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11Cellular Therapy and Transplantation (CTT). Vol. 12, No. 2, 2023 doi: 10.18620/ctt-1866-8836-2023-12-2-40-50 Submitted: 23 March 2023, accepted: 26 May 2023
Double layer tissue-engineered vascular graft of small diameter based on electrospun polylactide and fluoropolymer microfibers
Pavel V. Popryadukhin Guriy I. Popov Galina Yu. Yukina Elena G. Sukhorukova Elena M. Ivan'kova 2, Valery N. Vavilov 1
1 Pavlov University, St. Petersburg, Russia
2 Institute of Macromolecular Compounds, Russian Academy of Sciences, St. Petersburg, Russia
Dr. Pavel V. Popryadukhin, Institute of Macromolecular E-mail: [email protected]
Compounds, Russian Academy of Sciences, 31 Bolshoy Ave, 199004, St. Petersburg, Russia
Citation: Popryadukhin PV, Popov GI, Yukina GY, et al. Double layer tissue-engineered vascular graft of small diameter based on electrospun polylactide and fluoropolymer microfibers. Cell Ther Transplant 2023; 12(2): 40-50.
Summary
Double layer non-woven tubular grafts 1.1 mm in diameter were prepared by electrospinning. The inner layer of the graft was made of bioresorbable poly(L-lactide) polymer; the outer layer consisted of a non-resorbable fluoropolymer, which served to reinforce the graft wall and to prevent appearance of aneurysm upon complete resorption of the inner layer. The grafts were implanted into rat abdominal aorta for different periods of time (from 1 hrs to 24 months); they showed no toxicity, demonstrated good biocompatibility and high athrom-bogenicity. Total graft permeability was equal to 96%. Morphological analysis of the samples demonstrated that two processes occurred simultaneously in the inner layer of the graft: bioresorption of poly(L-lactide) fibers and formation of connective tissue. Free volume of the outer layer was filled with connective tissue; no signs
of bioresorption were revealed. An intermediate layer consisting of connective tissue was formed between the two polymer layers; thickness of this interlayer increased gradually with time. The tissues formed on the surface of graft walls included mainly fibroblasts, CD68+ cells, a-actin-containing cells and multinucleated foreign body giant cells. It was shown that the non-resorbable outer layer reliably prevented appearance of aneurysms at all the studied time points.
Keywords
Tissue engineering, vascular graft, small-diameter blood vessels, bioengineering, poly(L-lactide), fluoropolymer, electrospinning, microfibers, bioresorption, aneurysm.
Introduction
According to the estimates of World Health Organization, high incidence of cardio-vascular diseases and high mortality rates caused by these disorders are still registered worldwide [1]. The common surgical methods aimed for treatment of these diseases include bypass grafting, stenting, and prosthetic repair of arteries affected by atherosclerosis. During bypass surgery, vascular shunting is performed by means of native vessels (autologous veins or arteries), or synthetic
polymer prostheses (grafts) in order to get around the occluded or thrombosed vessel. A stenting procedure includes insertion of a metal frame into the damaged vessel, thus extending vascular lumen and providing free blood flow. During prosthetic repair, a portion of the damaged vessel is substituted with a graft made of synthetic polymers, or native materials [2].
When autologous vessels are used for bypass grafting or prosthetic treatment, the implant usually becomes integrated
Abbreviations: TEVG, tissue engineering vascular graft; PLLA, poly(L-lactide); FP, fluoropolymer; MFBGC, multinucleated foreign body giant cells.
into the living body. Since these implants contain functional endothelial layer and cause only minimal inflammation, they are successfully used for bypassing the low-diameter vessels. However, the amount of autologous material is limited, especially when several arteries require intervention, or in cases of repeated surgery. Autologous arteries are rarely used, because of considerably disturbed local blood supply when taking these vessels for surgery. Therefore, autologous veins are mainly used; their absence is relatively easily compensated by other veins, and, therefore, no pronounced tissue damage occurs.
However, thickness of a vein wall is significantly lower than that in arteries, and degenerative processes are observed in the walls of autologous veins within a prolonged period of time after the operation, with potential development of aneurysms [3, 4]. Moreover, taking autologous material is an additional trauma to the patient, thus complicating the entire surgical procedure.
