Accepted Manuscript Scaffolds for tissue engineering of cardiac valves S. Jana, B.J. Tefft, D.B. Spoon, R.D. Simari PII: DOI: Reference:

S1742-7061(14)00125-1 ACTBIO 3165

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Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

27 November 2013 25 February 2014 12 March 2014

Please cite this article as: Jana, S., Tefft, B.J., Spoon, D.B., Simari, R.D., Scaffolds for tissue engineering of cardiac valves, Acta Biomaterialia (2014), doi:

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Scaffolds for tissue engineering of cardiac valves Running title: Heart valve scaffolds

Jana S, Tefft BJ, Spoon DB, and Simari RD* Division of Cardiovascular Diseases, Mayo Clinic, 200 First Street SW, Rochester, MN 55905, USA

*Corresponding Author: Robert D. Simari, M.D., Division of Cardiovascular Diseases, Mayo Clinic, 200 First Street SW, Rochester, MN 55905, USA, E-mail: [email protected], Fax: 507-538-6418



Tissue engineered heart valves offer a promising alternative for the replacement of diseased heart valves avoiding the limitations faced with currently available bioprosthetic and mechanical heart valves. In the paradigm of tissue engineering, a three-dimensional platform — the so-called scaffold — is essential for cell proliferation, growth, and differentiation as well as the ultimate generation of a functional tissue. A foundation for success in heart valve tissue engineering is recapitulation of the complex design and diverse mechanical properties of a native valve. This article reviews technological details of the scaffolds that have been applied to date in heart valve tissue engineering research.

Keywords: Heart valve, tissue engineering, scaffold, fiber, hydrogel


Introduction Valvular heart disease (VHD) is a major health problem that results in substantial morbidity and death worldwide [1]. In the western world, 2.5% of the population have a dysfunctional or diseased valve [2, 3]. Secondary to the aging of the population it is predicted that there will continue to be an increase of VHD in industrial nations, owing primarily to an increase in degenerative pathology [2]. In the United Kingdom alone, more than 4 million people from 75 to 84 years of age could develop VHD by 2018, and this figure could double by 2028 [4]. In developing countries, VHD is primarily caused by the persistent burden of rheumatic fever rather than degenerative pathology and tends to effect younger individuals [5, 6].

The pathophysiology of valvular heart disease is broad and the specific etiology varies by the particular valve affected. The semilunar valves, consisting of the aortic and pulmonic valves, are commonly affected and have distinct primary pathologic mechanisms of failure. Pulmonic valve disease is most commonly related to congenital abnormalities and tends to present early in life. Aortic valve disease most commonly presents as calcific aortic valve stenosis secondary calcific degeneration [7, 8] while the presence of a congenitally bicuspid aortic valve predisposes to subsequent valvular stenosis and regurgitation [9].

Calcific aortic valve stenosis is the most common valvular pathology requiring valve replacement and is present in to some degree in 2.8% of adults over the age of 75 years and a far larger population has evidence of some aortic valve thickening known as valvular sclerosis [10, 11]. Despite the frequency of calcific aortic valve stenosis, our understanding of its pathogenesis remains incomplete. While there are similarities


between the risk factors and mediators between calcific aortic valve disease and atherosclerosis as many as 50% of patients with calcific aortic valve disease do not have evidence of significant atherosclerosis [12, 13]. Recent data demonstrate that valvular calcification is not a passive process, as originally thought, but rather an active process that relies on activation of pro-osteogenic signaling cascades such as bone morphogenetic protein and Wint/β-catenin for the induction and progression of disease [14, 15]. Additionally, our understanding of the cellular mediators of valvular calcification continues to expand. Conventionally, the differentiation of valvular interstitial cells into an osteoblast-like phenotype with the capacity to produce calcification has been thought to be the primary cellular driver of valvular calcification [16]. Recently, valvular endothelial cells have been implicated through a process of endothelial-mesenchymal transformation as have circulating progenitor cells through differentiation or paracrine signaling [17-20]. The calcification process results in mechanical disruption of valve function, which can lead stenosis, regurgitation, or the combination.

Unfortunately, the treatment of dysfunctional heart valves requires surgical or interventional repair or replacement. Replacement options currently include mechanical or bioprosthetic valves. Mechanical valves have excellent durability; however, the risk of thromboembolism necessitates the use of anticoagulation therapy and its attendant morbidity. Bioprosthetic valves are less thrombogenic; however, they are less durable and more prone to degeneration particularly when implanted in younger individuals [21]. Bioprosthetic valves are generally treated with glutaraldehyde (GA) to stabilize the tissue by preventing rejection of the xenogenic scaffold. However, such treatments stiffen the fiber network and diminish the cushioning function of the spongiosa layer [22]. In addition, GA is toxic and inhibits the repopulation of cells after implantation [23]. Both mechanical and bioprosthetic valves share another disadvantage. They cannot grow and remodel which may necessitate sequential surgeries in pediatric patients [24].


Nevertheless, the current generation of bioprosthetic pericardial valves are adequate substitutes for the majority of elderly patients as they do not require anticoagulation in the majority of patients and their durability is usually sufficient for the lifespan of this population. In the pediatric and young adult populations requiring aortic valve replacement, the Ross procedure has been used and has been shown to have low perioperative mortality and rates of re-operation [25, 26]. In this procedure, a patient’s disease aortic valve is replaced with their own modified pulmonic valve (autograft) and then a cadaveric pulmonic valve allograft is used to replace the pulmonic valve. This procedure has several advantages including minimal thromboembolism, favorable hemodynamics, and potential for valve growth. A disadvantage of this procedure is the harvesting of the healthy pulmonic valve leading to the development of pulmonary valve disease in addition to aortic valve disease. As an alternative, tissue engineering is a promising approach for the treatment of defective or diseased heart valves [27]. In this method, living cells are grown (in vitro or in vivo) onto a supporting three-dimensional (3D) biocompatible structure to proliferate, differentiate, and ultimately grow into a functional tissue construct (Figure 1) [28-30]. Importantly, a tissue engineered valve may be capable of growth and remodeling and may mitigate the need for anticoagulation.

The scaffold is one of the most important entities to be considered for efficient tissue engineering 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 [31, 32]. All scaffolds designed for tissue engineering applications must meet basic requirements such as biocompatibility, sterilizability, and mechanical integrity. Scaffolds intended for heart valve tissue engineering face additional distinct challenges due to their direct contact with blood. Specifically, the construct should be


resistant to calcification and thrombosis [33]. In addition, the construct must withstand the unique hemodynamic pressures and flows of the cardiac environment from the moment of implantation. These unique challenges underscore the importance of carefully considering the materials and design when fabricating a scaffold for tissue engineered heart valves.

Semilunar valves in human (pulmonic and aortic) consist of three semicircular leaflets (also called cusps) attached to a fibrous annulus called the root [23]. The leaflets are less than 1 mm thick and have a flexible structure consisting of three distinct layers: the fibrosa, spongiosa and ventricularis (Figure 2). These layers are composed of valvular interstitial cells (VICs) within a matrix of collagen, elastin and glycosaminoglycans (GAGs). Normal leaflets are virtually avascular and obtain nutrients and oxygen from the bloodstream via hydrodynamic convection and diffusion. In contrast, the aortic or pulmonary root is a bulb-shaped fibrous structure with intimal, medial, and adventitial layers. They are primarily populated with endothelial cells in the intima, smooth muscle cells in the media, and fibroblasts in the adventitia.

In brief, the structure of semilunar heart valves is complex. To mimic this organization in heart valve engineering, the first requirement is an appropriate scaffold structure comprised of suitable morphologies, surface properties, mechanical properties, and pore sizes. In this respect, a broad knowledge of potential heart valve tissue engineering scaffolds is important for their application. However, to date only a limited number of studies have examined outcomes of different scaffolds used for heart valve tissue engineering. This review thus focuses on the scope of usefulness of different scaffolds to regenerate heart valve tissue. It first provides a brief summary of native valve’s physiological properties that can be mimicked into the scaffolds intended for heart valve engineering. It then covers the most common different scaffolds and


scaffolds with combination of various structures used to regenerate heart valve tissue [34]. At the end, it focuses on their drug delivery capabilities which are important for regeneration of functional heart valves and for mitigation of thrombogenicity. The article also outlines the strategies of in vitro and in vivo population of cells and those cell types that have tested.

