Tissue engineered heart valve




Heart valves role importantly in both heart function and circulating system. The heart valves allow a unidirectional blood flow through the heart and distribution throughout the body subsequently. The valve dysfunction might result in catastrophic effects, as the blood backflow would limit the oxygen delivered to the tissues. In the United States, 20,000 people die annually as a direct result of valvular dysfunction and 60,000 valve replacement operations are required in 2001 {{161 Neuenschwander, S. 2004; 165 Fisher, J.P. 2007}}.

Before talking about the Tissue Engineered Heart Valve (TEHV), it is important for us to insight the terminology and the microstructure of the heart valves. Four types of valves are in the heart, they are aortic valve, pulmonary valve, mitral valve and tricuspid valve. Among them, aortic valve and pulmonary valve are classified into semilunar valve, while the left two types of valve are considered asatrioventricular (AV) valves {{162 Mendelson, K. 2006; 165 Fisher, J.P. 2007}}. The semilunar valves contain three histological layers: fiberosa, spongiosa, and ventricularis. The fiberosa layer is located at the upper surface, and composed by collagen fibers (primary type I). The collagen in fiberosa aligned largely circumferentially and acts as the main source of strength in the diastolic pressure. Beneath the fiberosa layer, it comes to spongiosa layer. It is a gelatinous layer, which contain loss connective tissue and is rich in proteoglycans. The main function of fiberosa layer in heart valve is acting as a compressive resistance as well as shear between fibrosa and ventricularis layers. The ventricularis forms the lower surface, and is constituted by elastic fibers and collagens. The elastic fibers in the ventricularis layers allow the layer expand in response to tension in the close state, and retract when valve is opened {{165 Fisher, J.P. 2007}}. The atrioventricular valves have the similar structure with semilunar ones, except the respective outlayers are upside down. In mitral and tricuspid valves, the thick, heavy collagen layer is in the ventricularis side, while the elastic fibers transferred to the atrial side. The appearance of chordate tendineae is also another difference between two classes of heart valves {{165 Fisher, J.P. 2007}}. The arrangement of the collagen and elastic fibers in different layers contribute the leaflet anisotropic behavior {{150 Balguid, A. et al. 2007; 162 Mendelson, K. 2006; 163 Fong, P. et al. 2006}}. The greater circumferential stiffness of the collagens restrict the leaflet move downward, while the lower radial stiffness permits the inward motion towards leaflet coaptation {{150 Balguid, A. et al. 2007}}.

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The majority of treatments for heart valve diseases involve replace the valves by a mechanical, bioprothesis, as well as allograft valves. However, the defects of mechanical valves on continuous injection of anticoagulant as well as inability on growing with the child patients and calcification for the bioprothesis ones limit their further applications. Therefore, TEHV, as a living, unfixed and low immuneresponse heart valve could be manufactured follow this technology, is considered as an alternative approach to replace the traditional therapies of the heart valve diseases. In this paper, we will first have a quick review of the advantages and disadvantages for both mechanical and biological prosthesis, and then discuss about the development of the THEV from several points of views, which include the materials for the scaffolds, cell sources for cell seedings as well as the bioreactors.

Types of prosthesis of heart valves:

Mechanical prosthesis of heart valves:

The first generation of the implantation for the diseased heart valve is the mechanical ones. 3 models of this kind of heart valves are named as Starr-Edward caged-ball valve, Bjork-Shiley tilting disc valve and St. Jude Medical bileaflet tilting disc heart valve.

The Starr-Edward caged-ball valve, the first generation of artificial valve, is composed by a silastic ball which seated at the bottom of the cage. Upon implantation, the valve will be close if the ball remains at the bottom, and open once the ball moves forward into the cage. Bjork-Shiley tilting disc valve is buildup by a single graphite disc between two struts. The graphite disc is coated with pyrolite carbon in order to prevent wearing out after years. It will open and close with the heart pump blood through the valve. St. Jude Medical bileaflet tilting disc valve, the third mechanical valve introduced here, is the most commonly applied mechanical prosthesis in the world. It is constituted by two semicircular leaflets and used as mitral valve replacement. St. Jude Medical is well tolerated by the body. After implantation, only a small amount of anticoagulant is required for the patients to prevent the thromboembolism{{159 Bloomfield, P. 2002}}{{165 Fisher, J.P. 2007}}.