On the contrary, stenting is a low-traumatic intervention. A metal stent in its collapsed form attached to the outside of a balloon catheter is threaded through the damaged part of an artery. Then the balloon is inflated, causing expansion of the stent followed by deflation of the balloon and its removal from the organism, with a metal stent remaining within the vessel. This surgery has its disadvantages. First, there is a risk of restenosis due to neointimal hyperplasia caused by the contact between stent and vessel wall. Secondly, repeated stenting (if required at the same site) is difficult, since the metallic frame remains in the organism of a patient lifelong. Recently, the novel stents have been developed, consisting of bioresorbable materials (polylactide and magnesium alloys). However, their surgical parameters are still inferior to those of common non-bioresorbable alloy stents. Thus, these innovative stents are yet not widely applied [5, 6]. Moreover, the stenting procedures become difficult in the cases of extensive atherosclerotic damage, at the sites with complex configuration (e.g., arterial bifurcation), or in the cases of complete luminal closure with atherosclerotic plaque. Stenting is also counter-indicated when the plaque is unstable, i.e, the atherosclerotic plaque cap becomes damaged and may be ruptured when inflating the stent. Both thrombosis and embolism may appear at the site of damage, followed by their migration in bloodstream, causing arterial and venous thrombosis. Their risk is especially high during surgical intervention in coronary and cerebral arteries.
Prosthetics and bypass surgery using synthetic polymeric grafts are conventional and conservative procedures. Advantages of synthetic grafts include good mechanical characteristics, wide range of sizes and shaping, like as commercial availability. They can be used for prosthetic repair of prolonged areas of a vessel, or for simultaneous treatment of several vessels; they are also suitable in the case of complete vascular occlusion. Synthetic grafts are also employed when the stenting is not possible, due to high risk of atherosclerotic plaque rupture. However, currently used grafts are made of polytetrafluoroethylene, lavsan, and their combinations, thus being unable for biological resorption in the body. No protective endothelial layer is formed on their surface, and, moreover, intimal hyperplasia is frequently observed in the area of anastomosis being a pre-requisite for thrombosis,
especially inside the narrow vessel grafts, where blood flows relatively slowly. The grafts under 5 mm in diameter are not used, due to development of early thromboses [7, 8]. When the synthetic grafts are implanted in children, complex repeated surgery is required, e.g., replacement of small grafts for larger ones [9].
One possible solution for these issues may be provided by the tissue-engineered vascular grafts (TEVG), which involves three main components: bioresorbable scaffold, cell material, mechanical and biological signaling [10]. To date, five main techniques for TEVG preparation have been developed: (1) use of bioresorbable polymeric scaffolds (grafts) [11-16]; (2) bioprinting [17, 18]; (3) layer-by-layer tissue engineering [19, 20]; (4) use of decellularized vessels [21]; and (5) use of granulation tissue [22]. Despite a variety of existing approaches, there is still no solution complying with all requirements for the vascular grafts. In particular, the need for usage cell materials at the in vitro preparation stage has not been proven. Some authors deliberately omit this stage due to technological complexity and low reproducibility [14, 23-25].
Therefore, we have developed the technique for preparation of polymeric bioresorbable TEVG with low diameters by electrospinning of microfibers from poly(L-lactide) (PLLA) solution, followed by their partial crystallization on the collecting electrode [14, 26]. The grafts prepared in such a way are biocompatible, non-toxic, showing high athrombogenic-ity (TEVG permeability exceeded 90%), and possess ability for complete bioresorption within 16 months, involving gradual replacement of polymeric fibers with native tissues [14]. In addition, these grafts demonstrate high porosity, which facilitates cell growth; the cells fill the whole intermediate volume of the graft in a short time. Since the pores have small diameters, bleeding does not occur during implantation and in the early postoperative period. However, we observed aneurysms of various sizes upon complete bioresorption of the grafts. Occurence of these defects is caused by formation of structures with low mechanical strength (as compared with that of a native vessel), instead of resorbed grafts. Aneurysm is a life-threatening condition, since the dilated vessel may rupture, thus leading to uncontrolled bleeding and death of a patient.
Therefore, the aims of the present work included development of a two-layer polymeric TEVG with low diameter consisting of a layer of bioresorbable PLLA microfibers reinforced with a layer of non-resorbable fluoropolymer micro-fibers, and the results of in vivo observations of these grafts.
