Drug delivery in aortic valve tissue engineering.

COREL-07406; No of Pages 19 Journal of Controlled Release xxx (2014) xxx–xxx

Contents lists available at ScienceDirect

Journal of Controlled Release

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Review

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Drug delivery in aortic valve tissue engineering

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Soumen Jana a, Robert D. Simari b, Daniel B. Spoon a, Amir Lerman a,⁎

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Article history: Received 5 September 2014 Accepted 9 October 2014 Available online xxxx

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Keywords: Drug delivery Fiber Heart valve Hydrogel Scaffold Tissue engineering

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Over the last 50 years medicine and technology have progressed to the point where it has become commonplace to safely replace damaged or diseased heart valves with mechanical and biological prostheses. Despite the advancements in technology current valve substitutes continue to have significant limitations with regards to thrombogenicity, durability, and inability to grow or remodel. In an attempt to overcome the limitations of currently available valve prosthesis, heart valve tissue engineering has emerged as a promising technique to produce biological valve substitutes. Currently, the field of tissue engineering is focused on delivering complex matrices which include scaffolds and cells separately or together to the damaged site. Additional functional enhancement of the matrices by exposing encoded biological signals to their residing cells in a controlled manner has the potential to augment the tissue engineering approach. This review provides an overview of the delivery of biological reagents to guide and regulate heart valve tissue engineering. © 2014 Elsevier B.V. All rights reserved.

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Introduction . . . . . . . . . . . . . . . Aortic valve . . . . . . . . . . . . . . . 2.1. Structure . . . . . . . . . . . . . 2.2. Development . . . . . . . . . . . 2.3. Molecular regulation in development Aortic valve disease and treatment . . . . . Drug carriers . . . . . . . . . . . . . . . 4.1. Materials for carriers . . . . . . . . 4.1.1. Natural polymers . . . . . 4.1.2. Synthetic polymers . . . . 4.1.3. Cells . . . . . . . . . . . 4.2. Design of carriers . . . . . . . . . 4.2.1. Substrate-based . . . . . . 4.2.2. Cell-based. . . . . . . . . Drug delivery for aortic valve generation . . 5.1. Delivery of growth factors . . . . . 5.2. Delivery of genes. . . . . . . . . .

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Division of Cardiovascular Diseases, Mayo Clinic, 200 First Street SW, Rochester, MN 55905, USA University of Kansas School of Medicine, 3901 Rainbow Blvd, Kansas City, KS 66160, USA

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Abbreviations: α-SMA, α-smooth muscle actin; AVHD, aortic valvular heart disease; bFGF, basic fibroblast growth factor; BMP, bone morphogenetic protein; CS, chondroitin sulfate; EC, endothelial cell; ECGF, endothelial cell growth factor; ECM, extracellular matrix; EMT, transformation of epithelial cells to mesenchymal cells; eNOS, endothelial nitric oxide synthase; EPC, endothelial progenitor cell; FGF, fibroblast growth factors; GA, glutaraldehyde; GAG, glycosaminoglycan; HA, hydroxyapatite; hDMC, human dermal mesenchymal cell; HGF, hepatocyte growth factor; hMSC, human mesenchymal cell; HSP47, heat shock protein 47; IGF, insulin-like growth factor; MSC, mesenchymal stem cell; NO, nitric oxide; NRG-I, neuregulin-I; PAA, poly(amido amine); PCL, polycaprolactone; PEG, polyethylene glycol; PEGDA, poly(ethylene glycol) diacrylate; PEI-CD-DNA, poly-ethyleneimine-plasmid DNA; PGA, polyglycolic acid; PGS, poly(glycerol sebacate); PHA, polyhydroxyalkanoate; PHB, polyhydroxybutyrate; PHBV, poly(hydroxybutyrate-co-valerate); PLA, polylactic acid; PLGA, poly(lactic-co-glycolic acid); PVA, polyvinyl acetate; REDV, Arg-Glu-Asp-Val; RGD, Arg-Gly-Asp; RGDS, Arg-Gly-Asp-Ser; SDS, sodium dodecyl sulfate; SMC, smooth muscle cell; Tβ4, thymosin β4; TGF-β, transforming growth factor-beta; TNC, tenascin-C; VEGF, vascular endothelial growth factor; VIC, valvular interstitial cell. ⁎ Corresponding author at: Division of Cardiovascular Diseases, Mayo Clinic, Rochester, MN 55905, USA. Tel.: +1 507 255 2398. E-mail address: [email protected] (A. Lerman).

http://dx.doi.org/10.1016/j.jconrel.2014.10.009 0168-3659/© 2014 Elsevier B.V. All rights reserved.

Please cite this article as: S. Jana, et al., Drug delivery in aortic valve tissue engineering, J. Control. Release (2014), http://dx.doi.org/10.1016/ j.jconrel.2014.10.009

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6. Conclusion . . . . 7. Future perspective Disclosures . . . . . . Acknowledgments . . . References. . . . . . .

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1. Introduction

2. Aortic valve

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A large number of people suffer from aortic valvular heart disease (AVHD) worldwide leading to cardiovascular morbidity and mortality [1,2]. In the United States alone, approximately 5 million people are affected by AVHD every year [3]. Although, number of deaths from AHVD is relatively small (~ 15000) 30% of people (age N 65) suffer from valvular sclerosis and 10% of them face stenosis with degenerative pathology being the predominant cause. [4]. The incidence of AVHD is much higher in developing countries secondary to the persistent burden of rheumatic fever where more than 15 million people worldwide are estimated to be have rheumatic valve disease [5,6]. Due to the hemodynamic consequences of valvular heart disease, systemic blood flow can become compromised leading to morbidity and mortality [7]. Valvular tissue has limited regenerative capacity, and in case of damage or loss, scar tissue forms and calcification occurs leading to valvular impairment [8]. At present, the only effective therapy for severe valvular stenosis is the replacement of the affected valve with mechanical or biological prostheses [7]. In the next thirty years, the number of aortic valve surgeries will be tripled due to increase in population and improved access to healthcare [9]. However, both biologic and mechanical prosthetic valves have significant limitations and degree of survival and clinical improvement after surgery depends on the type of aortic valve substitute used [10]. Mechanical valve prosthesis requires life-long anticoagulation therapy (currently with Vitamin K antagonists) to decrease the risk of blood clotting on the valve surface which can result in valve dysfunction and thromboembolism [11]. On the other hand, biological prostheses have a low risk of thromboembolism but have limited durability with a significant failure rate at 10 years [12]. Biologic valves are generally treated with glutaraldehyde (GA), sodium dodecyl sulfate (SDS), and other substances to reduce their antigenicity and to prevent calcification. These treatments stiffen the prostheses and disrupt the collagen and glycosaminoglycan (GAG) structure leading to loss of functionality contributing to their decreased durability [13]. Importantly, neither mechanical nor biological prosthesis has the capacity to grow or remodel and thus are even less optimal for pediatric patients [14]. The field of regenerative medicine offers an alternative approach in which living cells, biomaterials, and soluble mediators are employed to create tissue with normal structure and function at the damaged site [15]. In native tissue development, including aortic heart valve generation, multiple growth factors are involved at different stages of growth and maturity [16]. In tissue restoration in vitro or in vivo, similar kinds of growth factors can be delivered in a controlled manner at the right phases to obtain a functional heart valve construct. Moreover, if the cells at the damaged site are not active or have low concentration, the delivery growth factors as well as specifically transfected cells, a new tissue could potentially be regenerated or damaged tissue could be remodeled [17]. This review summarizes comprehensive information on drug delivery in aortic heart valve tissue engineering. The review first discusses the structure, development, and molecular regulation in the development of aortic valve as well as the disease that cause the damage or impairment of aortic valve function. It then addresses design of drug carriers and materials used to prepare them. Lastly, it describes the delivery of key biomolecules including growth factors, genes and cytokines in heart valve generation through different approaches including substrate-based and cell-based deliveries.

The aortic valve is located central to the pulmonary, mitral, and tricuspid valves and between the left ventricular outflow tract and the ascending aorta [18]. It allows the blood flow from the left ventricle into the aorta during systole and prevents flow in the opposite direction during diastole. The aortic valve is one of the two semilunar heart valves—the other being the pulmonary valve.

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2.2. Development

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Embryologically, the heart is the first organ to form and supports continuous growth of the embryo. Initially, the heart is a primitive tube consisting of myocardial cell layer surrounded by endocardial EC layer [21]. At the fourth week of gestation, dextrosuperior and sinistroinferior endocardial cushions and two intercalated endocardial cushions located 90° from the previous ones form in the distal portion of outflow tract [25]. The endocardial cushion formation is characterized by transformation of epithelial cells to mesenchymal cells (EMT) in the presence of signaling factors emanated from myocardium cells [26]. At this stage, mesenchymal cells are highly proliferative and loosely connected in the cardiac gel (ECM), although in remodeling or mature valves, they recycle very little [27]. Mesenchymal cells are responsible for the creation of valvular interstitial cells (VICs). The cardiac gel, which contains proteoglycans, glycosaminoglycans, and other structural proteins, provides the morphology of the endocardial cushions and other parts of the heart [28]. Rhythmic contraction and expansion of extracellular matrix (ECM) gel in the endocardial cushions acts as primitive regulatory controls allowing for unilateral blood flow [29]. In a subsequent stage, the endocardial cushions fuse to form valve primordia, which arise from truncal septum. The valve primordia become thinner and elongate to form valve cusps. At this stage, the ECM goes through major changes in the form of patterning evidenced by gene expression from the surface of the valve. The patterning of the ECM is influenced by direction of blood flow. In the cusps, the ECM remodels into three layers

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The aortic valve consists of three cusps: left coronary, right coronary, and non-coronary— named according to their relationship with the coronary arteries (Fig. 1a) [19]. The aortic valve cusps are supported by the aortic valve annulus and commissures. The valve is connected to the heart muscle through an annulus, which is comparable to a tendon that links skeletal muscle and bone. Cusp thickness is generally less than 1 mm, although it varies by region—thicker at the base and tip compared to other areas [20]. Each cusp consists of complex stratified connective tissues forming three layers: fibrosa, spongiosa, and ventricularis (Fig. 1b) [21]. The fibrosa, located nearest the aorta, is composed of circumferentially oriented fibrillar collagens (types I and III), which provide tensile stiffness to the valve [22]. The middle layer (spongiosa) consists of proteoglycans interspersed with collagen fibers. This layer works as a cushioned interface between the two outer layers to provide compressibility and integrity to the valve. The ventricularis, named because of its continuity to the ventricle, is composed of radially oriented filamentous elastic fibers and enables extension and recoiling of the valve under diastole and systole pressures, respectively. The cusps are covered with a layer of endothelial cells (ECs) with fibroblast/myofibroblast like interstitial cells inside [23].

