**3. Biomechanics of heart valve prostheses**

During their motion natural valve leaflets imposed to different mechanical loading and corresponding stress fields. Mechanical bending is developed during opening especially at sites near their attachments to valvular ring, while shear stress is gradually developed at their sides faced blood flow. A near parabolic velocity profile is produced in fully developed central axial blood flow across the valve which fast stabilizes leaflets in a position parallel to its axis. Upon starting of closing phase of the valves a reverse leaflet movement is

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 69

**ΕΗ (MPa)**

**COLLAGEN MODULUS**

BPN BPG

0,1 0,5 1 5 10 20 **Cyclic Frequency (Hz)**

Fig. 7. (a) Typical stress-strain diagram of a pericardial tissue under uniaxial tensile cyclic loading 0.1 Hz demonstrating non linear mechanical behaviour. EH is the characteristic high

Pathology of natural heart valve calcification causes valve dysfunction (stenosis, regurgitation, tissue rupture). A lot of possible aetiologies and mechanisms seem to be implemented in its initiation and development in human body tissues. Chronic pathologies, infections, metabolism, long term drag therapies and age related tissue degeneration, as well biochemical compounds involved in the structure of implanted biomaterials may contribute in different ways in growth of different types of calcium phosphate crystals from ions of electrolytes diluted in biological fluids, that finally deposited in valve tissue structure (Gross, 2003; Schoen & Levy, 2005). However, despite the complicated aetiologies and mechanisms of tissue calcification, crystal growth is basically a physicochemical process of calcium and phosphate ion crystallization under certain physicochemical conditions that

Calcium phosphate deposition on implants may result from the presence of high phosphorus levels in the biological fluids in contact with the implanted surface. Due to their very low solubility products, a number of phosphate scale minerals may form in aqueous supersaturated solutions. In the order of decreasing solubility, they are listed in the following table 2. At high solution supersaturations it is possible that a number of precursor phases may be formed, depending on the solution pH, which finally transform into the thermodynamically more stable HAP, in accordance to Ostwald's rule of stages which predicts that the least stable phase having the highest solubility is formed preferentially during a stepwise precipitation process. It is well established that kinetic factors may be more important in determining the nature and, hence, the characteristics of the solid deposits formed during the precipitation process than the respective equilibrium consideration complications that may arise from the formation of mixed solid phases,

(collagen) modulus, the tangential modulus of the second linear phase. (b) Collagen modulus of fresh natural (BPN) and glutaraldehyde treated (BPG) bovine pericardial tissue

(a) (b)

**EH**

(mean ± SE) at different cyclic loading rates

0 0,05 0,1 0,15 0,2 0,25 **Strain ΔL/L0**

may be satisfied during the function of a living organism.

**4.1 Biomineralization: Physicochemical background** 

caused of the overgrowth of one crystalline phase over another.

**4. Heart valve calcification** 

**BPG 0,1 Hz #DHML5B**

0 0,2 0,4 0,6 0,8 1 1,2

**Stress MPa**

performed with changing stress fields applied to their structure, till the closing position during which these membranous tissues, supported at their sites of attachment on valvular rings and their coaptation areas, are imposed in surface tensile loading, while obstructing blood backflow and big pressure differences between their sides. As a result of all these different mechanical loadings imposed, leaflet tissue is remodelled in a multilaminate, anisotropic (see diagram in figure 6) reinforced composite biomaterial, demonstrating a structure of multilayered 3D collagen and elastin fibrous networks, integrated and moved into an amorphous organic protein matrix, filled with cells and extracellular water, electrolytes and soluble proteins of low molecular weight. This tissue structure, different in individual valve leaflets, is specifically able to function at the specific anatomic position. Been in a different position, like, for example, from pulmonary artery to aorta as in Ross procedure, tissue remodelling started to make the pulmonary valve leaflets able to resist higher mechanical loading in aortic position, compared with that of pulmonary circulation.

Fig. 6. Diagram (from measurements in histological sections) of the laminate and total leaflet thickness changed under different pressure levels applied during their fixation with glutaraldehyde. In dynamic mode, the valve was under normal function during fixation. Anisotropic deformation of the different tissue laminates is demonstrated as pressure increases

Natural heart valve leaflets exhibit non linear viscoelastic mechanical behaviour under mechanical loading (figure 7a). A similar mechanical behaviour is demonstrated by tissue heart valves of all species, as well all membranous soft tissues. Chemical modification of bioprosthetic heart valves, needed for removing antigenic factors and stabilization against enzyme biodegradation, resulted in significant stiffening of leaflet tissue compared with its natural state, as demonstrated by increase of high (collagen) modulus (figure 7b).

Other viscoelastic mechanical parameters of valve leaflet tissue, like low (elastin) modulus (the slope of the first linear part of the loading stress-strain curve (fig. 7a)) and relaxation index may also changed, although hysteresis, another viscoelastic characteristic of membranous soft tissues (defined as the ratio of dissipated to the loading energy in every loading-unloading cycle) demonstrating energy dissipation inside the tissue during cyclic deformation, measured at 20/35% of loading energy depending on cyclic frequency seems to be unchanged after chemical modification.

