Recent advancements in the bioprinting of vascular grafts

Recent advancements in the bioinks and three-dimensional (3D) bioprinting methods used to fabricate vascular constructs are summarized herein. Critical biomechanical properties required to fabricate an ideal vascular graft are highlighted, as well as various testing methods have been outlined to evaluate the bio-fabricated grafts as per the Food and Drug Administration (FDA) and International Organization for Standardization (ISO) guidelines. Occlusive artery disease and cardiovascular disease are the major causes of death globally. These diseases are caused by the blockage in the arteries, which results in a decreased blood flow to the tissues of major organs in the body, such as the heart. Bypass surgery is often performed using a vascular graft to re-route the blood flow. Autologous grafts represent a gold standard for such bypass surgeries; however, these grafts may be unavailable due to the previous harvesting or possess a poor quality. Synthetic grafts serve well for medium to large-sized vessels, but they fail when used to replace small-diameter vessels, generally smaller than 6 mm. Various tissue engineering approaches have been used to address the urgent need for vascular graft that can withstand hemodynamic blood pressure and has the ability to grow and remodel. Among these approaches, 3D bioprinting offers an attractive solution to construct patient-specific vessel grafts with layered biomimetic structures.


Introduction
Cardiovascular disease (CVD) is the principal cause of death globally, with the annual number of mortalities expected to reach 23.3 million by 2030 [1].Coronary heart disease, peripheral arterial disease, and aortic disease are the common forms of CVD [2].In the UK, more than 180 000 patients are brought to hospitals each year as a result of myocardial infarction [3].The patients suffering from CVD are treated with various vascular repair treatment techniques, including angioplasty, stent placement in the blocked artery, and bypass surgery [4].Annually, the average number of coronary artery bypass grafting (CABG) operations done in the United States and the United Kingdom are ∼400 000 and ∼20 000, respectively [5,6].Currently, autologous grafts, which include saphenous vein, radial artery, or internal mammary artery, are utilized for CABG procedures [7].Although these autologous vessels represent a gold standard, unfortunately, 30% of patients lack a suitable vessel for bypass grafting and 50% of the patients do not have good quality grafts [8,9].Moreover, venous grafts such as a saphenous vein can result in intimal hyperplasia, limiting the patency rates and showing failure rates of around 50% at ten years [10].
The shortage of autologous vessels has led to the use of synthetic (mainly polyethylene terephthalate and expanded polytetrafluoroethylene) grafts.Commercially, Gore-Tex® and Dacron® synthetic vascular grafts are available which work well as replacements for large conduit arteries (such as the abdominal aorta) but suffer a high incidence of thrombosis, stenosis, and infection when the diameter is smaller than 6 mm [11].Further, these grafts are not suitable for pediatric patients due to their inability to grow [12].The tissue engineering approach provides a solution for the biofabrication of small-diameter vascular grafts to meet the patient-specific requirements [13].The various tissue engineering techniques used to construct small-diameter blood vessels can be categorized as scaffold-based processing [14][15][16][17][18], decellularization-based processing [19][20][21], molding techniques [22][23][24][25], and cell-sheet engineering [26,27], as shown in Figure 1.These methods have been useful in preclinical trials, but they possess several disadvantages.Scaffold based processing has a long processing time, poor micro-architecture control, and non-homogeneous distribution of cells during the cell-seeding step [28].Decellularization-based techniques have high chances of immune rejection and generally involve high cost [21].Molding methods perform poorly to fabricate patient-specific sized vascular grafts, whereas cell-sheet engineering methods possess disadvantages such as size scalability, high cost, and long maturation time.Lifeline vascular grafts made by Cytograft® take up to nine months to produce an implantable graft from a fibroblast cell sheet [29].Three-dimensional (3D) bioprinting offers an attractive solution for the direct fabrication of vessel constructs using different cell types [30].
This review focuses on the recent bioinks and bioprinting techniques used for the biofabrication of blood vessels.In the end, the properties of an ideal vascular graft and various testing methods are summarized as per the guidelines of Food and Drug Administration (FDA), the United States, and International Organization for Standardization (ISO) 7198:2016 Cardiovascular implants and extracorporeal systems-Vascular prostheses-Tubular vascular grafts and vascular patches.

