Bioinks

According to Moroni et al., material (s) and biological molecules or cells can be composed for the formulation of a bioink [23]. A myriad of current reviews summarized the recent achievements in the field of bioinks [20,42,60–62]. According to these studies, natural biomaterial-based bioinks, especially alginate, gelatin and fibrin, are the most cited for vascular tissue engineering applications.

While designing a bioink, key features, namely, printability, stability, biology and rheology issues, should be seriously considered and balanced [20]. The viscosity, gelation and crosslinking capabilities are the basic characteristics to consider when selecting a bioink [19]. The deviation of the produced construct from the design depends on the bioink properties [27]. For example, an increased step in viscosity leads to an improved fidelity, but also an increased shear stress leads to cell damage and activation of misleading biophysical cues related to the ECM elasticity and pores' characteristics [20].

The fidelity of the 3D manufactured construct depends on the rapidness of transition to the solid state of the bioink after the deposition. After the ejection, the decrease of gelification time improves the structure's resolution [20]. Rheological and mechanical properties can be enhanced due to nanoparticles, with the inherent characteristic of drug delivery [63,64].

In Zhang et al.'s research, human umbilical vein smooth muscle cells and sodium alginate were combined, and vasculature conduits were printed through an extrusion printer, resulting in ECM formation and in increased proliferation rate [52]. Zigzag vascular tubes were fabricated through an inkjet-based bioprinter [65]. The viability of fibroblasts was at least 80% within 72 h of culture. In addition, the laser bioprinting technique improves the interplay between different types of cells and the formation of a vascular-like network [66].

The cell sources used in the bioprinting process could be classified into allogenic and autologous. Cardiomyocytes, human umbilical cord and embryonic stem cells were categorized in the first one, whereas adipose stem cell, skeletal stem cell, induced pluripotent stem cells (iPSCs) and bone marrow derived stem cells were put in the second category [67].

#### Maturation Methods of 3D Printed Vascular Grafts

The fabrication of vascular scaffolds is usually accompanied by post-curing methods for successful cell delivery. ECM proteins are frequently used to create a cell-supporting environment [68]. Jordahl et al. reported that extended 3D fibrillar fibronectin networks improved cell invasion and proliferation [68].

Efficient graft maturation favored by si RNA and poly-L-lysine (PLL) multilayers which deposited on polydopamine-coated substrates, thus a remarkable cell adhesion was noticed [69]. Low temperature plasma treatment can also be used for the treatment and maturation of polymeric scaffolds to obtain enhanced cell proliferation. More precisely, according to the research of Liu et al., nanofiber vascular scaffolds exhibited plasma treatment and the resulting hydrophilicity of these scaffolds effectively promoted vascular endothelial cell adhesion and proliferation [70]. Biocompatible photoabsorbers favor intricate scaffold maturation during the printing process. Grigoryan et al. used tartrazine, curcumin or anthocyaninc as photoabsorbers and improved the stereolithographic production of hydrogels, hence acquiring multilateral and functional vascular architectures [71]. Moreover, bioactive soft materials with enhanced biomimetic mechanical properties may result in graft maturation. Interestingly, in the Sun et al. study, magnesium ion incorporated into 3D printed polymer, where cell adhesion and proliferation were significantly promoted [72].

#### **3. In Vivo Applications of 3D Bioprinting in CVD**

The main aim of 3D bioprinting is to design functional tissues or parts of organs in situ for in vivo applications. The pivotal problem in terms of in vivo application is the compliance of cells and hydrogels, where cells need to precisely assemble themselves together exactly after printing, to achieve an adequate cell viability and vascularization of printed tissues. Cell–cell interaction for oxygen and nutrient interchange is mandatory to promote paracrine activity and homeostasis [73].

#### *3.1. Cell Viability and Biocompatibility*

Adequate cell viability is more than debatable in printed scaffolds due to high shear stresses on the cells delivered from extremely small diameter needle tips [62]. Cell viability decreases as the wall shear stress increases and the nozzle diameter of the deposition 3D bioprinting system decreases [74]. Overall, researchers should carefully select the cell density, the alginate concentration and dispensing pressure, and the coaxial nozzle size to obtain optimum cell viability on 3D bioprinted constructs [75].

