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Ciolacu, D.E.; Nicu, R.; Ciolacu, F. Natural Polymers in Heart Valve Tissue Engineering. Encyclopedia. Available online: https://encyclopedia.pub/entry/47050 (accessed on 20 June 2024).
Ciolacu DE, Nicu R, Ciolacu F. Natural Polymers in Heart Valve Tissue Engineering. Encyclopedia. Available at: https://encyclopedia.pub/entry/47050. Accessed June 20, 2024.
Ciolacu, Diana Elena, Raluca Nicu, Florin Ciolacu. "Natural Polymers in Heart Valve Tissue Engineering" Encyclopedia, https://encyclopedia.pub/entry/47050 (accessed June 20, 2024).
Ciolacu, D.E., Nicu, R., & Ciolacu, F. (2023, July 20). Natural Polymers in Heart Valve Tissue Engineering. In Encyclopedia. https://encyclopedia.pub/entry/47050
Ciolacu, Diana Elena, et al. "Natural Polymers in Heart Valve Tissue Engineering." Encyclopedia. Web. 20 July, 2023.
Natural Polymers in Heart Valve Tissue Engineering
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The new generation of heart valves developed by tissue engineering has the ability to repair, reshape and regenerate cardiac tissue. Achieving a sustainable and functional tissue-engineered heart valve (TEHV) requires deep understanding of the complex interactions that occur among valve cells, the extracellular matrix (ECM) and the mechanical environment.

heart valve tissue engineering polysaccharides proteins scaffold heart valve replacement regenerative medicine

1. Introduction

Cardiovascular diseases are the leading cause of death globally, and among them, aortic stenosis, determined by the calcification of a trileaflet valve (degenerative calcific aortic valve stenosis, CAVS) or stenosis of a congenital bicuspid valve (congenital bicuspid aortic valve, CBAV), is the most prevalent form of cardiovascular disease in the world, after hypertension and coronary artery disease [1][2][3].
The common treatment for heart valve disease is surgical replacement, but because of the lack of organ donors, alternative approaches are essential for restoring the cardiac function after a heart attack. Surgical replacement of diseased heart valves has been widely performed, primarily with mechanical valves and bioprosthetic heart valves. All these devices have significant limitations with risks of further morbidity and mortality: mechanical valves may cause hemorrhage and thromboembolism, and thus, they require lifelong anticoagulation treatment; bioprosthetic valves have relatively poor long-term durability because of degeneration, calcification and fibrosis, and may cause immunogenic complications [4][5][6].
These difficulties have motivated the development of tissue engineering strategies for valve substitution, which are intended to achieve valve replacements that are based on a three-dimensional (3D) structure capable of supporting cell proliferation, differentiation and growth (in vitro or in vivo) in a functional tissue construct [7]. The main function of heart valves (HV) is to maintain unidirectional blood flow during cardiac systole and diastole, knowing that the normal heart valves open and close about 4 million times a year without obstruction or regurgitation [8][9]. Thus, heart valve tissue engineering (HVTE) requires complex substrate geometries to provide for optimal opening and closing behavior of the valve leaflets [10].
Over the past few decades, several studies have been performed to clarify the desirable characteristics of tissue-engineered heart valves (TEHV) and to develop strategies for generating these valve substitutes [11][12][13][14].

