Translational Pathway for Tissue Engineered Aortic Valves
A. The Need for Translational Platforms
The Renaissance era redefined scientific methods through the convergence of observation, engineering and mechanistic approaches. Modern sciences have evolved and matured into basic and applied research. As a result, applying breakthroughs of fundamental research into meaningful outcomes in practice is seldom a seamless process.
In medicine, translational research proposes a paradigm shift to bridge basic and applied research by establishing a multidisciplinary vision that combines disparate fields of laboratory science with the applicability of clinical research. In other words, translational research aims to identify the most effective path from laboratory bench to patient bed.
To effectively develop strategies and solve clinical problems, translational research needs multi-skilled ecosystems in which experts of different fields generate and cross-analyze data through various methods and techniques and through feedback cycles between basic, applied and clinical scientists. By its structure and panel of assets, such a translational platform defines and implements the necessary actions, process or methods.
In the context of aortic valve replacement, specific aspects have to be considered to efficiently link clinical needs and expertise with basic and applied sciences. The aim of this chapter is to provide an overview of such aspects, related to the specificity of the aortic valve tissue, its pathologies and diseases, and the resulting clinical need for replacement. The different replacement approaches that have been investigated and applied in recent decades, from mechanical valves to the most recent and promising valve tissue engineering strategy will be introduced in relation to their original clinical objectives and resulting assets and liabilities. Finally, the specific components of a translational platform for aortic valve replacement will be detailed, from assessment criteria, markers and techniques, to preclinical testing.
B. Aortic Valve Characteristics: Structure and Function
Aortic valves perform an extremely complex functional role. Not only do the valves ensure the unidirectional flow of blood out of the left ventricle into the ascending aorta, but they are also crucial for optimising coronary blood flow and preserving myocardial function. It is now widely accepted that the function of the aortic valve is regulated by complex mechanisms that are initially evident during embryonic development and during adaptation of the valve after birth and growth into adulthood. Clinical experience with the Ross Procedure illustrated the importance of a ‘living valve’ (1). The sophisticated function of valve and its durability depend largely on the biological properties of the valve at cellular and molecular level (2). An elegant example of this can be seen in the observations of Hashadigate et al. who shown that the valve anticipates the movement of blood out the left ventricle by beginning to open prior to the forward flow of blood through the valve (3). The role of living cells with the valve has recently been extensively reviewed (4).
Similarly to the three other types of heart valves in the human heart, aortic valves provide the crucial function of ensuring a unidirectional blood flow through the cardiovascular system. However, unlike the right-sided valves of the heart (tricuspid and pulmonary) that direct venous blood to the lungs at low pressure (less than 20 mm Hg), the aortic valve conveys oxygenated blood to the body at a high pressure (~ 120 mm Hg). Considering that heart valves open and close over 30 million times per year and that 3 to 5 litres of blood are delivered by the heart each minute, the aortic valve is subjected to important and repetitive tissue strains and surface shear stresses (5). To cope with such mechanical requirements, the aortic valve possesses a very specific and adaptive architecture that is continuously remodelled by the valve constituting cells.
Aortic valves are macroscopically composed of supporting structures in the aortic root, a fibrous annulus connected to the left ventricular outflow tract, and three semilunar cusps (or leaflets) (Fig. 1-A).
The aortic root is a fibrous tissue on which are attached the aortic leaflets. Directly above each cusps are three depressions (sinuses of Valsalva), two of which connect to the coronary arteries. The tri-layered structure of the root is akin to large arteries, with an intima, media and adventitia. Similarly, the cell population is comprised of intimal endothelial cells, medial smooth muscle cells and adventitial fibroblasts.
The aortic valve leaflets are thin and compliant tissues that join during diastole to close the valve. They are attached to the annulus of the aortic root in a parabolic fashion, meeting each other at the commissure. They are composed of three layers of distinct organisations and compositions (Fig. 1-B). Facing the aortic root is a fibrous layer, the fibrosa, mainly composed of collagen bundles organized anisotropically in the circumferential direction, and a less organized elastin structure (6). The opposite side of the leaflet (towards the left ventricular chamber) is the ventricularis fibrous layer characterized by a richer content of radially orientated elastin, and a less organized collagen structure. A loose and gelatinous third layer, the spongiosa, separates the two fibrous layers and is rich in glycosaminoglycans.
Each layer plays an important role during the systole/diastole cycle of opening/closing of the valve. Through its collagen content and circumferential organization, the fibrosa provides the strength and stiffness of the aortic valve during diastole, minimizing flopping of the cusp centers, preventing regurgitation and allowing to withstand repetitive biomechanical loads (7). The spongiosa acts as a shock absorber by accommodating the shape changes of the cusp during the valve cycle while the elastin content of the ventricularis allows for valve extension in diastole to maximize coaptation area and valve recoil during systole to minimize surface area (8).
Two principal types of cells are present in the healthy aortic valve: valve interstitial cells (VICs) that are disseminated throughout the entire valve and valve endothelial cells (VECs). Small populations of smooth muscle cells and nerve cells are also present (9-11).
Interstitial cells are responsible for the synthesis and reorganization of the valve extracellular matrix (ECM), allowing to continuously cope with mechanical strains applied on the aortic valve. Although different sub-cell populations have been described (fibroblast, smooth muscle cells and myofibroblast) (12) VICs generally exhibit a quiescent fibroblastic phenotype in the healthy valve. In response to injury, or during growth and remodeling, VIC phenotype changes to an activated myofibroblastic state characterized by an increased production of ECM proteins as well as ECM degrading enzymes (matrix metalloproteinases, cathepsins and tissue inhibitor of metalloproteinases) (8,13,14). Myofibroblast activation can be as well triggered by mechanical stimulations (15,16) and valve stiffness (17). The important proliferation, ECM protein and glycosaminoglycan turnover of VICs in comparison to other cell types in the body (18) tend to confirm that VICs are recurrently activated to repair mechanically induced tissue damages and ensure valve durability.
Endothelial cells, in direct contact with the blood flow, ensure non-thrombogenic properties while regulating immune and inflammatory responses (19,20). VECs are constantly exposed to shear stress of different natures, which have been related to different cellular behaviors (21). Indeed, the venticular side is exposed to a unidirectional shear stress of an averaged 20 dynes/cm2 while the fibrosa surface is subjected to an oscillatory stress of lower magnitude (22,23).
The homeostasis of the leaflets seems to be regulated by a crosstalk between interstitial and endothelial valve cells. For instance, neuronal stimulation of aortic cusps has been shown to be effective only if the endothelial layer was intact (10). Furthermore, studies in vitro have shown the VECs had an effect on VICs proliferation, protein synthesis and decrease in myofibroblast activation (24). Accordingly, loss of VECs by mechanical damage or pathology could result in myofibroblastic activation and ECM remodeling.
During the cardiac cycle, the aortic root and leaflets act in coordination to efficiently close and open the aortic valve and convey a unidirectional blood flow into the aorta (25). During diastole, the aortic cusps are closed by apposition and ensure the filling of the left ventricle. During filling, the fibrous annulus is dilated, which results in a supplementary strain on the leaflets and prevents a prolapse. At the end of diastole, the isometric contraction of the ventricle further increases the annulus expansion that pulls further the leaflets from their commissure. The valve then quickly opens during systole into a circular pattern while annulus retracts and the root resumes a cylindrical shape to ensure an optimal ejection of blood. When the leaflets open, eddy currents form in the sinuses of Valsalva (26) that prevent the leaflets striking the aortic wall and hold them away from the aortic wall in a position promoting valve closure and helping transfer blood in the coronary arteries (27). The closing of the valve is slower than its opening, possibly due to the vortexes that create a pressure gradient across the leaflets and promote a smooth closure (5).
