**1. Introduction**

Small-diameter vascular grafts (SDVGs) with inner lumen diameter (d) less than 6 mm are required in vascular reconstructive surgery. Tissue engineering (TE) represents an emerging research field where the production of vascular grafts utilizing state-of-the-art manufacturing methods has gained grea<sup>t</sup> attention from the scientific society [1,2]. In contrast to large (d > 8 mm) and medium (d = 6–8 mm) diameter vascular grafts, which have currently been applied in a wide variety of vascular applications, such as carotid and aorta replacement, the production of SDVGs (d < 6 mm) requires further improvement [1–3]. Indeed, synthetic vascular grafts, derived from expanded polytetrafluoroethylene (ePTFE) and Dacron, serving as medium- or large-diameter vessel transplants, have shown interesting results in reconstructive surgery [4]. Long-term results of large diameter vascular grafts (LDVGs), e.g., when applied as aortoiliac substitutes, have exhibited good patency rates (90%) within the first year of implantation [2,5,6]. Additionally, medium-diameter vascular grafts, such as the carotid substitutes, are characterized by patency rates greater than 60% after the 1st year of

implantation [2,7]. On the other hand, the proper production and use of small-diameter vascular grafts in reconstructive surgery are still under evaluation.

SDVGs are initially aimed to be used in coronary artery bypass grafting (CABG), issued by manifestations of cardiovascular disease (CVD). Regarding non-communicable diseases, CVD is the most leading cause of death globally [8,9]. CVD is a group of complex disorders, including peripheral arterial disease (PAD), coronary heart disease (CHD), cerebrovascular disease, and rheumatic heart disease [8,10]. It has been estimated that in the European Union (EU), CVD causes more than 3.9 million deaths, which accounts for 45% of all deaths each year [11]. Moreover, 11.3 million new cases of CVD are reported in the EU annually [12,13]. Furthermore, the United States is characterized by an increased percentage of CVD cases and deaths [14,15]. It is estimated that more than 400,000 CABG procedures are performed in the USA annually [14,16]. The CVD occurrence is mostly related to changes in dietary habits, reduced exercise, increased working time, depression, national health care deficiencies and the occurred financial crisis [17–20]. In terms of economic burden, it has been estimated that in Greece, the mean annual healthcare cost per patient is 5495 €, 4594 €, and 8693 € for CHD, CVD, and PAD, respectively [21]. Therefore, the proper development and clinical utilization of functional SDVGs is of paramount importance.

Nowadays, a grea<sup>t</sup> number of treatments can be effectively applied in CVD. These treatments may include the change of dietary–lifestyle habits or the application of pharmaceutical and surgical approaches. In the context of vascular surgery intervention, endovascular approaches such as angioplasty, atherectomy, and stent insertion can be performed. Additionally, vascular graft transplantation may be applied as an alternative option to replace or bypass the injured vessels.

To date, the gold standard procedure for CABG is the use of autologous vessels, such as the internal thoracic artery, radial artery, and saphenous vein [22]. Among them, the saphenous vein (SV) is the most widely used graft in SDVGs replacement [23–27]. The first use of saphenous vein in the clinical setting has been reported in 1951 by Kunlin and his colleagues [28]. The SV is characterized by greater patency rates (90% after the 1st year of implantation), compared to synthetic grafts (>60%, within the first year) [7,29,30]. However, significant drawbacks also accompany the use of autologous vessels. It is estimated that >30% of patients with CVD lack suitable vessels [1,31]. Moreover, in the case of the performance of second bypass surgery, the possibility of finding another suitable vessel decreases dramatically [32]. The latter can be performed within 10 years after the initial implantation, considering that the patency rate of autologous vessels (saphenous vein) after the 5 years is approximately less than 50% [2]. Moreover, biomechanical incompliance between arteries and veins can result in neointima formation, immune system activation, and finally graft failure and rejection [32].

Taking into account the above information regarding the use of SDVGs for bypass surgeries, alternative strategies for the development of vessel conduits must be evaluated and established. Tissue engineering may assist significantly to this issue by providing evidence and new ideas for the manufacturing of suitable SDVGs, which will be capable for cell homing, growth, and differentiation, and also characterized by improved in vitro and in vivo remodeling properties. In this review, we will highlight the state-of-the-art methodologies, while the future perspectives of SDVGs will be presented in detail.