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enriched with proteoglycans (spongiosa), fibrillar collagen (fibrosa), and elastin (ventricularis) [23].

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As above, there are three main stages of development: 1) formation of endocardial cushion and occurrence of EMT, 2) development of valve primordia after the growth of endocardial cushion and 3) creation of valve cusps through elongation and thinning of valve primordia. Through gene analysis using the embryos of mice, zebra fish, and human, the molecular regulation in each stage of development has been elucidated [16,21,23]. Inception of endocardial cushion formation and EMT occurrence is regulated by the superfamilies of transforming growth factor-beta (TGF-β) cytokine including bone morphogenetic proteins (BMPs) and TGF-βs [30]. Their receptors and ligands are also expressed during EMT. TGF-β induces the expression of Slug transcription factors through signaling of SMADs 2/3 resulting in endocardial cushion EC invasion [31]. TGF-β takes part in signaling of Wnt/β-catenine which promotes the EMT occurrence, growth of cushion and formation of valve primordia [32]. In the developing chicken and mouse, it has been demonstrated that hyaluronan and versican deposition that are controlled by BMP-2 and BMP-4 and occur during cushion formation in the outflow tract region [30]. Expression of Tbx2 in the myocardium coincides with BMP-2 expression to regulate the suppression of chamber-specific gene expression and ECM deposition in the region [33]. Growth of the endocardial cushions at the end of EMT and fusion of cushions to form valve primordia is controlled by appropriate level of BMP signaling [34]. Fibroblast growth factors (FGFs) family and their receptors also take part in these developments [35]. Endothelial progenitors influence the proliferation cushion ECs through the expression of vascular endothelial growth factor (VEGF) for cushion growth; however, VEGF works against EMT [23]. Notch signaling induces Snail transcription factor, which represses VE-cadherin expression and stimulates EMT. VEGF/NFATc1 is found responsible in valve primordia development [36]. In addition to these cytokines,

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Fig. 1. Images of aortic valve with trilayered structure. (a) Photograph of cadaverous aortic valve (cusps pointed by black arrows). (b) Photograph of dissected valve showing cusps (pointed by black arrows) and sinuses (pointed by white arrows) in an aortic valve. (c) Immunostained image of trilayered structure of an aortic leaflet consisting of fibrosa (F), spongiosa (S) and ventricularis (V) [24]. Reprinting permission will be obtained upon acceptance.

several other signals including Sox9, SHP2, ERK1/2, Msx1/2, cadherin11, and Twist1 also take part in primordia development [37]. During cusp formation through elongation and thinning of primordia, notch signaling occurs due to shear stress on valve at the flow side, i.e. ventricularis, and causes valve polarity and stratification [38]. In the spongiosa, ECM structure is influenced by BMP2 signaling, which promotes expression of transcription factor Sox9 and structural proteins collagen type II, alpha1 and aggrecan [37]. Although less is known about the gene regulation on fibrosa layer development, Wnt signaling is mainly responsible for expression of osteogenic-related ECM proteins such as osteonectin, periostin, and fibronectin in fibrosa layer [39]. During the maturation of cusps and their supporting structure, signaling of FGF4 for the expression of different transcription factors and ECM proteins have been identified as important steps [40].

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Pathological abnormalities in aortic valve arise when valve fails to perform routine systolic and diastolic flows that depend upon the mechanical integrity and pliability of the cusps [41]. In other words, scars and calcification stiffen the cusp, resulting in development of abnormalities. Stenosis and regurgitation are the major pathological abnormalities in aortic valve (Fig. 2) [42]. Stenosis occurs when valve orifice narrows and obstructs blood flow during systole cycle [43]. In contrast, regurgitation occurs due to incomplete closure of cusps, causing the backward blood flow during diastole cycle. Congenital defects in valve structure, involving either with one cusp or two cusps (unicuspid, bicuspid), lead to altered hydrodynamic and mechanical stress distributions [44]. The consequence is calcification and gradual cusp degeneration leading to stenosis and regurgitation in childhood and young adulthood. In addition to deviations in valve structure and stiffness, other causes of valve disease may be systemic infectious and inflammatory conditions. Rheumatic fever — the most common systemic inflammatory condition triggered by bacterial infection (Streptococcus pyogenes) — could initiate heart valve disease [45]. Infection on cusp surface may originate from the bacteria present in circulating blood [46]. In response to this

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4.1. Materials for carriers

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Polymers are gaining importance in pharmaceutical application, including drug delivery, due to their desirable physical and chemical properties such as lack of leachable impurities, chemical inertness, minimal undesired aging, and easy processing [55]. Although both natural and synthetic polymers have been applied in drug delivery carriers, natural polymers are preferred due to their nontoxic biodegradability and bioresorbability [53]. In addition to polymers, some inorganic materials such as tri-calcium phosphate, hydroxyapatite (HA), and semi-synthetic materials such as cross-linked thiolated HA and esterified hyaluronan are used in scaffold forms for drug delivery [54]. Cells also have been applied as carriers in drug delivery. For this purpose, cells are transfected through viral and non-viral techniques and delivered to the engineering site either in vitro or in vivo where cells express the required genes and produce growth factors [56,57]. Here, we discuss the several materials that have been used for drug delivery to or have promise for heart valve tissue engineering.

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Drug carriers consist of biomaterials, cells and viruses that can be used to deliver the drugs or cytokines in a controlled manner to the

4.1.1. Natural polymers Natural polymers are from biologic sources and similar to biomacromolecules recognized by our body, they are useful as drug carriers [58]. Although, natural polymers are biocompatible and biodegradable with

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target site with limited drug metabolism and toxicity. The goal is not only to influence host healing response the damaged site, but also to support the stimulation of tissue growth for functional tissue construct formation [53,54]. Therefore, the carriers for drug delivery should have specific characteristics to allow for: 1) satisfactory biocompatibility and toxicity, 2) appropriate binding and loading capacity, 3) not damage the drug structure and efficacy when drug is incorporated into carrier, 4) allow for uniform drug dispersion throughout the carrier, 5) protect drug in physiological condition, and 6) release the drug controllably over a given period of time [55].

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infection, autoimmune reactions include several complex biological activities resulting in inflammation, scarring, degeneration, bacterial vegetation and adverse remodeling [47]. Ultimately, stenosis occurs as stiffness of the cusps increases due to thickening, fibrous matrix production and calcification to protect from dynamic environment [7]. Heart valve degeneration may occur from mechanical fatigue of the cusps due to hemodynamic stress and systole–diastole cycle repeated approximately 3 billion times over an average lifespan [49]. To protect valve from degeneration and its malfunction, cusps self-cure through calcification and fibrosis. In reality, these valve abnormalities are the sum of various complicated processes including manifestation of various residing cells and different bioactive molecules (Fig. 3). Several clinical symptoms including shortness of breath (dyspnea), fatigue, dizziness, chest pain and presyncope/syncope are observed when heart valve degeneration occurs [50]. Turbulent blood flow in valve due to stenosis and regurgitation create a sound (murmur) that can be detected in a physical test [51]. Echocardiography can be applied to confirm it and further interpret the valve condition [45]. Cardiac catheterization, an invasive procedure, is used to evaluate and treat aortic valve disease [52]. Although, some circulating biomarkers related to valve defectiveness can be applied to identify valve disease and to measure its severity, they are not used clinically [7]. Unless the valve disease is severe, patients may remain relatively asymptomatic; thus, monitoring of patients diagnosed with aortic valve disease is important. At present, valve replacement through open-heart surgery is the only effective treatment for severe disease. Unfortunately, no pharmacological drugs are available to cure valve stenosis and regurgitation [7]. Due to limitations of valve replacements, heart valve tissue engineering is a promising alternative.

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Fig. 2. Aortic valve disease. (a) Photograph of calcified aortic valve. Echocardiographic appearance (Parasternal short-axis views) of aortic stenosis in: (b) calcific degenerative disease, (c) bicuspid aortic valve, (c) rheumatic disease [43,48]. Reprinting permission will be obtained upon acceptance.