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The advantages of the mechanical prosthesis are obviously, the mechanical valves composed by some strong materials like carbon, titanium and Dacron, are durable, the probability of the reoperation is relatively low. However, the blooding because of the continuous injection of anticoagulant after implantation and the "clicking" noise inside the patients limit their application for the heart valve replacement {{167 Schoen, F.J. 1999}}. Most importantly, for the kids, the implantations of mechanical heart valve are not favourable as the device could not grow up with kids {{165 Fisher, J.P. 2007}}. Therefore, bioprothesis were developed to get over some of the disadvantages.

Bioprothesis for heart valves:

Comparing with the mechanical valves, biological valves are more similar with the nature heart valves. The most commonly used biological valves in present are Hancock porcine one and Caroentier-Edwards bovine pericardial valve. The base rings of the biological valves are coated by an ePTFE covered sewing cuff to facilitate surgical implantation and heeling. Before implantation, biological valves required being treated with glutaraldehyde, which could preserve the tissue as well as kill the cells from other species to prevent the immune response. Therefore, the bioprothesis doesn't need worry about the immune problems. The cusps in biological valves are usually mounted on a stent, which is made by metal or plastic, to simulate the geometric shape of the native heart valve.

Although the biological valves solve the problem on thromboembolism of the mechanical valves, the poor durability limits their application. On the other hand, the biological heart valves required pretreatment with glutaraldehyde before implantation, which will lead calcification and noncalcific degradation.

Tissue Engineered Heart Valve (TEHV):

You cannot have your cake and eat it too. The mechanical valves exhibit excellent resistant on the cyclic loads in vitro, but the problems on thromboembolism limits their applications in vivo. Biological valves, although can cover the shortages of the mechanical valves on immune reactivity, the poor mechanical behaviour caused by calcification also prevent their further application. Therefore, a new approach on valve prosthesis that can overcome the disadvantages on both mechanical valves and biological valves are demanded.

Professor Robert Langer, a chemical engineer, and Dr. Joseph Vacanti, a medical doctor, proposed the concept of tissue engineering in 1993 and believe that one day the organs in our body can be replaced or repaired, and still behavior as the native ones {{160 Langer, R. 1993}}. The biocompatibility (the autogeous cells are used for cell seeding) and durability (the scaffolds, unlike the pretreated biological valves, have excellent mechanical properties) of the TEHV makes it possible have the similar properties with native healthy valves. The general approach for TEHV involved cell seeded onto or within polymeric scaffold, moved and cultured into a bioreactor, which simulates a physiological metabolic and mechanical environment. After implantation, the tissue will be allowed growth and remodeling in vivo {{160 Langer, R. 1993; 162 Mendelson, K. 2006}}.

Scaffold for HETV:

The biodegradable polymeric (either synthetic or natural polymers) scaffold is used to guide the neotissue develop into the right shape as well as mechanical strength for the final implantation in vivo. The degradation time of the candidate polymer should match the tissue formation by cellular components of the tissue-engineered valve {{163 Fong, P. et al. 2006; 162 Mendelson, K. 2006}}.

Synthetic polymers:

Polyglycolic acid (PGA) is the first generation of the synthetic biodegrable polymers that used as scaffold for tissue engineering. It is a highly crystalline, aliphatic polyester, which has a high melting point as well as low solubility in organic solvents {{161 Neuenschwander, S. 2004}}. The hydrophilic PGA homopolymer will lead the loose of the mechanical properties rapidly by the water resorption {{165 Fisher, J.P. 2007}}. Therefore, hydrophobic Poly (Lactic acid) (PLA) was introduced to the PGA backbone to improve the hydrophobicity of the scaffold {{136 Mooney, D.J. et al. 1995; 165 Fisher, J.P. 2007}}. The implantation of the TEHV replaced a single leaflet in the posterior position of a pulmonary artery was placed by a TEHV which used PGA and PLA copolymer (poly(GA-co-LA)) as the scaffold. The resulting device was completely degraded after 6 weeks, and a new viable tissue had been developed in vivo, but the scaffold was too stiff and thick to develop a long term application {{137 Sodian, R. et al. 2000}}.

