Patent Application
20250161015
The present invention belongs to the field of biomedical materials and medical apparatuses and instruments, and specifically relates to a biodegradable ureteral stent and a preparation method therefor.
Ureteral stents have been widely used in urological surgery, which are applicable to upper urinary tract surgery, lithotripsy using lithotripters, dilation of ureterostenosis, and other treatment processes, and can play important roles of draining urine and preventing ureterostenosis after implantation into ureters.
Existing clinically used ureteral stents are all non-degradable and are prepared from soft polyurethane elastomers or silicone rubber materials. The polyurethane stents are harder and easy to catheterize; and the silicone rubber stents are softer and have slightly poor catheterization performance. After catheterization of these two types of non-degradable ureteral stents, the hardness of the stents will remain unchanged. When a drainage function is completed, the ureteral stents need to be pulled out through an invasive operation, namely through a cystoscope, and patients will be in pain and are possible to have infections and other complications. Therefore, research and development of biodegradable ureteral stents have important clinical values.
According to the needs of clinical cases, the ureteral stents will have differences in drainage time. Short-term temporary drainage of urine usually takes about 1-2 weeks, such as mild ureteral injuries, simple stones, pre-stenting, etc.; and common temporary drainage usually takes 3-6 weeks. Thus, different degradation materials need to be used. However, common basic requirements for the stents include that they should have mechanical performance of elastic materials, and their degradation fragments can be completely excreted out of the body as soon as possible, otherwise the risk of a series of complications will be increased.
It has been reported (Lumiaho, J, J. Endourol. 1999, 13, 107-112; Laaksovirta, S Laurila M. et al. Jurol, 167:1527, 2002) that the ureteral stents are manufactured by using biodegradable polylactide (PLLA) or a lactide/glycolide copolymer (PLGA) as a raw material. However, the PLLA and PLGA materials are plastic, which lack softness, have harder degraded fragments, and have a great possibility of being embedded or getting stuck in the renal pelvis, leading to various complications.
A Chinese patent No. CN1672739A and a Chinese patent No. CN112516390A disclose a biodegradable ureteral stent, which relates to a glycolide-ε-caprolactone copolymer material and has mechanical performance of elastic materials within a certain composition proportion range. However, such copolymer formed by a soft chain monomer (ε-caprolactone) and a hard chain monomer (glycolide) usually becomes increasingly hard with degradation in water. This is because a hard chain structure in the material, such as a glycolide chain segment, usually has a crystallization trend in a degradation process, and thus, the material is easy to get stuck or remain in the renal pelvis and other parts.
In summary, the biodegradable materials obtained in the prior art cannot achieve the material properties which have not only mechanical performance similar to that of silicone rubbers and polyurethane elastic materials. but also can make its degradation fragments not to be stuck or not to remain in the renal pelvis. That is still a key point restricting the research progress in the art up to now and is also a primary reason why the degradable ureteral stents are currently not commercially available on the market.
In view of the above disadvantages of the prior art, the present invention provides a biodegradable ureteral stent. The stent has mechanical strength of elastic materials, higher initial hardness (or modulus) and convenience in catheterization, gradually becomes increasingly soft with degradation in urine, and is easier to be excreted out of the body.
Technical solutions of the present invention are as follows.
A biodegradable ureteral stent prepared from a composite material, wherein the composite comprises at least a glycolide-ε-caprolactone copolymer, an ethylene oxide polymer, and barium sulfate, which are blended together, with their relative contents as follows:
According to the ureteral stent of the present invention, the above object is achieved by using the composite material formed by blending the degradable glycolide-ε-caprolactone copolymer, the ethylene oxide polymer, and the barium sulfate.
The present invention and related studies find that when the stent is harder, degraded fragments are easy to get stuck at the position of the renal pelvis. Thus, the softness and hardness of a substrate is a key influencing factor. Common degradable elastic materials, such as glycolide-ε-caprolactone copolymers, L-lactide/ε-caprolactone copolymers, etc., usually become increasingly hard in a degradation process, and degraded fragments or segments are more difficult to pass through narrow sites of the ureteral. However, studies of the present invention find that within an appropriate comonomer proportion range, by using an appropriate polymerization process, the glycolide-ε-caprolactone copolymer can gradually become soft with degradation in water or urine, and is an elastic material with good mechanical strength and suitable softness and hardness, and the appropriate comonomer proportion range is that the weight percentage content of the glycolide is 51-58%. When the weight percentage content of the glycolide is greater than 58%, the copolymer has an obvious crystallization trend and becomes hard in a degradation process; and when the weight percentage content of the glycolide is less than 51%, the copolymer has too poor mechanical performance and is too soft.