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comparatively low or no toxic waste, they are limited by low stability, low yield in some cases, variation in physical properties due to age of the source as well as variation of chain length of the polymer molecules. Living organisms are able to synthesize a variety of polymers in several chemical structure groups: protein-based, polysaccharides, and polyhydroxyalkanoate-based polymers. 4.1.1.1. Proteins. Proteins are large biological macromolecules that contain amino acid residues. Protein-based polymers mimic many features of extracellular matrix—the natural medium for cell proliferation, migration, and differentiation and thus hold promise for tissue engineering application through scaffolds and drug delivery vehicle systems [53]. In several forms such as gel, foam, nanofibers and nanoparticle, protein-based polymers are used for drug delivery and their gel, foam and nanofibrous forms can also be used as scaffold substrate. A brief description of different protein-based polymers that can be applied for drug delivery follows. 4.1.1.1.1. Collagen. Collagen has been used for tissue engineering including drug delivery intended for heart valve tissue engineering [53,59]. Collagen has good biocompatibility, low antigenicity, and high mechanical strength [60]. Growth factors and other active agents are generally combined into the collagen as either scaffold or gel to extend slow release to obtain maximum therapeutic effect on tissue engineering in the desired site [53,61]. As the collagen degrades and releases the drug in its surrounding area. Therefore, to tailor the collagen degradation, i.e., release of drug, cross-linking of collagen is necessary and the type of application including the microenvironment must be considered. As an example, penta-galloyl glucose—a collagen-binding polyphenol treated collagen scaffold showed slow and controlled degradation [62]. Thus,

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Fig. 3. Schematic diagram on participation of various cells in calcification and fibrosis formation in aortic valve.

the release of conjugated drugs can be precise by meticulously controlling the collagen gel degradation. After three weeks of implantation, rhBMP-2soaked collagen sponges induced osteoblastic differentiation in muscle tissue [63]. Controlled release of thymosin β4 (Tβ4) from collagenchitosan hydrogel over one month showed increased migration of ECs and tube formation from epicardial explants [64]. 4.1.1.1.2. Fibrin. Fibrin is natural protein polymer produced during the blood coagulation cascade [65]. It is a complex polymer network formed by polymerization of fibrinogen in the presence of thrombin. It is generally used for hemostasis in surgical procedures and wound healing applications. In addition to application as a matrix component in tissue engineering, it has been applied in drug delivery applications in heart valve generation [66,67]. Fibrin can be applied as gel matrix, nanostructured particles, and cross-linked scaffold systems [65,68]. Fibroblast and smooth muscle cell (SMC)-laden fibrin gel was applied to produce tricuspid heart valve through molding system [69]. Drugs can be conjugated into fibrin gel systems before mixing with cells to improve the functionality of engineered tissues. For example, delivery of BMP-2 through heparin-conjugated fibrin showed excellent efficacy in sustained release of BMP-2 over a period of time [70]. Nanostructured fibrin tube with a length of 700 nm, a diameter of 150–300 nm, and a wall thickness of 50 nm demonstrated 66% drug encapsulation capacity and sustained but complete drug release over one week at the physiological pH of 7.4 [71]. 4.1.1.1.3. Fibronectin. Fibronectin is a high-molecular weight glycoprotein of the extracellular matrix and binds cells through membranespanning receptor proteins called integrins [72]. It was observed that in response to injury, secretion of fibronectin from cardiac valve interstitial cells occurred and formed fibrillar adhesions [73]. Fibronectin is mainly


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4.1.1.3. Polyhydroxyalkanoates. Polyhydroxybutyrate (PHB) is among the group of natural polyesters with the generic name polyhydroxyalkanoates (PHA). They are produced by a variety of microorganisms (bacteria) in nature as a part of their survival mechanisms [118]. Although, various natural PHA polymers are available, PHB and poly(hydroxybutyrate-co-valerate) (PHBV) are mainly used in tissue engineering and drug delivery [119,120] and can be produced by chemical synthesis. They are semicrystalline and thus the degree of crystallinity influences the rate of degradation, drug compatibility, drug diffusion, and drug release [121]. PHA family members are soluble in organic solvent; their solubility depends on molecular weight, crystallinity, and surrounding temperature [122]. They degrade naturally, especially in hydrolytic and enzymatic environment, mainly through surface

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deacetylation of chitosan generally used for biomedical application is usually 70–85% with a molecular weight in the range of 10–1000 kDa [103]. Chitosan is comprised of copolymers of β(1 → 4)-glucosamine and N-acetyl-D-glucosamine—the ingredients found in ECM [104]. Chitosan has amino and hydroxyl groups that are important for chemical modification and attachment of biomacromolecules [105]. It has biocompatible, biodegradable, non-immunogenic, and nontoxic properties suitable for biomedical applications including heart valve tissue engineering and drug delivery [106]. It can be fabricated into nanoparticles, nanofibers, micropores, gel, and 3D solid porous scaffolds for heart valve tissue engineering and all of which can be used as drug carrier [105]. Heart valve scaffolds with molded leaflets made of chitosan have shown improved functionality when the fabricated scaffold was wrapped with chitosan nanofiber due to an increase in mechanical properties [106]. The advantage of this scaffold system is that drug delivery can occur from the scaffold surface or from the scaffold through diffusion and degradation or both as necessary. 4.1.1.2.3. Alginate. Alginate is naturally derived polysaccharide block with copolymers consisting of M-blocks (β-D-mannuronic acid monomers), G-blocks (α-L-guluronic acid), and interspersed M and G blocks [107]. It is found in marine brown algae as structural components and in some soil bacteria as capsular polysaccharide [53]. In the presence of divalent-cations such as Ca2+, bonds form between G-blocks of adjacent chains and alginate hydrogel forms [108]. In addition, nanoparticles, nanofibers, micropores, gel, and 3D solid porous scaffolds can be prepared from alginate alone or with other materials. Alginate in combination with gelatin has been used for heart valve scaffold fabrication [80] . Alginate scaffolds alone produced compatible environment for cardiac cells intended for cardiac patch generation [109]. Different growth factors such as VEGF, endothelial cell growth factors (ECGFs), and basic fibroblast growth factor (bFGF) have been delivered through alginate carriers during tissue engineering [110]. At low pH solutions, release of biomolecules from alginate decreases and this phenomenon can be used for drug delivery. Through 3D bioprinting, a living (cells included) alginate-based heart valve conduit has been produced [80]. 4.1.1.2.4. Chondroitin sulfate. Chondroitin sulfate (CS) is a GAG made of repeating disaccharide units D-glucuronic acid and N-acetyl galatosamine [111]. Due to its GAG nature, CS is often used for drug delivery and heart valve tissue engineering [112,113]. CS can bind different proteins and is used for different biomedical applications. As CS is easily dissolved in water, other polymers such as chitosan, hyaluronan, collagen, poly(vinyl alcohol), and ploy(lactic-co-glycolic acid) can be added to CS to produce more stable materials [114,115]. Moreover, use of crosslinking materials on dissolved CS can control the physical properties of hydrogels and is suitable for various applications [116]. Because CS is negatively charged, attaching a positively charged biomacromolecules is the predominant method for delivering drugs [115]. Collagen scaffolds in combination with CS closely resemble native heart valve structures and have similar properties [117]. Additionally, the presence of CS in collagen scaffolds improved the porosity which can help with better nutrient flow and facilitate a more effective release of drugs [112].

T

372 373

C

370 371

4.1.1.2. Polysaccharides. Polysaccharides are natural, stable, nontoxic, hydrophilic, and biodegradable biomaterials [87]. They have different derivable groups on molecular chains and thus they can be modified chemically. Most polysaccharides contain hydrophilic functional groups such as hydroxyl, amino, and carboxyl, which assist the polymers with drug delivery by attaching the biomacromolecules to those functional groups [88]. 4.1.1.2.1. Hyaluronic acid. Hyaluronic acid (HA) is a linear polysaccharide polymer that has been used for various medical applications including wound healing, tissue augmentation, and ocular surgery for its biocompatible, biodegradable, non-immunogenic, and nontoxic properties [89]. It is distributed throughout the extracellular matrix and organs. HA has received significant interest for use in drug delivery materials development due to its mucoadhesive properties [90]. It has its own binding capabilities to drugs and to cell surface receptors such as CD44 and RHAMM. During valve development, endocardial cushion cells migrate from endocardium to myocardium in the presence of HA [91]. HA is conjugated with other polymers to augment the sustained release properties and therapeutic effect [92]. Both parenteral and non-parenteral methods are applied for sustained release of protein and gene, and drug delivery [93,94]. It can be fabricated as nanocarriers, micelles, film, hydrogel, and 3D solid porous form for drug delivery applications [95,96]. For cardiac valve development, HA is used as scaffold material to prepare complex heart valve structures by applying photo-polymerization cross-linking method [97,98]. Results from synergistic relationship between HA and VICs have shown promise for the use of HA a heart valve drug delivery material [99,100]. 4.1.1.2.2. Chitosan. Chitosan is a cationic polymer that is produced from chitin by deacetylation [101,102]. Crustaceans, cuticles of insects, and cell walls of fungi are the major sources of chitin. The degree of

E

368 369

R

366 367

R

364 365

O

362 363

C

360 361

involved in cell adhesion, migration, growth, and differentiation in the normal function of vertebrate organisms. Despites these capabilities, other growth factors are combined with fibronectin to provide enhanced growth in tissue engineering. VEGF was mixed with fibronectin to deliver of VEGF for therapeutic angiogenesis [74]. The results showed that a VEGF-fibronectin mixed protein was less harmful than VEGF alone, and controlled delivery of VEGF through fibronectin was more effective in therapeutic angiogenesis. Fibronectin scaffolds designed to bind soluble growth factors could stimulate bone and wound healing and repair [75]. Growth factor fused fibronectin proteins were used to coat scaffolds for different aspects of tissue regeneration including cardiovascular applications [76,77]. Fibronectin–hepatocyte growth factor combinations improved the re-endothelialization of decellularized biologic scaffolds [78]. 4.1.1.1.4. Gelatin. Gelatin, a naturally occurring polymer derived from collagen, is generally used in medical and pharmaceutical applications due to its biocompatibility and biodegradability [53,79]. Owing to possibility of polyion complexation formation, gelatin is being used in drug delivery for different types of tissue engineering including cardiac tissues [80,81]. Complex heart valve structure with trileaflets was printed from VIC and SMC-laden gelatin/alginate gel and showed high cell viability, good cell spreading and appropriate phenotype expression. Although gelatin scaffolds in other structures including sponge and foam can be useful for tissue engineering, in drug delivery, the scaffold has been applied mainly in two ways: microsphere and coating [61]. The microspheres containing biomacromolecules such as BMP-2, VEGF are normally incorporated into another scaffold such as hydrogel for drug delivery [82,83]. By including cell adhesion peptides, gelatin microspheres are inserted in 3D porous scaffold system for in vitro and in vivo tissue engineering [84]. Through cross-linkers, gelatin hydrogels are prepared and used for drug delivery directly [85]. Other proteinbased polymers such as elastin and glycosaminoglycans are also applied to some extent for drug delivery applications [86]. As they have limited applications in drug delivery application of heart valve tissue engineering, they are not considered in this review.