Physicochemical properties of native valve The physicochemical properties of a scaffold including architecture/morphology, mechanical properties, and drug delivery capabilities are important parameters in tissue engineering as they together create a microenvironment on which cell adhesion, proliferation, migration, differentiation, and subsequent tissue development depend [35]. In order to optimize results in heart valve tissue engineering, scaffolds should mimic the physicochemical properties of native heart valve. In this respect, a brief discussion on the formation and subsequent physicochemical properties of native valve could define the set of goals in designing of scaffolds intended for heart valve tissue engineering. Heart valve cusps are thin; however, they reliably perform under high blood flows and pressures. This is possible due to the organization and orientation of collagen, elastin and other extracellular proteins within the heart valve. The valve cusp structure is not just a single organized ECM layer inhabited by cells. It is more complex and organized into three layers — a layer of inflow surface (ventricularis) with a radially aligned fibrous morphology, a central layer (spongiosa) containing loose collagen and abundant proteoglycans in a random manner, and a layer of outflow (fibrosa) surface consisting of circumferentially aligned collagen fibers (Figure 2) [22, 23, 36].


During development, heart valves go through several essential stages: 1) the endocardial cushion is formed and endothelial-mesenchymal transformation (EMT) takes place; 2) endocardial cushion growth and subsequent formation of the valve primordia and 3) thinning and elongation of the valve primordia to form valve cusps. At each stage, different signaling and transcriptional molecules regulate valvulogenesis [36-38]. Transforming growth factor beta (TGF-β) cytokine superfamily including bone morphogenetic proteins (BMPs) act to support cushion formation initiation by increasing the ECM synthesis and inhibiting the expression of chamber-specific genes [39]. BMP2 and BMP-4 are responsible for increased deposition of hyaluronan and versican in cushion morphogenesis regions of the outflow tract [39, 40]. Vascular endothelial growth factor (VEGF) from endothelial cells encourages proliferation of endocardial cushion endothelial cells [38]. Endocardial cushion EMT is initiated by TGF-β family [41]. TGF-β receptors and ligands also are expressed during EMT [38]. During endocardial cushion EMT, signaling of WNT/β-catenin is associated with TGF-β and increases endocardial cushion EMT [42]. Expression of VEGF terminates EMT and promotes endothelial cell proliferation leading to cushion growth [38]. At the end of EMT, through cell proliferation and ECM deposition, endocardial cushions continue to grow and ultimately, endocardial cushions fuse to form valve primordia [37, 43]. BMP signaling and fibroblast growth factors (FGFs) including FGF4 and FGF receptors 1, 2 and 3 are responsible for both these developments [44, 45]. Wnt/β-catenin and VEGF/NFATc1 are involved in these processes [46, 47]. In thinning and elongating of valve primordia into valve cusps, Notch signaling, BMP2 signaling and EGF4 expression have been regarded as important [48]. In addition, large numbers of intermediate signals such as Tb×20, Twist1, periostin, cadherin-11, Sox9, and Msx1/2 are responsible for valvulogenesis [49-52]. Thus the ultimate structure of the adult valve is the result of a complex interplay between cells, matrix and cytokines and growth


factors over relatively long periods. Engineering approaches must consider this interplay throughout the process.

Mechanical properties of human aortic valves in both the circumferential and radial directions have been shown to differ [53]. The modulus of elasticity, ultimate stress, and ultimate strain were 15.34±3.84 MPa, 1.74±0.29 MPa, and 18.35±7.61%, respectively, in the circumferential direction. In the radial direction, these values were 1.98±0.15 MPa, 0.32±0.04 MPa, and 23.92±4.87%, respectively. The mechanism responsible for these dissimilar properties is the orientation of the collagen and elastin fibers in the valve. Circumferential alignment of collagen in the fibrosa causes increased stiffness and strength in that direction whereas radial alignment of elastin in the ventricularis causes increased elasticity in that direction [54]. Proper function of the valve relies heavily upon these anisotropic mechanical properties as too much stiffness can hinder coaptation of the cusps whereas too little stiffness can allow for distortion and regurgitation. It is also necessary to achieve proper mechanical properties because VICs are highly sensitive to local tissue strains and they respond by modulating behavior such as collagen synthesis [55]. Thus, it is imperative to produce scaffolds made of either biologic or synthetic materials with similar anisotropic mechanical properties when engineering a valve.

Types of scaffolds for heart valve engineering The goal of heart valve tissue engineering requires designing a scaffold that provides physiological support for cell attachment, proliferation, and development. The complex structure of a heart valve includes a spongy middle layer sandwiched by two outer laminar anisotropic fibrous layers [56]. To mimic the native heart valve structure, multiple scaffolds designs have been proposed and tested. Two main types of scaffolds have been developed: (1) acellular native heart valve scaffolds from


allogeneic/xenogeneic sources and (2) fully artificial scaffolds fabricated from synthetic and natural (biologic) polymers [57]. The fabricated scaffolds can be further categorized as porous, fibrous, and hydrogel scaffolds. The following sections assess these distinct types of scaffolds used in heart valve tissue engineering.

Decellularized valve scaffolds Xenogeneic heart valves from pigs, cows and sheep are the main sources of acellular heart valves because human valves – the allogeneic valves that are potentially less antigenic compared to their xenogeneic counterparts -- are in short supply especially for pediatric patients (Figure 3a). Decellularized scaffolds retain the original valve structure and many of the ECM molecules, providing potential advantages over fabricated synthetic scaffolds (Figure 3b). In this approach, applying one of many tested decellularization methods, the cells are removed minimizing damage to the original structure to avoid problems during subsequent recellularization or implantation [58, 59]. In some cases, the decellularized tissue scaffolds are modified by crosslinkers such as GA and pentagalloyl glucose (PGG) to sterilize the valve, minimize disease transmission, reduce immunogenicity, and avoid negative outcomes such as calcification and thromboembolism [60].

Different single agents or combinations of agents have been used to decellularize valve tissues, and each treatment has shown advantages and disadvantages [61]. The most commonly used single-agent treatments are nonionic detergents, ionic detergents, and chelating agents [62, 63]. Nonionic detergents such as Triton X-100, sodium deoxycholate acid (SDC), sodium dodecyl sulfate (SDS), and deoxycholic acid disrupt lipid–lipid and lipid–protein connections but not protein–protein connections [64, 65]. They lyse nuclear materials but cannot completely dissolve the fragments, which leads to residue in the scaffolds. In contrast, ionic detergents completely remove


cytoplasmatic and nuclear cellular materials [66, 67]. Loss of glycosaminoglycans, chondroitin sulfate, and other proteins including collagen, laminin, and fibronectin is reduced by treatment with ionic detergents as compared to nonionic detergents [68, 69]. Thus, the scaffolds are more intact compared to those treated with nonionic detergent. Chelating agents, such as ethylenedaminetetracetic acid (EDTA), remove the cellular material from the tissue by isolating bivalent cations such as Mg2+ and Ca2+, which are necessary for cells to attach to collagen within the extracellular matrix [70, 71].

In the enzymatic method of single agent decellularization, proteolytic enzymes such as trypsin cleave peptide bonds to decellularize the valves. Endonucleases and exonucleases degrade ribonucleic acid (RNA) and deoxy-ribonucleic acid (DNA) [72, 73]. Endonucleases catalyze the disruption of the interior bonds of ribonucleotide or deoxyribonucleotide chains whereas exonucleases catalyze the disruption of the terminal bonds of these chains. Trypsin-treated scaffolds showed fragmentation and distortion of elastin fibers [74]. Collagen distribution is also distorted, and glycosaminoglycans, laminin, fibronectin, and chondroitin sulfate were almost completely washed out after treatment. However, resynthesis of chondroitin sulfate, laminin, and fibronectin has been achieved by seeding endothelial cells in the scaffolds. Most of the enzymes used in the decellularization process come from non-human sources and non-recombinant sources raising the possibility of disease transmission through the decellularization process [75].