Fig. 7. (a) Typical stress-strain diagram of a pericardial tissue under uniaxial tensile cyclic loading 0.1 Hz demonstrating non linear mechanical behaviour. EH is the characteristic high (collagen) modulus, the tangential modulus of the second linear phase. (b) Collagen modulus of fresh natural (BPN) and glutaraldehyde treated (BPG) bovine pericardial tissue (mean ± SE) at different cyclic loading rates

## **4. Heart valve calcification**

68 Aortic Valve Surgery

performed with changing stress fields applied to their structure, till the closing position during which these membranous tissues, supported at their sites of attachment on valvular rings and their coaptation areas, are imposed in surface tensile loading, while obstructing blood backflow and big pressure differences between their sides. As a result of all these different mechanical loadings imposed, leaflet tissue is remodelled in a multilaminate, anisotropic (see diagram in figure 6) reinforced composite biomaterial, demonstrating a structure of multilayered 3D collagen and elastin fibrous networks, integrated and moved into an amorphous organic protein matrix, filled with cells and extracellular water, electrolytes and soluble proteins of low molecular weight. This tissue structure, different in individual valve leaflets, is specifically able to function at the specific anatomic position. Been in a different position, like, for example, from pulmonary artery to aorta as in Ross procedure, tissue remodelling started to make the pulmonary valve leaflets able to resist higher mechanical loading in aortic position, compared with that of pulmonary circulation.

2 mm Hg 5 mm Hg 16 mm Hg 45 mm Hg 90 mm Hg dynamic

Fig. 6. Diagram (from measurements in histological sections) of the laminate and total leaflet

Other viscoelastic mechanical parameters of valve leaflet tissue, like low (elastin) modulus (the slope of the first linear part of the loading stress-strain curve (fig. 7a)) and relaxation index may also changed, although hysteresis, another viscoelastic characteristic of membranous soft tissues (defined as the ratio of dissipated to the loading energy in every loading-unloading cycle) demonstrating energy dissipation inside the tissue during cyclic deformation, measured at 20/35% of loading energy depending on cyclic frequency seems

thickness changed under different pressure levels applied during their fixation with glutaraldehyde. In dynamic mode, the valve was under normal function during fixation. Anisotropic deformation of the different tissue laminates is demonstrated as pressure increases Natural heart valve leaflets exhibit non linear viscoelastic mechanical behaviour under mechanical loading (figure 7a). A similar mechanical behaviour is demonstrated by tissue heart valves of all species, as well all membranous soft tissues. Chemical modification of bioprosthetic heart valves, needed for removing antigenic factors and stabilization against enzyme biodegradation, resulted in significant stiffening of leaflet tissue compared with its

natural state, as demonstrated by increase of high (collagen) modulus (figure 7b).

**Laminate and total thickness of aortic valve leaflet** 

**(Circumferencial)** Fibrosa

Spongiosa Ventricularis

**Fixation pressure**

to be unchanged after chemical modification.

**Thickness (1/100 mm)**

Pathology of natural heart valve calcification causes valve dysfunction (stenosis, regurgitation, tissue rupture). A lot of possible aetiologies and mechanisms seem to be implemented in its initiation and development in human body tissues. Chronic pathologies, infections, metabolism, long term drag therapies and age related tissue degeneration, as well biochemical compounds involved in the structure of implanted biomaterials may contribute in different ways in growth of different types of calcium phosphate crystals from ions of electrolytes diluted in biological fluids, that finally deposited in valve tissue structure (Gross, 2003; Schoen & Levy, 2005). However, despite the complicated aetiologies and mechanisms of tissue calcification, crystal growth is basically a physicochemical process of calcium and phosphate ion crystallization under certain physicochemical conditions that may be satisfied during the function of a living organism.

#### **4.1 Biomineralization: Physicochemical background**

Calcium phosphate deposition on implants may result from the presence of high phosphorus levels in the biological fluids in contact with the implanted surface. Due to their very low solubility products, a number of phosphate scale minerals may form in aqueous supersaturated solutions. In the order of decreasing solubility, they are listed in the following table 2. At high solution supersaturations it is possible that a number of precursor phases may be formed, depending on the solution pH, which finally transform into the thermodynamically more stable HAP, in accordance to Ostwald's rule of stages which predicts that the least stable phase having the highest solubility is formed preferentially during a stepwise precipitation process. It is well established that kinetic factors may be more important in determining the nature and, hence, the characteristics of the solid deposits formed during the precipitation process than the respective equilibrium consideration complications that may arise from the formation of mixed solid phases, caused of the overgrowth of one crystalline phase over another.

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 71

DCPD

OCP

HAP

TCP

1/

ν

<sup>∞</sup> *<sup>s</sup>* (2)

(1)

*o*

refer to solution and equilibrium conditions respectively, *α* denote

4.0 5.0 6.0 7.0 8.0 9.0 10.0 11.0

pH

Fig. 8. Solubility isothems for calcium phosphate phases at 25°C, Ionic strength 0.1M NaCl Spontaneous precipitation investigations have been faced with the problem that at high supersaturations the ACP forming initially results in a rapid decrease of the calcium and phosphate ion activities which fall bellow the values needed for the spontaneous formation of other calcium phosphate phases. A careful analysis of the precipitation of calcium phosphate over the pH range 5.0-8.0 suggested that the limiting ion activity product for ACP was constant as expected for a discrete mineral phase. Moreover it has also been shown that the presence of magnesium in the precipitation medium promoted the formation

The driving force for the formation of a solid phase in a continuous aqueous phase is the solution supersaturation which can be developed in many ways including temperature fluctuation, pH change, mixing of incompatible waters, increasing the concentration by evaporation or solid dissolution etc. Although supersaturation is the driving force for the formation of a salt, the exact values in which precipitation occurs are quite different from salt to salt and as a rule, the degree of supersaturation needed for a sparingly soluble salt is orders of magnitude higher than the corresponding value for a soluble salt. For sparingly

of DCPD at the expense of ACP (Feenstra & de Bruyn, 1979; Posner et al., 1984).