N

359

U

358


E

6


424 425 426 427 428 429 430 431 432 433 434 435 436 437 438 439 440 441 442 443 444 445 446 447 448 449 450 451 452 453 454 455 456 457 458 459 460 461 462 463 464 465 466 467 468 469 470 471 472 473 474

477 478 479 480 481 482 483 484 485 486 487


506 507 508 509 510 511 512 513 514 515 516 517 518 519 520 521 522 523 524 525 526 527 528 529 530 531 532 533 534 535 536 537 538 539 540 541 542 543 544 545 546 547 548 549

4.1.2.1. Polyesters. Polyester is a collection of polymers that contain an ester functional group in a main chain [127]. Polyglycolic acid (PGA), polylactic acid (PLA), their copolymer poly(lactic-co-glycolic acid) (PLGA), polycaprolactone (PCL), and poly(glycerol sebacate) (PGS) are some examples of biocompatible and biodegradable polyesters that have been used in tissue engineering, drug delivery, and absorbable implants. The advantage of these polymers is that their degradability can be controlled as needed for drug delivery [128,129]. The degradability rate depends on molecular weight, crystallinity, ratio of copolymers, and site of implantation. However, they have some disadvantages: hydrophobicity, acid formation during degradation, and self-accelerated degradation. To improve the hydrophilicity of these polymers, PEGylation method is often used to attach PEG, which is a biocompatible nontoxic synthetic polymer. PEG also modifies the conformation of copolymers and reduces protein adsorption [130,131]. PEGylated-PLGA microparticles containing VEGF have demonstrated successful long term drug delivery [132].

4.1.2.2. Amine-based. Amine group containing polymers are useful in drug delivery because the amine groups trigger protein/drug adsorption on their delivery vehicle. Poly(amido amine)s (PAAs) and poly(urethane) are some polymers in this group. PAAs contain amido and tertiary amine groups that are regularly arranged along their polymer chain [55]. PAAs are being used in biomedical applications due to their biocompatibility, biodegradability, hydrophilicity, and low hemolytic activity [133]. PAAs can selectively adsorb heparin from plasma or blood and create a stable environment without any adverse effect on blood plasma or cells in vitro [134]. They can be created by stepwise addition of primary or secondary aliphatic amines to bis(acrylamide)s [135]. These polymers have a large number of surface sites compared to their molecular volumes. They are being applied as non-viral vectors and PAA/drug conjugates for intracytoplasmic deliveryLike polyethylenimine, bisacrylamide segment carrying carboxylic acid side group showed good endosomolytic properties and transfection efficiencies. The heart valve tissue engineering scaffold can be coated with PAA polymers and drugs can be attached to the polymer molecules.

Table 1 Stem/progenitor cells used in aortic valve tissue engineering.

t1:1 t1:2

O

F

560

R O

505

4.1.3. Cells Intact cells may be used as a means for efficient drug delivery [54, 142]. Cell-based drug delivery is an aspect of cell therapy which includes the administration of whole live cells as well as transfected or induced cells and drug encapsulated cells [143]. In broad terms, cell-therapy includes isolation of stem cells, administration of original stem cells and effector cells, transfection of stem cells, induction of stem cells to pluripotent cells and reprogramming of mature cells, encapsulation of drugs into the cells and administration of these modifies cells to sites of interest and their monitoring. Cell-based drug delivery can occur in two ways: transfected cells and cell carriers. Erythrocytes and macrophages are the commonest cell type currently used as carriers [144]. The drug can be released at the site or can be targeted to different cells while circulating. They are mainly used for therapeutic drug delivery. Different cell sources including stem and progenitor cells may be transfected with selected genes (e.g. complementary DNA) via viral or non-viral vectors [158]. These genes are selected for encoding secreted therapeutic products. Autologous stem or progenitor cells can be harvested from a variety of different sources (Table 1) and can be engineered genetically via the introduction of transgenes [144]. These transfected cells can then deliver the immunomodulatory molecules to the target site. They can be differentiated into cells with specific phenotypes and thus, can then be used to express and release different gene biomolecules at the tissue engineering site [159]. Dendritic cells – the antigen-presenting cells – can be genetically engineered through viral and non-viral means, mostly for therapeutic applications [160]. Transfected cells can be delivered to sites by attaching to the implanting scaffolds or injectable hydrogels, or through cardiac catheterization. Some disadvantages of this approach include lack of control over the duration of effect of the transfected cells and their distribution in the body and safety of the vector used for transfection. In heart valve tissue engineering research, although stem cells and progenitor cells have been applied, their application for drug delivery has not been fully explored.

P

503 504

550

D

501 502

the building blocks, the required mechanical properties, and biodegradability. To improve the biocompatibility of polyurethane, another natural polymer such as collagen is added during scaffold or drug delivery vehicle preparation [139]. Other synthetic polymers such as polyethers, polyanhydrides, and polyorthoesters are also used in drug delivery but in a very limited form [140]. For cardiovascular including aortic valve engineering, polyurethane is frequently used. For example, polyurethane fibrous scaffolds have been used to create artificial heart valve prostheses populated with human valvular interstitial cells and ECs which were then cultivated in a pulsatile bioreactor [141].

E

499 500

T

497 498

4.1.2. Synthetic polymers Synthetic polymers are produced in the laboratory and have controlled properties including molecular weight, mechanical properties, and degradability [125,126]. Most synthetic polymers are hydrophobic and have fewer cell-binding functional groups. Some synthetic polymers produce toxic elements when they degrade and thus are less optimal. The benefits of synthetic polymers lie in the facts that they are stable, available in large amounts and can be processed quite easily. In order to apply them as drug carriers, they must be compatible with the specific biological environments of interest.

C

495 496

E

494

R

492 493

R

490 491

erosion. Although some initial inflammation has been reported, they are biocompatible and biodegradable, leading to applications in cardiovascular tissue engineering and drug delivery [123]. PHA polymers can be fabricated into 3D porous scaffolds, nanofiber, nanoparticle, and microsphere forms, and drugs can be inserted into matrices at the time of fabrication or after fabrication. Copolymers such as polyethylene glycol (PEG) can be used for better sustained drug delivery [124].

N C O

489

U

488

7

553 554 555 556 557 558 559

561 562 563 564 565 566 567 568 569 570 571 572 573 574 575 576 577 578 579 580 581 582 583 584 585 586 587 588 589 590 591 592 593

t1:3

Animal sources Cell type

Organ source (tissue/animal)

References

t1:4

Mesenchymal stem cell Endothelial progenitor cell Bone marrow progenitor cell Autologous amniotic fluid cell

Bone marrow/sheep Peripheral blood/sheep Bone marrow/lamb Amniotic fluid/sheep

[145] [146] [147] [148]

t1:5 t1:6 t1:7 t1:8 t1:9 t1:10

Human sources

4.1.2.3. Hydroxyl-based. Hydroxyl group polymers contain OH groups, are very hydrophilic, and induce cell adhesion [55]. Thus, polymers containing OH groups such as polyvinyl alcohol are useful in tissue engineering and drug delivery applications. Hydrogels made of polyvinyl acetate (PVA) contracts or expands depending on pH of the culture media, allowing for control of drug release [136]. In addition to amine group discussed above, polyurethane has hydroxyl group which is useful for drug molecules binding [137]. It consists of a hard segment made of alternating diisocyanate and chain-extender molecules and a soft segment made of linear diols [138]. Thus, it is possible to tailor

551 552

Cell type

Organ source

Reference

t1:11

Mesenchymal stem cell

Bone marrow Adipose tissue Umbilical cord matrix Umbilical cord blood Amniotic fluid Chorionic villi Amniotic fluid Peripheral blood Umbilical cord blood Skin (fibroblast)

[149] [150] [151] [152] [153] [154] [155] [156] [151] [157]

t1:12 t1:13 t1:14 t1:15 t1:16 t1:17 t1:18 t1:19 t1:20 t1:21

Endothelial progenitor cell

Induced pluripotent stem cells



Drug carriers have different forms: 3D solid porous scaffolds, hydrogels, nanofibers/nanoparticles, microsphere/microparticles, membranes and cells (Fig. 4) [161]. All drug carriers should possess specific qualities including: drug binding affinity, homogeneous drug dispersion capability, drug retention capacity, drug release capability at a predetermined rate, and stability for a required period of time [81]. Importantly, the drugs should not be damaged from time of attachment to the carrier until it has been delivered. Drug release occurs by diffusion of biomolecules through pores, degradation of carriers, or swelling of materials in carriers [162]. The release of drug from carrier materials depends on physiochemical properties of polymer and drug, and physiological properties of release area.

