RU2764051C1 - Method for in vitro production of a personalized cell-populated vascular prosthesis - Google Patents

Abstract

FIELD: medicine.

SUBSTANCE: invention relates to medicine, namely to vascular surgery. A method for the manufacture of a personalized cell-populated vascular prosthesis in vitro consists in the formation of a prosthesis frame from a mixture of polyhydroxybutyrate/valerate at a concentration of 3 to 7% (m/v) and c-polycaprolactone at a concentration of 9 to 14% (m/v) by the method of electrospinning, coating framework with autologous fibrin with further colonization of the inner surface with autologous colony-forming endothelial cells isolated from peripheral blood in an amount of at least 700,000 cells per 1cm² and preparation of the endothelial layer under conditions of a pulsating flow-through bioreactor for further implantation into the vascular bed.

EFFECT: method makes it possible to obtain tissue-engineered small-diameter vascular prostheses covered with autologous fibrin and populated with the patient's own endothelial cells and can be used to replace damaged blood vessels and to perform shunting operations.

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Description

SUBSTANCE: invention relates to medicine, namely to a technology for creating a personalized small-diameter vascular prosthesis, and can be used in vascular surgery to replace damaged blood vessels and to perform shunt operations.

Cardiovascular diseases, including ischemic heart disease, cerebrovascular disease and peripheral arterial disease, are the leading causes of death worldwide. Despite the widespread use of interventional therapies, vascular bypass remains the most commonly used strategy to effectively treat cardiovascular disease and ensure long-term patency of blood vessels. For these operations, it is preferable to use autografts, such as the internal thoracic artery and the great saphenous vein of the thigh. These vessels demonstrate excellent patency after implantation, but are not always available in the required quality in different patients. Synthetic vascular prostheses made of polytetrafluoroethylene, polyethylene terephthalate and polyurethane are successfully used for prosthetics of large diameter vessels (>6 mm). At the same time, synthetic vascular prostheses with an internal diameter of less than 4 mm are not used in clinical practice due to the risk of thrombus formation and neointima hyperplasia, leading to implant occlusion. Therefore, the development of new effective vascular prostheses or artificial blood vessels is of great importance for cardiovascular surgery.

Today, tissue engineering offers alternative approaches for repairing damaged tissues. Vascular tissue engineering strategies are based on the use of two key components: an artificial extracellular matrix (scaffold) and cells. The ideal scaffold creates both mechanical support and a microenvironment favorable for cell colonization. The choice of material and method of scaffold fabrication directly affect the temporary mechanical support and microenvironment it provides to cells until they synthesize their own extracellular matrix. The use of synthetic materials (biodegradable polyesters, polyester amides, polyurethanes) allows you to control the mechanical properties of the scaffold to a greater extent and set the necessary parameters. In turn, the best microenvironment for cells, which creates signals for migration, proliferation, and differentiation, is formed using materials of biological origin (collagen, elastin, gelatin), since they have potential binding sites for cell receptors. However, these materials have insufficient strength and in some cases (allogeneic and xenogenic origin of materials) can cause an immune response of the body after implantation.

Scaffolds can be populated with cells both in the patient's body, when a cell-free tubular scaffold is implanted at the site of blood vessel reconstruction, and in vitro in a bioreactor using autologous (own) or allogeneic (from a donor) endothelial and/or smooth muscle cells. The main problem of cell-free vascular scaffolds is the insufficient thromboresistance of their inner surface, which cannot ensure 100% patency of these prostheses after implantation. In turn, pre-population of the inner surface of tissue engineering scaffolds with endothelial cells prevents their thrombosis after contact with blood. The main problem at this stage is the choice of the source of endothelial cells. To date, the possibility of using mature, progenitor and stem cells, as well as induced pluripotent cells has been shown. Stem and induced pluripotent cells, although they are very promising sources, their clinical use is currently limited by the imperfection of the production technology and ethical issues. In turn, mature endothelial cells are characterized by relatively low proliferation. Thus, the complexity of the development of tissue-engineered vascular prostheses lies in the choice of materials for creating a scaffold with optimal mechanical properties and high bio- and hemocompatibility, as well as a source of endothelial cells that will allow obtaining autologous cells in the amount necessary for endothelialization of a vascular implant.

