Patent Application
20250170303
The present disclosure relates to the field of medical device technology, and particularly, to a tissue engineering blood vessel for in vivo rapid recellularization and a method for preparing same.
Along with social progress and the improvement of living standards, cardiovascular diseases have become the leading cause of death, threatening people's health, and there is an increasing clinical demand for small-diameter artificial blood vessels (less than 6 mm in diameter). At present, China sees about 870,000 peripheral artery disease surgeries each year, while the United States sees over 100,000 peripheral artery disease surgeries each year. In China, 150,000 vascular trauma surgeries are performed annually, while in the United States, 73,000 are performed. In China, the number of coronary artery bypass surgeries performed per year exceeds 50,000 and is expected to reach 150,000 within the next decade, and the counterpart in the United States is 500,000, with each patient requiring an average of about 2.5 bypass blood vessels. Autologous vein grafts have to be collected surgically, which may easily lead to extended surgery time, graft incision infection, and prolonged ICU stays. Additionally, in many patients, the great saphenous vein does not meet the clinical requirements. Each year, about over 3.5 million hemodialysis treatments are given. The patients require a safe, reusable arteriovenous (AV) channel as dialysis access. According to the Chinese National Renal Data System (CNRDS), the rate of hemodialysis is increasing annually by 13-14%. In China, there are over 120 million patients with chronic kidney disease, with more than 800,000 registered for long-term dialysis in 2020, and projections indicate that this number will exceed 2.3 million by 2030. Of these, 87% use arteriovenous fistulas, 10% use central venous catheters, and about 1% use artificial vascular access. In the United States, about 600,000 hemodialysis treatments are given to patients with chronic diseases each year. Of these, 65% use arteriovenous fistulas, 15% use central venous catheters, and 20% use artificial vascular access, with the use of artificial vascular access on the rise. Patients undergoing arteriovenous fistula surgery need to wait for 3-6 months before the arteriovenous fistula is available as dialysis access. However, for about one-third to half of these patients, the arteriovenous fistula becomes unusable after the 3- to 6-month waiting period, and such an access may degenerate due to its susceptibility to infection, thrombosis, or aneurysm formation and may thus be clinically unusable. Central venous catheters can only be used as temporary dialysis access, with an annual infection rate of up to 200% in patients. Conventional expanded polytetrafluoroethylene (ePTFE) artificial blood vessels are prone to infection, with a 2-year secondary patency rate of about 50% and a very low 5-year patency rate.
Existing commercially available small-diameter artificial blood vessels (such as those made of Teflon and polyurethane or decellularized blood vessels from allogeneic or xenogeneic sources) feature low patency rates and are prone to thrombosis and infection, which affect clinical outcomes. The primary reason is the failure in effective vascular wall recellularization, including endothelialization and vascular smooth muscle cell recellularization. Therefore, there is a pressing clinical need to develop high-performance, clinically translatable, small-diameter artificial blood vessels.
Conventional tissue engineering techniques require several months to construct the small-diameter tissue engineering vascular graft (TEVG). These processes are usually cost-inefficient, with uncertain safety profiles for use in humans. Moreover, they may require special methods for storage and transportation before surgery, making these techniques unsuitable for widespread clinical use. The in situ tissue engineering technology avoids complex processes such as in vitro cell isolation, culture, and cell-scaffold integration in the conventional tissue engineering techniques through direct implantation of a tissue engineering blood vessel into the body, where it recruits endogenous cells to regenerate the vessel in situ. Moreover, this technology meets the clinical need for “on-demand availability”, offering significant advantages for promoting clinical applications. However, current research on in situ tissue engineering faces notable challenges, such as insufficient vascular endothelialization, which leads to thrombosis, and insufficient vascular wall recellularization with vascular smooth muscle cells, which causes infection or aneurysm formation. These issues result in low long-term patency rates, significantly limiting the clinical translation.
