**5. Physiological behavior of heart valves under loading**

The ultrastructural histoarchitecture of the valvular tissue and the interaction of its collagen fibers with its non-collagenous components form the basis of the mechanical properties of heart valves. Similarly to other biological tissues that are subjected to high deformation and loading, collagen and elastic fibers act synergistically in heart valves to provide valvular tissue with strength and elasticity, respectively, which are required for their efficient function. The strength and stiffness of the collagen fibers prevent fracture of the valvular components when they are subjected to the peak transvalvular pressure during the time the valves are fully closed. The tensile strength of collagen is approximately 120 MPa, which is only one order of magnitude lower than high tensile steel (about 1110 MPa) and an elastic modulus of about 1.2 GPa, which confers substantial stiffness (**Table 4**) [35, 87]. However they stretch only minimally (about 13%) [35, 88]. On the other hand, elastic fibers are the most extensible biosolids known with very low modulus of approximately 0.3–1.1 MPa (**Table 4**) and capable of reaching failure strains in excess of 150%. However, elastic fibers demonstrate a low tensile strength of approximately 2 MPa, which limits their load-bearing capacity [35].

During the cardiac cycle, valvular tissue deforms to relative large strains, exhibiting non-linear stress-strain behavior. A typical stress-strain behavior for SL valve leaflet tissue within its physiological range is shown in **Figure 6a**. The graph describes the behavior of SL valve leaflets


\*Data obtained under uniaxial tensile testing.

collagen I, III, and V in normal human MV leaflets to be approximately 74, 24 and 2% of the total collagen content, respectively, whereas collagen IV and VI were below the detection limit of the technique used in those studies [72, 77]. A similar study by Lis et al. [78] reported a collagen I to collagen III ratio of 2.4 and 2.7 for human MV leaflets and chordae, respectively.

The results indicate mean ± 95% confidence intervals, or mean ± standard deviation (\*), or means ± standard error (\*\*).

**Leaflet**

70.7±6.2 (P) [71]

504.8±44.3 (P) [71]

20.2±0.6 (P) [71]

**Mitral Leaflet Mitral Chordae**

88.6±7.1 (P) [68]

632.6±50.7 (P) [68]

15.5±5.4(H)\* [78]

9.7±2.1 (H) \* [78]

88.2±3.9 (P) [68]

629.7±17.8 (P) [68]

26.8±3.5 (P) [68]

[78]

N.A. 10.0±17.8 (H) \*

**Content**

71.6±26.1 (H) [79]

511.2±186.3(H) [79]

11.4±0.4 (H) [79]

11.2±1.0 (H)\*\* [81]

H: Human; P: porcine; N.A.: not available. Numbers in brackets indicate source document.

**Table 3.** Hydroxyproline, collagen, sulfated GAG and elastin content of normal heart valves.

**Constituent Aortic Leaflet Pulmonary** 

Hydroxyproline (μm/mg dry

20 Structural Insufficiency Anomalies in Cardiac Valves

Total collagens (μm/mg dry

Sulfated GAGs (μm/mg dry

(% of dry tissue weight)

tissue)

tissue)

tissue)

Elastin

Three major GAGs have been identified in varying concentrations in heart valves, including chondroitin sulfate, dermatan sulfate and hyaluronic acid, together with decorin, biglycan and versican, which are chondroitin and dermatan sulfate proteoglycans [51, 84, 85]. The total content of sulfated GAGs (chondroitin and dermatan sulfate) in normal heart valves has been reported to range between 11 and 27% (dry tissue weight ratio), depending on valve type and site (**Table 3**) [68, 71, 78, 79]. Hyaluronic acid is the most abundant GAG in normal heart valves, accounting for up to half of the total leaflet GAG content, with chondroitin and dermatan sulfate accounting for about a quarter of the total GAG content of the leaflet each [86]. The relative percentages reported for normal human MV leaflets were 49, 25 and 23% of the total GAG content, for hyaluronic acid, chondroitin sulfate and dermatan sulfate, respectively, with the remaining 3% attributed to hyper-sulfated chondroitin/dermatan sulfate. In the case of the normal MV chordae, the corresponding percentages have been reported to be 24 (hyaluronic acid), 26 (chondroitin sulfate), 43 (dermatan sulfate) and 7% (hyper-sulfated chondroitin/dermatan sulfate) [84–86]. Moreover, differences in the absolute and relative GAG contents have been reported for different regions of the valves, subjected to different modes of loading during the cardiac cycle. Specifically, regions of the valves that are predominantly subjected to tensile loading, such as the leaflet belly and chordae (in the case of the AV valves), present a reduced overall GAG content compared to regions such as the free edge of the valve leaflets, which are predominantly subjected to compressive loading during coaptation with the other valve leaflets when the valve is fully closed. In addition, the regions that are predominantly subjected to tensile loading present significantly increased levels of

chondroitin and dermatan sulfate compared to hyaluronic acid [85].

