*5.2. Electrospinning*

The electrospinning method comprises the first attempt to mimic the complex structure of natural ECM. This method was introduced in 1930, providing an economical solution for sca ffold fabrication [231]. Nowadays, its use has been expanded, thus sca ffolds for bone and cartilage regeneration can be manufactured e fficiently. Electrospinning has relied on the production of nanoand microfibers derived from a viscoelastic solution, where a high electrostatic force is applied. More specifically, the material that will be electrospun is pumped at a slow rate, ending in a high

voltage electrical field [34,232]. This, in turn, leads to charging the polymer material during its exit from the syringe, which results in the production of the Taylor cone. A narrow jet of liquid is generated from the Taylor cone, which is further collected to a specific set up, known as the collector (Figure 3). Finally, the production of a scaffold, characterized by adequate ECM structure and fine-tuning mechanical properties, is produced [2,34]. The formation of the produced fibers is affected by various parameters, which are specific for the material, used each time, including molecular weight, surface tension, density, and viscosity [233]. Except for those, other parameters that can affect the fiber composition mostly include the applied electrostatic field, temperature, humidity, and flow rate of the polymers [233].

**Figure 3.** Fabrication methods for the production of SDVGs. (**A**) Production of SDVGs with the originally proposed method of L'Heureux et al. In this method, the production of SDVGs was relied on the self-assembly of cell sheets using a tubular mandrel. (**B**) Production of SDVGs with the electrospinning method. This methodology can produce complicated extracellular matrices (ECMs). In addition, combination with cellular populations can lead to the development of cellularized structures. (**C**) Production of SDVGs with the bioprinting method. Bioprinting offers the potential for the production of either acellular or cellularized complicated structures. Moreover, when used "smart" materials in the production process, the final product can assembly on the desired structure upon external stimulation (e.g., temperature, pH, humidity, and magnetic field).

The polymer materials used in the electrospinning approach could be either degradable or natural derived materials [2,34]. However, important di fferences between the di fferent materials exist. In the past, degradable materials such as PLA, PGA, PCL, PU/silk fibroin have been used for the production of sca ffolds and specifically tubular conduits utilizing the electrospinning approach [2,34]. Vascular grafts have also been fabricated with the electrospinning method. Importantly, the proper combination of PLGA with collagen type I and elastin can improve the mechanical properties of the produced sca ffolds, and their use is preferred for the production of electrospun blood vessels [234]. Moreover, it has been shown that the addition of naturally derived materials, such as collagen, gelatin, and fibronectin, may provide more RGD-binding sites, thus improving the cellular functions, like adhesion, growth, and di fferentiation [235].

In the context of electrospun tubular sca ffold application, both acellular and cellularized conduits have been evaluated. Wise et al. [144] developed a tubular sca ffold consisted of tropoelastin and PCL with the electrospinning method. The produced sca ffold was characterized by similar biomechanical properties as the internal mammary artery (IMA). Further investigation involved the implantation of the acellular conduit in animal models [144]. Furthermore, the biomechanical analysis was performed in electrospun vascular grafts pre- and post-implantation. Specifically, acellular electrospun vascular conduits were implanted as carotid artery interposition grafts in rats for a total period of 1 month. Histological analysis in the explants showed the successful recellularization of the vascular grafts with ECs. Moreover, the explanted electrospun vascular grafts were able to preserve the initial vessel morphology and characterized by similar biomechanical properties as the pre-implanted grafts. In this study, the successful accumulation of tropoelastin in PCL sca ffolds was shown for the first time, resulting in the production of vascular grafts, which were characterized by impaired platelet adhesion and increased endothelialization [144]. Additionally, Soletti et al. [236] provided substantial evidence regarding the proper development and production of anti-thrombogenic vascular conduits. Soletti et al. [236] showed that the acellular poly(etherurethane urea) (PEUU) grafts coated with the non-thrombogenic 2-methacryloyloxyethyl phosphorylcholine copolymer showed better patency and mechanical properties compared to uncoated PEUU vascular grafts [236]. Unlike Wise et al. [144] and Soletti et al. [236], Min Ju et al. [237] managed to develop electrospun bilayer tubular sca ffolds consisted of PCL and collagen type I. Then, ECs and SMCs obtained from female Dorper Cross Sheep were seeded onto the tubular sca ffolds, followed by maturation in the pulsatile flow bioreactor. The seeded vascular grafts were implanted as carotid artery substitutes in the sheep model and remained for 6 months. The electrospun vascular grafts were remained patent and the histological analysis revealed the production of collagen, elastin, and glycosaminoglycans within 6 months of implantation [237]. This study provided valuable data regarding the production and application of the electrospun vascular grafts. Moreover, Du et al. [238] used the electrospinning method to fabricate a 3D vascular microenvironment. In this approach, immobilization of VEGF onto the electrospun tubular sca ffold consisted of gradient chitosan and PCL nanofibers was performed. The controlled release of VEGF potentially can enhance the adhesion of ECs and SMCs and further promote their rapid proliferation. In this way, engineered SDVGs with improved anti-thrombogenic properties could be developed, leading to the avoidance of lumen occlusion and thrombus formation, a series of common manifestations which are presented several days after the vessel implantation [238]. Taking into consideration the above data, it was clearly shown that electrospinning could be applied for the e fficient production of engineered SDVGs. The produced electrospun SDVGs could be successful in vitro seeded with cellular populations and maintain further their graft patency, mechanical properties, and vessel integrity over a long time period.