Bioinks for vascular graft
Bioink contains bio-printable materials, cells, and other biologics, to fabricate tissues or organs using 3D bioprinting processes [31].These bioinks provide structural integrity and a viable environment for the cells to grow and proliferate into 3D-bioprinted living tissues.An ideal bioink should have favorable biological, rheological, and mechanical properties [32].Bioinks are generally made from natural and synthetic biomaterials [33].Bioinks prepared from natural biomaterials have several advantages over synthetic materials, which include biomimetic structures similar to the extracellular matrix (ECM) of the native tissues, biocompatibility, and self-assembly.In contrast, synthetic biomaterials possess better and tunable mechanical properties compared to natural materials [34].
Stiffness (elasticity) of a cross-linked bioprinted construct is an important property in deciding the cellular activity [35,36].The stiffness of the cross-linked bioprinted structure influence the cell migration and proliferation [37].The cell migration reduces as the stiffness of the hydrogel is increased, whereas the rate of cell proliferation is directly related to the stiffness of the cross-linked construct [38].It has been demonstrated that the elasticity of alginate gel has significant effect on the fate of mesenchymal stem cells (MSCs) [39].Another study showed that endothelial cells (ECs) change their cell stiffness in response to the stiffness of the surface on which they were grown [40].Degradability of biomaterials is also critical, as it affects the mechanical properties of the printed constructs over time [35].The vascular cells can produce enzymes to degrade the matrix to make their migration easy [41].Therefore, it is very important that bioinks must have a range of stiffness and biodegradability similar to the native tissue so that cells can grow and spread naturally.Table 1 shows the summary of bioinks used in the bioprinting of vascular grafts.Agarose is a natural polysaccharide used as a bioink for bioprinting applications.The thermal gelation temperature of agarose is 30 • C-40 • C. It has good mechanical properties and biocompatibility characteristics [42] but lack cell attachment and proliferation [43].Agarose was blended with collagen, alginate, and fibrin to improve the cell adhesion properties [42][43][44][45].Agarose hydrogel was used to print support for cell aggregates to form straight as well as branched vessels [46].In this work, a mold was produced using agarose rods.Cell aggregates were then printed to get a vascular structure.
Alginate is commonly used for the bioprinting of vascular constructs.It has good biocompatibility and can be cross-linked using calcium chloride and calcium sulfate solution [7,47].Alginate bioink was used in the liquid support-based inkjet bioprinting process to fabricate vascular structures with bifurcations [48].The calcium chloride solution provided support to the printed structure and acted as a cross-linking agent for alginate bioink.Alginate bioink has also been printed using a co-axial nozzle printing setup [49,50].The shell and core of the printing nozzle contain alginate bioink and a cross-linking agent, respectively.Thus, hollow vascular channels could be printed directly without any postprinting processes.The ionic gelation process makes alginate bioinks a suitable material for bioprinting.However, low in vivo degradability, poor cell adhesion, and migration are the drawbacks associated with these bioinks.The oxidized alginate bioinks have been developed to address these shortcomings.Along with the advantages of alginate hydrogels, oxidized alginate offers faster degradation rates [51].A composite bioink, made from a combination of alginate di-aldehyde and gelatin (ADA-GEL), was tested for the fabrication of small-diameter vascular grafts using a bioplotter and double-needle system [52].
Fibrin is a protein found in blood and plays an important role in blood clotting [34].Fibrin can be converted into hydrogel after treating it with thrombin.It offers good biocompatibility and degradation characteristics but possesses weak mechanical properties.Fibrin has been used as a bioink to print microchannels using human ECs [53].A tri-layered blood vessel model was also fabricated using a drop-ondemand bioprinting technique [54].In this study, human umbilical vein endothelial cells (HUVECs) were loaded in gelatin and printed as sacrificial material.The next layer containing smooth muscle cells (SMCs) in fibrinogen was printed mimicking tunica media of a natural blood vessel.The stiffness of these printed samples was close to that of a natural blood vessel.Recently, a composite bioink, blending fibrinogen with gelatin was used to make vascular constructs using rotary bioprinting technique [55].The addition of gelatin improved the rheological properties and printability of fibrinogen.
Collagen is one of the most important proteins found in the ECM of mammals, which makes collagen an attractive biomaterial for 3D bioprinting applications.Rat tail tendon, porcine skin, and bovine skin are the main sources of type I collagen [56].Collagen bioinks show excellent cell attachment and proliferation characteristics, but the gelation takes a long time (approximately an hour at 37 • C), which results in poor print shape fidelity and non-homogenous cell distribution [57].However, the drawbacks associated with the long gelation time can be overcome by blending collagen bioinks with other biomaterials.This blending improves both the printability and mechanical properties of bioprinted constructs [58].The photocrosslinking agent, riboflavin, was used in collagen hydrogel, which improved the printability and increased the storage modulus of the bioink [59].Recently, the viscoll collagen solution has been used as a bioink in the extrusion bioprinting process [60].This highly purified collagen solution showed shear-thinning properties with improved shape fidelity and homogeneity.Thus, complex-shaped constructs can be printed with this viscoll collagen solution.
Gelatin is a denaturized form of collagen.Gelatin hydrogels are viscous biomaterials with thermal gelation characteristics along with good biocompatibility and degradability [61].These bioinks have a problem of thermostability, i.e. the hydrogel changes to the solution above 37 • C [62].In order to address this shortcoming, gelatin is either combined with other biomaterials or chemically modified.Gelatin has been blended with alginate [52], hyaluronic acid [61], silk fibroin [63], fibrinogen [55], and chitosan [64] for various bioprinted constructs.Gelatin bioink can be modified by the addition of the methacryloyl group to form a photo cross-linkable hydrogel called gelatin methacryloyl or GelMA [65].GelMA hydrogel possesses a high level of thermostability without compromising the salient features of gelatin, including biocompatibility and degradability [66].Irgacure 2959 or lithium phenyl-2, 4, 6-trimethylbenzoylphosphinate are used as photoinitiator during the synthesis of GelMA hydrogels to cross-link under ultraviolet (UV) light or visible light, respectively.The visible light crosslinking can help to reduce the detrimental effect of UV light on cell viability [67].For extrusion bioprinting applications, the low viscosity of GelMA bioink remains a challenge, which can be improved by carefully controlling the printing parameters such as nozzle temperature, printing pressure, bed temperature, bioink concentration, and cell density [68].Another way of improving the rheological properties and printability of GelMA bioink is to use Gellan gum to make composite bioink [69].Tubular constructs with high aspect ratios were bioprinted using GelMA-Gellan gum composite bioink [37].A layer-by-layer UV curing was done to print taller constructs.In another study, type I collagen was blended in GelMA bioink [70].This addition of collagen improved the rheological properties of the bioink and storage modulus of the bioprinted construct without affecting the characteristic low gelation time of GelMA bioink.Recently, a bi-layered vascular graft was printed using a composite bioink containing GelMA, hyaluronic acid, glycerol, and gelatin [71].High concentration bioink was loaded with HUVECs, whereas low concentration bioink contained SMCs for the outer layer.In order to print a 20 mm long graft, the construct was cured with UV light (365 nm, 200 mW cm −2 ) for 100 s.The tensile testing of the bioprinted graft revealed a J-shaped stress-strain curve, but the strength values were still well below the target values of the native arteries.In another recent study, a novel bioink blend containing GelMA, polyethylene(glycol)diacrylate, alginate, and lyase was synthesized to print small-diameter vessels using the co-axial 3D bioprinting technique.The lyase catalyzed the degradation of alginate in the bioink blend, generating spaces for the growth of vascular smooth muscles cells, and improving the cell proliferation [72].
Natural bioinks provide a favorable environment for the cells to migrate and proliferate but lack suitable mechanical properties.On the other hand, bioinks made from synthetic polymers can easily be tuned to achieve the desired structural strength of the bioprinted constructs.Low cell biocompatibility and lack of cell binding sites are the major disadvantages of synthetic bioinks [73].Pluronic F127 as fugitive ink is employed extensively in the field of vascular tissue engineering [74].Polyethylene glycol (PEG) polymer was used as a bioink in 3D bioprinting because of its high hydrophilicity, good strength, and low immunogenicity [75].Low biodegradability and lack of cell binding sites were the major drawbacks of PEG bioinks [76].PEG was combined with other natural biomaterials to improve the cell attachment characteristics of hydrogels [77].A bioink blend consisting of GelMA, sodium alginate, and 4-arm poly(ethylene glycol)-tetra-acrylate (PEGTA) was also used to print vascular constructs [30].GelMA and sodium alginate in the bioink blend resulted in two stages of cross-linking, whereas PEGTA improved the printability and mechanical strength of vascular structures.In recent work, a gelatin-PEG-tyramine (GPT) bioink was synthesized to construct vascular grafts using a co-axial printing setup [78].Gelatin with HUVECs was used as core sacrificial material, and human dermal fibroblast (HDF) cells in GPT bioink were printed as a shell to make a bi-layered vascular construct.
The commercially available bioinks have shown some good results for various tissue engineering applications.NovoGel® from Organovo was utilized in the bioprinting of aortic vascular graft [79].As the field of bioprinting is advancing rapidly, there are many companies including, Cellink, Regemat3D, and Allevi3D, which have developed bioinks for specific tissue engineering applications [77].