( ) ( )

ν

*<sup>a</sup> IP <sup>S</sup> a a K* α

+ − α

*m*

( ) ( )

Δμ = − μ μ

ν

*m a M s s*

+ −

*<sup>s</sup> M A*

the activities of the respective ions and *ν++ν-*= *ν*. *IP* and *<sup>o</sup> Ks* are the ion products in the

The fundamental driving force for the formation of a salt from a supersaturated solution is the difference in chemical potential of the solute in the supersaturated solution from the

+ − Α + − ∞ ∞ <sup>⎧</sup> <sup>⎫</sup> ⎪ ⎪ ⎛ ⎞ <sup>=</sup> ⎨ ⎬ <sup>=</sup> ⎜ ⎟ ⎜ ⎟ ⎪ ⎪ ⎝ ⎠ <sup>⎩</sup> <sup>⎭</sup>

ν

ν

soluble salts Mν+Aν- the supersaturation ratio, *S*, is defined as:

∞

supersaturated solution and at equilibrium respectively.




log ( Ca

where subscripts *s* and

respective value at equilibrium:

t / M )




0.0


Table 2. Calcium phosphate crystalline phases, formulae and corresponding thermodynamic solubility products. \*(Hench & Wilson, 1991), \*\*(Eanes et al., 1965), \*\*\*(LeGeros et al., 1975), \*\*\*\*(Betts & Posner, 1974)

The tendency for a particular calcium phosphate phase to form in supersaturated aqueous media may be determined from the solubility phase diagrams such as the diagram shown in figure 8. It has been reported that when calcium phosphate is precipitated from highly supersaturated solutions forms an unstable precursor phase. This phase is characterized by the absence of peaks in the powder x-ray diffraction pattern and is known as the amorphous calcium phosphate (ACP). The composition of ACP appears to depend upon the precipitation conditions and is usually formed in supersaturated solutions at pH >7.0 (Betts & Posner, 1974; Eanes et al., 1965; LeGeros et al., 1975; Newesely, 1966). In slightly acidic calcium phosphate solutions the monoclinic DCPD is formed (Bets & Posner, 1974; Brown & Lehr, 1959). OCP is formed by the hydrolysis of DCPD in solutions of pH 5-6 and may also be precipitated heterogeneously upon TCP (Brown et al., 1957). HAP is the thermodynamically most stable phase and often, when precipitated in the presence of foreign ions, substitution of calcium, phosphate and/or hydroxyls by some of these ions take place. Thus, substitutions of OH- by F- or Cl- ions, of the phosphate by sulfate and carbonate and of the calcium by Sr2+, Mg2+ and Na+ ions have been reported (Heughebaert et al., 1983; Legeros R.Z & Legeros J.P., 1984; Moreno & Varughese, 1981; Nathan, 1984).

A considerable amount of the work done for the identification of calcium phosphate minerals which precipitates spontaneously has been based on the stoichiometric molar ratio of calcium to phosphate calculated from the respective changes in the solutions. This ratio has been found in several cases to be 1.45±0.05 which is considerably lower than the value of 1.67 corresponding to HAP which is generally implied as the precipitating mineral. A number of different precursor phases have been postulated to be formed including TCP (Montel et al., 1981; Narasaraju & Phebe, 1996; Walton et al., 1967), OCP (Eanes & Posner, 1968, Posner, 1969) and DCPD (Furedi-Milhofer et al., 1976). On the basis of the analysis of the induction times preceding the spontaneous precipitation of calcium phosphate, it was concluded that at high solution supersaturations the initially forming ACP was converted into an apatitic mineral through an OCP precursor phase formation (Fransis & Webb, 1971).

**Product** 

1.87x10-7(mol L-1)\*

9.2x10-7 (mol L-1)\*

2.8x10-9 (mol L-1)\*\*

2.5x10-99(mol L-1)\*\*\*

5.5x10-118 (mol L-1)\*\*\*\*

**Solid Phase Abbrev. Formula Therm.Solub.** 

CaHPO4.2H2O

CaHPO4

Ca3(PO4)2

(o≤x≤2)

Ca8H2(PO4)6.5H2O

Ca10-x(HPO4)x(PO4)6-x(OH)2-x

The tendency for a particular calcium phosphate phase to form in supersaturated aqueous media may be determined from the solubility phase diagrams such as the diagram shown in figure 8. It has been reported that when calcium phosphate is precipitated from highly supersaturated solutions forms an unstable precursor phase. This phase is characterized by the absence of peaks in the powder x-ray diffraction pattern and is known as the amorphous calcium phosphate (ACP). The composition of ACP appears to depend upon the precipitation conditions and is usually formed in supersaturated solutions at pH >7.0 (Betts & Posner, 1974; Eanes et al., 1965; LeGeros et al., 1975; Newesely, 1966). In slightly acidic calcium phosphate solutions the monoclinic DCPD is formed (Bets & Posner, 1974; Brown & Lehr, 1959). OCP is formed by the hydrolysis of DCPD in solutions of pH 5-6 and may also be precipitated heterogeneously upon TCP (Brown et al., 1957). HAP is the thermodynamically most stable phase and often, when precipitated in the presence of foreign ions, substitution of calcium, phosphate and/or hydroxyls by some of these ions take place. Thus, substitutions of OH- by F- or Cl- ions, of the phosphate by sulfate and carbonate and of the calcium by Sr2+, Mg2+ and Na+ ions have been reported (Heughebaert et al., 1983; Legeros R.Z & Legeros J.P., 1984; Moreno & Varughese, 1981; Nathan, 1984). A considerable amount of the work done for the identification of calcium phosphate minerals which precipitates spontaneously has been based on the stoichiometric molar ratio of calcium to phosphate calculated from the respective changes in the solutions. This ratio has been found in several cases to be 1.45±0.05 which is considerably lower than the value of 1.67 corresponding to HAP which is generally implied as the precipitating mineral. A number of different precursor phases have been postulated to be formed including TCP (Montel et al., 1981; Narasaraju & Phebe, 1996; Walton et al., 1967), OCP (Eanes & Posner, 1968, Posner, 1969) and DCPD (Furedi-Milhofer et al., 1976). On the basis of the analysis of the induction times preceding the spontaneous precipitation of calcium phosphate, it was concluded that at high solution supersaturations the initially forming ACP was converted into an apatitic mineral through an OCP precursor phase formation (Fransis & Webb, 1971).

