;

[0180] (4) Set the gripping cavity to return to its initial diameter at a speed of 0.02 mm / s to simulate the process of the stent being released from the sheath after being delivered to the target lesion vessel.

[0181] The radial stiffness and peak radial force of the stent were obtained after testing, and the data are shown in Table 2. The blank stent, drug-loaded stent, and post-drug-release stent used were the same as above.

[0182] Table 2 Radial performance test data

[0183]

[0184] From Table 2 and Figure 3 The data show that, compared to the blank stent, the radial stiffness and peak radial force of the drug-loaded stent did not show significant differences, indicating that the drug loading did not significantly weaken the stent's anti-collapse ability. With the extension of immersion time, the radial stiffness and peak radial force of the stent gradually decreased from 7 to 30 days after drug release, reflecting the impact of polylactic acid hydrolysis and degradation on mechanical properties; the changes were relatively rapid in the first 7 days, and then the rate of decrease slowed down.

[0185] 2. Biocompatibility

[0186] Human coronary artery endothelial cells, catalog number: CP-H087, purchased from Wuhan Pusaino Life Science Co., Ltd.

[0187] Cell culture: After purchasing the cell line, culture it according to the instructions, passage the purchased cells twice, and prepare a cell suspension.

[0188] Stent sterilization: The stent is placed in the sterilization chamber, evacuated to -0.1 MPa, and liquid CO2 is injected to 8 MPa. The temperature is then raised to 40°C to form a supercritical state. Maintaining 40°C and 8 MPa, 0.1% acetic acid vapor is injected and maintained for 90 minutes. The pressure is then slowly released to atmospheric pressure (at a rate of 1 MPa / min) to prevent microchannel collapse. The stent is rinsed three times with nitrogen to ensure no CO2 / acetic acid residue remains.

[0189] Cell seeding: After removal, place the scaffold on a clean bench and insert it into a 24-well culture plate, one scaffold per well. Then, add human tissue cell suspension at a ratio of 1.5 × 10⁶ cells per well. 4 / mL, the wells without the support were used as the control group. The control group and the experimental group (Examples 1-3) were performed in 5 parallel experiments. Then the 24-well plate was placed in an incubator for incubation (the temperature of the incubator was adjusted to 37°C, the CO2 concentration to 5%, and the humidity to 100%).

[0190] Cell metabolic fluids from days 1, 2, 3, 4, 5, 6, and 7 of in vitro cell culture were collected. The wells were rinsed three times with phosphate buffer, and then digested with trypsin solution for 3 minutes using a 1 ml pipette. The digestion solution was then discarded, and fresh cell culture medium was added to terminate the digestion. 300 μL of the cell suspension was pipetted and mixed with 300 μL of 0.4% trypan blue solution for staining. After thorough staining, three drops of the cell suspension were placed on a hemocytometer, and the cells were counted under an inverted biological microscope. The total cell count was then calculated. The results are shown in Table 3.

[0191] Table 3. Cell proliferation count results (×10) 4 (cell / mL)

[0192]

[0193] The concentration of albumin (ALB) in coronary artery endothelial cells was determined using a standard ALB assay. On days 1, 2, 3, 4, 5, 6, and 7 of human coronary artery endothelial cell culture, the culture medium from both the experimental and control groups was collected after cell consumption. A certain amount of the supernatant of the cell culture medium was then taken and an equal volume of ALB standard solution and a equal volume of double-liquid water were added to different 5 ml conical centrifuge tubes. Bromocresol green solution was then added to each centrifuge tube, allowing for thorough reaction and mixing. The wavelength of the UV spectrophotometer was adjusted to 628 nm, and the albumin concentration was measured using a 1 cm thick cuvette. The formula for calculating the albumin content is as follows:

[0194] Albumin content (g / L) = (A1-A0) / (A2-A0)×C0

[0195] A1: Absorbance of the solution in the sample tube; A2: Absorbance of the solution in the standard tube; C0: Concentration of the standard albumin solution; A0: Absorbance of distilled water. The results are shown in Table 4.

[0196] Table 4. Results of albumin (ALB) concentration determination (g / L)

[0197]

[0198] According to the data in Table 3, cells exhibited logarithmic growth from day 1 to day 4, resulting in a significant increase in albumin levels on day 4. Proliferation slowed from day 4 to day 6; and from day 6 to day 7, cells entered apoptosis, leading to a decrease in cell number. Cell data from Examples 1-3 showed no significant difference from the control group (P>0.05), remaining within the normal range of cell growth fluctuations, demonstrating that the scaffold did not release toxic substances and did not affect cell growth. According to the data in Table 4, cells were in a period of vigorous synthesis from day 1 to day 4, with albumin gradually accumulating with cell proliferation. On day 5, due to contact inhibition, cell metabolism shifted to a dormant state, reducing the albumin synthesis rate; and from day 6 to day 7, cell apoptosis occurred, resulting in a decrease in albumin content. The albumin content from Examples 1-3 showed no significant difference compared to the control group, demonstrating that cells still possessed high metabolic activity after the scaffold was added. In conclusion, this scaffold exhibits excellent biocompatibility, with Example 1 showing superior performance.