Vascular cells: morphology and their sources
The types and sources of cells used in the vascular constructs influence the structure and in vivo performance of the implanted grafts [5].The blood vessel contains ECs, SMCs, and fibroblast cells in the tunica intima, tunica media, and tunica adventitia, respectively [81].In an ideal vascular graft, ECs should be aligned longitudinally, whereas the SMCs should have the circumferential orientation, similar to the cellular structure found in a natural blood vessel [82].It is essential to consider all three cell types to achieve the required performance of a vascular graft.The ECs in the vascular graft undergo a morphological change and form a functional layer of elongated cells aligned in the direction of blood flow [83].This layer of ECs help resist blood coagulation and control the tone of the vascular construct [84].
In the medial layer, SMCs are oriented in a concentric manner [85].The SMCs control the contraction and dilation of a blood vessel in response to the changes in the blood flow characteristics [86].The cellular functions of SMCs are greatly influenced by cyclic stresses and hydrostatic pressure [87].The repeated stretching also increases the ability of SMCs to synthesize ECM, which improves the mechanical strength of the graft [88].Next to SMCs, fibroblast cells are distributed randomly in the tunica adventitia of a vessel.The fibroblast cells produce collagenrich ECM to prevent the vessel from bursting at high blood pressures [89].
The human umbilical cord is one of the major sources of vascular cells [90].HUVECs and human umbilical vein SMCs (HUVSMCs) can be obtained from umbilical vein and then stored using cryopreservation techniques [90,91].Another important source of cells to biofabricate a vascular construct is autologous vascular cells [5].These cells are taken directly from the patients and cultured for an extended time before using them in a tissue-engineered vascular graft (TEVG) [28].Although these autologous cells have been employed extensively, there are some drawbacks associated with this cell source.The vascular cell extraction techniques are invasive and often result in donor site morbidity [29].The capacity of these autologous cells to proliferate and regenerate is also limited [92].To address these limitations, stems cells have been investigated for both in vitro fabrication methods and in vivo regeneration of vascular tissues [93].Stem cells possess good proliferation characteristics and high differentiation rates along with multiple sources, including stems cells from bone marrow, induced pluripotent stem cells, adipose tissue, and embryonic stem cells [94,95].