#### **2. Characteristics of Engineered SDVGs**

The manufacturing of SDVGs with the TE methodologies has been improved significantly since the first attempts for production and application of synthetic vascular grafts used in bypass surgeries in the late 1980s [33]. Several years later, the first commercially available tissue-engineered vascular grafts (TEVGs) appeared, including Synergraft® (CryoLife, Inc., Kennesaw, GA, USA), Artegraft® (LeMaitre Vascular, Inc., Burlington, MA, USA), Procol® (LeMaitre Vascular, Inc., Burligton, MA, USA), and Cryovein® (CryoLife, Inc., Kennesaw, GA, USA) [34]. The majority of these grafts have received approval from the Food and Drug Administration (FDA) and the European Medicinal Agency (EMA) for human applications.

The proper design of the vascular grafts ensures successful cell seeding at the pre- and post-implantation stage. Cellular populations may positively influence the vessel graft functionality [35]. The most applied cellular populations are the endothelial cells (ECs) and vascular smooth muscle cells (VSMCs) [36]. ECs are located in the internal layer of the vascular wall, known as tunica intima, forming the endothelium [37]. The endothelium has unique anti-thrombogenic properties, avoiding the platelet aggregations and clots formation [38]. VSMCs are responsible for vasoconstriction and vasodilation, located in the media layer of the vessel wall, which is known as tunica media [39]. Dependent on microenvironment stimuli, the ECs can elevate the levels of endothelial nitric oxide synthetase (eNOS), leading to NO production, which downstream induces the VSMCs-dependent vasodilation [38]. Importantly, VSMCs also support the vascular remodeling and regeneration with the production of extracellular matrix (ECM) proteins such as collagen and elastin [39]. Besides, the beneficial e ffects of the cellular populations may occur to the vascular graft, their successfully seeding and proliferation may require long-term cultivation periods. Additionally, the isolation and expansion of specialized cellular populations from patients with CVD is a demanding challenge [40]. To date, there is a tendency for developing readily available acellular vascular sca ffolds with improved anti-thrombogenic properties [41–44]. Indeed, these pioneering studies are focusing on the fabrication of a negatively charged synthetic surface in order to avoid red blood cells and platelet aggregation. In this way, the SDVGs must satisfy certain design criteria to be clinically available [45]. Specifically, SDVGs must have similar biomechanical properties (burst pressure, high-stress deformation, and suture strength) with the substituted vessels to avoid aneurysm and neointima development [46]. In addition, regardless of the vascular graft material, engineered vessels must be non-cytotoxic and support cell growth [45]. Engineered SDVGs must be characterized by specific ultrastructure, ensuring the cell seeding, proliferation, and di fferentiation [2]. Finally, the engineered SDVGs must not be immunogenic, and also must be characterized by in vivo remodeling and regeneration properties [2].

Nowadays, a wide variety of manufacturing techniques for SDVGs such as the use of synthetic polymers, decellularized natural matrices, bioprinting, and 4D printing have been developed, although the majority of them require further evaluation and optimization.

#### **3. TEVGs Derived from Synthetic Polymers**

Manufactured TEVGs from polymer materials have been widely used in vascular reconstructive surgery in the last years [47,48]. The use of synthetic polymers has brought a new era in surgery, decreasing the time needed for vessel conduit manufacturing. Vascular grafts produced from synthetic materials can be manufactured with state-of-the-art tissue engineering methods, including tissue engineering by self-assembly (TESA), electrospinning, and bioprinting. Among them, bioprinting has gained grea<sup>t</sup> attention from the scientific society due to the production of high-quality tissue engineering vascular sca ffolds. The manufactured sca ffolds (acellular or cellularized conduits) can be implanted in the patient to replace the damaged vessels (Figure 1). Synthetic conduits can be divided into non-degradable, degradable polymers, and biopolymers. Each category is characterized by specific characteristics, which will be further explored in this review article.

**Figure 1.** Development and implantation of engineered small-diameter vascular grafts (SDVGs). (**A**) The first approach comprises the production of acellular SDVG derived from polymer materials using the state-of-art bioprinting approach. Then, the manufactured SDVG can be implanted immediately into the patient. In this approach, the patient's body will serve as a bioreactor for the recellularization of the implanted vascular graft. However, some major disadvantages, including the time period needed for the proper cellularization or the impaired functionality of the produced vascular grafts, maybe existed. (**B**) The second approach comprises the combination of cellular populations with the polymer derived SDVGs. In this approach, the cells can be isolated from the patient's tissue biopsy, in vitro expanded, and seeded onto the SDVG. Finally, the engineered SDVG can be implanted back to the patient. The advantage of this approach is the production of compatible SDVGs with the patients, avoiding in this way any potential adverse reactions.