Decellularization with a combination of reagents has shown advantages over singleagent treatments. Cadaveric valves treated with trypsin and Triton X-100 revealed maintained fibrous scaffold structure and mechanical strength with low antigenicity [66]. Detergent (e.g. EDTA/sodium-deoxycholate and Triton X-100) and enzymatic (e.g. RNase, and DNase) reagents together were also used to remove cells from cadaveric


heart valves. The mechanical properties of the decellularized scaffolds showed no significant change with retention of sufficient collagen and elastin along and glycosaminoglycans [76]. Zhao et al. prepared decellularized porcine heart valves in the same manner and cultured them with canine endothelial cells (ECs) [22]. The results showed that leaflet scaffolds were covered with cells [77]. A combination of EDTA/sodium-deoxycholate, Triton X-100, RNase, and DNase treatments was used to decellularize porcine heart valves, which were then cultured with ECs [78]. The ECs formed a layer on the surface of the scaffolds. Korossis et al. treated fresh porcine aortic valve leaflets with hypotonic buffer, agarose gel, trypsin, and SDS to obtain acellular scaffolds [79]. The scaffolds showed significant loss of elastin fibers in the ventricularis layer and increased transition strain compared to fresh porcine samples. Therefore in acellular scaffolds, the mechanical strength and other leaflet competencies under different systematic treatments were more or less compromised.

The effect of different reagents on cells used for heart valve decellularization is different and thus, these reagents need varioable time periods for complete cell removal. In addition to removing cells, these reagents damage various proteins in the valve in a time-dependent manner. Therefore, researchers have investigated different time periods of valve decellularization to achieve cell removal with minimal damage to ECM components. For example, Cebotari et al. applied 1% sodium deoxycholate (SDC) solution, 1% sodium dodecyl sulfate (SDS) solution, and a mixture of 0.5% SDC solution and 0.5% SDS solution for 24 hours under continuous shaking condition at 22oC [80]. Decellularized valves were then washed with phosphate buffered solution in 10 steps with each step for 12 hours at 22oC under continuous shaking. Sacks and colleaguesaltered the time periods of three previously investigated decellularization methods (anionic detergent-based: SDS, enzymatic agent-based: Trypsin, and nonionic detergent-based: Triton X-100) to 48 hours to compare their effects on the


mechanical and structural properties of decellularized valves [81]. T hey soaked cadaverous valves in a hypotonic Tris buffer (10 mM Tris, pH 8.0) with 0.1% EDTA and 10 kIU/ml aprotinin for 1 hour and then treated with in 0.1% SDS in the same hypotonic Tris buffer, protease inhibitors, RNase A (20mg/ml), and DNase (0.2 mg/ml) for 48 hours (original was 72 hours) during continuous shaking at room temperature [82]. Another combination, Trypsin and EDTA (0.5% Trypsin and 0.2% EDTA) in a hypotonic Tris buffer with RNase A (20mg/ml) and DNase (0.2 mg/ml) was used for decellularization for 48 hours during continuous shaking at 37oC [83]. On the other hand, 1% Triton X-100 was mixed with 0.2% EDTA (Sigma) in a hypotonic Tris buffer with RNase A (20mg/ml) and DNase (0.2 mg/ml) and then applied for decellularization for 48 hours (original was 24 hours) during continuous shaking in a cell culture incubator (i.e. 5% CO2/95% air atmosphere at 37oC) [84]. In all of the above methods, decellularized valves were washed with copious PBS during shaking, usually at room temperature. The SDS-based method seemed to best preserve the required mechanical and structural properties of valves. Some investigators have modified these above methods further to obtain better results in terms of decellularization with less damage to the existing proteins [66].

After decellularization, acellular scaffolds are sometimes treated with cross-linkers such as GA, PGG, and NDGA (nordihydroguaiaretic acid) prior to implantation. This treatment is performed to stabilize the collagen matrix and decrease antigenicity [85, 86]. GA has been commonly used for cross-linking of ECM molecules in acellular scaffolds. However, these scaffolds have reduced cell compatibility with cell toxicity and reduced proliferative capacity of cells following repopulated. The use of PGG, a collagen-binding polyphenol cross-linker, results in greater retention of biaxial mechanical and other biological properties in leaflets when compared to GA-treated leaflets [86]. In vivo, these PGG-treated leaflets did not calcify, and they supported


infiltration by host fibroblasts and subsequent matrix remodeling. To improve cell attachment and seeding, cells can be encapsulated in polyethylene glycol (PEG) hydrogel and then seeded into acellular scaffolds [87]. PEG hydrogel was shown to assist the scaffold in retaining seeded cells when the scaffold was tested in a bioreactor with pressurized medium flow. Baraki et al decellularized ovine aortic valve conduits with detergents and implanted as an aortic root in lambs up to 9 months [88]. The controls were fresh native ovine aortic valve conduit. Explants showed trivial regurgitation and no sign of graft dilation, degeneration or rejection. Minute calcium deposits at the anastomosis and thrombi formation on the leaflets were observed. There were endothelial monolayer at the luminal side and neovascularization at the adventitial site. Control samples also demonstrated some adverse signs including calcification and degeneration. These findings suggest that methods of decellularization and associated treatment affect the mechanical and biological function of implanted scaffolds.

If cross-linking agents are used, scaffolds exhibit less porosity, smaller pore size, and stiffer extracellular matrix. A harsh treatment such as acetic acid can be applied to acellular scaffolds to increase their pore size and porosity and the resulting scaffolds can be conjugated with RGD for higher cell adhesion [89]. When cultured with human mesenchymal stromal cells in vitro, scaffolds treated with acid showed increased proliferation and migration of cells deep into the scaffold.

In addition to chemical and biological reagents, physical methods such as temperature, force, flow/pressure, and non-thermal irreversible electroporation have also been applied to remove cells from heart valves. A freeze-thaw process has been shown to lyse cells while subsequent processing was required to remove the intracellular contents [90]. This process also disrupted the ECM ultrastructure and this disruption


must be considered [91, 92]. Cells from the surface of a tissue have been removed by physical or chemical means [92]. Hydrostatic pressure may be more effective than detergents and enzymes for removing cells from the surface [93]. Electrical pulses have also been applied to kill cells by inducing the formation of micropores in the cell membrane [94]. Following exposure, the cellular remnants need to be removed from the tissue.

As decellularized scaffolds are made of ECM, original scaffolds still degrade and are replaced by new ECM deposition [95, 96]. The degradation rate depends on a variety of factors including decellularization treatment, cross-linking technique, and the use of a structural modification agents such as epigallocatechin-3-gallate (EGCG) [97-99]. It is understood that each cell removal agent and method causes some degree of structural disruption and alteration of ECM composition. Thus, the objective of decellularization should be to minimize the undesirable effects and to improve the recellularization efficiency so that recellularized bioscaffold implants can be applicable for tissue engineering.

Despite promising results in ovine implantation models, clinical outcomes of decellularized valves were disastrous [100, 101]. The SynerGraft (Cryolife, Inc., USA) trial involved the implantation of decellularized porcine valves into the right ventricular outflow tract of four children and all four grafts failed due to inflammatory response (at day 2) leading to structural failure (at day 7) and degeneration of leaflets (after 6 weeks) and wall (after 1 year) [100]. The decellularization process was predicted to remove the majority of antigenic material from the tissue rendering it relatively inert. However, subsequent analysis of the pre-implantation SynerGraft scaffolds revealed they were inadequately decellularized and still contained immunogenic cell fragments and DNA [100]. Thus, a strong inflammatory response occurred and no host cells were found to


repopulate the tissue. Interestingly, favorable results have since been reported with decellularized allogeneic valves in terms of immunologic response, durability, and overall clinical performance [102-105]. There has even been a report of host cell repopulation [106]. Nevertheless, pre-seeding these scaffolds with host cells is likely to yield more favorable results.

As a decellularized construct relying on host recellularization, allogeneic valves have proven to be far superior to xenogeneic valves. Nevertheless, xenogeneic valves have the distinct advantage of being in plentiful supply and strategies are being explored to mitigate the immunologic response to decellularized xenogeneic valves. For example, alpha-Galactosidase and gene knock out technology have been used to remove the critical alpha-Gal xenoantigen [107, 108]. Antigen masking by gluteraldehyde fixation is undesirable for a tissue engineered valve and has been found to be ineffective [109]. Ideally, a decellularized scaffold, whether xenogeneic or allogeneic, will be seeded with autologous cells that will degrade the foreign matrix and synthesize an entirely autologous tissue to eliminate immunological responses following implantation.