3D bioprinting of vascular graft
3D bioprinting is an advanced and emerging technology having great potential in the fields of tissue engineering and regenerative medicine.Bioprinting uses cell-laden biomaterials, generally called bioinks, to deposit in a layer-by-layer fashion [96].The ultimate goal of 3D bioprinting is to offer an alternative to autologous or allogeneic tissue grafts for the replacement or treatment of damaged tissues [97].In 3D bioprinting, the process starts with data acquisition or 3D models of the affected tissue or organ developed with the help of MRI or computer tomography (CT) scans.This data is processed by open-source or bioprinter company's software to create a G-code file for the bioprinter to print the same shape construct using bioinks [98].Extrusion-based bioprinting, inkjetbased bioprinting, stereolithography-based bioprinting, and laser-based bioprinting are the common As the physiological conditions and anatomy of vascular conduits vary from patient to patient, there is a high demand for patient-specific and tailor-made vascular grafts [101,102].3D bioprinting has a great advantage over other conventional methods to fabricate patient-specific vascular constructs having intricate geometries.Furthermore, bifurcated grafts, accurate distribution of cell types, and the capability to produce layered biomimetic vascular tissues are the other major advantages of 3D bioprinting technology [12].

Inkjet-based bioprinting
Inkjet bioprinting, derived from commonly available 2D inkjet printers, is a fast, low in cost, and programmable technique that offers high resolution and accurate cell distribution [103].In the inkjet bioprinting process, a bioink droplet is generated by making use of energy sources.Based on the type of energy source, inkjet bioprinting is divided into four major modalities, which are thermal, piezoelectric, electromagnetic, and acoustic [104].This method is suitable only for low viscosity bioinks, normally in the range of ∼30 cP [105].The thermal, mechanical, or electrical shock to cells during droplet formation and low cell concentration are the major drawbacks of the inkjet bioprinting process [106].In 2005, the inkjet bioprinting process was used to construct tubular structures from hydrogels [107].These early studies showed that the 3D structure could be printed in a bath for rapid cross-linking of hydrogels [108,109].Alginate bioink was used to print tubular constructs in a cross-linking solution containing calcium chloride on a platform assisted 3D inkjet bioprinting system [110,111].The horizontal and zigzag vessels were printed to reveal the capability of inkjet bioprinting technology to produce different shaped vascular grafts.In another study, the horizontal and vertical vascular networks with bifurcations were produced [48].As a layer of alginate bioink is printed, the platform moved in the Z-axis, dipping and cross-linking the printed layer in the solution of calcium chloride.The cell viability of fibroblast cells in the printed structure was reported to be above 90%.
A tri-layered blood vessel was printed using a drop-on-demand printing system [54], as shown in Figure 3.In this work, a custom-built printing setup was utilized to print ECs in the sacrificial gelatin core, SMCs in fibrinogen cross-linked by thrombin, and fibroblast cells in collagen solution.The sacrificial layer was removed after a few hours to ensure proper attachment of ECs on the lumen surface.The cell viability was ∼83%, and the expression of collagen IV was observed.In a recent study, print resolution and fidelity are improved by using an upward bioprinting technique [112].This method used a drop-on-demand bioprinting process with the nozzle at the bottom, as shown in figure 4. A droplet was produced by a digitally actuated mechanical valve under the air pressure of 0.1 MPa.The bioink droplet (GelMA and alginate) moved against the gravitational force, and then it was attached to a substrate.The droplet spreading diameter in upward bioprinting is less than that of the conventional downward bioprinting, thus, improving the shape fidelity of the tubular construct.
Although inkjet bioprinting provides an attractive solution to make straight and bifurcated 3D conduits, the mechanical strength of these constructs still needs a lot of improvement to qualify for the clinical trials.