#### *5.3. Three Dimensional (3D) Bioprinting*

In the last decade, 3D printing technology has gained significant attention and has been utilized with grea<sup>t</sup> success in a wide range of applications [239]. Using this technology, complex structures and materials can be produced e fficiently, thus can be further used by the scientific society. The evolution of printing technology is 3D bioprinting, which has currently been applied in various tissue engineering approaches [239]. 3D bioprinting can produce complex structures, utilizing non-degradable/degradable and naturally derived polymers [240]. The significant potential of this methodology is the production of ready to use transplantable scaffolds and tissues. Currently, the 3D bioprinting approaches such as inkjet, extrusion, and laser-assisted bioprinting are mostly used for the production of the majority of the scaffolds [240]. A grea<sup>t</sup> series of materials are compatible with the bioprinter applications, although the polymer materials are mostly preferred in comparison with the naturally derived materials [240,241]. The bioprinter materials can be distinguished into three categories: (a) fibrous materials, (b) powder materials, and (c) bioinks. The use of the starting material is dependent on the characteristics of the produced scaffold [240,241].

3D bioprinting approaches and the proper combination of the aforementioned materials have been successfully applied in the production of LDVGs and SDVGs. [240,242–244]. In this direction, Freeman et al. [245] presented for the first time a new approach for the development of SDVGs using a custom-made 3D bioprinter. In this study, gelatin and fibrinogen were properly combined, producing a bioink with good rheological and printability properties. The produced vascular graft provided a favorable ECM for cell attachment. However, comprehensive in vitro and in vivo evaluation is further needed to be performed [245]. Jia et al. [246], in their study, used a multilayer coaxial nozzle device to produce vascular grafts. Moreover, human umbilical vein endothelial cells (HUVECs) and MSCs were expanded and encapsulated in a gelatin methacryloyl (GelMA), sodium alginate, and 4-arm poly(ethylene glycol)-tetra-acrylate (PEGTA) based bioink. Using the current bioprinter set-up in combination with the developed bioink resulted in the printing of highly organized vascular structures. No sign of cytotoxicity was reported, and after a time period of 21 days, the cells filled the entire printed vascular grafts. To evaluate better the cell behavior into the vascular wall, immunofluorescence was performed, showing the positive expression of α-SMA and CD31, in MSCs and HUVECs, respectively [246].

#### *5.4. Four-Dimensional (4D) Bioprinting*

Next-generation bioprinting demands the use of materials capable of self-transform into a prerequisite shape in order to exert their key functional properties. This state-of-the-art approach is known as 4D bioprinting and has gained increased attention in the last decade by the entire scientific community [247]. 4D bioprinting uses the same materials as conventional 3D printing approaches [240]. The major difference between 3D and 4D bioprinting is that the latter exerts a "smart" behavior of the produced scaffolds [248]. 4D produced scaffolds are superior to the conventionally bioprinted scaffolds. The "smart" behavior corresponds to "materials that can change their physical or chemical properties in a control and functional manner upon exposure to an external stimulus" as has been referred to by Tamay et al. [240]. In this way, 4D printed materials upon exposure to external stimuli such as pH, heat, magnetic field, light, and humidity can adopt effectively different shapes, exhibiting different properties [249]. The above-mentioned factors are playing important role in scaffold's shape-changing properties. There exists a grea<sup>t</sup> variety of materials that achieve shape-transformation in response to temperature stimuli. Thermoresponsive materials are the most commonly used in 4D bioprinting applications [250]. These materials can be distinguished into (a) shape memory polymers (SMP) and (b) responsive polymer solutions (RPS). The first category involves polymers consisting of two distinct components, the elastic segmen<sup>t</sup> which is characterized by high glass transition temperature (Tgh), and the switching segment, characterized by intermediate glass transition temperature (Tgi). When the applied temperature is above the Tgh, the produced scaffold adopts its permanent shape. On the other hand, when the temperature is between Tgi and Tgh, the switching segmen<sup>t</sup> becomes soft, while the elastic segmen<sup>t</sup> resists any shape-changing [251]. Additionally, if the material is cooled below the Tgh, then the elastic segmen<sup>t</sup> cannot return to its initial shape and the produced scaffolds acquire its final definitive form. SMPs include mostly the poly(ε-caprolactone) dimethacrylate (PCLDMA), polycaprolactone triol (Ptriol), and poly(ether urethane) (PEU). These materials have been used mostly in applications such as bone and cartilage engineering. RPS is characterized by a critical solution temperature, where if the applied temperature is above the aforementioned temperature (critical solution temperature), the polymer chains are contracting and the overall solution is adopting a solid form [247].