To solve the shortcomings in the prior art, such as the recellularization failure or slow recellularization on artificial blood vessels, low patency rates, susceptibility to infection, and high degeneration rates, the present disclosure provides a tissue engineering blood vessel for in vivo rapid recellularization and a method for preparing same. By employing biomaterial modification and tissue engineering biomimetic techniques, the present disclosure significantly enhances the adhesion, growth, and proliferation of seed cells on degradable scaffold materials. During culture, the degradable polymer composite scaffold completely degrades, reducing the culture period. Through decellularization techniques, a decellularized matrix-based tissue engineering blood vessel with a special biomimetic extracellular matrix 3D structure is acquired. The biomimetic extracellular matrix structure allows the blood vessel, when implanted into the body, to realize rapid in vivo recellularization, construct a new blood vessel in situ in the body, and improve the blood vessel's performance. The present disclosure meets the clinical need for “on-demand availability”, facilitating clinical translation.
To achieve the above objectives, according to one aspect of the present disclosure, provided is a method for preparing a tissue engineering blood vessel for in vivo rapid recellularization, including:
Preferably, in step S1, the mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for the inner tube wall of the tubular polymer composite scaffold is (15:1)-(50:1), and the mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for the outer tube wall of the tubular polymer composite scaffold is (5:1)-(15:1).
Preferably, the tubular polymer scaffold has an inner diameter of 1-35 mm, a tube wall thickness of 0.01-3.5 mm, a length of 0.5-120 cm, and a tube wall porosity of 85-99.9%.
Preferably, the culture condition simulating the arterial pulsatile flow in the body in step S2 is as follows:
Preferably, the decellularization in step S3 includes the following steps:
Preferably, the seed cells are uniformly seeded on the tubular polymer composite scaffold at 0.1-3.0×106 cells/centimeter.
Preferably, the seed cells include, but are not limited to, vascular smooth muscle cells or fibroblasts from an adult human or mammal, vascular smooth muscle cells from an umbilical vein/umbilical artery, and vascular smooth muscle cells or fibroblasts differentiated from a mesenchymal stem cell and an induced pluripotent stem cell.
Preferably, the seed cells are primary to 10th-passage cells, preferably primary to 4th-passage cells, from a single donor or a cell bank.
According to another aspect of the present disclosure, further provided is a tissue engineering blood vessel for in vivo rapid recellularization, which is prepared through the following steps:
Preferably, proteins in the extracellular matrix of the tissue engineering blood vessel include, but are not limited to, collagen I and collagen III.
According to a third aspect of the present disclosure, provided is use of the tissue engineering blood vessel described above. The use includes use as a replacement of an esophagus, a trachea, a ureter, a urethra, a bile duct, a fallopian tube or a vas deferens, use in preparing an allogeneic biological heart valve, and/or use as a surgical patch.
In general, compared to the prior art, the above technical solutions contemplated by the present disclosure can achieve the following beneficial effects:
(1) According to the tissue engineering blood vessel for in vivo rapid recellularization and the method for preparing same provided by the present disclosure, the degradable polymer composite scaffold with the special tube wall structure is prepared from polyglycolic acid and polyethylene glycol in a certain ratio by electrospinning. The fibers on the outer side of the vascular scaffold contain a higher proportion of polyethylene glycol than the fibers on the inner side of the tube wall. When the seed cells are seeded on the scaffold and subjected to in vitro perfusion culture in a certain condition that simulates the arterial pulsatile flow in the body, the degradable polymer scaffold completely degrades, and an engineering blood vessel composed of the seed cells and the extracellular matrix is acquired. Then the engineering blood vessel is decellularized to give the tissue engineering blood vessel consisting of the extracellular matrix with a biomimetic structure. The decellularized matrix-based tissue engineering blood vessel has a special biomimetic extracellular matrix 3D structure. The extracellular matrix with the biomimetic structure allows the decellularized matrix-based blood vessel, after being implanted into the body, to be rapidly recellularized and secrete a new extracellular matrix, facilitating the self-repair and renewal of the vascular wall's extracellular matrix to resist infection. Meanwhile, the lumen of the decellularized matrix-based blood vessel is rapidly endothelialized (1 week to 1 month), improving the patency rate of the blood vessel. The blood vessel of the present disclosure is superior to conventional artificial blood vessels.