Extrusion-based bioprinting
Extrusion-based bioprinting is the most widely used bioprinting method in the field of tissue engineering [113].This technique involves the dispensing of a bioink through a nozzle under the action of force created by pneumatic pressure, linear movement of the piston, or rotating screw [114].This method can print with a wide range of bioink viscosities (6-30 mPa•s) and printing resolution from 100 µm to mm [115].The major advantages of extrusion bioprinting over other types of bioprinting are high mechanical integrity of printed constructs, low cost, high cell densities, and ability to print with multiple cell types and bioinks [116][117][118].Whereas, the biggest limitation of extrusion bioprinting is the high level of shear stresses experienced by the cells during the extrusion of bioink through the small diameter nozzle.This reduces cell viability and their functions [119].The extrusion bioprinting technique has been utilized extensively to construct vascular tubular structures.The methods reported in the literature can be categorized into support bath method, coaxial printing method, mandrel-based method, and direct-bioprinting method [115].The support bath bioprinting method, introduced by Hinton et al, is also termed as 'freeform reversible embedding of suspended hydrogels (FRESH)' [120].In this printing method, bioink is extruded into a hydrogel bath to support the printed structure and provide a sacrificial material to create hollow features in the printed construct, such as lumen in blood vessels.The support bath, which behaved like a Bingham plastic, was made from gelatin microparticles.This technique is similar to indirect bioprinting methods for vascular grafts, in which sacrificial material is removed after the printing of tubular construct [46,79].The support bath also provided a hydrated buffered environment for the cells to maintain viability.The right coronary arterial structure with multiple bifurcation and hollow lumen was successfully printed using the FRESH technique, as shown in figure 5.The blood vessels in a thick tissue were printed in a support bath using personalized ECs hydrogel [121].The tri-axial lumen structure within a thick tissue construct revealed the potential of this method in creating bioprinted vascularized tissues.In a recent study, a 3D vascular construct with biofabrication was printed using GelMA-alginate interpenetrating polymer networks (IPNs) in a microgel bath [122].The tensile testing results showed that the ultimate tensile strength of vascular constructs using IPNs is higher than those of GelMA and sodium alginate hydrogels.
The co-axial printing of blood vessels was first introduced by Ozbolat et al [49].The core-shell nozzle arrangement demonstrated the capability of printing hollow vascular constructs using two biomaterials.Ozbolat et al used alginate in the shell and calcium chloride (cross-linking agent) in the core of the co-axial nozzle to print tubular vessels.Another research group printed a vascular structure using a multi-layered co-axial nozzle with a blend bioink containing GelMA, sodium alginate, and 4-arm PEGTA [30].A continuous generation of hollow conduits with various diameters was demonstrated in a single step using the multi-layered co-axial nozzle.A commercially available, open-source 3D printer was modified by replacing the print head with a co-axial nozzle to print hollow alginate structures, cross-linked by calcium chloride [123].A freestanding vascular tube was printed in a single step using HUVECs loaded vascular tissue-derived ECM (VdECM) and alginate bioink in the shell and fugitive pluronic F127 with calcium ions in the core [50], as shown in figure 6(a).The calcium ions cross-linked the alginate bioink, followed by the thermal gelation of collagen found in VdECM.By changing the nozzle sizes and flow rates, the inner diameter and wall thickness of the vessels can be changed.A group demonstrated the use of stereolithography 3D printers to print customized co-axial nozzles for the extrusion of hollow tubular constructs [124].This methodology proved helpful in making custom-designed nozzles to control the critical dimensional features of vascular grafts.In a recent study, GPT bioink was used in the shell and gelatin in the core [78].The tubular structure was printed on a bio-paper surface made from fibrin.Two variants of blood vessels were printed.First, HUVECs were loaded in GPT, and gelatin was used as the sacrificial material, while a bilayer structure was generated by adding HDF cells in GPT and HUVECs in gelatin.A combination of ADA-GEL bioink was also used in the co-axial extrusion bioprinting, with gelatin in the core as sacrificial material [52].Although the co-axial printing technique represents a rapid and direct fabrication of long blood vessels, there are some limitations, including a lack of ability to make bifurcated and patient-specific vessels.
The mandrel-based or rod supported bioprinting method has also been investigated for the biofabrication of vascular grafts.It includes a rotating rod, which acts as a support for the hydrogel.After the hydrogel is printed on the rotating rod, the hollow construct is removed from the rod or mandrel [106].Gao et al printed a blood vessel on the rotating rod, using two co-axial nozzles, one for SMCs and the other for fibroblast cells, as shown in figure 6(b) [125].The endothelialization was done at the end.The construct had both macro and microchannels.The tensile testing of the constructs showed the linear stress-strain behavior, as compared to the J-shaped curve of the natural blood vessels.Freeman et al used gelatin with fibrinogen to print a vascular graft onto a rotating polystyrene rod, which was pre-treated with pluronic F-127 (10% w/v) [55].After 60 days of culture, collagen deposition was observed, and the burst pressure was found to be around 1000 mmHg, which is well below the required target value in natural blood vessels.To improve the mechanical properties gelatin, sodium alginate, and carbon nanotubes hybrid bioink was printed onto a rotating rod [126].The addition of carbon nanotubes improved the mechanical properties at the expense of mild cytotoxicity.A recent study showed the biofabrication of trilayered vascular graft on a rotating rod using alginate bioink and polycaprolactone (PCL) [127].The MSCs in alginate were sandwiched between the inner and outer layers of biodegradable PCL.The innermost PCL layer was cross stripped to make it porous for the easy exchange of materials to the middle layer of cells, whereas the outer layer was helically printed to give strength to the graft and make it leak proof.The mechanical testing was not done.Rather these grafts were implanted in dogs to evaluate their in vivo biocompatibility.These rod-based methods provide a good basis to develop vascular constructs, but this technique cannot be used to print patientspecific graft and bifurcated vessels in a single step.
In order to print a vascular graft from the patient's MRI data, a direct bioprinting method can be used in which bioink is stacked layer by layer to make a 3D structure [108].The shape fidelity is the most important factor in the direct bioprinting methods, and the bioink should have sufficient mechanical strength to withstand and maintain a 3D structure [106].A layer by layer UV assisted extrusion-based bioprinting method was developed to directly print high aspect ratio constructs from GelMA-gellan gum bioink [37].The mechanical testing of these tubular constructs revealed the J-shaped stress-strain curve.Recently, a bi-layered vascular graft is printed using GelMA, hyaluronic acid, glycerol, and gelatin, as shown in figure 7 [71].The inner layer and outer layer of the printed graft contain ECs and SMCs.A similar J-shaped stress-strain curve was obtained, but the strength is well below the target values of natural blood vessels.