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To avoid the mechanical deterioration in vivo by using poly(GA-co-LA ) scaffold, Sodian et al. {{137 Sodian, R. et al. 2000; 138 Sodian, R. et al. 2000}}developed a new scaffold by using a thermoplastic polymer - poly(hydroxyalkanoate) (PHA), which was constructed by valved-conduit-shaped molding and generated a porous structure by a salt leaching technique. The heart valve was seeded with ovine vascular myofibroblasts and endothelial cells and incubated for 2 weeks in cell culture medium. The TEHV could open and close in the pulsatile flow bioreactor in vitro, and functioned normally in a sheep model for 17 weeks. However, the deficiencies of elastin and endothelial cell coating in the neotissue as well as the insufficient degradation suggested the PHA is not a perfect material for scaffold in TEHV, the revolution is still required.

The shortcoming of the PGA derived scaffolds were degradation too fast, while the PHA derived ones, on the other hand, were degradation too slow. Then the combination of the PGA and PHA like polymer poly-4-hydroxybutyrate (P4HB) (also a biodegrable thermoplastic polymer) should exhibit the excellent durability as well as the appropriate degradation rate. Hoerstrup et al. {{169 Hoerstrup, S.P. et al. 2000}} used the nonwoven PGA coated with a thin layer of P4HB to sever as a new scaffold for TEHV. Like the PHA scaffold, the porous structure in PGA/P4HB was generated by salt leaching. Like the PHA scaffold, the PGA/P4HB then was undergone cell seeding with ovine vascular myofibroblasts and endothelial cells and incubated for 2 weeks in a pulsatile flow bioreactor. After implantation for 20 weeks, the uniform layered cuspal tissue with endothelium was found during the histological exam. The degradation of the polymers was completely by 8 weeks, but the extracellular matrix content like collagen, elastin, and glycosaminoglycans, as well as DNA were higher than the native tissue.

Hydrogels, like polyvinyl alcohol (PVA), are bioinert, they have a high water content and tissue-like elasticity. In addition, the hydroxyl groups on PVA can be modified to attach growth factors, adhesion proteins, as well as other molecules. However, due to the relative high hydrophilicity, the ability on the cells attachment to hydrogels is limited. Therefore, Charles et al. {{139 Charles, R. et al. 2002}} grafted PLA with PVA and photocrosslinked it to form a biodegrable hydrogel and planed new class of hydrogel used as scaffold for TEHV. The incorporation of hydrophobic PLA would slow the degradation of the PVA hydrogel by excluding water and preventing the hydrolysis. The result showed much more cells were attached and assumed to the long, spindle morphologies which were similar to the valve cells. The completely degradation of the crosslinked PVA scaffold was 80 days, which was longer than the PGA/P4HB, but the weaknesses on lower proliferation rates of valve interstitial cells , and protein expression levels limit its applications.

Electrospinning is an efficient technique to produce polymeric fibers with diameters in nano or micrometers. The high surface tensions and porosity as well as the perfect mechanical properties of the produced nanofibers make the electrospinning be possible applied into the tissue engineering field. Gaudio et al. {{155 Gaudio, C.D. 2007}} reported an early results of a hydrodynamic characterization of a stentless electrospun PCL heart valve. During the in vitro test in a bioreactor, the leaflets could were able to open synchronously and a correct close at the diastolic phase. The present studies on the application of electrospinning in TEHV is still immature, some drawbacks, like the slight rotation of the leaflets in the diastolic period {{155 Gaudio, C.D. 2007}} implied, for the success future surgical application, this technique required further improvement.

