Calcification of the valve leaflet, endocarditis, and congenital valve anomalies are the main causes of heart valve dysfunction [1]. In pulmonary and systemic circulations, valves open in such a way that does not prevent flow to the ventricles under systolic pressure. Aortic valves (AVs) are exposed to shear stresses with high mechanical values due to the blood flowing through them in each cardiac cycle [2, 3]. All four types of heart valves are vulnerable to valve disease, but the most common are the AV and mitral valve (MV). Disorganization of the valve interstitial cells (VICs) and disruption of the ECM layer stratification are characteristics of diseased heart valves [4]. These structural defects have an impact on the function of heart valves. They become stenotic, restricting unilateral blood flow, affecting the function of the heart valves, and causing backflow of blood.

Anatomical, biomechanical, and structural properties of heart valves must be well understood in tissue engineering applications to mimic their function and physiology [5]. The heart is one of the most physiologically active organs in the human body. It nourishes every cell and removes waste products from metabolism by pumping blood that is rich in nutrients and oxygen in unidirectional throughout the body. There are four major chambers in the heart: The right and left atria are the upper chambers, and the right and left ventricles are the lower chambers [6]. The right ventricle receives blood from the atrium through the tricuspid valve. The tricuspid valve, also known as the pulmonary valve, regulates the flow of blood from the right ventricle into the pulmonary arteries and then into the lungs, where it is oxygenated again. Nutrient-rich blood enters the left atrium by the four pulmonary veins and exits the left ventricle via the left MV, also known as the bicuspid valve. The left ventricle pumps nutrient-rich blood through the AV (tricuspid valve) to the aorta and then to the whole body [7]. The heart valve opens and closes approximately 40 million times annually, and more than three billion times during a normal life span of 75 years [8]. The heart valve is an essential dynamic functional component that ensures unidirectional and unimpeded blood flow. The complexity of the macro- and micro-structure of heart valves plays a major role in their ability to fulfill their important properties [9]. The cusps, commissures, and supporting structures in the aortic and pulmonary roots are crucial components of the trileaflet semilunar valves, which include the aortic and pulmonary valves. The main components of atrioventricular valves (tricuspid and MVs) are the leaflets, commissures, annulus, chordae tendineae, papillary muscles, and arterial and ventricular myocardium [10, 11]. The heart valve leaflets are exposed to mechanical stress in the forms of shear, bending, and tension [12]. Each heart valve has unique anatomical features to promote optimal performance, even though valves and atrioventricular valves have comparable microstructures.

The AV consists of three half-moon leaflets enclosed in a connective tissue sheath [13]. Heart valve leaflet tissue is complex and multilayered, approximately 300 to 700 µm thick. The AV consists of three semilunar leaflets contained within a connective tissue sheath. In cross-section, the leaflet has three distinct layers: fibrosa (~45%), spongiosa (~35%) and ventricularis (~20%) [14]. In particular, the biomechanical function is met by fibrosa, spongiosa and ventricularis. The fibrosa, as the first layer, contains dense collagen fibers with endothelial cells lining the surface. Tensile strength is provided by collagen fibers directed toward the periphery when the valve is closed [15, 16]. In addition, the aligned collagen fibers make fibrous the main load-bearing layer. Collagen fibers in valve bend under non-stressful conditions and straighten under pressure. At the closure phase of the valves, the pressure is approximately 80 mmHg. During that time, the valves extend themselves to prevent backflow leakage. Collagen matrix receives the greatest stress in this phase [17]. Collagen shows high resistance to this elevated pressure and, therefore, is the essential component of the heart in mechanical performance.

The middle layer of spongiosa, consisting of a fiberless gel-like layer, present in the middle of the leaflet. Also, proteoglycans, hydrated glycosaminoglycans (GAGs), VICs, and valvular endothelial cells (VECs) [18, 19] are primary components [20] of spongiosa. These properties produce heterogeneous mechanical forces upon the cells, while providing strength and flexibility in the leaflets upon tension. The last layer is ventricularis, which contains a dense layer of collagen and elastic fibers, lined with endothelial cells [21, 22]. The radially aligned elastin and collagen fibers provide flexibility in leaflet formation. Therefore, it can be deduced that the composition and alignment of collagen and elastin fibers play an important role in the strength and biomechanical behavior of heart valves [23].