Ca10(PO4)6(OH)6

Table 2. Calcium phosphate crystalline phases, formulae and corresponding thermodynamic solubility products. \*(Hench & Wilson, 1991), \*\*(Eanes et al., 1965), \*\*\*(LeGeros et

DCPD

DCPA

TCP

OCP

HAP

al., 1975), \*\*\*\*(Betts & Posner, 1974)

Dicalcium phosphate dihydrate Dicalcium phosphate anhydrous β.Tricalcium phosphate Octacalcium phosphate Hydroxyapatite Defect Apatites

Fig. 8. Solubility isothems for calcium phosphate phases at 25°C, Ionic strength 0.1M NaCl

Spontaneous precipitation investigations have been faced with the problem that at high supersaturations the ACP forming initially results in a rapid decrease of the calcium and phosphate ion activities which fall bellow the values needed for the spontaneous formation of other calcium phosphate phases. A careful analysis of the precipitation of calcium phosphate over the pH range 5.0-8.0 suggested that the limiting ion activity product for ACP was constant as expected for a discrete mineral phase. Moreover it has also been shown that the presence of magnesium in the precipitation medium promoted the formation of DCPD at the expense of ACP (Feenstra & de Bruyn, 1979; Posner et al., 1984).

The driving force for the formation of a solid phase in a continuous aqueous phase is the solution supersaturation which can be developed in many ways including temperature fluctuation, pH change, mixing of incompatible waters, increasing the concentration by evaporation or solid dissolution etc. Although supersaturation is the driving force for the formation of a salt, the exact values in which precipitation occurs are quite different from salt to salt and as a rule, the degree of supersaturation needed for a sparingly soluble salt is orders of magnitude higher than the corresponding value for a soluble salt. For sparingly soluble salts Mν+Aν- the supersaturation ratio, *S*, is defined as:

$$S = \left| \frac{\left(a\_{M^{m+}}\right)\_{s}^{\nu+} \left(a\_{A^{a-}}\right)\_{s}^{\nu-}}{\left(a\_{M^{m+}}\right)\_{\infty}^{\nu+} \left(a\_{A^{a-}}\right)\_{\infty}^{\nu-}} \right| = \left(\frac{IP}{K\_{s}^{\circ}}\right)^{1/\nu} \tag{1}$$

where subscripts *s* and ∞ refer to solution and equilibrium conditions respectively, *α* denote the activities of the respective ions and *ν++ν-*= *ν*. *IP* and *<sup>o</sup> Ks* are the ion products in the supersaturated solution and at equilibrium respectively.

The fundamental driving force for the formation of a salt from a supersaturated solution is the difference in chemical potential of the solute in the supersaturated solution from the respective value at equilibrium:

$$
\Delta \mu = \mu\_{\text{o}} - \mu\_{\text{s}} \tag{2}
$$

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 73

biological environment but in accelerated process conditions, in order to study calcification mechanisms. In vitro calcification models have also proposed. These models have the advantage of studying the role of isolated or group of parameters that may contribute in calcification. Both models are very useful especially as pre-screening methods for studying the efficacy of various anticalcification treatments proposed (Bailey et al., 2004; Gross, 2003;

**OUT**

**Magnetic Stirrer**

**H H**

**T**

**H H**

**T N**

**G S**

**Electric Burette**

**A**

**<sup>B</sup> <sup>C</sup>**

The use of various model experimental procedures has resulted in the suggestion that the calcification is initiated at the matrix vesicles (Anderson, 1983; Wuthier, 1982), by acid phospholipids (Boskey, 1981; Boskey & Posner, 1977) or by heterogeneous nucleation of various calcium phosphate phases in the body fluids considered as aqueous solutions supersaturated with respect to the salts formed (Kim & Trump, 1975; Dallas et al., 1989;, Schoen et al., 1985; Brown et al., 1988; Grott et al., 1992). Despite intensive research, it seems that there is no agreement on the mechanism of formation of calcific deposits. As a result, the evaluations of bioprosthetic materials used for cardiac valve replacements are often met

Fig. 9. Experimental set-up for the quantitative investigation of the calcification of heart valves in vitro. (A) Automatic titrator apparatus (B) Mounting of heart valves (C) Lid of the

**Glass / SCE electrode**

**Reactor**

**Supersaturated solution**

**Thermostated water IN**

Kapolos et al., 1997; Krings et al., 2009; Schoen & Levy, 2005).