Decellularized leaflet scaffolds:

The synthetic polymers for THEV scaffold perform excellently on mechanical properties, but the degradation of the polymers might fill the space among the cells and extracellular matrix (ECM) by fibrosis or scars, the replaced heart valves might then contracted or distort during maturation {{162 Mendelson, K. 2006}}. While the bioprothesis methods, as we mentioned before, were suffered on the drawbacks like calcification because of treating with glutaraldehyde before implantation to fix the cells. As a revolutionary method on heart valve prosthesis, the decellularized leaflet scaffolds were designed to remove either exogenous or autogenous cells, then reseeding autogenous cells onto the ECM {{165 Fisher, J.P. 2007}}. This approach, unlike the bioprothesis, it does not require fabricating or molding procedures, and allow the implanted device grow, repaired and remold in vivo, and elicit almost no antigenic reactions.

A major area as well as the most important one in the development of this kind of scaffold is the method to remove the cells from the initiate heart valves. To meet such requirement, approaches involve using ionic and nonionic detergents, as well as the enzyme treatment were developed recently {{165 Fisher, J.P. 2007}}. To developed such a scaffold, Zhao et al. {{143 Zhao. D.E. et al.}} attempted apply the detergent and enzymatic extraction to remove the cellular components in porcine aortic valves. In detail, the porcine valves were firstly treated with 1% triton X-100 at 4 oC for 12 hours, and immerse into a tris chloride buffered solution contained DNAase I, RNAase and 0.02% EDTA at 37 oC for another 12 hours. Later, the treated ECM was seeded with endothelial cells and myofibroblasts from canine right carotid artery and implanted into dogs. Although there was no calcification found after implantation in 70 days, the dogs scarified in 10 weeks suggested the decellularization procedure may introduce the change on the mechanical properties of the ECM. In order to discover the relationship between agents used in decellularization and changes of valve properties, Liao et al. {{140 Liao, J. 2008}} ran experiments by applying ionic detergent (SDS), non-ionic detergent (Triton X-100) and enzymatic agent (trypsin) as decelluarized agents. Each sample that treated by the different agents were subsequently undergone mechanical tests as well as histological analyse to investigate the defects cased by the different decelluarization procedure. The results suggested all these three treatments were able to remove the cells in the original valves completely, but their effects on the mechanical properties were quite different. The valves decellularized with SDS did not cause any statistical differences of thickness compared with the native valves, while the treatment by the trypsin induced the leaflet thickness increased and Triton X-100 treated valves were thinner compared with the native ones. The histological results indicated the trilayer structure of the SDS treated leaflets were similar with the native ones, however, the in the Triton X-100 and trypsin treated samples, the spongiosa layer (function as dampen the vibrations in the fibrosa associated with leaflet flexion during closure) was depleted, which implied compared with the SDS treated ones, a significant change on mechanical properties of Triton X-100 and trypsin treated samples would be occurred {{163 Fong, P. et al. 2006; 140 Liao, J. 2008; 165 Fisher, J.P. 2007}}.

Nature polymers:

Although the majority of scaffolds used on TEHV are composed by either biodegradable synthetic polymers or decellularized leaflet, scaffolds manufactured by nature polymers are developed as the substitution for above two classes of scaffolds. The biodegradable polymers, such as poly (lactic acid), might elicit cytotoxicity because of their degradation products would lower the pH of the culture medium {{154 Ye, Q. et al. 2000}}. For the decellularized leaflet, the reduction of the mechanical properties of the scaffold might be observed after the decellularization treatment {{140 Liao, J. 2008}}. Nature polymers, unlike synthetic polymers, are completely compatible with the body. Therefore, the application of nature polymers on the scaffolds for TEHV is likely to be the next direction of TEHV development.

Collagen and elastin are two classes of proteins, which are discovered abundant in ECM. Although collagen in early works was used as the three-dimensional scaffold for TEHV and providing good cell attachments, the application collagen scaffold in TEHV was limit. Since the induction of the platelets activation and the thrombus was lead by the open surface of collagen, it was always coated with synthetic polymers, like P4HB, to improve the biocompatibility {{153 Stamm, C. et al. 2004; 154 Ye, Q. et al. 2000}}. On the other hand, the elastin, was found as a factor which would elicit calcification once it was biodegraded {{149 Bailey, M.T. et al. 2003}}, therefore, other types of natural polymers, like fibrin gel, were undergone several researches to develop a natural polymeric scaffold for TEHV {{154 Ye, Q. et al. 2000}}. Ye et al. incubated human myofibroblasts within fibrin gels. The concentration of proteinase inhibitor aprotinin was controlled to optimize the degradation rate of the three-dimensional fibrin gel structure. In addition, the results showed the gel structure served as semi-permeable membranes, which separate the cell contact with the medium so the collagen and other species ECM components could accumulate inside the gel rather than diffusing into surrounding medium {{154 Ye, Q. et al. 2000}}. However, the unfavorable thickness of the developed tissue was unstable to create cardiovascular graft on arterial side, another natural polymer, chitosan, was then introduced into TEHV.