2. Scaffolds for Tissue Engineering: General Concepts

Tissue engineering is a rapidly advancing field in regenerative medicine, with various research papers directed toward the production of new biomaterial scaffolds with tailored properties which can be used to restore, maintain or improve damaged tissues or even whole organs [15]. A key concept in tissue engineering is to restore and improve the function of the tissues by preparing porous three-dimensional scaffolds, and seeding them with cells and growth factors. These three things, i.e., scaffolds, cells and growth factors, are known as “the tissue-engineering triad” and this system is set up in an appropriate environment in a bioreactor [16].
One of the most important entities to be considered for efficient tissue engineering is the scaffold, because its external geometry, surface properties, pore density and size, interface adherence, biocompatibility, degradation and mechanical properties affect not only the generation of the tissue construct in vitro, but also its post-implantation viability and functionality [17][18].
Multiple scaffolds have been designed, developed and tested, and thus, nowadays, different types of scaffolds are available in the field of tissue engineering. In general, these can be classified into two main groups: (i) acellular scaffolds, such as decellularized human or animal tissue, and (ii) artificial scaffolds, fabricated from natural or synthetic polymers and composites.
Acellular scaffolds are the ideal bio-scaffolds necessary to guide host or donor cells toward the regeneration of new and functional tissues and are obtained upon the removal of nuclear content and cellular elements, the scaffolds retaining the architecture and complexity of the native tissues, including vasculature and bio-factors present in the extracellular matrix (ECM) [19]. The obtained acellular or decellularized matrices slowly degrade upon implantation and are generally replaced by the ECM proteins secreted by the in growing cells. The advantages of these scaffolds lie in the removal of all foreign cells and immunogenic compounds, and the retention of their correct anatomical structure and the similar bio-mechanical properties to those of native tissues (such as signaling for cell adhesion and induction of cell migration, proliferation and differentiation), which are critical for the long-term functionality of the grafts [20]. Acellular tissues are biocompatible and the absence of rejection after allogeneic or xenogeneic transplantation makes them the ideal scaffolds for translational medicine applications and organ replacement [21][22]. The obvious advantage of this scaffold is that it is composed of ECM proteins typically found in the body. Naturally derived materials and acellular tissue matrices have the potential advantage of biological recognition. Polymer coating of a tissue-derived acellular scaffold can improve the mechanical stability and enhance the hemocompatibility of the protein matrix.
The decellularization process consists of removing the cellular material from the ECM of biological tissues, leading to a semiporous scaffold (remaining ECM), minimizing damage to the original structure and maintaining the same complex geometry of the native tissue. The scaffold obtained contains natural components (collagen, elastin and glycosaminoglycans) that provide clues for cell migration and differentiation, resulting in a constructive remodeling.
Decellularized heart valves have been more clinically relevant than polymeric valves, due to (i) their positive answer regarding cell differentiation (natural components that can positively impact cell differentiation), (ii) the remodeling process, when these serve as building blocks, (iii) maintaining the mechanical anisotropy of the native valves and, furthermore, (iv) they do not necessitate complete biodegradation. However, decellularized heart valves require human or animal tissue for manufacture, which is limited in supply, and necessitates cryopreservation for storage. The successful use of decellularized heart valves depends on the decellularization process and on the immune response following implantation. The freeze-drying method of biologic heart valves has been used to facilitate long-term storage. Unfortunately, certain limitations of this method have been found, specifically, the collapse of the ECM structure and disruption of biomolecules during the freeze-drying process. To overcome these limitations, the use of lycoprotectants has been proposed [23].
The search for alternative solutions to replace acellular scaffolds leads the research toward the scaffolds fabricated from polymeric materials, which can be categorized into porous, microsphere, hydrogel and fibrous scaffolds.
Porous scaffolds are a 3D structure with an interconnected homogeneous pore network, providing a continuous flow of nutrients and metabolic waste to enable growth and vascularization of engineered tissues. Porous scaffolds can be manufactured using biopolymers with a specific surface-area-to-volume ratio, crystallinity, pore size and porosity [20]. The preparation techniques can be divided into two categories: (i) non-designed manufacturing techniques, which include freeze drying or emulsion freezing, melt molding, phase separation, solvent casting or particulate leaching, gas foaming or high-pressure processing, electrospinning and combinations of these techniques, and (ii) designed manufacturing techniques, which includes rapid prototyping and 3D printing [24]. Generally, conventional fabrication techniques do not enable precise control of internal scaffold architecture (pore size, pore geometry, pore interconnectivity, spatial distribution of pores and construction of internal channels within the scaffold) or the fabrication of complex architectures that could be achieved by rapid prototyping techniques, for example [25]. Rapid prototyping (RP), generally known as solid free-form fabrication or additive manufacturing, is a group of advanced manufacturing processes in which objects can be built layer by layer in additive manner directly from computer data, such as computer-aided design (CAD), computed tomography (CT) and magnetic resonance imaging (MRI) data [26]. Recently, 3D printing has emerged as a promising technology for fabricating geometrically defined porous architectures in 3D, thereby efficiently improving the physiological relevance of tissues and overcoming the significant limitations of various scaffold-based approaches [27]. Regarding the design of 3D printed porous scaffolds that simulate tissues, some properties to keep in mind are: surface area and interconnectivity, which are related to cell growth; permeability, which governs nutrient transport; and mechanical strength, which assures support and protection, among other properties [28]. The most commonly used approaches in developing 3D printed models include selective laser sintering (SLS), fused deposition modeling (FDM), inkjet printing (IJP), multi-jet modeling (MJM), extrusion-based approach and laser-based stereolithography (SL) [29][30]. Porous scaffolds exist in different forms, such as sponge, foam, mesh and nano- and microscale biodegradable fibers; the last two types can indeed be categorized under fibrous scaffolds [31]. Within this category of scaffolds, sponge or foam porous scaffolds have been used in tissue engineering applications [20].
An ideal porous scaffold in heart valve tissue engineering should exhibit a native extracellular matrix (ECM) texture to support repair and regeneration processes. The tissue-engineered valve scaffolds obtained by the conventional techniques, such as particulate leaching, solvent casting, gas foaming, vacuum drying, thermally induced phase separation, melt molding, high internal phase emulsion and microfabrication [32][33], have pores with irregular sizes, which are not interconnected, and more importantly, lack features such as shape and elastomeric flexibility. Recently, in order to create anatomic models, the scaffolds have been prepared by using computer-controlled tools for layer-by-layer deposition of materials or 3D printing [34]. With the advancement of 3D printing technique, a heterogeneous 3D scaffold with strong mechanical strength and with all required characteristics of an ideal scaffold for cardiac tissue engineering, such as the morphology and accuracy of native ECM, can be fabricated. In order to develop the scaffolds intended for heart valve engineering, a bioink composed of cells and desired biomaterials is used to print the specific shape of the organ. Three-dimensional printing-based applications of tissue engineering in combination with stem cell technology have the potential to address the shortage of donor organs for transplantation and provide patient-specific tissue replacement [35].
Microsphere scaffolds are increasingly used as drug delivery systems and in advanced tissue engineering applications such as gene therapy, antibiotic treatment of infected bone and so forth [36]. Regarding the methods used to fabricate microspheres, these are the emulsion-solvent extraction method, precision particle fabrication (PPF) and thermally induced phase separation (TIPS), while the methods used to produce microsphere-based scaffolds as a single macroscopic unit are: (i) heat sintering, (ii) solvent-based sintering (solvent vapor sintering and weak solvent sintering), (iii) subcritical CO2 sintering and (iv) selective laser sintering (SLS) [37]. Microspheres as building blocks have various benefits, such as simple method of preparation, controlled morphology and physico-chemical characteristics and controlled release of encapsulated factors [20]. Densely packed microsphere-based porous scaffolds can both serve as a template for cell proliferation and act as a guide for establishing intricate cell–cell/cell–ECM connections, which permits their utilization in regenerative engineering.
In cardiac tissue engineering, an important challenge is the design of myocardium, which must be highly porous to allow the nutrients’ passage to the cells and to enable formation of aligned and electrically interconnected cardiomyocytes. The spherical nature of microspheres permits a dense packing in regular arrangements, which can be tailored to meet the specific tissue requirements [37].
Hydrogel scaffolds. Over the past decades, an increasing demand for scaffolds to guide the growth of new tissues has led to the development of new strategies for the production of hydrogels with applications in the revolutionary field of tissue engineering. These can be prepared from synthetic or natural polymers, which are physically cross-linked (reversible) or chemically cross-linked (irreversible), and the cross-linking bonds could be covalent or non-covalent (hydrogen bonds, ionic or hydrophobic interactions) [38][39][40]. Hydrogels based on natural polymers have various advantages, such as biocompatibility, cell-controlled degradability and intrinsic cellular interaction, while synthetic polymer-based hydrogels can be prepared with precisely controlled structures and functions [20]. In addition, the combination of natural and synthetic polymers can be used to provide proper scaffold degradation behavior after implantation. Hydrogels are considered biocompatible, due to the structural similarity to the ECM found in tissues, and need specific requirements to function appropriately and promote new tissue formation. These requirements include both physical parameters (in vivo swelling properties, mechanical strength, biodegradation properties), as well as biological performance parameters (cell adhesion and proliferation). Their compatibility with biological tissues, high water content and good mechanical properties make hydrogels particularly attractive for tissue-engineering applications. By adding cells to a hydrogel before the gelling process, these can be distributed homogeneously throughout the resulting scaffold. Fibroblasts, osteoblasts, vascular smooth muscle cells and chondrocytes successfully immobilize and attach to these hydrogel scaffolds [41].
Tissue engineering techniques used three types of hydrogels for cardiac tissue engineering, and those are: (i) natural polymer-based hydrogels, materials derived from a biological source, either animals, plants or algae, such as collagen (COL), fibrin (F), hyaluronic acid (HA), alginate (Alg), gelatin (Gel), chitosan (CH), etc.; (ii) synthetic polymer-based hydrogels, such as poly(ethylene glycol) (PEG), poly(ethylene glycol) diacrylate (PEG-DA), polycaprolactone (PCL), polylactic acid (PLA), poly(lactic-co-glycolic acid) (PLGA), polyacrylamide (PAM), polyurethane (PU), etc.; and (iii) composite hydrogels, which combine the advantages of both synthetic and natural polymers [18][42][43][44][45]. These materials are used to fabricate hydrogel scaffolds that mimic the native ECM and present similar morphology.
Fibrous scaffolds are superior scaffolds in terms of cell adhesion, migration, proliferation and differentiation, due to the high aspect ratio of fibers, growth factor loading efficiency and sustained release capacity. Different techniques are available for preparation of nanofibrous materials, such as electrospinning, self-assembly, phase separation, jet-spraying, jet-spinning, double component electrodeposition and, more recently, melt electro-writing [46][47][48][49][50]. Among these, electrospinning is the most widely used technique and with the most promising results for tissue engineering applications, due to easy handling, applicability to most polymers and cost-effectiveness. The development of nanofibers has enhanced the scope for fabricating scaffolds that can potentially mimic the architecture of natural human tissue at the nanometer scale.
For heart valve tissue engineering, fibrous scaffolds would provide an ideal environment for cells, if they could form 3D structures with porosity, pore size and mechanical characteristics comparable to native heart valves. Various polymers have been used for HVTE, such as polyglycolic acid (PGA), PLGA, PLA, poly L-lactic acid (PLLA), PCL, poly(L-lactic acid-co-ε-caprolactone) (PLCL) and PU as synthetic polymers and COL, Gel, CH and HA as natural polymers [18][49][50][51][52].
Regardless of the scaffold specific properties, a number of key considerations are important when designing or determining the suitability of a scaffold for use in tissue engineering, as described below.
The porous architecture of scaffolds used for tissue engineering should have an interconnected pore structure and adequate mean pore sizes, large enough to ensure cellular penetration and small enough to establish a sufficiently high specific surface [25]. If pores are too small, cell migration is limited, resulting in the formation of a cellular capsule around the edges of the scaffold, which can limit the diffusion of nutrients and the removal of waste, resulting in necrotic regions within the construct [53]. If pores are too large, limited cell adhesion was observed due to a decrease in surface area. Therefore, the critical dimension of pores may vary depending on the cell type used and the tissue being engineered. In addition, the scaffold must allow an adequate diffusion of nutrients to cells and the ECM formed by these cells, as well as the diffusion of waste products out of the scaffold [54].
  • The produced scaffold should have adequate mechanical properties, to mimic the anatomical site where it is intended to be implanted, and to function from the time of implantation to the completion of the remodeling process [55]. A scaffold’s mechanical properties (strength, modulus, toughness and ductility) are determined both by the material properties of the bulk material and by its structure (macrostructure, microstructure and nanostructure). Matching the mechanical properties of a scaffold to the graft is critically important, so that the progression of tissue healing is not limited by its mechanical failure prior to complete tissue regeneration [56]. Many materials have been produced with good mechanical properties, but to the detriment of retaining high porosity. In addition, many of these materials, with demonstrated in vitro potential, have failed when they were implanted in vivo because of insufficient capacity of vascularization [53]. Thus, to achieve a suitable scaffold, it is necessary to balance the mechanical properties with a porous structure, sufficient to allow cell infiltration and vascularization.
  • Interface adherence of the scaffold referred to the interactions between cells and their environment, which play a critical role in determining cell fate and physiological functions, so as to maintain normal phenotypic shape within the scaffold. An ideal scaffold should provide informative microenvironments mimicking physiological niches to direct advanced cell behaviors, such as differentiation, proliferation and apoptosis, without inducing pathological outcomes, such as calcification [4].
  • The scaffold’s biocompatibility is related to the cell’s adherence, which should function normally, migrate onto the surface or even through the scaffold, begin to proliferate and, finally, have a negligible immune reaction. Thus, to be accepted in vivo, the host immune response should be minimal for the scaffold. The biocompatibility of the cross-linking agent used is particularly important, especially in cases where reactive groups of the cross-linker are incorporated into the hydrogel network and might then be released upon degradation. Although unreacted chemicals are usually eliminated after cross-linking through extensive washing in distilled water, as a rule, toxic cross-linkers should be avoided, in order to preserve the biocompatibility of the final scaffold [53].
  • A scaffold should be biodegradable and the degradation products should be non-toxic and able to be eliminated from the body without interference with other organs. There are different mechanisms for in vivo degradation, such as hydrolysis, oxidation, enzymatic and physical degradation [57]. The biodegradation process permits to the cells to produce their own extracellular matrix and finally to replace the implanted or tissue-engineered constructed scaffolds, eliminating the need for further surgery to remove it. The scaffold’s degradation rate should be adjusted to match the rate of tissue regeneration so that it has disappeared completely once the tissue is repaired [58][59].
The advantages and disadvantages of the above presented scaffolds, such as porous, microsphere, hydrogel and fibrous scaffolds, are summarized in Table 1 [18][20][37].
Table 1. Comparative analysis of porous, microsphere, hydrogel and fibrous scaffolds.
Scaffolds intended for heart valve tissue engineering face additional distinct challenges owing to their direct contact with blood. Specifically, the construct should be resistant to calcification, should have a minimal thromboembolism risk and must withstand the unique hemodynamic pressures and flows of the cardiac environment from the moment of implantation [18][35]. Moreover, the scaffold should imitate the natural myocardial ECM and should possess adequate porosity that promotes vascularization.
It should also allow continuous diffusion of oxygen and nutrients to the seeded cells and it must mimic the mechanical properties of the native cardiac tissue and bear the cyclic strains and stresses exerted upon transplantation, and must also be sufficiently thick to contract with proper strength and beat synchronously with the neighboring cardiomyocytes [35][57].
These unique challenges underline the importance of carefully considering the materials and design when fabricating a scaffold for tissue-engineered heart valves.
Ideal heart valve tissue-engineered scaffolds are defined as three-dimension porous solid biomaterials designed to perform some or all of the following functions: (i) promote cell-biomaterial interactions, cell adhesion and ECM deposition, (ii) permit sufficient transport of gases, nutrients and regulatory factors to allow cell survival, proliferation and differentiation, (iii) biodegrade at a controllable rate that approximates the rate of tissue regeneration under the culture conditions of interest and (iv) provoke a minimal degree of inflammation or toxicity in vivo [20]. Scaffolds can be seeded with embryonic or adult stem cells, progenitor cells, mature differentiated cells or co-cultures of cells to induce tissue formation in vitro and in vivo. While the specific functions vary with tissue type and clinical need, scaffolds may potentially coordinate biological events at the molecular, cellular and tissue levels on time and length scales ranging from seconds to weeks and nanometers to centimeters, respectively. A central theme in designing tissue-engineered scaffolds is to understand the correlations between scaffold properties and biological functions [60].