This problem can be solved by combining synthetic and biological materials in the manufacture of the scaffold, as well as by populating it with autologous cells with a high degree of proliferation. This approach will make it possible to create a personalized vascular prosthesis with optimal mechanical properties and a thromboresistant inner surface.

A known method of manufacturing a multilayer tissue-engineered vascular prosthesis, the inner layer of which is formed by mature endothelial cells, the middle layer by mature smooth muscle cells and the outer layer by mature fibroblasts (GB 2489081 A: Int C1. C12N 5/071; A61F 2/00; A61L 27/38 ; A61L 27/50. Multilayered vascular tubes. Organovo Inc (US); Inventors: Khatiwala CB (US), Murphy K. (US); appl. no. 1203622.4; filed 05.26.2010; pub. date 05.26.2010). Allogeneic or autologous cells and tissues such as aortic or umbilical cord cells can be used as cell sources. Also, mature cells can be obtained by differentiation from multipotent stem cells or progenitor cells. The cell lines are cultured separately until the formation of confluent layers, which are collected in layers around a pin or cylinder gel-coated with agarose, agar or other hydrogels such as polyethylene glycol to create a tubular product. The resulting multilayer cell prosthesis is cultivated until it is fully formed, and then placed in a pulsating bioreactor that provides mechanical load to mature the prosthesis and improve its mechanical properties. The disadvantages of this method include the low strength of the products obtained as a result of the absence of a scaffold that would provide mechanical support to the cells. Tissue-engineered vascular prostheses implanted into the bloodstream, consisting only of cells, have a high risk of aneurysm formation and wall rupture. In addition, this approach requires a large amount of cellular material, and obtaining it from the proposed sources, for example, from the aorta, is an invasive procedure with additional damage to the patient's or donor's vessels.

Another way to create tissue-engineered blood vessels of small diameter is based on seeding cells on tubular forms or pins, such as fibrin threads or collagen-coated silicone tubes (US 20110033927 A1: Int C1. CI2N 5/071; CI2N 5/0775; CI2N 5/ 00. Methods of generating small-diameter tissue engineered blood vessels, Worcester Polytechnic Institute (US), Inventors: Rolle M. (US), Parekh DP (US), Doshi KJ (US), Gwyther T. (US), appl. no. 12/752708; filed 04/01/2010; pub. date 02/10/2011). In the first case, bone marrow-derived mesenchymal stem cells are cultured on fibrin threads in a v-shaped chamber coated with agarose, which prevents cell adhesion to the chamber surface. In the process of cultivation, the cells form the wall of the vascular prosthesis, while fibrin is gradually resorbed. In the second case, aortic smooth muscle cells are cultured on the outer surface of a silicone tube coated with collagen. After the cells form the wall of the prosthesis, the silicone tube is removed. The disadvantages of these products are the absence of an endothelial layer, the presence of which is critical for thromboresistance of a vascular prosthesis, as well as their low strength, since the intercellular matrix formed by cells during in vitro cultivation is not able to withstand the load exerted by blood flow on the vascular wall.

