The mechanical properties of vascular grafts should be as close as possible to those of the internal mammary artery, the obvious choice for a natural artery. In patients with severe hypertension, the systolic pressure may be as high as 180 mmHg, while the burst pressure of natural blood vessels exceeds 3,000 mmHg. According to the bioengineering consensus, the graft with burst pressure (>1,000 mmHg) is desirable (L'Heureux et al., 2006; Konig et al., 2009). Concurrently, compliance should match as much as possible to those of the natural blood vessels. The suture retention strength exceeds 1 N, and the ultimate tensile strength is higher than 1 MPa, with the compliance being 10%–20%/100 mmHg (Girerd et al., 1992; L'Heureux et al., 2006; Konig et al., 2009; Stekelenburg et al., 2009; Seifu et al., 2013). A sufficient extracellular matrix can provide appropriate ultimate tensile strength, suture retention strength, and fracture strength for vascular grafts.

To ensure vascular patency, the lumen surface of vascular grafts should have proper blood compatibility to prevent or reduce the adhesion of platelets and the activation of the coagulation system; this is a necessary condition, especially for small-diameter vascular grafts. Likewise, vascular grafts also need excellent cell compatibility on the lumen surface for cell adhesion and proliferation. If there are endothelial cells—from autologous veins or circulating progenitor cells—on the lumen surface, their presence reduces the probability of thrombosis. Tissue engineering can promote early endothelialization of the vascular wall to thus improve the biocompatibility of vascular grafts. Moreover, vascular grafts need to be able to be remodeled as part of the body after implantation. Only when smooth muscle cells and endothelial cells proliferate, the structure and function of vascular grafts can gradually approach the natural blood vessels.

The extracellular matrix of vascular grafts needs to be human-derived to minimize the risk of inflammation, foreign body reaction, and immune recognition. If the cells are not autogenous and have not been completely removed, the recipient’s immune system recognizes antigens from the donor’s graft surface after implantation, triggering an autoimmune response. Vascular grafts can be divided into allografts and xenografts depending on the cell source.

The antigens that cause allograft rejection are histocompatibility antigens. Among them, the histocompatibility antigen that can cause strong rejection is called major histocompatibility antigen. The histocompatibility antigen that causes weak rejection is called secondary histocompatibility antigen (mH antigen). The key to successful transplantation depends on whether the tissue capacitive antigens between the donor and recipient are consistent or similar. After transplantation, T cells of the recipient recognize MHC antigen on the surface of the graft through direct and indirect ways to cause the immune response. Human leukocyte antigen (HLA), encoded by the major histocompatibility complex (MHC), is the most important antigen of allograft rejection (Hamada et al., 2021). Especially the antigen molecules located in the HLA-DR, followed by HLA-A, HLA-B, HLA-DQ, and HLA-DP (Hubscher et al., 1990; Fleischhauer et al., 2001; Lim et al., 2016; Wiebe et al., 2017; Féray et al., 2021). There was no significant relationship between HLA-C and transplant rejection (Piccinni et al., 2021).

The biggest problem caused by xenotransplantation is immune rejection, especially hyperacute rejection. Hyperacute rejection can kill cells from the donor vascular graft within 24 h (Tector et al., 2020). The reason is that natural antibodies in human blood bind to antigens on the surface of the xenograft to activate the complement system. When activated, the complement system forms membrane-attacking complexes, triggering a series of chain reactions including endothelial cell dysfunction, platelet aggregation and thrombosis, leading to transplantation failure. Xenograft immunogenicity is directly related to the presence of several antigens including the α-Gal epitope, the linked N-glycolyl neuraminic sialic acid (Neu5Gc) and the Sd(a) (Byrne et al., 2011; Naso et al., 2013; Gao et al., 2017). Mammals have these antigenic molecules except for humans, apes and certain monkeys.

Consequently, the decellularized vascular matrix can only be used as a vascular graft in the clinical application if it meets all the above-discussed criteria. In general, these factors should also consider market regulation, large-scale production, and other requirements. So far, this goal has not been achieved and is still in progress.

Linked/adhesive proteins refers to a series of proteins with adhesiveness and connectivity. They are usually spherical, responsible for connecting the fibrous protein structure in the decellularized vascular matrix with cells.