The primary design objective for heart valve tissue engineering scaffolds is to mimic the structure and material characteristics of the natural AV, especially its anisotropy (Figure 1). It is crucial for developing tissue engineering constructs that replicate the structure of a natural leaflet in order to develop a heart valve. To enhance the anistropic feature of native mimicking structure, the electric conductive carbon fibers were previously used for cardiac tissue engineering [24]. Numerous biomedical fields, such as tissue engineering [25], drug delivery [26], wound healing [27], and regenerative medicine [28], have advanced significantly as a result of the versatility of electrospun polymer nanofibers. Recent advancements in electrospun heart valve scaffolds, in particular, scaffold types, cell types, electrospinning settings of aligned nanofibers, and tissue engineering applications are covered in this study.

The development of a scaffold that offers physiological support for cell adhesion, proliferation, and growth is necessary for heart valve tissue engineering. A heart valve’s architecture consists of two outer laminar anisotropic fiber layers layered on top of a spongy inner layer [30]. In an attempt to replicate the structure of the natural heart valve, several scaffold designs have been tested and proposed. For this objective, two primary scaffold types have been developed: (i) native heart valve scaffolds that have been decellularized (acellular) from allogeneic/xenogenic sources, and (ii) completely artificial scaffolds composed of natural (biological) and synthetic polymers [31].

Polymer solutions have been electrospun effectively, notably synthetic ones that dissolve readily in organic solvents, to produce random and aligned fiber mats to promote heart valve cell proliferation (Table 3). Electrospinning is a highly adaptable and effective method for producing materials that can be scaled up with ease. Continuous production of fibers with micron or nanometer dimensions from a variety of polymeric materials is now feasible thanks to electrospinning [115, 116]. Through the optimization of the solvent system and manufacturing conditions, a wide variety of synthetic and natural biopolymers are employed in electrospinning [117]. Biopolymers such as collagen, chitosan, gelatin, fibrinogen, chitin, HA, and silk are examples of biopolymers that are electrospun. Despite their potential, only a limited number of these polymers have been utilized in producing heart valves. Mechanical anisotropy can be achieved by manipulating the alignment of electrospun microfibres [118]. Electrospinning is an adaptable and promising method for producing scaffolds in tissue engineering. In electrospinning, a high-voltage power supply, a collector, and a syringe pump are the basic requirements of the electrospinning apparatus (Figure 2) [119]. A force transfers the polymeric solution in a syringe to the needle, causing the conical protrusion to produce fibers (Figure 2a). 3D configuration of scaffolds is crucial in the manufacturing of heart valves, which may require the development of specialized tools like innovative collectors for electrospinning machinery [120]. For example, the motorized mandrels, which rotate at high speed could align fibers during deposition (Figure 2b). Furthermore, 3D printing of molds and subsequent metalization could offer a novel way of alignment as recently studied in research group (Figure 2c) [121]. The fiber is accelerated and collected on a grounded target [122]. A variety of microstructures can be produced through electrospinning, depending on the parameters selected. These include polymer type, solvent, flow rate, physical distance between the target and collector, voltage difference, and rotational speed of the mandrel [123]. The 3D configuration of scaffolds is crucial in the manufacturing of heart valves, which may require the development of specialized tools like innovative collectors for electrospinning machinery [124].