Both hydrophobic and hydrophilic interactions are existing between the polymer chains. In addition, a change in temperature may a ffect the behavior and the interaction of the above polymer chains. This, in turn, leads to shrinkage or expansion, which is a characteristic of each polymer material. In this direction, a material with a critical solution temperature above 25◦C, when implanted to a mammalian organism, would expand, acquiring its definitive form. Poly(N-isopropylacrylamide), poly(ethylene glycol), collagen, gelatin, and methylcellulose are some of the most used RPS [249].

Besides the temperature stimuli, materials that can respond to pH changes also can be widely applied in the clinical setting [247]. The initial structure of these materials is consisting of acidic or basic groups, which are the main players in proton exchange upon pH changes [252]. In this way, polymers consisting of acidic groups, when exposed to pH > 7, act as anionic compounds, while polymers with a basic group, exposed to pH < 7, act as cationic compounds. Therefore, these materials upon pH stimuli can acquire di fferent structural and functional properties, including change in solubility, degradability, swelling, etc. [248]. Like thermoresponsive polymers, pH-responsive polymers also can be utilized in a wide range of applications. Indeed, di fferent human body compartments are characterized by di fferent pH in order to serve properly their initial function, including the gastrointestinal tract, stomach, small intestine, and di fferent regions of the vascular system including the kidney vascular network. Additionally, many solid tumors induce pH changes upon their growth. In this way, pH-responsive polymers can act as DDS, delivering tumor-specific therapy, such as signaling and cell proliferation inhibitors or monoclonal antibodies [253]. Examples of the most commonly used materials in this category are poly(acrylic acid), poly(aspartic acid), poly(L-glutamic acid), and poly(histidine), which can be combined e ffectively with naturally derived materials such as collagen, gelatin, and chitosan [252].

Besides the aforementioned, other categories of responsive materials have also been manufactured. These categories mostly include the photoresponsive, magneto-responsive, and humidity-responsive materials. However, their potential use is limited and, therefore, further evaluation of their properties is clearly needed. Briefly, photoresponsive materials can change their structural and functional properties, including wettability, solubility, degradability upon photo-stimulation [240]. Considering this, polymer materials with photosensitive groups can be manufactured, where the produced sca ffolds can swell or shrink when specific photo-stimulation is applied. The combination of magnetic particles with polymer materials results in the development of magneto-responsive materials. The most commonly used magnetic particles are iron (Fe), nickel (Ni), cobalt (Co), and their oxides [249] Currently, magneto-responsive materials have been used as targeted therapeutic vehicles, carrying anti-tumor drugs. On the other hand, significant adverse reactions may be induced by their use in living organisms [254]. It has been shown that magnetic particles with a size less than 50 nm are transportable through the biological matrix, which can further cause inflammation and cell death due to high reactive oxygen species (ROS) production, DNA damage, and cytochrome C release. In this category, materials such as Fe3O4/PCL, Fe3O4/poly (ethylene glycol diacrylate), PCL/iron doped hydroxyapatite (PCL/FeHA) are currently evaluated for their potent use in living systems [254]. Lastly, humidity responsive materials also have been proposed for their use in 4D bioprinting and the production of tissue-engineered sca ffolds. Interestingly the change in humidity could result in shape-change modification, which can act as a driving force for movement. These materials have not received grea<sup>t</sup> attention from the scientific community due to their limited use. Humidity responsive materials include poly(ethylene glycol) diacrylate, cellulose, polyurethane, and their combinations [255].