An approach to the manufacture of a tissue-engineered prosthesis is described, which consists in the use of a supporting form, a matrix, and endothelial and smooth muscle cells (US 8192348 B2: Int C1. A61F 2/04. Engineered blood vessels. Regents of the University Minnesota (US); Inventors: Tranquillo RT (US); Ross J. (US); Reyes M. (US); appl. no. 10/562955; filed 06/01/2004; pub. date 01/13/2005). To do this, a matrix containing cells is applied to the tubular form of the carrier. The tubular form has a porous structure, is permeable to nutrients and gases and can be made from various biocompatible materials, such as lactic and polyglycolic acids, polydimethylsiloxane, polyglycerol sebacate, polylacide-glycolide copolymer, polydioxanone, polygluconate, polycaprolactone, polycarbonates, polyamides, polyanhydrides others, as well as their combinations. The matrix into which endothelial and smooth muscle cells are introduced has a fibrillar structure and is made predominantly of fibrin or fibronectin, amphiphilic diblock copolymers or amphiphilic triblock copolymers or peptides. Also, the matrix can be made from laminin, proteoglycans or matrigel. The cells are injected directly into the matrix and the resulting composition is applied to the carrier form. Vessels (arteries or veins), microvascular tissue, or umbilical vein can be used as cell sources. Cells can be either autologous or allogeneic. When using donor cells, therapy with the use of pharmacological drugs such as prednisone and prednisolone, azathioprine, mycophenolate mofetil, methotrexate, cyclophosphamide, monoclonal antibodies and others is envisaged to suppress the immune rejection reaction of the tissue-engineered prosthesis after its implantation. In addition to differentiated cells, embryonic stem cells, mesenchymal stem cells, endothelial progenitor and smooth muscle progenitor cells can be used. Stem cells can be obtained from any suitable tissue such as bone marrow, brain, spinal cord, cord blood, peripheral blood, liver, placenta, adipose tissue or muscle. Stem cells are differentiated in vitro into mature endothelial smooth muscle cells using differentiation factors. This approach has disadvantages associated with the location of the supporting tubular shape relative to the endothelial layer. Since the matrix with endothelial cells is applied to the outer surface of the carrier form, if the prosthesis is used together with this supporting form, the endothelium will not come into contact with the blood, which increases the risk of thrombosis, and when the carrier form is removed, the endothelial layer may be damaged, which can also promote thrombosis. Moreover, endothelial cells are enclosed in a fibrin matrix, which limits the formation of a continuous endothelium on the inner surface of the prosthesis.

A known method of manufacturing a tissue-engineered vascular prosthesis by seeding cells on the outer surface of a tubular scaffold, in the process of culturing the cells infiltrate the scaffold and completely cover it (US 20030068817 A1: Int C1. A61K 45/00; C12N 5/06. Vascular tissue engineering. Inventors: Gazit D. (IL); Gafni Y. (1L); Turgeman G. (IL); Pelled G. (IL); appl. no. 10/234707; filed 05.09.2002; pub. date 10.04.2003). Cultivation can be carried out both under static conditions and in a bioreactor. Vascular cells, mesenchymal stem cells, endothelial progenitor cells, and others from the bone marrow or peripheral blood can be used to populate the scaffold. The first layer is formed from endothelial cells or cells capable of differentiating into endothelial cells. The scaffold has a fibrillar structure and is made of a bioresorbable polymer, such as cellulose, polylactide-polyglycolide copolymer, and others. To form a new tissue around the fibrillar scaffold, an additional coating with polymers that support cell proliferation, such as dextran, chitosan, fibrinogen, fibrin, collagen, gelatin, and others, can be used. The support polymer is applied to the scaffold already populated with cells. The disadvantages of this method include the colonization of the outer part of the scaffold by cells, which limits the possibility of endothelial formation on the inner surface of the prosthesis. Also, the application of a supporting polymer to a scaffold populated with cells prevents the flow of gases and nutrients to the cells, which negatively affects their viability.

