Heart valve pathologies have been described as a cause of disability and death since the seventeenth century. They still remain today a relevant contributor of loss of physical comfort and reduction of longevity and result in a considerable socio-economic burden (1, 2). Diseases of the cardiac valves can be divided into two main categories, namely congenital pathologies (e.g., the malformation of the aortic and pulmonary valves, the Ebstein's Anomaly, the Fallot tetralogy or the bicuspid aortic valve), with an impact especially during the neonatal period and infancy, and acquired pathologies which, depending on the etiology, can have an impact at all ages (i.e., the rheumatic or the infectious heart valve disease) or in the elderly (e.g., calcification of the mitral and aortic valves) (1, 3, 4). In this framework, a major contribution to the increase in the overall impact of cardiac valves pathologies worldwide is the rapid increase of conditions leading to the aortic valve (AoV) stenosis, specifically “calcific aortic valve disease” (CAVD)-a disease correlated primarily to aging (1, 3, 5, 6) with an important sex-related component (7). We will refer below to prostheses to treat CAVD, considering that those employed to treat other pathologies are very similar in design and performance.

The AoV is composed of three leaflets, each of which comprises three laminas: the fibrosa, the spongiosa and the ventricularis, each with different structural and mechanical characteristics. The fibrosa, the layer associated with the outflow, or aortic, side of the leaflet, is predominantly composed of collagen fibers arranged circumferentially in parallel bundles and crossing with a typical “χ” geometry at the leaflet “belly” portion, surrounded by a matrix rich in elastin (8) (Figure 1). This is the layer that confers the maximal resistance of the leaflets to the compression forces acting on the aortic side, when the valve closes in diastole, and which can reach 80–120 mmHg (8). The ventricularis, the layer associated with the inflow side of the leaflet and facing the ventricular cavity, is a curvilinear structure mostly composed of elastin fibers oriented along the radial direction (Figure 1). The recoiling of these fibers supports the closure of the valve during the transition from systole (valve open) to diastole (valve closed) in the cardiac cycle (9). The spongiosa, finally, contains primarily glycosaminoglycans, a material with relatively low elastic modulus and an essential isotropic structure, which provide the deformability function of the valve leaflets and which serve to absorb the excessive mechanical forces (10). The three AoV layers are populated by specialized fibroblast-like cells, the so-called valve interstitial cells (VICs) (11). Although they are present in each of the valve layers, VICs are mostly abundant in the spongiosa, where they contribute to repair the abundant extracellular matrix continuously exposed to mechanical workload. VICs have heterogeneous phenotypes depending on the developmental and pathologic status of the valve (12).

The evolution of aortic valve pathology begins with the occurrence of micro-ruptures of the endothelial layer covering the leaflets, especially on the outflow surface of the valve, due to perturbations of the shear forces. As with the initial events of atherosclerosis, these ruptures cause lipid infiltration and recruitment of inflammatory cells (13–15). This, in turn, determines a release of inflammatory cytokines and an oxidative stress burst that lead to pathological activation of the resident VICs whose functions are altered under these conditions. Specifically, these cells participate directly in the inflammatory process by responding via Toll-like receptors 2/4 to inflammatory signaling and, in turn, secreting an array of inflammatory cytokines (16). VICs also undergo a modification in their phenotype from one of matrix-repairing to one of matrix accumulating/remodeling with the potential to cause thickening of the leaflets-the so-called valve “sclerotic” phase-which is considered the first pathologic event in valve disease (13). The phase of AoV sclerosis persists in a clinically silent fashion until the beginning of the more rapid phase of valve calcification, characterized by transformation of VICs into calcific cells. These “osteoblastic” VICs have the ability to secrete initially small, but subsequently larger, calcific nodules that progressively deform the leaflet structure (17). This causes variations in the motion of the leaflets, incomplete valve closure at diastole and regurgitation with ensuing compromisation of heart function. While inflammatory signaling is generally connected to the initial VICs matrix remodeling activity in the sclerotic phase, transformation from “activated” to osteoblastic “VICs” has been also linked to mechanical factors. In this respect, it has been hypothesized that the progressive hardening of the matrix surrounding VICs due to their remodeling activity prompts the activation of mechanosensitive-dependent pathways setting progression of the cells toward a calcific phenotype (18, 19).

Until the beginning of the current century, surgical aortic valve replacement (SAVR) has been considered the elective option for surgical treatment of heart valve pathologies. Although highly effective, this is an invasive procedure requiring temporary cardiac arrest and extracorporeal circulation, which exposes patients to complications and side effects (20). For the majority of patients, the choice of the replacement device for SAVR is either a mechanical or a bioprosthetic valve. In a minority of patient's other options are adopted. These include the recuspidalization with autologous pericardium (the so-called Ozaki procedure) or transposition of the autologous pulmonary valve into the aortic position with replacement of the pulmonary valve with an aortic homograft-the Ross procedure adopted commonly for infants and children with congenital valve defects/stenosis and young adults (21–25).