615 616 617

D

613 614

E

612

T

610 611

C

608 609

4.2.1. Substrate-based This approach encapsulates or attaches protein drugs in or onto the biomaterial carrier in the surrounding medium or scaffold matrix and controls release of the drug at the targeted site [163]. During carrier fabrication, the drug is directly incorporated by mixing it with carrier materials. Fabrication methods that involve high temperature, organic solvents, and that generate shear stress or free radicals should be compatible to drug integrity. Moreover, for controlled release with appropriate drug concentration, carrier degradation rate and drug diffusion rate should comply with rate of tissue growth [164]. Drug release efficiency also depends on local microenvironment, chemistry, and interaction of

628

E

607

4.2.1.1. 3D solid porous scaffolds. Implantable 3D solid porous scaffolds are one of the most frequently used drug delivery systems in engineering tissues, including heart valves [81]. Because they provide support and guidance to cells their mechanical properties, pore size, porosity, and degradation rate should be conducive to the specificity of the tissue. The scaffolds can be prepared using several major methods such as porogen leaching, phase separation, gas foaming, and solid freeform fabrication [167]. Natural polymers such as collagen, gelatin, chitosan, and PHAs or synthetic polymers such as PLA, PGA, PLGA, and PGS have been used to prepare 3D scaffolds. Laser microfabricated PGS scaffold with biomimetic structure was found mechanically compatible to heart valve leaflet [168]. Drugs are included into the raw materials before fabricating the scaffold or drugs can be attached on the scaffold surface through surface functionalization, graft polymerization and drug immobilization.

R

606

R

604 605

O

602 603

C

600 601

N

599

U

597 598

618 619

F

595 596

drugs. In addition to encapsulation in the carrier, drugs can be attached onto carrier polymers [54]. In this process, drugs (growth factors) are immobilized onto the surface of carrier by their adsorption or chemical crosslinking. In the adsorption process, physiochemical interaction includes as ion complex formation [165]. The amount of drug binding depends on number of binding sites and affinity of drug to the binding sites. The release of drugs can occur upon either degradation of linking tether or carrier itself. In some cases, non-biodegradable materials such as ethylene–vinyl acetate are used for continuous, diffusion-controlled drug release [166].

O

4.2. Design of carriers

R O

594

P

8

Fig. 4. Different forms of drug carriers. They include solid 3D porous scaffold, 3D hydrogel, nanofibrous scaffold, nanoparticles, microspheres and cells. Drugs are encapsulated into or adsorbed onto the carriers.


620 621 622 623 624 625 626 627

629 630 631 632 633 634 635 636 637 638 639 640 641 642


663 664 665 666 667 668 669 670 671 672 673 674 675 676 677 678 679 680 681 682 683 684 685 686 687 688 689 690 691 692 693 694 695 696 697 698 699 700 701 702 703 704 705

4.2.1.3. Microspheres. Biodegradable polymeric microspheres are used for drug delivery due to advantages such as easy encapsulation of drugs into the microsphere and easy administration through a syringe needle [81,175]. Microspheres are fabricated by several methods including emulsion/solvent evaporation, phase separation, or precipitation and spraying method [176,177]. The drug release from microspheres depends on fabrication method in addition to type and molecular weight of polymer, copolymer ratio, size of microsphere, and drug stabilization agent. Drugs can be incorporated into microspheres or onto surface of the microspheres. Despite their success in drug delivery, they have some key drawbacks including low scale of production, potential to damage the drug during fabrication, and poor release rates [178]. Both synthetic and natural polymers are used to fabricate microspheres [179]. Direct injection of drug-loaded microspheres made from poly(cyclohexane-1,4-diylacetone dimethylene ketal) polymer showed improvement in cardiac function following myocardial infraction [180].

4.2.1.4. Nanofibers. Nanofibers have received much attention in drug delivery due to their extremely high surface to weight ratio [181,182]. Moreover, they provide ECM structure that is compatible to tissue engineering. Nanofibers can be produced through several processing techniques such as drawing, template synthesis, phase separation, electrospinning, and self-assembly [183]. A wide variety of polymers, both natural and synthetic, are used to prepare nanofibers intended for drug delivery. Drugs are mixed into polymer solution before making nanofiber or drugs can be attached to the surface of nanofibers through surface functionalization or graft polymerization and drug immobilization. These attached drugs are released in three ways: immediate, extended, and triggered release [184]. Nanofibrous membrane patches are an efficient form for drug delivery because the drugs are attached to the surface of the patch and then directly applied to the damaged site [185]. Till now, most of the heart valve scaffolds applied for valve tissue engineering are fabricated from nanofibers [41]. 4.2.1.5. Nanoparticles. Nanoparticles can be envisioned as a future drug delivery system due to potential as therapeutic and diagnostic agents [98,186]. Nanoparticles demonstrate rapid onset of therapeutic action as drug particles, are resistant to settling, obtain higher solubility, rapid dissolution, and enhanced adhesion to biological surfaces [187]. Moreover, due to a higher surface area, biomacromolecules can reside on the surface and easily reach the target, leading to efficient delivery of drugs, proteins, and polynucleotides to cells and tissues [188]. Nanoparticles can be synthesized through processes such as self-assembly, electrostatic or vapor deposition, aggregation, nano-manipulation, and imprinting [188]. Viral and non-viral vectors are considered to be nanoparticles that are used to transfect cells to express specific phenotype

5. Drug delivery for aortic valve generation

735

The main challenge of drug delivery in aortic valve tissue engineering has been lack of an ideal scaffold or platform that can mimic the aortic valve structure in shape, mechanical properties, and capability to withstand harsh physiological conditions [148,149,151]. Only decellularized biologic scaffolds and nanofibrous valve shaped scaffolds have been applied for heart valve tissue engineering as they have the capacity to mimic the heart valve structure and properties [192]. Thus, all the delivery systems or vehicles that have been developed and applied thus far in heart valve tissue engineering take an indirect approach or just characterize the efficacy of the drugs in vitro on stimulating different cell types to improve collagen formation on a substrate. It is difficult to mimic the native molecular regulation in valve genesis; however, efforts are being made to deliver drugs (Table 2) in multiple ways to optimize valve engineering [193]. There are two methods that have been used for drug delivery in aortic valve tissue engineering: substrate-mediated drug delivery and cell-mediated drug delivery. Among them, substrate-based drug delivery is more commonly investigated as the inclusion of growth factors can be done with relative ease to a scaffold matrix or culture media. The other method, cell-mediated delivery, has been tested less extensively. Transfected cells that are administered in vivo may not survive, may be damaged, may be washed away, may be immunogenic or may lose bio-functionality before reaching to the target. Additionally, gene vectors can directly injected into the damaged sites of the valve for tissue engineering, although this method is used more in engineering of other tissues such as bone, cartilage and skeletal muscle [194].

736

5.1. Delivery of growth factors

762

In heart valve engineering, cell adhesion is vital for tissue regeneration. Extracellular receptors such as proteoglycans play important roles in cell adhesion [210]. These receptors can bind to other proteins such as collagen, fibrin, and fibronectin, which can influence cell adhesion.

763

O

F

713 714

R O

661 662

4.2.2. Cell-based Proper selection of cells and growth factors/genes is important in designing cell carriers for drug delivery to the targeted site. The more difficult task is the allowing for the efficient introduction of selected genes to the cells while avoiding adverse issues such as toxicity, immunogenicity and over expression. The genes are introduced to cells through viral and non-viral methods [142,190]. A variety of viruses including adenovirus, retrovirus, and lentivirus have been used for viral delivery. Viral transfection is an efficient method but there have been multiple adverse issues which create problems for in vivo tissue engineering. In non-viral transfection method, non-living objects such as nanoparticles (liposomes, micelles, polymer nanoparticles, iron oxide nanoparticles, gold nanoparticles etc.) and microspheres made of organic or inorganic materials are used as gene carriers [191]. These gene-attached viral/non-viral carriers are internalized by cells which are then transfected by the attached genes, respectively. In non-viral methods, although immunogenicity and toxicity can be avoided by using biocompatible materials, transfection efficiency is very low. Ongoing efforts are being made to improve non-viral systems for efficient cell transfection and improved drug delivery. Efficiency in gene expression of transfected cells depends on transfection efficiency of the cells, their maturation and the microenvironment surrounding them.

P

660

706

D

658 659

and release protein drugs at the target site. The efficiency of a particular nanoparticle depends on its size, surface charge, encapsulation efficacy, and release capacity. Drugs are attached to the nanoparticle surface through various methods such as adsorption and covalent bonding. Due to biocompatibility and biostability, composite with polyhedral oligomeric silsesquioxane nanoparticles in polycarbonate soft segment has been evaluated as synthetic material for heart valve leaflets [189].

E

656 657

T

654 655

C

652 653

E

650 651

R

649

R

647 648

N C O

645 646

4.2.1.2. Hydrogels. Another form of matrix for drug delivery is hydrogels that can be composed of chemically or physically cross-linked, watersoluble polymers [169]. In most cases, at fabrication step, drugs are mixed with raw ingredients and then proceed for hydrogel preparation. This way, a homogeneous concentration of drug throughout the hydrogel can be achieved. Due to their biocompatibility and high water content, they are preferred for biomedical applications. Hydrogels can be prepared from natural polymers such as collagen, fibrin, chitosan, and gelatin or from synthetic polymers such as PVA, poly(ethylene glycol) diacrylate (PEGDA) and PEG [80,170,171]. In addition, other parameters such as ionic concentration, temperature of the surroundings, and type and concentration of cross-linker control the hydrogel permeability and swelling and thus regulate drug release. Bioprinted aortic heart valve structured hydrogel constructs from alginate and PEGDA polymers showed cells survival and respective gene expression in static culture [80,172]. Thin hydrogel layer membranes containing drugs are applied in some applications such as cardiac patch and intravascular tissue engineering [173,174].