The present invention also finds that by adding a certain content of the ethylene oxide polymer, such as polyethylene glycol or polyoxyethylene, to materials of the glycolide-ε-caprolactone copolymer, the composite material can be further promoted to gradually become softer in a degradation process, smoother in surface, and easier to disintegrate and fracture, thereby being conducive to excretion of fractured fragments of the tubular stent.
According to the biodegradable ureteral stent of the present invention, the barium sulfate used is a commonly used radiopaque material in medical developer, which has good compatibility with the above materials and has a certain reinforcing effect.
According to the biodegradable ureteral stent of the present invention, an intrinsic viscosity of the glycolide-ε-caprolactone copolymer measured at 25±1° C. in hexafluoroisopropanol at a concentration of 0.1 g/dl is 1.30-3.00 dl/g, and when the viscosity is greater, a degradation maintenance time is longer. Thus, a drainage time of the stent may be controlled by the intrinsic viscosity.
According to the biodegradable ureteral stent of the present invention, the ethylene oxide polymer includes polyethylene glycol, polyethylene glycol monomethyl ether, polyethylene glycol dimethyl ether, and polyoxyethylene.
As a preference, a weight proportion of the ethylene oxide polymer in the composite material is 2%-8%, and when the content of the ethylene oxide polymer is higher, the composite material is softer and faster to degrade.
According to the biodegradable ureteral stent of the present invention, a molecular weight of the polyethylene glycol, the polyethylene glycol monomethyl ether, the polyethylene glycol dimethyl ether, or the polyoxyethylene is 1,000-1,000,000 Da.
As a preference, according to the biodegradable ureteral stent of the present invention, the molecular weight of the ethylene oxide polymer including the polyethylene glycol, the polyethylene glycol monomethyl ether, and the polyethylene glycol dimethyl ether is 5,000-40,000 Da.
As a preference, according to the biodegradable ureteral stent of the present invention, the ethylene oxide polymer is the polyoxyethylene, and the molecular weight is 50,000-400,000 Da.
According to the biodegradable ureteral stent of the present invention, the barium sulfate may be replaced by other medical developers, including bismuth subcarbonate and a metal developer, and the developer used may include one or more.
According to the degradable ureteral stent of the present invention, a preparation method for the glycolide-ε-caprolactone copolymer includes:
under the protection of nitrogen, sequentially adding stannous octanoate with a mass ratio of 0.005%-0.1%, ε-caprolactone, and glycolide to a reactor with a stirrer; raising the temperature of a reaction system from room temperature to 165° C.-200° C. within 30 minutes under stirring, maintaining the temperature for 18-30 hours, and then holding the system under vacuum for 1-4 hours; and transferring a copolymer out of the reactor, further breaking the copolymer, and placing the copolymer in a vacuum oven for vacuum drying at 50° C.-110° C. for 8-24 hours.
The biodegradable ureteral stent of the present invention is of a hollow circular tubular structure (1), a fixing structure (2) for preventing sliding is disposed at two ends or one end, the fixing structure is preferably in a circular tubular coil shape, a tube wall is also provided with a plurality of penetrating drainage side holes (3), and a tube diameter is 1.0 mm-4.0 mm (millimeters).
The biodegradable ureteral stent of the present invention is prepared by a melting extrusion method. The method specifically includes:
An initial modulus at 100% deformation of the degradable ureteral stent of the present invention is 2-10 MPa, and the modulus at 100% deformation after degradation is not greater than an initial value. An initial Shore hardness (A) of the material used is 70-95, and the Shore hardness (A) after degradation is not greater than an initial value.
According to the biodegradable ureteral stent of the present invention, during melting extrusion of a tube, various additives may be conventionally added to achieve different purposes, including but not limited to, a plasticizer, a lubricant, a dye, an antioxidant, an anti-hydrolysis agent, a melt thickener, a chain extender, a reinforcing agent, and a polymer modifier. These additives are conducive to improving the processing performance, degradation performance, surface performance, and mechanical performance of the stent.
Compared with the prior art, the present invention has the following beneficial effects.
In the present invention, by controlling the content of the glycolide monomer in the glycolide-ε-caprolactone copolymer, the copolymer has appropriate mechanical strength and softness and hardness and gradually becomes soft with prolonging of the degradation time. Meanwhile, a certain proportion of the ethylene oxide polymer is added, such that the polymer has a lower constant stretch modulus after degradation and is softer and smoother, thus is more suitable for preparing a biodegradable ureter stent.
A 0.04% stannous octanoate catalyst, 570 gram (g) of an ε-caprolactone monomer (CL), and 580 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 180° C. within 25 minutes under the protection of nitrogen, a reaction was carried out for 25 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 2 hours. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 1 (PGC1).
Weight percentage contents of the glycolide and the ε-caprolactone in the copolymer were determined by a 1H nuclear magnetic spectrum, and hexafluoroisopropanol was used as a solvent.