#### *3.1. Non-Degradable Polymers*

Non-degradable polymers were among the first materials used as a source for the production of vascular grafts that have been employed in bypass surgeries (Table 1). Historically, the first attempt for the production of ePTFE material has been performed by Robert Gore in 1969 [4]. Several years later, Campbell et al. reported promising results regarding the use of ePTFE vascular grafts in 15 patients as a femoropopliteal bypass graft [49]. In 1986, Weinberg and Bell [33] developed the first tissue-engineered blood vessel substitute through culturing of bovine ECs, VSMCs, and fibroblasts to a Dacron derived conduit. Since then, a grea<sup>t</sup> effort by the research teams has been performed establishing new strategies to obtain functional TEVGs. EPTFE, Dacron, and polyurethanes are the most used materials for the production of non-degradable vessel conduits [48]. Compared to autologous vessels, synthetic non-degradable conduits are characterized by a lower percentage of patency rates when used as SDVGs [50]. To date, Dacron is preferred to be used as a material for the production of vessel conduits due to improved biomechanical properties [48,51]. However, both of them exhibit significant adverse reactions. Specifically, a generalized immune response toward the polymers is exerted mostly by macrophages and T cells [52,53]. This could lead to lumen occlusion, which may be treated with new cardiovascular reconstructive surgery. Moreover, most of these grafts lack arginine-glycine-aspartic acid (RGD) binding sites in order to promote cell adhesion [54]. In this direction, several alternative strategies have been employed such as the chemical modification or pre-coating of the polymer materials toward favoring the cell adhesion. Indeed, the addition of P15 peptide, pre-coating with fibronectin, or cross-linked RGD binding sites have been suggested as alternative strategies for improving ECs and VSMCs seeding on polymer scaffolds [55]. A number of research groups have performed pre-coating of polymer vascular grafts with fibroblast growth factor (FGF), vascular endothelial growth factor (VEGF), and epidermal growth factor (EGF), improving in this way the ECs, VSMCs, and fibroblasts mobilization, seeding, and proliferation onto the produced graft [56–58]. Randone et al. [59] reported the efficient production of VEGF pre-coated ePTFE vascular grafts. The results of this study showed increased ECs proliferation and endothelium formation

in VEGF pre-coated grafts compared to non-pre-coated vascular grafts. In addition, Randone et al. reported that the microporous structure of ePTFE was ideal for ECs seeding, thus grafts with high porosity ( >90 μm) may have better endothelialization outcomes [59]. It is known that VEGF exerts chemoattractant and mitogenic abilities on ECs. In this way, the ECs can be attracted by the VEGF pre-coated graft [60]. During ECs proliferation, a significant amount of growth factors are released, which can further regulate the function of vessel resident cellular populations, such as the VSMCs and the fibroblasts [61].

Another important issue that should be addressed is the possibility of thrombus formation. Typically, the polymer acellular vascular grafts are preferred mostly due to the short manufacturing time that is needed. On the other hand, the absence of an organized endothelium could result in increased platelet aggregation and thrombus formation [62,63]. This series of events can cause serious adverse events to the patients that might be even life-threatening. A possible solution to this issue could be the production of polymers with anti-thrombogenic surface or polymers with the substantial release of anti-thrombogenic molecules. Hoshi et al. [64] have reported the successful production of heparin-modified ePTFE vascular grafts. Moreover, Hoshi et al. managed to develop an easily implemented approach, including the covalent link of heparin to the inner side of the ePTFE grafts, to produce vascular grafts with anti-thrombogenic properties [64]. The produced graft inhibited successfully the platelet adhesion; however, a minor negative e ffect in endothelial cell function was evident. Furthermore, heparin-modified ePTFE vascular grafts were characterized by the high stability of their modified surface area, which was retained for a long time period (28 days) [64]. Moreover, it should be noted that non-degradable polymers are characterized by specific biomechanical properties. Mismatch of tubular compliance may exist in vascular grafts derived from non-degradable polymers. This phenomenon is mostly occurred due to the pre-existing di fferences in elasticity between the TEVG and the native artery. It is known that small diameter arteries, which are characterized by specific mechanical properties, can absorb energy (pulsatile energy) during the vasoconstriction, which is further released during vasodilation, contributing to the pulsatile blood flow. In this way, a vascular graft, which is characterized by a sti ffer behavior than the native ones, can diminish the pulsatile energy by 60%. This compliance mismatch between the two vessels can lead to intima hyperplasia, immune system overactivation, and final graft failure.


