Introduction and review of literature
1.2 Review of literature
1.2.2 An overview of vascular tissue engineering
believed to be solely originating from the deposition of cholesterol plaques along the wall and is now known to be associated with inflammation, is the most common cause of artery blockage [17, 18]. In 1960, Robert Goetz and colleagues performed the first coronary artery bypass grafting (CABG) using Rosenak (tantalum) rings [19], and later in 1962, David Sabiston successfully anastomosed right coronary artery with saphenous vein graft without rings in a completely hand sewn manner [20]. It was the starting point of bypass surgery using autologous grafts, which remains the gold standard after almost six decades. The saphenous vein, internal mammary artery, and radial artery are the most frequently used autologous grafts; however, the literature suggests that the chances of vein graft failure are more likely than arterial grafts for CABG [21]. The anatomical differences between veins and arteries make vein grafts more susceptible to aneurysms and thrombosis when used for an arterial replacement. The longevity of arterial grafts is markedly superior. The outperformance of arterial grafts could be attributed to either higher prostacyclin production [22] or a mature vasa vasorum network in the arterial adventitia, which ensures better blood supply in the vessel wall [23]. Among various autografts, the left internal mammary artery (LIMA) remains the most preferable and successful choice of clinicians for CABG because of its patency rates (better than SVG) post revascularization. The superior patency of LIMA is attributed to lesser fenestrations in the endothelial layer, higher production of NO, and other antithrombotic molecules (tissue plasminogen activator and heparin). The limited permeability of endothelial cell layer makes it impervious for lipoproteins, hence minimal chances of atherosclerosis [24].
Soon after CABG became the standard therapeutic approach for coronary artery disease, doctors realized the limited availability of autografts and started exploring synthetic grafts (most commonly used are: Dacron and PTFE). Synthetic grafts were developed before the routine use of autologous grafts. Almost a decade after the development of Dacron by J.T. Dickinson and J.R.
Whinfield, DeBakey was the first to use them for aortic reconstructive surgery in the early 1950s.
Since then, Dacron grafts have been the regularly used synthetics for large diameter vessel revascularization, including lower extremity bypass surgery and aortic replacement [25]. Dr. Roy Plunkett developed another key player in the field, PTFE, at DuPont in the late 1940s; however, its expanded derivative (ePTFE) was launched in the market because of its better compliance and porosity than the former one. ePTFE grafts witnessed first implementation in the vascular surgery field in the mid-1970s as a lower extremity bypass conduit. At present, ePTFE grafts are frequently used for arterio-venous and lower extremity bypass grafting [25]. While synthetics were successful
for replacing larger vessels (>6mm diameter), they performed poorly when used for smaller vessels. Hence continued the search for a suitable grafting vessel. In the 1960s, Charles Sparks first started using silicone tubular structures for arterial replacements [26] and later moved on to a tissue-engineering approach to developing autologous biological tubes. These patient-specific grafts were prepared by subcutaneous implantation of a silicone mandril in the patient’s leg, which produced fibrotic tissue tubes as a result of foreign body response and referred to as ‘Sparks mandril’. The clinical trials of Sparks mandril were continued up to the late 1970s, but their dismal patency rate and aneurysmal behavior forced researchers to abandon them [27].
In the 1980s, in a seminal study, Weinberg and Bell came up with the idea of producing the first in vitro TEVG consisting of vascular cells [9]. They created a native-like multilayered vessel composed of collagen gel embedded with bovine aortic vascular cells (fibroblasts, smooth muscle cells, and endothelial cells). The incorporation of Dacron sleeves facilitated tubular integrity and improved mechanical strength. Organogenesis, Inc. tried taking up this technology but failed to bring out any viable product [5]. Nonetheless, this was the starting point for tissue- engineered vascular grafts, and the idea still exists today after 34 years. Following the groundwork laid by Weinberg and Bell, many investigators started looking into various aspects aiming to create a functional blood vessel analog. Zilla, Deutsch, and colleagues showed strong evidence of synthetic grafts' clinical success (ePTFE) seeded with autologous endothelial cells [28]. In a long- term study over 15 years, they implanted 341 ePTFE grafts in a cohort of human subjects at femoropopliteal and femorodistal locations. The grafts' lumen was first coated with fibrin glue, followed by lining with autologous endothelial cells obtained from segments of saphenous, basilica, external jugular, and cephalic veins. After years, the primary patency rates of endothelialized ePTFE grafts were comparable with vein grafts validating their feasibility in clinical settings [28]. Despite encouraging outcomes, this technology's prevalent adoption is limited, likely due to the need to harvest an optimal amount of autologous endothelial cells and their expansion. Moreover, ensuring the long-term adherence of endothelial cells to synthetic surfaces is another challenging task.
In 1998, Nicolas L’Heureux and colleagues came up with the idea of creating a completely biological tissue-engineered blood vessels without using any supporting material [29]. The concept relied on the self-assembly of cell sheets, which were prepared by maintaining a confluent monolayer culture of SMCs and fibroblasts for extended periods in the presence of sodium
ascorbate. The obtained cell sheets were rolled sequentially onto inert mandrel and matured in a pulsatile bioreactor for at least 8 weeks. The long-term maturation allowed cohesion of consecutive cell-sheet layers providing a self-assembled tubular graft. Endothelial cells were seeded in the lumen at the later stage. The burst strength of these mechanically robust grafts was nearly 2000 mmHg, attributed to an organized collagen matrix [29]. Interposition grafting in the canine model demonstrated optimal suturability. Completely biological TEBVs supposedly have multiple theoretical advantages, including inherent self-renewing potential, on-demand remodeling, precludes foreign body reaction, and minimal chances of graft infection. Cytograft Tissue Engineering, Inc. took up this technology for phase I clinical trials in a hemodialysis access setting.