The 4D bioprinting comprises an important evolution in the fabrication of tissue-engineered sca ffolds. Vascular grafts can be developed with this next-generation approach. In this way, we can imagine the development of a 4D bioprinted vascular graft (with a large or small diameter), which can acquire its specific shape inside the living organism upon temperature stimulation. Moreover,

changes in pH of the vascular network may stimulate the implanted vascular graft in a way either to acquire a di fferent shape or to substantially release key therapeutic agents in order to reduce or even to reverse the occurred situation. In the future, "smart" telebiometrics vascular grafts will be plausible to be employed, which can detect the changes of human body conditions, like temperature, pH, osmolarity, and will be able to notify or even to reverse a health issue. Currently, the utilization of "smart" materials and the manufacturing of those sca ffolds is under the developmental stage [247]. Therefore, no significant number of publications is currently existing, with the only exceptions of reviews and opinion articles in this field. In this way, the development of "smart" materials that can be used in vascular engineering is quite important, but further exploration of this research field is needed.

#### **6. Concluding Remarks**

Globally, there is an increasing demand for SDVGs, as they are employed primarily in cardiovascular reconstruction surgeries. Indeed, more than 400,000 bypass surgeries are performed each year [12–15]. Parameters such as the modern way of life, increased working hours, overall stress, lack of physical exercise, and smoking comprise important risk factors for CVD development [18,20]. From an economical point of view, CVD is also a serious burden for all countries; therefore, novel and better treatment options must be utilized [21]. In this direction, the production of SDVGs and their e fficient application in patients su ffering from disorders that are belonging to CVD could be an important alternative strategy. One of the applied therapeutic approaches that are currently followed is the replacement of damaged vessels with autologous vascular grafts, such as the SV, mammary artery, and others [208]. However, CVD can a ffect the entire circulatory system, therefore, less than 60% of the patients have suitable vessels. Moreover, when second vascular reconstruction is needed to be performed in the same patient, this percentage is lower than 15% [31]. Blood vessel compatibility is another important parameter that should have in mind. Compliance mismatch between native and implanted graft (mostly at the anastomosis site) could induce unfavorable results, including calcification initiation, intima hyperplasia, lumen occlusion, platelets aggregation, and thrombus formation [32]. To avoid the above manifestations and in order to the availability of the vascular grafts to be increased, alternative strategies must be explored. Currently, the fabrication of LDVGs is efficient, using the latest TE approaches; therefore, the utilization of these methods could be applied in SDVGs production [2]. Although the significant drawbacks which engineered SDVGs may present, the interest of the scientific society is increasing day by day, exploiting better strategies to improve further the development of those grafts. Nowadays, the production of the SDVGs relies on the use of synthetic (non-degradable/degradable), naturally derived materials, and decellularized ECM. These materials can be combined with state-of-the-art manufacturing approaches (TESA, electrospinning, bioprinting) to produce vascular grafts with improved properties [2,245]. In most of these approaches, cell seeding and maturation in bioreactors (mostly pulsatile flow bioreactors) are needed in order to produced vascular grafts to e ffectively cellularized and acquire the proper biomechanical properties. Additionally, the decellularization of tissues and organs is a very promising approach, especially for the development of SDVGs. Indeed, decellularization can e fficiently be applied in vessels such as the umbilical arteries or vessels derived from cadaveric donors to produce properly defined SDVGs.

In vascular engineering, di fferent cellular populations have been proposed such as ECs, VSMCs, and stem cells, derived mostly from the recipient, avoiding in this way any unfavorable immune response and possibly graft failure and rejection [34]. However, a grea<sup>t</sup> number of cellular populations are needed for tissue engineering approaches. Terminally di fferentiated cells such as ECs and VSMCs can be isolated from a vessel biopsy, although to reach cell numbers >10 × 10<sup>7</sup> requires extended in vitro manipulation and cultivation. On the other hand, the use of stem cells such as Mesenchymal Stromal Cells (MSCs) may be a more feasible approach. MSCs initially were isolated from bone marrow aspirates, while currently other sources including the adipose tissue and stromal vascular fraction can be used [256,257]. MSCs is a heterogenic multipotent stem cell population derived from mesoderm, and capable to di fferentiate e ffectively to "chondrocytes", "adipocytes", and "osteocytes" [258]. Immunophenotypically, these cells express (>95%) CD73, CD90, and CD105, while lacking the expression (<3%) of CD34, CD45, and HLA-DR, as has been indicated by MSCs committee of the International Society of Cell and Gene Therapy (ISCT) [259]. MSCs can be easily in vitro handled, while their stemness (specific gene expression and protein production) can be retained for an increased number of passages (>P8). Another candidate stem cell population for vascular engineering may be the induced pluripotent stem cells (iPSCs). In 2006, for the first time, Shinya Yamanaka managed to gain the pluripotent state of terminally differentiated cells by introducing a set of specific genes, including OCT4, SOX3, KLF4, and C-MYC [260]. Currently, different strategies have been developed for the production of iPSCs, even avoiding the use of *C-MYC*, a known oncogene. The efficient differentiation of iPSCs into various cell populations such as neural cells, cardiomyocytes, hepatic cells, ECs, VSMCs, etc. has been demonstrated in literature [261,262]. However, this technology has not ye<sup>t</sup> received FDA approval for human clinical use, and therefore their applicability is limited even in vascular engineering [263,264].