Known tissue engineering vascular prosthesis, consisting of a biocompatible bioresorbable scaffold and cells or cell layers, cultured in a bioreactor or static conditions (US 9290742 B2: Int C1. CI2N 5/00; CI2N 5/071; A6L 27/8; A6L 27/36; A6L 27/38, A6L 27/50, Tissue engineered blood vessel, Cordis Corporation (US), Inventors: Cooper K. (US), Chun I. (US), Colter DC (US), Dhanaraj S. (US); Gosiewska A. (US); Seyda A. (US); Fang CH (US); Yang C. (US); appl. no. 13/799520; filed 04.13.2013; pub. date 08.01.2013). Natural (collagen, fibrin, elastin, and their combinations) and synthetic polymers (lactides, glycolides, polycaprolactone, dioxanone, and their copolymers) can be used as scaffold polymers. Scaffolds can be made in the form of woven materials, meshes, foams, non-woven fibrous materials obtained by extrusion, electrospinning, melting, and other methods. The scaffold is populated with one or more cell types: multipotent or pluripotent stem cells, such as progenitor smooth muscle and progenitor vascular endothelial cells, placental and umbilical cord cells. The scaffold is populated with cells immediately before the surgical operation, or several days before, and cultured in a bioreactor. The main disadvantage of this prosthesis is the use of blood vessel cells, which require additional damage to the patient's vessels. In addition, progenitor smooth muscle and progenitor vascular endothelial cells cannot be obtained in the required amount, as a result of their limited proliferation. And cells of the placenta and umbilical cord are of allogeneic origin, and their use as part of a tissue-engineered construct can lead to an immune response.

Also, a tissue-engineered vascular structure based on a bioresorbable polymer scaffold made of polyglycolic acid, or polylactic acid, or polysebacate, or their copolymers is described (CA 2426819 A1: Int C1. A61F 2/06; A61f 2/36. Tissue-engineered vascular structures. Childrens Medical Center Corp (US), Inventors: Bischoff J. (US), Kaushal S. (US), Mayer J. (US), Perry TE (US), appl. no. 60/244277, filed 10/30/2001, pub .date 05.10.2002). Alternatively, the scaffold can be made from polymers that make up the intercellular substance: collagen, laminin, elastin, fibronectin. The inner surface of the graft is populated with progenitor endothelial cells from the patient's peripheral blood under conditions of a pulsating bioreactor. In this case, the shear stress applied to the cells is gradually increased from 1 to 25 dynes/cm ² during the first two days of culturing the prosthesis in the bioreactor. The disadvantages of this design and its manufacture include a rapid increase in shear stress to high values. With this mode of cultivation in a flow bioreactor, most of the cells are washed off the surface of the prosthesis by a liquid current. In addition, collagen and laminin are less effective in retaining cells on the surface of the material under shear stress, compared with fibrin (Chlupáč J, Filová Ε, Riedel Τ, et al. Attachment of human endothelial cells to polyester vascular grafts: pre- coating with adhesive protein assemblies and resistance to short-term shear stress Physiol Res. 2014; 63(2): 167-177).

The closest technical solution is a method for manufacturing a tissue-engineered vascular prosthesis using a mixture of fibrinogen, thrombin and cells (US 20040115176 A1: Int C1. A61K 45/00; C12N 5/08. Fibrin-based tissue-engineered vasculature. Inventors: Swartz DD (US ); Andreadis ST (US); appl. no. 10/692381; filed 10/23/2003; pub. date 06/17/2004). The prosthesis has a tubular shape, and its wall consists of a fibrin gel containing cells of autologous origin. To obtain fibrin, fibrinogen and thrombin are obtained from the patient's blood. Vascular smooth muscle cells and/or fibroblasts are introduced into the fibrin gel and formed around a cylindrical pin. For optimal cell orientation in the wall of the prosthesis, a voltage can be applied to it, for example, using a fluid flow through the lumen of a tubular product in a pulsating bioreactor. The inner surface of the prosthesis can be populated with endothelial cells. Also, fibrin gel can be combined with a porous scaffold made of decellularized elastin or a copolymer of polylactide and polyglycolide. The main disadvantage of this approach is the use of mature smooth muscle and endothelial cells of the patient, since their proliferative capacity is very low and their availability is extremely limited. In addition, fibrin is subject to rapid degradation by proteases secreted by smooth muscle cells. Therefore, this prosthesis is characterized by a rapid loss of strength. At the same time, the combination of fibrin with rapidly resorbable scaffolds made of elastin or a copolymer of polylactide and polyglycolide will not sufficiently enhance the strength of the structure.