3. Heart Valve Replacements

The heart contains four chambers (two atria and two ventricles) and four valves: (i) the tricuspid valve, serving as blood flow regulator, from the left atrium to the ventricles, (ii) the pulmonary valve, which controls the blood flow from the right ventricle to the pulmonary artery, (iii) the mitral valve, which regulates the blood inflow from the right atrium to the ventricles, and (iv) the aortic valve, having the role of regulating the flow from the left ventricle to the aorta [61].
This research dwells on two valves of the four (aortic valve and pulmonary valve), with a major focus on the aortic valve. Both these valves have similar structures and mechanical characteristics, the differences appearing in the thickness of the layer and its density. As a structure, they consist of three semicircular leaflets (cusps), which are connected to a fibrous annulus (root).
The most common valve disease is the aortic valve stenosis (determined by the calcification and thickening of the cups), which is presented concomitantly with aortic regurgitation (because of the loss of stretch in calcified cups) [62]. Generally, the valve dysfunction is caused by either aging or congenital defects, and the severe complications of these degenerative diseases can seriously affect the structure and function of heart valves. In the short term, medication may be used to improve the health of patients, but for patients with severe valvular pathologies, the best option is to do surgery in order to repair or even replace the valve.
It is well known that the first step toward heart valve replacements is to ensure long-term functionality of implantation and a competent and stable structure with specific anatomical and histological features [63][64].
There are three categories of heart valve replacements, of which the most two common categories are the mechanical and bioprosthetic valves [4]. The development of the polymeric valves was intended to overcome the problems characteristic of the above-mentioned valves, related to regeneration, growth potential and durability.
The classification of heart valve replacements, as well as their main advantages and disadvantages, are presented in Table 2.
Table 2. The classification and the advantages/disadvantages of heart valve replacements.