From a mechanical point of view, stresses on the leaflets are most important during diastole and early systole. Bidirectional measurements in vitro of the cusps mechanical properties have shown that they were strongly anisotropic (circumferential strains being much smaller than radial strains) and compliant until a critical strain where they become rigid (28) (Fig. 2-A). This behavior can be related to the structure of the different cusps layers (6,29-31). At lower strain during systole, non-elastic collagen fibers of the fibrosa are crimpled while elastic fibers of the ventricularis layer provide the leaflets elastic response. During diastole, however, collagen fibers unfold and stiffly resist the pressure load (Fig. 2).
The complex coordination of the aortic valve elements and microstructures during the cardiac cycle is a strong remainder of the necessity to consider multiple fields of research in view of its replacement and proper recapitulation of its function.
C. Aortic Valve Disease and Clinical Needs for Replacement
Although aortic valve diseases (AVD) are responsible for a limited fraction of the deaths caused by cardiovascular diseases (about 2%), their prevalence is important, notably during aging, as 29 % of people over 65 have aortic valve sclerosis (33,34). The presence of a sclerotic aortic valve is not benign and is associated with an increase of the risk of death from other cardiovascular diseases by 50 % (35). In Western countries, valve diseases are mostly degenerative and related to aging. Conversely, in developing countries, valve diseases mostly affect children and young adults (36). In the latter, even if surgical interventions can save lives, pathological remodeling usually occurs, leading to premature valve failure.
The most common effects of valve pathologies are stenosis, insufficiency (regurgitation or incompetence), or both (8). Aortic stenosis, the most common pathology, is characterized by a thickening and stiffening of the valve, leading to a narrowing of the valve and an increased workload of the heart. Aortic stenosis is usually the result of calcification of the cusp tissue. The mechanisms of calcifications are not fully understood, but could be linked with inflammation, lipids accumulation, trans-differentiation of VICs towards osteoblastic phenotype and the presence of specific spherical calcium phosphate particles (37,38-40). Aortic insufficiency is usually the result of a dilatation of the ascending aorta, related to hypertension, aging, or can be caused by rheumatic fever (41).
AVD diagnostic is usually a two-step process. A first hint of valve dysfunction is obtained during physical examination by the presence of a high-pitched murmur (42,43). A confirmation is then established by ultrasound echocardiography, which measures hemodynamic values (peak ejection velocity, valve opening area and transvalvular pressure gradient) and determines the severity of the dysfunction (44,45).
Once diagnostic and severity have been confirmed, the two only therapies available to aortic valve disease are valvuloplasty (repair) and valve replacement. However, valvuloplasty usually relies on the replacement of the aortic root or cusps (46). Without replacement, the development of aortic stenosis symptoms shorten life expectancy to about 3 years (40).
D. Approaches to Reproduce and Replace Aortic Valves
Historical paradigm: reproduce function
Historically, the objective sought through valve replacement was to reproduce and restore the mechanical function of healthy valves. Logically, the first attempts to reach this goal have concentrated on the use of prosthetic devices using the heart cycle and blood flow pattern to act as mechanical valves. As early as 1951, Hufnagel reported his precursor design, consisting in a polyethylene ball entrapped in a methyl methacrylate tube (47), which he implanted in 1952 in a patient thoracic aorta to prevent aortic backflow and treat aortic insufficiency. With the advent of heart and lung machine and open heart surgery in the 50’s, more designs were successfully evaluated as replacement of the aortic valve such as caged-ball (Starr-Edwards valve, 1960 (48), tilting-disc (Björk-Shiley Delrin, 1969 (49) and pivoting bileaflet devices (St. Jude Medical valve, 1976 (50) (Fig. 3).
Nowadays, mechanical prosthetic valves are frequently used in clinic. However, although they are effective, structurally sound and durable, they suffer from diverse shortcomings (51). Their components are recognized by the body as foreign material while their geometry in the blood flow results in non-physiological shear stresses that induce platelet lysis and give rise to thromboembolic complications. As a result, and despite more than 50 years of research, patients require life-long anticoagulation therapy, which are limiting and increase risk of serious bleeding events (52-54)
In parallel to the development of mechanical valves, tissue valves homografts from cadaveric aortic valve were investigated as implants by Dr. Ross in 1962 (55,56) followed by xenografts from porcine (Hancock glutaraldehyde-fixed valve, 1969 (57) and bovine (Carpentier Edwards pericardial valve, 1976 (58) origin (Fig. 3). These tissue valves advantageously solved the hemocompatibility issues of mechanical valves by providing more physiological hemodynamics and cancelled the need for anticoagulant therapies. However, being composed of fixed tissues, without any living cells, they have a strong risk of structural deterioration. In addition, the lack of living cells prevents valve calcium homeostasis and favours calcium nucleation, nodule formation and calcification (59) that ultimately lead to more re-operations (60). In fact, the tissue valves are now increasingly considered by clinicians to “replace one disease with another” (41, 61).
Another critical drawback of either mechanical or tissue valve is the passive role that they exert in restoring the valve function. Since they are not living tissue, they lack the ability to grow, repair or remodel. While this passivity is acceptable for sedentary or older patients, it becomes problematic for young and active adults or for infants and children. To compensate for the growth of children conduits, the only solution for valve replacement is to implant larger prosthesis than ideal and/or re-operate frequently to resize the conduit (63). Furthermore, as xenogenic tissue valves in children are notorious for rapid calcification and resulting early reoperation, mechanical valves are used instead together with anticoagulation therapy and the associated risks for children (Fig. 4).
Paradigm shift: reproduce function with living tissues
The drawbacks and difficulties associated with mechanical or decellularized tissue prosthesis, especially in paediatric patients, have motivated the development of other strategies aiming not only to restore proper aortic valve function, but as well its living regulatory and adaptive properties.
Surgically, a technique has been proposed as early as 1967 by Dr. Ross, consisting in replacing the aortic valve with the patient’s own pulmonary valve, which can in turn be replaced by a decellularized tissue valve (homograft) (64). This pulmonary autograft, commonly termed “Ross procedure”, circumvent the limitations of prosthetics in the aortic position by advantageously transposing them to the less demanding pulmonary position (Fig. 5-A). This operation has been performed for decades and although clinical studies indicate its superiority to homografts in adults (1) (Fig. 5-B), its use still remains controversial (65). For infants and children, the advantages of the Ross procedure as cell-containing living tissue being able to grow with the patient have made it the operation of choice.
Nevertheless, clinical acceptance of the Ross procedure is still not consensual due to the technical difficulty of the operation and a risk of aortic root dilatation instead of growth leading to valve insufficiency(66-68), especially in young patients with bicuspid aortic valve (69).