Chitosan is a potential natural polymer that capable to be used in the tissue engineering field owing to its properties on biocompatibility, solubility, and porosity. Cuy et al. {{156 Cuy, J.L. et al. 2003}} designed casting the chitosan solution into Polystyrene (PS) multiwall plates to form a porous film. Before casting, the PS plates were treated with proteins in purpose to enhance the adsorption of valvular endothelial cells (VEC) on the chitosan surface. By compared with the untreated chitosan film, the cell adhesion of the protein precoated chitosan was significantly higher. For the gelatine, PLGA and PHA treated as same as chitosan did, chitosan exhibit highest cell adhesion, but minimally enhanced VEC growth. By incorporation of collagen IV into chitosan solution and then casting onto the PS plated, the film was observed improve the VEC morphology and growth over chitosan alone. Therefore, chitosan/collagen IV composite film combined with appropriate protein pretreatment might be suggested as a prospect scaffold for TEHV.

Cell Sources:

Once the material of the scaffold is decided, it then comes to the cell seeding step. The source of cells for TEHV is supposed to be autologously derived from the patent and elicits little or no immunoresponse after implantation {{165 Fisher, J.P. 2007}}. In the early stage of the development of TEHV, the scientists tried to seed more than one class of vascular cells within the scaffold {{137 Sodian, R. et al. 2000; 138 Sodian, R. et al. 2000}}, but genetic studies pointed out that the valvular endothelial cells (VEC) have many transcriptional differences from vascular endothelial cells. Some works also blamed the failure of the TEHV seeded with the vascular endothelial cells on the difference from the cell proliferation between VEC and vascular endothelial cells. Therefore, VEC or valvular interstitial cells (VIC) seeding was developed to optimize the biocompatibility of the THEV {{165 Fisher, J.P. 2007; 139 Charles, R. et al. 2002; 156 Cuy, J.L. et al. 2003}}. However, cell very few researches attempted to develop VEC or VIC in TEHV. The harvesting of either two classes of cells requires the sacrifice of intact valvular donor tissues. Moreover, most of patients do not have the health valve suitable for cell harvesting {{144 Fang, N.T. et al. 2007; 157 Kadner, A. et al. 2002}}. Stromal cells and mesenchymal stem cells (MSC) were then reported to be the optimal cell sources in THEV.

Kadner et al. {{157 Kadner, A. et al. 2002}} seeded human bone marrow stromal cells on a PGA/P4HB scaffold and incubated in static state for 14 days. The cells exhibited resembling phenotypic characteristics of myofibroblasts and ECM as vascular cells seed TEHV did. Modified experiments were later conducted to evaluate the in vivo test results as well as mechanical properties of stromal cells seeded TEHV. The stromal cells were isolated from ovine bone marrows, then seeded onto the PGA/P4HB scaffold and incubated at static state. One week later, the TEHV was moved in bioreactor and growth under the mimicked physiological environment. The mechanical test revealed that the TEHV exhibited a resembling behavior as native heart valve did {{142 Perry, J.E. et al. 2003}}. Fraser et al. {{158 Fraser, W.H. et al. 2005}} also observed the similar results by using the same cell source seeding with PLGA scaffold. Also, they performed the Noninvasive echocardiographic imaging to assess the geometry of the implanted TEHV. The trivial regurgitation of the implant indicated the TEHV manufactured by seeding MSC in a synthetic biodegradable scaffold might develop as a potential therapy for heart valve diseases. Moreover, the recent observation of the cell proliferation, myofibroblasts like morphology, histological three layers structure, collagen (I, III, IV), laminin as well as alpha-actin after the 12 weeks implantation of the endothelial progenitor cells (EPC) seeded P4HB, indicated the EPC derived from human umbilical cord blood could also be considered as a candidate cell source for TEHV {{145 Sodian, R. et al. et al. 2006; 144 Fang, N.T. et al. 2007}}.