The copolymer was prepared into a hexafluoroisopropanol solution at a concentration of 0.1 g/dl, and the intrinsic viscosity was tested by an Ubbelohde viscometer at 25° C.
The copolymer was prepared into a sheet with a thickness of 2 mm by a hot pressing molding method on a press vulcanizer at 140° C., and the Shore hardness (A) of the material was tested by a Shore hardness tester.
By using the same method, the copolymer was prepared into a dumbbell plate with a thickness of 2 mm, and the tensile strength at break and elongation at break of the material were tested on a universal mechanical tester at a speed of 200 mm/minute.
An in vitro degradation experiment of the material was carried out in simulated urine at 37° C., and the Shore hardness (A) of the material was tested regularly.
A 0.02% stannous octanoate catalyst, 540 g of an ε-caprolactone monomer (CL), and 590 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 170° C. within 20 minutes under the protection of nitrogen, a reaction was carried out for 28 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 1 hour. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 2 (PGC2).
Performance tests were carried out on the copolymer by using the method in Example 1. Test results are listed in Table 1.
A 0.03% stannous octanoate catalyst, 530 g of an ε-caprolactone monomer (CL), and 600 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 190° C. within 30 minutes under the protection of nitrogen, a reaction was carried out for 22 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 4 hours. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 3 (PGC3).
Performance tests were carried out on the copolymer by using the method in Example 1. Test results are listed in Table 1.
A 0.05% stannous octanoate catalyst, 480 g of an ε-caprolactone monomer (CL), and 600 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 195° C. within 30 minutes under the protection of nitrogen, a reaction was carried out for 18 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 4 hours. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 4 (PGC4).
Performance tests were carried out on the copolymer by using the method in Example 1. Test results are listed in Table 1.
A 0.01% stannous octanoate catalyst, 470 g of an ε-caprolactone monomer (CL), and 630 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 190° C. within 30 minutes under the protection of nitrogen, a reaction was carried out for 20 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 2 hours. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 5 (PGC5).
Performance tests were carried out on the copolymer by using the method in Example 1. Test results are listed in Table 1.
A 0.01% stannous octanoate catalyst, 470 g of an ε-caprolactone monomer (CL), and 660 g of a glycolide monomer (GA) were sequentially placed in a 3 L reactor. The temperature of a system was raised to 184° C. within 26 minutes under the protection of nitrogen, a reaction was carried out for 23 hours at a stirring speed of 10-20 r/min, and the system was held under vacuum for 3 hours. A copolymer was transferred out of the reactor, further broken, and placed in a vacuum oven for vacuum drying at 90° C. for 24 hours to obtain a glycolide-ε-caprolactone copolymer 6 (PGC6).
Performance tests were carried out on the copolymer by using the method in Example 1. Test results are listed in Table 1.
The above results indicate that when the content of glycolide in the glycolide-ε-caprolactone copolymers of the present invention is higher, the hardness is higher; when the weight content of glycolide is 51%-58%, the Shore hardness (A) is equivalent to that of polyurethane (75-95) and silicone rubbers (50-70), and the copolymers are more suitable for use as ureteral stents; and degradation results in the simulated urine at 37° C. show that the hardness of PGC1 to PGC5 is decreased in a degradation process. The weight content of glycolide in PGC6 is greater than 58%, and the hardness is higher and is increased with increase of the degradation time. When the content of glycolide is lower than 51%, the tensile strength is lower.
200 g of the glycolide-ε-caprolactone copolymers (PGC) prepared above, a certain amount of polyethylene glycol 20000 (PEG2) or polyethylene glycol 5000 (PEG5), and medical grade barium sulfate (Ba) were evenly mixed, and further blended and granulated through a twin-screw extruder at an extruder temperature of 120° C.-150° C. Then, granules of resulting composite materials were prepared into sheets of the composite materials of the above three materials with a thickness of 2 mm by a vulcanizing press at 140° C., which were used for tensile strength and hardness tests. Dynamic friction coefficients of the composite materials were tested by a method specified by ASTM-D1894. Results are shown in Table 2.
The above results indicate that changes in the hardness of the composite materials containing PEG are obviously decreased in a degradation process in simulated urine with increase of the degradation time, and the higher content of PEG leads to a greater decrease of the hardness; the hardness of the composite material 3 without addition of PEG is decreased to a lower extent compared with that of the composite materials with addition of PEG; for the composite material 7 and the composite material 8, in the glycolide-ε-caprolactone copolymers used, the weight content of glycolide is greater than 58%, and thus the Shore hardness (A) is increased or slightly changed with prolonging of the degradation time; and the composite materials with addition of PEG have smaller dynamic friction coefficients, suggesting that surfaces are smoother.