Materials and methods
Preparation of grafts
Poly(L-lactide) (PLLA) Purasorb PL-10 (Corbion Purac, Netherlands) was used in preparation of porous tubular grafts. Electrospinning was performed in the following manner. PLLA was dissolved in trichloromethane (chloroform, Sigma-Aldrich, USA); concentration of the solution was 15%. Using an injection pump, the prepared solution was fed through a metallic tubular electrode into electric field (E = 1.5-104-4.0-105 V/m, the distance between electrodes
15 cm). Microfibers were precipitated on a grounded metallic cylindrical electrode 1.1 mm in diameter; rotation rate was 1500 rpm. The produced grafts fixed on the cylindrical electrode were subjected to thermal treatment at 70°C for 10 min, resulting into partial crystallization of PLLA, thus leading to considerable enhancement of its mechanical and operating parameters. Upon thermal treatment, the thickness of graft wall was equal to 200 pm. The graft preparation procedure developed by our group is described in detail elsewhere [14, 26]. The PLLA grafts (still located on the electrode) were then placed into electrospinning setup again, in order to apply the layer of fluoropolymer (FP) microfibers on their outer surface. We used Fluoropolymer F-32LV (poly(1-chloro-1,2,2-trifluoroethylene-1,1-difluo-roethylene), (-CFCl-CF2-)n[-CF2-CH2-]m) produced by AO "GaloPolymer" (Russia). Our preliminary studies demonstrated that this FP could be easily dissolved, possesses good mechanical parameters and very high athrombogenicity. The grafts prepared of this material remain permeable in 98% of cases. FP was dissolved in ethyl acetate (Sigma-Aldrich, USA), and 15% solution was obtained. Electrospinning parameters were similar to those described above for PLLA. A 50-pm thick layer of microfibers was deposited, and the resulting two-layer graft was taken off the collecting electrode.
Electron microscopy studies of the samples were performed using a Supra 55VP scanning electron microscope (Carl Zeiss, Germany) in the secondary electron imaging mode. Before SEM study, a thin platinum layer was sprayed onto the sample surface. The images were taken at 1 hour, 2 days, 1, 2, 4, 12, 24, 48, 56, 64, 72, 80, and 96 weeks after implantation.
Assessment of mechanical properties
Mechanical characteristics of the grafts were determined by means of an Instron 5943 universal testing machine (Instron, UK) in the uniaxial tension mode; the extension rate was 10 mm/min. Young's modulus, tensile strength and tensile strain were measured for PLLA and PLLA-FP grafts (internal diameter 1.1 mm, wall thickness 250 pm), and native rat aorta. PLLA grafts were also partially crystallized on the collecting electrode. The sample base length was 20 mm in all cases.
Contact angle determination
The experiments were performed using a DSA 30 setup (Kruss, Germany) on the surface of four types of samples: non-porous PLLA, FP films obtained by pouring polymer solution onto glass support followed by drying, porous PLLA, and non-woven FP films prepared by electrospinning. Compositions of solutions and electrospinning parameters are given above ('Preparation of grafts').
Studies of barrier properties of grafts
Two types of grafts (PLLA and two-layer PLLA-FP samples) were tested, with following sizes: inner diameter, 1.1 mm; length, 30 mm; wall thickness, 250 pm. A special setup was designed for these measurements: an NE-1000 programmable single syringe pump (a roller pump for blood perfusion) (New Era Pump Systems, Inc., USA) equipped with a 20 mL syringe was connected with a graft via tubular adapter; another end of the graft was connected via adapter with
a transparent tube with an inner diameter of 4 mm and a length of 2 m. This tube was placed in vertical position, and a scale was set against the tube in order to measure the height of a liquid column. Two types of liquids were used in the studies: (i) water colored with a green dye to improve visualization and (ii) rat blood supplemented with sodium citrate (3.95% concentration) as anticoagulant. A liquid was fed through the studied graft at a constant rate into the tube; then it rose through the tube. When the first drops of a liquid appeared on the outer surface of a graft, the experiment was stopped, and the height of the liquid column was measured. The volume feed rate was 10 mL/min, thus corresponding to linear rate of liquid in the graft lumen equal to 0.18 m/s. This value is typical of blood flow rate in arteries 1.5-2 mm in diameter that are included in human systemic circulation [27]. Thus, in these experiments, the value of hydrodynamic pressure remained constant, while hydrostatic pressure increased gradually. Each measurement was made in five repeats.