A recent novel possibility to restore the functionality of a diseased aortic valve with reduced peri-procedural side effects involves trans-catheter aortic valve replacement (TAVR). This technology, exploiting the pliability of pericardial tissue, allows deployment of a completely functional prosthesis using a minimally invasive procedure. These valves are currently approved for patients with severe, symptomatic aortic stenosis in all surgical risk categories given the favorable outcomes in the postoperative period (26, 27). Despite that TAVRs are based on a novel design and can be implanted with enormously lower risks, they still carry severe problems related to structural deterioration analogous to that of bioprosthetic grafts. For this reason, they are rarely implanted in patients younger than 60–65 years of age in accordance with the guidelines (28).

Polymeric valves are manufactured with elastomeric polymers by a simple fabrication procedure using molds. Typically, the process involves “injection molding” whereby synthetic (e.g., polyurethane or polystyrene) (29, 30) or natural polymers (e.g., fibrin) (31) are injected into tri-leaflet molds that give rise to complete sutureless valves, and can be readily mounted onto posts for implantation. This design provides advantages including easy scalability, low cost, natural hemodynamic performance and a relatively high long-term durability comparable to that of mechanical prostheses (32). On the other hand, when employed in animal valve replacement studies, polymeric valves, at least initially, did not lead to encouraging results due to calcification, thrombus and fibrous capsule formation, resulting in implant failure (33–35). Despite these shortcomings, in 2010, Quintessenza et al. (36) published a clinical study performed on 126 patients surgically treated with bicuspid pulmonary polytetrafluoroethylene (PTFE) valves. In particular, two types of valves with different thicknesses were used; the first was made with porous 0.6 mm PTFE while the second with non-porous 0.1 mm PTFE. Six patients treated with the porous PTFE valves needed reoperation due to leaflet calcification. In contrast, non-porous valves were less prone to stenosis as the lack of porosity prevented cellular in-growth and thickening (37). Moreover, 0.1 mm PTFE valves were characterized by higher leaflet mobility and lower transvalvular gradients (36). Stasiak et al. (29) recently introduced the so-called Poli-Valve (38)-a styrene triblock copolymer valve obtained by injection molding. This technique, besides being inexpensive and highly reproducible, appears to allow an optimal anisotropic distribution of forces on the leaflets and the polymeric fibers (39, 40) resulting in maximal mechanical durability due to a similar collagen fiber orientation to that of native heart valve tissue (39). The valve was bench-tested and validated according to ISO standards. Moreover, preliminary ex vivo and short-term in vivo feasibility tests were done, showing a good biocompatibility in the absence of mechanical failure and regurgitation. The lack of long-term tests in vivo still raises the question as to whether these valves offer an advantage over the most advanced mechanical replacements, in particular concerning the need for anti-coagulation therapies to prevent formation of thrombi on the surface of the leaflets.

The general strategy to derive living replacements resembling native tissues was introduced in 1993 by Langer and Vacanti (41). These Authors, in their initial proposition of tissue engineering, defined three essential steps consisting of, (i) to seed autologous or allogenic-compatible cells inside scaffolds pre-fabricated with biocompatible/biodegradable materials, (ii) to enable tissue formation in bioreactors by exposing the tissue constructs to controlled mechanical loading and, (iii) to promote final tissue maturation, exploiting the ability of the pre-seeded cells to interact with circulating cells to complete the final evolution of the tissue constructs toward native leaflets (41). If induced to mature with appropriate instructing stimuli, the tissue constructs could therefore have regeneration and growing capacity. In the valve pathology scenario, this ability to grow would be especially indicated for use in pediatric patients, for whom the possible failure of the implants is compounded by the inability of the new valve to grow with the individual, making continuous surgical procedures necessary (42).