Types of Valves

Definition

Advantages

Disadvantages

Mechanical valves

-

made entirely from metal, pyrolytic carbon and expanded polytetrafluoroethylene (ePTFE or teflon).

-

limitless supply;

-

lack of structural deterioration.

-

risk of thrombosis;

-

requires anticoagulant drugs for life;

-

not available in small size;

-

possible mismatch with patients.

Bioprosthetic valves

Autograft valves

-

made from another valve within the patient’s own heart (such as the removal of the pulmonary valve to fix the aortic valve).

-

not immunogenic;

-

no risks of thrombosis;

-

growth potential.

-

high probability of replacement after 12 years;

-

difficult to handle.

Allograft valves

(homograft)

-

transplanted within the same species;

-

from a deceased human donor.

-

good hemodynamic profile;

-

preservation of the morphology;

-

no risks of thrombosis;

-

low risk of infection.

-

limited availability;

-

lack of growth potential;

-

decellularization weakens ECM;

-

immunogenic response if decellularization not complete.

Xenograft valves

(heterograft)

-

transplanted from one species to another

-

derived from porcine aortic valve or bovine pericardium, implanted in humans.

-

limitless supply;

-

adequate anatomic structure;

-

optimal biological properties.

-

lack of growth potential;

-

decellularization weakens ECM;

-

immunogenic response, if decellularization not complete.

Polymeric valves

Natural

polymeric scaffolds

-

made by cross-linking, photo-polymerization, pressure casting, injection molding, 3D printing, etc.

-

limitless supply;

-

ease of shaping;

-

polymers combination to meet specific mechanical properties;

-

combination with stem cells to obtain a living graft.

-

degradation by hydrolysis can affect mechanical properties;

-

possible cytotoxicity of degradation products.

Synthetic polymeric scaffolds

Composite polymeric scaffolds

Abbreviations: ECM—extracellular matrix; ePTFE—expanded polytetrafluoroethylene.