Aortic valve tissue engineering
To overcome the limitations of prosthetic and surgical approaches, a multidisciplinary strategy has emerged in the past two decades that aims to combine the beneficial aspects of living transplants (biocompatibility, growth, repair and remodeling) with a proper and controlled reproduction of the aortic valve structure and function. Heart valve tissue engineering follows the general paradigm of tissue engineering that was first defined in 1988: “the application of principles and methods of engineering and life sciences towards the fundamental understanding of structure-function relationships in normal and pathological mammalian tissues and the development of biological substitutes that restore, maintain or improve tissue function” (70). In theory, aortic valve tissue engineering could be particularly beneficial for young patients who require valve growth and cannot tolerate the side effects of non-living prosthesis (71). If this population is mostly confined to birth malformations in the western world (about 2 % of life birth), it could amount millions in developing countries due to higher incidence of rheumatic fever (52).
Although different methodologies have been applied and evaluated for aortic valve tissue engineering, an accepted paradigm comprises a natural or synthetic scaffolds that is pre-seeded with cells, followed by an in vitro step of tissue formation conducted in a bioreactor, and subsequent in vivo implantation allowing for tissue growth and remodeling (72). This global approach allows for extensive variations of parameters such as scaffolds properties, architecture and degradation, cell type and in vitro stimulations. The resulting key physiological processes evaluated are cell proliferation and migration, ECM production and organization, tissue remodeling and recruitment of host cells, either inflammatory or precursor, after implantation. Interestingly, the recruitment of host cells by unseeded scaffolds, designed to attract endothelial and mesenchymal circulating precursor cells in vivo and orchestrate tissue formation is as well considered as tissue engineering (Fig. 6).
Although this chapter will introduce the different aspects of aortic valve tissue engineering in the following paragraphs, the reader will find more details in extensive reviews on this subject (73-76).
Cell supportive matrices
The first important component in aortic valve tissue engineering is the development of a cell supportive scaffold which shape mimics either a valve leaflet (77-79) or a full conduit comprising leaflets and aortic root (80). Both approaches have advantages and limitations. A semi-lunar valve leaflet shape is certainly of easier conception and realisation but forecasting exact dimension for a given patient is difficult and might result in improper coaptation and closure after suturing to the existing aortic root. Conversely, a full conduit simplifies hemodynamic issues, but is technically difficult considering that the aortic root and leaflets are of different organizations and mechanical properties (81). Regardless of the approach considered, the tissue engineered valve must resume a perfect hemodynamic function immediately upon implantation and keep relatively constant mechanical properties over time despite tissue remodeling and eventual scaffold resorption.
The role of the scaffold is therefore complex. It must possess a number of qualities that will allow and guide the formation of a new tissue such as biocompatibility, elevated and interconnected porosity (to allow cell growth, nutrient diffusion and waste removal), cellular attachment, migration, proliferation and extracellular matrix formation; while providing an initial mechanical support after implantation that must gradually be compensated by the neo-tissue as the scaffold is degraded and resorbed.
Intuitively, the best scaffold to comply with these requirements would be the native aortic valve extracellular matrix itself. In theory, using detergents washings, cellular content can be removed from valve tissue to obtain an intact extracellular matrix structure of low immunogenicity. However, despite experimentation of various decellularization protocols using detergents, enzymes or nucleases (82) (83), some features of the extracellular matrix such as collagen pattern, distribution and glycosaminoglycan content can be modified (84) and lead to altered biomechanical and hydrodynamic properties (85,86). The balance between extent of decellularisation and conservation of mechanical properties is further complicated as remaining cellular debris are known to elicit a strong immune response and promote calcification (87). Despite these difficulties, decellularized valves in pulmonary (88,89) or aortic (90) positions have reached clinical evaluations in adults and children. Although recellularization of the grafts was marginal, the outcome for adults was positive in regards of valve function. Conversely, it resulted in the tragic death of 3 out of 4 children who were treated in the study, due to a swift and strong inflammatory reaction (87). As a result, the prospective use of decellularized valves has been set back and current strategies focus on improving recellularization in vivo to enhance biointegration (91-93). Without minimizing the potential of such an approach, the use of decellularized matrices will nonetheless always rely on a strong supply of human valves of good quality, which is lacking (Table 1).
Table 1. Comparison of cell supportive matrices used in aortic tissue engineering.
In view of overcoming the limited access to native valve extracellular matrix while controlling a wide range of properties, synthetic or natural biodegradable polymers have been as well investigated. Synthetic polymers have the advantages of limitless supply, well controlled and tailored intrinsic properties and formulation ease towards highly porous and interconnected structures through various manufacturing processes such as salt leaching (94), knitting (95) stereolithography (96), electrospinning (97) or jet-spraying (98, 99). On the other hand, they must be biocompatible, resorbable without major local side effects, allow cell attachment, migration and proliferation, differentiation, ECM formation and remodeling and exhibit mechanical and structural properties comparable to aortic valves, which narrows the list of potential candidates (100). As a result, various polyesters (PGA, PLLA, PLGA, P4HB, PHA and PCL) have been used, without consensus on a unique polymer presenting all necessary requirements (83,101-107). Indeed, each polymer presents different interactions with cells through intrinsic surface properties. Additionally, sponge-like or fibrillar and nanofibrillar formulations of varying porosities influence cells migration, deep colonization of the structures and synthesis of a homogeneous tissue (99, 108). As degradation occurs through hydrolysis at various rates (from weeks for PGA to years for PCL), acidic molecules are released that can influence cell fate over time and induce local tissue inflammation. Concurringly with degradation, polymer matrices mechanical strength and rigidity are decreased while valve function must remain constant over time. This central requirement implies to find the perfect balance between tissue formation, remodeling and polymer resorption. So far, the most advanced results in large animals in vivo have been obtained with fast degrading polymers (PGA and P4HB) cultured in vitro during 2 weeks with myofibroblasts and endothelial cells or associated with bone marrow cells prior implantation as pulmonary valve replacement (95,109).
As an alternative to synthetic molecules, natural polymers issued from ECM have the advantages of being inherently biocompatible and degradable through enzymes, which allows for a cell-driven remodeling. The most commonly used proteins are type 1 collagen and fibrin, which form a fibrillar hydrogel spontaneously at neutral pH or through thrombin polymerisation, respectively. The direct encapsulation of cells within the hydrogels is another advantage over synthetic polymers, which rely on cellular invasion to achieve cellularity of the constructs. Furthermore, all mesenchymal cells seem to flourish within such structures (110-112) that can easily be shaped as aortic valves by injection moulding (113,114). However, these hydrogels lack the mechanical strength required for valve replacement. As a result, dynamic conditioning has been successfully evaluated to induce sufficient strength prior implantation (115). However, severe gel contraction induced by the cells resulted in a quasi-disappearance of the cusps after three months (116). In the future, natural proteins might find a suitable role in valve replacement in combination with synthetic polymers, to improve cellularity while maintaining mechanical strength (117, 118). Nevertheless, for the body, a fibrin matrix still represents a wound granulation tissue that needs to be remodelled in fine in a scar tissue dissimilar to the native one; therefore, the biological rationale of using fibrin gels as substrate for functional tissue formation in vivo remains to be demonstrated. A promising alternative could be to use fibrin matrices as a template for cells to form a controlled collagen matrix that can be subsequently implanted after decellularisation (119).