**pH meter**

**Impulsomat**

reactor (Kapolos et al., 1997)

Since the chemical potential is expressed in terms of the standard potential and the activity, *α*, of the solute:

$$
\mu = \mu^o + RT\ln a \tag{3}
$$

where *R* and *T* are the gas constant and the absolute temperature respectively. Substitution of eq. (3) to eq.(2) gives for the driving force for solid deposition (Mullin, 1993):

$$\frac{\Delta\mu}{RT} = \ln\left(\frac{a\_s}{a\_\infty}\right) = \ln S \tag{4}$$

For electrolyte solutions the mean ionic activity is taken:

$$a = a\_{\pm}^{\ \nu} \left( \nu = \nu\_{+} + \nu\_{-} \right) \tag{5}$$

and

$$\frac{\Delta\mu}{RT} = \ln\left(\frac{\alpha\_{\pm,s}}{\alpha\_{\pm,o}}\right)^{\frac{1}{\nu}} = \ln S \tag{6}$$

#### **4.2 Crystal growth in heart valves**

Degeneration of the leaflet tissue, together with calcification constitutes the principal reasons for bioprosthetic valve failure (Schoen et al., 1992). Calcification consists as already mentioned in the formation of sparingly soluble salts of calcium phosphate due to the presence of high levels of calcium and phosphate in blood serum (Schneck, 1995). Although the calcific deposits consist of apatitic calcium phosphate (HAP containing mainly carbonate, fluoride, magnesium and sodium) the formation of transient precursor phases such as DCPD and OCP is possible, as in vitro studies have shown (Brown et al., 1957; Heughebaert et al., 1983; Moreno & Varughese, 1981). The formation of calcium phosphates on porcine heart valves to a percentage of 30-50% is responsible for their dysfunction after 12-15 years due to stenosis or insufficiency (Narasaraju & Rao, 1979).

Despite the fact that the thermodynamic driving force in blood serum for nucleation and growth is sufficiently high for the homogeneous formation of calcium phosphates, the process is believed to be heterogeneous as dead cell remnants, lipids or degenerative collagen fragments may provide active sites for heterogeneous nucleation. Chemical treatment of porcine and pericardial bioprosthetic valves with glutaraldehyde is considered as one of the main causes of valve calcification (Hammersmeister et al., 1993; Schoen & Levy, 1992). Despite intensive research of the past few decades, the mechanism of initiation and development of calcific deposits on tissues in contact with blood is still poorly understood. As a result the development and production of biological valves resistant to calcification is still a major challenge (Grabenwoger et al., 1996; Schoen et al, 1987). Biological, chemical and mechanical factors seem to play a significant role in the kinetics of the process of calcification (Zipkin, 1970).

#### **4.3 Models of heart valve calcification**

Heart valve calcification is a slow process, difficult to be studied in vivo in humans. Various animal models, like the rat subcutaneous implantation, have been used for simulation of

Since the chemical potential is expressed in terms of the standard potential and the activity,

where *R* and *T* are the gas constant and the absolute temperature respectively. Substitution

ln ln *<sup>s</sup> <sup>a</sup> <sup>S</sup>*

1 , , ln ln *<sup>s</sup> S*

ν

∞ <sup>Δ</sup> ⎛ ⎞ = = ⎜ ⎟ ⎝ ⎠

*a a* ( ) ν = = + <sup>±</sup> ν ν ν

> α

± ± ∞ <sup>Δ</sup> ⎛ ⎞ <sup>=</sup> ⎜ ⎟ <sup>=</sup> ⎜ ⎟ ⎝ ⎠

Degeneration of the leaflet tissue, together with calcification constitutes the principal reasons for bioprosthetic valve failure (Schoen et al., 1992). Calcification consists as already mentioned in the formation of sparingly soluble salts of calcium phosphate due to the presence of high levels of calcium and phosphate in blood serum (Schneck, 1995). Although the calcific deposits consist of apatitic calcium phosphate (HAP containing mainly carbonate, fluoride, magnesium and sodium) the formation of transient precursor phases such as DCPD and OCP is possible, as in vitro studies have shown (Brown et al., 1957; Heughebaert et al., 1983; Moreno & Varughese, 1981). The formation of calcium phosphates on porcine heart valves to a percentage of 30-50% is responsible for their dysfunction after

Despite the fact that the thermodynamic driving force in blood serum for nucleation and growth is sufficiently high for the homogeneous formation of calcium phosphates, the process is believed to be heterogeneous as dead cell remnants, lipids or degenerative collagen fragments may provide active sites for heterogeneous nucleation. Chemical treatment of porcine and pericardial bioprosthetic valves with glutaraldehyde is considered as one of the main causes of valve calcification (Hammersmeister et al., 1993; Schoen & Levy, 1992). Despite intensive research of the past few decades, the mechanism of initiation and development of calcific deposits on tissues in contact with blood is still poorly understood. As a result the development and production of biological valves resistant to calcification is still a major challenge (Grabenwoger et al., 1996; Schoen et al, 1987). Biological, chemical and mechanical factors seem to play a significant role in the kinetics of

Heart valve calcification is a slow process, difficult to be studied in vivo in humans. Various animal models, like the rat subcutaneous implantation, have been used for simulation of

α

= + (3)

(4)

(6)

+ − (5)

ο μ μ

of eq. (3) to eq.(2) gives for the driving force for solid deposition (Mullin, 1993):

*RT* μ

12-15 years due to stenosis or insufficiency (Narasaraju & Rao, 1979).

*RT a* μ

ln *RT a*

For electrolyte solutions the mean ionic activity is taken:

**4.2 Crystal growth in heart valves** 

the process of calcification (Zipkin, 1970).

**4.3 Models of heart valve calcification** 

*α*, of the solute:

and

biological environment but in accelerated process conditions, in order to study calcification mechanisms. In vitro calcification models have also proposed. These models have the advantage of studying the role of isolated or group of parameters that may contribute in calcification. Both models are very useful especially as pre-screening methods for studying the efficacy of various anticalcification treatments proposed (Bailey et al., 2004; Gross, 2003; Kapolos et al., 1997; Krings et al., 2009; Schoen & Levy, 2005).