U

643 644

9


707 708 709 710 711 712

715 716 717 718 719 720 721 722 723 724 725 726 727 728 729 730 731 732 733 734

737 738 739 740 741 742 743 744 745 746 747 748 749 750 751 752 753 754 755 756 757 758 759 760 761

764 765 766

10 t2:1 t2:2


Table 2 Bioactive agents applied in aortic valve tissue engineering.

t2:3

Agents

Major functionalities

Cells/agents carriers

References

t2:4

Bone morphogenetic proteins (BMPs) Fibroblast growth factors (FGFs)

Culture media, collagen type I gel, transfected porcine valvular interstitial cells, fibrin matrix gel and scaffold, gelatin scaffold and microsphere Culture media, transfected mesenchymal stem cell, alginate gel and scaffold, chitosan nanofibers, collagen gel

[195–197]

t2:5

t2:6

Insulin growth factor-I (IGF-I)

t2:7

Platelet-derived growth factor (PDGF)

t2:8 t2:9

Transforming growth factor-β1 (TGF-β1)

t2:10 t2:11

Vascular endothelial growth factor (VEGF)

Activation of osteoblast-associated transcription factors, augmentation of osteogenic proteins expressions, upregulation of Wnt3a, Runx2, Msx2 and β-catenin Increment in cell viability and proliferation, stimulation to the residing cells for pro-fibroblast phenotype expression, reduction in TGF-β influence, cell growth in combination with other growth factors, increase in myofibroblastic differentiation Increment in growth, proliferation and differentiation of different cells, stimulation to other growth factors production, reduction of apoptotic cell death, boosting in calcification Upsurge of SMC proliferation. Reduction in the smooth muscle actin (SMA) expression, higher growth of cell in combination with other growth factors, inhibition of negative effect of mechanical stimulation on aortic valve tissue engineering. Activation of myofibroblast phenotype in VICs, formation of apoptosis-driven dystrophic calcific nodules, higher collagen/ecm deposition Inhibition in calcification, improvement in proliferation, growth and differentiate of stem cells to ECs, induction of vascular environment

782 783 784 785 786 787 788 789 790 791 792 793 794 795 796 797 798 799 800 801 802 803 804 805 806 807 808 809 810

F

O

R O

P D E

T

C

780 781

E

778 779

[207–209]

R

776 777

Collagen, fibronectin and laminin coated surface, poly-glycolic acid/poly-4-hydroxybutyrate scaffold, fibronectin gel, gelatin scaffold and microsphere

R

774 775

[204–206]

O

772 773

Culture media, collagen type I gel, poly-glycolic acid/ poly-4-hydroxybutyrate scaffold

C

770 771

Integrins, another family of proteins that work as cell surface receptors, also mediates the binding of cells through different cytokines and growth factors. For improved cell adhesion to scaffolds or substrate materials, collagen, fibrin, and other proteins have been incorporated onto the scaffolds or substrates by coating, adsorption, or binding. Likely, secondary to better cell adhesion, higher growth, differentiation, ECM production, and maturity occurred on those substrates [211]. Some other adhesive ligands such as Arg-Gly-Asp (RGD) peptides, Arg-Gly-Asp-Ser (RGDS), and Arg-Glu-Asp-Val (REDV) also increased the cell adhesion. However, these peptides inhibited ECM production on scaffolds, which limits their application in tissue engineering [212]. As substitutes, seeded cells might be modified to secrete appropriate growth factors. In the aortic valve, VICs with the characteristics of both fibroblast and myoblast (termed myofibroblast) are mainly responsible for the collagen deposition in heart valve tissue formation. TGF-β1—a leading growth factor in heart valve generation, induces differentiation of different cell lineages into myofibroblast leading to a collagen deposition [213,214]. Lijnen et al. and other researchers have applied TGF-β1 on fibroblasts to differentiate into myofibroblasts and induce collagen deposition confirmed by analysis of alpha-smooth muscle actin (α-SMA) [215,216]. The rate of collagen production and fibroblast differentiation were time dependent as well as TGF-β1 concentration dependent. As cell growth is inhibited by dose-dependency of TGF-β1, TGF-β1 in combination of other growth factors such as insulin-like growth factor (IGF), EGF, bFGF, and platelet-derived growth factor (PDGF) induces similar differentiation but overcomes TGF-β1-induced growth inhibition while simulating fibroblast proliferation and invasion (Fig. 5) [198]. Another growth factor cocktail consisting of TGF-β1, EGF, bFGF, and PDGF induced proliferation of vascular smooth muscle cells (VSMCs) and expression of ECM matrix components in aortic valve [217]. However, some researchers have reported that bFGF blocks TGF-β1-mediated myofibroblast activation [218]. On the other hand, TGF-β1 inhibited osteogenic differentiation of human mesenchymal stem cells [219], while TGF-β1 in the presence of Wnt3A caused higher stiffness in fibrosa through greater myofibroblast differentiation than TGF-β1 alone [220]. Beside fibroblasts, other cells such as VICs, embryonic valvular cells, human mesenchymal cells (hMSCs), VSMCs, and ECs also exhibited higher collagen deposition in the presence of TGF-β1 [217,219,221]. However, all of the cell lineages did not show similar performance in heart valve tissue engineering. For example, compared to VICs, mesenchymal stem cells (MSCs) represented a promising cell type for valve tissue engineering [222]. Addition of heparin improved the activity of

[202,203]

N

768 769

[200,201]

U

767

Transfected mesenchymal stem cell, culture media, polyglycolic acid/poly-4-hydroxybutyrate composite scaffold Culture media, collagen type I gel, chitosan nanofibers

[198,199]

Fig. 5. Proliferation and differentiation of human dermal mesenchymal cells (hDMC1.1) in the presence of different single growth factors or their combination. (a) The graph shows the relative proliferation and standard deviation in terms of optical density. * indicates the statistically significant results (p b 0.05). The standard deviation was calculated relative to either control or TGF-β1-treated sample. (b) Expression of α-SMA in hDMC1.1. The growth factors in these samples were similar to shown in (a) but clockwise order. (c) Expression of tenascin-C (TNC) in hDMC1.1. The growth factors in these samples are similar to that shown in (a) but in clockwise order. Scale bars indicate 50 μm [198]. Reprinting permission will be obtained upon acceptance.



832 833 834 835 836 837 838 839 840 841 842 843 844 845 846 847 848 849 850 851 852 853 854 855 856 857 858 859 860 861 862 863 864 865 866 867 868 869 870 871 872 873 874 875 876

Gene-based delivery provides an alternative approach to induced protein expression within tissue engineered site [231,232]. In this method, the genes DNA or RNA are packed with proteins, lipids, or polymers to create nanovectors-based drug particles that can then be encapsulated in biomaterials with scaffold, hydrogel, or other forms [233]. When scaffold substrates are implanted at the tissue engineering site, genes can be directly released from the scaffold carriers to transfect cells [234,235]. In targeting a cell population or anatomical location, direct gene delivery from scaffold has some advantages over a traditional delivery system such as intravenous administration. Direct delivery of gene/ drug from scaffold is localized, and scaffolds can protect vectors against the ECM barrier that protects the drug from immune response and limits the degradation from serum nuclease or protease [233]. Moreover, release of vectors depends on the degradation of scaffold materials; thus, an effective level of vector release for an extended time leads to efficient drug internalization in cells and gene transfer. Another advantage is that the scaffold prevents clearance of drug due to prolonged release. Like other drugs, gene vectors are incorporated in scaffold systems in two ways: encapsulation and surface immobilization [233]. In scaffold preparation, the gene vectors are included in the materials and applied for fabrication. For hydrophilic polymer scaffolds, vector release occurs through scaffold degradation and vector dissolution and diffusion [233]. Vector release from hydrogels relies upon hydrogel structure, its interaction with hydrogel, and hydrogel degradation strategies. Diffusion of the vector is the only option for non-degradable hydrogel. Vector immobilization follows the mechanisms of natural virus binding to extracellular matrix proteins. However, relative affinity of vectors to the scaffold surface should be modulated to achieve local delivery at the targeted site by considering the influence of its physiological environment [233]. In general, non-specific binding between viral or non-viral vectors containing negatively charged DNA or RNA and polymeric biomaterials uses the immobilization technique; specific binding can be ensured through complementary functional groups in each of the vectors and polymeric biomaterials. Hydrophilic and ionic materials show better immobilization characteristics compared to their counterparts. The strength of immobilization depends not only on the affinity of the vectors toward the biomaterials, but also on the environmental condition such as pH and ionic strength. Although, growth factors have been delivered through scaffolds or other carriers, only a few have been described in gene delivery for aortic valve tissue engineering. Thus, in addition to review these, gene delivery for related tissue engineering such as myocardium, vessel that can potentially be applied to heart valve tissue engineering is discussed here. A study applied thiol-based chemistry by attaching polyurethane films coated with Type I collagen to anti-adenovirus antibodies to enable gene delivery via vector tethering [236]. They also activated the polyurethane film surface by the alkyl-thiol-conjugation method to directly attach anti-adenovirus (AdGFP) antibodies to the polyurethane surfaces (Fig. 6). These two gene delivery systems were implanted in pig using a button configuration that simulated the blood-contacting environment, and in sheep valve leaflet replacement studies, respectively. In

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with endothelial progenitor cells and aortic valvular ECs for their transformation to mesenchymal cells in the presence of VEGF and TGF-β1 [208]. The growth factor-enriched environment showed higher cellular proliferation, differentiation, and tissue development. Like hydrogel, 3D solid microspheres carrying different drugs were also applied noninvasively to the heart valve damage site [219]. With respect to other scaffolds, those in nanofibrous form showed better results due to nanoscale environment that mimics the ECM structure; however, in most cases, nanofibers were limited to in vitro studies only [230].