**Table 4.** Indicative low-strain and high-strain modulus of valvular tissue and constituents.

from peak systole, when the SL valves are fully open, to peak diastole, when the SL valves are fully closed and loaded by the maximum transvalvular pressure. The graph also shows the contributions of the elastic and collagen fibers towards the overall behavior of the tissue and can be better comprehended in conjunction with **Figure 5c** and **d**. This type of stress-strain behavior has three distinct phases [65]. During the first phase (elastic phase), the leaflet offers little resistance to elongation since force transmission and load bearing is provided mainly by the elastic fibers. During this phase the collagen layer in the fibrosa unfolds and the collagen fibrils change their angular distribution [75]. Owing to these, the collagen fibers have minimal contribution to force transmission, resulting in a stress-strain response for this phase that is characterized by a low slope (low modulus). In the elastic phase the leaflet tissue behaves almost as an elastic solid with the stress increasing linearly with the strain. Under further loading, the leaflet enters the transition phase, during which the collagen fibers uncrimp and gradually align and uncoil, increasing their contribution to the force transmission. In the collagen phase, all the collagen fibers are uncoiled (recruited) and the load is entirely borne by them. Further extension in the collagen fibers occurs by extrafibrillar (between molecules) and interfibrillar shear in the fibers, as well as molecular distortion [35]. Interfibrillar shear has been reported to be dependent on the amount of proteoglycans associated with the surface of the fibrils, which influences the extent of GAG association and electrostatic interactions between the fibrils and the surrounding ECM [35]. The slope of the stress-strain curve for the collagen phase is steep (high modulus) and almost constant, reflecting the material properties of the collagen fibers, which allow limited elongation to fracture [89]. Although the collagen phase of the leaflet continues well beyond the physiological range before failure, corresponding to the reserve strength of the collagen fibers, after peak systole the valve starts opening again and the stress is relieved whilst the leaflets recoil back to their original shape/size at peak systole. A similar non-linear stress-strain profile can be observed for the case of the leaflets and chordae of the AV valve during their closing phase from peak diastole to peak systole. Moreover, experimental studies have indicated that the stress-strain response of valve leaflets were independent of strain rate [75]. The Young's modulus of different valvular tissue constituents, together with the elastic and collagen phase slopes of pulmonary and aortic valve leaflets, and MV leaflets and chordae, are listed in **Table 4**.

necessary energy for rapid retraction of the valve leaflets during opening [35, 75, 91]. The energy storage capacity of the elastic fibers has been linked the high entropy of the elastin molecules. In the case of collagen, the intrafibrillar and interfibrillar shear, and molecular distortion that occur during the collagen fiber stretching has been suggested to contribute to the elastic energy storage [35]. Overall, valve leaflets have been reported to demonstrate a relatively low hysteresis of approximately 12% (in the case of the MV), which is independent

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**Figure 6.** (a) Structure-property relation of SL valve leaflets within physiological ranges of stress and strain. Redrawn and modified from Schoen (1999) [64]. (b) Successive loading and unloading of biological tissue, showing the hysteresis

loop (hashed area). (c) Stress relaxation under constant strain. (d) Creep under constant stress.

of the strain rate [75].

Similarly to other biological tissues, the stress-strain behavior of valvular tissue also demonstrates viscoelastic behavior. Generally, viscoelasticity is manifested by a number of different features, including hysteresis, preconditioning, stress relaxation and creep. Viscoelasticity is a fundamental characteristic of biological materials, which exhibit both viscous and elastic behavior, depending on their constitution, temperature and the time over which the tissue is observed. Under cyclic loading (**Figure 6b**), valvular tissue exhibits a hysteresis loop (a phase lag) between loading and successive unloading. The area under the loading curve represents the energy stored during the extension of the tissue, whereas the area under the unloading curve represents the energy recovered during the recoil of the tissue back to its unloaded state. The hysteresis is the area between the loading and unloading curves of the stress-strain response and it is proportional to the mechanical energy dissipated. The hysteresis provides a measure of the energy storage efficiency of the tissue, with the larger the hysteresis the less efficient is the energy storage capacity of the tissue during loading and the less energy is returned to the system on unloading [90, 91]. Both collagen and elastic fibers do not show significant hysteresis, which makes them efficient in energy storage and able to provide the necessary energy for rapid retraction of the valve leaflets during opening [35, 75, 91]. The energy storage capacity of the elastic fibers has been linked the high entropy of the elastin molecules. In the case of collagen, the intrafibrillar and interfibrillar shear, and molecular distortion that occur during the collagen fiber stretching has been suggested to contribute to the elastic energy storage [35]. Overall, valve leaflets have been reported to demonstrate a relatively low hysteresis of approximately 12% (in the case of the MV), which is independent of the strain rate [75].

