Valvular heart valve disease caused a major problem worldwide, result in increasing rate of morbidity and mortality (Rashid et al., 2004). Each of the four valves of heart can be affected by dysfunction (Mol et al., 2009). The common solution is the replacement of diseased valves either mechanical or bioprostheses. However, they are associated with shortcoming. Patients with mechanical valves are associated with long term of anticoagulation therapy which leads in higher risk of complications (Hjortnaes et al., 2009). Bioprosthetic valves are less durable, enable to grow and repair and lead to calcification. Moreover, they are associated with reoperation especially with young patients. None of these valves are capable to restore native valves functions.
To overcome replacement shortage of heart valve, in the mid of 1980s in the USA, tissue engineering concept was proposed as alternative solution and has the capability to produce autologous heart valves (Matsumura, 2003).
Tissue engineering of heart valves is “manipulation of biological molecules and cells for the purpose of creating new structures capable of metabolic activity”. (Vesely, 2005).
Worldwide, two main approaches have been used to achieve the desirable and clinical needs of heart valves; regeneration and repopulation (Vesely, 2005).
In the first approach, decellularized xenograft (ECM proteins; collagen or fibrin) or allograft served as valve biological scaffold, mixed with autologous cells (myofibroblast and endothelial) has been seeded in vitro (Mendelson and Schoen, 2006), (Neuenschwander, 2004).
However, limitations of this approach are; the construct lack the ability to grow, long term mechanical properties could alter physical properties of the native valves and a possibility of disease transmission (Neuenschwander, 2004). Moreover, inflammatory reaction due to ECM proteins used may deteriorate the scaffold result in biomechanical damage (Mendelson and Schoen, 2006).
A biodegradable synthetic scaffold is the traditional approach in which cells of a specific phenotype are transplanted onto a bioresorbable scaffold in the shape of heart valves (Vesely, 2005).These biocompatible materials offer several advantages; Due to degradation products, they have neither side effects nor immunogenic reactions. They are harmless and easy to handle. At the site of implantation, the remaining autologous living structures after complete biodegradation of the scaffolds have the ability to remodel, to adapt and to grow (Neuenschwander, 2004).
The danger of the transmission of new infections and immunogenic reactions is minor as a result of utilization of autologous cells which lead to identical bioengineered tissues to host cells (Neuenschwander, 2004).
Extracellular matrix material is formed by the cells after attachment and the polymer scaffold degrades (Fuchs et al., 2001).
As shown in (Shinoka et al., 1995) mixed cell population of endothelial cells and from ovine arteries were isolated and separated from each other by fluorescence-activated cell sorting. An acetylated low-density lipoprotein marker was used to label endothelial cells. Myofibroblasts were seeded onto polyglycolic acid scaffolds which then seeded with endothelial cells. Before new tissues being regenerated, the degrading scaffold cannot tolerate pressures of the left ventricle (Vesely, 2005). Therefore, the constructs were implanted in sheep in place of the native right posterior leaflet of the pulmonary valve. This approach results in trivial pulmonary regurgitation in autografts but moderate in allografts and doesn’t show stenosis (Shinoka et al., 1995).
Repopulation is the less popular approach. Complex structure (collagen) is fabricated by manipulating biological molecules. Prior to implanting, cleaning off the porcine aortic valve from cells is required, and leaving intact the connective tissue matrix. Then, the acellular matrix of the patients’ cells is repopulated, stimulated and thus result in creating a living tissue resemble to the native tissue (Vesely, 2005).
In tissue engineered valve applications, mesenchymal stem cells and differentiated tissue-specific cells (including circulating endothelial progenitor cells or smooth muscle cells) are the two main types used. They are harvested from either patients or experimental animals (Vesely, 2005). Stem cells along with the appropriate matrix are expected to provide a broader source of either autologous or allogenic cell lines once differentiated to the proper end point. Thus, they are beneficial for therapeutic use in the cardiovascular field (Vesely, 2005).
Animals’ tissues such as canine and pigs are conducting in a research with variability in successful implantation. Mesenchymal stem cells for TE of ECM scaffolds have been obtained from canine tissues. Also, due to presence of type 1 and IV collagen and fibronectin, ECM from pig’s tissues showed ability to grow to a single layer (Rashid et al., 2004).
In most approaches, veins (saphenous vein) or peripheral arteries (radial artery, mammary artery) are efficient sources of cells. Mixed vascular cell populations give rise to myofibroblast and endothelial cells lines. myofibroplasts derived from arteries show decreased proliferation in monolayer culture and ECM formation when cultivated on three dimensional structure compared to cells obtained from veins (Neuenschwander, 2004).
Another promising alternative source of heart valves can be derived from progenitor cells derived from peripheral blood and bone marrow cells (mesenchymal stem cells). MSC can be differentiated into different tissues, are easy to obtain, shows high proliferating capacity in vitro and faster growing than vascular derived cells.
They produce well developed ECM after cultivation under bio-mimetic conditions (Neuenschwander, 2004).
There are two principal choices of scaffold; naturally or synthetic polymeric scaffolds. Exogenous ECM scaffolds are required to provide mechanical support until the target newly tissue is formed and become stable. Designing is based on the target cells population needed to implant at the injury or diseased sites. In vivo, they are degraded alongside the implanted functionally tissues grow and organize their matrix structure (Kim et al., 2010).