Fig. 9. Experimental set-up for the quantitative investigation of the calcification of heart valves in vitro. (A) Automatic titrator apparatus (B) Mounting of heart valves (C) Lid of the reactor (Kapolos et al., 1997)

The use of various model experimental procedures has resulted in the suggestion that the calcification is initiated at the matrix vesicles (Anderson, 1983; Wuthier, 1982), by acid phospholipids (Boskey, 1981; Boskey & Posner, 1977) or by heterogeneous nucleation of various calcium phosphate phases in the body fluids considered as aqueous solutions supersaturated with respect to the salts formed (Kim & Trump, 1975; Dallas et al., 1989;, Schoen et al., 1985; Brown et al., 1988; Grott et al., 1992). Despite intensive research, it seems that there is no agreement on the mechanism of formation of calcific deposits. As a result, the evaluations of bioprosthetic materials used for cardiac valve replacements are often met

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 75

where *kg* is the rate constant for crystal growth, *f(S)* a function of the total number of the growth sites available and *n* the apparent order of the crystal growth process. Logarithmic plots according to eq.7 yielded the kinetics shown in figure 12. The practical implications of this finding is that the deposition of the mineral phase on the membrane matrix is controlled

Fig. 11. Scanning electron micrographs of : (a) surface of the valves (b) OCP deposited on the

Possible strategies aiming at the retardation of the calcification process should therefore rely on the alteration of the surfaces so as to make surface integration more difficult. An additional feature revealed by the kinetics plots at constant supersaturation (figure 12) is that the glutaraldehyde treated porcine valves are substrates favoring the mineral

Fig. 12. Logarithm of the rate of OCP formation on glutaraldeyde treated porcine valves (black circles) as a function of the relative solution supersaturation; pH 7.4, 37° C, 0.15M NaCl. Open squares and triangles refer to literature results. Experiments performed in our laboratory with the same methodology on fresh, untreated porcine valves have shown that these tissues failed to induce any formation of calcium phosphate deposits although they

were kept in the mineralizing solution for as long as four days

valves at higher relative supersaturation and (c) OCP+HAP on valves at low relative

by the diffusion of the growth units on the OCP nuclei been formed.

**a b c** 

supersaturation

nucleation and growth.

with not unfounded criticism. One of the most successful experimental models presented, which is appropriate for the in vitro investigation of mineralization processes is the constant supersaturation model introduced and further developed by Nancollas and co-workers (Thomson & Nancollas, 1978; Amzad et al., 1978). This methodology allows for the solution supersaturation to be kept constant by the addition of titrant solutions. The concentrations of the reagents in these solutions are calculated so that full replacement of the precipitated mass is ensured. In the case of tests performed on the mineralization of heart valves the rates obtained from this system were proportional to the solution's supersaturation.

An in vitro model, based in the constant supersaturation model was introduced in 1997 in our laboratories for the investigation of heart valve calcification in vitro. The experimental set-up is shown in figure 9. Porcine aortic heart valves, treated with glutaraldehyde were sutured on Plexiglas® frames to be kept in flat position during stirring. The frames were immersed in supersaturated solution with respect to calcium and phosphate ions in near physiological concentrations, at 37°C. Solution's pH drop, associated with initiation of crystal growth deposited out of solution, triggered titrant addition with movement of computer controlled syringe pumps.

Fig. 10. Titrant addition for maintaining constant supersaturation during the course of mineralization of artificial heart valves. The rates were increased with increasing solution supersaturation (σ=0.72 , 1.09, 1.25 from the lower to the upper curve respectively)

Figure 10 presents experimental recordings of titrant volume added with time, in order to maintain solution's supersaturation constant. The signal from a pH drop sensitive device resulted from the growth of calcium phosphate crystals precipitated out of the solution, triggered titrant addition in small time steps. The diagrams correspond to different solution supersaturations. The mineral phase deposited on the glutaraldehyde treated porcine valves was identified as OCP, hydrolysed partially to HAP. The extent of hydrolysis was larger at the lower supersaturations, while at higher supersaturations with respect to OCP, this phase was stabilized, as may be seen in the scanning electron micrographs shown in figure11a-c. The characteristic plate like crystals are clearly seen in figure 11b.The rates, calculated from the titrants addition rates and normalized per unit geometric surface area of the exposed valves, were fitted to the semi empirical equation:

$$R\_{\chi} = k\_{\chi} f(S) \sigma^n \tag{7}$$

with not unfounded criticism. One of the most successful experimental models presented, which is appropriate for the in vitro investigation of mineralization processes is the constant supersaturation model introduced and further developed by Nancollas and co-workers (Thomson & Nancollas, 1978; Amzad et al., 1978). This methodology allows for the solution supersaturation to be kept constant by the addition of titrant solutions. The concentrations of the reagents in these solutions are calculated so that full replacement of the precipitated mass is ensured. In the case of tests performed on the mineralization of heart valves the

An in vitro model, based in the constant supersaturation model was introduced in 1997 in our laboratories for the investigation of heart valve calcification in vitro. The experimental set-up is shown in figure 9. Porcine aortic heart valves, treated with glutaraldehyde were sutured on Plexiglas® frames to be kept in flat position during stirring. The frames were immersed in supersaturated solution with respect to calcium and phosphate ions in near physiological concentrations, at 37°C. Solution's pH drop, associated with initiation of crystal growth deposited out of solution, triggered titrant addition with movement of

rates obtained from this system were proportional to the solution's supersaturation.