[0199] 3. Drug controlled release research

[0200] Plotting the standard curve of rapamycin concentration

[0201] A certain amount of rapamycin (RAPA) was dissolved in dimethyl sulfoxide (DMSO) to prepare RAPA / DMSO solutions with concentrations of 0.1 μg / mL, 0.5 μg / mL, 1 μg / mL, 5 μg / mL, 10 μg / mL, 20 μg / mL, and 40 μg / mL. Subsequently, these standard solutions were injected into a high-performance liquid chromatograph (HPLC) for detection, and the peak areas corresponding to each concentration were recorded. A RAPA concentration standard curve was plotted using the linear relationship between known concentrations and peak areas.

[0202] The drug-loaded scaffolds of Examples 1-3 and Comparative Examples 1-5 were placed in a certain amount of release medium (PBS) and incubated at 37°C. At the set time points, the supernatant of the release medium was taken as the test sample. Each time, 1 mL of release solution was taken (with the addition of an equal volume of fresh PBS at the same temperature), and the results were detected by HPLC. The cumulative release rate was calculated using the following formula, and the data are shown in Table 5.

[0203] Cumulative release rate calculation formula:

[0204] C i : Concentration of the i-th sample (μg / mL)

[0205] V: Volume of release medium (20 mL)

[0206] V 补 Cumulative fluid replacement volume (1 mL each time)

[0207] W 载药量 Initial drug loading of the stent (calibrated by solvent extraction method)

[0208] Solvent extraction method Weigh each group of drug-loaded stents and prepare 10 mL of a THF:water (8:2) mixed solvent containing 0.1% antioxidant (such as BHT) to prevent drug degradation. Cut the stents into 2-3 fragments, immerse the fragments in the extraction solution, and sonicate at 37°C (300W, 40kHz) for 30 minutes. Transfer to a 50°C water bath and shake for 4 hours. Centrifuge at 12000 rpm for 10 minutes, and filter the supernatant through a 0.22 μm PTFE membrane. Quantitatively analyze the concentration using HPLC. Prepare a series of rapamycin standards using the same extraction solvent. The drug loading formula is:

[0209]

[0210] C 测得 Drug concentration (μg / mL) determined by HPLC

[0211] V 提取液 Extraction volume (10 mL)

[0212] m 取样 Mass of the extracted stent (mg)

[0213] m 支架 Drug-eluting stent (mg)

[0214] Table 5. Cumulative release rate of rapamycin (%)

[0215]

[0216] From Table 5 and Figure 4 The results show that, experimentally verified, the use of 25-micron microchannels combined with 0.2 to 0.5 MPa pulsed pressurization technology can significantly improve the controlled release effect of rapamycin. Test data shows that this method (Example 1) can stabilize the drug release rate on the first day at approximately 23.1%, which is about 46% lower than the burst release rate of 42.8% of the traditional coating technology in Comparative Example 1, effectively avoiding the risk of initial burst release. Furthermore, it can achieve nearly complete release of 92.1% within four weeks. Model fitting analysis shows that the release curve (Example 1) is in high agreement with the Weibull equation (correlation coefficient R0). 2 The calculated shape parameter β value is approximately 0.89, falling within the range of 0.8-1.0. This β value characteristic indicates that the drug release of the stent of this invention is not a simple diffusion process, but rather exhibits an ideal diffusion-dissolution synergistic mechanism: that is, the diffusion behavior of the drug within the microchannel and the degradation / relaxation behavior of the polylactic acid scaffold synergistically drive drug release, thereby achieving an ideal curve of low initial burst release and long-term stable release.

[0217] This superior performance stems from the precise coordination between the microchannel structure design and pressure parameters. Regarding channel size, the 25-micron (Example 1) channel diameter perfectly balances drug delivery efficiency, preventing the excessively high burst release of 42.3% caused by the 40-micron (Comparative Example 5) channel, while avoiding the incomplete release problem of only 72.4% within 28 days in the 10-micron channel (Comparative Example 4); and it is superior to Examples 2 and 3 in balancing drug delivery efficiency. In terms of infusion process, the pulsed pressurization technology exhibits superior performance compared to constant pressure infusion. Compared to the release lag and incompleteness observed in Comparative Example 2 (0.2 MPa constant pressure) (cumulative release of 86.2% over 28 days), and the high burst release caused by Comparative Example 3 (0.5 MPa constant pressure) (38.5% on the first day), the pulsed pressure process used in Example 1 effectively overcomes capillary resistance and bubble lock-in effects within the microchannel, achieving full penetration and stable release of the drug within the microchannel. Only when the microchannel diameter is controlled at around 25 micrometers, and specific pulse pressure parameters are used, can the following three key indicators be achieved simultaneously: controlling the first-day burst release rate to less than one-quarter; ensuring that the release amount reaches about 90% within three weeks; and obtaining a stable and controllable release curve. This technical solution achieves a breakthrough performance balance in the field of drug sustained-release and controlled-release.