(2) According to the tissue engineering blood vessel for in vivo rapid recellularization and the method for preparing same provided by the present disclosure, the degradable polymer composite scaffold, compared to composite scaffold fibers and pure polyglycolic acid fibers in in vitro culture, can promote cell adhesion, increasing the success rate of cell seeding. During culture, particularly in the early stages of culture, the tube wall structure, compared to pure polyglycolic acid materials, allows the seeded seed cells to adhere to scaffold fibers containing polyethylene glycol, so that the cells will not easily fall off. In the static and dynamic perfusion culture processes, the outer side of the tube wall degrades first due to the higher proportion of polyethylene glycol, resulting in pores that facilitate the migration and proliferation of the seed cells on the tube wall toward the inner side of the tube wall. The inner side of the tube wall degrades continuously and slowly due to the high proportion of polyglycolic acid, thereby maintaining the mechanical performance of the tube wall of the scaffold material. The present disclosure overcomes the shortcomings of conventional tissue engineering techniques, such as insufficient cell viability and slow growth after seed cells are seeded on the scaffold material.
(3) According to the tissue engineering blood vessel for in vivo rapid recellularization and the method for preparing same provided by the present disclosure, a combination of static culture and perfusion culture in a condition that simulates the arterial pulsatile flow in the body is used to promote the uniform growth, migration, and proliferation of seed cells on the fiber surfaces of the inner and outer sides of the vascular scaffold material, accelerating the synthesis of extracellular matrix proteins. The proteins in the extracellular matrix secreted by the cells on the scaffold include, but are not limited to, collagen I and collagen III. Thus, the inner wall structure of the lumen is smooth and uniform, and the vascular wall does not easily break and has good mechanical performance. This significantly increases the success rate of in vitro blood vessel culture and reduces culture costs. The present disclosure has more advantages of being commercialized.
(4) According to the tissue engineering blood vessel for in vivo rapid recellularization and the method for preparing same provided by the present disclosure, the seed cells include, but are not limited to, adult human vascular smooth muscle cells, human fibroblasts, human umbilical artery and human umbilical vein vascular smooth muscle cells, and human vascular smooth muscle cells and human fibroblasts differentiated from human mesenchymal stem cells including human adipose-derived stem cells and human induced pluripotent stem cells, vascular smooth muscle cells and fibroblasts from other adult mammalian sources (including pigs, cattle, sheep, etc.), and vascular smooth muscle cells and fibroblasts differentiated from mesenchymal stem cells including adipose-derived stem cells and induced pluripotent stem cells. The sources are diverse and unrestricted, thus facilitating large-scale, commercial cultures.
To make the objectives, technical solutions, and advantages of the present disclosure more clear, the present disclosure is further described in detail below with reference to the drawings and examples. It will be appreciated that the specific examples described herein are merely illustrations of the present disclosure and do not limit the present disclosure. In addition, the technical features described below in the embodiments of the present disclosure can be combined with each other as long as they do not conflict with each other.
In the present disclosure, a degradable polymer composite scaffold with a special tube wall structure is constructed from polyglycolic acid and polyethylene glycol by electrospinning, with the fibers on the outer side of the vascular scaffold containing a higher proportion of polyethylene glycol. When seed cells are seeded on the scaffold and subjected to in vitro perfusion culture in a certain condition that simulates the arterial pulsatile flow in the body, the degradable polymer degrades, and an engineering blood vessel composed of the seed cells and an extracellular matrix is acquired. The engineering blood vessel is further decellularized to give a tissue engineering blood vessel composed of the extracellular matrix with a biomimetic structure. The method specifically includes the following steps:
S1: Preparation of Tubular Polymer Composite Scaffold with Special Tube Wall Structure
Polyglycolic acid and polyethylene glycol are separately dissolved in an organic solvent to give uniform, stable spinning solutions. The two spinning solutions are used for electrospinning in a certain ratio and received by a Teflon-coated metal rod, so as to give an oriented fiber tubular polymer composite scaffold having a special tube wall structure. The mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for an inner tube wall of the tubular polymer composite scaffold is higher than the mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for an outer tube wall of the tubular polymer composite scaffold.