4. Heart Valve Tissue Engineering: Cells and Strategies

The highly complex architecture of heart valves includes an extracellular matrix (ECM) populated by valvular interstitial cells (VICs) and encapsulated by valve endothelial cells (VECs). All these are in a continuous reorganization, as a response to the changes during the cardiac cycle.
ECM is an extremely organized network, composed of three closely linked layers, arranged according to the blood flow, with unique properties that vary continuously throughout the cross section of the leaflet[65], namely:
-
The fibrosa layer is located near the outflow surface and is made of collagen (COL) and represents densely aligned fibers that ensure the primary strength of the valves;
-
The ventricularis layer is located on the opposite surface of the entrance and is made of elastin (EL), with an important role in stretching and retraction during the cardiac cycle;
-
The spongiosa layer is located between the two layers mentioned above and is made of proteoglycans (PG)—glycosaminoglycans (GAG), with the role of loose connective tissue to facilitate the relative movements of the adjacent layers.
The quantity, quality and the structure of ECM depend on the viability and function of the VICs, this cell–matrix interaction being determined by a dynamic and complex mechanical stress state during every cardiac cycle [66]. Cell adhesion on the surface of ECM is mediated by the ECM components of the valve leaflet and consist of small amino acid sequences that mediate cell attachment, the most popular being the arginine-glycine-aspartic acid (RGD) domain [67].
While the ECM plays a critical role in the structure–function relationship of the valve, the VICs cells have an important role in preserving its architecture for functional biomechanics and maintaining homeostasis and also have a crucial role in some pathological valve processes. Moreover, the valve cusp is encapsulated by a single cell layer of VECs, which creates a functional barrier between the blood and the inner tissue of the valve, acting as protection against physical and mechanical stress of the hemodynamic environment, and continuously communicating with VICs for regulating their phenotype [65].
The most numerous valvular cell types are VICs, which present particular characteristics and functions depending on the environmental conditions, and can be classified as: embryonic endothelial progenitor cells (eEPCs), quiescent VICs (qVICs), activated VICs (aVICs), progenitor VICs (pVICs) and osteoblastic VICs (obVICs). At different cycles of development, VICs show different phenotypes. In adult heart valves cultured in situ, VICs are quiescent and display a fibroblast-like phenotype, characterized by the presence of vimentin, and very low levels of α-smooth muscle actin (α-SMA), metalloproteinases (MMP-13) and SMemb (non-muscle myosin heavy chain) [66]. In contrast, in heart valves cultured in vitro, 50–80% of VICs isolated express high levels of myofibroblastic markers such as αSMA [68].
Biomechanical and biochemical factors have an important role in VICs response, so that VICs from aortic and mitral valves are more rigid than those from pulmonary and tricuspid valves, which suggests that VICs respond to local tissue stress by altering their stiffness [69]. VICs have a fusiform, ellipsoidal shape and contain a large amount of cytoplasm, rich in mitochondria, rough endoplasmic reticulum and exocytic vesicles. VECs have the role of maintaining a nonthrombogenic blood–tissue interface, for the transport of nutrients, regulating immune and inflammatory reactions and ensuring the transduction of biochemical and mechanical signals in the heart valve. Additionally, they have a cobblestone-like morphology and are aligned perpendicular to the blood flow direction and parallel to the collagen fibers from ECM [70].
Engineering of heart valves is greatly influenced by the type of the used cells, and the three most frequently used cell types TEHV are: xenogeneic (from a different species), allogeneic (same species) and autologous (same living being). If in the case of autologous cells, they have a high activity and are best suited for use in TEHV, allogeneic and xenogeneic cells invoke an immune response and cannot be used without immunosuppressive therapy [71].
However, it has been observed that analogous cell types from animal and human sources showed almost identical phenotypes in TEHV, which led to the possibility of using cells from animal sources in both in vitro and preclinical research. Cells from animal sources have several advantages, such as wide availability, being cheaper and not being subject to the same level of safety and ethical regulations as human cells. Another advantage of animal cells over human cells is the fact that they can be isolated from all four valves of the heart, while in humans, it is usually obtained from a single valve [72]. The cells used in TEHV, both from animal and human sources, are: mesenchymal stem cells, valvular interstitial cells, valvular endothelial cells, endothelial cells, miscellaneous cells and fibroblasts [71].
Starting from various scaffolds (1), including autografts, allografts, xenografts and polymeric scaffolds, and from different types of cells (2), appropriate to maintain and remodel the ECM, several strategies (3) have been established, in order to obtain living tissue valve replacements that can function like the native heart valve.
Regarding the scaffolds used in TEHV, it should be mentioned the fact that allographs and xenografts are pre-processed by decellularization, to ensure immunocompatibility and a standard availability of valve tissues, mostly preserving the integrity and functionality of the ECM. The use of bioresorbable polymers in TEHV has attracted special attention due to the possibility of quickly manufacturing scaffolds with reproducible architectures, with controllable degradation rates and mechanical and chemical properties adapted to the desired purpose [73][74][75]. In addition, these scaffolds have the advantage of being absorbed and metabolized by the body.
In vitro TEHV strategy consists of the incorporation of autologous cells into a bioresorbable scaffold, which can be either biological or polymeric. This cell–scaffold system is usually cultured in a bioreactor to allow ECM deposition and to promote the new tissue synthesis with an adequate elasticity and strength for implantation [76]. To in vitro cellularize the scaffold before implantation, autologous cells are used with a view to preventing an immunogenic response, and these are [77][78][79][80][81][82]:
  • (Myo)fibroblasts isolated from harvested vascular or dermal tissues;
  • Mesenchymal stromal cells (MSCs) from bone marrow or adipose tissue;
  • Prenatal or early postnatal sources of MSCs, where cells are harvested before or immediately after birth and used toward the synthesis of autologous valve tissue for replacement in early childhood;
  • Amniotic membrane sources of MSCs (AM-MSCs);
  • Amniotic fluid sources of MSCs (AF-MSCs);
  • Chorionic villi sources of MSCs (CV-MSCs);
  • Umbilical cord sources of MSCs (UC-MSCs), from the cord blood, Wharton’s jelly or perivascular tissue;
  • Stem cell (iPSC)-derived endocardial cells with the potential to provide VIC-like cells by undergoing endothelial-to-mesenchymal transition, with the best potential to obtain the native VICs population, compared to other mesenchymal cells.
The optimal biological scaffolds, from a geometrical and hemodynamical point of view, are decellularized heart valves (allogenic or xenogenic). However, they also have an important number of negative effects, such as microstructural changes and altered protein composition, as a response to the cryopreservation process or decellularization, the limited availability of human tissue, the residual immunogenicity of animal tissue and the limited cellular infiltration [23]. Both synthetic and natural polymer-based scaffolds offered an attractive solution for TEHVs achievement, due to their unlimited availability, their tunable architectures and mechanical properties, and their inherent lack of xenogeneic disease transmission [83]. In addition, so far, in vitro TEHV has not advanced into routine clinical use.
The in vivo TEHV strategy uses the body’s ability to encapsulate foreign material and use fibroblasts to produce ECM proteins. In this sense, a valve-shaped mold is implanted subcutaneously, and this is covered over time by a fibrous tissue, which is then removed and used as a replacement valve. Even if the method seems accessible, unfortunately, apart from obtaining an adequate geometry of the valves, there is no control over the cells present in the tissue, nor over the ECM composition or its mechanical properties [84].
The in situ TEHV strategy consists of the direct implantation of an acellular resorbable scaffold, which can be either biological or polymeric, and which is designed to induce the potential for endogenous regeneration, directly at the functional site of the valve. In this type of strategy, the scaffold must ensure an optimal environment for the adhesion, differentiation and growth of the host cells, to support the formation of the new tissue, while the controlled degradation of the initial scaffold takes place, and that the newly formed tissue has mechanical properties similar to those of native functional tissue [85]. Moreover, the scaffolds that are used for in situ TEHV may be either newly fabricated (natural or synthetic polymers) or decellularized from native valves or bioreactor grown valves. For synthetic polymers, an alternative method to enable them to mimic native heart valves are their biofunctionalization by the incorporation of peptides, proteins or recognition sequences [84]. In the case when TEHV is grown in vitro and then is decellularized before further implantation, the choice for an autologous cells source is not absolutely necessary, because the tissue produced by allogeneic cells is immunocompatible, if adequately decellularized. Other cell sources may be the UC-MSCs cells and the induced pluripotent stem cells (iPSCs), which offer different advantages, in terms of accessibility, expandability and capacity for tissue synthesis [77]. To date, the in situ TEHV strategy recorded the highest progress, on the basis of different in vivo studies in animals and with delivery of the valve using transcatheter implantation, showing encouraging results [46][86][87].
The most relevant experimental approaches related to in vitro and in situ TEHV, together with cell source, the type of scaffold and the main results of the studies, are presented in Table 3 for the aortic valve replacements and Table 4 for the pulmonary valve replacements.
Table 3. Overview of cell sources and scaffold materials used for aortic valve replacement.
Abbreviations: BPUR—biodegradable poly(ether ester urethane) urea; DNA—deoxyribonucleic acid; ELR—elastin like recombinamer; EOA—effective orifice area; GAG—glycosaminoglycans; Me-Gel—methacrylate gelatin; Me-HA—methacrylated hyaluronic acid; MPG—mean pressure gradient; P4HB—poly-4-Hydroxybutyric acid; PAN—polyacrylonitrile; PCL—polycaprolactone; PEG—poly(ethylene glycol); PEG-DA—poly(ethylene glycol) diacrylate; PET—polyethylene terephthalate; PGA—polyglycolic acid; PLDL—poly(L/D,L-lactide); PLGA—poly(lactic-co-glycolic acid); SMC—aortic root sinus smooth muscle cells; α-SMA—α-smooth muscle actin.
Table 4. Overview of cell sources and scaffold materials used for pulmonary valve replacement.

Abbreviations: CD31—clone JC/70A; CD44—clone G44–26; CD44—fluorescein isothiocyanate [FITC]-conjugated (Clone MEM-85); DNA—deoxyribonucleic acid; eNOS—endothelial nitric oxide synthase type III; GAG—glycosaminoglycans; P4HB—poly-4-hydroxybutyric acid; PCL—polycaprolactone; PGA—polyglycolic acid; PLA—polylactic acid; PLLA—poly L-lactic acid; α-SMA—α-smooth muscle actin.

5. Natural Polymer-Based Scaffolds for Heart Valve Tissue Engineering

Heart valve tissue engineering scaffolds based on natural polymers have the advantage to be prepared from environmentally friendly, renewable and low-cost raw materials, with appealing properties for biomedical applications, such as biocompatibility, biodegradability and intrinsic cellular interaction [20][111][112][113][114]. Although natural polymers provide excellent cell attachment and growth, they have many disadvantages, such as immune response problems or poor mechanical properties [16]. All these will be discussed in detail for each category of natural materials (i.e., polysaccharides and proteins), taking into account different examples for each polymer.

5.1. Polysaccharide-Based Scaffolds for Heart Valve Tissue Engineering

Polysaccharides are the most abundant biomaterials in nature that meet several criteria for eligible supports for tissue engineering, which include biocompatibility, biodegradation and the ability to support cell development [115][116]. Due to their biological properties and their structural and functional similarities to ECM, it is reasonable to use them in tissue engineering [117][118][119]. In combination with appropriate cells or bioactive molecules, the polysaccharides become an important asset to promote heart valve tissue regeneration [120]. Their applications for heart valve tissue engineering are vast and varied, and approximately 70% of all studies in this field focus on chitosan, alginate, hyaluronic acid and cellulose, respectively [117]. Table 5 presents several examples from the multitude of applications in valve engineering, for each of these polysaccharides.
Table 5. Polysaccharide-based scaffolds for tissue-engineered heart valves.