**Table 1.** Representative applications of tissue-engineered vascular grafts (TEVGs) derived from non-degradable polymers.


**Table 1.** *Cont.*

#### *3.2. Degradable Polymers*

Degradable polymers can be used as an alternative strategy for the production of SDVGs (Table 2). These materials can be substantially degraded, forming a proper ECM [47]. Hydrolysis of the ester bonds of the scaffolds and the metabolism of polymers into H2O and CO2 comprises the main degradation mechanism. The most known degradable materials are the poly (lactide-co-glycolide) (PLGA), polyglycolic acid (PGA), poly-lactic acid (PLA), poly-l-lactic acid (PLLA), polyglycerol sebacate (PGS), and polycaprolactone (PC) [47,48]. The above materials have been extensively used for the production of TEVGs with large and medium lumen diameter. Currently, these polymers have been proposed as starting materials for the production of SDVGs, while their efficient in vivo application is still under evaluation. Each material is characterized by unique properties. Indeed, the molecular structure, the polymerization transition temperature, and biomechanical behavior are some of the different properties that may exist among the materials [73]. For instance, PGA is characterized by rapid degradation time, which affects its biomechanical properties [47,48]. For this purpose, the degradation time can be controlled through polymerization with other materials such as PLA. PGS, another material that is used for the fabrication of TEVGs, can be fully degraded within 30 days [47,48]. PLA is a material whose complete degradation may last over years [2]. This material is characterized by a stiffer behavior than the PGA and also by improved endothelialization and patency rates. PCL is a hydrophobic material with long-term degradation time and, due to this, can sustain better initial biomechanical properties [47,48]. The first report regarding the biocompatibility and biodegradability of the polymer materials was performed in 1966 by Kulkarni and his colleagues [74]. Specifically, it was shown that PLA does not bear any cytotoxic factors and could be used in various applications, such as the production of surgical implants, without causing any tissue reaction.

Degradable polymers represent a valuable source for the production of acellular large, medium, and small diameter vessel conduits, reducing the manufacturing time even more. On the other hand, significant adverse reactions have been reported regarding their use. One major drawback is the lack of RGD-binding motifs, leading to ineffective cell seeding and proliferation [75]. As a consequence, organized endothelium cannot be formed, which can result in platelet aggregation, clot formation, and lumen occlusion [43]. For this purpose, several research groups are evaluating novel strategies for the efficient endothelialization of the luminal surface of the polymer-derived vascular grafts [2,47,48,76]. Previous strategies including chemical modifications and lumen surface pre-coating have also been

employed to scaffolds derived from degradable polymers to improve further their functionality. Wang et al. [77] managed to develop an SDVG using a combination of PCL and gelatin. In addition, surface modification with heparin was also performed [77]. The produced vascular conduits were implanted in rats as an abdominal artery graft and remained patent for 12 weeks [77]. These grafts were proven capable of efficient recellularization by ECs. In the same way, Quint et al. [78] used a PGA vascular graft as a scaffold for in vitro recellularization with aortic SMCs. Then, these grafts were placed in a pulsatile bioreactor system for 10 weeks followed by decellularization [78]. The occurred acellular vascular graft was reseeded with ECs and endothelial progenitor cells (EPCs) in order to avoid thrombus formation. Finally, the vascular conduit was implanted to a porcine model as a common carotid artery interposition graft and remained for 30 days [78]. The results of this study showed the efficient production of a personalized vascular graft, which has retained its ability for in vivo remodeling [78]. To date, a small number of clinical trials with degradable SDVGs have been performed (Table 2). Specifically, Lawson et al. [79] developed PGA-based SDVGs that were initially repopulated with VSMCs in a bioreactor setting. Then, pulsatile cyclic distension for 8 weeks, was applied to the SDVGs, followed by decellurization procedure. The occurred acellular SDVGs were applied as an arteriovenous graft in 60 patients (divided into two studies). In both studies, the average primary patency rate was 58% and 23%, after 6 and 12 months, respectively [79]. No aneurysm formation or immune response against the SDVGs was observed in all patients. In total, 4 patients died from end-stage renal disease (ESRD) manifestations rather than vascular graft complications. Moreover, histological analysis in SDVGs segments after 16 weeks of implantation showed infiltration by CD68+ monocytic cells, SMA+ VSMCs, and CD31+ ECs [79]. On the contrary, no T or B cells were evident in the histological analysis. The above outcome is quite promising, widening in this way the clinical feasibility of degradable SDVGs.


**Table 2.** Representative applications of TEVGs derived from degradable polymers.


**Table 2.** *Cont.*