Experiments with animals
The in vivo experiments involved 52 male white Wistar rats (age: 3 months, weight: 200-250 g); 4 animals were used in each series of experiments. The surgical manipulations were performed under general anesthesia [Zoletil 100 dissolved in 20 mL of physiological solution (0.1 mL) and Rometar (20 mg/mL, 0.0125 mL of solution per 0.1 kg of animal weight), intraperitoneally, once]. Y-shaped incision for laparotomy was made; microvascular surgery was used to mobilize in-frarenal portion of the abdominal aorta and to insert a prosthetic graft; 8 sutures were put at each anastomosis using at-raumatic needles with Prolen 9-0 threads. In all experiments, no significant bleeding through graft wall or along the lines of anastomosis was observed after restoration of blood flow. No anticoagulants or disaggregant drugs were used. Vascular permeability was estimated according to the classical method [28]. Then the front abdominal wall was sutured in layers using atraumatic needles with Prolen 9-0 threads. Upon suturing, the rats were caged individually, had free access to water and were fed a standard diet. Color and temperature of skin of hind extremities of the animals were monitored; their physical activity was estimated.
Compliance with ethical standards
The animal experiments were carried out in accordance with the regulations concerning use of laboratory animals (principles of European Convention (Strasbourg, 1986) and the Declaration of Helsinki developed by the World Medical Association concerning humane treatment of animals (1996)), and State Standard 33216-2014 ("Guide to keeping and care of laboratory animals. Regulations for keeping and care of laboratory rodents and rabbits").
Morphological studies
At 2 days, 1, 2, 4, 12, 24, 48, 56, 64, 72, 80 and 96 weeks, TEVG with fragments of native aorta were excised and fixed in 10% neutral solution of formalin in phosphate buffer (pH=7.4) for, at least, 24 hrs. Then the samples were dehydrated using a series of ethanol solutions at increasing concentrations, and enclosed in paraffin blocks according to the standard histological technique. The paraffin slices (5 pm thick) were obtained with the use of an Accu-Cut SRT 200 microtome (Sakura, Japan) and stained with Mayer hema-
toxylin and eosin (BioVitrum, Russia). The connective tissue elements were visualized according to the Mallory and Mas-son technique (BioVitrum, Russia).
For immunohisto chemical detection of macrophages and multinucleated foreign body giant cells (MFBGC), mouse primary monoclonal antibodies [Anti-CD68 antibody (ab 31630), Abcam, UK] was used (dilution 1:1000, 20°C, exposure time: 1 h). To reveal bound primary antibodies, multimeric biotin-free detection system was used (D&A, Reveal-Biotin-Free Polyvalent DAB, Spring Bioscience Corporation, USA). The preparations were additionally stained with Mayer hematoxylin (BioVitrum, Russia). In order to detect actin-containing cells (smooth muscle cells and mi-ofibroblasts) after standard procedure of deparaffinization, the slices were treated with mouse monoclonal smooth muscle a-actin antibodies (clone 1A4, dilution 1:2000) (AbCam, UK) for 10 min at room temperature. A MACH2 Mouse kit (Biocare Medical, USA) was used as secondary reagent. To visualize the product of immunohistochemical reaction, the preparations were treated with 3',3'-diamino-benzidine (DAB+, Dako, Denmark). Microscopic analysis of the TEVG-containing preparations was performed using a Nikon Eclipse Ni light microscope (Nikon, Japan) with a 10x ocular, and 4, 10, 20, and 40x objectives. Digital images were recorded with a Nikon DS-Ri2 camera (Nikon, Japan). In all series of experiments, thickness of connective tissue interlayer between the two TEVG layers was measured sequentially ten times along the longitudinal histological section, and the average values were deduced. Statistical treatment of the obtained data was performed using the standard software package (Statistica 7.0, Stat.Soft for Windows). The arithmetic mean value and its standard deviation (M±SD) were calculated; significance of differences was estimated using the Wilcoxon criterion and the Mann-Whitney U test. The significance of differences was determined at P < 0.05.
Results
Fig. 1 presents SEM images of a two-layer PLLA-FP graft. The inner PLLA layer consists of microfibers with round cross-sections, 3-5 ^m in diameter, at the pore size between fibers varying from 5 to 40 ^m. This layer is significantly thicker than the outer layer; it contacts blood directly and undergoes bioresorption. The outer FP layer consists of mi-crofibers with dumbbell-shaped cross-sections 2 to 10 ^m wide, 1-3 ^m high, with pore sizes ranging from 3 to 30 ^m. This layer is significantly thinner, being not susceptible to bioresorption. Its main purpose is to enhance mechanical characteristics of TEVG after complete bioresorption of the inner layer, and, thus, to prevent formation of aneurysms. The amount of non-resorbable polymer retained in the body should be minimal. Therefore, this layer is relatively thin, but it should have a sufficient thickness to perform its mechanical function. Both layers possess high porosity and pore sizes suitable for migration of cells into graft wall. After implantation, the migrating cells fill free volume in the pores between fibers of polymeric graft thus creating a natural vascular graft without a need for preliminary cell repopulation of the grafts.