Scaffold Types

Preparation Methods

Results

Ref.

Chitosan-Based Scaffolds

CH films

Casting method to form films;

Adsorption of protein sol. on CH films (4 °C, overnight).

CH films: FI/SMCs is less spread and more elongated;

CH/AP: modest VECs growth, altered elongated morphology, low spreading;

CH/COL IV composites: enhanced VECs growth, superior cell morphology.

[121]

CH/AP

CH/COL

composites

(bFGF-CH-P4HB)/DPAV hybrid scaffolds

Coating DPAV with bFGF-CH-P4HB by electrospinning technique (20 kV, room temp.)

bFGF-CH-P4HB fibers form membranes with uniform thickness, firmly attached on DPAV surface;

bFGF has a positive effect on the MSCs proliferation.

[122]

CH/BP

scaffolds

Immersion of BP tissues in CH/H2CO3 sol. (pH 3, 2 h, 30 MPa, room temp.)

CH/BP are less rigid and the risk factor of fatigue failureis reduced;

Calcification and bacterial strains adhesion are attenuated;

In vivo: no inflammatory reaction, after 4 months of implantation in rats.

[123]

CH-PU-GEL

nanofibrous scaffolds

Electrospinning technique

(16 to 20 kV, room temp.)

OCAs adhered preferentially on CH-GEL-PU, are flattened, spread across the surface and have cobblestone morphology; able to withstand shear-stresses ranging from 0.062 to 0.185 N/m2 for up to 3 h;

[124]

CH fibers with immobilized HEP

Extrusion method;

HEP immobilization with EDC

Crosslinking degree influences fiber diameters, strength and stiffness; CH-HEP promotes VIC attachment and growth (cell viability ~ 95%, 10 days).

[125]

CH-PCL/DBP biohybrid scaffolds

Electrospinning technique

(27–32 °C, 15 kV)

hVICs viability on CH-PCL/DBP (A&R) ~ 90%;

Biohybrid (A) has better uniaxial mechanical properties and higher alignment of hVICs compared to a randomly electrospun sample (B).

[126]

Hyaluronic Acid-Based Scaffolds

Me-HA, Me-HA/PEG-DA hydrogels

Photopolymerization

(UV light, 5 mW/cm2, 3 min, photoinitiator)

Degradation rate: Me-HA/PEG-DA—1 week; Me-HA—2 days;

VICs remain viable following photopolymerization; high proliferation after exposure to LMW HA degradation products.

[127]

(Me-HA+CD34)/Me-Gel

hydrogels

Photopolymerization (UV light, 180 s, 5.5 mW/cm2); CD34 immobilization by EDC/NHS.

Increasing CD34 conc. increases EPC attachment (25.3 ± 5.3 EPCs/mm2 at 10 μg/mL; 52.2 ± 5.0 EPCs/mm2 at 25 μg/mL);

(Me-HA+CD34)/Me-Gel promoted cell elongationand higher spreading.

[128]

SilylHA-CTA/LLDPE IPNs

Silylation of HA-CTA;

LLDPE films swollen in silylHA-CTA/xylene (50 °C/1 h).

HA/LLDPE exhibit lower contact angles and less blood clotting than LLDPE alone, which led to considerable thrombus formation; PHVs showed acceptable values for RF (4.77 ± 0.42%) and EOA (2.34 ± 0.5 cm2).

[129][130]

HA-LLDPE IPNs/CoCr-MP35N stent

Swelling process was used to obtain IPNs; fixing by PP sutures on the stent frame.

Hemodynamic parameters (EOA, RF, PI) have values comparable with those of commercial transcatheter valves;

Turbulent flow tests show a decrease of RSS at each cardiac phase.

[131]

Me-HA/Me-Gel

MOHA/Me-Gel

hybridhydrogels

Molding technique and exposure to UV light

(2 mW/cm2; 5 min)

Me-Gel stimulates VICs spread and migration from spheroids; Cell circularity was much lower in low stiffness hydrogels than in stiffer ones;

VICs have a spindle-like morphology only in hydrogels with Me-Gel.

[132]

Me-HA/Me-Gel/PGS-PCL

hybrid hydrogels

Immersion of electrospun PGS-PCL into hydrogel;

Photocrosslinking

(UV light, 45 s, 2.6 mW/cm2).

MVICs have an initial rounded shape and low spread;

MVICs are predominantly spread over the surface of PGS-PCL fibers only;

21 days: MVICs spread is complete into hybrid hydrogels, with non-homogenous distribution at different depths.

[133]

Cellulose-Based Scaffolds

CA coatings

for metallic valves

Electrospinning technique;

Surface functionalization with RGD and YIGSRG

CA coatings promote cardiac cell growth on valve surface;

CA ensures the control of endothelialization and reduction of thrombosis.

[12]

CNF/PU films

nanocomposites

Film-stacking method;

Compression molding

Prosthetic valves have good biological durability, fatigue resistance and hemodynamics properties;

no failure is registered after accelerated fatigue tests, equivalent of 12-year cycles.

[134]

mNG composite hydrogels

Covalent conjugation of mNCC on Me-Gel backbone via

NHS/EDC crosslinking

Encapsulated HADMSCs on mNG displayed phenotypic properties found within the heart valve spongiosa;

lower expression of osteogenic genes indicates resistance toward calcification.

[135]

BC/PVA

anisotropic

nanocomposites

Physical crosslinking by freeze-thaw cycles

(20 °C/−20 °C);

molding technique

Mechanical properties are similar to valve leaflet tissues, in both principal directions; the composition and number of freeze-thaw cycles substantially influence the tissue properties.

[136][137]

Thermal processing;

molding technique

Trileaflet mechanical heart valve mimics the non-linear mechanical properties and anisotropic behavior of the porcine heart valves.

[138][139]

Alginate-Based Scaffolds

PEG-DA/Alg

hydrogels

Simultaneous 3D printing/photocrosslink ingmethods

The scaffolds with 10% Alg allow PAVICs to grow along the conduits surface, but less on the root and leaflet interstitium;

high cell viability: 91.3 ± 10.7% (day 1) and 100% (day 7 and 21).

[94]

Alg/GEL

hydrogels

3D bioprinting with mold

extrusion technique

Printing accuracy: 84.3 ± 10.9%;

Cell viability (7 days): 81.4 ± 3.4% (SMCs); 83.2 ± 4.0% (VICs);

SMCs express α-SMA in stiff matrix;

VICs express VIM in soft matrix.

[90]

Dop-Alg

hydrogel

coatings

Covalent bonding of Dop to Alg (EDC/NHS route);

Crosslinking with GA

In vitro: only Dop-Alg determines a decrease in the Ca content:

2.919 ± 0.252 mg/L—day 3; 0.725 ± 0.012 mg/L—day 6;

In vivo: the largest decrease in Ca content for Dop-Alg:

1.737 ± 0.124 mg/L—day 20; 0.675 ± 0.084 mg/L—day 30.