Anisotropy describes the way physical properties depend on direction and is a crucial feature of specific tissues such as cartilage, muscle, tendon, and ligament [134]. ECM can optimize its function along the direction of anisotropy. For instance, these special structures can offer efficient contractility and force transfer to support the regeneration of healthy muscle fibers. Structurally and mechanically anisotropic scaffolds are necessary to mimic the properties of the target tissue and provide a distinct microenvironment for cell growth. By controlling the alignment and arrangement of the electrospun nanofibers, many 3D electrospun nanofibers with the essential anisotropy features have been established [135–137] (Table 3). For instance, the myocardium, which is the primary tissue involved in systole and diastole, has a hierarchical design with aligned cells established in an anisotropic and micropatterned ECM [138]. Heart valve tissue engineering is a promising approach for the repair of damaged or diseased heart tissue. In a study by Wu et al. [139], an electrospun fiber network with heterogeneous and anisotropic properties was formed by textile methods to replicate the natural properties of heart tissue. When combined with a hydrogel made from methacrylated HA and GelMA that contains cells, this composite scaffold demonstrated superior mechanical properties, improved cell growth, and consistent ECM remodeling, showing resistance to the degradation and shrinkage in comparison to individual materials [139]. The arrangement of cells in 3D nanofibre/cell constructs, achieved through layering enabled by nanofibres, can be hindered by the small inter-fibre pore sizes in electrospun nanofibre scaffolds, which restricts cell movement between layers and slows down the development of a cohesive tissue structure [140]. To overcome this challenge, efforts are being undertaken to enhance cell penetration through nanofibre scaffolds by changing the density distribution and spatial arrangement of the fibers during electrospinning.

A grounded conductive surface is typically utilized to collect the electrospun nanofibres during the electrospinning process. By adjusting the electric field distribution during the fiber collection process, a special spatial arrangement of the fibers can be formed to mimic the anisotropy of the native ECM. For instance, the orientation of fibers accumulated on a revolving mandrel is parallel to the rotational direction. Additionally, fiber deposition can be influenced appropriately to give distinctive fiber organization within the electrospun scaffolds by interfering with the electric field distribution utilizing a patterned collection surface [141, 142]. Research has demonstrated that these patterned fiber constructions can improve mechanical strength and cause noticeable biological reactions on their surface [143]. However, it is still uncertain if the electrical field’s influence on fiber configuration can actually increase the distance between fibers enough to allow for cell infiltration. Therefore, the main goal of these type of research is to find whether periodic changes to the organization and distribution of fibers within electrospun nanofiber constructs can result in high cell penetration and, as a result, make it easier to create incorporated tissue constructs using nanofiber-enabled cell layering technology [144]. In one study, a combination of PCL and collagen was utilized to enhance cell adhesion and proliferation. The mesh-patterned nanofiber scaffolds’ shape and pore size were assessed under a microscope, and the number of cells that passed through them was used to count the degree of cell infiltration. The results showed that the random, aligned, and patterned nanofiber scaffolds had respective values of 2.2 µm, 4.1 µm, and 35.4 µm. The highest cell penetration is observed in the patterned nanofiber structure [145]. It was thought that the control of proper cell phenotype for the development of functional tissue depended significantly on the restoration of the fundamental morphological characteristics of the native ECM. The arrangement of the collector determines whether the electrospun fibers in the tissue scaffold are anisotropic or isotropic, depending considerably on how the electric field is distributed between the grounded surface and the syringe tip [146, 147]. Cell morphology provides knowledge about the condition and behavior of cells. According to reports, nanofibers can alter the cytoskeleton’s organization and cell shape. In a study, normal human oral keratinocytes were used to electrospin silk fiber nanofibres and examine the impact of cell spreading. SEM pictures revealed that the cells formed a 3D network of the nanofibrous structure by interacting and integrating well with the surrounding fibroin nanofibers, and indicated growth in the direction of fiber orientation [148]. In an another study, fibroblasts grown on polyamide nanofibres were shown to have a longer shape than those grown on a glass surface, which led to a significant difference in actin organization and cell spreading [149]. The arrangement of cells is a biological process in various tissues across the body, and the structural organization of the ECM significantly influences cellular behavior and functionality in organisms.