Although the vast majority of valve tissue engineering approaches associate cells with supportive structures in vitro and stimulates them through chemical and/or mechanical conditioning prior in vivo implantation, cell-free strategies are as well investigated to design smart materials that could attract and retain host circulating progenitor cells and form a new valve in situ. As synthetic polymers lack the appropriate signalling molecules naturally present in the ECM for cell recruitment, the challenge here is to integrate this bioactivity into the materials. Various bioactive cues and moieties can be envisioned to capture, recruit and guide differentiation of progenitor cells, such as antibodies (CD34 and vascular endothelial-cadherin (120,121), aptamers (122), and growth factors (123) for endothelial progenitor cells. Similarly, coatings with mimetic peptides or isolated ECM proteins are known to promote cellular adhesion (124), but could as well direct biological events such as apoptosis (125) or cell differentiation (126). It should be noted that in the context of in situ valve engineering, the polymeric matrices will require appropriate mechanical properties and slower degradation rates as compared to in vitro cellularized constructs, to allow for cell recruitment, differentiation and tissue formation.
As promising as these precursor strategies might be to bypass the in vitro stage of cellularisation and conditioning, they are still in their infancy and the in vitro association of cells with supportive structures remains the most employed approach.
Cell sources and Challenges
Optimally, the cells used for association with supportive structures and development of a new valve tissue should be non-immunogenic, proliferative, easy to harvest and either keep their specialized function or be able to gain such a specialization through differentiation. The most logical cells to harvest are autologous cells from the patient to be treated, to remove any risk of immune rejection. The use of autologous aortic valve cells being impossible, other cell sources have been evaluated, such as fibroblasts, myofibroblasts and endothelial cells isolated from blood vessels or subdermis (77,80,95,127-129). However, autologous material from patient requiring heart valve replacement is usually too sparse or with diminished proliferative capacity, especially for elderly patients. As a result, the use of autologous mesenchymal stem cells (MSCs) has been actively investigated, either from bone marrow or adipose tissue origin (99,130-133). Additionally, specific cell sources for pediatric patients, such as Wharton jelly, umbilical cord blood, peripheral blood, (129,134,135) and prenatal cells issued from chorionic villi (136) and amniotic fluid (137,138) have shown proof of concept results.
Important problematic linked to ce llularization of constructs in vitro are cell retention and endothelial function. Retention of cells after implantation has not been extensively investigated, although a study indicated that canine bone marrow cells associated with decellularized porcine pulmonary valve were still present after 3 weeks of implantation and could have therefore contributed to the regeneration process (139). Endothelial function of the engineered valve is critical to provide anti-thrombotic properties and ensure survival of recipients. Isolation of endothelial cells from blood vessels or differentiation towards endothelial phenotype of MSCs by chemical and cytokine cocktails (140) is possible while re-endothelialisation of supportive structures (141) seem to improve endothelium coverage and prevention of thrombus formation as was shown within 3 months after implantation in pulmonary position in sheep (142). In humans, reconstruction of right ventricular outflow tract of adults undergoing the Ross procedure performed with re-endothelialized pulmonary allografts showed no calcification and thrombogenesis after 10 years (143,144). More importantly, re-endothelialisation with endothelial progenitor cells from peripheric blood of decellularized pulmonary valve allografts implanted in 2 children seemed to have prevented valve degeneration over 3.5 years of implantation (135).
Another important aspect of the in vitro expansion and association of cells with scaffolds is the conservation or acquisition of a phenotype similar to native valve cells in vivo. While cultured MSCs display a spindle morphology characteristic of fibroblasts and myofibroblasts, express markers shared by VICs (72) and have the ability after implantation to form endothelial and interstitial-like layers (145-148), concerns remain regarding the stability of the acquired phenotype and potential unwanted fibrotic overgrowth (149). Similarly, while myofibroblastic and endothelial cells can be maintained in culture, lack of phenotypic control once implanted can result in formation of fibrotic-like tissues causing retraction and regurgitation of the engineered valves (128,129,150).
With regard to cell origin, it should be noted that animal cells are often used in prevision of proof-of-concept animal experiments. However, such cells do not always behave similarly to human ones. For instance, ovine vascular-derived cells show different sizes, proliferation rates and contractile behaviour as compared to human cells of similar source, which result in a stronger scaffold contraction (73). As a result, specific protocols have to be developed for each cell type and origin, and should always be validated again with human cells.
Regardless of cell sources, the favoured way to induce a strong cell commitment toward a specific phenotype and monitor undesired orientations consist in mimicking the natural valve mechanical and flow environment and pre-conditioning the tissue engineered constructs in vitro prior implantation.
Bioreactors and conditioning
The design and use of bioreactors for conditioning heart valve engineered tissue find multiple rationales in biological mechanisms. Indeed, during heart valve adaptation or repair after injury, VICs are known to change phenotype and become activated in myofibroblasts (151). Such activation has been shown to be regulated by environmental stimuli such as dynamic loading of the valve (14,152), profibrotic cytokines (transforming growth factor-beta 1) (153) and elevated matrix stiffness (13). Furthermore, mechanical conditioning can influence extracellular production as was demonstrated with native porcine aortic valve leaflets exposed to hypertensive cyclic pressure (154) or isolated circumferential cyclic stretch (155). The former study indicated an increase of collagen and GAGs synthesis related to magnitude and frequency while the later showed an increase of collagen and a decrease of GAGs after stimulation.
Bioreactors allowing a variety of conditioning strategies in terms of cyclic flow and pressure changes (95,156-159), strain (132,160,161) and shear stress (162) have been developed (Fig. 7) and many studies have highlighted the benefit of mechanical and flow conditioning. For instance, a moderate pulsatile flow reached in small increments over endothelial cells seeded on decellularized pulmonary valves revealed of better quality for cellular proliferation than a rapid increase to physiological flow (141). Straining of tissue engineered valves (163) and collagen cross-linked constructs seeded with smooth muscle cells (164) resulted in more pronounced and organized tissue formation with superior mechanical properties as compared to unstrained controls. Similarly, collagen synthesis by porcine VICs was shown to be dependent on degree and duration of stretching (165) while stretching of adipose and bone-derived MSCs cultures on substrates stretched at 14 % for 3 days showed an increase of collagen synthesis (132). Dynamic straining (4% for up to 10 days) of myofibroblasts-seeded polymer constructs indicated a decrease of collagen production but an increase of collagen cross-links, GAGs production and remodeling markers (166). Prolonged dynamic straining (4% for 2 weeks) were shown to induce the production of collagen with a higher degree of crosslinks while inducing a specific cell orientation parallel to strain direction and a layered collagen organization (167). Interestingly, excessive dynamic strain magnitudes (8%) resulted in a decrease of the mechanical properties of the constructs while intermittent dynamic straining seemed to improve tissue formation and reduce the risk of deterioration (168).
Although the optimal conditioning protocol will depend on different parameters linked to the cell phenotype, the type of scaffold used and the magnitude and type of mechanical cues, the length of stimulation that has been generally applied in various studies is 3 to 4 weeks (80,136,145,160).
Validation of tissue engineered valves?
Various animal experiments performed mostly on sheep have clearly demonstrated the proof of concept of tissue engineered valves (Fig. 8) (72,77-79,145). However, although most of these experiments have been conducted in the pulmonary position, which is biomechanically less demanding and therefore more attainable than the aortic, translation to human has not yet been achieved. The few experiments performed in human with decellularized allo- or xenograft (170) have indeed showed either promising results (135,143,144) or catastrophic outcomes (87).