In order to prevent formation of aneurysms and their rupture, the initial graft should have good mechanical characteristics. In this work, comparative analysis of mechanical properties
A
Figure 1. Scanning electron microscopy (SEM) images of the PLLA-FP graft. A and B, cross-section; C, internal surface; D, outer surface
of PLLA-FP, PLLA grafts and native rat aorta was performed. It was demonstrated that mechanical strength and Young's modulus of PLLA-FP and PLLA grafts were considerably higher than those of the native aorta. The grafts also demonstrated higher elasticity (tensile strain) as seen from Table 1. Rupture of PLLA-FP grafts proceeds in two stages. At the first stage, the internal layer of PLLA undergoes destruction, while the FP layer retains integrity and continues to stretch until the elongation value reaches 273±28%. This property of the graft provides additional protection from destruction and bleeding.
The obtained grafts are highly porous products. Therefore, one of the most important characteristics of a graft is permeability of its wall for various liquids, mainly their blood permeability. Our studies of barrier properties involving water and stabilized rat blood showed that the two-layer PLLA-FP graft possessed considerably better barrier properties than the monolayer PLLA graft (Table 2). This is due to the two-stage impregnation of PLLA-FP graft with a liquid: the inner PLLA layer is impregnated followed by considerable rise in hydrostatic pressure, thus causing saturation of the outer FP layer (Fig. 2). Moreover, both layers consist of highly hydro-phobic substances. The non-woven materials prepared by electrospinning have higher hydrophobicity (due to their peculiar surface relief) than non-porous films of the same
materials (Table 3). After implantation of PLLA-FP grafts into rat aorta, visual observation revealed that the internal PLLA layer was soaked with blood, whereas the outer FP layer was not impregnated; this observation confirms good barrier properties of the bilayer graft.
One hour after implantation of the graft into rat aorta, no macroscopic thromboses were revealed, and blood pulsation above and distally from the implantation site was clearly seen. Scanning electron microscopy studies showed that the whole internal surface of the graft was coated with a thin fibrin layer 3-5 pm thick. The fibrin film was coated with a 10-15 pm thick layer of erythrocytes. Fibrin covered the major part of graft surface, without predominant localization. Fibrin was also found within the graft wall; it filled the pores and formed bridges between polymeric fibers (Supplementary file, Fig. 3 a, b). The amount of fibrin was higher in the pores located near the inner graft surface. Virtually no fibrin was found near the outer graft surface. The appearance of fibrin plugs in the pores inside the graft after starting blood flow led to enhancement of barrier properties and, as a consequence, decreased risk of bleeding.
Two days after surgery, during explantation of grafts, no visual signs of active inflammatory response were revealed in abdominal cavity and in retroperitoneal space. The grafts
Table 1. Mechanical properties of vascular grafts and rat native aorta
Sample Wall thickness, pm Strength, MPa Young's modulus, MPa Elongation, %
Native aorta 150 2.27±0.56 17±5 139±32
PLLA graft 250 3.63±0.41 145±11 175±18
PLLA-FP graft* 250 4.19±0.34 184±23 131±15
* Note: The parameters were obtained before starting the destruction of PLLA layer
Table 2. Barrier properties of PLLA-FP and PLLA grafts
Critical pressure, mm Hg PLLA graft PLLA-FP graft
Water 16.91±2.67 40.45±2.82
Blood 25.75±2.91 63.25±2.57
Table 3. Water contact angles of the materials based on PLLA and FP
Sample PLLA FP
Film 79°±1.7° 80°±0.6°
Non-woven material 118°±2.3° 122°±1.4°
Table 4. Morphometric analysis of the Time-dependent changes of thickness of connective tissue interlayer between PLLA and FP layers of the graft (n=4)
Time post grafting, weeks 1 2 4 12 24 48 56 64 72 80 96
Thickness, ^m 2.9±1.1 6.0±3.2* 12.2±3.3 18.6±8.2 20.6±7.9** 25.1±8.7 20.0±6.3 24.0±11.8 23.3±6.3 26.7±8.1 30.2±6.9
Note: * , the parameters significantly differ from those observed after 1 week of experiment, p <0.05 (p = 0.0001207); **, the parameters significantly differ from those after 2 weeks of experiment, p <0.05 (p = 0.0000001)
were not adherent to the surrounding tissues. Neither their external appearance nor manually determined mechanical characteristics were changed. SEM and histological studies showed that the inner surfaces of grafts were covered with a thin fibrin layer (like as samples from the previous series), but no erythrocyte layer was revealed (Fig. 3 c, d), which indicates the involvement of blood anticoagulation system. Isolated leukocytes and erythrocytes were found between PLLA and FP fibers; the cells migrated into this space due to impregnation of the graft with blood. CD68+ cells (macrophages, MFBGC) and a-actin-containing cells were not observed.