Integrin is one of the most important globular proteins that serve multiple functions. For example, as a “bond,” integrin connects the cytoskeleton to the vascular matrix. This connection not only stabilizes cells, but also allows vascular cells to feel the changes of external mechanical information (such as pressure), thus regulating the biological function of cells (Sun et al., 2016). In addition, integrin, as a receptor, can realize the information transmission between vascular matrix and cells (Singh et al., 2010).

Fibronectin is widely distributed and has the ability to combine with a variety of vascular matrix components. The multiple structures of fibronectin enable it to simultaneously combine with cell surface receptors (such as integrin), collagen, proteoglycans, and other adhesion molecules (Nelson and Bissell, 2006; Barkan et al., 2010). This property also enables it to mediate the assembly of a variety of extracellular matrix proteins. In short, fibronectin thus shapes the matrix structure and plays a connecting role (Avila Rodriguez et al., 2018).

Laminin is the main component of basement membrane. The α, β, and γ peptide chains are interlinked to form a cross structure. It is closely connected with basal cells to fix them. Laminin affects adhesion, differentiation, migration and phenotype stability, and provides resistance to apoptosis of related cells (Domogatskaya et al., 2012).

In addition to the components mentioned above, there are proteoglycans, glycoproteins, amino polysaccharides, proteases, bioactive molecules, electrolytes, and water in invisible gelatinous structures of the decellularized vascular matrix.

Proteoglycan is a huge molecule with a protein chain as the core and glycosaminoglycans (GAGs) as the side branches. Because GAGs can attract water, it forms a gelatinous structure to fill the vascular matrix (Scott and Panitch, 2013). This gel maintains the stability of microenvironment, provides nutrition for cells and transmit signals (Koyama et al., 1998).

Hyaluronic acid is a huge single chain molecule only composed of disaccharides. It can combine with many proteoglycans to form a huge proteoglycan polymer and work together with elastin and collagen to resist damage to tissues and body caused by external impact (Li et al., 2020; Li et al., 2021).

None of the components are isolated. Collagen and elastin provide mechanical strength and flexibility to the decellularized vascular matrix. Linked/adhesive proteins participate in mechanical information transmission, act as receptors to transmit chemical information and regulate the activity of biological molecules. Invisible gelatinous structures can transmit signals, store important bioactive substances, and endow blood vessels with flexibility and buffering capacity against external forces. All of these components play indispensable roles in creating the most suitable environment for the survival of cells and the body. Thus, the decellularized vascular matrix retains the complex extracellular three-dimensional structure and biological activity, much like the structure and function of the native vessels. Undoubtedly, it is more in line with the standards of TEVGs and has broad application prospects.

Early protocols regarding the preparation of decellularized vascular matrices were mainly based on the decellularization of blood vessels obtained directly from animals or humans to obtain decellularized vascular matrices.

Besides the decellularization of obtained vessels, a more popular attempt is to prepare decellularized vascular matrix by tissue engineering.

In the 1970s, Charles Sparks made one of the early attempts to prepare decellularized vessels through tissue engineering to obtain the “Sparks Mandril” structure of artificial blood vessels (Sparks, 1969). The process involved subcutaneous implantation of Mandril of 5.1 mm diameter in the legs of patients, followed by covering with loosely woven synthetic polymer Dacron. A tube composed of fibrous tissue was formed outside Mandril. However, this method was abandoned because of the formation of a thrombus or aneurysm post-implantation. Weinberg and Bell published a report in 1986 announcing the in vitro production of the first tissue-engineered vascular graft (TEVG) (Weinberg and Bell, 1986). They cast collagen gel around the mandril supported using a polyester mesh sleeve and implanted fibroblasts into the outer layer. The tube wall was filled with smooth muscle cells, and the inner wall comprised endothelial cells. The artificial blood vessel could not be successfully implanted in vivo because of the low rupture strength and the inability to conduct in vivo tests without a polyester stent; however, this approach represents important conceptual progress. It encourages the co-culturing of other cells and endothelial cells to prepare TEVGs further.