Moreover, the estimation of cell viability is of paramount importance in order to decipher the interactions and stimulations between bioinks and cells, in a way that cells will satisfactorily adhere and survive [76]. Available methods for the evaluation of cell viability in 3D printed constructs are the common assays of trypan blue, release of LDH (lactate dehydrogenase), early apoptosis detection (Annexin V), Tetrazolium dye (MTT), study of DNA damage at the chromosome level (micronucleus assay) and other similar methods [77]. The optimum method to estimate cell viability, though, is fluorescent-based probes in the form of live/dead cells. Liu et al. utilized an improved in situ microscope method, where 3D constructs were split in order to investigate layer by layer the fluorescent number of cells and categorize live/dead cells [78].

Regarding in vivo studies, Bejleri et al. used bioprinted cardiac patches composed of native decellularized ECM and human cardiac progenitor cells (hCPCs). This specific combination of bioinks achieved cell viability of over approximately 75% [79]. Moreover, patches were retained on rat hearts and show vascularization over 14 days in vivo, indicating that the patches integrate well with the native myocardium inducing nutrient exchange with implanted cells.

Ong et al. suggested that in vivo implantation promoted vascularization of 3D bioprinted cardiac patches with engraftment into native rat myocardium [80]. In this study, multicellular cardiospheres consisted of human induced pluripotent stem cell derived cardiomyocytes (hiPSC-CMs), human adult ventricular cardiac fibroblasts (FBs) and human umbilical vein endothelial cells (ECs) assembled using a 3D bioprinter, and simultaneously the cell viability, in this patch, surpassed 90%.

Biocompatibility and circumvented cell cytotoxicity are mandatory in the field of 3D bioprinting materials as mentioned before. The in vivo study of Maxson et al. supports the potential use of a collagen-based bioink as an alternative for a tissue engineered heart valve implant [81]. Results of this study showed increased host cellularization potential, biocompatibility and biomechanical behavior results. The bioink was successfully printed with MSCs and showed remodeling.

#### *3.2. Microarchitecture and Composition of 3D Construct Vascular Network*

Three-dimensional bioprinting technology aims to combine different cell types and biomaterials heading to an enhanced cell repopulation within a 3D structure. An integrated vascular network is necessary to achieve cell viability in cardiovascular 3D printed tissues. Via that network, the influx and outflow of nutrients, metabolites and regulatory molecules are achieved. Large blood vessels ensure the flow in remote distances, whereas molecular diffusion occurs between capillaries and the surrounding tissue. In addition, the size of pores of 3D bioprinted constructs plays a major role for host cell recruitment. A pore size scaffolding >1 mm enables diffusion of nutrients until sufficient vascularization is achieved [82]. In the study of Shao et al., large scale constructs with mesoscale pore networks (100 μm to 1 mm) were successfully printed and the encapsulated vein endothelial cells were spread more efficiently compared toconstructs without mesoscale pore networks [82]. In hydrogel-based scaffolding the preferable pore size of 1–150 μm provided structural support and adequate nutrient diffusion; specifically, in the study of

Zhang et al., 120–150 μm pore size resolution encouraged cells to gradually migrate into the microfibers to form a layer of confluent endothelium [83].

In the study of Maiullari et al. hydrogels and cells were printed layer by layer, thus emulating the native tissue architecture. Specifically, heterotypic human umbilical vein endothelial cells (HUVECs) and induced pluripotent cell-derived cardiomyocytes (iPSC-CMs) were transplanted hypodermically in mice and the bioprinted engineered tissue effectively merged with the host vasculature by providing enriched vascular networks [84].

Angiogenic factors play a pivotal role in the neovascularization of bioprinted cardiac tissues [85]. Notably, the tissue-engineered constructs need blood vessel development in the core. The Vascular Endothelial Growth Factor (VEGF) is used as such a regulator. VEGF regulates the vascular development and its therapeutic overexpression by the cells loaded into the construct. In this way, blood vessels sustainably grow directly into the core of the bio-engineered graft. Poldervaart et al. underlined the VEGF secretion from gelatin microparticles into the 3D constructs and the following vascularization was widely examined [85]. Further in vivo studies, regarding the effectiveness of 3D bioprinted materials, need to be implemented in order to overcome the challenge of VEGF overexpression with the intertwined side effect of vascular tumor growth (angioma) in the myocardium and other tissues [86].