This discrepancy underlines the need for integrated translational tools and platforms. Many unanswered questions remain in regards of biological mechanisms, interactions of cells with supportive materials, animal models and discrepancy between animal and human cells behaviour, before successful human implantation can be achieved. Furthermore, criteria and quality markers need to be established to determine when a tissue engineered heart valve can be implanted in a patient, with a satisfactory prognosis for long-term survival. Similarly, predictive models issued from animal experiments and from in silico models of tissue growth, remodeling and clinical epidemiology remain to be developed.
In this context, it is vital to find and establish suitable criteria to assess the efficacy of engineered aortic valves and translate in vitro and animal experiments to human prognostic, as will be presented in the following part of this chapter.
E. Translational Platform: Establish Criteria and Quality Markers
The need of a translational platform for aortic valve replacement is strongly linked to the valve tissue engineering approach. While mechanical and biological cardiac valve prosthesis design, production and risk management can be evaluated through International Standard Organization (ISO) guidelines (ISO 5840:2005, “Cardiovascular implants — Cardiac valve prostheses”), tissue engineered valves have to comply with more requirements, being living tissues. The need exists for engineered valve to evaluate their functionality prior implantation through quality criteria indicative of a positive clinical outcome.
So far, the path to evaluate aortic tissue engineered valve has been to primarily focus on pulmonary valve engineering and replacement in conjunction with Ross procedure, as pulmonary valves share the same structure as aortic valves but are less hemodynamically demanding (5). The constructs are first evaluated in vitro for biological and biomechanical properties, often in bioreactors allowing hemodynamic stimulations (116,146,171). They are then applied in small animal models, such as rabbit or rats to provide information on biocompatibility, or in large animals such as sheep for functionality assessment. The next evaluation is in adult humans, as replacement of the pulmonary valve during the Ross procedure, prior to the final evaluation in children that will require valve growth.
Before reaching evaluation in infants and child, many steps must therefore be validated. But the question remains as to the validation criteria and markers allowing to indicate a satisfactory prognosis for heart valve replacement in patients. From the previous studies outlined in this chapter, the function of aortic valve is correlated to its structure; therefore, validation criteria of engineered valves must be orientated towards physiological structure and function, and assessed always before and after animal implantation. Accordingly, a set of major criteria to consider for the clinical relevance of engineered construct can be defined as follows:
- shape and structure
- biocompatibility, biodegradation
- cellular biomarkers, tissue formation and quality
- remodeling towards native structures, structural integrity and endothelial function
- hemodynamic performances and biomechanical properties
- functional assessment and durability
These criteria and markers must be established both in vitro and in vivo, in long term follow-up preclinical studies. The provided data can then be used to derive predictive models in humans in regards of:
- hemodynamic and functionality
- in vivo remodeling, repair and growth
- lack of pathological behaviour of the implanted valve (calcification, stenosis)
Finally, the collected data and model predictions must be compared to the function observed in native aortic valves and with mechanical or biological prosthetic valves (172). A visual summary of such a translational approach is presented in Fig. 9.
To provide an overview of this translational approach, the last part of this chapter will describe in more details the different quality criteria/markers and the corresponding invasive and non-invasive tools that can be applied for the development of a translational platform for aortic valves replacement.
Evaluation Criteria in vitro: supportive matrices and cells
With regards to parameters and criteria of the cell supportive structures that can be evaluated at early stage of development in vitro, shape and structural organisation are the most straight forward. While shape and exact dimension of the constructs can be easily measured directly or through macroscopy, microscopic structure and organization are commonly characterized with scanning electron microscopy (SEM) followed by image analysis (98). Such analyses are useful in developing matrices that aim to mimic the native anisotropic organization of aortic valves ECM (99,173).
Porosity, pores sizes and pores interconnection of the matrices directly relate to cellular colonization and tissue in-growth and as such can be used as predictive tools. A pore size value of 10 micrometres has indeed been proposed as threshold permitting cellular infiltration (174,175). While gravimetric measures allow to calculate porosity, SEM images analysis (176) or mercury intrusion porosimetry (99,177) can be used to evaluate pores sizes and interconnection. However, results obtained with these methods should be analysed with caution as they do not provide fully accurate results. Image analysis is rendered imprecise by the depth of field of scanning electron microscopes that prevents to isolate pores in a single plane, while the pressure applied on soft matrices to force mercury intrusion will induce structural deformations. As such, these methods remain useful for crossed comparison of various supportive structures and should always be confirmed by cellular infiltration assays.
Biocompatibility of manufacturing materials is a mandatory criterion that is usually evaluated both in vitro and in vivo. Direct cell toxicity of the materials and leached out or degradation products can be determined using in vitro cell viability assays based on the quantification of cellular metabolic activity (MTT, MTS assays or alamar blue) (176,178). Direct visualization of cell mortality is as well possible through life/dead assays that clearly differentiate metabolically active cells from dead cells with ruptured cytoplasm under fluorescence microscopy (92,179).
In vivo studies provide further information on the foreign body reaction of the native immune system. Small animal models, such as rabbits or rats, are implanted with the materials to minimize the costs and ethical issues associated with the use of large animal models. Subcutaneous implantations during various times allow to observe capsule formation and identify infiltrating cells such as inflammatory cells, macrophages or fibroblasts by mean of histology and immunohistochemistry (91). Systemic toxicity is indicated by weight changes of the animals after implantation while remote organ toxicity can be evaluated by collecting liver, heart, and kidney samples from each animal and controlling their histological structures (92). Aside from subcutaneous models, alternatives have been proposed to assess biocompatibility in environments closer to the heart valve such as implantation of materials in the blood flow of rats abdominal aorta (180).
Similarly to biocompatibility, materials biodegradation is an important parameter to take into account during development of cell supportive structures, in prevision of its replacement by neo-tissue after implantation.
For synthetic polymers, in vitro assays can be performed by incubating the structures in buffers either mimicking physiologic conditions or allowing an accelerated hydrolytic degradation (178). Mass loss, molecular weight decrease, measure of degradation products, crystallinity, and evolution of macro- and microscopic morphology by SEM over time provide valuable information of polymer degradation rates. Variations of mechanical properties over time indicate degradation as well and are particularly useful when combined with in vitro cell culture to infer that ECM synthesis, and hence tissue formation, is sufficient to maintain the constructs mechanical properties. However, degradation rates obtained in vitro cannot be readily translated to the in vivo situation and should always be confirmed.
For matrices made of natural materials, biodegradability is mostly assessed directly in vivo in subcutaneous small animal models (91,181). Histological evaluations of cellular infiltration and type of cells associated with quantification of the remaining implant area over time are then reliable indicators of biodegradation.
It is worth noting that the characterization of polymers biocompatibility and biodegradation are especially valid for novel compounds or compositions. Over the past decades many biodegradable polyesters have already been validated for human use and approved by the FDA (100), such as poly(glycolide), poly(lactic acid), poly(lactic-co-glycolic acid) and poly(ε-caprolactone), which facilitates their use in aortic valve engineering.
Although the most common strategy for aortic tissue engineering relies on in vitro cell culture and bioreactor conditioning to improve mechanical properties prior implantation, the gross properties of the supportive structure should optimally mimic some of the valve features such as anisotropy. Moreover, in the specific in situ approach where the implants are colonized by the patient own cells after implantation, the supportive structures should be sufficient to withstand the ranges of stresses and pressures applied on a native aortic valve immediately after implantation and remain until complete tissue formation and remodeling.