Unfortunately, the outcomes were not very encouraging, as the implanted grafts showed signs of dilatation over time [30-32].
The groundwork of the currently most advanced engineered blood vessel was laid by Niklason et al. in the late 1990s [33]. For the first time, the investigators used a rapidly degrading polymer, PGA, as supporting material for growing vascular cells. The degraded byproducts of PGA are metabolized by cells without inducing a chronic immune response. Bovine SMCs and ECs were cultured onto tubular PGA scaffold under physiologically relevant pulsatile perfusion system for extended periods, creating native-like blood vessel constructs. Implantation in the porcine right saphenous artery revealed good patency after four weeks. Derivatives of this technology were further developed to render immune compatibility and ready availability. In 2003, the same group showed cells’ removal from engineered vessels without compromising the mechanical properties [34]. The underlying idea was to obtain a tubular construct consisting of natural ECM derived from allogenic or xenogenic SMCs, followed by decellularization. The resulting vessels can be stored for longer periods. When required, the grafts can be retrieved, seeded with autologous ECs, and implanted in the patient. The technology was transferred to Humacyte, Inc, which initiated the clinical trials in hemodialysis access setting [35]. Notably, the HAVs were tested without autologous ECs seeding but were shown to have a mature endothelial layer in 16 weeks’ explants. HAVs are currently in phase III clinical trials, wherein their applicability as hemodialysis conduit is being tested compared with ePTFE grafts and fistulas [3].
Tissue-engineered vascular grafts composed of biodegradable polymers became a trend after Shinoka first implanted a synthetic biodegradable scaffold in 2001 [36]. The patient was born with a single ventricle and pulmonary atresia. The PCL-PLA copolymer-based tubular scaffold (10 mm
diameter) was seeded with autologous vascular cells isolated from a peripheral vein and used for pulmonary reconstruction with no postoperative complication. The intention was that the implanted graft would grow with patient age and would not require repetitive surgical interventions. Following this breakthrough, Shinoka and Breuer have been exploring the polymeric biodegradable scaffolds in combination with autologous BM-MNCs. It brings down the graft fabrication time from days to hours rendering clinical feasibility. These scaffolds were used in a clinical trial in Japan targeting extracardiac Fontan procedure in children born with single ventricle physiology. The TEVGs were implanted in a low pressure, high flow system, and long-term (10 years) outcomes showed positive results [37-39]. In continuation, an FDA approved U.S. phase I clinical trial investigated the use of these TEVGs in congenital heart surgery [40]. Researchers observed frequent graft narrowing post six months’ transplantation requiring angioplasty, which led to a premature clinical trial closure. When analyzed with advanced computational modeling, the observed data revealed the eventual reversal of stenosis over time. Hence, the model was further validated, wherein seeded TEVGs were implanted in 24 lambs as Fontan conduits connecting the pulmonary artery with IVC [40]. The hypothesis stands true even in animal models suggesting the possibility of avoiding angioplasty for patients showing early asymptomatic stenosis; however, regular medical monitoring is required. In addition to the material centric and self-assembly approaches, allogenic and xenogenic grafts are also under active investigation by several research groups [41, 42]. The primary setback of these vessels is acute/chronic rejection, which leads to allograft vasculopathy [43].
Significant efforts have been invested over the past few decades, yet the search for a translatable ‘off the shelf’ small-diameter vascular substitute continues. On a positive note, various prototypes are in various stages of clinical trials, and hopefully, few are on the verge of success.
With the improved understanding gained through past experience, it is crucial to identify and avoid the critical failure mechanisms in the future. One imperative notion is that engineered vessel need not to recapitulate all compositional, physiological, and cellular aspects of native tissue, rather focusing on a few critical aspects should serve the purpose. Foundational and follow up studies in the field indicate a few indispensable criteria for a successful tissue-engineered vascular graft.
Especially, while designing a low flow high-pressure arterial substitute, it is necessary to provide blood compatible/anti-thrombogenic lumen. The blood-contacting surface should also prevent platelet adhesion and activation to maintain long-term patency. Immune compatibility is another
crucial aspect. Autologous implants are the only candidates having the privilege of uncontested acceptance; whereas, any foreign (non-self) object activates a cascade of innate/adaptive immune system, leading to rejection. With the latest scientific advancements, the role of the host immune system in constructive graft remodeling is now established [44-46]; however, regulating the immune response in a balanced manner remains a pre-requisite. An exuberant reaction may lead to eventual hyperplasia and stenosis [40]. Another notable aspect is to ascertain the mechanical stability of the implanted graft during remodeling, wherein simultaneous occurrence of matrix degradation and neo-tissue formation takes place in a dynamic microenvironment. An imbalance of these two would succumb graft integrity and may result in aneurysm formation in the long run.
Biomechanics of the bioengineered graft needs to be optimized, keeping in mind the initial (including suture retention, dynamic compliance, tensile properties, and burst strength) and post- implantation factors (degradation and remodeling). Besides these mechanical and biological pre- requisites, successful clinical translation demands their ready availability and automated reproducibility. Economic factors like low cost and affordability should not be avoided during the course of graft manufacturing.