The technical result of the proposed solution is the technology for manufacturing in vitro tissue-engineered vascular prosthesis of small diameter, covered with autologous fibrin and populated with the patient's own colony-forming endothelial cells (Fig. 1). The technical result is achieved by combining the following parameters:

1. The main part of the prosthesis, the so-called scaffold, is made from biocompatible bioresorbable polymers: polyhydroxybutyrate / valerate (PHBV) and ε-polycaprolactone (PCL). These polymers have a strength that is optimal for the manufacture of a vascular prosthesis that can withstand the load exerted by the blood flow. In addition, products made from PHBV and PCL have a long bioresorption period, which ensures the preservation of the integrity of a tissue-engineered prosthesis based on a scaffold from a mixture of these polymers until the formation of its own tissues.

2. A tubular scaffold is made from a mixture of PHBV and PCL using the electrospinning method. The scaffold material is multidirectional polymer fibers with a diameter of 1 to 3 microns, forming a porous structure, while the porosity of the material is from 45% to 90%.

3. To increase the biocompatibility and adhesive properties of the inner surface of the polymer scaffold, it is coated with an autologous fibrin polymer. Fibrin is the end product of the coagulation cascade and is a natural bioresorbable matrix (Matveeva V.G., Khanova M.Yu., Antonova L.V., Barbarash L.S. Fibrin is a promising material for tissue vascular engineering. Bulletin of Transplantology and Artificial Organs 2020;22(1):196-208). Biomaterials based on it have high biocompatibility and controlled bioresorption through fibrinolysis with the formation of non-toxic products. Fibrin fibers contain cell adhesion sites that promote cell adhesion, migration, and proliferation. The three-dimensional porous structure of fibrin is easily populated by cells, as it is able to support their migration and vital activity due to the diffusion of nutrients and oxygen into the fibrin matrix and the removal of metabolic products (Aper T., Teebken O.E., Steinhoff G., Haverich A. Use of a fibrin preparation in the engineering of a vascular graft model Eur J Vase Endovasc Surg 2004;28:296-302).

4. Fibrin polymer is made from fibrinogen obtained from the patient's own peripheral blood. The use of autologous fibrinogen avoids an immune reaction to the vascular prosthesis after its implantation.

5. To create a fibrin layer on the surface of the vascular prosthesis, the PHBV/PCL scaffold is impregnated with a fibrinogen solution and polymerized by adding thrombin and calcium chloride (CaCl ₂ ).

6. To ensure thromboresistance of the vascular prosthesis, its inner surface is populated with autologous colony-forming endothelial cells (CPEC). CPEK can be isolated from peripheral blood. These cells belong to progenitor cells, in terms of morphology, phenotype, genome and transcriptome they are close to mature endothelial cells, while they have a high proliferative potential. In addition, it has been shown that CPEC can be isolated from the blood of patients with coronary heart disease (Matveeva V.G., Khanova M.Yu., Velikanova E.A., Antonova L.V., Sardin E.S., Krutitsky S. .S, Barbarash O.L. Possibility of obtaining and characterization of colony-forming endothelial cells from the peripheral blood of patients with coronary heart disease. Cytology 2018; 60(8): 598-608). Availability and high proliferative activity make CPEC a promising candidate for endothelialization of tissue-engineered vascular prostheses.

7. Autologous CPECs are isolated from the patient's peripheral blood and cultured to obtain the required number of cells.

8. Endothelization of a tissue-engineered vascular prosthesis is carried out in a pulsating flow bioreactor. The colonization of the CFEC prosthesis in the flow bioreactor promotes the orientation of cells in the direction of the flow of the culture medium and the maintenance of a functionally active endothelial phenotype by the cells.

The proposed technical solution is a personalized biodegradable small-diameter vascular prosthesis populated with endothelial cells, the manufacture of which is carried out as follows.

Stage 1. Production of a polymer tubular scaffold.