Specifically, the spinning solutions have a polyglycolic acid concentration of 1-500 mg/mL and a polyethylene glycol concentration of 1-500 mg/mL; the mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for the inner tube wall of the tubular polymer composite scaffold is (15:1)-(50:1), and the mass ratio of the polyglycolic acid spinning solution to the polyethylene glycol spinning solution for the outer tube wall of the tubular polymer composite scaffold is (5:1)-(15:1). The electrospinning fibers have a diameter of 5 nanometers to 50 micrometers; the organic solvent includes, but is not limited to, one or more of hexafluoroisopropanol, trichloromethane, and the like.
In the preparation of tissue engineering blood vessels, scaffold preparation is crucial. In the present application, two polymers, polyglycolic acid and polyethylene glycol, are selected as spinning solutions to prepare the degradable polymer composite scaffold with the special tube wall structure. The fibers on the outer side of the vascular scaffold contain a higher proportion of polyethylene glycol. In the early stages of culture, when the tube-wall cells have not yet fully grown and proliferated, the polyglycolic acid fibers can maintain the scaffold's mechanical performance. The polyethylene glycol fibers can promote the adhesion, migration, and proliferation of seed cells in the early stages of culture and ensure sufficient cell growth on the tube wall in static and pulsatile flow conditions. The two composite material fiber structures gradually degrade during the culture and cell growth, and the polyethylene glycol fibers degrade faster than polyglycolic acid. Some of the fibers on the outer side of the tube wall degrade preferentially due to the higher proportion of polyethylene glycol, resulting in pores that facilitate the migration and proliferation of the seed cells on the tube wall toward the inner side of the tube wall. The inner side of the tube wall degrades continuously and slowly due to the high proportion of polyglycolic acid, thereby maintaining the mechanical performance of the tube wall of the scaffold material. This further facilitates the migration and growth of cells in the middle and inner layers without affecting the mechanical performance of the tube wall of the composite fiber scaffold. Consequently, the vascular cells secrete more extracellular matrix more rapidly, improving the mechanical performance of the wall of the cultured blood vessel, significantly increasing the success rate of culture, and reducing the culture time.
Seed cells are seeded on the tubular polymer composite scaffold, cultured statically for 1-10 days, and then cultured for 3 weeks to 3 months in a certain perfusion culture condition simulating the arterial pulsatile flow in the body to allow the seed cells to adequately grow and proliferate on the composite scaffold. The tube-wall polymer gradually degrades, while the vascular cells secrete the extracellular matrix to gradually enhance the mechanical performance of the tube wall. Finally, an engineering blood vessel composed of the seed cells and the extracellular matrix is acquired.
Specifically, the culture condition simulating the arterial pulsatile flow in the body is as follows: firstly, a static in vitro culture of 1-10 days; secondly, a culture of 6-8 days in an arterial pulsatile flow at a pressure of 10-90 mmHg; thirdly, a culture in an arterial pulsatile flow at an increased pressure of 90-110 mmHg; and finally, after the scaffold material completely degrades, a culture of one week in an arterial pulsatile flow at an increased pressure of 110-160 mmHg to significantly improve the mechanical performance of the vascular wall. The in vitro culture period is 3 weeks to 3 months, preferably 6-12 weeks; all dynamic cultures are perfusion cultures.