Gross mechanical properties of supportive structures are most commonly evaluated by tensile assays. The force necessary to produce a given deformation rate of the sample is recorded and stress/strain curves calculated from the measured force and the sample cross-sectional dimension. From this data, the tissue strength, stiffness, relaxation and elongation before rupture can be calculated and compared to native tissues measured in a similar way (99,182). The material stretching can be applied in one direction (uniaxial), or simultaneously in perpendicular directions (biaxial). Anisotropic properties of the tested structures can be calculated using both approaches, but biaxial tests are closer to physiological situations and as such should be preferred (183, 184). For materials derived from natural ECM, additional information concerning fibre alignment and deformation during stretching can be obtained by combining tensile tests with the quantification of polarized light angles (184,185).
As a guideline, computational modelling can as well be employed to determine specific range of mechanical properties, anisotropy and shape that would allow proper radial stretch and coaptation of the supportive structure in either pulmonary or aortic conditions for instance (186).
Aside from measure of gross mechanical properties, local stiffness can as well be measured by atomic force microscopy. This is particularly relevant to compare different structures to the local mechanical properties of the different valve layers and hence mimic the multifactorial properties of the ECM (176).
The in vitro association and culture of cells with supportive matrices is crucial to identify cellular interactions with the supportive structure, allow the formation of ECM, and provide endothelial functions and sufficient hemodynamic/biomechanical performances.
The ability of cells to recognize the provided structures as suitable for their proliferation is usually assessed by classical biochemical techniques. MTT/MTS assays (132) or DNA quantification (99) are commonly used to measure proliferation.
Cellular penetration and colonization of the constructs is an important criterion to assess as it will reflect the possibility of forming a homogenous tissue prior and after implantation. Histological preparations and sectioning of the constructs at different times followed by image analysis are effective to determine the rate of cellular invasion. Cell nuclei are stained using either fluorescent (DAPI) or colorimetric (Movat pentachrome) dyes and cell distributions within the matrices can be deduced (99,132,187). To further evaluate that cellular colonization is homogeneous throughout all available volume, three dimensional reconstructions can be performed by assembly of multiple contiguous cross sections (Fig. 10-A) (188).
Cellular morphology and organisation can be guided by mechanical stimulations (189) and by the type and structural features of supportive matrices, such as anisotropy (99,173). Scanning electron microscopy analyses provide convenient information of cells located at the surface of the matrices, but not of cells within (7). To visualize these cells, confocal microscopy can be used after fluorescent labelling of the cytoskeleton with phalloidin for instance (Fig. 10-B). However, the attainable depth depends strongly on the scaffold material and remains confined to hundreds of micrometres at best.
Evolution of the cell phenotype during culture with the supportive structures should be evaluated over time, to assess uncontrolled differentiation or transdifferentiation when stem or mesenchymal cells are used. The expression of stem or differentiated cell surface markers can be followed over time by immunohistochemistry on cross sections (99), or by flow cytometry (fluorescence-activated cell sorting, FACS) when cells can easily be extracted from the supportive matrices (179). The acquisition of a myofibroblast phenotype is of particular importance as VICs are naturally activated into myofibroblasts during valve growth or remodeling (13,190). Myofibroblast activation of fibroblastic cells in general and VICs in particular is known to be mediated through physical (substrate rigidity) cues in their microenvironment (13,191,192,193) and as well through mechanical conditioning (155). However, excessive and uncontrolled activation of various cell types can lead to excessive deposition of unorganized ECM of fibrotic properties (194,195). Immunohistochemical stainings of sections for alpha-smooth muscle actin (α-SMA), desmin and vimentin indicate myofibroblast-like cells (99,106,196), while western blotting of α-SMA can be used in a semi-quantitative fashion (155).
ECM synthesis and remodeling
The composition and organization of the cell-produced ECM is fundamental for the structural integrity and biomechanical properties of aortic tissue engineered valves after implantation and for its long-term performance, similarly to native valves. It is therefore of high interest to quantitatively and qualitatively assess ECM components deposited by the cells during culture and compare them to the native tissue.
Cross-sections of the constructs at different culture times and histological and immunohistological methods are commonly used to evaluate the deposited ECM within the supportive structures. Classical colorimetric stains such as hematoxilin and eosin are useful for the general ECM distribution and morphology for instance (80, 106, 155) while ECM composition can be discriminated using specific antibodies directed against collagen, elastin and chondroitin sulfates (99, 146). Particular staining procedures such as picrosirius red (155, 196) and Masson’s trichrome stains (80) or alcian blue (182) are also useful to stain collagen or glycosaminoglycan content, respectively. Movat pentachrome allows one to visualize proteoglycans (in green), collagen (in yellow), and elastin (in black) on the same section (14, 145, 196). ECM-related gene expression can also be detected using in situ hybridization (197). With regard to matrix remodeling, information is provided by immunohistological techniques targeting matrix metalloproteinases (MMPs), cysteine endoproteases (cathepsins) and tissue inhibitors of MMPs (TIMPs) (196-198).
All these characterizations require the destruction of the analysed sample by embedding, sectioning and staining. In prevision of quality control of the constructs prior implantation, non-destructive methods for ECM formation analysis have been developed. For instance, multiphoton excitation microscopy allows to visualize collagen structures and elastin fibers with great details at high resolution (Fig. 11) (199, 200). This technique based on the absorption of multiple photons carrying approximately less energy than necessary to excite a molecule results in the subsequent emission of a fluorescence photon at a higher energy. The infrared excitation used results in deeper penetration in the sample than confocal microscopy while being of low cytotoxicity (201). Collagen-specific fluorescent probes have as well been developed to evidence collagen synthesis and organization under confocal microscopy (202, 203).
It is worth noting that the use of polymer matrices often limits histological, immunohistological, confocal and multi-photon techniques due to polymer dissolution in histological solvents, antibodies retention, light diffraction and auto-fluorescence that will interfere with samples preparation and examination.
Quantitative assessment of ECM formation can be obtained by biochemical assays specific for collagen, GAGs and elastin followed by normalization per cell using DNA content (80, 84, 106, 134, 146). Quantification of different collagens, lysyl oxydase and tropoelastin mRNA by real-time polymerase chain reaction provides as well information on ECM proteins gene expressions levels during in vitro culture (132). Alternatively, collagen formation can be precisely quantified by radio-isotopic detection of (3H)-proline incorporation (132, 165). Aside from the amount of ECM proteins, collagen and elastin crosslinks can as well be monitored by respectively measuring pyridinoline and desmosine in protein hydrolysates using liquid chromatography and mass spectrometry (132). The amount of crosslinks of collagen have been demonstrated to be directly correlated with the resulting tissue strength and stiffness (7) and as such is an important criteria to consider.
The formation of an ECM comparable to the native aortic valve is crucial for the biomechanical performances of the constructs prior and after implantation. Furthermore, not only the repartition and amount of ECM proteins are of importance, but as well their maturation and architecture as a whole. Although tissue engineered constructs in vitro can present a tissue formation of composition similar to native heart valve leaflets (106, 136), different aspects are still not fully satisfactory. For instance, anisotropic organisation of collagen fibres is not yet achieved. It was once thought that once implanted in the aortic valve hemodynamic environment, collagen will be able to mature and gain increased alignment (13). Elastin content is also problematic and not sufficiently synthesized in engineered valves (73, 80), even though its critical role in providing a proper biomechanical function, ensuring optimal valve closure and preventing valve replacement failure is acknowledged (204, 205).