- Prepare a solution of a mixture of polymers containing PHBV at a concentration of 3 to 7% (m/v) and PCL at a concentration of 9 to 14% (m/v), in trichloromethane.

- Tubular scaffold with wall thickness from 50 to 600 microns. are produced by electrospinning of a PHBV/PKL solution at a voltage of 15 to 25 kV, a polymer solution feed rate of 0.5 to 3 ml/h, using a needle with an inner diameter of 20G to 27G and a collector - a rotating metal pin with a diameter of 1 to 6 mm.

- After fabrication, the scaffold is dried until the solvent has completely evaporated.

Stage 2. Modification of the surface of the polymer scaffold with fibrin.

- Take blood from the patient in sterile test tubes with sodium citrate.

- Isolate the precipitate of fibrinogen from the blood. To reduce time costs, achieve the most efficient fibrinogen yield (up to 80% of the fibrinogen present in the blood plasma), preserve the physico-mechanical and biological properties of the future fibrin (Weis-Fogh, US Fibrinogen prepared from small blood samples for autologous use in a tissue adhesive system Eur Surg Res 20,381,1988 Dietrich M, Heselhaus J, Wozniak J, et al Fibrin-based tissue engineering: comparison of different methods of autologous fibrinogen isolation Tissue Eng Part C Methods 2013;19(3):216 -226. doi:10.1089/ten.TEC.2011.0473) use the low ethanol ethanol precipitation method. To do this, the blood is centrifuged at 2000 x g for 10 min to obtain plasma. The resulting plasma is cooled to 4°C. 16% ethanol (4°C) is added to the plasma with constant stirring so that the ratio of plasma to ethanol solution is 1:1. Immediately after this, the plasma solution in ethanol is shaken on the neck at a temperature of 4°C, then centrifuged at 1000 × g and 4°C for 20 minutes. The supernatant is decanted, the precipitate is dissolved in a saline solution of 0.9% NaCl with HEPES to a fibrinogen concentration of 30-40 mg/ml, and a solution of aprotinin 100 CIE/ml is added to suppress fibrinolysis.

- Carry out the polymerization of fibrin on the surface of PHBV/PCL scaffold. To do this, the PHBV/PCL scaffold is impregnated with the resulting fibrinogen solution. After that, a solution of thrombin at a concentration of 500 U/ml and CaCl ₂ at a concentration of 40 mmol/l is applied to the surface of the prosthesis to polymerize fibrin. After polymerization, the prostheses are immersed in phosphate-buffered saline with ε-aminocaproic acid at a concentration of 2 mg/ml and stored until the cells are populated.

Stage 3. Endothelization of the inner surface of the fibrin-modified PHBV/PCL scaffold and obtaining a personalized vascular prosthesis.

- Take blood from the patient in sterile test tubes with heparin in a volume of at least 20 ml.

- Isolate CPEK from the blood. To do this, a mononuclear fraction of peripheral blood is isolated on a Histopaque 1077 density gradient. The cells obtained from the interphase are washed with phosphate-buffered saline and centrifuged (this procedure is carried out twice). The cell suspension was resuspended in complete EGM-2MV medium containing 5% fetal bovine serum, 2% penicillin/streptomycin and 0.25 g/ml amphotericin B and seeded in collagen-coated culture flasks. Cultivation is carried out in CO ₂ -incubator at 37°C and 5% CO ₂ . The first 2 days, the medium is changed daily, in the future - after 2-3 days. At the end of the week of cultivation, the cells are removed with accutase, washed with phosphate-buffered saline with 5% fetal bovine serum, resuspended in complete nutrient medium and transferred to a fibronectin-coated culture plate. The cells are cultivated until the formation of 70-80% confluent culture, after which the passage is performed. Cultivation is continued until the required number of cells is obtained.

