Abstract

Vascular tissue engineering has the potential to make a significant impact on the treatment of a wide variety of medical conditions, including providing in vitro generated vascularized tissue and organ constructs for transplantation. Since the first report on the construction of a biological blood vessel, significant research and technological advances have led to the generation of clinically relevant large and small diameter tissue engineered vascular grafts (TEVGs). However, developing a biocompatible blood-contacting surface is still a major challenge. Researchers are using biomimicry to generate functional vascular grafts and vascular networks. A multi-disciplinary approach is being used that includes biomaterials, cells, pro-angiogenic factors and microfabrication technologies. Techniques to achieve spatiotemporal control of vascularization include use of topographical engineering and controlled-release of growth/pro-angiogenic factors. Use of decellularized natural scaffolds has gained popularity for engineering complex vascularized organs for potential clinical use. Pre-vascularization of constructs prior to implantation has also been shown to enhance its anastomosis after implantation. Host-implant anastomosis is a phenomenon that is still not fully understood. However, it will be a critical factor in determining the in vivo success of a TEVGs or bioengineered organ. Many clinical studies have been conducted using TEVGs, but vascularized tissue/organ constructs are still in the research & development stage. In addition to technical challenges, there are commercialization and regulatory challenges that need to be addressed. In this review we examine recent advances in the field of vascular tissue engineering, with a focus on technology trends, challenges and potential clinical applications.

Introduction

Vascularization is vital during development and wound healing, where formation of hierarchical networks enables efficient transfer of oxygen, nutrients, immune cells and other biological factors to the tissues for maintaining growth and homeostasis. Ischemic diseases, including atherosclerotic cardiovascular disease (CVD) remain one of the leading causes of mortality and morbidity across the world [1,2]. Treatments for such diseases have resulted in persistent demand for vascular conduits to reconstruct or bypass vascular occlusions and aneurysms. Also, the demand for functional bioengineered tissues and organs for addressing shortage of transplant tissue and organs will increase in future. However, a major bottleneck in this approach still is the vascularization of these bioengineered constructs. Without an extensive and functional vasculature, these thick bioengineered constructs will face the risk of ischemia even before implantation into a host. In order to acquire an adequate flux of oxygen and nutrients, cells in a tissue have to be in close proximity to a capillary. In fact, highly metabolic cells such as the cardiomyocytes must be within 100–200 µm of a capillary [3,4]. Since the presence of a developed vascular network would have significant effects on the success of an engineered tissue, one of the major current goals of tissue engineering is the vascularization of three-dimensional (3D) tissue constructs to mimic the networks of host tissue.

The regenerative potential of the human body after injury or disease is variable across different organs and tissues. In any healing scenario, there is a need for a well-formed local microvasculature so that cells and proteins can be transported to the site of injury for initiating and supporting the wound-healing response [5]. Furthermore, a hierarchical vascular bed is necessary to efficiently deliver blood to tissues, which necessitates a balance of capillaries and larger arteriolar vessels. For example, the human heart has a dense vessel network of 2000–3000 capillaries/mm2 that supports the terminally differentiated, non-proliferative cardiomyocytes that mediate contractile function [6]. Major coronary events such as a myocardial infarction (MI) can cause localized cell death if the tissue is not quickly supplied with adequate oxygen and nutrients [7]. Therefore, the ultimate aim of cardiac tissue engineering is to regenerate the contractility of the heart that is lost following MI. However, currently implantable cardiac constructs have ∼7-fold lower microvascular densities than that of the native heart [8]. Hence, development of well vascularized tissues and organs (including the heart) is a top medical necessity.

The field of tissue engineering has turned toward biomimicry to solve the problem of tissue oxygenation, nutrient exchange and waste exchange through the development of vasculature. To engineer blood vessels or vascular networks, it is vital to understand its basic biology and physiology. The formation of vascular networks in the body begins early in embryonic development and dynamic changes to the network occur throughout life in response to both stimulatory and inhibitory signals. Neovascularization (formation of new blood vessels) happens through two distinct processes, vasculogenesis [9] and angiogenesis [10] (Figure 1A). The process of angiogenesis is more commonly utilized in tissue engineering applications. There is another process called arteriogenesis, which involves the compensatory remodeling of existing collateral blood vessels into functional arterial vessels in response to physical stimuli such as flow-related stresses [11]. This effectively increases vessel diameter and wall thickness through endothelial cell (EC) proliferation, migration, and recruitment of supportive cell types for further autocrine and paracrine signaling. This remodeling process is an important event in vascular biology and will also be critical for vascular tissue engineering. For engineering vascular tissues, it is important to take the structural components of a blood vessel into consideration (Figure 1B). Understanding these elements is required for rational biomaterial design and choosing an appropriate cell source. Blood vessels can markedly vary in size within the vasculature hierarchy. The arteries and veins (large vessels) are arranged to ensure efficient transport over long distances while the capillaries (smaller vessels) are organized for optimal exchange of oxygen, nutrients and waste products. Therefore, it is no surprise that design requirements and approaches for engineering large vessels and smaller vessels (<1 mm in diameter) are quite different. Many of the different blood vessel beds also share some common structural features.

Figure 1

The process of vasculogenesis and angiogenesis (A) [12]. The structure of blood vessels (B)

© 2015 Pearson Education, Inc.

Figure 1

The process of vasculogenesis and angiogenesis (A) [12]. The structure of blood vessels (B)

© 2015 Pearson Education, Inc.

Vascular tissue engineering has now evolved to incorporate these structural variations, which in turn offers the right functionality to the engineered vasculature. In addition to functional outcomes, the integration of TEVCs with the host tissue without mounting immunologic rejection is of primary concern in vascular tissue engineering. A suitable cell source is also critical to help impart structural stability and facilitate in vivo integration. Use of patient-derived autologous cells has garnered interest because of their potential to minimize graft rejection. However, isolating and expanding viable primary cells to a therapeutically relevant scale has its own challenges. With the advancement of stem cell (SC) technology and gene editing tools such as CRISPR (clustered regularly interspaced short palindromic repeats) [13], use of autologous adult stem cells and induced pluripotent stem cells (iPSCs) are emerging as promising alternative sources of ECs and perivascular smooth muscle cells (SMCs) for engineered vasculature [14,15]. Biomaterials, cells and chemical signals together can provide appropriate matrix support and pro-angiogenic signals to form functional blood vessels and vascular networks. Presenting angiogenic signals in a spatiotemporal fashion can also help generate better vascular networks, which will allow for the creating large and dense engineered tissues. Recent studies have also attempted to introduce more sophisticated uses of physical and chemical cues to selectively induce angiogenesis in engineered tissues. These include in vitro pre-vascularization and encouraging penetration of host vasculature into implants via in situ tissue engineering methods [16].

In this review, we provide an overview of critical topics in engineering vascular tissues, with a focus on the technology trends related to the material, cells and fabrication methods. Recent developments in tissue engineered vascular grafts (TEVGs) and vascularization of bioengineered tissues are presented, along with other important topics such as importance of anti-thrombogenicity, the nature of blood contact surfaces and understanding host-implant anastomosis. Vascular tissue engineering is one area of regenerative medicine that has advanced early into clinical applications and this aspect is also discussed here. Finally, current challenges in engineering vascular tissues are presented along with potential solutions and future perspective about this exciting field.

Challenges in engineering vascular tissues

With several decades of exploratory studies and development, the field of vascular tissue engineering is advancing quickly and has also reached clinical evaluation. However, it also faces many challenges. Some of the major challenges include finding best combinations of materials and cell sources for making TEVGs, engineering blood vessels with varying lumen diameters, incorporating vascular networks in thick bioengineered tissues, ensuring appropriate remodeling responses in TEVGs, adequate burst pressure and compliance in engineered vascular tissues. Additional challenges include non-thrombogenicity (both short-term and long-term) and saturability of TEVGs to the host vasculature. One of the fundamental hurdles that faces the development of bioengineered tissues and organs for transplantation is supplying microvasculature to thick tissues (>100–200 μm) [17]. Hence, producing a perfusable microvascular network will be critical not only for the engineering of transplantable organs, but also for the treatment of ischemic disease [18].