Bioreactors are widely used in THEV and aim to provide a mechanical stimulation before implantation. Compare with the static incubation, the THEV cultured in the bioreactor demonstrated great improvements on mechanical properties, collagen formation, cell proliferation, as well as the augment of the ECM {{145 Sodian, R. et al. et al. 2006}}. As the bioreactors provide a dynamic environment rather than a static one, the bulk degradation rate of the scaffold is also increased. The most popular used bioreactor in the world was first developed by Hoerstrup et al. It was made by PMMA and composed by two principal chambers: the bottomed air chamber and the upper level cell media fluid chamber. Between these two chambers, there was a silicone diaphragm composed of "super stretch" silicone rubber. A respirator pump was designed to connect with the air chamber while the fluid chamber was separated into two compartments. The connection between the lower compartment and topper compartment was a tube, where a second removable silicon tube was mounted at the top. Once in operation, the pulsatile fluid would be created by pumping air into the air chamber and propel the cells in fluid chamber to contact with the TEHV in the silicon tube {{146 Hoerstrup, S.P. et al. 2000}}.

The bioreactor developed earlier was on a fluid base and failed to simulate the loading condition at the implantation site. Moreover, the mechanical properties of the valve were changing during the incubation, the induced tissue deformations were unknown, and varied during the tissue culture {{163 Fong, P. et al. 2006}}. Therefore, another kind of bioreactors, the controllable strain based diastolic pulse duplicator, was invented to give the more optimal incubation conditions. The new bioreactor was composed by two identical chambers, and each chamber contained 6 culture wells. In each well there are 4 stainless-steel stationary posts used to fix the TEHV and perpendicular to the central channel. Between the two chambers, there assembled an actuator, and the piston of the actuator was rigidly coupled to a cross-arm in the form of a T-junction. Each arm was then bifurcated and extended into a chamber and terminated in six fingers. The flexure pin was attached onto the end of each finger and inserted to bracket the samples in the middle. Therefore, the samples in wells would be subjected to either unidirectional or bidirectional three point flexure. Flexure strain can be controlled by the parameters, like frequency, amplitude, acceleration and deceleration of the actuator and mimicked the physiological conditions. The result tissue exhibited more pronounced and organized formation with superior mechanical properties compared with the unstrained controls {{170 Engelmayr,George C. 2003}}.

In vivo tests in animal models:

The assessments of the biocompatibility of TEHV cannot be fully established in vitro, the physiological environment is more complicated than the conditions mimicked in bioreactors. The cells, salts, as well as other biomolecules would affect the biocompatibilities of the implants with the organism. Therefore, the tests on the inflammatory, immunologic and calcific responses in vivo are extremely important to evaluate the implants' biocompatibilities {{166 Ratner, B.D. et al. 2004; 165 Fisher, J.P. 2007}}.

As we mentioned earlier, both the mechanical and biological prosthesis of heart valve subjected to the problems on the poor biocompatibility, however, their replacement TEHV exhibited overall better biocompatibility but still have some drawbacks. Even the TEHV exhibit good biomechanical properties in vitro, we still don't have the confidence that the valve would perform as good as health heart valve in vivo. For example, collagen, the main component for the ECM and exhibit ideal mechanical property as well as the cell adhesion in vitro. However, the in vivo test indicated the induce of the thrombus, which was lead by the activation of platelets by the collagens, make it inappropriate as a candidate for the TEHV scaffold {{153 Stamm, C. et al. 2004}}.

A good in vivo test should be established on right experimental object at long term duration {{153 Stamm, C. et al. 2004; 158 Fraser, W.H. et al. 2005}}. In most of the cases, the tests were built on ovine model {{136 Mooney, D.J. et al. 1995; 158 Fraser, W.H. et al. 2005}}, because of the mechanical loading in ovine is similar to that of human beings. There did have some tests ran successfully in the rats' model, but when it came to the ovine model, it failed because of the implants suffered higher mechanical loading in ovine {{153 Stamm, C. et al. 2004}}. The test duration should be long enough (usually 4 months in ovine model) {{158 Fraser, W.H. et al. 2005}} to get the appropriate information on device's biocompatibility.