Seed cell seeding is one of the key steps in tissue engineering blood vessel preparation. In the present application, the seed cells include, but are not limited to, adult human vascular smooth muscle cells, human fibroblasts, human umbilical artery and human umbilical vein vascular smooth muscle cells, and human vascular smooth muscle cells and human fibroblasts differentiated from human mesenchymal stem cells including human adipose-derived stem cells and human induced pluripotent stem cells, vascular smooth muscle cells and fibroblasts from other adult mammalian sources (including pigs, cattle, sheep, etc.), and vascular smooth muscle cells and fibroblasts differentiated from mesenchymal stem cells including adipose-derived stem cells and induced pluripotent stem cells.
More specifically, the cells are primary to 10th-passage cells, preferably primary to 4th-passage cells, from one or more donors or a cell bank and are cultured in vitro for 3 weeks to 3 months, most preferably for 6-12 weeks. The cell culture medium contains a high content of sugar, insulin (0.1-5.0 mg/mL), the growth factors bFGF and/or EGF, penicillin (1-300 U/mL), streptomycin (1-500 mg/mL), and vitamin C (0.5-800 mg/mL), and may be, for example, DMEM. In weeks 0-4 of the culture, 5-50% human or mammal serum is added to the culture medium. On the following days, upon the culture medium replacements once every 3-7 days, the human or mammal serum concentration is reduced by 1-20%. The procedures continue until the culture is finished.
More specifically, the cells are uniformly seeded on the composite scaffold at 0.1-3.0×106 cells per centimeter, and the optimal seeding density is 1.0-2.0×106 cells/centimeter.
The degradable polymer composite scaffold fibers gradually degrade in the culture condition described above, and an engineering blood vessel consisting of the seed cells and the extracellular matrix secreted by the seed cells is acquired. In the present application, by combining static culture with culture in a condition simulating the arterial pulsatile flow in the body, the uniform growth and proliferation of the seed cells on the inner and outer sides of the vascular scaffold material and thus the synthesis of the extracellular matrix are promoted. The extracellular matrix secreted by the cells on the scaffold includes, but is not limited to, collagen I fiber and collagen III fiber, ensuring that the tube wall has good mechanical properties, including a burst pressure of 1000-5000 mmHg, a suture tension of 100-250 g, and an elastic modulus of 0.5-3.0 MPa. Particularly, during early dynamic culture, the tube wall does not easily break, and the structure of the inner wall of the lumen is smooth. The success rate of in vitro culture has increased from 80-85% of conventional techniques to 95-100%. Therefore, the present disclosure has more advantages for commercialization.
The engineering blood vessel is decellularized to give an extracellular matrix 3D structure that is looser than arterial vessel walls in the body, which is conducive to vascular wall recellularization after the implantation of the blood vessel into the body and ensures that the vascular wall composed of the extracellular matrix has excellent mechanical performance.
The decellularized tissue engineering blood vessel is mainly composed of the extracellular matrix and may include trace amounts of residual polyglycolic acid and/or polyethylene glycol. The proteins of extracellular matrix of the vascular wall include, but are not limited to, collagen I and collagen III. The amount of cells in the vascular wall is less than 1-2%. If the blood vessel is completely decellularized, the cell DNA content is less than 0.1%-0.001%. Meanwhile, the decellularization does not affect the mechanical performance of the blood vessel. Unlike conventional tissue engineering blood vessels that require a prolonged culture period and special methods for preservation and transportation, the blood vessel of the present disclosure can be preserved at room temperature for a long time, meeting the clinical need for “on-demand availability”.
The decellularization may be conducted using a chemical reagent, electric treatment, etc. In one preferred example, the decellularization specifically includes the following steps:
The tissue engineering blood vessel for rapid recellularization and the method for preparing same provided by the present disclosure are further described in detail below using different specific examples. The specific examples are as follows:
Polyglycolic acid and polyethylene glycol were separately dissolved in hexafluoroisopropanol solvent to prepare uniform, stable spinning solutions. The tube wall of the scaffold was 480 μm thick. During the spinning, spinning solutions of two mass ratios were used. The mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 30:1 for the 240-micrometer thick tube wall on the inner side, and the mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 5:1 for the 240-micrometer thick tube wall on the outer side. An oriented fiber tubular polymer composite scaffold was obtained.