Cell source is an important component in fabricating TEVGs and also tissue or organ substitutes. Some challenges here include quantity of harvestable primary cells, effect of culture conditions on functionality of derived cells, potential to become patient-specific cells (in case of stem cells or iPSCs) and controlling the remodeling response during the time needed to fabricate TEVGs or tissue substitutes in vitro. All cell sources have their own limitations. Primary mature vascular cells were the first cell sources used and significantly contributed to the establishment of the vascular tissue engineering paradigm. However, the quantity and quality of isolated vascular cells are far from what is needed for clinical use. Additionally, in case of CVD, most patients who require vascular bypass and replacement surgery lack proper blood vessel segments that can be used for cell harvesting.

Adult stem cells are another source and can be used immediately after isolation. However, cell viability would depend on patient’s health condition and age. The use of embryonic stem cells (ESCs) is increasing in vascular tissue engineering studies [19]. For clinical use, there are also ethical concerns and questions on their tumorigenic potential. Use of iPSCs has also widely increased for a variety of regenerative medicine applications [20]. They have the advantages of unlimited cell quantity, vascular lineage-specific differentiation, patient-specific cell sourcing and even the ability to correct genetic mutations before use. However, use of iPSCs must be investigated further for use in vascular tissue engineering and also for concerns similar to those for ESCs. Developing a biocompatible blood-contacting surface is still a major challenge for both small and large diameter TEVGs. Understanding and replicating the blood-contacting surface and anti-thrombogenic mechanisms will be critical to the clinical success of TEVGs. Due to the complexity of the structure and composition of blood vessel inner walls, mimicking the blood-contacting surface is still a challenge. Use of specially engineered biomaterials, cells and biological molecules (including pro-angiogenic factors) might help address this challenge.

One challenge in generating TEVGs has been how to distributed cells within the graft so as to recapitulate their natural arrangements. In blood vessels, the SMCs and ECs are resident in medial intimal space. Static (or passive) method of seeding cells onto engineered vascular grafts can only achieve a sub-optimal distribution. Use of biological molecules such as collagen, fibronectin and cell adhesion peptides (RGDC, heparin-binding sites of fibronectin, etc.) on scaffold surfaces have also helped improve the adherence of seeded cells [21,22]. However, from a vascular engineering point of view, these may not be enough. An alternative to static seeding of cells on vascular grafts is using dynamic seeding. One type of dynamic seeding is the rotational seeding, where centripetal force is used to insert cells onto scaffolds [23]. Another way is by using internal or external vacuum pressure to seed grafts along a pressure gradient [24]. This method offers an inexpensive alternative to the rotational seeding technique in terms of speed and higher seeding efficiency. The most recent and popular method for cell seeding using a top-down approach is using 3D bioprinting to deposit the cells along with other materials, in a layer-by-layer fashion while fabricating the vascular graft [25].

Just fabricating a vascular graft using all the components may not be enough for it to function normally, particularly in vivo. In most cases, a TEVG would require maturation in a bioreactor.

One method to achieve this is by using a pump to introduce pulsatile flow through the inner wall of the graft. One of the earlier studies that demonstrated that a pulsatile flow is necessary for the SMCs to migrate throughout the scaffold of an engineered vascular graft was by Niklason et al. [98]. It is now known that ECs can also respond to fluid shear stress. This stretching activates signaling pathways that regulate proliferation, maturation, quiescence of neighboring SMCs, in addition to playing an important role in vascular barrier functions [26,27]. Recent research has also shown that external stimuli (such as matrix stretching and flow-induced shear stress) can enhance differentiation of progenitor cells, matrix deposition and contractility [28,29]. Looking at these results, use of mechanical pre-conditioning of TEVGs is justified as they can further improve vascular graft structure and function, particularly in vivo.

Engineering artificial vascular grafts and vascular networks

A basic requirement for engineered vascular tissues for potential clinical use is that it must be transplantable or it should stimulate the formation of new vasculature at the transplant site. For TEVGs, the implanted construct must be able to withstand physiological pressures without leakage or aneurysm formation. Additionally, they should not be thrombogenic and should not elicit an immunological response in the patient. The most common strategies for vascular tissue engineering include implantation of ECs within a biomaterial matrix, re-endothelialization of decellularized tissues/organs and stimulation of angiogenesis in vivo. Recently, uses of microfabrication technologies such as stereolithography, additive manufacturing, microfluidics have been showing promising results for both in vitro and in vivo vascular tissue engineering.

The main criteria for engineering a long-lasting blood vessel includes suitable mechanical (tensile) properties, high compliance, low thrombogenicity and regenerative remodeling for successful integration within the host [30]. For example, tissue engineered arteries that are implanted must be tensile enough to prevent rupture and compliant enough to accommodate flow with highly pulsatile pressure. Also, the implanted blood vessel must survive the host’s response, which often results in graft failure due to inflammation and thrombus formation. Finally, remodeling events within the engineered vessel and at the site of anastomosis should allow integration with the host tissue. Engineering such biomimetic vascular tissues would require the correct choice of materials, cell types, biological molecules/factors, fabrication methods and pre-implantation culture conditions. They should also match to the type of vessel or vasculature being replaced.

Choice of materials

Small vessels like arterioles, venules and capillaries present a different set of challenges for bioengineering because they have diameters in the sub-millimeter range. Fabrication methods for larger arteries are not usually applicable for microvessels because the clinical need here is for a vascular bed network with high endothelium surface area-to-volume ratios. From a materials point of view, scaffolds (both natural and synthetic) are used to provide regenerative cues that induce formation and maintenance of the vascular bed. Natural materials such as collagen and fibrin are commonly used for engineering microvessels. In particular, fibrin has been widely used for this due to its natural angiogenic properties [31]. Other naturally-derived materials being used in engineering microvessels include dextran, agarose, hyaluronic acid and gelatin.

Synthetic materials such as polyethyleneglycol (PEG) have been chemically modified to support vascularization and also for making them more pro-angiogenic [32]. An example of this is a study where a hyaluronic acid-based hydrogel was chemically modified with fibronectin motifs that explicitly promoted EC binding; thereby resulting in better vascularization compared with a non-modified hydrogel in a mouse stroke model [33]. Such modified biomaterials can support microvessel formation and function and have great potential for engineering clinically relevant vascular tissues. Collagen and elastin (which are natural components of an artery) have been widely used as biomaterial backbones of TEVGs [34]. Synthetic polymer materials are also good candidates for use in vascular tissue engineering. The properties needed for these materials include biocompatibility, biodegradation and non-immunogenicity. Non-thrombogenicity is also an important property and any material that has it would be desirable for vascular tissue engineering. Synthetic polymers that been used for vascular tissue engineering include polycaprolactone (PCL), poly-l-lactic acid (PLLA), polyglycolic acid (PGA) and their mixtures. When using these materials, the mechanical strength of the resulting graft is dependent not only on the initial mechanical properties of the materials, but also on the new extracellular matrix (ECM) proteins deposited by the seeded cells. Some of the materials, fabrication methods, and cells used for vascular tissue engineering are listed in Table 1.

Table 1

Common materials, cells, and fabrication methods used for vascular tissue engineering

