F/Free Essay-Direct-584.Txt 584

1. Introduction

The four valves in the mammalian heart are responsible for controlling the one-way direction of the blood stream from the heart to body and vice versa [1,2]. For this purpose, the inlet and outlet valves open and close continuously during the cardiac cycles. The malfunction of a valve has resulted in blood circulation disorders and may cause serious heart disease and even death [3]. Currently, when one of the valves malfunctions, the end step of medical choice is to replace it with an artificial one that is state of the art. The drawbacks of artificial valves (either mechanical or biological) such as infection, inflammation, thromboembolic, anticoagulation medication requirement and low durability compel the biomedical engineer to introduce a new concept of tissue engineering heart valve (TEHV) [4]. TEHV is an advanced principle to develop a neo-valvular tissue that can mimic the original tissue characteristics with the capacity to grow, repair and remodel in vivo. Based on tissue engineering (TE) principle, a three-dimensional (3D) scaffold is fabricated using the biomaterials as the initial template for the cells. The shape, structure and mechanical properties of fabricated scaffold should resemble the original aortic heart valve [2,5]. Thus the utilized materials as well as fabrication technique can significantly influence the scaffold properties.

Electrospinning is a versatile method for fabricating high quality of continuous nanofibers. The nanofibers are accumulated over each other and shape the electrospun mats. The mats obtained have hierarchy structure that is suitable for scaffolding engineered tissue. The 3D scaffold is then seeded with proper source of cells and biological behaviors are investigated. The hierarchy structure of the scaffold facilitated the nutrient supply to the cells (particularly those far away from the surface) as well as waste removal. Later, the cell seeded 3D scaffold is developed in a bioreactor (in vitro) prior to implementation into the body (in vivo). Obtaining a proper structure and morphology of the nanofibers mats via electrospinning requires optimum setting for electrospinning parameters involved as well as solution parameters. Previously, initial evaluation of electrospinning process was performed as reported in several publications [6,7,8,9,10]. In our latest publication [10], the parameters involved such as flow rate, voltage, collector rotating speed and solution parameters such as concentration and ratio of compounds were evaluated to determine the process behavior in terms of mechanical strength. The analysis of results indicated significance and importance level of the investigated parameters. The maximum elastic modulus obtained was roughly 20 ± 1.2 MPa in such a way that elasticity was around 20%, which is not quite desirable for heart valve tissue engineering. It is expected that the scaffold may lose mechanical properties after cell seeding when the degradation rate starts (roughly 10–12 MPa loss is expected during one-month incubation of the scaffold in particular condition) [11,12,13]. However, statistical analysis shows the optimum point is achievable somewhere near the selected range. For this purpose, mathematical techniques such as steepest ascent/descent and response surface methodology are useful to explore this region. These methods enable us to move around the previous experimental range to find the optimum point. Furthermore, a quadratic regression model will be obtained that can be used for further point prediction. The advantage of TE concept compared to the current prosthesis is the potential to mimic the original tissue and no further medication treatment, inflammation and reoperation during the long time is required.

A significant difference between the left and right side of the heart is that the left heart distributes blood to a wider part of the body, naturally achieving a peak pressure six times more than the right side. Subsequently, the mitral and aortic valves on the left side of the heart are exposed to much higher pressure than the pulmonary and tricuspid valves. The wall of the left side is thicker than the right side. Generally, the majority of valvular dysfunctions are related to the left heart’s valves [14,15,16]. The aortic heart valve operates under a dynamic tensile–shear–flexural loading, and tolerates an elastic modulus within 10–15 MPa [11,17]. The heart pumps around 3–5 L blood with velocity around 1.35 ± 0.35 m/s every minute [18]. The suggested value for shear stress value for aortic heart valve is approximately 1–8 Pa, while the peak of this value may increase to a range of 3–150 Pa. In heart valve function, elasticity as well as strength are the two undeniable characteristics [19,20].