Endothelialisation of engineered aortic valves is a crucial parameter to assess prior implantation, in view of decreasing thrombogenicity after implantation. Immunohistochemical detection of CD31, vascular endothelial grow factor receptor FLK-1 and von Willebrand factor (vWF) is useful to determine the presence of an endothelial layer at the surface of the engineered constructs (Fig. 12-A) (135, 141, 142). The detection of endothelial nitric oxide synthase (eNOS) production by immunohistochemistry provides as well an indication of the endothelial layer functionality (134, 141, 142).
In conjunction with immunohistochemical techniques, SEM microscopy allows to determine the homogeneity of the endothelial coverage and its density (Fig. 12-B) (134).
Evaluation Criteria in vivo
Considering the critical and vital mechanical and hemodynamic characteristics of aortic heat valves, it is mandatory to validate that the implanted constructs are able to provide functional closing and opening motions and cope with the aortic valve environment, at least immediately after implantation and prior to a complete tissue formation and remodeling.
The same tensile tests as for the characterization of acellular supportive structures are commonly used to determine tissue strength, extensibility, stiffness and anisotropic properties (80, 136, 182). In addition, the evolution of the constructs mechanical properties over culture time is a good indicator of ECM production and quality (7, 99). In this regard, measure of anisotropic properties of the tissue engineered aortic valves should preferentially be performed through biaxial testing to ensure that mechanical behaviour in radial and circumferential direction are not inconsistent with those of native valves (80, 183).
The evolution of the constructs mechanical properties combined with the characterization of ECM production at different culture times could eventually lead to models of tissue formation and biomechanical properties in function of scaffold and cell type (81, 206). Although those approaches are still in their infancy, the prospects are high to allow the preparation of constructs perfectly matching the aortic valve.
To validate hemodynamic performances, bioreactors have been developed that simulate the pressure cycles applied on the aortic valve (159) or replicate the complete physiological heart cycle and pulsed aortic flow (80, 134, 160). Aside from the validation of closing and opening motions, other parameters such as transvalvular flow velocity, pressure gradients and effective orifice area can be assessed (80, 207, 208). The constructs can be tested over many weeks in such systems to provide information of durability, aging of the constructs and compliance with heart pressure cycles (209). Long test periods always increase the risk of bacterial and fungi contaminations of the samples.
F. Preclinical Testing
Functionality assessment of living engineered valves needs to be performed in pre-clinical models prior human to implantation. Although many animal models have been developed to assess function, integrity, thrombogenicity, hemodynamics and pathological behaviors, no standard model is accepted for aortic valve replacement (210). The ISO guideline 5840:2005 mentions that evaluation length should be no less than 90 days, but it does not indicate a preferential animal, possibly as each model possesses inherent advantages/limitations and does not completely replicate the human anatomy, biochemistry and physiology.
Sheep is by far the most employed animal model for pre-clinical valve studies (210, 211). Lambs have cardiac size and output similar to humans under 20 years old (5) and attain full growth within 2 years, which provides fast and valuable information in regards of growth and remodeling (212). Furthermore, they tolerate well cardiothoracic surgery and their elevated calcium metabolism allows to evidence degenerative processes in a short time (213, 214). However, sheep endothelialisation of implanted constructs is more rapid than humans, which renders extrapolation of the results to humans difficult (215).
Porcine models are less employed but present several advantages over sheep. Pigs have a fast growth, possess cardiac sizes and outputs similar to humans and do not readily endothelialize implants, which make them more human-like models. However, the surgical intervention for valve replacement in pigs is technically as stringent and difficult as for humans, and is therefore challenging. Furthermore, the costs associated with animals (especially Yucatan mini-pigs), husbandry and surgery are very high. As a results, pig models are preferred for acute and not long term studies (210).
Other models include dogs, calves and primates. Dogs have a convenient size, are easy to work with and do not have a rapid growth, which would make them interesting for long-term studies. However, animal rights concern are important for dogs (211) and they present a cardiac output much lower than humans. Conversely, calves rapid growth is accompanied by a dramatic increase of cardiac output, well above human values, which makes them more suited for very acute studies or for studying patient-prosthesis mismatch (210). Primates, aside from ethical concerns and very elevated costs (216), present annulus diameters of lower dimension than humans. Therefore, implants of relevant size for humans cannot be directly evaluated. However, the model has be successfully used as proof of concept for transapical tissue engineered pulmonary valve implantation (109).
Regardless of the animal model employed, the strategy followed consists in comparing post-operative measurements over time to reference data obtained immediately prior implantation and to control valves (mechanical or biological) (217). To obtain information of the implants structure and function during implantation, the animals are either euthanized at different time points and the explants analysed in a destructive fashion. Alternatively, noninvasive techniques can be used instead. The latter option is preferred for ethical and cost reasons but cannot be applied for all necessary characterizations.
An important parameter to consider in pre-clinical studies is site-specificity, since implants placed in positions other than the aortic one may behave differently due to the different biomechanic and hemodynamic environments. As a matter of fact, heart valve replacement pre-clinical studies have indeed been conducted preferentially in the pulmonary position (77, 95, 102, 105, 109, 129, 139, 145, 150, 171, 218). These studies provide valuable information but which cannot be readily extrapolated to the systemic pressure environment of the aortic valve. However, they can be used as guidelines for the necessary validation of aortic valve replacement strategies.
Tissue formation, remodeling and endothelial function
The different techniques employed during in vitro testing can similarly be employed to evaluate tissue formation, structure, organisation and remodeling over the course of implantation. Hematoxilin/eosin or Movat pentachrome staining procedures are useful to determine cellular ingrowth and tissue formation and structure (95, 129, 142, 150, 171, 218). Elastin and collagen deposition are more specifically visualized with van Gieson (129, 139) and Masson’s trichrome stainings (129), respectively, or using immunohistochemical techniques (142, 150). Quantification of the different ECM proteins can be performed with biochemical assays normalized by cell number (DNA quantification) (105, 129, 218)
Tissue remodeling within the implanted constructs can similarly be followed histologically, by comparing volume, ECM composition and cell phenotype of explants retrieved at different times. The presence of myofibroblastic-like cells similar to activated VICs is an important marker of tissue remodeling and should be actively sought. Immunohistochemistry directed against αSMA provides valuable information of the presence and distribution of such activated cells over time (129, 139, 142, 150).
The presence or restoration of a functional endothelial layer at the implant surface can be evaluated by SEM analysis evaluating cell morphology and the presence of adherent thrombocytes (142). Alternatively, immunohistological detection of CD31 (139), vWF (129, 142, 150) and eNOS (129, 142) are commonly employed.
Retention of the implanted cells after implantation is an important parameter to consider for understanding of their role in tissue formation. However, only few studies have sought to discriminate cell origin, using fluorescent labelling prior implantation (139).
The ultimate goal of these characterizations is to evidence the formation of a neotissue structurally and physiologically similar to native aortic valves. The formation of a tri-layered structure rapidly after implantation will undeniably be beneficial for the long term stability and function of the implant (95, 145). This criterion is nevertheless not sufficient. The consistent synthesis of elastin is as well a very important criterion for the proper function and longevity of the replaced valve (204, 205). Finally, the presence of activated myofibroblastic-like cells should not remain for too long periods, to prevent the excessive deposition of unorganized ECM and cause compaction and retraction of the constructs within the first 20 weeks of implantation, resulting in blood regurgitation and impaired function (Fig. 13) (14, 106, 129, 160, 219-221).