Materials Fabrication methods Cell used Development level 
PCL Electrospinning – using a spinning counter electrode Autologous BM-MNCs In vivo (human) 
PGA Mesh sewn into a tube SMCs and ECs In vivo (including human trial) 
PCL-PGA Polymer sheets with cells concentrically wrapped Fibroblasts, ECs and SMCs In vitro 
PLLA-PGA Woven mesh None In vivo 
PCL/PLA Rapid prototyping, salt leaching, dip coating None In vivo 
PLLA-PCL/PLA or PGA-PCL/PLA Dual cylinder chamber molding None In vivo 
PGA-PCL/ PLA Non-woven mesh sequentially wound, coated, and reinforced None In vivo 
PCL-PGS heparin-coated Electrospinning and salt fusion None In vivo 
PEUU Electrospinning, thermally-induced phase separation Muscle-derived stem cells; pericytes In vivo 
PU Porogen leaching SMCs In vitro 
PGA-P4HB Non-woven PGA coated in P4HB Fibroblasts and ECs In vivo 
PCL-PGS Electrospinning and pore leaching None In vivo 
Collagen Rolling SMCs, EC, aortic SMCs In vitro 
Collagen-elastin Freeze-drying, crosslinking Umbilical vein SMCs In vitro 
Chitosan-gelatin Knitting, freeze-drying Vascular SMCs In vitro 
Fibrin Rolling SMC, nDF, vascular SMCs, BM-SM progenitor cells, ECs In vivo 
Silk fibroin Electrospinning or gel spinning None In vivo 
Silk fibroin with Collagen type I Electrospinning NIH 3T3 In vitro 
Materials Fabrication methods Cell used Development level 
PCL Electrospinning – using a spinning counter electrode Autologous BM-MNCs In vivo (human) 
PGA Mesh sewn into a tube SMCs and ECs In vivo (including human trial) 
PCL-PGA Polymer sheets with cells concentrically wrapped Fibroblasts, ECs and SMCs In vitro 
PLLA-PGA Woven mesh None In vivo 
PCL/PLA Rapid prototyping, salt leaching, dip coating None In vivo 
PLLA-PCL/PLA or PGA-PCL/PLA Dual cylinder chamber molding None In vivo 
PGA-PCL/ PLA Non-woven mesh sequentially wound, coated, and reinforced None In vivo 
PCL-PGS heparin-coated Electrospinning and salt fusion None In vivo 
PEUU Electrospinning, thermally-induced phase separation Muscle-derived stem cells; pericytes In vivo 
PU Porogen leaching SMCs In vitro 
PGA-P4HB Non-woven PGA coated in P4HB Fibroblasts and ECs In vivo 
PCL-PGS Electrospinning and pore leaching None In vivo 
Collagen Rolling SMCs, EC, aortic SMCs In vitro 
Collagen-elastin Freeze-drying, crosslinking Umbilical vein SMCs In vitro 
Chitosan-gelatin Knitting, freeze-drying Vascular SMCs In vitro 
Fibrin Rolling SMC, nDF, vascular SMCs, BM-SM progenitor cells, ECs In vivo 
Silk fibroin Electrospinning or gel spinning None In vivo 
Silk fibroin with Collagen type I Electrospinning NIH 3T3 In vitro 

PLA: polylactic acid; PGS: poly(glycerol sebacate); PEUU: poly(etherurethane uraa); PU: polyurethane; P4HB: poly-4-hydroxybutyrate; nDF: neonatal dermal fibroblasts.

Cell sources for vascular tissue engineering

The development of an intact and functional vascular endothelium is critical for successful vascular tissue engineering. When choosing an EC lineage for vascularization, it is important to know certain characteristics such as biology of the individual cell types, its interaction with the materials of the vascular graft, the effect of the fabrication techniques etc. Also, understanding the interactions between ECs and other cell types being used and how these will interact with host cells in vivo would be valuable. The different cell types that are currently being used for vascular tissue engineering include ECs, endothelial progenitor cells (EPCs), vascular smooth muscle cells (VSMCs), pericytes and stem cells [35]. These cells can also be derived from different sources (Figure 2) and play specific roles in development and maintenance of vasculature. For ECs, the human umbilical vein endothelial cells (HUVECs) remain one of the more popular EC lines being used due to their low cost, simple isolation and high angiogenic potential [36]. However, these are not the most physiologically relevant cell type to be used for vascular tissue engineering. In different tissues, microvascular bed development is mediated by different organ-specific EC lineages. Therefore, several studies are focusing on use of organ-specific microvascular ECs (such as primary aortic, vein, lymphatic, etc.). Also, ECs derived from human embryonic or iPSCs are gaining popularity for vascular tissue engineering [37].

Figure 2

Different cell types and sources for vascular tissue engineering

Figure 2

Different cell types and sources for vascular tissue engineering

EPCs, which are unspecialized but determined vascular cells, have become a subject of interest because they can recapitulate the embryonic vasculogenesis processes both in vitro and in vivo. As these cells can be sourced from various organs, their EC differentiation efficiency is variable and thus, might contribute differentially to an engineered vessel development [38,39]. The main criticisms behind the use of EPCs are their variable lineage and relatively unknown characterization. Therefore, it is difficult to predict their efficacy in culture and for widespread use. Also, sourcing of EPCs for commercial use can be a problem. If the source of EPCs is a human donor, then this process is also time-consuming and expensive. Another concern with the use of EPCs in vascular tissue engineering is that while most endothelium in the body is naturally quiescent and proliferates only in response to injury, the EPCs being used for vascular tissue engineering are expected to be actively proliferating. Proliferating EPCs undergo a change in surface marker expression and can lose function and organ-specific characteristics. Additionally, the specific characteristics and responses of EPCs are dependent on the tissue they originate from [40,41]. Hence, the need for proliferative cells to form a vascular lining in tissue engineered constructs conflicts with the aim of using these cells to recapitulate the structure and function of the target tissue or organ. This will be a major challenge when using these cells.

Vascular smooth muscle cells (VSMCs) are present around all vessels, except the capillaries [42]. They contract circumferentially to regulate transport of blood around the body. Pericytes are the cells that maintain the basement membrane of capillaries and provide a number of supportive roles for ECs depending on the organ. In general, they mediate paracrine signaling, hemostasis and vessel stabilization [43]. Pericytes can be recruited and differentiate into vascular mural cell types [44]. Historically, autologous vascular cells harvested from primary tissues (such as vein segments) have been a common source for vascular tissue engineering. However, there are significant challenges in their sourcing and use, including harvesting, prolonged in vitro expansion due to their reduced proliferative capacity and overall lack of cell availability from certain patients due to age or disease [45]. Advancements in stem cell technology have made use of adult and induced stem cells as promising alternatives to autologous vascular cells [15]. Additionally, iPSCs are increasing being used for vascular tissue engineering because they have shown superior in vitro proliferative capacity compared with both adult primary and stem cells. The different cell types used for vascular tissue engineering are listed in Table 2.

Table 2

Cell types and their advantages or disadvantages for use in vascular tissue engineering

Cell category Cell types Advantages Disadvantages 
Embryonic stem cells ESCs Can be differentiated into vascular ECs and SMCs
ECSs-derived ECs and SMCs are capable of capillary tube formation 
Difficult cell sourcing
Efficiency of generating vascular ECs is very low (1–3%)
Ethical issues
Regulatory issues 
Progenitor cells Vascular EPCs
BM-derived SM progenitor cells 
Relatively easy availability (from blood or BM)
Culture for multiple passages
Greater replicative and regenerative capacity 
Some types of patients (including elderly ) can have very low availability 
Adult stem cells MSCs
Adipose-derived MSCs
Muscle-derived MSCs
Umbilical-cord derived MSCs
Hair-follicle derived MSCs
BM MNCs 
Multiple sources (blood, BM, adipose tissue, muscle, live, etc.)
Easy culture; can be differentiated into SMCs and fibroblasts for blood vessels
Can support establishments of ECs within TEVGs
Can possess anti-thrombogenic properties 
Harvesting from many sources is invasive
Differentiating them into ECs can be challenging 
Autologous somatic cells Vascular-derived ECs,
Vascular-derived SMCs
Vascular-derived fibroblasts
Dermal fibroblasts 
Easy sourcing; easy culture and expansion
Have widely been use for making TEVGs and for cell seeding in vascular networks 
Sourcing of vascular cells needs invasive procedures
Quality of cells would depend on the health of the source vasculature
Limited regenerative potential 
Non-autologous somatic cells Dermal fibroblasts
Vascular-derived fibroblasts 
Can be used when autologous cells are limited or difficult to obtain
Can help avoid time delays associated with sourcing and expanding autologous cells
Can be used for off-the-shelf products
A variety of non-autologous cell types can be used 
Can have immune compatibility issues
Getting regulatory approval for their use can be challenging 
Cell category Cell types Advantages Disadvantages 
Embryonic stem cells ESCs Can be differentiated into vascular ECs and SMCs
ECSs-derived ECs and SMCs are capable of capillary tube formation 
Difficult cell sourcing
Efficiency of generating vascular ECs is very low (1–3%)
Ethical issues
Regulatory issues 
Progenitor cells Vascular EPCs
BM-derived SM progenitor cells 
Relatively easy availability (from blood or BM)
Culture for multiple passages
Greater replicative and regenerative capacity 
Some types of patients (including elderly ) can have very low availability 
Adult stem cells MSCs
Adipose-derived MSCs
Muscle-derived MSCs
Umbilical-cord derived MSCs
Hair-follicle derived MSCs
BM MNCs 
Multiple sources (blood, BM, adipose tissue, muscle, live, etc.)
Easy culture; can be differentiated into SMCs and fibroblasts for blood vessels
Can support establishments of ECs within TEVGs
Can possess anti-thrombogenic properties 
Harvesting from many sources is invasive
Differentiating them into ECs can be challenging 
Autologous somatic cells Vascular-derived ECs,
Vascular-derived SMCs
Vascular-derived fibroblasts
Dermal fibroblasts 
Easy sourcing; easy culture and expansion
Have widely been use for making TEVGs and for cell seeding in vascular networks 
Sourcing of vascular cells needs invasive procedures
Quality of cells would depend on the health of the source vasculature
Limited regenerative potential 
Non-autologous somatic cells Dermal fibroblasts
Vascular-derived fibroblasts 
Can be used when autologous cells are limited or difficult to obtain
Can help avoid time delays associated with sourcing and expanding autologous cells
Can be used for off-the-shelf products
A variety of non-autologous cell types can be used 
Can have immune compatibility issues
Getting regulatory approval for their use can be challenging 