Human aortic vascular smooth muscle cells at the 3rd passage were resuspended in a low sugar DMEM culture medium containing 10% FBS, seeded on the tubular polymer composite scaffold at a density of 1.0×106 cells/cm, statically cultured for 5 days in a 37° C., 5% CO2 incubator, then cultured for 1 week in a pulsatile flow at 30 mmHg, cultured for 1 week in a pulsatile flow at 60 mmHg, cultured for 1 week at 90 mmHg, and cultured for 3 weeks at 120 mmHg to give an engineering blood vessel, with the culture time being less than 7 weeks.
According to the formula shown in the table above, exact amounts of the reagents were added to a clean wide-mouth glass flask. The flask was shaken on a shaker for 2-3 h, until the solution became clear, to fully dissolve the reagents, and thus the decellularization reagent was prepared. The decellularization reagent was preserved at 4° C. before use.
The engineering blood vessel was placed into a sterile centrifuge tube and perfused at room temperature with the decellularization reagent for 5 h. After the decellularization was complete, the blood vessel was perfused and washed with a sterile PBS solution for 3 h to remove the decellularization reagent. Finally, the resulting decellularized tissue engineering blood vessel was placed in normal saline at 4° C. for preservation before use.
Polyglycolic acid and polyethylene glycol were separately dissolved in hexafluoroisopropanol solvent to prepare uniform, stable spinning solutions. The tube wall of the scaffold was 480 μm thick. During the spinning, spinning solutions of two mass ratios were used. The mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 15:1 for the 240-micrometer thick tube wall on the inner side, and the mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 5:1 for the 240-micrometer thick tube wall on the outer side. An oriented fiber tubular polymer composite scaffold was obtained.
Human umbilical vein vascular smooth muscle cells at the 3rd passage were resuspended in a low sugar DMEM culture medium containing 10% FBS, seeded on the tubular polymer composite scaffold at a density of 0.8×106 cells/cm, statically cultured for 5 days in a 37° C., 5% CO2 incubator, then cultured for 1 week in a pulsatile flow at 30 mmHg, cultured for 1 week in a pulsatile flow at 60 mmHg, cultured for 1 week at 90 mmHg, and cultured for 2 weeks at 120 mmHg to give an engineering blood vessel, with the culture time being not greater than 8 weeks.
The procedures for decellularization are the same as those in Example 1. Please refer to Example 1.
Polyglycolic acid and polyethylene glycol were separately dissolved in hexafluoroisopropanol solvent to prepare uniform, stable spinning solutions. The tube wall of the scaffold was 480 μm thick. During the spinning, spinning solutions of two mass ratios were used. The mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 30:1 for the 240-micrometer thick tube wall on the inner side, and the mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 5:1 for the 240-micrometer thick tube wall on the outer side. An oriented fiber tubular polymer composite scaffold was obtained.
Human fibroblasts at the 3rd passage were resuspended in a low sugar DMEM culture medium containing 10% FBS, seeded on the tubular polymer composite scaffold at a density of 0.5×106 cells/cm, statically cultured for 5 days in a 37° C., 5% CO2 incubator, then cultured for 1 week in a pulsatile flow at 30 mmHg, cultured for 1 week in a pulsatile flow at 60 mmHg, cultured for 1 week at 90 mmHg, and cultured for 2 weeks at 120 mmHg to give an engineering blood vessel, with the culture time being less than 6 weeks.
The procedures for decellularization are the same as those in Example 1. Please refer to Example 1.
Polyglycolic acid and polyethylene glycol were separately dissolved in hexafluoroisopropanol solvent to prepare uniform, stable spinning solutions. The tube wall of the scaffold was 480 μm thick. During the spinning, spinning solutions of two mass ratios were used. The mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 50:1 for the 240-micrometer thick tube wall on the inner side, and the mass ratio of polyglycolic acid spinning solution to polyethylene glycol spinning solution was 5:1 for the 240-micrometer thick tube wall on the outer side. An oriented fiber tubular polymer composite scaffold was obtained.