The first case of a completely biological TEVG was proposed by L'Heureux et al. (1998). It comprised concentric layers of cell sheets, which rolled to form a tubular structure and matured in the bioreactor finally (Figure 3). First, fibroblasts and smooth muscle cells were co-cultured into monolayer cells for 30 days. The fused fibroblasts were separated from the matrix, rolled around the Mandril to form tubular structures, and then dehydrated as the intima of blood vessels. Multiple layers of smooth muscle cells were wrapped on the intima as the media layer, and the lumen was inserted into a silicone tube and placed into a bioreactor for periodic expansion. To be exact, the bioreactor was filled with culture medium and equipped with a pulsatile pump instead of heart. Pulsatile radial stress was applied to smooth muscle cells at certain beats per minute. These conditions not only provided mechanical support, but also mimicked the pulsation of blood in vessels. After maturing in the bioreactor for a week, smooth muscle cells appeared as elongated cells with circumferential or longitudinal orientations. Finally, fibroblasts were wrapped around smooth muscle cells to form an outer membrane, and endothelial cells were planted on the lumen surface. The whole process took nearly 3 months. Compared with the bursting pressure of the saphenous vein, the bursting pressure of this vascular graft was more than 2,500 mmHg, with better mechanical properties. The vascular graft was implanted in 10 patients with end-stage renal disease as a hemodialysis pathway. Three months post-implantation, the patency rate was 60% (McAllister et al., 2009; L'Heureux et al., 2007).

In 1999, (Niklason et al., 1999), were the first to describe the latest preparation process of tissue-engineered arteries. PGA scaffold, which degraded rapidly, was wrapped around the silicone tube and placed in the bioreactor. In the bioreactor, smooth muscle cells were planted on the PGA scaffold, and mechanical stimulation of vascular pulsation and stretching was simulated. After 8 weeks of culturing in the bioreactor, the matrix secreted by smooth muscle cells was deposited on the scaffold, and PGA was degraded. The generated tissues were decellularized to generate decellularized vascular matrix. This technology is used to produce human acellular vessels (HAVs) (Figure 4). It is also being tested for hemodialysis access and vascular reconstruction following trauma.

Bioreactors for the production of TEVGs can be designed for different functions and types in accordance with research purposes (Supplementary Figure S1). The main body of the bioreactor is a glass reservoir fitted with a silicone stopper. The vessels are cannulated with glass pipettes, secured with silk sutures, and placed into the bioreactor. Silicone tubing is attached to the glass pipettes, reservoir inlet, and outlet to complete the perfusion loop. The flow rate, intramural pressure, and wall shear stress can be tuned effectively by adjusting various conditions, for instance, pump speed, tubing resistance, media viscosity, and so on.

Since these groundbreaking studies, significant breakthroughs have been made in the research on vascular grafts using tissue engineering techniques. Some of these have shown promising results in animal models; nonetheless, supporting clinical research data is still required.

Different decellularization techniques often provide different decellularization results. Post-decellularization, a small amount of residual cellular structure remains, making it difficult to obtain decellularized vessels comprising entirely of extracellular matrix. Hence, a set of reasonable and sound evaluation criteria are needed to judge the effectiveness of decellularization protocols. DNA quantification and qualitative evaluation of cell removal are common methods (Naso and Gandaglia, 2022). Particularly, Gilbert et al. (2006), Crapo et al. (2011), and Keane and Badylak (2014) have conducted numerous trials in different animals and tissues and have proposed the following evaluation criteria for the effectiveness of decellularization: <50 ng/double-stranded (ds) DNA/mg extracellular matrix dry weight; <200 bp DNA fragmented length; Lack of visible nuclear materials after staining either with 40,6-diamidino-2-Phenylindole or hematoxylin and eosin. However, the effect of cell residues on host response is not only caused by DNA, but may be the result of the interaction of multiple structures. Therefore, we summarize a set of common but comprehensive evaluation criteria and test methods for the decellularization effect (Table 2) (Naso and Gandaglia, 2022).

The advantage of physical decellularization is the relative simplicity of the procedure. Nevertheless, if the goal is to minimize the destruction of the extracellular matrix, physical decellularization cannot be used alone and needs to be used in combination with other decellularization protocols to achieve good results (Table 3).