In vitro differentiation of pluripotent stem cells into ECs and SMCs is also becoming common due to use of special media and growth factors. For example, in one study, selective media supplemented with growth factors and metabolites [such as platelet-derived growth factor (PDGF-BB), retinoic acid (RA) and transforming growth factor-β (TGF-β)] have been shown to promote differentiation of progenitor cells into a SMC phenotype [46]. Recently, another study demonstrated that exposure to shear stress can further specify iPSC-derived ECs into an arterial lineage [47]. The phenotypic plasticity and function of iPSC-derived ECs in response to humoral, biomechanical and pharmacological stimuli was shown to be similar to that seen in adult tissue ECs [48]. These studies highlight the amenability of iPSC-ECs to vessel-specific functions and physiological cues; thereby supporting their use in engineering vascular tissues.

An advantage of using stem cells is that they also produce ECM components (such as collagen and elastin) with higher yields in comparison with autologous cells. In one study, when tubular PGA scaffolds were seeded with SMCs differentiated from human iPSCs, abundance of collagen deposition was observed and the resulting TEVG was mechanically strong enough to be surgically sutured in vivo [49]. Stem cells have also been used as a source of cells to create microvascular beds using a bottom-up approach. For example, human iPSCs (hiPSCs) were differentiated into early vascular cells and were seeded onto synthetic hyaluronic acid-based hydrogel [50]. Results at day 3 showed the presence of complex vascular networks with patent luminal structures, ECs containing lumen lining and pericytes encircling the ECs. In another study, EPCs generated from hiPSCs were used to form a vascular network. When co-implanted with mouse fibroblasts in collagen in a mouse cranial model, the vessels remained stable for 280 days [51].

Use of biological factors for vascular tissue engineering

Angiogenesis is dependent on cell–cell and cell–matrix interactions. Therefore, careful orchestration of the appropriate signaling factors and pathways is required to induce neovascularization in engineered tissues. The most well-known molecules used to induce vascularization are angiogenic growth factors. These are morphogenetic proteins that induce and support angiogenic processes such as proliferation, differentiation and migration of cells through a variety of unique signal transduction pathways [52]. The process of angiogenesis is very complex and stems from the multiple isoforms of growth factors and cytokines that interact with target receptors in time-dependent and spatially-controlled ways [53]. Recapitulating these signaling pathways is a great challenge for translational application of pro-angiogenic factors and also their use in vascular tissue engineering. Pro-angiogenic factors also play important roles in vasculature development and therefore can have an important role in vascular tissue engineering. These include vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), angiopoietin (Ang) family of growth factors, transforming growth factor β (TGF-β), and sonic hedgehog (SHH). The VEGF family of angiogenic growth factors is considered as a gold standard and they play significant roles in both vasculogenesis and angiogenesis [54]. The VEGF family consists of six main isoforms of VEGF (VEGF-A, VEGF-B, VEGF-C, VEGF-D, VEGF-E, and VEGF-F) as well as placental growth factor (PlGF) and their associated receptors. VEGF-A remains the most commonly used angiogenic growth factor in vascular engineering applications due to its ubiquitous presence in initiating angiogenic processes and its potent effects. However, it is now known that VEGF alone is limited in its ability to induce the formation of functional, dense,and persistent vessel networks [55]. Additional modulation from other factors or cells will also be needed.

PDGF family of growth factors is closely related to VEGF [56]. While normally secreted by circulating platelets in the bloodstream, other cell types have been shown to release specific isoforms of the factor. PDGF-AA, BB, and CC are the most well-known supporting factors for angiogenesis. Endothelial cells, fibroblasts, macrophages and other vascular support cells can release this growth factor after neovessels are established. This generates a chemotactic gradient, which further attracts pericytes, leukocytes, fibroblasts and vascular SMCs that can stabilize, remodel, and mature the vessels [57]. The FGF family consists of 22 growth factors and 4 associated receptors (including their alternative splicing variants) that mediate a wide variety of functions from angiogenesis to neural development [58]. Among these, FGF-1 (also known as acidic FGF, or aFGF) and FGF-2 (also known as basic FGF, or bFGF) are of significant relevance to angiogenesis, while other isoforms have also been implicated in angiogenic pathways [59]. FGF has been shown to promote tubulogenesis and migration of ECs and EPCs. It can also assist in signaling of both ECM proteolysis and synthesis; in addition to enhancing VEGF production. All these indicate a synergistic effect on angiogenesis [59,60]. Further, just like VEGF, FGF provides cues for initiating angiogenesis by inducing EC proliferation and migration. However, both these molecules act through different pathways.

Many other cytokines participate in angiogenesis and produce downstream effects on angiogenesis in tissues. This depends on specific conditions such as location, cell type and timing. An example is angiopoietin (Ang) family of growth factors that are crucial for embryonic vascular development and also necessary for sprouting angiogenesis, vessel wall remodeling and EC survival [61]. TGF-β is another cytokine that is released by ECs as part of the inflammatory response to recruit macrophages. It mediates the differentiation of vascular SMCs and contributes to EC proliferation, migration and tubulogenesis [59]. SHH is a morphogen that is present in both embryonic development as well as postnatal angiogenesis and has been shown to mediate expression of angiogenic growth factors like VEGF-A and Ang isoforms. It also promotes angiogenic cytokine production from stromal populations and enhances recruitment of EPCs to remodel microvascular networks [62]. Matrix metalloproteases (MMPs) play an important role in angiogenesis and vessel stabilization; thereby deserving important considerations during vascular tissue engineering. MMPs initiate ECM remodeling events that are required for EC sprouting [59]. MMP-14 (also called MT1-MMP) in particular is a membrane-bound MMP that has been shown to play a crucial role in facilitating angiogenesis by degrading the ECM for cell migration and activating membrane-bound growth factors [63].

Technology trends for artificial vascular grafts and vascularized tissues

There is an urgency to develop engineered vascular grafts (both small and large diameter) in order to address multiple clinical needs. Tissue engineering of small vessels (<5 mm in diameter) is relatively more challenging due to the high failure rates of current synthetic arterial grafts. One of the early attempts at engineering vascular replacements was the “Sparks’ Mandrel” construct [64]; which was made by subcutaneously implanting a 5.1 mm diameter mandrel covered by a loosely knitted synthetic polymer polyethylene terephthalate (Dacron) in the patient to generate autogenous conduits made of fibrous tissue. It was a novel method for generating a vascular graft at that time. However, there were several issues, including thrombosis and aneurysms formation. The first report of an in vitro generated blood vessel was published in 1986, where Weinberg and Bell created a blood vessel construct where collagen gels were cast around a central mandrel supported by a Dacron mesh sleeve and fibroblasts were seeded as an outer layer [65]. Although this engineered artery was never implanted successfully in vivo due to low burst strength and other issues, it represented an important conceptual advancement that encouraged further development of engineered blood vessels using co-culturing of ECs with other cells. Other approaches toward development of small diameter vascular grafts include use of fibrin-based hydrogels containing fibroblasts to cast into a tubular mold [66] and formation of a matrix rich graft followed by decellularization to remove the cellular component (and associated antigens) [67]. Recent advancements in additive manufacturing are making it possible to create a wide range of vascular architectures, from straight vessels to highly branched tubular networks. For engineering microvasculature, the current approaches can be broadly divided into two types: bottom-up and top-down methods (Figure 3).

Figure 3

Two different approaches for engineering vascular grafts or tissues [68]

(A) Bottom-up methods, where physical or chemical stimulants are used to induce angiogenic sprouting (i) or vasculogenic self-assembly of ECs (ii) that create a microvascular network. (B) Top-down methods, where pre-designed structures are fabricated using sacrificial 3D printing (i), spatial laser-degradation (ii), or layer-by-layer assembly (iii).