In view of the growing awareness of the maladaptive interactions between cells and scaffolds used to produce TEHVs, more complex manufacturing concepts are now emerging based on more systematic views of the cells/scaffolds interactions (66) and the recognition of the role of the forces dominating cellular mechanosensitivity of the cells (20, 67). Central to this second-generation design is the change from scaffolds made of randomly interleaved fibers, to a design that is more compliant with the distribution of the strain and compression forces acting in the kinematics and mechanical loading of the natural valves. The idea underlying this new concept derives from the evidence that the natural ECM fibers in the valve are deposed from the very beginning of valve development mainly with anisotropic patterns instructed by mechanical forces, and that the cells residing within the anisotropically deposited fibers are adapted to maintain a quiescent phenotype (Figure 1) (68). In order to achieve this aim, one of the current trends is to manufacture scaffolds with oriented deposition of ECM components (e.g., collagen) by mechanically forcing cells to deposit fibers according to defined geometric patterns, and/or to employ polymers that can be deposited with anisotropic patterns in 3D. A further step in this biomimicry approach is the attempt to implement the natural tri-layered valve structure in scaffold design. An example of this new design has been provided by Masoumi et al. where a three-layer scaffold included an anisotropic fibrous layer deposited between two coatings of electrospun fibers. Cells were then seeded in an attempt to obtain a fully engineered heart valve with layers resembling the native structure of the aortic valve tissue (69). Despite the fact that the resulting valve differed from the native valve in its organization of the ECM (for example, that an anisotropic layer of PGS represented the highly isotropic spongiosa layer), this type of scaffold gave good results when mechanical performance and maintenance of cellular viability was considered (70). Unfortunately, the lack of in vivo translation of this valve to date does not allow inferences on its behavior in a living organism. In a second example, Eslami et al. (71) employed a hydrogel made with a mixture of methacrylated hyaluronic acid and methacrylated gelatin, into which mitral valve interstitial cells were incorporated followed by its incorporation into a PGS-PCL electrospun scaffold. The author's speculated that this approach produces a more favorable environment for the remodeling of the ECM by the cells after in vivo implantation of the valve. Indeed, encasing cells into hydrogels before seeding a scaffold might mitigate the matrix digestion/remodeling activity of the cells and at the same time would favor the de novo deposition of matrix without affecting the mechanical function of the PGS-PCL layer. This may be particularly interesting considering that the behavior of the cells (and the resultant activation status) can be potentially modulated by mechanical tuning of the hydrogels characteristics, thereby crucially contributing to maintain them in a quiescent/self-renewing phenotype.

Numerous efforts toward the production of valve scaffolds with anisotropic mechanical characteristics have been made using novel fabrication techniques. For example, in a recent study by Wunner et al. (72) a polymeric scaffold with a highly controlled microarchitecture was manufactured using an “electrowriting” technique, which involves high voltage guided printing of a solvent-free, melted polymer onto a laterally sliding aluminum collector. Using this approach, the authors were able to orientate the polymeric fibers (made with medical grade PCL) to mimic that of the collagen and elastin fibers of the natural valve, resulting in a mechanical behavior comparable to that of native valve leaflet (73). Another method that has been exploited with the same aim has been developed by Moreira et al. (74) who introduced textile reinforcements into a scaffold containing a thin valve electrospun layer and fibrin cell-laden gel to confer anisotropic resistance against the forces acting in the valve motion cycle, which resembled the arrangement of collagen bundles of the native fibrosa layer. Preliminary bench testing of the resulting valve after 21 days dynamic conditioning, proved that mechanical stimulation enhanced matrix deposition (in particular collagen) by the cells, thus showing the versatility of a “mixed” fabrication approach to elaborate a design that more closely resembled the natural valve architecture. The utility of a tailored deposition of valve scaffolds fibers is, however, still under question, especially from the perspective of long-term scaffold remodeling after in vivo implantation. In fact, for example, Uiterwijk et al. (75) showed that the orientation of the collagen fibers in in situ TEHVs that were manufactured with isotropic or anisotropic fibers deposition and implanted for 1 year in sheep did not resemble the original arrangement of the fibers in the scaffold, but was instead dictated by the prevailing mechanical forces after implantation. While the authors concluded that the fiber's anisotropic deposition is insufficient to dictate the way the scaffold-populating cells mechanically adapt and deposit new matrix components, it has been discussed that other factors, such as the relatively rapid degradability of the scaffold and inflammatory response, may contribute to override any instructive signals provided by the original geometry of the implant (76).

Finally, a promising technology that is still in its infancy but will undoubtedly provide a decisive future impact in cardiovascular medicine, is that of three-dimensional (3D) printing of valve scaffolds or direct bioprinting of valve leaflets using cell-laden polymers (77). The advantage of this manufacturing technology is that 3D printing/bioprinting allows deposition of matrix components with precise patterning and also exploits a layer-by-layer positioning of materials and cells with a high control of the output geometry (with μm accuracy). Potentially, given the possibility to translate the actual geometry of valves via imaging system data (i.e., CT-scan), this technique could be used for personalized manufacturing of valves tailored to the individual patient with maximal hemodynamic performance and adaptability (78). Despite these advantages, 3D printable materials or, in particular, bio-printable hydrogel/cellular mixtures still suffer from the lack of mechanical strength (79–81), especially with respect to the need to withstand an intense mechanical workload. In keeping with this conclusion, for example, the collagen Type I bioprinted valve described by Lee and co-authors (77) was efficiently cellularized with human endothelial cells, but its mechanical characteristics were insufficient to meet the standards required for mitral and aortic valves in vivo. While other valve-specific printable polymers are currently undergoing investigation, including methacrylated hydrogels, such as gelatin and hyaluronic acids (82), and PEG-DA hydrogels (83), further work is necessary to improve the mechanical characteristics of 3D/bioprinted valves to produce realistic alternatives.