Fabricated scaffolds Direct fabrication of a scaffold mimicking the structure and function of a heart valve from raw synthetic or biological materials is an approach distinct from decellularization of biological scaffolds [34, 110]. Potential advantages of fabricated polymeric scaffolds may include less immunogenicity and thrombogenicity [111]. In addition, the biodegradability, durability, and mechanical properties of these scaffolds can be controlled. The disadvantages are that their structure may not fully mimic the complex structure and function of native tissue with varied mechanical properties. In general, three types of scaffolds have been fabricated and applied to heart valve tissue engineering: porous scaffolds, fibrous scaffolds, and hydrogels.


Solid 3D porous scaffolds Highly porous three-dimensional scaffolds with an interconnected homogeneous pore network and high pore size provide continuous flow of nutrients and metabolic waste to enable growth and vascularization of engineered tissues [34]. A number of fabrication methods have been applied to natural and synthetic polymeric materials to produce porous scaffolds meant for heart valve tissue engineering. The conventional techniques for scaffold fabrication are particulate leaching, solvent casting, gas foaming, vacuum drying, thermally induced phase separation, melt molding, high internal phase emulsion, and microfabrication [112-115]. Although various groups have applied these relatively straightforward methods, they have some drawbacks. Most importantly, the pores are of irregular size and are not interconnected. Solid free-form (SFF) fabrication is a technique to prepare 3D porous scaffolds with interpore connections. The scaffolds are created by using computer-controlled tools for layer-by-layer deposition of materials [116, 117]. The geometry can be obtained from a solid model file or digital data produced by imaging sources such as microcomputer tomography and magnetic resonance imaging. Three dimensional printing is the advanced form of SFF with one or several cartridges containing different biomaterials, biomolecules, or other required materials [118, 119]. With any of these methods, a broad range of materials including synthetic or natural and organic or inorganic could be used to produce 3D porous scaffolds for the engineering of tissues including heart valves.

Polyhydroxyalkanoates (PHAs) are perhaps the materials most often used to produce solid 3D scaffolds applied to heart valve tissue engineering [120-123]. PHAs are a class of natural polyester polymers produced by different microorganisms [124]. Initially, polyglycolic acid (PGA) and polylactic acid (PLA) were used to prepare valve leaflets as promising alternatives to the decellularized biomatrix scaffolds. However,


their relatively high stiffness and rigidity made it impossible to obtain adequate cell proliferation of applied cells [125-127]. Stock et al. then introduced polyhydroxyoctanoate (PHO), a soft elastomeric polymer and member of the PHA family, to PGA materials to overcome the issues with mechanical properties [127]. They prepared porous (pore size 80–180 µm) PHO leaflets (120–µm thickness) using the salt leaching method and coated the construct with PGA felt. The scaffold construct was then cultured with autologous smooth muscle cells and endothelial cells and their progressive cellular and ECM deposition indicated desirable remodeling of the tissueengineered structure. After 24 weeks of implantation in a sheep model, there was no thrombus formation and only mild, nonprogressive valvular regurgitation in the seeded implant whereas the unseeded implant demonstrated thrombus formation after 4 weeks.

Besides PHO, other members of the PHA family have demonstrated promising results in heart valve tissue engineering [23]. Various hydroxy acid monomers are incorporated to make PHA polymers, and these monomers can be substituted with other groups such as alkyl and aryl. PHAs have been of interest in biomedical applications and the emerging field of tissue engineering due to their high biocompatibility and tunable mechanical properties [122]. Several companies, e.g., Metabolix, Inc. and Tapha, Inc., have been developing PHA-based polymers including poly-3-hydroxybutyrate (poly(3HB)), poly-3-hydroxybutyrate-co-3-hydroxyvalerate (poly(3HB-co-3HV), poly-R-3hydrooxyoctanoate-co-R-3-hydroxyhexanoate (poly(3-HO-co-3HH)), poly-4hydroxybutyrate (poly-(4HB)), and poly-R-3- hydroxybutyrate-co-4-hydroxybutyrate (poly(3-HO-co-4HB)) [128]. Their wide range of thermal properties allows for the use of different processing methods for a variety of product development needs including scaffolds for tissue engineering. Mechanical properties of the PHA polymers vary widely and thus make them suitable for engineering both soft and hard tissues (Table 1) [128]. Their mechanical properties can be varied further by extending the pendant


group. For example, when the hydroxyvalerate group of poly-3-HB is extended, the polymer becomes more elastomeric (compare 2nd column and 4th column in Table 1). Trileaflet heart valve scaffolds were prepared from porous PHA foam (Tepha, Inc.) with pore size of 180–240 µm and compared with acellular biomatrix scaffolds (Figure 4) [129]. Vascular cells from ovine carotid arteries were seeded onto these scaffolds, which were then implanted in a lamb model. Interestingly, all samples except the acellular scaffolds were covered with tissue and exhibited collagen and ECM deposition. No thrombi formed on any PHA scaffolds. In another study, vascular cells were seeded into PHA leaflet scaffolds (Meatabolix Inc, 0.3 mm thick) with pore size 100 to 240 µm and similar results were observed [130]. The cells were viable and ECM deposition was observed following pulsatile flow exposure within a culture system. Expression of crosslinked elastin by vascular smooth muscle cells was also observed in PHA scaffolds [131]. Some researchers tried to improve the biocompatibility of P-3HB-co-3HH porous scaffolds by coating with silk fibroin and found improved growth of smooth muscle cells on the silk fibroin modified hybrid scaffolds [132].

Beside PHAs, other natural polymers such as chitosan and collagen have been used to prepare porous scaffolds for heart valve engineering. A collagen sponge with irregular connecting pores of 58 ± 24 µm diameter facilitated proliferation of seeded porcine aortic myofibroblasts [133]. The cells also synthesized a considerable amount of ECM such as collagen and elastin. Chitosan-modified PCL porous scaffolds were fabricated to improve the attachment of fibroblast cells for heart valve tissue engineering [134]. This experiment also demonstrated that chitosan enhanced the biocompatibility of the PCL scaffolds, which has since been used extensively in valve and blood vessel tissue engineering.


A composite scaffold prepared in combination with poly(DL-lactide-co-caprolactone), poly(D,L-lactide-co-glycolide) (PLGA), and type I collagen have been tested to determine their efficacy in heart valve engineering [135]. The scaffold had a void volume of 80%. Seeded neonatal rat heart cells showed higher cardiac markers expression and contractile properties in composite scaffolds as compared with controls made of each individual material. Appropriate porosity, hydrophilicity, structural stability, elasticity, and degradability were the main reasons for the improved results.

Scaffolds with proper morphology can guide cells and consequently the depositing proteins that replace the gradually degrading scaffold to achieve this specific organization [136]. To obtain native collagen fiber orientation and mechanical properties in engineered heart valves multiple approaches have been used. Engelmayr and colleagues demonstrated the use of microfabricated scaffolds with diamond shaped pores and observed a correlation between the aspect ratio of the pores and the alignment of collagen fibers [137]. Microarchitecture of the scaffold influenced the structural orientation of deposited collagen fibers through contact guidance of cellular orientation [138]. Accordion-like honeycomb microstructured scaffolds made of polyglycerol sebacate (PGS) were found to generate native-like ventricular myocardium tissue constructs with proper stiffness and anisotropy due to the scaffold’s precise biophysical environment directing the seeded neonatal rat heart cells [139]. The architecture guided cell alignment whereas the controllable stiffness of PGS in conjunction with electrical stimulation offered the in vivo biophysical environment to the cells to enable their growth and maturation into a desirable tissue construct.

Some current research indicates that PGS could be a worthwhile material for heart valve tissue engineering [140]. It is a biocompatible, biodegradable, and, more importantly, a tough elastomer with mechanical and degradation properties that can be


controlled by manipulating the polymerization parameters. Diamond shaped microporous PGS scaffolds were used for heart valve tissue engineering by culturing VICs up to 28 days. The cells were able to secrete collagen and the deposited collagen was aligned inside the diamond shaped micropores [141]. Despite interesting biomechanical properties, PGS has poor water uptake capacity (2%), which limits its utility for tissue engineering especially of soft tissues. Patel et al. modified PGS with PEG to prepare PGS-co-PEG block copolymers which have controllable mechanical, degradation, and water uptake properties [142]. Due to the presence of PEG, the mechanical properties improved from 13 kPa to 2.2 MPa and elongation increased sixfold. Coating laminin, fibrin, collagen-I, fibronectin, and elastin on PGS scaffolds increased their cellularity, ECM production, and regulation of cell phenotype [143].