#### *3.3. Improved 3D Prined Grafts in Animal Models*

Three-dimensional bioprinted cardiovascular grafts require robust control over a range of physical and mechanical properties that will enable bioink tailoring to a specific clinical application [62]. Overall, the greatest post-implantation challenge of 3D construct in cardiovascular tissue engineering is to maintain integrity and durability over time. Therefore, studies with animal models are necessary to improve the sustainability of 3D bioprinted cardiovascular grafts.

In the study of Melchiorri et al., 3D fabricated poly (propylene fumarate) PPF graft maintained mechanical properties, long-term mechanical support and physical parameters of graft (inner diameter and wall thickness) post six months of implantation in the venous system of the mice-selected animal model, while no thrombosis, aneurysm or stenosis were obtained [87]. In a rat animal model, 3D printed polyvinyl alcohol (PVA) mimicking 3D vascular grafts showed increased postoperative endothelialization during 30 days with significant decreased thrombogenesis [88]. Another study regarding a porcine animal model, utilized tissue engineered vascular graft (TEVG) with optimum anatomically fit and hemodynamic properties and adequate physical properties in a low-pressure venous system within one month [89].

#### **4. Future Perspectives**

New techniques to improve 3D bioprinting emerged due to intrinsic limitations of exogenous scaffolds or ECM-based materials [37]. Scaffold-free way of 3D bioprinting is one upcoming challenging approach to this endeavour. Tissue strands, cell sheets and spheroids, as a prefabricated block can be used for this purpose [90–92]. The "Kenzan" method is thought to be a pioneering method for bioprinting scaffold-free vascular grafts [93]. More precisely, spheroids are combined via micro-needles into contiguous structures. Thus, the achieved precision in a micron-level renders the method capable for tissue engineering purposes. In addition, the studies of Tseng et al. and Maina et al. introduce the magnetic 3D printing method [94,95]. Three-dimensional cellular blocks, which secrete their own ECM proteins, can be assembled with magnetic levitation. Bioinks of fibroblasts and smooth muscle cells are used for bioprinting cylindrical vessels 10 nm to 10 cm in length. Via this method, the scaffold degradation toxicity is remarkably eliminated.

Research on 3D printing and bioprinting has rapidly grown with the collaboration of various fields of expertise. Current breakthroughs in 3D bioprinting continue to broaden the spectrum of bioprinting methods and applications introducing nowadays 4D bioprinting which is expected to become the evolution of bioprinting and the next generation

technology, as one more dimension of transformation over time is added [96]. In this way, dynamic 3D-patterned biological constructions could alter their microarchitecture by responding to external stimuli [97].

Merging 4D time controlled bioprinting features with innovative shape memory polymers (SMPs) paves the way to enhanced treatment in CVDs while maturation and functionalization of cells in 3D constructs alters over time [98]. Hence, the necessity of self-monitoring by regaining and maintaining their bioprinted properties over time may establish a remarkable evolution, especially in the field of personalized medicine.

## **5. Conclusions**

In the realm of cardiovascular medicine, 3D bioprinting methodology leverages engineering-controlled viable biomimetic products to incorporate into clinically applicable cardiovascular grafts and tissues, heart patches, valves and other relative constructs. This review briefly summarizes the benefits and drawbacks of the 3D bioprinting method upon CVD treatment. To sum up, in order to treat a wide field of CVDs via the bioprinting method, a 3D bioprinted construct should meet the criteria of non-cytotoxicity, biodegradation, biocompatibility with preserved mechanical strength and structural integrity. Therefore, biomimicking the patient's tissue and thoroughly incorporating into surrounding tissues and organs, thus enhancing homeostasis and construct durability and viability. In conclusion, the 3D bioprinting method still has some limitations, but has mainly tangible improvements with in vivo application for clinical translation.

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

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