Figure 3

Two different approaches for engineering vascular grafts or tissues [68]

(A) Bottom-up methods, where physical or chemical stimulants are used to induce angiogenic sprouting (i) or vasculogenic self-assembly of ECs (ii) that create a microvascular network. (B) Top-down methods, where pre-designed structures are fabricated using sacrificial 3D printing (i), spatial laser-degradation (ii), or layer-by-layer assembly (iii).

The bottom-up methods depend on supporting vascular-related cells to recapitulate the physiological mechanisms for new vessel formation; similar to events seen during development or events such as wound healing. The first type of bottom-up fabrication method is vasculogenesis-driven vascularization. This involves self-assembly of ECs and/or EPCs to form vascular lumen; which then transforms into multicellular networks [69]. One of the earlier studies of vasculogenesis-driven vascularization methods was by Davis et al. [70], who used HUVECs and mesenchymal cells embedded in a collagen matrix to demonstrate formation of improved human EC containing capillary network. Chen et al. used a similar method to create a vascular network, but instead used HUVECs and fibroblasts in a fibrin gel [71]. The second type of bottom-up fabrication method is angiogenesis-driven vascularization. Several studies have tried to induce angiogenic sprouting in vitro using fluid shear stress [72,73] and chemical gradients [74]. In one study, Nguyen et al. used a cocktail of angiogenic factors including VEGF, phorbol 12-myristate 13-acetate (PMA), sphingosine-1-phosphate (S1P) and monocyte chemoattractant protein-1 (MCP-1) to show that the neo-vessels formed have lumen and can be perfused [75]. One advantage of using the vasculogenesis-driven approach over the angiogenesis-driven approach is its scalability.

In a top-down approach, the architecture and geometry of the vasculature are pre-designed and fabricated before cells are introduced. 3D bioprinting represents the most recent trend in fabricating vascular tissues using a top-down approach; where patterning of biomaterials and cells at high resolutions in 3D is possible [76]. Extrusion-based and Inkjet 3D bioprinting techniques have been used to directly fabricate tubes that can act as vasculatures. For example, Zhang et al. used an extrusion-based 3D bioprinter to create chitosan-alginate tubes (with diameters as low as 200 µm) and variable pathways within alginate hydrogels [77]. In a different study, DiVito et al. used a hydrodynamic shaping device to directly fabricate 125 µm vessels within in a matrix of methacrylated gelatin, PEG, collagen, fibronectin and hyaluronic acid along with HUVECs or a combination of HUVECs, SMCs and vascular pericytes cell [78]. Another similar approach was used by Miller et al., who used a thermal-based bioprinter to create a sacrificial lattice (made of sucrose, glucose, and dextran) and embed them in a hydrogel (made from fibrin, alginate, agarose, matrigel or PEG-based gel) [79]. When the sacrificial lattice was removed, channels with diameters as low as 150 µm were formed that could be perfused with HUVECs to generating a confluent vascular wall.

Laser-based 3D bioprinting is a sub-class of 3D bioprinting which uses laser pulse on a bioink (with cells) that is coated with an energy-absorbing layer (usually a metal like titanium) [80]. When used in a layer-by-layer process, these bioink droplets can build a 3D structure such as a vascular graft. Wu and Ringeisen used this technique to pattern a vascular network, where HUVECs and SMCs were directed into branch and stem structures using laser-assisted bioprinting [81]. The deposited HUVECs could form inter-connected lumen, while the deposition of SMCs near this lumen supported SMC proliferation and stabilization of the HUVEC lumen. Another related laser-based printing technology is the multiphoton excitation-based laser bioprinting which uses principles similar to multi-photon microscopy in order to print 3D scaffolds out of biological materials [82]. However, in spite of enabling very high resolutions (close to 3 µm), it currently does not support the printing of cells. Laser induced degradation is another top-down approach for fabricating a vascular network where selectively regions are degraded within a bulk using a focused laser beam. In this method, a biomaterial (with or without cells) is first crosslinked to form a hydrogel and then exposed to a laser in a pre-designed geometry (as per the vasculature design) to locally degrade the biomaterial [83]. In one such study, Heintz et al. used a structure that was derived from scanning in vivo cerebral cortex vasculature to pattern a photodegradable material [84] with channels as small as 3 mm diameters (scale of the smallest capillary vessels). However, there were issues in perfusing the smallest channels with HUVECs without clogging. Nevertheless, this strategy provided an attractive alternative to other methods in creating microvessels for engineered tissues.

Use of sacrificial materials or fugitive inks has interesting applications for vascular tissue engineering. Sacrificial materials can be printed into filamentous networks with fine resolution and then later removed from the construct (by dissolution, heating, etc.) to create microfluidic channels. These channels can then be seeded with ECs to form a functional vasculature. The first demonstration of using a sacrificial material and 3D printing for vasculature fabrication by Miller et al., where a water-soluble sugar called carbohydrate glass was used as an ‘ink’ to print channels with diameters as small as 150 µm [79]. Then a soluble ECM was printed around this sacrificial lattice and crosslinked, followed by dissolving and flushing out of the carbohydrate glass to leave behind hollow microfluidic channels. A final vascular bed was created when ECs were seeded and cultured in the microfluidic channels. Another type of sacrificial materials being used for creating vascular networks is alginate, which can be washed off with a sodium citrate solution once embedded in a hydrogel/polymer matrix. Abaci et al. created hollow channels within a collagen gel this way and seeded them with ECs differentiated from iPSCs [85]. A major achievement in this area was reported by Kolesky et al., who used sacrificial printing to generate a thick (>1 cm) engineered tissue construct with human fibroblasts and human mesenchymal stem cells (MSCs) in a fibrin hydrogel [86]. An interconnected vascular network was generated whose channels were lined with HUVECs. Top-down fabrication approaches for vascularization are attractive due to their flexibility in accommodating a variety of ECM materials, sacrificial materials and cell types.

Use of topological engineering and additive manufacturing technologies represent the latest trends in vascular tissue engineering. Topographical engineering has been used to generate conduits resembling vessels in vitro for cardiac tissues engineering by creating channels in a fibrin-based scaffold by embedding alginate fibers and seeding them with a mixture of cardiomyocytes, SMCs, fibroblasts and ECs [87]. Lumens with an estimated diameter of 100 µm were generated that could be perfused for approximately 25 days. In another study, Thomson etal. developed a porous fibrin scaffold using a template of parallel-aligned 60 µm polycarbonate fibers bundled in a poly(methyl methacrylate) (PMMA) shell and spaced with PMMA beads (27 µm diameter) [88]. Cells (including ECs) were seeded into this construct and in vitro cellular organization was observed with the ECs localizing to the lumen structures. More recently, Zhang et al., developed a synthetic microvascular chip called ‘Angiochip’ [89], which has branching networks of channels that are as small as 50–100 µm in order to facilitate molecule diffusion and exchange. This construct could also be surgically anastomosed to host vasculature to allow for immediate blood perfusion.

Use of 3D bioprinting for engineering vascular grafts has been discussed before in this review. However, this fabrication technology is also popular for pattering vascular networks in vitro and for creating tissue constructs with vascular channels. For example, using 3D bioprinting, Lee et al. created a vasculature within the printed liver tissue by using a collagen bioink with HUVEC cells, in addition to hepatocytes and lung fibroblasts [90]. HUVECs developed interconnected capillary networks with lumens, which can be suported by the fact that fibroblasts can promote angiogenesis by secreting soluable pro-angiogenic factors. In a different study, Poldervaart et al. used encapsulation techniques along with 3D bioprinting to create a Matrigel-alginate scaffold, where VEGF was encapsulated and released from gelatin micro-particles for inducing local angiogenesis [91]. Recent studies have attempted to create vascular networks by combining vasculogenesis-driven methods with microfluidics [74,92]. The use of supporting cells such as fibroblasts, pericytes and factors such as VEGF and FGF2 have been used to aid the vasculogenesis process in these studies. For example, Belair et al. used a microfluidics device to demonstrate that when human iPSC-derived ECs were co-cultured with primary human lung fibroblasts, they not only formed interconnected channels, but also perfusable vascular networks in vitro [93].

Another practical way of recreating the natural vascular organization of a specific tissue or organ is by decellularizing that tissue or organ using mild detergents and other reagents to remove all cells and associated antigens; while preserving the entire ECM, including the vascular architecture. Decellularization protocols for different tissue/organs have already been described, including liver [94], kidney [95], heart [96] and lungs [97]. The advantages of this approach is that the integrity of the vascular channel is preserved that makes it suitable for cellular repopulation (recellularization). Once the vascular networks of these acellular constructs are re-populated with ECs, microvascular coagulation can be prevented and a barrier function to the niche of parenchymal cells within the organ can be achieved. This will be a critical aspect of whole organ recellularization and bioengineering.