Laser-assisted and stereolithography-based bioprinting
In a laser-assisted bioprinting process, a laser impulse is generated to create a small droplet from a bioink coated ribbon, which is directed towards a substrate to form 3D constructs [128].It is a nozzle-free technique capable of printing with a wide range of bioink viscosities with great precision [96].A limited number of studies have been reported on the laser-based bioprinting of vascular constructs.A biological laser printing method was used to print a vascular network with HUVECs and HUVSMCs [91].In a later study, a bifurcated vascular construct was bioprinted in a solution bath using alginate bioink and fibroblast cells [129].This work successfully demonstrated the use of a solution bath for the fabrication of Y-shaped vascular grafts.
Stereolithography bioprinting is an emerging technique in the field of tissue engineering, in which a digital light cures the hydrogel resin layer by layer, and the solid cross-linked constructs are raised from the reservoir [130].High resolution and printing of complex structures are the main advantages of stereolithography [131].A bifurcated vascular construct was printed using acellular, and HUVECs loaded hydrogels using visible light stereolithography [132].The visible light exposure time increased the mechanical stiffness of the printed grafts, but these stiffness values reduced significantly due to degradation after 24 days.Another group used polyacrylate resin to fabricate complex acellular vascular networks using UV irradiation [132].In a later study, a Yshaped cell encapsulated 3D vascular graft was printed using GelMA hydrogel, which is cured by UV light [133].After printing, cell viability was found to be 75%.The UV intensity and time are important parameters for fabrication and medium intensity and exposure time revealed optimal results.The effect of UV exposure on cell viability needs more consideration.A recent study defined two important parameters, which include depth of UV penetration and critical energy level, during the stereolithography of GelMA vascular grafts [134].After 48 h of cell incubation, the cell viability was 80%.
A major limitation of stereolithography in the field of vascular tissue engineering is its inability to produce a layered graft, as found in the natural blood vessels.

Testing of vascular graft
An ideal vascular graft should have matched biomechanical properties with the native blood vessel.
Although long-term patency, biocompatibility, and biodegradability are important traits, a functional layer of ECs and a similarity between the mechanical properties of a vascular graft and native blood vessel are the most critical requirements of an ideal vascular graft [13].The structure of natural blood vessels consists of three types of cells; ECs lined at the lumen surface, SMCs found in the tunica media, and fibroblast cells located in the outermost tunica adventitia layer [81].Among the three cell types of a blood vessel, the layer of ECs provides a smooth, semi-permeable, and thrombus-resistant surface that facilitates the laminar flow of blood [135].Thus, a vascular graft should have a functional layer of ECs so that it can resist thrombosis and control the smooth flow of blood through it.
A vascular graft is a construct to transport blood at hemodynamic pressures; therefore, it must have enough strength to withstand these pressures without rupturing or permanent deformation.Ideally, a vascular graft should have similar mechanical properties as that of a native blood vessel, which requires replacement.Burst pressure, compliance, fatigue resistance, and suture retention are the most important criteria for the selection of a vascular graft [136].Burst pressure is the maximum pressure that a vascular graft can withstand before the development of an acute leakage [137].The maximum circumferential force per unit area, σ, of graft can be related to the burst pressure, P, as, σ = Pd/2t, where d and t are the graft diameter and thickness, respectively.It is obvious that the burst pressure increases with the decreasing diameter.Therefore, the burst pressure of a smallcaliber graft will be higher than the normal range of physiological blood pressures in humans [138].For instance, the burst pressures of a human saphenous vein and internal mammary arteries are 2476 mmHg and 4460 mmHg, respectively [139].
During the pulsatile flow of blood, the walls of a blood vessel undergo deformation.This deformation is always gradual, both during the application and removal of the force.Thus, a natural blood vessel demonstrates viscoelastic behavior under normal physiological conditions [140].In a simplified approach, compliance is an important property to describe the complex mechanical behavior of the artery or vein experiencing hemodynamic pressure.Compliance of a blood vessel with diameter d, Young's modulus E, and thickness t is given as C = d/2Et.This shows that compliance is inversely proportional to wall thickness.The arteries with thick walls are always less compliant than thin-walled veins [141].Compliance of a vascular graft can be measured using ANSI/AAMI/ISO 7198:2016 'Cardiovascular implants-tubular vascular prostheses' (ANSI 7198) [142] as, (1) where P 2 and P 1 are higher and lower pressures, respectively, and R 2 and R 1 are corresponding greater and smaller radius of the vascular graft.Table 2 shows the burst pressure, compliance, and suture retention values of human saphenous vein, radial artery, internal mammary artery, and the coronary artery.The target burst pressure and compliance values of a TEVG must be ∼2031 mmHg and 7.25%/100 mmHg, respectively [28,141].
The other important requirements of an ideal vascular graft include non-toxicity, nonimmunogenicity, and low cost.It must be able to be produced in a wide range of diameters and lengths, as per the patient-specific requirements.It should be kink-resistant and has the ability to be handled and manipulated during the surgical procedures.Lastly, the TEVG should be able to grow and remodel itself after the surgery, especially in pediatric patients [13].Figure 8 represents the requirements of an ideal vascular graft.
Despite the significant research on the biofabrication of vascular grafts, autologous grafts remain the gold standard for coronary artery bypass surgery (CABG) [145].In order to reach the clinical trial stage, the TEVGs must pass through the biological and mechanical testing as defined by the FDA, the United States, and ISO 7198:2016 Cardiovascular implants and extracorporeal systems-Vascular prostheses-Tubular vascular grafts and vascular patches [142,146].According to the FDA guidelines, a vascular graft, either smaller or greater than 6 mm internal diameter, must be tested for thrombosis, embolic events, occlusion, stenosis, leakage, biocompatibility allergic reaction, graft disruption, seroma, and aneurysm.The mechanical performance of a vascular graft is a pre-requisite for further graft assessment.Therefore, the following tests have been performed in the literature to evaluate the TEVGs.