One week after the surgery, no visual signs of active inflammation were observed; a thin connective tissue capsule was formed on the outer surface of the graft. Histological and SEM studies revealed a homogeneous fibrin layer on the inner graft surface. In the anastomosis area, gradual transition of aortal intima and media to inner graft surface was observed; endothelium and smooth myocytes were grown in this area. The PLLA layer was homogeneously populated with small amount of CD68+ cells. The FP layer was densely populated with CD68+ cells; at the periphery of the outer surface, single phagocytizing MFBGC were detected. In both graft layers, small amounts of a-actin-containing cells were distributed homogeneously. At one week of experimental observation, a thin connective tissue inter layer 2.9±1.1-pm thick appeared between PLLA and FP layers of the graft (Table 4). Moderate amounts of CD 68+ cells were present in the formed connective tissue capsule (neoadventitia).
No visual signs of inflammation were observed at histolog-ical section within 2 weeks after implantation. The grafts were surrounded with connective tissue that virtually did
not grow into the graft walls. Histological analysis of aortal anastomosis area revealed the presence of endothelium and smooth myocytes; which have grown within intraluminal surface of the scaffold and formed the neointimal structures. In the PLLA layer, moderate amounts of CD68+ cells and a-actin-containing cells were found. On the contrary, the FP layer was populated with a large amount of CD 68+ cells, including numerous MFBGC located at peripheral areas; no a-actin-containing cells were revealed. The connective tissue interlayer between graft walls became significantly thicker (6.0±3.2 pm) than the layer formed at 1st week (Table 4). Neoadventitia consisted of young collagen fibers, fibroblasts and CD 68+ cells.
Four weeks after surgery, the connective tissue capsule around the graft was seen more clearly; no signs of inflammation were observed. The capsule was adhered to the surrounding tissues, but remained movable, being penetrated with blood vessels. Histological analysis showed the presence of completely formed neointima over the whole internal surface of the graft. Neointima consisted of endothelial cells, subendothelial layer and smooth muscle cells. A moderate amount of CD 68+ cells still remained in the PLLA layer, while the amount of a-actin-containing cells decreased. Like as by 2 weeks after operation, the FP layer was populated with numerous CD 68+ cells and MFBGC without a-ac-tin-containing cells. Within both graft walls, collagen fibers appeared between the polymeric microfibers. The fibrils were synthesized by fibroblasts that migrated from the ad-ventitial side into the graft. The connective tissue structures continued to grow between PLLA and FP layers, and their thickness reached 12.2±3.3 pm (Table 4). The neoadventi-tial structures were similar to those observed at earlier terms (i.e., 2 weeks after surgery).
Figure 2. Studies of barrier properties of PLLA-FP grafts. A, original view; B, impregnation of the PLLA layer with colored water; C, impregnation of two layers of the graft with colored water; D, impregnation of two layers of the graft with blood
■ f, M Wài
Fig. 3. Scanning electron microscopy (SEM) images of PLLA-FP graft taken at 1 hour (A, B) and 2 days (C, D) after implantation. A and C, internal surface; B and D, graft wall (PLLA layer, longitudinal section)
By 12 weeks (3 months) after implantation, no visual signs of inflammation were found, whereas the connective tissue capsule surrounding the graft became more densely connected with native tissues than in 4 weeks. The capsule had a smooth shiny surface and was penetrated with numerous small blood vessels. Histological pattern observed in this period of time was virtually similar to that revealed in 4 weeks after operation. The thickness of connective tissue interlayer between PLLA and FP layers was 18.6±8.2 pm (Table 4).
In 24 weeks (6 months) after implantation, we, generally, observed a similar tissue pattern (Fig. 4 a, b). However, it should be noted that neointima contained small calcifications. Thickness of the connective tissue interlayer between two graft walls was 20.6±7.9 pm (Table 4). In the neoadven-titia, a continuous cluster of MFBGCs was formed which closely fitted to the outer wall of the FP layer.