Use of pulsatile flow bioreactors in culturing engineered vascular grafts was a major step in fabrication of in vivo complaint grafts and vascular tissues. One of the earlier studies using bioreactors to expose engineered vascular grafts to pulsatile flow and radial strain was carried out by Niklason et al. [98], where a biodegradable PGA scaffold was seeded with VSMCs and cultured in a pulsatile flow bioreactor using medium that contained growth factors and biochemical supplements to support ECM synthesis. These engineered vessels achieved burst pressures of >2000 mmHg during in vitro tests. This was the first demonstration of fabrication and in vitro maturation of an engineered autologous artery, followed by successful in vivo implantation in a large animal model. To achieve better remodeling and development of TEVGs, juvenile animal models have also been used. One example of such a study is by Brennan et al. who used a degradable, non-woven PGA mesh tube coated with caprolactone and l-lactide copolymer and seeded with autologous bone marrow (BM) derived mononuclear cells (MNCs). After implantation into the inferior vena cava of seven juvenile lambs [99] for over a 6 months period, the graft volumes increased by about 127%; which was roughly proportional to the growth of the native pulmonary artery. This was a clear indication of native-like vascular and growth potential of these implants.

Induction of vascularization by hypoxia is an area that has not been explored much from a vascular tissue engineering point of view. It is well known that low oxygen tension can induce vascularization through the expression of hypoxia inducible factor (HIF) [100]. The HIF transcriptional complex can induce expression of multiple pro-angiogenic factors, including VEGF, PDGF, FGF and Ang [101]. Hypoxia is also known to support differentiation of progenitor cells and stem cells into the endothelial lineage [102]. In one study, Park and Gerecht [103] developed a gelatin-based hydrogel that induces acute hypoxia and seeded it with EPCs to demonstrate that hypoxia can accelerate vasculogenesis in vitro. This was mediated through the induction of a HIF-dependent increase in the expression of VEGF (along with its receptor), Ang and MMPs. This study suggested that fabrication of a vascular bed can be achieved through HIF-based vasculogenesis without the use of high-dose angiogenic drugs or other perivascular cell types. Therefore, looking at hypoxia as a strategy for engineering vascular tissues can have potential benefits; particularly for therapeutic applications based on angiogenesis-driven platforms.

Another potentially useful technology for vascular tissue engineering is a gene editing technology called CRISPR; which has been gaining popularity for treating several disorders and diseases at the genetic level [13]. Generating an ‘immune-evasive endothelium’ is a strategy that can be used to developed vascular grafts that are not subjected to immune attach from the host after implantation. In this method, vascular grafts or vascular networks within a construct is generated from ECs that have been genetically modified using the CRISPR technology. An example is a study by Abrahimi et al., who used this technology to ablate the class II major histocompatibility complex (MHC) trans-activator (CIITA) in endothelial colony forming cells (ECFCs) [104]. This ablation provided the ECFCs with very high proliferation potential that is not seen in other ECs (such as HUVECs). When these mutated EPCs were used to generate micro-vascular networks, these structures were protected from acute rejection when the host animals (immunodeficient mice) were challenged with human peripheral blood MNCs that were allogeneic to the mutated cells [105]. Therefore, this ‘immune-evasive’ engineering strategy has the potential to overcome immune destruction of non-autologous microvasculature; thereby enabling use of a wide range of cells and materials for vascular tissue engineering. Efforts are also underway for ‘humanizing’ animals using gene editing so that one day it would be possible to reduce or completely remove xenograft immunogenicity. If a similar strategy can be used to inactivate immunogenic biomolecules or make them human-compatible, then vascular xenografts and off-the-shelf vascular grafts could become reality and have a major positive impact on health care.

Importance of blood-contacting surface and anti-thrombogenicity for vascular engineering

Understanding the biology of the blood-contacting surface and anti-thrombogenic mechanisms will be critical for developing functional vascular grafts or vascularized tissues. The blood vessel lumen is covered with a confluent monolayer of ECs and a healthy endothelium plays a critical role in maintaining vascular health and hemostasis [106]. It directly contacts and interacts with substances within the circulating blood. Hemostasis and thrombus formation, which occur naturally in the vascular system, are regulated by several biological factors and interconnected processes. This includes biomechanical forces, the blood borne cells and molecules, the coagulation and complement systems [107]. From a vascular tissue engineering point of view, the coagulation cascade is of primary importance as a bioengineered graft or vasculature should be able to support it (or avoid it) as per the requirement. The coagulation cascade itself can be divided into an intrinsic and extrinsic pathway [108]. Due to its unique anti-thrombogenic, fibrinolytic and anticoagulant properties, the endothelium plays a central role in the coagulation cascades. For development of useful TEVGs and their successful in vivo use, avoiding activation of the coagulation cascade can reduce thrombus formation; thus ensuring the long-term viability of the graft. The natural anticoagulation mechanism involves several factors that can be used or mimicked to improve patency of TEVGs [109]. Some of the anticoagulation factors present on ECs includes heparans (also called heparin-like glycosaminoglycans (GAG), ADPase and thrombomodulin. Anti-thrombogenic properties of heparin-coated vascular grafts have been demonstrated in many studies [110,111].

Mechanical stresses and pulsatile flow in the blood vessels lumen also influence vascular patency. The endothelium acts as a mechanoreceptor, which senses changes in blood flow and pressure. This triggers the release of vasodilation or vasoconstriction signaling molecules that affect the VSMCs. The blood flow produces shear stress on ECs and causes hyperpolarization of the cell membrane via potassium channels. The endothelium-derived hyperpolarizing factor (EDHF) mediates this. This hyperpolarization stimulates the release of nitric oxide (NO) and prostacyclin (PGI2), leading to vasodilation [112]. Recent studies have shown that NO release can also decrease thrombus formation and improve vascular patency. Therefore, eNOS (endothelial nitric oxide synthase) and NO are important in regulating the anti-thrombogenic properties of ECs [113]. ECs can also be activated by several cytokines, including VEGF. Another observation that would be valuable for vascular tissue engineering is that laminar shear stress can have a role in maintaining natural endothelial morphology. Therefore, a better understanding and use of endothelial function, activation and control of coagulation cascades can help design TEVGs with improved anti-thrombogenic properties.

To reduce thrombogenicity and maintain patency, the TEVGs would require processing of the blood-contacting surface. For current clinical applications, use of systemic anticoagulant drug regimens is common to maintain graft patency. Suitable modification of the blood-contacting surface can reduce or eliminate the need for such drugs, thereby decreasing the corresponding risk of systemic bleeding. Heparin is the most widely used anti-thrombogenic agent for coating blood-contacting TEVG surfaces and vascular devices [114]. Also, coating of graft and scaffold surfaces with heparin and other pro-angiogenic factors have been used to enhance neovascularization. For example in one such study, decellularized liver scaffolds with heparin-bound VEGF were implanted subcutaneously in a rodent model [115]. This resulted in an increase in blood vessel formation within the scaffold at 28 days post-implantation. Another molecule with similar function is Argatroban [116]. The characteristics of a material–blood interface are determined by the material properties, effects of blood flow and the biological environment where the graft is placed [117]. A starting point for TEVG engineering should be proper material selection and understanding its interaction with blood and blood borne cells. Therefore, recreating a hemocompatible blood-contacting interface on vascular grafts should be a major goal of vascular tissues engineering.