200 g of the glycolide-ε-caprolactone copolymers (PGC) prepared above, a certain amount of polyoxyethylene (PEO, molecular weight: 200,000 Da), and medical grade barium sulfate (Ba) were evenly mixed, and further blended and granulated through a twin-screw extruder at an extruder temperature of 120° C.-140° C. Then, granules of resulting composite materials were prepared into sheets with a thickness of 2 mm by a vulcanizing press at 140° C., which were used for tensile strength and hardness tests. Dynamic friction coefficients of the composite materials were tested by a method specified by ASTM-D1894. Results are shown in Table 3.
Changes in the hardness of the above composite materials of the glycolide-ε-caprolactone copolymers, the polyoxyethylene, and the medical grade barium sulfate in a degradation process have similar rules, which are obviously decreased with increase of the degradation time. The higher content of polyoxyethylene leads to a greater decrease of the hardness. However, when the content of polyoxyethylene is too high, the strength is greatly reduced, which is not conducive to fixing. The composite materials with addition of polyoxyethylene have smaller dynamic friction coefficients, suggesting that surfaces are smoother.
200 g of the glycolide-ε-caprolactone copolymers (PGC) prepared above, a certain amount of polyethylene glycol 20000 (PEG2) or polyethylene glycol 5000 (PEG5), and medical bismuth subcarbonate (Bi) were evenly mixed, and further blended and granulated through a twin-screw extruder at an extruder temperature of 120° C.-140° C. Then, granules of resulting composite materials were prepared into sheets of the composite materials of the above three materials with a thickness of 2 mm by a vulcanizing press at 140° C., which were used for tensile strength and hardness tests. Dynamic friction coefficients of the composite materials were tested by a method specified by ASTM-D1894. Results are shown in Table 4.
Properties and change rules of the above composite materials of the glycolide-ε-caprolactone copolymers, the polyethylene glycol, and the bismuth subcarbonate are similar to those of the above composite materials, that is, changes in the hardness of the composite materials in a degradation process in simulated urine are obviously decreased with increase of the degradation time, the composite materials with addition of PEG have smaller dynamic friction coefficients, and the above rules are not related to the type of a developer used.
The composite materials prepared in the above examples were subjected to extrusion molding by an extruder at 120° C.-150° C. to obtain a degradable elastic tube body 1, respectively, wherein the tube body 1 had an outer diameter of 2 mm and an inner diameter of 1.1 mm. The tube body was subjected to bending shaping at 50° C.-80° C. to form a fixing structure with a coiled tubular coil 2 formed at one end or two ends. Then, the tube body was subjected to punching (longitudinal aperture: 1.0 mm) at an equal distance (hole distance: 50 mm) by a punching device to form drainage holes 3, so as to obtain a degradable ureteral stent shown in
On a universal mechanical tester, the tensile strength at break and constant stretch modulus (tensile strength at the elongation of 100%, same as the Shore hardness for characterizing the softness and hardness of a material) of materials were tested at a speed of 200 mm/min, and changes in the constant stretch modulus in an in vitro degradation process in simulated urine at 37° C. were tested. Results are shown in Table 5.
After miniature pigs were selected and subjected to general anesthesia, ureteral stents were implanted into left and right ureters of the miniature pigs by a ureteroscopy method, in which tubular coils at upper ends were fixed in the renal pelvis, and tubular coils at lower ends were fixed in the bladder. Whether the ureteral stents with the situations of slipping, fracturing, and excretion at different time points were observed by an X-ray after ureteroscopy. The constant stretch modulus of the stents was tested by a universal mechanical tester at different time points to characterize changes in the softness and hardness. Results are shown in Table 5.
The above results show that the ureteral stents 1, 2, and 3 are added with the ethylene oxide polymer, while the ureteral stents 4 and 5 are not added with the ethylene oxide polymer. All the degradable stents can be completely excreted from the urinary system of animals, and the fracture time depends on the molecular weight of materials, the proportions of comonomers, the amount of the ethylene oxide polymer added, and other factors, including process conditions during processing, etc. However, since the ureteral stents 1, 2, and 3 with addition of the ethylene oxide polymer have lower constant stretch modulus after degradation and are softer and smoother in surface, these ureteral stents have a shorter complete excretion time in vivo than the ureteral stents 4 and 5 without addition of the ethylene oxide polymer, and thus have a smaller possibility of incrustation in vivo. The various groups of ureteral stents have certain but not large differences in initial modulus, and can all undergo catheterization smoothly. The modulus of a commercial silicone rubber stent and a polyurethane stent is not changed with the degradation time. The tensile strength and modulus of the degradable ureteral stents of the present invention are between those of the silicone rubber stent and the polyurethane stent, and are closer to those of the polyurethane stent.
| Number | Date | Country | Kind |
|---|---|---|---|
| 202210848552.2 | Jul 2022 | CN | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/CN2023/102900 | 6/27/2023 | WO |