from peak systole, when the SL valves are fully open, to peak diastole, when the SL valves are fully closed and loaded by the maximum transvalvular pressure. The graph also shows the contributions of the elastic and collagen fibers towards the overall behavior of the tissue and can be better comprehended in conjunction with **Figure 5c** and **d**. This type of stress-strain behavior has three distinct phases [65]. During the first phase (elastic phase), the leaflet offers little resistance to elongation since force transmission and load bearing is provided mainly by the elastic fibers. During this phase the collagen layer in the fibrosa unfolds and the collagen fibrils change their angular distribution [75]. Owing to these, the collagen fibers have minimal contribution to force transmission, resulting in a stress-strain response for this phase that is characterized by a low slope (low modulus). In the elastic phase the leaflet tissue behaves almost as an elastic solid with the stress increasing linearly with the strain. Under further loading, the leaflet enters the transition phase, during which the collagen fibers uncrimp and gradually align and uncoil, increasing their contribution to the force transmission. In the collagen phase, all the collagen fibers are uncoiled (recruited) and the load is entirely borne by them. Further extension in the collagen fibers occurs by extrafibrillar (between molecules) and interfibrillar shear in the fibers, as well as molecular distortion [35]. Interfibrillar shear has been reported to be dependent on the amount of proteoglycans associated with the surface of the fibrils, which influences the extent of GAG association and electrostatic interactions between the fibrils and the surrounding ECM [35]. The slope of the stress-strain curve for the collagen phase is steep (high modulus) and almost constant, reflecting the material properties of the collagen fibers, which allow limited elongation to fracture [89]. Although the collagen phase of the leaflet continues well beyond the physiological range before failure, corresponding to the reserve strength of the collagen fibers, after peak systole the valve starts opening again and the stress is relieved whilst the leaflets recoil back to their original shape/size at peak systole. A similar non-linear stress-strain profile can be observed for the case of the leaflets and chordae of the AV valve during their closing phase from peak diastole to peak systole. Moreover, experimental studies have indicated that the stress-strain response of valve leaflets were independent of strain rate [75]. The Young's modulus of different valvular tissue constituents, together with the elastic and collagen phase slopes of pulmonary and aortic valve leaflets, and MV leaflets and chordae, are listed in **Table 4**. Similarly to other biological tissues, the stress-strain behavior of valvular tissue also demonstrates viscoelastic behavior. Generally, viscoelasticity is manifested by a number of different features, including hysteresis, preconditioning, stress relaxation and creep. Viscoelasticity is a fundamental characteristic of biological materials, which exhibit both viscous and elastic behavior, depending on their constitution, temperature and the time over which the tissue is observed. Under cyclic loading (**Figure 6b**), valvular tissue exhibits a hysteresis loop (a phase lag) between loading and successive unloading. The area under the loading curve represents the energy stored during the extension of the tissue, whereas the area under the unloading curve represents the energy recovered during the recoil of the tissue back to its unloaded state. The hysteresis is the area between the loading and unloading curves of the stress-strain response and it is proportional to the mechanical energy dissipated. The hysteresis provides a measure of the energy storage efficiency of the tissue, with the larger the hysteresis the less efficient is the energy storage capacity of the tissue during loading and the less energy is returned to the system on unloading [90, 91]. Both collagen and elastic fibers do not show significant hysteresis, which makes them efficient in energy storage and able to provide the

22 Structural Insufficiency Anomalies in Cardiac Valves

**Figure 6.** (a) Structure-property relation of SL valve leaflets within physiological ranges of stress and strain. Redrawn and modified from Schoen (1999) [64]. (b) Successive loading and unloading of biological tissue, showing the hysteresis loop (hashed area). (c) Stress relaxation under constant strain. (d) Creep under constant stress.

Following altered loading conditions, valvular tissue reorganizes itself to compensate for the altered mechanical stress, demonstrating an initial period of adjustment in its stress-strain behavior under cyclic loading [75]. This adjustment is manifested by an increased hysteresis loop, which subsequently decreases, tending to a steady state after a number of loading/ unloading cycles. Once this steady state is reached, no further change occurs in the stressstrain behavior, unless the loading routine is changed again. This period of adjustment after a large disturbance is called preconditioning and occurs due to internal changes in the structure of the tissue during cycling [6]. The viscoelastic feature of stress relaxation is manifested by the reduction of the stress generated in the tissue over time under constant strain (**Figure 6c**). Specifically, if the tissue is suddenly loaded to an initial stress σ<sup>o</sup> and its length held constant, the stress relaxes asymptotically to a limiting value σ<sup>1</sup> following an exponential decay. Creep is the counterpart of stress relaxation in the sense that the tissue is loaded to a strain ε<sup>o</sup> and the stress is held constant. Under these conditions the specimen continuous to deform asymptotically to a limiting value ε<sup>1</sup> (**Figure 6d**) [92]. Studies have reported that valve leaflets exhibit significant stress relaxation, but negligible creep over time, suggesting that they behave more like anisotropic quasi-elastic materials rather than viscoelastic materials. This behavior suggests that valve leaflets exhibit a load-locking behavior under maintained loading conditions that enables them to withstand high loading without any time-dependent deformation effects [75].