Although, positive results were found during both in vitro and in vivo testing in terms of cell adhesion, proliferation, differentiation, and ECM deposition on fabricated 3D solid porous scaffolds, these tissue engineered valve scaffolds lacked two important features of native valves: 1) shape and 2) elastomeric flexibility. It is very challenging to fabricate 3D porous scaffolds exhibiting the shape and elasticity of native valves. Bioprinting is a leading technology that may be used to fabricate scaffold with above requirements.

Fibrous scaffolds Fibrous scaffolds have achieved great importance in tissue engineering since they resemble the structure of extracellular matrix [144]. Due to the high aspect ratio of fibers, fibrous scaffold are superior in terms of cell adhesion, migration, proliferation, and differentiation, which are important in tissue engineering applications [145-147]. Moreover, fibers have a high growth factor loading efficiency and sustained release capacity to specific sites of application. For heart valve tissue engineering, fibrous


scaffolds would provide an ideal environment for cells if they can form 3D structures with porosity, pore size, and mechanical characteristics comparable to native heart valves [148].

Several techniques including electrospinning, phase separation, and self-assembly have been applied to fabricate nanofibrous materials. Electrospinning is the most commonly used technique to prepare tissue engineering scaffolds due to its versatility, applicability to most polymers, easy handling, and cost-effectiveness [149]. In electrospinning, a high electrical voltage is applied to a polymer solution and when the voltage is strong enough, built-up charges on the surface of the polymer droplets overcome the surface tension and a liquid jet forms. Then, via electrostatic force, the jet accelerates toward the grounded collector and is stretched to produce continuous ultrathin fibers [150]. Applying collectors of appropriate shapes and arrangements allows fabrication of nanofibrous scaffolds with desirable forms and fiber orientations (e.g. random or aligned).

A majority of the nanofibrous scaffolds intended for heart valve tissue engineering have been fabricated from PGA polymer. PGA indeed has many favorable properties such as biocompatibility and bioabsorbility; however, the mechanical stiffness is high (modulus of 7 GPa) [151]. More importantly, due to its hydrophilicity, PGA monofilaments can be degraded in less than 4 weeks in vivo, thus providing a constructive environment and influence for higher collagen and ECM production [152]. To prepare nanofibers, PGA has been dissolved in organic solvents such as 1,1,1,3,3,3 hexafluoro-2-propanol (HFP) or chloroform to make a solution that can be electrospun by applying a voltage of 15-20 kV. Hoerstrup et al. cultured human fibroblast cells on electrospun PGA fibrous meshes with fiber diameter of 12–15 µm to construct autologous cardiac valves [153]. On a similar kind of PGA mesh, human fibroblast cells


followed by human endothelial cells were cultured to obtain similar constructs (Figure 5) [154]. Endothelial cells formed a monolayer without formation of capillaries was detected [125]. In native valves, valvular endothelial cells (VECs) form the outer layer and work in a similar fashion, and thus layering of endothelial cells onto a tissue construct seems a rational approach towards functional valve tissue engineering.

The compatibility of PGA non-woven fibrous meshes for heart valve tissue engineering was further improved by dip-coating the meshes with a 1% solution of P4HB, a homopolymer that has a lower degradation rate compared to PGA [155-157], and thus P4HB restricted degradation of PGA. After culturing autologous ovine fibroblast and endothelial cells for 20 weeks, the surface of the tissue construct was smooth, like that of a valve. The ECM and DNA contents were higher on modified PGA meshes compared to PGA meshes alone [155]. Cultured bone marrow-derived mesenchymal stromal cells showed differentiation to the smooth muscle cell lineage characterized by their expression of markers including alpha-smooth muscle actin, desmin, and calponin [157]. Hoerstrup et al. tested PGA/P4HB hybrid scaffolds seeded with human bone marrow stromal cells which showed results in terms of myofibroblast cell marker expression, tissue and ECM formation, and biomechanical properties. In addition, the surface of the engineered leaflet was smooth [158]. Umbilical cord cells (UCC) exhibited proliferation, growth, and differentiation towards a smooth muscle cell phenotype. A layered tissue formed, which contained collagen I, collagen III, and glycosaminoglycans. Collagen fibril formation was also observed [156]. Due to higher compatibility of the hybrid scaffold compared to the PGA scaffold alone, the former resulted in engineered tissue constructs with superior mechanical properties. Further studies on leaflet tissue engineering with this PGA/P4HB hybrid scaffold revealed that addition of the growth factor VEGF and the differentiation factor TGF-β1 increased proliferation, differentiation,


and tissue development [159]. TGF-β1 was suggested to be responsible for transdifferentiation from endothelial to mesenchymal cells.

Several investigators have used PGA mesh to mold into valve shaped scaffolds by applying P4HA. Molded scaffolds were then cultured with appropriate cell lines including autologous ovine myofibroblasts and endothelial cells and autologous bone marrow-derived mononuclear cells in vitro and implanted in vivo to determine the biocompatibility of the scaffold systems [155, 160, 161]. The transplanted valve constructs demonstrated maintenance of valvular structure and showed adequate functionality (leaflet mobility and coaptation) up to 4 weeks. There were no sign of thrombus formation and structural damage. From histology and immunohistochemistry, valve constructs were found to obtain layered endothelialized tissue due to substantial cellular remodeling and in-growth into the scaffold material.

Besides PGA, other synthetic polymeric materials have been applied to heart valve engineering. Poly(ester urethane) urea (ES-PEUU)-based fibrous scaffolds with specific anisotropic mechanical properties could be appropriate for valve tissue engineering [148]. This anisotropic behavior of such a scaffold has been obtained by high-speed electrospinning. PGA/PLLA composite fibrous scaffolds were studied to evaluate post-implant characteristics in heart valves [162]. A natural polymer chitosanbased composite scaffold was prepared by incorporating chitosan fibers into a 3D chitosan-porous scaffold. The fibers augmented the modulus and strength of the scaffold [163]. Engineered heart valves with this composite scaffold showed tensile strength as high as 220±17 kPa, which is comparable to the radial values of human valve leaflets.


The rate of biodegradation of scaffold materials is an important parameter in tissue engineering because biodegradation reduces the mechanical properties of scaffold materials and may be balanced by ECM production [34]. Thus, proper material selection or synthesis in a specific tissue engineering application is necessary to achieve this balance. Sant et al. applied a blend of fast-degrading polyglycerol sebacate (PGS) and comparatively slow-degrading polycaprolactone (PCL) to produce scaffolds with controlled biodegradation for heart valve tissue engineering [164]. The PGS–PCL hybrid scaffolds showed a gradual decrease in mechanical properties due to biodegradation, which was compensated by secretion of new ECM by cells within the scaffold. The unmet challenge of nanofibrous scaffolds is their very low pore size (less than 10 µm) and thus cells may not penetrate into the scaffolds. Polymers with appropriate biodegradable properties could be useful in nanofibrous heart valve tissue engineering if the rate of biodegradation can be synchronized to cell proliferation and ECM production. Thin coating of decellularized scaffolds with nanofibers could possibly bring their efficiencies together in heart valve tissue engineering. Hydrogel scaffolds Hydrogels are swollen hydrophilic polymer chain networks with high water content. Due to their structural similarity to extracellular matrix, they have been used as polymeric materials in tissue engineering and regenerative medicine [165, 166]. Hydrogel materials generally show high permeability to oxygen, nutrients, and water-soluble metabolites. They also can be delivered percutaneously, promoting their use for cell encapsulation in tissue engineering, [167]. These materials have many other applications as well including drug delivery, biosensors, and linings for artificial implants [168-171].