Clinical applications of engineered vascular grafts and vascularized tissues

The first reported successful clinical application of TEVG in patients was performed by Shin’oka et al. who implanted a biodegradable pulmonary conduit construct (made of a mixture of l-lactide and e-caprolactone, reinforced with PGA and seeded with autologous bone marrow-derived mesenchymal stem cells or BM-MSCs) in a child with pulmonary atresia [118]. This clinical study reported patency and survival of the construct in the patient for 7 months. Later, this study was expanded to 23 patients who had TEVGs implanted and 19 other pediatric patients who had tissue patch repairs [119]. Another TEVG product called Lifeline graft that was made using cultured fibroblast cell sheet [120] was implanted in nine patients with end-stage renal disease (ESRD) and who were on hemodialysis [121]. Autologous engineered grafts generated from human skin fibroblast cell sheets have been implanted into patients as arteriovenous grafts for hemodialysis access. This product, made by a company named Cytograft Inc. (CA, U.S.A.), was used in a multicenter cohort study in Argentina and Poland on 10 patients with ESRD [121]. However during the study three grafts failed during the initial 3 months because of thrombosis, graft dilatation and aneurysm. An ‘off the shelf’ vascular graft produced by a company called Humacyte Inc. was implanted into 60 ESRD patients for two single-arm phase II clinical trials at six centers in the U.S.A. (with 20 patients) and Poland (with 40 patients) [122]. At 12 months follow-up, 28% of the 60 study patients had functional access patency without intervention (primary patency) while 89% of the patients had functional patency with or without intervention (secondary patency). Also, there was no evidence of a rejection or significant immunological response nor any aneurysms were reported with these acellular grafts.

Stimulation of angiogenic responses in vivo has been examined clinically using a variety of approaches, including cell-based [123], gene-based [124] and cytokine-based [125].

Heparin-coated vascular grafts have also been used in clinical trials. An example is a heparin sodium coupled-recombinant human albumin-coated vascular graft called FUSION BIOLINE (MAQUET Cardiovascular, LLC, NJ, U.S.A.). In a multicenter, randomized clinical trial with 203 patients, the FUSION BIOLINE was compared with expanded polyterafluoroethylene (ePTFE) grafts for above and below knee [126]. The primary patency was 86.4% at 6 months for the heparin modified graft and 70% for the ePTFE graft in the same time. Another approach for vascular tissue engineering using a combination of synthetic and biological modifications is currently being explored in pediatric patients with congenital heart disease. A study was reported in 2001, where a 3-year-old girl with a single right ventricle and pulmonary atresia was treated with a 1 cm diameter biodegradable vascular graft that was made from a caprolactone-lactide-PGA copolymer and seeded with the autologous peripheral vein cells (mostly myofibroblasts and SMCs) [127]. The engineered vascular graft showed patency even after 7 months and the patient did not show any signs of aneurysm or graft occlusion. This scenario is clinically important because synthetic grafts lack the ability to grow with their pediatric recipients. In another Japanese cohort clinical trial with 25 patients, all grafts were patent at 30 day follow-up without any evidence of thrombosis, stenosis, or aneurysm [119]. To assess the safety of these engineered vessels in pediatric patients, a nonrandomized Phase 1 clinical trial in currently going on in the United States. Some of the clinical studies using TEVGs are listed in Table 3.

Table 3

Ongoing, recently completed, or new clinical studies using TEVGs

Product/intervention Condition(s) being treated Sponsor(s) Status 
Personalized tissue engineered veins (P-TEV) Chronic venous insufficiency Verigraft AB, Göteborg, Sweden Phase-I
Estimated start date: March 2019 
Lifeline (tissue engineered blood vessel) ESRD (Phase I)
Hemodialysis (Phase II) 
Cytograft Tissue Engineering, CA, U.S.A. Phase-I and Phase-II Completed: December 2012 
Tissue engineered vascular grafts Single ventricle cardiac anomaly Nationwide Children’s Hospital, OH, U.S.A. Phase-I
Start date: December 2009
Completed: January 2018 
Human acellular vascular graft (HAVG) Peripheral arterial disease; peripheral vascular disease Humacyte, Inc., NC, U.S.A. Start date: October 2013
Estimated completion date: June 2019 
Human acellular vascular graft (HAVG) Chronic kidney failure; end-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-I
Start date: May 2013
Estimated completion date: May 2019 
Human acellular vascular graft (HAVG) Chronic kidney failure; end-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-II
Start date: April 2016
Estimated completion date: April 2026 
Human acellular vessel (HAV) Vascular access; hemodialysis; renal failure; end-stage kidney disease Humacyte, Inc., NC, U.S.A. Phase-III
Start date: May 2016
Estimated completion date: September 2021 
Human acellular vessel (HAV) End-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-III
Estimated start date: December 2019
Estimated completion date: December 2023 
Human acellular vessel (HAV) Vascular system injury; vascular trauma Humacyte, Inc., NC, U.S.A. Phase-II
Start date: September 2018
Estimated completion date: September 2023 
Product/intervention Condition(s) being treated Sponsor(s) Status 
Personalized tissue engineered veins (P-TEV) Chronic venous insufficiency Verigraft AB, Göteborg, Sweden Phase-I
Estimated start date: March 2019 
Lifeline (tissue engineered blood vessel) ESRD (Phase I)
Hemodialysis (Phase II) 
Cytograft Tissue Engineering, CA, U.S.A. Phase-I and Phase-II Completed: December 2012 
Tissue engineered vascular grafts Single ventricle cardiac anomaly Nationwide Children’s Hospital, OH, U.S.A. Phase-I
Start date: December 2009
Completed: January 2018 
Human acellular vascular graft (HAVG) Peripheral arterial disease; peripheral vascular disease Humacyte, Inc., NC, U.S.A. Start date: October 2013
Estimated completion date: June 2019 
Human acellular vascular graft (HAVG) Chronic kidney failure; end-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-I
Start date: May 2013
Estimated completion date: May 2019 
Human acellular vascular graft (HAVG) Chronic kidney failure; end-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-II
Start date: April 2016
Estimated completion date: April 2026 
Human acellular vessel (HAV) Vascular access; hemodialysis; renal failure; end-stage kidney disease Humacyte, Inc., NC, U.S.A. Phase-III
Start date: May 2016
Estimated completion date: September 2021 
Human acellular vessel (HAV) End-stage renal disease Humacyte, Inc., NC, U.S.A. Phase-III
Estimated start date: December 2019
Estimated completion date: December 2023 
Human acellular vessel (HAV) Vascular system injury; vascular trauma Humacyte, Inc., NC, U.S.A. Phase-II
Start date: September 2018
Estimated completion date: September 2023 

Source: clinicaltrials.gov; information current as of December 2018.

Summary and future outlook

The shortage of autologous blood vessel for transplantation and the disadvantages associated with current synthetic vascular grafts has been driving the development of TEVGs. TEVGs are being developed to mimic the structure and function of native blood vessels. Also, the focus is also now on generating vascular networks in vitro and fabricating 3D tissues and organs with a fully perfusable vasculature. Since its beginning, the field of vascular tissue engineering has seen many milestones. They include the first construction of a biological blood vessel containing cells [65]; first report of an in situ generated living pulmonary artery conduit made using cells and synthetic biodegradable scaffolds [128]; development of small diameter arteries that were matured in a pulsatile perfusion bioreactor [98]; the first clinical application of a TEVG [127]. In spite of all the challenges that still exist for vascular tissue engineering, the above-mentioned milestones represent significant advances toward the clinical application of TEVGs.

The uses of biomaterials and 3D fabrication techniques have advanced the field of vascular tissue engineering. The choice of appropriate biomaterials is critical because there will be interactions when placed into a physiological environment (cell–material, material–immune cells, material–body fluids etc.). Multiple molecules participate in angiogenesis via different pathways. This means that in order to stimulate a preferred vascularization response, the right choice of molecules will have to be used. For inducing vasculature in bioengineered tissues, exposure to multiple pro-angiogenic factors might be a useful strategy. Incorporation of growth and pro-angiogenic factors in the matrix to support neo-vascularization has been a well-studied approach. However, it has its inherent drawbacks. Due to immobilization within the matrix and random distribution, the release of these factors can generate random chemotactic gradients; thereby leading to an ill-defined spatial control and guided vessel formation [3]. Hence, more sophisticated approaches are needed for modulating angiogenesis within a scaffold matrix. These can include seeding biomaterials with relevant cells, spatiotemporal arrangement of growth factors and encapsulation of the factors to control its distribution and release.

Topographical engineering has become an effective technique to use biomaterials for inducing and supporting angiogenesis. In this direction, 3D bioprinting has emerged as a promising technology in tissue engineering; particularly for vascular tissues. Many different cell types can now be printed by mixing them with hydrogels (referred to as bioink) that have all components to support their nourishment and survival during and after printing. Stem cells seem to support better proliferative and morphogenetic capacity for vascular development in the tissue engineered constructs [129]. Moving forward, a better understanding of the role of stem cells in engineered vascular tissue environments will be needed for better outcomes. A current challenge is engineering of microvasculature, particularly as part of a larger tissue construct. Complex microvascular networks have been fabricated and tested in pre-clinical studies [130, 131]. Their successful integration into the host tissue without significant immunologic reactions is an encouraging indication that this technology can move forward toward clinical testing. As we move toward engineering of human size organs, the biomolecular pathways and environmental factors that control organ specific heterogeneity of vascular cells (including ECs) will have to be deciphered and defined. Immune response against vascular grafts is an issue that needs to be addressed urgently. Some studies have shown that use of stem cells in vascular constructs can make them immunotolerant [132]. On a broader level, the ability to control immune acceptance of cell-based vasculature will be important for their success in vivo.