Non-toxic, biodegradable and biocompatible scaffolds serve as a temporary matrix for the seeded cells due to their own a highly porous microstructure that supply nutrients required for growth and necessary for waste removal. They possess structural integrity to withstand in vitro and in vivo loading. The most widely used in HVE are polyglycolic acid (PGA), polylactic acid (PLA) and their copolymers (PGLA). PGA is highly resorbable, linear, a highly crystalline and have a high melting point. PLA reduce the rate of hydrolysis and have limited water uptake. In all of that, to facilitate tissue growth, growth factors are integrated in the scaffold. However, a number of disadvantages must be overcome; slow or incomplete degradation that cause inflammation, limited nutrients and oxygen delivery to deep cells and fibrosis (scar) occupied the space formed after scaffold degradation (Mendelson and Schoen, 2006).
Natural scaffold composed of decellularized xenograft or allograft tissue or ECM components such as collagen or fibrin (Mendelson and Schoen, 2006). Fibrin gel scaffold can be designed into a valve structure to form a biodegradable, autologous scaffold. Blood is the source of fibrin gel and mould into autologous manner to work against immunogenicity. Seeding cells into fibrin scaffold results in good tissue development with viable fibroblast. However, fibrin scaffold would lack mechanical strength (Knight, 2004). Furthermore, in a process similar to the contraction of a wound healing, cells entrapped in collagen gels compact the gel, improve their property and increase density (Mendelson and Schoen, 2006). But, any scaffold made from collagen alone would like to degenerate very quickly due to in vivo forces affected heart valves (Knight, 2004). Moreover, hyaluronan is a glycosaminoglycan polymer with a repeating disaccharide structure and being used as biocompatible material ECM scaffold. It imparts viscoelastic properties and applies pressure that gives tissues compressive resistance (Vesely, 2005).
Several methods are used to sterile materials utilized in process of seeding and tissue harvesting. 1) Dry heat sterilization in which items are placed in an oven for one hour at 160T to be sterilized by dry heat. 2) Moist heat sterilization is used to autoclave objects and solutions not suitable for dry heat sterilization for 20 minutes at 121 T, 15 pounds per square inch (psi). 3) 0.2ptm filters are used to sterilize solutions not suitable for autoclaving (Knight, 2004).
Animal model is an essential part of biomedical research to approve tissue engineered devices by FDA in order to carry on clinical trials. Using animal tissues as sources relies on many factors; cost, ethical considerations, availability and the nature of the tested tissue. In Zilla study (Zilla et al., 1994), baboons were used to study the proliferation of seeding (endothelial cells) EC on polytetrafluoroethylene (PTFE) grafts. Compared to the control of unseeded graft, seeded graft showed persistent confluent EC layer through time with the aid of fibrin glue enriched with RGD (Zilla et al., 1994).
Due to similar anatomy and physiology to that of humans, pigs have been used widely for experimental study. Also, they are cooperative without general anesthesia. They are capable of rabid growth so limiting the time required for TE construct (Rashid et al., 2004).
Biodegradable polyglycolic acid (PGA) scaffold was treated by sodium hydroxide and modified seeded bovine SMC and EC were used by Niklason and co-workers in after 24 days of implantation in swine model.
Ovine and caprine are also models to study TE due to large size and easy access to the carotid artery in the long neck. They can be used for long term study because adult animal can’t grow (Rashid et al., 2004). In a number of researches, sheep model was used widely but ordained for failure as a result of exuberant fibrotic response to implants (Vesely, 2005). Compared to humans, implants grow rapidly with fibrotic tissue in sheep (Schoen, 2011).
In preclinical testing, the choice of animal model is a challenge owing to immunologic considerations (Mendelson and Schoen, 2006).
Various parameters determine the optimal conditioning protocols; the scaffold, the magnitude and types of mechanical cues, the sensitivity of cell to the used scaffolds. Bioreactors in TEHV have been developed to improve tissue formation, organization and functions and to stimulate dynamic mechanical of the TEHV. Moreover, to mimic native excitation-contraction coupling, electrical stimulation has been used. Additionally, to mimic the diastolic phase of the cardiac cycle, a diastolic pulse duplicator bioreactor has been developed which result in dynamic tissue straining (Sacks et al., 2009).
In October 2000, CE Mark approved use of the CryoLife Synergraft. It was similar to decellularization matrix approach by removing cellular antigens using extraction and dissolution. It was expected to sound mechanically as acellular matrix. Unfortunately, complications related to stenosis, inflammation and valve rupture result in death and thus withdrew from the market (Vesely, 2005).
Till date, no EHV constructs have been commercialized. Before translation of the construct to patients, numerous steps must be considered and assured laboratory. For example; ethical issues, safety, efficacy and quality of the product should be evaluated. Additionally, medical devices interactions results such as; thrombosis, infection and inflammation will have to be accepted. There is a need to develop tools to monitor the fate of transplanted and endogenous cells, biomarkers to evaluate the patient’s variability to implantation. in all of that, suitable approach is required that ensure efficiency and safety (Schoen, 2011).
Three main issues determine the success of tissue engineered heart valve; 1) sources of cells, 2) the (matrix) scaffold that serves as a guiding structure and determines the three dimensional shapes of tissue development and cell attachment and 3) the optimal culturing condition for cell growth. State of art of TEHV today is still on research, significant challenges must be solved before start in clinical application.
FUCHS, J. R., NASSERI, B. A. & VACANTI, J. P. 2001. Tissue engineering: a 21st century solution to surgical reconstruction. The Annals of thoracic surgery, 72, 577-591.
HJORTNAES, J., BOUTEN, C. V. C., VAN HERWERDEN, L. A., GRUNDEMAN, P. F. & KLUIN, J. 2009. Translating autologous heart valve tissue engineering from bench to bed. Tissue Engineering Part B: Reviews, 15, 307-317.
KIM, B. S., PARK, I. K., HOSHIBA, T., JIANG, H. L., CHOI, Y. J., AKAIKE, T. & CHO, C. S. 2010. Design of artificial extracellular matrices for tissue engineering. Progress in Polymer Science.
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