Hydrogels are synthesized by cross-linking hydrophilic homopolymers, copolymers, or macromers [172]. Through free radical polymerization, the polymer chains propagate by consuming vinyl monomers or other functional groups such as acrylate base derivatives [173]. The reaction proceeds quickly and uncontrollably leading to a wide distribution of molecular weights and inhomogeneous properties throughout the hydrogel. An alternative synthesis method is conjugate addition reaction based on acrylated monomers, esters, or amides combined with thiols [174-176]. This method brings the risk of side reactions by competing nucleophiles from biological compounds including living cells. Some hydrogels are prepared by click chemistry, which offers mild reaction conditions with high chemical selectivity [177-179]. Thus, the quality of hydrogels prepared by this method is far superior to that of hydrogels produced by conventional methods. Poly(vinyl alcohol) (PVA), PEG, and polyacrylates such as poly(2-hydroxyethyl methacrylate) (PHEMA) are some of the synthetic monomers used to prepare hydrogels. Biological hydrogels can be formed from collagen, fibrin, hyaluronic acid, alginate, and chitosan.

Encapsulation of cells within a hydrogel is a standard technique for cell culturing and tissue engineering [167]. The polymer chains in a hydrogel mimic the structure of proteins and other biomolecule chains within ECM; thus, cell-seeded hydrogels offer high cellular efficiency in tissue engineering through higher cell adhesion, proliferation, and differentiation. PEG-based hydrogels were used as seeding media for bone marrow mesenchymal stromal cells (BMSCs) in decellularized porcine aortic valves for better cell attachment [87]. Differentiation of BMSCs into endothelial and myofibroblast cells was higher in scaffolds with PEG hydrogel compared to only decellularized scaffolds. Decellularized aortic valves seeded with PEG-encapsulated cells demonstrated higher tensile strength compared to the aortic valves seeded without PEG


hydrogel. Culture of porcine VICs and VECs in glycosaminoglycans (GAG)-collagen hydrogels showed enhanced surface coverage of VECs on collagen-GAG constructs compared to collagen-only constructs [180]. Expression of elastin and laminin by VICs and expression of vasoactive molecules and endothelial nitric oxide synthase (eNOS) by VECs were higher on collagen-GAG constructs compared to collagen-only constructs. Bovine aortic valve endothelial cells, bovine valve interstitial cells, and bovine aortic endothelial cells were viable in porous amino acid-hydrogels; thus, these hydrogels could be used to engineer heart valves [181].

Hydrogels have weak mechanical properties, and with the addition of cells, their stiffness further decreases [59, 182, 183]. Therefore, to make cell-encapsulated hydrogels work as a standalone scaffold for tissue construct development, the mechanical properties of hydrogels need to be augmented. Modified PEG hydrogel — polyethylene glycol diacrylate (PEGDA) — is an attractive material for heart valve tissue engineering due to its tunable mechanical and biological properties [184, 185]. Three layers of different PEGDA hydrogels, each with mechanical properties that correspond to the three layers of a native heart valve, were sandwiched to prepare cusp scaffolds. Under simulated flow and pressure conditions in a pulsatile bioreactor, the interfaces and the sandwich remained intact [186]. Another hydrogel system was developed consisting of gelatin macromers synthetically modified with methacrylate functionalities. The stiffness could be regulated by adjusting the porosity which was in turn dependent upon polymerization rate [169]. Encapsulated VICs within the composite hydrogel were able to achieve native morphology within two weeks of culture. PVA hydrogels have mechanical properties similar to soft tissue [169]. By regulating freeze-thaw cycles, the PVA hydrogels obtained mechanical properties similar to that of native porcine aortic roots within the physiological pressure range [187]. To improve the mechanical properties of PVA hydrogels directly without repeated freeze-thaw cycles, PLA has been


included for cross-linking [188]. The complex geometry of artificial aortic valve could be molded to produce scaffold for heart valve engineering. Despite supporting cell adhesion and proliferation deposited ECM and collagen have no specific orientation since hydrogels do not possess any definite structure. Hockaday et al. used a 3D printing/photo-cross-linking system with a PEGDA hydrogel to prepare complex and heterogeneous aortic valve scaffolds (Figure 6) [189]. The scaffolds were supplemented with VIC-encapsulated alginate gel. After 21 days in culture, it was found that due to printed orientation and presence of alginate gel, the engineered scaffold constructs from blended hydrogel had stiffness 10 times greater than did engineered scaffold constructs from original hydrogel. This indicates the former scaffold system can withstand dynamic physiological pressures. Duan et al. used a 3D bioprinter to prepare scaffolds from alginate/gelatin blended hydrogels [190]. Aortic root sinus SMC and aortic VICs were encapsulated into alginate and gelatin gels separately and then mixed to prepare the blend before printing. After seven days of culture, the mechanical properties of the cell-laden scaffolds were much higher than that of acellular 3D printed hydrogels. Bioprinting of cellular microspheroids which can work as bio-ink particles may be useful to fabricate three dimensional structure like heart valve through selfassembly of cells [191]. Challenges remain for the use of extrusion-based to produce heart valves due to issues like flexibility and mechanical properties of current printed. Current printers can use only low viscosity solution that may not duplicate valve structure and function. Cross-linking of component materials in low-viscosity solution can be a solution but that works against cell compatibility.


A more functional bioartificial valve was made by Tranquillo and colleagues using cell encapsulated fibrin gel generated in a mold that applied appropriate mechanical constraints to the gel in order to obtain both proper fiber alignment of deposited fibers and proper cusp geometry (Figure 7) [192, 193].

Hydrogels are of particular interest in biomedical applications due to their efficacy in growth factor delivery [168]. Inclusion of TGF-β1 in metalloproteinase degradable PEG hydrogels increased the expression of alpha smooth muscle actin (α-SMA) and collagen-1 after two days of culture indicating differentiation of myofibroblasts [194]. Similar observations occurred with gelatin-methacrylate hydrogels in the presence of TGF-β1 [169]. Incorporation of fibronectin-derived RGD peptides, which promote integrin binding, increased VIC process extension and integrin alpha(v)beta(3) binding [194]. A metalloproteinase (MMP) degradable PEG hydrogel system could be useful for characterizing the transient state of VICs – either fibroblast or myofibroblast – a useful cell type for heart valve tissue engineering [194]. PEG hydrogels have been shown to reduce calcification on engineered leaflet scaffold tissues [195]. Addition of fibrin or fibronectin to PEG hydrogels further reduced calcification. Furthermore, their addition did not change the expression of α-SMA, a marker for myofibroblastic activity.

Promotion of endothelialization of implanted devices is another goal for heart valve tissue engineering. Antibody-modified polysaccharide-based hydrogels are capable of capturing circulating endothelial progenitor cells (EPCs). Applying this technique, Camci-Unal et al. prepared CD34 antibodies immobilized on hyaluronic acid (HA) hydrogels and used them for valve tissue engineering with seeded EPCs [196]. After two days, the cells spread better on modified HA hydrogels compared to unmodified hydrogels. This technique could be useful for improving the biocompatibility of implants including artificial heart valves by endothelialization.


Hydrogels, including protein gels, were found to be effective materials for bioprinting complex valve shapes. The elastomeric flexibility of the printed valves could be adjusted through appropriate crosslinking concentration and time, either by radiation or chemicals. If the cells are encapsulated, then there can be a compromise between crosslinking concentration and time for flexibility and cell survivability. Crosslinked scaffolds can have issues with sustainability in a dynamic environment. Several weeks of static incubation of printed hydrogel scaffolds may increase their mechanical characteristics through cell proliferation and ECM deposition. Ultimately, dynamic conditioning of the cultured valves could align the ECM protein fibers and bring the overall structural morphology closer to that of a native valve. All of these hurdles have not yet been overcome and thus more studies need to be done to bring printed hydrogel scaffolds to clinical trials, at least for in vivo experiments. Tranquillo et al. were successful in generating valves from cell encapsulated protein gel by utilizing a molding method and they obtained almost all the characteristics of native valve including bileaflet shape [192, 193].