[0218] This invention addresses the core problems faced by traditional surface-coated drug-eluting stents, such as drug burst release, easy coating peeling, mutual constraints between mechanical properties and drug release effect, and inflammation caused by long-term foreign body residue. With the core research and development direction of "structural separation of drug load and mechanical support", this invention innovatively designs a continuous hollow microchannel network inside the stent strut as a structural drug reservoir. A three-dimensional tubular template is constructed using water-soluble sacrificial template fibers. A dynamic impregnation-dip slurry process allows polylactic acid (PLA) solution to fully penetrate and solidify, forming a composite embryo. The template is then gently dissolved to create a microchannel network. Finally, a pulsed pressure perfusion process achieves uniform drug loading, resulting in significant technical achievements: the stent's mechanical support performance remains stable after drug loading, fully meeting the short-term support requirements of interventional vascular treatment and effectively avoiding the weakening of the stent's mechanical load-bearing capacity by traditional drug loading methods; the drug release process is stable and controllable, significantly reducing the risk of burst release and achieving a long-term stable release effect that meets clinical treatment needs; the stent uses bioabsorbable PLA-based materials and a gentle preparation process, exhibiting excellent biocompatibility, avoiding significant inflammatory reactions or cytotoxicity, and gradually degrading and absorbing after completing its therapeutic mission, reducing long-term foreign body-related risks. This invention is not an optimization or improvement of a single performance aspect, but rather a synergistic improvement and breakthrough balance across multiple core dimensions, including mechanical support stability, controllable drug release, biocompatibility, and degradation safety, making it suitable for interventional treatment scenarios in cardiovascular, peripheral vascular, and non-vascular lumens.

Claims

1. A bioabsorbable polylactic acid scaffold based on microchannel drug delivery, characterized in that, include: A tubular mesh scaffold body formed of polylactic acid-based bioabsorbable material, the scaffold body being composed of multiple interconnected scaffold rods; wherein at least a portion of the scaffold rods form a continuous hollow microchannel network inside; the microchannel network has an inner diameter of 20-30 micrometers, and the microchannel network is used to contain therapeutic drugs or their formulations.

2. The bracket according to claim 1, characterized in that, The therapeutic drugs are antiproliferative drugs and / or anti-inflammatory drugs.

3. The bracket according to claim 1, characterized in that, The polylactic acid-based bioabsorbable material is selected from poly-L-lactic acid, poly-racemic lactic acid, and their copolymers.

4. A method for preparing a stent as described in any one of claims 1-3, characterized in that, Includes the following steps: S1: Preparation of water-soluble sacrificial template fibers; S2: A three-dimensional tubular template is prepared using water-soluble sacrificial template fiber, and then polylactic acid solution is infiltrated into the three-dimensional tubular template and cured through impregnation or dynamic impregnation-dip slurry process to obtain a three-dimensional tubular composite preform; S3: The composite preform is subjected to strengthening, shaping and curing treatment; S4: Immerse the composite preform in a solvent to dissolve and remove the sacrificial template fibers, thereby forming the microchannel network inside the support rod; S5: The therapeutic drug solution is infused into the microchannel network using pressure-driven perfusion and the solvent is removed to complete the drug loading.

5. The method according to claim 4, characterized in that, The water-soluble sacrificial template fiber is prepared from polyvinyl alcohol, alginate, sucrose fiber, or gelatin.

6. The method according to claim 4, characterized in that, The polylactic acid solution in step S2 has a mass fraction of 8-12 wt%, and the solvent is selected from dichloromethane, chloroform, or a mixture thereof.

7. The method according to claim 4, characterized in that, The temperature of the strengthening, shaping and curing treatment in step S3 is higher than the glass transition temperature of polylactic acid material but lower than its melting temperature.

8. The method according to claim 4, characterized in that, The solvent used in step S4 is deionized water, and the water is changed multiple times until the template fiber is completely dissolved and removed.

9. The method according to claim 4, characterized in that, The active pressure-driven perfusion in step S5 adopts pulsed pressurization, with the pressure cycle being an intermittent cycle of 0.2 MPa × 30 s and an alternating cycle of 0.5 MPa × 10 s.

10. The use of the stent according to any one of claims 1-3 as a material for the preparation of cardiovascular, peripheral vascular and / or non-vascular luminal interventional therapies.

Images