Fig. 10. Titrant addition for maintaining constant supersaturation during the course of mineralization of artificial heart valves. The rates were increased with increasing solution supersaturation (σ=0.72 , 1.09, 1.25 from the lower to the upper curve respectively)

Figure 10 presents experimental recordings of titrant volume added with time, in order to maintain solution's supersaturation constant. The signal from a pH drop sensitive device resulted from the growth of calcium phosphate crystals precipitated out of the solution, triggered titrant addition in small time steps. The diagrams correspond to different solution supersaturations. The mineral phase deposited on the glutaraldehyde treated porcine valves was identified as OCP, hydrolysed partially to HAP. The extent of hydrolysis was larger at the lower supersaturations, while at higher supersaturations with respect to OCP, this phase was stabilized, as may be seen in the scanning electron micrographs shown in figure11a-c. The characteristic plate like crystals are clearly seen in figure 11b.The rates, calculated from the titrants addition rates and normalized per unit geometric surface area of the exposed

σ

(7)

computer controlled syringe pumps.

valves, were fitted to the semi empirical equation:

() *<sup>n</sup> R kfS g g* =

where *kg* is the rate constant for crystal growth, *f(S)* a function of the total number of the growth sites available and *n* the apparent order of the crystal growth process. Logarithmic plots according to eq.7 yielded the kinetics shown in figure 12. The practical implications of this finding is that the deposition of the mineral phase on the membrane matrix is controlled by the diffusion of the growth units on the OCP nuclei been formed.

Fig. 11. Scanning electron micrographs of : (a) surface of the valves (b) OCP deposited on the valves at higher relative supersaturation and (c) OCP+HAP on valves at low relative supersaturation

Possible strategies aiming at the retardation of the calcification process should therefore rely on the alteration of the surfaces so as to make surface integration more difficult. An additional feature revealed by the kinetics plots at constant supersaturation (figure 12) is that the glutaraldehyde treated porcine valves are substrates favoring the mineral nucleation and growth.

Fig. 12. Logarithm of the rate of OCP formation on glutaraldeyde treated porcine valves (black circles) as a function of the relative solution supersaturation; pH 7.4, 37° C, 0.15M NaCl. Open squares and triangles refer to literature results. Experiments performed in our laboratory with the same methodology on fresh, untreated porcine valves have shown that these tissues failed to induce any formation of calcium phosphate deposits although they were kept in the mineralizing solution for as long as four days

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 77

generation biologic valve for patients at about 60 years of age derive improved life expectancy and event-free expectancy regardless of the need for concomitant coronary artery surgery (Birkmeyer et al., 2000). Of course, for special patient groups the indications should be changed. Patients who do need a long-term anticoagulation such as those with chronic atrial fibrillation, intracardiac thrombus, history of thromboembolic events, hypercoagulable state or low ejection fraction, should receive a mechanical valve regardless of age. In the contrary, patients with contraindication to anticoagulants, with bleeding disorders, women of child-bearing age, should receive a biologic valve (Bonow et al., 2006). Patients with chronic renal failure have a higher risk of earlier bioprosthetic valve degeneration, and also an increased incidence of anticoagulation-related complications. For that reason, the current ACC/AHA guidelines (2006) do not recommend the routine use of a

Artificial devices designed and manufactured for mid and long term implantation in patients has to satisfy quality criteria of biocompatibility and function during all implantation time expected. This time period is varied, from a few months, for temporary used prostheses, like some orthopaedic fixation plaques and screws, to long life function, as in the case of prosthetic heart valves. Despite of their evolution and future trends, even if medical technology could make implants satisfying that criteria, prosthetic devices made of, in its best, biologically inert biomaterials cannot meet a serious clinical problem: they cannot follow changes in patient's body from the time of implantation to end of their expected life. In other words, they cannot grow up and remodelled with patient. Tissue engineering is a recent technological approach in the construction of artificial implants that can be gradually remodelled into the patient in real living tissue and organs, following regeneration and auto repair capabilities similar to that of the other natural patient body components (Kretlow & Mikos, 2008; Zilla et al., 2008). Attempts for the construction of TE implants are spread to different tissues and organs, like dermal parches, cartilage, bone and cardiovascular implants and TE or hybrid organs like pancreas or liver. Design and construction of cardiovascular TE implants, like heart valves and blood vessels, is still a challenge because

As a general rule TE valves composed of two groups of biomaterials. One group is composing the scaffold, a structure having the morphology of natural heart valves, usually a biodegradable flexible composite synthetic membrane of a polymeric fibber network embedded in amorphous organic matrix. Different structure of valve parts like valve ring, wall stent and leaflets give to the synthetic valve mechanical strength and flexibility identical to function like natural heart valves. However, scaffolds have a temporary role as, in addition and in parallel with their normal physiological function as heart valves, they may have the ability to support cell adhesion in their structure. Different cells, like fibroblasts, smooth muscle cells and endothelia may adhere, proliferate, stimulated and function into the scaffold valve structure to produce different tissue components that will synergy to compose suitable valvular tissue. Some of these valve cells are transported into scaffold material prior implantation, followed by in vitro cell culture and parallel mechanical valve function in special designed devices, bioreactors. A hybrid structure of synthetic and living biomaterials is made in vitro, which by implantation in the living organism is expected to continue remodelling into real living tissue and organ. As the final

mechanical prosthesis (Bonow et al., 2006).

**6. The future: Tissue engineering (TE)** 

of numerous worldwide needs and the severity of possible failure.