Scaffolds of combined morphologies Scaffolds may be generated through the combination of materials with distinct morphologies and/or mechanical properties. Specially, presence of nanofibrous structure in combined scaffolds might provide both nanoscale and microscale structures. Moreover, inclusion of a nanoscale fiber coating on scaffolds showed an increase in mechanical properties compared to original scaffolds (without a nanofiber coating) when both were cultured with mesenchymal stromal cells, although cell proliferation was the same in both scaffolds [110]. Nanofibers present a similar structure as ECM, and this plays a key role in tissue architecture by providing structural


support and tensile strength [34]. Chitosan fibers were produced through solution extrusion and neutralization method and incorporated into chitosan scaffold [163]. This fiber-reinforced engineered heart valve scaffold achieved leaflet tensile strength values of 220 ± 17 kPa, which are comparable to the radial values of native valve leaflets. A nanofiber coating on the scaffold surface increased the strength of the scaffold, which caused higher deposition of collagen fibers leading to increased mechanical properties of the tissue-engineered valves. In some cases, decellularized scaffolds were coated with nanofibrous matrix to improve the mechanical properties and the degradation kinetics of the original scaffolds. Hong et al. coated acellular heart valve scaffolds with poly (3-hydroxybutyrate-co-4-hydroxybutyrate) (P(3HB-co-4HB)) fibers by electrospinning and seeded with mesenchymal stromal cells (MSC). The hybrid scaffold exhibited better mechanical properties, higher cell proliferation, and increased ECM production compared to the acellular biomatrix scaffold [110].

Most of the heart valve tissue engineering substrates including porous 3D scaffolds have been based on single-layer designs because generating appropriate trilayered scaffolds is a daunting task [197-200]. Results of animal studies revealed that, in most cases, only the outer layer was remodeled, indicating that a more sophisticated scaffold design may be necessary. One of the approaches to making a scaffold with a trilayered structure consisted of creating a porous spongiosa layer sandwiched between two collagen layers [30]. The porous spongiosa layer was obtained from decellularized and elastase-treated porcine pulmonary arteries. One outer layer, the ventricularis, was also collected from decellularized porcine pericardium and the other outer layer, the fibrosa, was created using a fibrous collagen scaffold. Both outer collagen layers were treated with PGG to stabilize them against rapid biodegradation without cross-linking. The middle layer spongiosa was seeded with stromal cells derived from human bone marrow and positioned between the two outer layers. The final construct was cultured


in a heart valve bioreactor and subjected to physiological pressure. The seeded stromal cells showed the phenotype of valvular interstitial cells indicating that mechanical loads in the bioreactor specific to an aortic valve induced the required phenotype without the presence of specific growth factors.

In another approach, trilayered scaffolds have been prepared from synthetic materials with appropriate mechanical properties in each layer. In the native heart valve, the two outer layers are stiffer than the soft middle layer. Tseng et al. have prepared such a multilaminate scaffold consisting of three layers from a hydrogel, poly (ethylene glycol) di-acrylate (PEGDA), which had tunable mechanical and biological properties (Figure 8) [186]. The two outer layers were stiffer compared to the middle layer. The mechanical properties of the layers were varied by adjusting the molecular weight and concentration of the polymer. Since all the layers were made of hydrogel, the scaffold system did not fail at the interfaces when subjected to force. This scaffold with a robust interface, integrated layers of different mechanical properties, and biofunctionalization was more appropriate than single-layered scaffolds for heart valve leaflet engineering.

Despite these good results, the trilayered constructs were simply three-layered membrane structure that may not develop into functional leaflets attached to a heart valve conduit. Indeed, functional leaflets from three-layered membrane structures have not been produced yet, possibly due to lack of integrity among the layers or due to the difficulty of molding it into the shape of a valve. Since the native heart valve has anisotropic mechanical properties, fibrous materials could be helpful in preparing scaffolds with anisotropic properties by aligning the fibers in the required directions [148].

Drug delivery capabilities


In the development process of various tissues including heart valve, growth factors are involved at different stages of maturity. In this process, growth factors are crucial for regulating cell migration into and differentiation within scaffolds. As cells within heart valves have limited proliferative capacity, application of growth factors might be used to aid functional tissue engineering in vitro, ex vivo, and in vivo. Examples of this approach include coating of decellularized aortic heart valves with heparin-vascular endothelial growth factor (heparin-VEGF) which not only acted as an antithrombotic agent but also induced adhesion, proliferation, and migration of endothelial progenitor cells on the decellularized valve [201]. In addition, heparin was responsible for a synergistic increase in alpha-smooth muscle actin (SMA) expression in valvular interstitial cells [202]. VEGF acted to increase proliferation of endothelial cells and reduce calcification generated by valvular interstitial cells. In contrast, it was observed TGF-β1 had a tendency in inducing the formation of calcific nodules, which did not form in the presence of both fibronectin and TGF-β1 [203].

The use of nanoscale morphology promotes growth factor release due to its high surface area with respect to volume, i.e. aspect ratio. Taking advantage of this feature, several researchers have modified scaffolds with nanoscale drug carrier coatings for use in heart valve tissue engineering. For example, decellularized matrix scaffolds coated with nanofibers were made by electrospinning a mixture of basic fibroblast growth factor (bFGF), chitosan, and poly-4-hydroxybutyrate (P4HB). The scaffold was cultured with mesenchymal stromal cells for 14 days. The results revealed a substantial increase in cell mass and strength due to 4-hydroxyproline and collagen in the hybrid


heart valve leaflets [110]. In another example, decellularized valves were modified with PEG nanoparticles loaded with TGF-β1. It was observed that in presence of both PEG and TGF-β1, hybrid scaffolds possessed superior biocompatibility, biomechanical properties, and ECM microenvironment compared to unmodified decellularized scaffolds [28]. Growth factors also influence the phenotype of delivered and retained cells [159, 204]. Both human and sheep fibroblasts were differentiated to metabolically active and functional myofibroblasts in the presence of TGF-β1. Similarly, TGF-β1 differentiated ovine aortic valvular endothelial cells and circulating endothelial progenitor cells to a mesenchymal phenotype [159]. VEGF was also applied within a scaffold system to induce seeded stromal cells (e.g., human mesenchymal stromal cells) to differentiate into a endothelial cells [205].

When unseeded scaffolds are implanted, a favorable environment is needed to be created to induce host cells to migrate from the surrounding tissues, be retained and proliferate and differentiate into specific cell types [206-208]. In an in vitro study, Somers et al. employed various growth factor cocktails within unseeded scaffolds to define their ability to induce proliferation, migration, and invasion of ovine mesenchymal stromal cells (oMSCs) [209]. Interestingly, different combinations of growth factors had different impacts on proliferation, migration and invasion. Growth factors may have a negative impact on tissue engineering. Granulocyte colony-stimulating factor (G-CSF) administration accelerated heart valve deterioration to a degree similar to that observed in fresh xenogeneic bioprosthetic valves [210]. Thus, prior to scaffold fabrication, careful planning is needed for the addition of relevant growth factors at different stages of development and maturity to ensure desirable results.


Thrombogenicity Thrombogenicity is one of the major issues with artificial heart valve scaffolds [211]. Thrombosis occurs due to adsorption of blood proteins such as fibronectin, vitronectin, fibrinogen, and von Willebrand Factor onto the scaffold surface causing blood contact activation, platelet activation, and thrombin and fibrin formation in blood plasma [212]. These phenomena are very complicated consisting of a cascade of reactions which can be divided into two main pathways: intrinsic and extrinsic [213]. I n this section, we do not discuss these pathways; rather, we briefly discuss properties of the scaffold biomaterial that cause thrombogenicity and ways to prevent it.

The surface chemistry of scaffold biomaterials, which correlates to material properties such as surface energy, intermolecular force, relative charge distribution on the surface, ionic interaction, and texture, is mainly responsible for protein adsorption leading to thrombogenicity [214].

Blood flow characteristics are also responsible for blood

coagulation to some extent. Proteins are comprised of amino acid subunits that gain or lose charge through their side chains interacting with the surrounding media depending on the pH of the media and which of these charged regions are exposed [214]. Different forces and interactions between the molecules on a biomaterial surface and the charged protein molecules in blood are involved in protein adsorption. As a result, there is a stronger tendency for protein adsorption onto a hydrophobic surface as compared to a hydrophilic surface.











endothelialization of the surface have been applied to improve the hemocompatibility of scaffold surfaces [212]. Anti- or noncoagulant materials such as polyethylene glycol (PEG), pyrolytic carbon, elastin, heparin, thrombomodulin, and phosphorycholine can be used for coating heart valve scaffolds to improve their hemocompatibility [215]. PEG (also polyethylene oxide) has a hydrophilic ether oxygen in its molecular repeat unit which discourages protein adsorption. To achieve a PEG coating, different PEGylation methods including covalent surface, chemisorptive surface, and physisorptive surface can be used [216]. Absorptive surface PEGylation methods are superior to covalent surface PEGylation methods because covalent surface PEGylation methods induce inflammation and lead to thrombosis.