The chemical decellularization techniques include a variety of reagents and is the main choice among current decellularization approaches (Table 4). Nonetheless, different reagents bring about different decellularization effects. For example, mild reagents are less destructive to the extracellular matrix, while the effects of powerful reagents are the opposite. Accordingly, the reagents should be carefully chosen after combining the structure and properties of the material to be decellularized.

Biological decellularization techniques are currently less diverse and seem to be less developed, and are less effective when used alone (Table 5). Thus, decellularization through a biological approach is usually performed in combination with chemical decellularization reagents to achieve satisfactory results similar to those obtained by physical decellularization techniques.

Liu and Zhang (2016) observed that the polyelectrolyte multilayers constructed with heparin and SDF-1α bind firmly to the surface of decellularized blood vessels in rats. Besides good hemocompatibility and anticoagulation, these polyelectrolyte multilayers could significantly enhance the migration ability of rat bone marrow MSCs and promote the recellularization of vascular grafts.

Huang et al. (2008) subjected decellularized vascular matrix to vacuum thermal cross-linking treatment; the process improved the mechanical strength and biological stability of the decellularized vascular matrix attributed to inter- and intramolecular dehydration of collagen and the formation of urethane bonds, concurrently maintaining a certain degree of hydrophilicity and porosity.

Xia et al. (2009) treated decellularized rabbit abdominal aorta with plasma and observed a significant reduction in contact angle, inhibition of platelet adhesion, and increased hydrophilicity of the decellularized vascular matrix.

The most commonly used chemical cross-linking agents are glutaraldehyde (GA) and 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC)/n-hydroxysuccinimide (NHS).

GA is the most commonly used cross-linking modification reagent in collagen-based tissue engineering materials due to its low cost and high efficiency. Compared to uncross-linking decellularized vascular matrices, vascular grafts with GA cross-linking show greater improvements in mechanical properties and immunogenicity. However, with the increase in the cross-linking time, the cytotoxic effects of chemical cross-linking agents become apparent and are more likely to cause tissue calcification (Braile et al., 2011; Sinha et al., 2012).

EDC cross-linking decellularized vascular matrices have better biocompatibility and higher cell differentiation potential compared to those obtained by GA treatment. These vascular matrices obtained after EDC treatment are catalytic; thus, EDC/NHS is also gradually and widely used for the chemical cross-linking of collagen biomaterials. Among them, NHS is used to stabilize the crosslinking effect of carbodiimide (Olde Damink et al., 1996).

Heparin is used to prevent thrombosis because it inhibits the thrombin coagulation cascade reaction. (Conklin et al., 2002) could inhibit thrombus formation by bridging heparin on decellularized treated porcine carotid arteries.

Zhang et al. (2010) used the sheep forelimb artery for acellular treatment, applied the light crosslinking method to fix the CD34 antibody on the inner surface of the acellular vessel, and then implanted it into the rabbit femoral artery. This experiment showed that the CD34 antibody-modified decellularized vascular scaffold was superior to the unmodified decellularized vascular scaffold in terms of early anticoagulation and patency rate.

Fu et al. (2021) introduced Silk Fibroin into decellularized vascular scaffolds to improve the physical properties of the scaffolds and also assessed the performance of the scaffolds. They found that Silk Fibroin effectively improved the overall mechanical properties of the stent and reduced the stent degradation rate.

With advances in biology, chemistry, and materials sciences, the improvement in performance and long-term patency by applying various modifications in the decellularized vascular matrix has positive implications, while the effects have yet to be validated in clinical trials.

Cytograft Tissue Engineering Inc., was one of the early companies to develop artificial vessels and The Lifeline™ graft as a typical product. Early research into the use of cell sheet engineering to reconstruct artificial vessels gave rise to the idea of developing “human textiles” (L'Heureux et al., 2006). Small biopsy skin samples of skin and superficial veins were obtained, and endothelial cells and fibroblasts were extracted. Fibroblast sheets were produced in as little as 6 weeks. The obtained TEVGs consisted of three components: a living adventitia, a decellularized internal membrane and an endothelium. This procedure cost upwards of $15,000 and required 6–9 months to produce patient-specific vascular grafts. However, it is unlikely that this vascular graft can be applied clinically owing to high production costs and long waiting times. Despite all this, its method of preparation provided a new idea for research on artificial vascular grafts.