disruptions and alterations in the ECM histoarchitecture and constitution, disruptions in the distribution and organization of the EC and VIC populations, and malformation of the heart valves, which can render them stenotic and/or regurgitant [74, 75]. Moreover, several ECM gene mutations have been linked to valve disease, including fibrillin 1 gene mutations, which have been associated with BAVs and MVP, and elastic fiber component gene mutations also associated with MVP (Williams and Marfan syndromes) [98–100]. MVP and pulmonary valve stenosis have also been linked to collagen III and tenascin X gene mutations [98], whereas Notch1 gene mutations have been associated with BAV development and early calcification [101, 102]. Several studies have suggested that the abnormal organization of the ECM induced by these mutations may lead to abnormal VIC signaling and subsequent dysregulation of

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BAVs in the most common form of congenital valve disease, affecting 1–2% of the general population and eventually leading to aortic valve stenosis or regurgitation, infective endocarditis, and aortic dilation and/or dissection later in life [80, 97, 98, 106, 107]. Stenotic BAVs explanted from pediatric patients have been reported to exhibit excessive ECM production and disorganization, and VIC disarray without calcification. [98]. Specifically, the valve leaflets demonstrated loss of the typical trilaminar structure of the normal aortic valve leaflets, with disorganized, fragmented and abnormally oriented collagen and elastic fibers, increased proteoglycan presence throughout the leaflets, leaflet thickening, and large areas relatively void of cells. Moreover, elastin content was decreased, whereas collagen and proteoglycan content were substantially increased [75, 98]. The abnormalities in BAV histoarchitecture, constitution and anatomy, have been show to affect leaflet kinematics and stress distribution in computational studies, suggesting that early occurrence of regurgitation or stenosis might strictly depend on those abnormalities [107]. Alterations in the mechanical loading of the heart valves, due to abnormalities in histoarchitecture, constitution and anatomy, induce tissue remodeling though abnormal VIC mechanotransduction, which can lead to further valvular disease and dysfunction. Several studies have characterized VIC response against alterations in the biomechanical environment, and demonstrated a clear link between abnormal VIC stimulation, valvular tissue deformations and disease development and progression, highlighting the fundamental role of the mechanical environment on the mechanobiology of VICs [108–111]. As a result, BAVs are highly susceptible to calcification in later life, due to an induced osteoblastic VIC phenotype and subsequent matrix mineralization [80, 112]. Calcification causes valvular tissue to become thicker and stiffer (representative of a higher modulus), which eventually leads to valve stenosis and inevitable valve replacement [75].

Calcified aortic valve disease (CAVD) is not restricted to BAVs alone. CAVD is a slow, progressive, multifactorial disorder that is frequently driven by aging and the obesity-associated metabolic syndrome, and affects 25–30% of the population aged over 65 years old [75, 84, 100, 105, 113, 114]. Initially, the disease in manifested by mild leaflet thickening of the leaflets alongside with increase in the proteoglycan and hyaluronic acid content, which progressively become more severe and lead to impaired leaflet motion, valvular tissue adaptation and stenosis, [75, 84]. In spite of the leaflet thickening, studies have indicated that there is little change in the mechanical properties of the valve at the early stages of the disease [75]. Although this makes the condition asymptomatic at its initial phases, 10% of the patients develop severe symptoms within 10 years

ECM synthesis [98, 103–105].

The degrees of non-linear stress-strain behavior, hysteresis, preconditioning, stress relaxation and creep are different for different tissues, depending on the type and amount of their individual constituents [6, 92]. Along these lines, the variability in the biomechanical properties between the different valves and valve components (**Table 4**) is predominantly due to the different fractions and organization of the major ECM constituents, including collagen fibers, elastic fibers, GAGs and proteoglycans, present in the different valves and valve components. These constitutional and organizational variations, which are dictated by the specific hemodynamic and biomechanical environment that the valves reside in, bequeath high directional and regional histoarchitectural and biomechanical anisotropy to the valvular tissue, assisting the heart valves to perform their specific function in the four different sites of the heart.