Several uniaxial tests were performed on both human and animal aortic heart valves to recognize the mechanical behavior. The studies mainly reported an elastic modulus for aortic heart valve in circumferential and radial direction. The reports indicate a much weaker elastic modulus in radial direction compared to circumferential direction. This can be attributed to the direction of collagen fibers that are dispersed circumferentially along the leaflets. In addition, the comparison between human and common animal valves mechanical properties exhibited a much weaker behavior for animals. Balguid and Rubbens [11] reported that the elastic modulus of native human aortic valve to be 15 MPa with ultimate tensile strength of 2.6 MPa and 22% of strain in circumferential direction. In the same report, the elastic modulus in radial direction was measured to be 2 MPa with ultimate tensile strength of 0.4 MPa and strain of 30%. Table 1 lists the related works pertaining to the uniaxial tensile tests of human and animal aortic valve. The average required elastic modulus seems to be approximately 14.5 MPa with at least 22% of elasticity in circumferential direction.

Previously, various types of synthetic/natural biomaterials have been used for scaffolding the heart valve. The synthetic/polymeric based materials must be biocompatible, biodegradable and fulfill the required mechanical properties for the dynamic function of heart valve. Synthetic biodegradable polymers such as polyglycolic acid (PGA) [24,25], polycaprolactone (PCL) [26,27], polylactic acid (PLA) [28,29] , and polyglycerol sebacate (PGS) [30,31] have already been reported for TEHV. Sodian et al. [32] engineered tissue scaffolds using PLA and PGA copolymers where the scaffolds obtained were thicker and less flexible than the individual polymers. Scaffolding PLA polymer has been reported by Armentano et al. [29] and, despite the desired tensile strength, it failed during dynamic mechanism of heart valve leaflets due to its high rigidity (brittleness). Van Lieshout et al. [33] reported the application of PCL in TEHV through the electrospinning technique. The biomechanical behavior was good but the low degradation rate of PCL (more than two years) is a hindrance. Masoumi et al. [30] investigated the scaffolding of PGS in TEHV which demonstrates good biodegradability, stiffness and cell adhesion compared to PGA. The PGS tensile strength tests exhibited nonlinear stress–strain behavior. The average elastic modulus of PGS was within the range of 0.025–1.2 MPa, which was not sufficient for heart valve. The main advantage of a synthetic scaffold is the fact that biomechanics and degradation properties can be chemically controlled according to the requirements. Although no biodegradable polymeric materials have been proven to be a desirable substitute for the native valves, work continues to be promising [1,15,24,25,26,27,28,29].

In our previous study [5], maghemite (γ-Fe2O3) loaded thermoplastic polyurethane (TPU)/poly-l-lactic acid (PLLA) was used as a novel mixture for fabricating aortic heart valve’s engineered tissue. The presence of maghemite nanoparticles in the nanofibers latter’s surface improved both cell proliferation rate and also the mechanical properties [10,34]. The elastic modulus of the 50:50% (v/v) neat TPU/PLLA was improved from 3.24 to 4.76 MPa (roughly 32%) when TPU/PLLA was impregnated with γ-Fe2O3. However, it is still far away from the required elastic modulus for a heart valve (around 10–15 MPa). Similarly, the cell viability improved from 72.12% to 95.41% during 72 h of incubation. The results indicated 50:50% TPU/PLLA containing 1% γ-Fe2O3 nanoparticles exhibited overall satisfaction in terms of structural, biological and mechanical properties. The Taguchi analysis was performed to determine the significant parameters influencing the behavior of the selected parameters and the linear regression model was obtained which can only justify the performed experimental range and cannot be generalized for every other process out of that range. In addition, the initial analysis represents half of the curvature (in 3D surface plot) within that range, which means the optimum point is somewhere out of that selected range. Therefore, in this research, the novelty lies in the utilization of initial evaluation and regression model to move from that range to near the optimum point and a second order regression model pertaining to elastic modulus and electrospinning process can be generalized. This was done using the steepest ascent and response surface methodology (RSM). In addition, the cell seeding over the 3D scaffold for 34 days is another novelty of this paper. Besides cell seeding, the new concept is roughness study by AFM techniques for preliminary cell attachment evaluation. Finally, the micro-indentation test was performed after cell seeding to elucidate the elastic modulus loss, which can be attributed to polymer degradation and other factors.

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