The most commonly used in vivo tests involve the exams on calcification, cell migration {{165 Fisher, J.P. 2007; 153 Stamm, C. et al. 2004}}, endothelialisation, inflammatory, as well as thrombus formation {{153 Stamm, C. et al. 2004; 158 Fraser, W.H. et al. 2005}}. Echocardiography was also applied sometimes to give undestructive visualized information on the implants in vivo {{158 Fraser, W.H. et al. 2005; 153 Stamm, C. et al. 2004}}. However, as most in vivo tests are destructive ones and require explantation to measure the influence of the physiological environment on the mechanical properties {{148 Kortsmit, J. et al. 2009}}. Therefore, an undestructive and noninvasive assessment for TEHV is eagerly needed for the future research.


After several years of development, TEHVs have become more and more important in the therapy of valve diseases. Compare with the allograft and bioprostheses, the TEHVs have advantages on ample supplement, less immunological rejection as well as the potentials on growing, remolding and repairing {{165 Fisher, J.P. 2007; 161 Neuenschwander, S. 2004}}. However, a successful TEHV should be composed by specialized cells and the ECM, which is remodeled in response to changes in local mechanical forces, should have comparative mechanical properties, so the origin of the cell source, in vivo analysis technology, choosing of the animal model as well as the development of the scaffold materials are important to meet the clinical requirements.

One of the important laboratory considerations for TEHVs is how to choose the right cells seeded with the scaffold and track them noninvasively. The seeded cells are supposed to remain living and attached to the scaffold after implantation. The molecular imaging might be used to realize tracking of the migration, proliferation as well as the function of the MSCs in vivo. It might be possible that the MSCs could be labeled with iron fluorophore particle(IFP) in vitro and provide a MRI contrast in vivo to assess short time andlong term localization {{171 Hill,Jonathan M. 2003}}.

Also as we discussed at the in vivo test section, the ovine are seemed to be the optimal experiment objects for evaluating the biocompatibility in vivo, but they produced an exuberant fibrotic tissue in response to the cardiovascular implants. Therefore, the implantation of valve in ovine model always overgrow more rapidly with fibrotic tissue than that in human cases and fail to provide accurate information {{161 Neuenschwander, S. 2004}}. More genetic and physical type similar animals, like monkeys and other primate animals might be used to replace the ovine in the in vivo experiments.

Finally, the scaffold used in TEHV is also needed to be improved. Two classes of materials, biodegradable synthetic polymers and natural polymers (include the decellularized leaflet) are used in the scaffold manufactures. However, both of them are not prefect and undergoing some problems on biocompatibility, so revolutions on the scaffold materials are still required. Smart polymers maybe applied in the future. One possible approach is applying the injectable gel as the novel scaffold to replace the conventional ones. For example, some copolymers, like PCL-b-PEG-b-PCL and PLGA-PEG, are thermal sensitive and can be potentially used in TEHV {{172 Bae,Soo Jin 2005}}. The transition from sol to gel around the body temperature allows the cells in polymer solution evenly at room temperature and inject into body to form a gel in situ. The highly encapsulated cells ensure there is no cell migration after implantation, and the small wound would reduce the risk on inflammatory after the surgery. Another potential smart polymer would be used in TEHV area is poly(N-isopropylacrylamide) (PNIPAM). The polymer would covalent grafting onto a PS plate, and showing hydrophobic at incubation temperature 37 oC to allow the cells attach and proliferate on the surface. However, upon the temperature reduce to the lower critical solution temperature at 32 oC, the surface would turned from hydrophobic to hydrophilic, and the cells would spontaneously detached from the plate without the enzymatic treatment. The nondestroyed cell sheets would be then manufactured to three dimensional neotissue. The enzyme free treatment in cell haverst as well as the scaffold free during the implantation would believe to enhance the biocompatibility of the artificial valve {{173 Yang,Joseph 2005}}.