- For endothelization, CPEK is applied to the inner surface of the fibrin-modified PHBV/PCL scaffold in an amount of at least 700,000 cells per 1 cm ² and cultivated under static conditions for 2 days. After that, the tissue-engineered structure is connected to the system of a pulsating flow bioreactor and cultivated at a shear stress of 1 to 1.5 dynes/cm ² during the day. Then gradually increase the shear stress to a value of 2 to 3 dynes/cm ² and cultivate in this mode from 2 to 10 days.

- Received personalized vascular prosthesis inner surface lined with a monolayer of endothelial cells, stored in a culture medium under sterile conditions at a temperature of 37°C and 5% CO ₂ .

The essence of the invention is illustrated by illustrations.

Fig. 1. Scheme of a longitudinal section of a personalized cell-populated vascular prosthesis of small diameter: 1 - porous wall of the prosthesis made of polyhydroxybutyrate/valerate and ε-polycaprolactone; 2 - fibrin coating; 3 - endothelial cells; 4 - internal lumen of the prosthesis. Scheme of a flow bioreactor: 1-capacity with a culture medium; 2-pump, creating a pulsating fluid flow; 3-cultural chamber filled with culture medium; 4-tube scaffold; 5 - direction of flow of the culture medium. Fig. 2. Scanning electron microscopy of the inner surface of a personalized cell-populated vascular prosthesis of small diameter: 1 - magnification ×100; 2 magnification ×500; 2 - magnification ×2000. Fig. 3. Atomic force microscopy of the surface of a personalized small-diameter cell-populated vascular prosthesis. Fig. 4. Viability of human colony-forming endothelial cells on the inner surface of a personalized small-diameter cell-populated vascular prosthesis. Combined staining with nuclear dyes Hoechst 33342 and ethidium bromide, magnification ×200, scale bar - 50 µm: 1 - cell stained with ethidium bromide; the rest of the cells in the image have only Hoechst 33342 staining.

An example of the invention "Method of in vitro manufacturing of a personalized cell-populated vascular prosthesis".

A vascular scaffold was fabricated by electrospinning on a NANON-01A device (MESS, Japan) from a mixture of 10% PCL (Sigma-Aldrich, USA) and 5% PHBV (Sigma-Aldrich, USA) in chloroform under the following conditions: needle for supplying the solution - 22 G, voltage - 20 kV, solution feed rate - 0.3 ml/h, collector - pin with a diameter of 4 mm.

The peripheral blood was taken from the donor in two test tubes: with sodium citrate and with heparin. Fibrinogen precipitate was obtained from citrated blood by ethanol precipitation. For this, the blood was centrifuged for 10 min at 2000 × g. The resulting plasma was cooled to 4°C and added 16% ethanol, also cooled to 4°C, with constant stirring in the ratio of plasma to ethanol solution 1:1. The resulting solution was centrifuged for 20 minutes at 1000 × g and a temperature of 4°C. The fibrinogen precipitate was dissolved in 0.9% NaCl with HEPES to a concentration of 30 mg/ml and aprotinin 100 cfu/ml was added

The tubular PHBV/PCL scaffold was impregnated with the resulting fibrinogen solution by immersion. Then, a thrombin solution at a concentration of 500 U/mL (T7009, Sigma-Aldrich, USA) and CaCl ₂ at a concentration of 40 mmol/L were applied to the surface of the scaffold to polymerize fibrin. After polymerization, the prostheses were immersed in phosphate-buffered saline with ε-aminocaproic acid at a concentration of 2 mg/ml.