Humacyte, Inc., was founded in 2004 by Laura E. Niklason of Yale University. The firm aims to develop innovative technologies for regenerative vessels and focuses on a new type of vascular graft known as HAVs (Dahl et al., 2011). HAVs are completely decellularized and mechanically robust with circumferentially aligned extracellular matrix without any chemical cross-linking or synthetic components (Niklason and Lawson, 2020). They have been favorably repopulated by host cells in multiple advanced-stage clinical trials in vascular trauma repair, arteriovenous access for hemodialysis, and peripheral arterial disease. In eight clinical trials to date, HAVs have been implanted in more than 500 patients. The firm is currently conducting three Phase Ⅲ clinical trials, one Phase Ⅱ clinical trial, and three Phase Ⅰ clinical trials worldwide (Kirkton et al., 2019; Lauria et al., 2022) (Figure 6).

Vascudyne, Inc., was founded by Jeff Franco and Kem Schankereli. It licensed its proprietary TRUE™ Tissue technology, developed in 2017 by Professor Robert Tranquillo and his colleagues from the University of Minnesota (Supplementary Figure S2). The current products are TRUE™ Tissue, including TRUE™ Graft, TRUE™ Valve, and TRUE™ Patch (Syedain et al., 2021). The biological and mechanical properties of TRUE™ Graft are very similar to those of native tissue; preclinical studies have shown that TRUE™ Graft regenerate and grow with the patient. In July 202, Vascudyne, Inc., announced the first successful in-human use of its TRUE™ Vascular Graft in patients with the end-stage renal disease requiring hemodialysis access.

Artegraft^® approved by FDA in 1970, is a bovine carotid vascular graft. Its decellularized matrix is treated for use as hemodialysis access and lower extremity bypass. The biological fibrous matrix of this graft has been processed to enhance long-term patency and provides a tightly woven, cross-linking conduit that is flexible and compliant. During the past 50 years, Artegraft^® has been implanted in more than 5,00,000 patients.

In October 2020, Hancock Jaffe Laboratories, Inc., (HJLI) announced the first use of grafts in CABG. The first-in-human CoreoGraft study of HJLI was completed, and the patient was discharged from the hospital. In May 2022, Medical 21, Inc., announced that they have been developing an artificial graft called MAVERICS, a small-diameter flexible tube encased in a nickel-titanium alloy stent that eliminates the need to harvest blood vessels from the patient’s legs, arms, and chest. The goal is to improve the quality of life of patients by reducing pain, shortening surgery and recovery time, and reducing the risk of infection and complications.

As already stated, these four representative corporations have been working towards the commercialization process of the decellularized vascular matrix as vascular grafts. Vascular grafts are class III implantable biological medical devices in the International Organization for Standardization (ISO), the US Food and Drug Administration (FDA) and EU classification standards. ISO is an independent non-governmental organization that has created thousands of international standards for many industries, including medical equipment. ISO standards are voluntary, consensus-based documents that provide guidance on specific aspects of technology and manufacturing. For medical device manufacturers, ISO standards are not only important for building high-quality medical devices, but also for maintaining compliance with regulatory requirements. Most of the ISO standards have been recognized by regulatory agencies such as FDA, or have been harmonized with regulations in other regions of the world (such as the EU). Even if ISO standards do not have legal force, they are also basic guidelines for medical devices and in vitro diagnostic equipment companies. We list the ISO Standard aspect to which commercial vascular grafts are subject before they can be put on the market (Supplementary Table S1). In addition, commercial decellularized vascular grafts need to be accord with an international standard publishes in 2021 (ASTM International, 2021). The commercialization process of vascular grafts is tortuous and arduous; in terms of clinical application, it still has a long way to go. For example, the research on product development is not mature enough, and the clinical trials have not been completed yet. On the whole, the commercialization process of vascular grafts is still improving and advancing. Through the active efforts of researchers across the world, the clinical needs of vascular grafts are expected to be met in the near future.