Although all of these anticoagulant materials

showed successful results in laboratory tests, most of them failed in clinical studies. This is especially true for heart valve scaffolds, likely due to the presence of a rigorous dynamic environment [212].

Endothelialization is another potentially useful method to prevent thrombogenicity of implanted valves because a layer of endothelial cells offers a non-thrombogenic surface similar to a native heart valve [217]. Two main methods, in vitro endothelialization and self-endothelialization, have been applied to form a layer of endothelial cells on scaffold surfaces [218]. For in vitro endothelialization, a patient’s endothelial cells are collected and cultured on the scaffold surface to form a layer of endothelial cells. S ome other cells such as fibroblasts and/or smooth muscle cells (SMCs) can be co-cultured with endothelial cells to form a better endothelial cell layer due to higher ECM production


from fibroblasts and SMCs [219]. To improve endothelialization, the surface chemistry of scaffolds can be changed through chemical phenomena such as plasma vapor deposition [220].

Self-endothelialization of implanted heart valve scaffolds is possible if the scaffold materials can induce cell adhesion [218]. Several cell types including endothelial cells and endothelial progenitor cells (EPCs) that circulate in the blood are capable of cell adhesion [221, 222]. To improve cell attachment, the coating of scaffold surfaces with ECM proteins such as collagen, laminin, fibronectin, and vitronectin may be useful [213]. Scaffold surfaces modified with RGD peptides have demonstrated improved cell adhesion. The orientation of immobilized RGD, its spatial arrangement, and ligand density were found to influence endothelial cell adhesion [223-226]. For example, cyclic RGD is more efficient compared to linear RGD for cell adhesion. proteins specific to EPCs are worthwhile for EPC adhesion.

Besides RGD,

However, due to the

presence of different proteins on scaffold surfaces, unwanted cells may adhere to scaffold surfaces and differentiate into inflammatory cells [212]. Therefore, coating of scaffolds should be performed with appropriate proteins for adhesion of endothelial cells and/or EPCs only.

Bioprosthetic valves that are currently used to replace diseased valves are mildly thrombogenic and anti-coagulants and antiplatelet agents (such as warfarin and aspirin) are prescribed for some period of time after implantation [227]. Although decellularized heart valves lose some proteins during the cell removal process, they retain most major


proteins including collagen.

However, thrombosis is a significant concern when

decellularized heart valves are implanted for in vivo studies [228, 229]. Fabricated heart valve scaffolds that are mainly made of polymeric materials are also thrombogenic depending upon the surface chemistry and texture.

Similar to bioprosthetic valves,

these fabricated scaffolds require anti-coagulants when implanted for in vivo studies [230].

Among all of the strategies for anti-thrombogenicity, endothelialization is

currently seen as the best option for all types of heart valve scaffolds, including decellularized valves, because this most closely resembles the native strategy.

Conclusions and future directions Tissue engineered heart valves are a promising alternative option to current mechanical and bioprosthetic valves. Constructing a heart valve requires a three-dimensional scaffold in which cells can grow, proliferate, and differentiate into a functional tissue construct. Researchers have attempted to find effective approaches to scaffold fabrication. Scaffolds should have appropriate morphology with trilayered structure, mechanical properties with bipolar stiffness (15.34 and 1.98 MPa in circumferential and radial directions respectively), and include expression (or delivery) of growth factors and generation of ECM for effective heart valve tissue engineering. Both decellularized and fabricated scaffolds have advantages and disadvantages. Decellularized scaffolds mostly retain the original valve structure and extracellular matrix molecules, which is a unique advantage over fabricated synthetic scaffolds. Their mechanical stiffness is close to native valve. However, their low pore size, porosity, and cell survivability limit their potential for valve engineering. Similarly, most of fabricated 3D solid porous


scaffolds, fibrous scaffolds, and hydrogel scaffolds cannot properly mimic the trilayered structure of native heart valves, although, there were some early studies of trilayered scaffolds. The mechanical properties and morphologies of 3D solid porous scaffolds can be tailored according to requirements. Both decellularized and fabricated scaffolds are able to deliver drugs. Building upon past achievements, we need to move forward to optimize and test scaffold systems that can be applied to heart valve tissue generation.

While the focus of this review was on scaffolds, any successful strategy will need to carefully consider whether cells will be added prior to implantation or simply allow for ingrowth following implantation. If direct application is utilized the source, methods to isolate, culture and deliver these cells will need to be optimized. Finally, whether the construct is conditioned or tested in vitro in a bioreactor must be considered. We anticipate that the coordinated efforts of engineers, biologists, surgeons and imagers will be required to achieve the goal of a clinically available tissue engineered heart valve.

Acknowledgements: This work is supported by the HH Sheikh Hamed bin Zayed Al Nahyan Program in Biological Valve Engineering, the Grainger Foundation, and the Mayo Clinic Center for Regenerative Medicine. Disclosures: None.


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Figure Legends Figure 1: Schematic diagrams of aortic heart valve tissue engineering. Living cells are grown onto a supporting three-dimensional (3D) biocompatible structure to proliferate, differentiate, and ultimately grow into a functional tissue construct. Figure 2: Schematic diagrams of an aortic heart valve. (a) Side section view of aortic valve in the heart. (b) Top view of aortic valve (seen from aortic side). (c) Side view of splayed-open valve showing semilunar shape of cusps. (d) Section view through the cusp and aortic wall showing the three-layered structure of the cusp. (e) Three cusp layers - fibrosa, spongiosa, and ventricularis showing their collagen fibril orientations.

Figure 3: Photographs of porcine aortic heart valves. (a) Native. (b) Decellularized.

Figure 4: Tissue engineered aortic valve from PHO polymer-based porous scaffold. (a) Photographic image of aortic valve after 5 weeks in-vivo implantation. SEM images of leaflet surfaces after (b) 1 week, (c) 5 weeks, and (d) 17 weeks of implantation. (e) Hematoxylin & eosin (H&E) stained conduit walls showing ingrowth of vascularized tissue islands and destruction of polymer (filled arrow). (f) Movat pentachrome stained leaflet showing significant collagen deposition (open arrow) [129]. Permission to reuse figure will be requested upon acceptance.

Figure 5: Nanofibrous scaffold for heart valve tissue engineering. (a) SEM image of continuous fibroblast/connective tissue on nanofibrous scaffold. (b) CD34 stain demonstrating the formation of an endothelial monolayer (A) on the surface of a core of fibroblasts (B) and hydrolysis of polymer fibers after 4 weeks (C) [154]. Permission to reuse figure will be requested upon acceptance.

Figure 6: Bioprinting of scaffolds for heart valve tissue engineering. (a) 3D printer setup showing the printing mechanism. (b) Porcine heart valve isolated for micro-CT scan. (c) Computerized 3D valve reconstruction from CT scan data. (d) Computerized slice of the valve reconstruction showing the calculated print head paths. (e) Printed valve where the root was formed with 700 MW PEG-DA hydrogel and the leaflets were formed with 700/8000 MW PEG-DA hydrogel. Scale bar = 1 cm [189]. Permission to reuse figure will be requested upon acceptance.

Figure 7: Photograph of mold and bileaflet valve after 6 weeks of compaction. (a) Side view of the bileaflet valve mold. (b) Top view of a valve in static conditions. (c) Ventricular view of a valve under pressurized conditions

53 (greater than 5 mmHg) showing coaptation of the leaflets [192]. Permission to reuse figure will be requested upon acceptance.

Figure 8: Schematic diagram depicting the fabrication of trilayer quasilaminates. Gel A represents 12.5% 3.4 kDa PEGDA and gel B represents 10% 6 kDa PEGDA [186]. Permission to reuse figure will be requested upon acceptance.

Table 1. Mechanical properties of PHAs.

Tensile strength






Native Cusp in

Native Cusp







in radial


















15.34 (MPa)

1.98 (MPa)








(MPa) Tensile modulus (GPa) Elongation at break (%)

# NA stands for not available

*Table was modified from Williams et al. [128].

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Scaffolds for tissue engineering of cardiac valves.

Tissue engineered heart valves offer a promising alternative for the replacement of diseased heart valves avoiding the limitations faced with currentl...
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