Despite advancements, production of a tissue substitute that mimics the molecular structure and architecture of the ECM to guide cell orientation remains a challenge. It has been widely reported that aligned nanofibres can guide cell morphology and direct cell orientation. Similarly, scaffolds serve as a structural template that facilitates the 3D arrangement of cells. One study showed that both aligned PCL and PCL/gelatin nanofibres were able to direct the orientation of rabbit cardiomyocytes on scaffolds [150]. However, compared to aligned PCL fibers, aligned PCL/gelatin nanofibers generated higher cell alignment, suggesting that scaffold topography, chemical and biological signals, and scaffold components all affect cell morphology.

In order to produce tissues that are structurally and functionally similar to their natural counterparts, they need also provide cells with the appropriate signals. It is generally acknowledged that the cellular architecture and ECM of natural tissues are important factors in defining their mechanical and biological function [151]. Moreover, most studies focused on the impact of nanofibers on cell behavior have utilized 2D cultures on nanofiber mesh surfaces, which significantly differs from the natural conditions where cells are integrated into a complex and fluid 3D fibrous network of the ECM. However, little work has been done to explicitly examine how the configuration of ECM fibers affects the behavior of cells that produce heart valve tissue because there is a lack of suitable model systems.

There are few studies on the electrical conductivity of scaffolds in the field of heart valve tissue engineering. Electrical conductivity was identified as a critical factor influencing cell regeneration. The electrical conductivity and mechanical contractility support cell adhesion, migration, differentiation, and proliferation [152]. For cardiac tissue engineering, a scaffold should be electrically conductive, mechanically stable, and have elasticity comparable to that of the native myocardium [153]. The electrical conductivity of the scaffold can be increased by functionalizing various materials that are frequently used for scaffold construction, such as collagen, silk, alginate, and chitosan, with nanostructures like graphene or carbon nanotubes [154]. Martinelli et al. [155] studied the behavior of newborn rat cardiomyocytes by depositing pure carbon nanotubes on glass surfaces. They found that there was increased proliferation and tight contact formation among the cardiomyocytes [155]. In another work, scientists developed an electrically conductive nanoscale scaffold by adding carbon nanotubes to electrospun PCL. They successfully differentiated human mesenchymal stem cells into cardiomyocytes by using electrical stimulation. In the presence of carbon nanotubes, cardiac markers such as Nkx-2.5, Cx43, GATA-4, and cardiac troponin T expressions were upregulated and the cell shape became elongated [156]. In a study by Mawad and colleagues [157], the conductive scaffold immobilized by PLA. Experiments have shown that PLA/polyaniline (PANI) nanofibrous sheets exhibit similar cell viability and proliferation to PLA nanofibrous sheets. However, they showed that the synchronized beating rate of cardiomyocytes produced on PLA/PANI nanofibrous sheets was significantly higher than that of PLA nanofibrous sheets [157].

Studies on conductive scaffolds have generally indicated encouraging findings and opportunities regarding maturation of cardiomyocytes cells and production of coordinated beats. The potential of injectable hydrogel scaffolds for myocardial repair following myocardial infarction was highlighted by Mihardja et al. [158] using polypyrrole (PPy) and alginate. The findings demonstrate that PPy and alginate together improve cell proliferation and adherence while facilitating myofibroblast penetration into the infarct region [158, 159]. Saghebasl et al. [160] developed four distinct electrospun scaffold types based on poly(glycerol sebacate)-polyurethane (PGU), PGU-soybean oil (Soy), gelatin (G) Soy/G, and simvastatin-loaded PGU-Soy/G scaffold as an effective in vitro scaffold to provide the ideal scaffold for cardiac regeneration. Their findings demonstrated that the electrical conductivity of nanofibrous scaffolds was enhanced by the addition of soy oil, a semiconducting substance. Furthermore, research showed that the simvastatin-loaded polymeric system improved the proliferation and attachment of cardiomyoblasts and could potentially be used as a drug release carrier in cardiac tissue engineering. In another study, an electrically conductive PPy-chitosan hydrogel was developed by Mihic et al. [161] conjugating the conductive PPy onto chitosan side chains. When injected into rat hearts following myocardial infarction, the conductive hydrogel greatly improved heart function and was able to coordinate myocyte function ex vivo.