The cellularization of TEVGs comprises a crucial step in the manufacturing process. Indeed, it has been shown that cell-seeded vascular grafts are characterized by better integration properties in the recipient's body [2]. In this way, properly cellularized vascular grafts can avoid the interaction with M1 macrophages, favoring in this way the attraction of M2 macrophages. M2 macrophages have been related to the tissue remodeling process, avoiding the activation of T and B cells [265]. Moreover, several research groups have observed the development of vasa vasorum, the responsible vessels for nutrient supplementation, into the transplanted vascular grafts, indicating its further proper adaptation by the studied living system [172,266].

The proper manufacturing of SDVGs is an important aspect, and for this purpose, specific evaluation tests must be performed before their final application. These tests include (a) histological analysis, (b) biochemical and DNA quantification, (c) cytotoxicity assay, (d) platelet adhesion assays, (e) biomechanical analysis, and (f) implantation in animal models in order to assess effectively functionality of the vascular grafts. The above processes represent the first line of evaluation tests that should be performed to assess the biocompatibility of the manufactured SDVGs. Furthermore, more tests need to be performed to properly define the produced TEVGs. Especially for the SDVGs, the performance of cytotoxicity and the platelet adhesion assays is of major importance. In contrast to LDVGs, manufactured SDVGs are characterized by an increased probability of platelet aggregation and thrombus formation. Moreover, ECs should be properly seeded in SDVGs to produce a functional endothelium; therefore, the establishment of a non-toxic vascular graft is highly recommended.

Taking into consideration the above information of this review, we can conclude that the production of SDVGs is requiring further improvement, which is performed by several research groups worldwide. Currently, the use of synthetic and decellularized vascular grafts has gained a significant advantage over other methods. Highly organized ECMs cannot be in detail reproduced with the bioprinting approaches. Indeed, additive manufacturing techniques such as 3D and 4D bioprinting are characterized by a few limitations. The inability of reproducing the highly organized structure of SDVGs may comprise the most significant drawback of the current approaches. SDVGs are characterized by a complex structure, where collagen, elastin, fibronectin, and other key ECM proteins have specific relation and orientation in the vascular wall, ensuring in this way the proper recellularization. Cellularization of the bioprinted vessel constructs may be related to improved biocompatibility and biomechanical properties. In order to produce highly organized constructs, crosslinking of the biomaterials, used in bioprinting approaches, is preferred. However, the use of fixative agents such as glutaraldehyde, can result in increased cytotoxicity and altered biomechanical properties to the produced vascular scaffolds [103,205]. Moreover, when naturally derived bioinks are used, crosslinking may hamper the in vivo remodeling process of the vascular graft, leading to unfavorable outcomes. Moreover, the manufacturing of vessel constructs with high resolution demands high-cost printing devices and experienced personnel. Besides these drawbacks, in the future, the quality of bioprinters and printed constructs will be improved, leading to a new era in SDVGs development. Besides the above

limitations, the introduction of 3D and 4D printing approaches may represent a new era regarding vascular scaffold production [240,247].

In conclusion, SDVGs now can be robustly produced and can be used in personalized medicine. Each production step must be specifically evaluated and the overall process must be performed in compliance with Good Manufacturing Practices (GMPs) conditions in order to produce readily available safe and fully functional grafts for patients suffering from CVD.

**Author Contributions:** Conceptualization, writing—original draft preparation, visualization, writing—review and editing, P.M.; supervision, project administration, C.S.-G. and A.K.; conceptualization, writing—review and editing, supervision, project administration, E.M. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Conflicts of Interest:** The authors declare no conflict of interest.