Tensile tests after explantation provide valuable information of the improvement of biomechanical properties linked to tissue formation and remodeling in the systemic flow. As in vitro characterizations, tensile tests are commonly used (95, 129, 218) and should preferentially be biaxial.
The evolution of anisotropic mechanical properties over implantation time should be particularly evaluated to evidence the structural maturation of the deposited collagen (13) and determine inconsistency with native valves and incoherent biomechanical behaviour (80).
Pre-clinical studies undeniably provide important information regarding the replaced valve functionality. They are as well of high interest to evidence pathological behaviors induces by the replaced valve and resembling natural valve pathologies (8).
Signs of thrombosis induced at the implant surface can be evidenced by SEM analysis (Fig. 14-A) (142) while inflammatory reaction due to the implant material or its degradation products can be assessed by immunostaining of macrophages (92).
Calcification and stenosis of the replaced leaflets can be investigated using various techniques. SEM combined with energy-dispersive X-ray spectroscopy (SEM–EDS) has been showed to confirm the presence of calcification nodules, isolated with focused ion beam, on valves (Fig. 14-B) (38). Raman microscopy can as well be employed to evidence the presence of calcified nodules (222) as well as classical histological stains targeting calcium (alizarin red) (92, 222) or phosphate (von Kossa) crystals (142).
The determination and measure of circulating biomarkers of aortic valve disorders would be highly beneficial to closely monitor the apparition of pathological behaviors due to the replaced valve. However, from the multiple early and late AVD markers candidates evaluated in clinical studies, only fetuin-A and osteopontin seem interesting as late marker for calcification (5). So far, these markers have never been monitored during pre-clinical studies of aortic valve replacement.
Non-invasive, in situ approaches for functionality, growth and remodeling evaluation
The aforementioned evaluation methods are endpoints that require euthanizing the animals. The development and use of non-invasive methods undeniably provides a better way of monitoring functionality and growth of the replaced aortic valve.
The most useful and straight forward method to non-invasively assess functionality and biomechanical/hemodynamic performance of replaced valves is Doppler echocardiography. This approach provides hemodynamic information over time on peak velocity across the valve, transvalvular systolic pressure gradient, left ventricular outflow tract velocity and insufficiency (223) as well as aortic valve area, effective orifice area, root dimensions, prosthesis size, morphology and function during growth of an animal (142, 224). Doppler is routinely used in post-operative follow up of patients after heart valve replacement to monitor the functional development of prosthetic heart valves and assess the severity of aortic regurgitation (225). It can therefore be used as well in pre-clinical studies to measure the intrinsic gradient after valve replacement and compare with subsequent measurements. From the orifice area obtained from Doppler, the presence and severity of stenosis can be assessed. Furthermore, the calculated orifice area and pressure gradients can be compared to the ones reported from commonly accepted biological prostheses (207). Two methods are currently used to perform echocardiographies: transoesophageal and transthoracic. Transoesophageal echocardiography is the most clinically used and more suited for diagnosing valvular diseases (226). However, transthoracic echocardiography is especially suitable for examination of replaced valves because of the proximity of the oesophagus to the heart and absence of interference with lungs and ribs (224). In summary, Doppler echocardiography is practical, cost effective, readily available and non-invasive. As a result it is a method of choice to study mechanical function and hemodynamic performance of replaced aortic heart valves. Nevertheless, some of the results obtained by echocardiography are underestimated and inter-operator variability is high (227).
Alternatives have been proposed to assess replaced valve functionality and hemodynamic performances such as magnetic resonance imaging (MRI) and multi–detector row CT (Fig. 15) (227). Multi-detector row CT has a superior resolution of aortic valve function and can quantify aortic valve calcification and stenosis (228, 229) while MRI can examine mechanical and hemodynamic properties of heart valves (blood flow velocity, valve opening and closing, valvular dysfunction, and especially regurgitation volumes by phase contrast MR). Other methods have been developed based on velocimetric techniques and computational analysis (laser Doppler anemometry, LDA, and particle image velocimetry) (230). However, these non-invasive techniques are so far limited to testing in vitro.
In addition to the evolution of valve leaflet over implantation time, the functionality and hemodynamic compliance of the aortic root should as well be monitored to determine the performance of replaced aortic valves. Indeed, flow pattern in the aortic root can affect valve function by allowance of load and stress sharing between leaflet and aortic wall. As a matter of fact, reduced aortic compliance and increased aortic stress lead to thickening of aortic valve leaflet (231). MRI combined with time-resolved particle traces and velocity vector fields can assess aortic root functionality and opening/closing motions of the leaflets simultaneously with morphological data (232, 233).
Valve growth and remodeling during implantation is more difficult to assess through non-invasive techniques although doppler echocardiography could possibly be used to deduce valve growth by monitoring of valve diameter and apparition of regurgitation over time (217). To determine ECM remodeling and organisation in situ, development of multi-photonic endoscope is currently investigated (234, 235). Collagen probes might as well be improved in the future to provide in situ markers of collagen formation (200, 202, 203). Other options relying on the use of activatable molecular imaging agent (near infrared fluorescence, NIRF) have been shown promising to follow proteolytic enzyme activity ex vivo. Matrix mettaloproteinase 2, (MMP2), matrix metalloproteinase 9 (MMP9), cathepsin B and cathepsin K have been evidenced with these techniques in early aortic valve disease (236) and atherosclerosis (237, 238). Future improvements might lead to the direct visualisation of proteolytic enzyme activity in the replaced valves, as was shown for cancer models in mice (239).
Finally, evaluation of endothelial function in situ is not yet foreseeable. However, expression of vascular cell adhesion molecule-1 (VCAM1) can be seen using VCAM1-targeted magnetofluorescent particles (Fig. 16) and MRI (240, 241). VCAM1 is present in inflammatory processes (242) and not in normal endothelium (236). As such, it could be used to detect pathological behavior of the implanted valves.
Aortic valve replacement has shown tremendous scientific and clinical progress in the last century. However, prosthetic heart valve replacement for the treatment of aortic valve disease has drawbacks and limitations. Tissue engineering of aortic valves has the potential to overcome these limitations by creating living substitutes with growth ability. If proof of concept has been demonstrated in animal models, the challenge lies in translating tissue engineered aortic valves to clinical practice. Aside from finding the optimal components for the tissue engineering approach, the elucidation of key biological processes in tissue formation and the definition of biomarkers linked to valve functionality must carefully be assessed prior implantation, to ultimately provide clinically valid engineered aortic valves.
The definition of effective translational platforms is an important step in this direction (Fig. 17). Such structure must regroup competences and expertise as multiple and complex as the aortic valve itself. To summarize, this translational effort should be based on a strong technological platform to design and evaluate biomaterials and their interactions with cells. Doing so, at a fundamental level, allows us to uncover and define key biological processes ruling tissue formation and regeneration. The corollary of such fundamental studies leads to the establishment of key biomarkers in animal models that provide the tools, through quality indicators and selection criteria, for monitoring and extrapolating the functionality of the replaced aortic valves. The role of clinical surgeons is critical to assess relevance of the results and provide enhancement feedback. Ultimately, such feedback must be the fertile soil on which functional predictive models can help establish the mandatory clinical safety and relevance prior human studies.
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