A conductive polymer that was designed for cardiac tissue repair is a hydrogel made of gelatin, aniline pentamer (AP), and glutathione (gelatin/AP-GSH) composite as shown by Li et al. [162]. Due to the antioxidant properties of glutathione (GSH), this construct may also lower local reactive oxygen species concentrations, which may improve the results of heart tissue healing. Based on the findings, the gelatin/AP-GSH scaffold’s conductivity was comparable to that of the native myocardium, ranging from 3.4 × 10⁻⁵ S/cm to 1 × 10⁻⁴ S/cm. According to these findings, gelatin/AP-GSH composite scaffolds showed potential in cardiac tissue engineering [162, 163]. The mechanical and electrical properties of PU/chitosan/CNT electrospun scaffolds were compared by Ahmadi et al. [164] by electrospraying of carbon nanotubes by (PU/Chitosan/CNT.sp) or blending (PU/chitosan/CNT.bl). The produced PU/chitosan/CNT composite nanofibrous scaffolds were shown to be electroconductive, and aligned nanofibers may be regarded as potential scaffolds with nanoscale properties for infarcted myocardial regeneration [164]. By mixing the conductive polymer PANI with poly(ethylene glycol) diacrylate (PEGDA), Haq et al. [165] produced hydrogels and a conductive ink for 3D printing. Their major objective was to use 3D printing projection micro-stereolithography to generate electrically conductive scaffolds that, from the perspective of technology, resemble the hierarchical orientation of heart myofibers. The resulting scaffolds had an average pore size of 300 nm and showed semi-conducting characteristics (~10^–6 S/m). Scaffolds with conductive properties exhibit adjustable conductivity and create an ideal setting for the growth of mouse cardiac progenitor cells.

The crucial concerns in the field of cardiac tissue engineering require further research on the application of electrically conductive nanostructured materials. All things considered, research on conductive scaffolds has produced encouraging findings and opportunities for applying these materials to produce mature cardiomyocyte cells and coordinated beats. In spite of significant efforts in cardiac tissue engineering, there is still a lack of conclusive materials, methodologies, and cellular resources, leading to ongoing research, particularly because in vivo studies have been limited.

Electrospinning is regarded as a beneficial technique for developing scaffolds in the field of tissue engineering, as it is simple, low-cost, adaptable, and capacity to form structures that mimic the ECM. A significant number of studies confirmed electrospun scaffolds were more similar to the ECM nanostructure. It has been applied in a variety of techniques to incorporate different morphological features with material properties for tissue engineering such as aligned nanofibers. By far, the post-potent technique was to employ patterned collectors which could orient the electric field effectively. In construction of heart valves aligned nanofibers proved the best approach to mimic anisotropy of valve leaflets. By mimicking the native heart valve anisotropic structure, they influence cell orientation and enhance tissue function. It was also observed that aligned nanofiber’s mechanical strength and anisotropic structure promote cell development. However, it is important to consider the constraints that come with using aligned nanofibers. Among the difficulties that must be taken into account while producing aligned nanofiber-based heart valve applications are reproducibility and scalability. Furthermore, the biocompatibility, hemocompatibility, and degradation characteristics of the materials chosen for electrospinning-based heart valve construction should be carefully considered. It was also concluded that nanofibers itself may not perform physiological functions arising from the lack of soft tissue properties. Some of the studies therefore focused on using nanofiber mats within hydrogel matrix to increase cell infiltration and provide more vascular structure. Another significant challenge detected in using nanofiber scaffolds was cell permissiveness between leaflets. Therefore, it would be essential to apply novel engineered collectors to adjust the inter-fiber distances and hence increase the infiltration rate of heart valve cells.