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growth factors on VICs and thus greatly induced their phenotypic expression of α-SMA [223]. Notably, heparin with shorter chain lengths showed higher activity. It was also observed that TGF-β1 induced calcination in VICs in the presence of ECM proteins such as collagen, fibronectin, or laminin [209]. In contrast, VEGF significantly inhibited the formation of calcific nodules independent of ECM protein presence. The outer layers of valves and cusps are covered by ECs, and with aging and the accumulation of medical comorbidities (diabetes, hypertension, hyperlipidemia etc.) they may be damaged or their functionalities affected [224]. VEGF was useful to differentiate MSCs into endothelial phenotype cells or to improve the activity of the existing ECs. In addition to delivering the appropriate drug, cyclic substrate deformation (equibiaxial mechanical strain) or other mechanical stimuli and electrical stimuli showed upregulation of Wnt/beta-catenin in mesenchymal progenitor cells [225]. Combination of mechanical strain or stimuli and growth factors such as TGF-β1 induced dual-mode endothelial– mesenchymal transformation in cardiac valves [226]. Although there may be impact of different growth factors on heart valve tissue engineering and its different cell types, the main challenge in developing delivery system for heart valve regeneration comes from the complexity of heart valve structure which hinders the fabrication of scaffolds with appropriate structure and mechanical properties. Moreover, the native aortic valve consists of highly elastomeric ECM materials. Therefore, researchers have tried to deliver drugs in the decellularized biologic scaffolds for heart valve regeneration [227]. Hong et al. coated the decellularized scaffold with bFGF/poly-4hydroxyburate by using electrospinning technique and reseeded and cultured MSCs for 14 days in vitro. As a result, improved cell mass, collagen deposition, and strength were observed in this hybrid scaffold compared to the decellularized scaffold control. PEG nanoparticles loaded with TGF-β1 were combined to decellularized scaffold for better biocompatibility and biomechanical properties [186]. It was also noted that fibronectin–hepatocyte growth factor boosted in situ recellularization in decellularized porcine aortic valves [78]. Heparin–VEGF multilayer film on decellularized aortic valve created not only an antithrombotic environment but also improved the adhesion of endothelial progenitor cells [228]. As mentioned earlier, cells can be encapsulated in the growth factormixed hydrogels and then be applied to the damaged site for regeneration. The advantage of thermoreversible hydrogels is that they can be applied noninvasively as it changes from solution to gel at high temperature (body temperature). In aortic valve engineering, natural polymers such as collagen, fibrin, fibronectin, gelatin, hyaluronic acid, alginate, and chitosan are extensively used in hydrogel forms [80,112]. They have thermoreversible capability; or with a cross-linker, their solutions can turn into gel. Those natural protein-based hydrogels generally contain different cytokines; thus, in some cases, they themselves provide growth factors. Synthetic polymer-based hydrogels have higher stability, and with incorporation of growth factors, is also efficient in tissue engineering. For example, TGF-β1 in PEG hydrogel improved the collagen deposition and α-SMA expression from the encapsulated VICs [221]. Photocrosslinked microspheric gelatin hydrogel in the presence of TGF-β1 and FGF showed higher proliferation and differentiation of VICs [229]. In addition to growth factors, the low mechanical properties of hydrogel reduced the risk of calcification in valves. Calcification studies through collagen gel showed similar impact. Three growth factors TGF-β1, BMP2, and VEGF in Matrigel showed different characteristics in endothelial to mesenchymal cell transformations [193]. Experience with drug delivery in aortic valves through a 3D-solid porous scaffold is very limited due to the noted difficulties in making valvular structured scaffold. In most cases, porous scaffolds of arbitrary shapes were applied to observe the efficacy of drug delivery from the scaffold or effect of drug on cells in the scaffolds. Drugs were incorporated onto the surface of solid scaffolds through heparin-mediated attachment or into the scaffold materials during fabrication. In in vitro experiments, 3D solid porous scaffold made of PGA/P4HB was seeded

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both cases, neointimal cells were observed on the implant surface after one week. Song et al. also used collagen-coated polyurethane films to assess their feasibility as a gene delivery platform. They attached antibody-tethered replication-defective adenoviruses encoding for green fluorescent to polyurethane film. After one week of implantation, the implant in pig demonstrated neointimal cells on the surface of the implant [237]. Acharya et al. demonstrated controlled and sustained release of nitric oxide (NO) for prevention of heart valve complications through gene delivery. They prepared NO donors DETA NONOate containing PLGA nanoparticles, which were incorporated in PLLA scaffold prepared by the salt leaching method. Its application in heart valve showed reduction of rate of calcification in VIC through NO-cGMP signaling pathway. Moreover, NO released from PLGA nanoparticles embedded in PLLA scaffold prohibited inflammation without causing any cytotoxic

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Fig. 6. Gene delivery using polyurethane (PU) replacement cusps with antibody-tethered AdGFP. (a) Photography of polyurethane (PU) replacement cusps with antibody-tethered AdGFP demonstrated a smooth blood contacting surface (shown by arrow). Scale bar indicates 2 cm. (b) Photography of non-modified polyurethane (PU) replacement cusps also demonstrated a smooth blood contacting surface (shown by arrow). Scale bar indicates 2 cm. (c) Presence of cells on explanted AdGFP–polyurethane cusp replacement was confirmed by immunostained (FITC and DAPI) image of GFP expressed cells (shown by arrow, X 400). (d) Absence of fluorescence (FITC and DAPI) confirmed the lack of cells on explanted non-modified polyurethane cusp replacement (X 400). (e) Cells on the myocardium near explanted AdGFP–polyurethane cusp replacement were confirmed by immunostained (FITC and DAPI) image of GFP expressed cells (shown by arrow, X 400). (f) Absence of fluorescence (FITC and DAPI) confirmed the lack of cells near explanted non-modified polyurethane cusp replacement (X 400) [236]. Reprinting permission will be obtained upon acceptance.

effects [238]. In another example of gene delivery, non-viral cationic poly-ethyleneimine-plasmid DNA (PEI-CD-DNA) condensates were used as gene delivery vehicles [239]. Porcine decellularized heart valve scaffolds were coated with PEI-CD-DNA embedded fibrin. VGEF and stromal derived factor-1 were used as functional genes. The results demonstrated high transfection efficiencies in porcine cells and blood vessel cells seeded in decellularized valve scaffold. Although polyurethane is a useful biomaterial for prosthesis implant development and scaffold fabrication in tissue engineering, thrombi formation and ectopic calcification on heart valve scaffolds make this material unreliable for heart valve tissue engineering [240]. Appending therapeutic moieties directly to polyurethane reduces the reactivity of its chemical groups. Fishbein et al. modified polyurethane by bromoalkylation and used this modified material as a gene delivery platform in heart tissue engineering [241]. They attached the gene to


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Table 3 Cell-based gene therapy applied to cardiovascular tissue engineering. Application area

Transfection agent/transfected cells

Therapeutic agent

Host

Benefit

Reference

t3:4

Myocardium

IGF-I

Rat

Myocardium

FGF-2

Rat

t3:6

Myocardium

Fused FGF4-bFGF

Rat

HGF-cDNA

Rat

VEGF

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Higher cell survival rate and resistance to apoptosis. Upregulation of VEGF, bFGF and HGF growth factors Increased levels of FGF-2 and displayed a threefold increase in viability, as well as increased expression of the anti-apoptotic gene, Bcl2, and reduced DNA laddering under hypoxic condition. Improved survival of the transplanted cells, diminish myocardial fibrosis, improved cardiac functions, and BFGF secretion HGF-modified ADSCs could migrate to ischemic myocardium, increase vascular density, and reduce fibrotic area Promoted cardiac stem cell (CSC) migration to infarcted area

[200]

t3:5

Adenoviral vector/bone marrowderived MSCs pEGFP-plasmid/adipose-derived MSCs

Heme oxygenase-1

Mouse

FnCBD64

In vitro

Enhances the tolerance of engrafted MSCs to hypoxia–reoxygen injury in vitro and improves their viability in ischemic hearts Increase in adhesion capacity

TGF-β1

In vitro

Higher proliferation and ECM production

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[251] [252] [253] [254]

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In most cases of cell-mediated drug delivery, drugs were internalized into the cells; only few drugs were attached to cell-membrane for their dispatch to target sites. Neuregulin-I (NRG-I) was encapsulated in PLGA polymer particle and was attached to adipose derived stem cells using collagen coating on the particle [258]. After 14 days of cellattached NRG-I drug administration, some particles were found partly degraded and cells were remained attached to them. However, drug release from the particles was sufficient and cell proliferation was augmented.

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the bromoalkylated polyurethane via anti-adenovirus antibodies. After implantation of the engineered scaffold in an ovine model, they observed local gene delivery. To address the thrombi formation issue, cholesterol moieties were attached to enhance EC retention in physiological condition [242]. Cholesterol has been recognized as a necessary component of cell-surface-mediated signaling pathways and adhesion. Experiments using the modified polyurethane with bovine aortic ECs showed improved cell retention properties. Attachment of cells to the scaffold systems is an importance issue in tissue engineering and could be enhanced through synthetic oligopeptides containing the adhesion site of fibronectin (the RGD sequence). The scaffolds could carry both the cells and their stimulating factors, including growth factors, prompting the cells to produce the specific growth factors at the implanted site. In some cases, cells such as MSCs are used as they are easy to manipulate and have potential to produce specific growth factors [243,244]. In aortic valve tissue engineering, there has been reports of cell-based therapeutic intent in different aspects of valve tissue engineering such as cell adhesion, proliferation, and differentiation with different cell lineages and genes [41]. A great number of cell-based gene therapy applied to cardiovascular tissue engineering (Table 3) can be useful for drug delivery in heart valve tissue engineering. Collagen deposition at differentiation stages of tissue engineering not only establishes the structure but also improves the mechanical properties of the valve. Rocnik et al. transfected the SMCs with a retrovirus containing heat shock protein 47 (HSP47)—a collagen-specific molecular chaperone applied for cardiovascular tissue engineering [245]. The results showed increased intra- and extracellular steadystate type 1 collagen production (Fig. 7). To improve the stiffness of the valve leaflet, Mironov et al. transfected the human fat tissue-derived stem cells with periostin and applied them on electrospun collagen nanofibers [246]. They demonstrated accelerated tissue formation and collagen deposition. In addition to collagen deposition, antithrombogenic responses and other developments are part of aortic valve genesis, and VECs are involved through endothelial nitric oxide synthase (eNOS) catalytic conversion of L-arginine to L-citrulline, termed nitric oxide signaling [247]. To assess nitric oxide release in VICs, they were transfected with a viral vector encoding eNOS (adeNOS) and release of nitric oxide was measured to find the level of Wnt3a secretion related to stenosis disease [248]. ECs play a vital role in the entire cardiovascular system including the aortic valve to make the system antithrombogenic [255]. In heart valve tissue engineering, generating of layer of ECs on the engineered valve is necessary and can be promoted by local delivery of growth factor genes or by growth factors themselves. ECs transfected with VEGF stimulated endothelialization and anti-thrombogenicity [256,257].