In 48 weeks (12 months) after implantation, the histologi-cal picture did not undergo any significant changes and remained virtually similar to the patterns observed in 12 and 24 weeks after implantation (Fig. 4 c, d). However, clear signs of bioresorption of PLLA microfibers were observed (fragmentation and formation of pores in the fibers, thus resembling a spongy structure). No signs of fluoropolymer (FP) biodegradation were revealed, which was the expected purpose of our work. One should note that bioresorption rate of PLLA microfibers incorporated in the double-layer PLLA-FP graft was significantly lower than that of the monolayer PLLA graft [14]. In the anastomosis area, a smooth transition of intima from aorta to the graft surface was observed. No signs of hyperplasia were revealed (similarly to the previous two series, i.e., 12 and 24 weeks). Thickness of the connective tissue interlayer between PLLA and FP parts was 25.1±8.7 pm. The clusters of MFBGCs still retained close to the outer FP wall.
Histological analyses performed within the period from 56 weeks (14 months) to 96 weeks (2 years) of the experiment showed that neointima was formed by continuous endothe-lium layer, smooth muscle cells grown from native aortal segments, and a thin layer of collagen fibers were observed (Fig. 5). It should be noted that calcifications were formed in the anastomosis zones in several animals, however, without signs of neointimal hyperplasia. Bioresorption of polymeric structures continued in the PLLA layer. Meanwhile, the amount of collagen fibers increased, and they filled the space previously occupied by PLLA microfibers. Moderate amounts of CD 68+ cells populated the entire layer. The numbers of a-actin-containing cells decreased, according to visual estimates. They were revealed at negligible amounts, predominantly in the outer part of the PLLA layer. As at earlier terms, there were no signs of bioresorption in the FP layer. This portion of graft was completely occupied with CD 68+ cells and surrounded with the MFBGC deposits; no a-actin-containing cells were revealed. Thickness of the connective tissue streak between the graft layers gradually increased up to 30.2±6.9 pm. Neoadventitia was represented by loose connective tissue, containing fibroblasts and CD68+ cells.
Figure 4. SEM images and histological sections of PLLA-FP grafts taken 6 (A, B) and 12 months (C, D) after the in vivo implantation.
A, the graft wall, cross section; B, longitudinal section; C, im-munohistochemistry (a-actin stained brown); D, staining by Mallory method (collagen stained blue). Ob. 10x.
Figure 5. SEM images and histological section of PLLA-FP grafts taken in 24 months after implantation.
A, C - graft wall, cross section, B, D - longitudinal section. (c) Immunohisto chemical detection (a-actin), (d) staining by Masson method (collagen stained green). Ob. 10x.
Discussion
In the current work, we report the results of experimental testing of an original two-layer PLLA-FP low-diameter vascular grafts. Their mechanical and barrier properties were investigated, hydrophobic characteristics were determined, and in vivo experiments were performed during long period of time (from 1 h to 2 years). The mechanical characteristics of the grafts were shown to be superior to those of the native rat aorta. This result is very important for development of vascular grafts (especially bioresorbable and partially bioresorbable products). Biological resorption is followed by renewal of native vascular tissues with low mechanical strength, which may be too thin to endure blood pressure. As a result, vascular aneurysms may appear. Spontaneous ruptures of such aneurysms may cause internal bleeding and threaten the patient's life. The inner PLLA layer of implanted graft is bioresorbable; it degrades slowly being replaced with native tissues. The outer non-bioresorbable FP layer imparts mechanical strength to the newly formed vessel walls. No aneurysms were found at any time during the whole observation period, thus supporting validity of the chosen approach (Fig. 6). However, the PLLA layer did not disappear completely even 2 years after graft implantation. The polymer fibers were partially disrupted, but were able to reinforce graft walls to a certain degree.
The barrier properties of two-layer PLLA-FP graft were considerably better than those of the monolayer PLLA graft. These characteristics are especially important for highly porous grafts (e.g., prepared by electrospinning). Bleeding through graft wall may lead to significant blood loss and, therefore, presenting a threat to patient's life. Therefore, usage of two-layer PLLA-FP graft it is preferable to the mono-layer PLLA prosthesis.
During the in vivo experiments, we studied thromboresis-tance properties of the grafts (absence of thrombi and degree of permeability) as well as biocompatibility, rate and mechanism of living tissue ingrowth into graft wall, and bioresorption of PLLA fibers were investigated. Total permeability of grafts was 96%, a very high value for the low-diameter vascular grafts (5 mm and smaller). However, we should consider implantation of the graft into abdominal aorta, where blood flow rate is higher than that in peripheral vessels, and, thus, a risk of thrombosis is lower.