CPEK was isolated from heparinized blood. For this, the peripheral blood mononuclear fraction was isolated on a Histopaque 1077 density gradient (Sigma-Aldrich, USA) according to the manufacturer's instructions. The cells obtained from the interphase were washed twice with phosphate-buffered saline followed by centrifugation, the cell suspension was resuspended in a complete nutrient medium EGM-2MV (Lonza, Switzerland) containing 5% fetal bovine serum (HyClone, USA), 2% penicillin/streptomycin (Invitrogen, USA) and 0.25 g/ml amphotericin B (Invitrogen), and populated in collagen-coated (Thermo Fisher, USA) T-25 culture flasks (surface area 25 cm ² ). Cultivation was carried out in CO ₂ -incubator at 37°C and 5% CO ₂ . For the first 2 days, the culture medium was changed daily to remove non-adherent cells and debris, then after 2-3 days. After a week of cultivation, cells were removed with accutase (Sigma-Aldrich, USA), washed with phosphate-buffered saline with 5% fetal bovine serum, resuspended in a complete nutrient medium, transferred to a culture plate coated with fibronectin (Sigma-Aldrich, USA) and cultivated until formation. 70-80% confluent culture, after which the passage was performed. Visual control of culture growth was carried out daily.

The inner surface of the tubular-modified fibrin PHBV / PCL scaffold KFEK settled in an amount of 700,000 cells per 1 cm ² and cultured in complete medium in static conditions for 2 days after which the system was connected to a pulsating flow bioreactor and cultured at a shear stress of 1.27 dyne/cm ² during the day. Then gradually increased the shear stress to a value of 2.85 dynes/cm ² and cultivated in this mode for another 4 days.

Evaluation of the effectiveness of the developed technology.

Evaluation of the shape and spatial orientation of CFEC on the inner surface of a personalized vascular prosthesis.

The inner surface of the personalized vascular prosthesis was assessed by scanning electron microscopy. All prosthesis samples were fixed in 2% glutaraldehyde solution with post-fixation in 1% OsO ₄ solution, after which they were washed with phosphate-buffered saline and dehydrated in a series of alcohols with an increasing concentration from 30% to 100% for 15 minutes each, followed by final drying at room temperature. The samples were mounted on stages for a scanning electron microscope using carbon tape and gold-palladium sputtering was carried out using an EM ACE200 vacuum unit (Leica Microsystems, Germany) to form a conductive coating. The resulting samples were studied on an S-3400N scanning electron microscope (Hitachi, Japan) under high vacuum conditions at an accelerating voltage of 5 kV.

The study showed that fibrin adhered to the surface of the PHBV/PCL scaffold with the formation of a uniform dense coating of a finely porous fibrous structure (Fig. 2). When cultivating CPEK on the surface of a prosthesis modified with fibrin, a high degree of their adhesion to the surface with the formation of a monolayer was noted, as well as good spreading of the cells and their tight fit to the surface.

Evaluation of the strength of cell adhesion and the relief of the inner surface of a personalized vascular prosthesis.

The study of the strength of cell adhesion and the relief of the inner surface of the prosthesis was carried out using atomic force microscopy. The samples were fixed in 4% paraformaldehyde solution and then dried. The study was carried out on an Asylum Research Cypher atomic force microscope in air, under normal conditions and at room temperature, in contact mode.

It was found that when CPEC vascular prostheses are populated under laminar flow conditions, an integral endothelial monolayer is formed (Fig. 3).

Evaluation of the quantity and viability of CPEC on the inner surface of a personalized vascular prosthesis.

Samples of the vascular prosthesis were stained with nuclear dyes Hoechst 33342 (Sigma-Aldrich, USA) in combination with ethidium bromide and analyzed using an inverted fluorescent microscope Axio Observer.Zl (Carl Zeiss, Germany).

The analysis of the results demonstrated the complete preservation of the viability of CPEC on the inner surface of the vascular prostheses (Fig. 4).

Claims (1)

In vitro manufacturing technology of a personalized cell-populated vascular prosthesis, which consists in the formation of a prosthesis framework from a mixture of polyhydroxybutyrate/valerate at a concentration of 3 to 7% (m/v) and c-polycaprolactone at a concentration of 9 to 14% (m/v) by electrospinning, covering the scaffold with autologous fibrin with further colonization of the inner surface with autologous colony-forming endothelial cells isolated from peripheral blood in an amount of at least 700,000 cells per 1 cm ² and preparing the endothelial layer under conditions of a pulsating flow bioreactor for further implantation into the vascular bed.

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