Burst pressure measurements
The burst pressure is measured as given in the ANSI 7198 standard [139].The graft samples with 6 cm in length are cannulated and pressurized with phosphate buffer saline solution.The pressure is increased at 80-120 mmHg s −1 until the graft's rupture.A digital pressure gauge and LabVIEW system can be used to record pressure inside the graft.The maximum pressure recorded is referred to as burst pressure.Another study utilized different techniques for the measurement of burst pressure using water as a pressurizing fluid [147].Figure 9(a) shows the schematics of burst pressure measurement with and without the use of latex sleeve.
Tensile testing methods have also been used for the estimation of burst pressure [55].The test specimen is cut into strips and then subjected to uniaxial tensile testing in the axial and circumferential directions of the vascular graft.The burst pressure was estimated using Barlow's formula given as, where t is the wall thickness of the graft, S is the circumferential ultimate tensile strength of the sample, and D is the outer diameter of the graft.Similar to uniaxial testing, ringlet tests can be also performed for the calculation of burst pressure using Barlow's equation [148], as shown in figure 9(b).

Compliance measurements
Compliance of a vascular graft can be measured using a setup, as shown in figure 9(c).This setup consists of a transparent chamber to house the vascular graft sample in a 9 g L −1 NaCl solution, called physiologic serum.The chamber is placed on a tensile testing machine to apply a constant axial load on the vascular graft during the test.The vascular graft is pressurized internally using another serum line, operated by compressed air.A high-pressure transducer and laser micrometer were used to record the applied internal graft pressure and respective graft diameter.This pressure-diameter data can be processed in a LabVIEW system.The compliance of the vascular graft can be measured using equation (1), as defined in the ANSI 7198 [144].
Similar to burst pressure, compliance has also been estimated using a tensile testing machine [55].The compliance was estimated as [149], where E c is the elastic modulus in the circumferential direction of a vascular graft, D is the outer diameter of the graft, t is the wall thickness of the graft.

Suture retention measurement
Suture retention test can be conducted as specified in ANSI 7198.A 15 mm length graft specimen was cannulated onto a metallic mandrel.A 2 mm bite of 5-0 suture is placed at the end of the graft specimen.The specimen is then pulled at a rate of 120 mm min −1 until failure.The force can be recorded digitally with the help of LabVIEW at a sampling frequency of 5 Hz.Each specimen should be tested at least three times with sutures at 120 • apart [139].