Vascular smooth muscle cells differentiated from human induced pluripotent stem cells, at the 3rd passage, were resuspended in a low sugar DMEM culture medium containing 10% FBS, seeded on the tubular polymer composite scaffold at a density of 1.0×106 cells/cm, statically cultured for 5 days in a 37° C., 5% CO2 incubator, then cultured for 1 week in a pulsatile flow at 30 mmHg, cultured for 1 week in a pulsatile flow at 60 mmHg, cultured for 1 week at 90 mmHg, and cultured for 3 weeks at 120 mmHg to give an engineering blood vessel, with the culture time being less than 7 weeks.
The procedures for decellularization are the same as those in the above examples. Please refer to Example 1.
Before and after decellularization, the engineering blood vessels in Examples 1˜4 were subjected to burst pressure, suture tension, maximum tensile stress, and maximum tensile strain tests, and the test results were compared. There were no statistically significant differences between the test results before and after decellularization. The specific testing method and results are as follows:
The burst pressure test was conducted on an in-house pressure measuring system equipped with a precise pressure gauge. The pressure measuring system was filled with a PBS buffer, and a test sample with a length of 8 cm was connected to the pressure measuring system. The pressure of the pressure measuring system was gradually increased until a sudden pressure drop was detected and a rupture in the vessel was observed by reviewing the video. The maximum pressure was recorded and used as the burst pressure.
The suture tension refers to the force when sutures tear the tissue. In this test, the suture tension was measured according to the ANSI/AAMI/ISO 7198 standard. A test sample with a length of 2 cm was prepared. One end of the vessel was sewn with 6-0 Prolene sutures at 2 mm from the edge, with a total of 3 stitches placed at 120-degree intervals. Each suture was tied independently to form a loop. The end of the vessel without sutures and one of the sutures were fixed to a universal tensile testing system. The stretching speed was adjusted to 30 mm/min, and the maximum force was recorded. The test was repeated on the other two sutures. The average of the 3 test results is the suture tension.
The test tissue vessel was cut into 5-mm wide rings, and their thickness and diameter were measured and recorded. The rings were hung by the two ends on paper clips, and then 2 paper clips were fixed to a universal tensile testing system. The initial stretch length was about 10% of the breaking length. The stretching speed was adjusted to 30 mm/min. The test was continued until the blood vessel ring was broken, and the maximum tensile stress and the maximum tensile strain were obtained.
According to the data from Table 1 and Table 2, the tissue engineering blood vessels prepared in the present disclosure possess good biocompatibility, good mechanical performance, ease to suture, and tolerance to high burst pressures and tensile forces, making them suitable for clinical applications. In addition, the decellularization method of the present disclosure has little impact on the mechanical performance of the vascular walls of the engineering blood vessels.
In conclusion, the tissue engineering blood vessel for in vivo rapid recellularization provided by the present disclosure has the advantages of:
It will be appreciated that the above examples are merely illustrative of the technical solutions of the present disclosure and are not intended to limit them. Although the present disclosure has been described in detail with reference to preferred examples, those of ordinary skill in the art will understand that modifications or equivalent replacements can be made to the technical solutions of the present disclosure without departing from the spirit and scope of the present disclosure, all of which should be included within the claimed scope of the present disclosure.
| Number | Date | Country | Kind |
|---|---|---|---|
| 202210906816.5 | Jul 2022 | CN | national |
This application is the national phase entry of International Application No. PCT/CN2023/098423, filed on Jun. 5, 2023, which is based upon and claims priority to Chinese Patent Application No. 202210906816.5, filed on Jul. 29, 2022, the entire contents of which are incorporated herein by reference.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/CN2023/098423 | 6/5/2023 | WO |