High biocompatibility between the polymer grafts and living tissues was observed during the whole experiment. No pronounced inflammation (macroscopic or microscopic) was revealed in the PLLA layer (only moderate or insignificant amounts of CD68+ cells were present). In the FP layer, the reaction was more pronounced, and the layer was populated with CD68+ cells, and the MFBGC clustering was observed around its wall, being a typical in vivo response to the foreign body invasion.
Immediately after starting blood flow in the graft, its porous walls became impregnated with blood (mainly, its liquid component). Thin fibrin depositions were observed inside the graft pores and its surface, as clearly seen within the first week post-implant. One week after grafting, smooth growth of intima from aorta to the inner side of the graft was observed in the anastomosis areas. During this period, thin connective tissue interlayer was developed between PLLA and FP layers of the graft. Thickness of this layer gradually increased until the end of observation period. This interlayer improves mechanical characteristics of the graft and its barrier properties. However, it prevents cell migration and, thus, potentially inhibits resorption of the PLLA layer.
Figure 6. Macroscopic views of PLLA-FP graft extracted 24 months after implantation. A, after explantation; B, complete clamping with a surgical instrument; C, graft reshaping after declamping
In four weeks, neointima was completely formed over the whole inner surface of the graft. It consisted of endothelial cells, subendothelial layer, and smooth muscle cells. The presence of neointima at this early stage provided thrombo-resistance of the graft. Moreover, no signs of neointimal hyperplasia were found throughout the experiment, thus being a very favorable prognostic sign, which also suggests high biocompatibility of the graft.
Ingrowth of native tissues into the graft started since the first week of the experiment. Across the whole width of the PLLA layer, virtually in all series and terms of experiments, we observed moderate amounts of CD68+ and a-actin-containing cells (presumably, myofibroblasts, in absence of smooth myocytes). The amounts of fibroblasts and collagen fibers gradually increased with time, while the fibers filled the space between polymer fibers and the volume that was earlier occupied with bioresorbable PLLA fibers. The FP layer was re-populated with numerous CD 68+ cells, starting from the 2nd week of the experiment.
Conclusion
The original vascular graft based on a double-layer structure (degradable and undegradable polymeric components) has demonstrated high biocompatibility, thromboresistance, mechanical strength, bioresorbability (in PLLA layer) during the 2-year experimental observation, and, thus, may be recommended for further studies and potential clinical use.
Conflict of interests
Authors do not declare any conflicts of interests.
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Резюме
Методом электроформования были получены двуслойные нетканые трубчатые протезы диаметром 1.1 мм. Внутренний слой протезов состоял из биоре-зорбируемого полимера поли(Е-лактида), наружный из нерезорбируемого фторполимера, который служил для укрепления стенки протеза и предотвра-шения развития аневризм, после полной резорбции внутреннего слоя. Протезы были имплантированы в брюшную часть аорты крысам на срок от 1 часа до 24 мес и продемонстрировали высокую биосовместимость, нетоксичность и выраженные атромбогенные свойства. Общая проходимость протезов составила 96%. Морфометрический анализ показал, что во внутреннем слое протеза происходят два параллельных процесса: биорезорбция волокон поли(Е-лактида) и образование соединительной ткани. Свободный объем наружного слоя заполнен соединительной тканью, признаков его биорезорбции не выявлено.
Между двумя слоями полимера образуется соединительнотканная прослойка, толщина которой постепенно увеличивается. Клеточный состав стенок протеза представлен преимущественно фибробластами, CD 68+ клетками, а-актин содержащими клетками и многоядерными клетками инородных тел. Было показано, что наружный нерезорбируемый слой надежно предотвращает появление аневризм на всех изученных сроках эксперимента.
Ключевые слова
Тканевая инженерия, сосудистый протез, сосуды малого диаметра, биоинженерия, поли(Е-лактид), фторполимер, электроспиннинг, микроволокна, биорезорбция, аневризма.
Двухслойный тканеинженерный сосудистый протез малого диаметра на основе полилактидных и фторполимерных микроволокон, полученных методом электроформования
Павел В. Попрядухин Гурий И. Попов Галина Ю. Юкина Елена Г. Сухорукова Елена М. Иванькова 2,
Валерий Н. Вавилов 1
1 Первый Санкт-Петербургский государственный медицинский университет им. акад. И. П. Павлова, Санкт-Петербург,
Россия
2 Институт высокомолекулярных соединений РАН, Санкт-Петербург, Россия