Accurately assessing vascularization in vivo is still a fundamental challenge in vascular tissue engineering. Use of non-invasive imaging techniques presents the best solution to assess neo-vascularization in implanted bioengineered tissues (particularly in deep tissues). Current methods of using static 2D imaging, often at a single time point, in combination with histological analysis cannot give the whole picture of the neo-vascularization or vascular remodeling happening in the implants. High resolution imaging, non-terminal imaging techniques and dynamic imaging modalities with multiple time point assessments will have to be further developed to account for the dynamic nature of angiogenesis. Combination of imaging techniques can also be beneficial [133, 134]. There are also no standardized methods for characterizing vascularization in bioengineered constructs. Current methods mostly rely on vague spatial and temporal metrics, such as number of vessels per unit area formed in the scaffold or construct. Proper quantification of new vasculature should include the presence of lumen within an endothelial network and not just EC aggregates [135]. Use of additional molecular markers such as NRP1 for arteries and NRP2 for veins can also be helpful [40]. The development of bioengineered tissues and organs as therapeutic replacement has become a feasible and realistic option. There are inter-related pathways for vascular development depending on the medical condition and tissue type. Identifying an optimal cocktail of growth factors, optimizing their spatiotemporal presentation, developing controlled-release strategies and use of cellular engineering strategies can all help to create highly vascularized networks within bioengineered tissues or organs. The ability of engineered vascular grafts and networks to handle physiological characteristics of in vivo vasculature (homeostasis, permeability, flow rates, immune response etc.) should also be explored in prospective studies.

Understanding the biology of the blood-contacting surface and anti-thrombogenic mechanisms will be critical for developing functional vascular grafts or vascularized tissues. The blood vessel lumen is covered with a confluent monolayer of ECs and a healthy endothelium plays a critical role in maintaining vascular health and hemostasis [106]. Due to its unique anti-thrombogenic, fibrinolytic and anticoagulant properties, the endothelium plays a central role in the coagulation cascades. For successful in vivo use of engineered grafts and vascular tissues, avoiding the activation of the coagulation cascade can reduce thrombus formation and thus ensure long-term viability of the graft or tissue. Having a functional blood-contacting surface is a necessity; irrespective of whether the engineered vascular tissue is made up of only biomaterials or has cellular components. Different strategies have been used to achieve this, including use of biomaterials that are resistant to thrombosis; chemical modification of the graft surface; using cell in the graft lumen that create an anti-thrombotic effect; and stimulating in vivo endothelialization [136]. Another strategy is in situ recruitment of EPCs and ECs. For this approach, the EPC/EC specific adhesion peptides, EPC/EC capturing antibodies and/or growth factors are being included in vascular scaffolds with the aim of enhancing endothelialization of the graft lumen [137139]. Currently, there is no commercialized small-diameter vascular graft. The difficulty in generating small diameter vascular grafts for clinical applications is generally attributed to issues at the blood-contacting surface, such as intimal hyperplasia, inflammation and thrombosis. These issues can also occur in engineered vasculatures with larger diameters. Therefore, the blood-contacting surface of engineered vasculature has become an important focus of research in this field. As new strategies for cell seeding for engineered vasculature are being developed [140], creating better blood-contacting surface and a more in vivo-like blood vessels is becoming a reality.

Pediatric patients present another challenge in vascular tissue engineering because their vasculature is constantly increasing in size and complexity. Vessel or vasculature remodeling is a major phenomenon in pediatric patients and this need to be understood in detail before engineering TEVGs for this class of patients. Mouse animal models have been mostly used to identify cellular populations in remodeling response and occasional stenosis that are seen in pediatric patients. An interesting observation from one such study was that cells seeded onto the graft scaffold did not get incorporated into the remodeled vessel in vivo [141]. Instead, host cells were recruited to populate the synthetic scaffold. Mechanisms of events such as these needs to be investigated and used appropriately in making vascular grafts for pediatric patients.

Challenges are the fuel for innovation and advancement. This has been true in many areas of science, technology, and society. One such challenge in the area of vascular tissue engineering that needs special mention here is the Methuselah Foundation/National Aeronautics and Space Administration (NASA) Vascular Tissue Challenge [142]. NASA, in partnership with the nonprofit Methuselah Foundation’s New Organ Alliance, is administering this Vascular Tissue Challenge that aims to award the top prize to the first three teams that ‘successfully create thick, metabolically-functional human vascularized organ tissue in a controlled laboratory environment’. The requirements that have to be met by the competing teams include producing vascularized tissue that is more than 1 centimeter in thickness; tissue should maintains more than 85 percent survival of the required cells throughout a 30-day trial period and demonstrate three successful trials with at least a 75 percent success rate before the competition end date. From NASA’s perspective, the successfully created vascularized tissue models from this challenge could be used for studying deep space environmental effects (cosmic radiation, loss of gravity, etc.) on humans and also for developing strategies to minimize this damage to healthy cells and tissues. Additionally, the efforts and outcomes from this challenge can complement current and future efforts at developing thick vascularized tissues for drug development, toxicity testing, disease modeling, organ transplant and other applications here on earth.

In summary, the ultimate goal of vascular tissue engineering is to create functional vascular grafts and vascularized tissues that have architectures tailored to the needs of the target tissues or organs. This can only be realized when we have a deeper understanding of basic biology that includes cell–material interactions; physical and chemical cues that drive cellular signaling in the vascular system; the host inflammatory and healing response; interactions and modulation between vascular cells etc. Innovations in the field such as biomaterials, fabrication technologies, cell culture, cellular engineering and non-invasive imaging technologies must continue to support this goal. This way, the native structure and function of the damaged tissues and organs can be recapitulated and many diseases and medical conditions could be treated. As in other areas of tissue engineering and regenerative medicine, standardization of materials and methods is of primary importance. So is the proper regulatory framework that can support clinical translations of vascular tissue engineering technologies. Looking at the current advancements in this field, the future looks very promising.

Competing Interests

The authors declare that there are no competing interests associated with the manuscript.

Author Contribution

P.C. and A.A conceived and wrote this article.

Abbreviations

     
  • 3D

    three-dimensional

  •  
  • BM

    bone marrow

  •  
  • BM-MSC

    bone marrow-derived mesenchymal stem cell

  •  
  • CRISPR

    clustered regularly interspaced short palindromic repeats

  •  
  • CVD

    cardiovascular disease

  •  
  • ePTFE

    expanded polyterafluoroethylene

  •  
  • EC

    endothelial cell

  •  
  • ECM

    extracellular matrix

  •  
  • EPC

    endothelial progenitor cell

  •  
  • ESC

    embryonic stem cells

  •  
  • ESRD

    end-stage renal disease

  •  
  • FGF

    fibroblast growth factor

  •  
  • hiPSC

    human induced pluripotent stem cell

  •  
  • HUVEC

    human umbilical vein endothelial cell

  •  
  • iPSC

    induced pluripotent stem cell

  •  
  • MI

    myocardial infarction

  •  
  • MSC

    mesenchymal stem cell

  •  
  • MMP

    matrix metalloprotease

  •  
  • MNC

    mononuclear cell

  •  
  • PCL

    polycaprolactone

  •  
  • PDGF

    platelet-derived growth factor

  •  
  • PEG

    polyethyleneglycol

  •  
  • PGA

    polyglycolic acid

  •  
  • PLLA

    poly-l-lactic acid

  •  
  • PMMA

    poly(methyl methacrylate)

  •  
  • RGDC

    (abbreviation for heparin-binding sites of fibronectin)

  •  
  • SHH

    sonic hedgehog

  •  
  • SM

    smooth muscle

  •  
  • SMC

    smooth muscle cell

  •  
  • TEVG

    tissue engineered vascular graft

  •  
  • TGF-β

    transforming growth factor-β

  •  
  • VEGF

    vascular endothelial growth factor

  •  
  • VSMC

    vascular smooth muscle cell

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