[140]

Abbreviations: A&R—aligned and random; ACAN—aggrecan; Alg—alginate; AP—adhesive proteins; BC—bacterial cellulose; bFGF—basic fibroblast growth factor; BP—bovine pericardium; CD34—mouse antibody; Ca—calcium; CA—cellulose acetate; CH—chitosan; CNF—cellulose nanofibrils; COL IV—collagen type IV; CTA—cetyltrimethylammonium bromide; DBP—decellularized bovine pericardium; Dop—dopamine; DPAV—decellularized porcine aortic valve; EC—endothelial cells; EDC—1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide; EOA—effective orifice area; EPC—endothelial progenitor cells; FI—fibroblast; Gel—gelatin; GA—glutaraldehyde; HADMSCs—Human adipose-derived mesenchymal stem cells; HEP—heparin; IPNs—interpenetrated networks; Me-GEL—methacrylated gelatin; Me-HA—methacrylated HA; mNG—mNCC—TEMPO-modified nanocrystalline cellulose; MOHA—methacrylated oxidized HA; MSCs—mesenchymal stem cells; MVICs—mitral valve interstitial cells; NHS—N-hydroxysuccinimide; OCAs—ovine carotid arteries cells; P4HB—poly-4-hydroxybutyrate; PAVICs—porcine aorta valve interstitial cells; PCL—polycaprolactone; PEG-DA—poly(ethylene glycol) diacrylate; PGS-PCL—poly(glycerol sebacate)-polycaprolactone; PHA—polyhydroxyalkanoates; PHVs—polymeric heart valves; PI—pinwheeling index; PP—polypropylene; PU—polyurethane; PVA—poly(vinyl alcohol); RF—regurgitant fraction; RGD—Arginine-Glycine-Aspartate peptides; RSS—Reynolds Shear Stress; SilylHA—silylated HA; α-SMA—α-smooth muscle actin; SMCs—smooth muscle cells; TCPS—tissue culture polystyrene; VICs—valvular interstitial cells; VIM—vimentin; YIGSRG—tyrosine-isoleucine-glycine-serine-arginine-glycine malinins.

5.2. Protein-Based Materials as Scaffolds in Heart Valve Tissue Engineering

Protein-based scaffolds have inherent biocompatibility, degrading under the specific action of enzymes and, thus, allowing cell-controlled tissue remodeling [141]. The most commonly employed proteins that may be suitable for HVTE are collagen and fibrin, but gelatin, elastin, keratin, silk fibrinoid or others may also be mentioned. Certainly, among all proteins, collagen is the most suitable in this regard, being the major component of ECM and providing most of the mechanical and tensile strength of the native valve. Moreover, all collagen types are biodegradable due to their protein nature and are poorly immunogenic mainly due to their homology across species [142]. Table 6 presents some results of relatively recent studies regarding the applications of protein-based scaffolds in HVTE, which will be further discussed in detail.
Table 6. Protein-based scaffolds in heart valve tissue engineering.

Scaffold Types

Preparation Methods

Results

Ref.

Collagen-Based Scaffolds

3D-COL

biological scaffolds

Decellularization by SDS extraction; crosslinking (EDC/NSH); enzymatic treatment to remove elastic fibers.

Mechanical properties of 3D-COL controlled by crosslinking degree;

3T3 cells adhere and proliferate on COL scaffolds and infiltrated to depth of about 20 mm after 7 days, and 40 mm after 28 days.

[143]

COL/NRASMC matrix

Collagen-cell suspension was cast into silicon rubber wells and cultured in an incubator.

Uniform tension, during COL compaction, increases the cell content, stimulates their metabolism and leads to stronger constructs;

NRASMCs are metabolically active proved by the elastin inside and around the COL fibers, and the proteoglycans at their interface.

[144]

3D COL disc scaffolds

Molding technique by rapid prototyping with 3D inkjet printer

VICs proliferate more on 1% w/v COL than 2% or 5%;

VICs remodel the scaffold and synthesize new matrix

(detection of remodeling enzymes, MMPs and ECM gene expression).

[145]

COL-EL; COL-C4S

heterogenous scaffolds

Molding technique using PTFE molds, followed by freeze drying

Good cell proliferation on COL, due to natural cells binding via integrin receptors;

C4S increase the cell metabolic activities;

Low cell proliferation on EL, due to its non-integrinsignaling pathway.

[146]

COL-EL

bilayer scaffolds

Solution casting into PTFE molds; freeze drying, repeated twice to obtain bilayer structure.

Bilayer scaffolds have anisotropic bending moduli similar to native valves CDCs prefer COL over EL when proliferating, resulting in asymmetrical cell distribution in the two different layers.

[147]

3D COL-EL

hydrogels scaffolds

COL-EL composition: 50% COL, 12% EL, 10% PBS, 28% equal parts of DMEM and FBS; pH 7.5; 37°C; 1 h.

3D COL-EL scaffolds support cell attachment, proliferation and differentiation: after 7 days, VICs double their number and exhibited stable levels of integrin β1 and F-actin expression; VECs have a very good proliferation, but the integrin β1 expression remained low.

[148]

3D COL-CH composites scaffolds

COL:CH (7:1, w/w)

The composites were seeded with 3 types of cells: SMCs, FIs and ECs.

3D COL-CH have good cells adhesion and support ECs differentiation;

SMCs group—large number of SMCs with dense disordered arrangement; SMC+EC group: large number of scattered

ECs with long shuttle shape.

[149]

COL-HA hybrid scaffolds

Crosslinking by EDC/NHS route.

Structure similar to fibrosa layer of the valve leaflets;

CDCs attachment not affected by the pore size and stiffness.

[150]

Fibrin-Based Scaffolds

Autologous fibrin-based heart valve scaffolds

Molding technique;

In vitro: bioreactor conditioning;

In vivo: implantation in sheep pulmonary trunk (3 months).

In vitro: well-organized structure of “conditioned samples”, aligned OCAs in leaflets; cellular detachment, possible cells death in “control samples”;

In vivo: fibrin scaffolds completely resorbed and replaced by ECM proteins; significant tissue development and cell distribution.

[102][151]

(SC-F)

composites biological valves

Coating DPPV with stem cells-fibrin complex

Static condition: 1st day—homogenous distribution of SC;

16th day—cell colony formation in SC-F compared to control (no cell clusters);

Dynamic conditions: starting with the 4th day, floating composite clots at the inner surface of the valve and leaflets are observed.

[152]

Fibrin-based tubular heart valves

The tube mounted on a frame with three struts which, upon back-pressure, cause the tube to collapse into three coating “leaflets”.

In vitro: excellent performance under hydrodynamic conditions, minimal RF (approx. 5%), excellent values for TGV and EOA;

In vivo (sheep, 2 months): substantial recellularization and no significant change in diameter or mechanical properties.

[153]

Tubular construct sutured at the root circumferential line and at three single points of sinotubular junction.

Advantage of one-piece construct manufacturing method without glue;

In vivo (sheep, 3 months): no thrombus formation, calcification or stenosis; formation of ECs confluent monolayer on the valve surface.

[97]

Fibrin-based tube-in-stent heart valves

Fibrin gel and HUVCs molded as tube-in-stent form and sewn into a self-expandable nitinol stent.