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Lentiviruses/bone marrowderived MSCs Ischemic heart Lentiviral vector/adipose-derived MSCs Left ventricle Adenoviral vector/bone marrowderived MSC Ischemic heart with a hypoxia- Hypoxia-inducible Tf PEI-plasmid/ regulated environment bone marrow-derived MSC Decellularized scaffold Adenovirus-coding Fn signal peptide/EPCs Decellularized scaffold Gene-plasmid pcDNA3.0/ Myofibroblasts

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6. Conclusion

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In the development of the native aortic valve, numerous growth factors are involved at different stages of growth and maturity. For successful tissue engineering, similar types of growth factors should be delivered in a controlled manner to the regeneration site. Drug carriers made of a wide variety of biocompatible and bioresorbable materials including natural polymers and synthetic polymers have been used for this purpose. To deliver the drug effectively, carriers have been designed in different forms such as hydrogels, 3D solid porous scaffolds, nanofibers, nanoparticles, and microspheres through manipulation of physiochemical properties. Moreover, drugs have been delivered in various ways: directly through carrier or through transfection of cells that can release drugs at the site or genes have been expressed or directly delivered from the carrier or vector to manipulate the local environment for efficient cell functioning. Although encouraging results have been obtained in many heart valve engineering experiments, efforts should move toward the development of a multifunctional platform to achieve native-like aortic heart valve tissue construct.

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7. Future perspective

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A better understanding of the molecular and cellular mechanisms specific to aortic valve development and regeneration must be achieved through further study. Special emphasis should be devoted to producing 3D scaffolds mimicking the heart valve structure to achieve clinical level valve constructs. Then, physiochemical properties of the scaffold materials can be tuned for better drug incorporation, protection, and controlled release. Additional exploration on the delivery of drugs through nanoparticle, microsphere, and hydrogel carriers can improve their efficacy and use in drug delivery. Both substrate-based therapeutic intent and cell-based therapeutic intent are promising methods for drug delivery; however, more widespread investigations are needed. In cell-based drug delivery, issues like poor cell adhesion to scaffold system, poor incorporation of systematically injected cells and loss or death of majority of adhered or injected cells are often observed.

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Native cardiac valves are developed in a multi-bioactive reagents environment and thus, to generate functional heart valve tissue construct in vitro, tissue engineering should be performed in similar multi-drug environment. However, most of drug delivery experiments deal with only one drug at a time to investigate specific issues. Potentiality of different natural and synthetic polymers can be used to develop multi-drug carriers in the forms of 3D scaffold, nanofibers, nanoparticles and hydrogel etc. With optimized multi-drug delivery system, cardiac valve can be developed in an in vitro dynamic environment and can then be tested in an animal model for real valve construct generation.

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Disclosures

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This work is supported by the HH Sheikh Hamed bin Zayed Al Nahyan 1078 Q3 Program in Biological Valve Engineering, the Grainger Foundation, and 1079 the Mayo Clinic Center for Regenerative Medicine. 1080

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[1] T. Treasure, Rethink on biological aortic valves for the elderly, Lancet 354 (1999) 964–965. [2] I. El-Hamamsy, Z. Eryigit, L.M. Stevens, Z. Sarang, R. George, L. Clark, G. Melina, J.J. Takkenberg, M.H. Yacoub, Long-term outcomes after autograft versus homograft aortic root replacement in adults with aortic valve disease: a randomised controlled trial, Lancet 376 (2010) 524–531. [3] J.L. d'Arcy, B.D. Prendergast, J.B. Chambers, S.G. Ray, B. Bridgewater, Valvular heart disease: the next cardiac epidemic, Heart 97 (2011) 91–93. [4] D. Messika-Zeitoun, L.F. Bielak, P.A. Peyser, P.F. Sheedy, S.T. Turner, V.T. Nkomo, J.F. Breen, J. Maalouf, C. Scott, A.J. Tajik, M. Enriquez-Sarano, Aortic valve calcification: determinants and progression in the population, Arterioscler. Thromb. Vasc. Biol. 27 (2007) 642–648. [5] A. Mol, A.I.P.M. Smits, C.V.C. Bouten, F.P.T. Baaijens, Tissue engineering of heart valves: advances and current challenges, Expert Rev. Med. Devices 6 (2009) 259–275. [6] B. Remenyi, N. Wilson, A. Steer, B. Ferreira, J. Kado, K. Kumar, J. Lawrenson, G. Maguire, E. Marijon, M. Mirabel, A.O. Mocumbi, C. Mota, J. Paar, A. Saxena, J. Scheel, J. Stirling, S. Viali, V.I. Balekundri, G. Wheaton, L. Zuhlke, J. Carapetis, World Heart Federation criteria for echocardiographic diagnosis of rheumatic heart disease—an evidence-based guideline, Nat. Rev. Cardiol. 9 (2012) 297–309. [7] J.T. Butcher, G.J. Mahler, L.A. Hockaday, Aortic valve disease and treatment: the need for naturally engineered solutions, Adv. Drug Deliv. Rev. 63 (2011) 242–268. [8] N.A. Boon, P. Bloomfield, The medical management of valvar heart disease, Heart 87 (2002) 395–400. [9] M.H. Yacoub, J.J. Takkenberg, Will heart valve tissue engineering change the world? Nat. Clin. Pract. Cardiovasc. Med. 2 (2005) 60–61. [10] P. Kvidal, R. Bergstrom, L.G. Horte, E. Stahle, Observed and relative survival after aortic valve replacement, J. Am. Coll. Cardiol. 35 (2000) 747–756. [11] S. Basude, C. Hein, S.L. Curtis, A. Clark, J. Trinder, Low-molecular-weight heparin or warfarin for anticoagulation in pregnant women with mechanical heart valves: what are the risks? A retrospective observational study, Bjog-an Int. J. Obstet. Gynaecol. 119 (2012) 1008–1013. [12] K. Hammermeister, G.K. Sethi, W.G. Henderson, F.L. Grover, C. Oprian, S.H. Rahimtoola, Outcomes 15 years after valve replacement with a mechanical versus a bioprosthetic valve: final report of the Veterans Affairs randomized trial, J. Am. Coll. Cardiol. 36 (2000) 1152–1158. [13] P.M. Dohmen, W. Konertz, Tissue-engineered heart valve scaffolds, Ann. Thorac. Cardiovasc. Surg. 15 (2009) 362–367. [14] M.N. Tillquist, T.M. Maddox, Cardiac crossroads: deciding between mechanical or bioprosthetic heart valve replacement, Patient Prefer. Adherence 5 (2011) 91–99. [15] C.V.C. Bouten, P.Y.W. Dankers, A. Driessen-Mol, S. Pedron, A.M.A. Brizard, F.P.T. Baaijens, Substrates for cardiovascular tissue engineering, Adv. Drug Deliv. Rev. 63 (2011) 221–241. [16] E.J. Armstrong, J. Bischoff, Heart valve development—endothelial cell signaling and differentiation, Circ. Res. 95 (2004) 459–470. [17] Y. Tabata, The importance of drug delivery systems in tissue engineering, Pharm. Sci. Technol. Today 3 (2000) 80–89. [18] H. Kim, R. Bergman, R. Matyal, K.R. Khabbaz, F. Mahmood, Three-dimensional echocardiography and en face views of the aortic valve: technical communication, J. Cardiothorac. Vasc. Anesth. 27 (2013) 376–380. [19] S.J. Choo, G. McRae, J.P. Olomon, G. St George, W. Davis, C.L. Burleson-Bowles, D. Pang, H.H. Luo, D. Vavra, D.T. Cheung, J.H. Oury, C.M.G. Duran, Aortic root geometry: pattern of differences between leaflets and sinuses of valsalva, J. Heart Valve Dis. 8 (1999) 407–415.

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Fig. 7. Characterization of type I collagen secretion by SMCs at different conditions by Western blot and immunostaining methods. (a) Type I collagen secretion in intracellular and extracellular matrices by SMCs transfected with retrovirus containing empty vector – HITB5 and vector containing cDNA encoding Hsp47 – HITB5-Hsp47 and cultured in ascorbate-free and ascorbate-supplemented media measured through western blot. SMCs transfected with Hsp47 gene showed higher collagen secretion. (b–e) Immunostained images of collagen type I fibril secreted by SMCs at the different conditions. (b) Amorphous aggregation of fibril was present in ascorbate free HITB5 SMCs culture. (c) Similar morphology was present in ascorbate free HITB5-Hsp47 SMCs culture. (d) More collagen fibril yielded by HITB5 SMCs cultured in ascorbate-supplemented media. (e) Combination of ascorbate and Hsp47 overexpression resulted in elaborate collagen fibril [245]. Reprinting permission will be obtained upon acceptance.

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Although, feedback on in vitro cell-based drug delivery experiments can be achieved fairly, feedback on survivability and functionality of injected cells is difficult to find. These issues should be addressed and 1062 more extensive experiments can be the only solution. Advancement in 1063 molecular imaging of single cell could be helpful in finding the reasons 1064 behind poor efficiency in cell-based drug delivery system. 1060 1061


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Pericyte-targeting drug delivery and tissue engineering.

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