Assessment of biological performance of vascular grafts
The biological performance assessment of bioprinted vascular grafts involves both in vitro and in vivo characterization techniques.These characterization methods are categorized under various risk groups associated with vascular prostheses, as proposed by FDA, USA.A clinically suitable vascular graft must resist thrombosis, embolic events, occlusion, and vascular stenosis [142].Cell penetration and viability studies, immunostaining, cell attachment and proliferation studies, cell tracking techniques, and other visual inspection techniques are often used to determine the biological performance of a vascular [150].In a recent study, the viability of cells was evaluated using a live/dead viability kit [71].Figures 10(a A layer of SMCs along with the elastin and collagen fibers in the vascular graft help to regulate the lumen diameter of the vessel [81].However, migration and abnormal proliferation of vascular SMCs cause vascular stenosis [151].The assessment of graft's patency can be done by studying the response of heparin on cultured SMCs [152,153].The long-term patency of the vascular grafts remains a challenge in the field of vascular tissue engineering.An active layer of ECs at the lumen surface is vital to resist coagulation and thrombosis [154].To evaluate the thrombogenicity of vascular grafts, in vivo assessment techniques were used [155].After six months, the implant's patency was checked using ultrasonography and CT techniques.The biocompatibility of a vascular graft must also be tested before using it to replace blocked native blood vessels.The in vitro assessment can be made by checking the activity of the cells using cell viability assay and mitochondrial metabolic activity test [156].The other in vitro techniques used to check the biocompatibility of scaffolds include cell morphology studies using an SEM and cell distribution studies using a confocal microscope [157].

Conclusions and future outlook
To summarize, we have discussed the properties of an ideal vascular graft along with the recent trends in bioinks and bioprinting methods for the biofabrication of blood vessel-like conduits.The conventional methods to make small diameter tissue-engineered blood vessels face various challenges which include, diameter and compliance mismatch between the native vessel and the graft, inability to grow and remodel, especially in the case of pediatric patients, and a possibility of immune reaction due to in vivo degradation of biomaterials after implantation.3D bioprinting is an attractive technology that has the potential to fabricate patient-specific grafts.Among the different categories of bioprinting methods, the extrusion bioprinting technique is widely explored in the field of vascular tissue engineering because of the superior mechanical properties as compared to other methods.The spatial control of cell types and the ability to print constructs in layers are the other main advantages of 3D bioprinting over other conventional processes.Despite the recent developments, the bioprintedvascular graft has not yet qualified for a clinical trial, mainly due to the challenges related to the bioinks.The bioprinted grafts made from hydrogels are very fragile and have insufficient strength to withstand hemodynamic pressures in vivo.Both natural and synthetic hydrogels have been tested, but there is still a requirement of a bioink that can provide a favorable environment to the cells and possess sufficient mechanical strength.More research is required to be carried out to design synthetic bioinks that can provide a bioactive and cell-friendly environment along with inherent and tailorable mechanical behavior.
Although recent studies show promising results in generating a graft with bifurcation and having similar features as found in the native arteries, many studies lack the biomechanical characterization results as suggested by ISO and FDA, United States.After following these guidelines, a vascular graft can be pushed towards clinical trials and long-term in vivo evaluation.Considering the current advancements, there is a possibility that the vascular surgeons may use handheld bioprinters or bio-pens for the quick reconstruction of damaged vascular tissues, as already seen in the cartilage and bone reconstruction treatment.

Figure 3 .
Figure 3. Printing of tri-layered blood vessel.(a) Schematic cross-section of custom-built bioreactor for vessel printing.(b) Schematic cross-section of the vessel showing layers of endothelial cells and fibroblast cells.(c) Schematic cross-section of the tri-layered vessel.Reprinted by permission from Springer Nature Customer Service Centre GmbH: Springer Nature, Scientific Reports [54] 2018.

Figure 5 .
Figure 5. 3D bioprinting of coronary artery tree using a FRESH technique.(a) A 3D model of a human right coronary arterial tree.(b) 3D bioprinted arterial tree using a FRESH technique in a gelatin support bath.(c) Printed arterial tree in fluorescent alginate in green showing hollow lumen.(d) A zoomed-in image to show vessel wall thickness and lumen.(e) Arterial tree mounted in a perfusion fixture.(f) A time-lapse showing the black dye perfused through the lumen of the arterial tree.Scale bars, 10 mm (b), 2.5 mm (c), 1 mm (d), and 2.5 mm (e) and (f) [120].

Figure 9 .
Figure 9. Mechanical testing of vascular grafts.(a) Schematic diagram of graft burst pressure measurement (i) without and (ii) with latex sleeve.(b) Ringlet mechanical testing of a vessel constructs.(c) Schematic diagram of setup for measuring vascular compliance.Reprinted from [144], Copyright (2016), with permission from Elsevier.
) and (b) show the viability of ECs and SMCs, bioprinted at two different printing pneumatic pressures.In the same study, the immunofluorescent technique was used to check the layer specificity of the bi-layered printed construct, as shown in the figures 10(c)-(e).

Figure 10 .
Figure 10.Analysis of cellular behavior of 3D bioprinted bi-layered vascular graft.(a) Live/dead assay of SMCs and HUVECs bioprinted at 160 kPa and 100 kPa pneumatic pressure.(b) Viability of cells at 160 kPa and 100 kPa.(c) Tracking of labeled cells in a vascular graft at day 7.(d) Hematoxylin and Eosin staining of the bioprinted graft after seven days in culture.(e) Immunofluorescent staining of bioprinted graft after seven days in culture.Scale bar: 200 µm [71].

Table 1 .
Summary of biomaterials used in the bioprinting of vascular graft.

Table 2 .
Mechanical properties of human saphenous vein, radial artery, internal mammary artery, and the coronary artery.