Homogeneous cells distribution throughout the valve;

The simulation of the catheter-based delivery (the valves crimping for 20 min) does not influence the valve mechanical properties or functionality.

[76]

F-ELR

biomimetic heart valves

Multi-step injection molding: the valve wall obtained from F gel and the leaflets from F-ELR gel.

Good structure cohesion and functionality (opening/closing cycles); Different cell type localization: the vessel-derived α-SMA negative (leaflets) and α-SMA positive cells (valve wall).

[32]

F/PLDL-PLGA anisotropic composites

BioTexValve

Molding of PLDL multifilaments and electrospun

PLGA fibers incorporated within fibrin gel.

Anisotropic Young’s moduli comparable with the native aortic leaflets;

The valve withstands aortic flow/pressure conditions in flow-loop system;

Homogeneous distribution of α-SMA, aligned with the longitudinal direction of the wall and leaflets.

[91]

SF/LDI-PEUU nanofibrous scaffolds

SF and LDI-PEUU prepared by electrospinning process.

Smooth and porous 3D structure of SF/LDI-PEUU scaffolds with randomly oriented fibers;

Good blood compatibility (hemolysis rate <5%);

HUVECs have spindle-shaped morphology and good spread.

[154]

Abbreviations: C4S—chondroitin-4-sulfate; CDCs—cardiosphere-derived cells; CH—chitosan; DMEM—Dulbecco’s modified Eagle medium; DPPV—decellularized porcine pulmonary valve; ECs—endothelial cells; EDC—1-ethyl-3-(3-dimethylaminopropyl) carbodiimide-hydrochloride; EL—elastin; ELR—elastin-like recombinamer; EOA—effective orifice areas; F—fibrin; FBS—fetal bovine serum; FIs—fibroblasts; HA—hyaluronic acid; HUVECs—human umbilical vein endothelial cells; HUVCs—human umbilical vein cells; LDI-PEUU—L-lysine diisocyanate poly(ester-urethane)urea; MMPs—matrix metalloproteinases; NRASMCs—neonatal rat aortic smooth muscle cells; NSH—N-hydroxysuccinimide; OCAs—ovine carotid artery-derived cells; PBS—phosphate buffered saline; PLDL—poly(L/D,L-lactide); PLGA—poly(lactic-co-glycolic acid); PTFE—polytetrafluoroethylene; RF—regurgitant fractions; SCs—stem cells; SDS—sodium dodecyl sulfate; SF—silk fibrinoin; α-SMA—α-smooth muscle actin; SMCs—smooth muscle cells; 3T3—mouse fibroblasts cells; TVG—transvalvular pressure gradients; VECs—valvular endothelial cells; VICs—valvular interstitial cells.

5.3. Structure-Properties-Functionality Correlations in HVTE

The scaffold, as mentioned earlier in this entry, is one of the most important entities to be considered for efficient tissue engineering, because its properties affect both the generation of the tissue construct in vitro and its post-implantation functionality. Along with the biological and mechanical properties, an important role is played also by the structural properties of the scaffolds, which refer mainly to external geometry, surface properties, pore density, pore size and interconnectivity, interface adherence, etc. [17][18][155].
From Table 5 and Table 6, it can be seen that the scaffolds based on polysaccharides and proteins, respectively, show a wide diversity, both compositional and structural. This is reflected in materials such as simple casting films, electro-spun membranes or fibers and continuing to the most complex hybrid, fibrous, nanostructured or multilayered scaffolds. Starting from this idea, in Table 7, the researchers tried to systematize these scaffolds, obtained by different techniques, and highlight how their structural properties influence their efficiency as scaffolding materials in HVTE.
Table 7. Structural properties of the scaffolds and their efficiency in HVTE.
Abbreviations: bFGF—basic fibroblast growth factor; ECM—extracellular matrix; EOA—effective orifice area; HV—heart valve; IPNs—interpenetrated networks; RF—regurgitation fraction.
The native heart valve is a complex trilayered structure consisting of collagen, elastin and glycosaminoglycans; thus, any hydrogel/composite/hybrid-like material that approaches its structure and composition is the most suitable scaffold for engineering a heart valve construct. Over time, various techniques have been used to obtain scaffolds with different structural features, in order to best perform the functions of the heart valve tissue. The conventional fabrication techniques, such as solvent casting to obtain films or membranes [121][122], do not enable the fabrication of complex architectures with precise control of pore size and geometry, pore interconnectivity or spatial distribution within the scaffold. In this case, although materials with a uniform structure, small thickness and good mechanical strength can be obtained, sometimes problems arise due to poor adhesion of the cells to scaffold surface or low spreading of the cells within the scaffold. In this case, it is necessary to immobilize on the surface of the materials clues, such as the adhesive proteins (fibronectin and mouse laminin) [121] or the growth factors (bFGF, basic fibroblast growth factor) that increase the cells adhesion to the surface [122].
With the development of new techniques (i.e., electrospinning, 3D printing, etc.), heterogeneous 3D scaffolds with strong mechanical strength and with the optimum characteristics of an ideal scaffold for cardiac tissue engineering, such as the morphology and accuracy of native ECM, can be fabricated. In this context, the feasibility and attractiveness of the electrospinning technique can be mentioned in the fabrication of porous 3D structures with small thickness, controlled porosity and optimum fibers diameter [124]. Generally, the porous scaffolds are distinguished by an interconnected homogeneous pore network, providing a continuous flow of nutrients and metabolic waste to enable cells growth, proliferation and spreading. Regarding this technique, it also allows obtaining fully aligned or randomly oriented fibrous scaffolds, which positively influence the mechanical properties, pores dimensions and interconnectivity, and implicitly, the proliferation and the distribution of the cells [126][154].
Hydrogels and IPNs hydrogels are complex structures, having ECM-like 3D microstructure and great mechanical strength that support both 2D surface-seeded cell culture and 3D cell encapsulation. They have the advantage of being obtained by a light technique, in different shapes, sizes and thicknesses, with porosity and mechanical strength controlled by changing the degree of crosslinking. Due to the porous 3D network structure, they are characterized by an excellent cell proliferation and extensive cell spreading with a faster migration rate. Their compatibility with biological tissues, high water content and relatively good mechanical properties make these materials attractive for HVTE applications. Moreover, by adding cells to hydrogel before the gelling process, these can be distributed homogeneously throughout the resulting scaffold [41][127][128][132][133].
Fibrous scaffolds are superior scaffolds in term of cell adhesion, migration, proliferation and differentiation, due to the high aspect ratio of fibers, growth factor loading efficiency and sustained release capacity. The development of nanofibers has enhanced the scope for fabricating scaffolds that can potentially mimic the architecture of natural human tissue at the nanometer scale. For HVTE, fibrous scaffolds provide an ideal environment for cells growth and proliferation, leading to 3D structures with porosity, pore size and mechanical characteristics comparable to native heart valves [125].
As a conclusion, each of the above presented scaffolds have advantages and disadvantages. However, the unique properties of the materials used either alone or in combination with other natural or synthetic polymers work for the purpose to develop new heart valves with the ability to repair, reshape and regenerate the cardiac tissue.

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