Comparison with data obtained for OCP synthetic crystals (Tomazic et al., 1989) and collagen of various types (Heughebaert & Nancollas, 1984; Combes, 1996) showed that glutaraldehyde treated porcine valves accelerated the formation of OCP. The significantly larger value of the rate constant suggested that in this case the number of active sites is not possibly a function of the available surface area. Structural factors on a molecular level should also be considered. Calcium phosphate crystal growth and crystal phases transformations obtained from the results of our in vitro model have been well correlated with similar results obtained from microscopic examinations, chemical analysis and spectrophotometric characterization of crystal phases on samples from calcified natural and bioprosthetic heart valves (Mikroulis et al., 2002). From the composition, the morphology and the size of the developed crystals their nature was determined by comparison with reference synthetic calcium phosphate phases. With this technique it was possible to determine the morphology of CDs developed at the internal sites of the tissue at high magnifications. It seems that in cases of natural heart valves the CDs are mixture of HAP (Ca:PO4=1.67) and OCP (Ca:PO4=1.33), while in bioprosthetic CDs the percentage of OCP is higher than of native in which the ratio Ca:PO4 (1,82) is close to the Ca:PO4 composition of mature physiological biomineral in bone (1,75).

As an overall conclusion from combined studies examined CDs from calcified valve leaflets in vitro, in animal models and in vivo, a model of development of calcification by crystal growth through the formation of precursor phases, which are gradually hydrolyzed in smaller in size, thermodynamically more stable crystal formations may be introduced. According to this, initiation of calcification may be supposed to take place in sites of heterogenous nucleation, formed in different tissue deficiencies, together with local changes in already highly supersaturated body fluids. This model can be very useful in the introduction of anticalcification therapies or techniques for better biomaterials.

#### **5. Conclusion**

Heart valve calcification is still a serious complication for a great number of patients, especially in economical active ages and the elderly. Although anticalcification therapies and procedures have been introduced for valve repair, valve replacement, especially that of aortic and mitral valves is the last choice. Unfortunately, till today there is no "ideal" aortic valve prosthesis. The latter would be easy to be implanted, possess long-term durability, would have no thrombogenicity, maximum effective orifice area, without haemolysis, "resistant" to infective endocarditis, and produce minima noise (Birkmeyer et al., 2000). Currently, available options for the patient include mechanical valve, stented or stentless biologic heterograft valves, allograft valves and pulmonary autograft valves. For the selection between mechanical and biologic valve the surgeon should balance the risks and benefits of each model. The mechanical valve has a long-term durability (till 35-40 years), but on the other hand its thrombogenicity is high, 2-4% per year. In addition, the administered anticoagulation has a significantly increased risk of bleeding. The biologic valve has an increased risk of degeneration, as its durability lasts not more than 10-12 years after implantation (Siddiqui et al., 2009). After this time, a significant regurgitation demands its replacement with an increased operative mortality in comparison to the initial implantation (about two-fold higher). On the other hand, the thrombogenicity of the biologic valve is lower than of mechanical ones, about 1% per year (Puvimanasinghe et al., 2003). Analyses based on mathematical models of data suggest that the selection of 3rd-

Comparison with data obtained for OCP synthetic crystals (Tomazic et al., 1989) and collagen of various types (Heughebaert & Nancollas, 1984; Combes, 1996) showed that glutaraldehyde treated porcine valves accelerated the formation of OCP. The significantly larger value of the rate constant suggested that in this case the number of active sites is not possibly a function of the available surface area. Structural factors on a molecular level should also be considered. Calcium phosphate crystal growth and crystal phases transformations obtained from the results of our in vitro model have been well correlated with similar results obtained from microscopic examinations, chemical analysis and spectrophotometric characterization of crystal phases on samples from calcified natural and bioprosthetic heart valves (Mikroulis et al., 2002). From the composition, the morphology and the size of the developed crystals their nature was determined by comparison with reference synthetic calcium phosphate phases. With this technique it was possible to determine the morphology of CDs developed at the internal sites of the tissue at high magnifications. It seems that in cases of natural heart valves the CDs are mixture of HAP (Ca:PO4=1.67) and OCP (Ca:PO4=1.33), while in bioprosthetic CDs the percentage of OCP is higher than of native in which the ratio Ca:PO4 (1,82) is close to the Ca:PO4 composition of

As an overall conclusion from combined studies examined CDs from calcified valve leaflets in vitro, in animal models and in vivo, a model of development of calcification by crystal growth through the formation of precursor phases, which are gradually hydrolyzed in smaller in size, thermodynamically more stable crystal formations may be introduced. According to this, initiation of calcification may be supposed to take place in sites of heterogenous nucleation, formed in different tissue deficiencies, together with local changes in already highly supersaturated body fluids. This model can be very useful in the

Heart valve calcification is still a serious complication for a great number of patients, especially in economical active ages and the elderly. Although anticalcification therapies and procedures have been introduced for valve repair, valve replacement, especially that of aortic and mitral valves is the last choice. Unfortunately, till today there is no "ideal" aortic valve prosthesis. The latter would be easy to be implanted, possess long-term durability, would have no thrombogenicity, maximum effective orifice area, without haemolysis, "resistant" to infective endocarditis, and produce minima noise (Birkmeyer et al., 2000). Currently, available options for the patient include mechanical valve, stented or stentless biologic heterograft valves, allograft valves and pulmonary autograft valves. For the selection between mechanical and biologic valve the surgeon should balance the risks and benefits of each model. The mechanical valve has a long-term durability (till 35-40 years), but on the other hand its thrombogenicity is high, 2-4% per year. In addition, the administered anticoagulation has a significantly increased risk of bleeding. The biologic valve has an increased risk of degeneration, as its durability lasts not more than 10-12 years after implantation (Siddiqui et al., 2009). After this time, a significant regurgitation demands its replacement with an increased operative mortality in comparison to the initial implantation (about two-fold higher). On the other hand, the thrombogenicity of the biologic valve is lower than of mechanical ones, about 1% per year (Puvimanasinghe et al., 2003). Analyses based on mathematical models of data suggest that the selection of 3rd-

introduction of anticalcification therapies or techniques for better biomaterials.

mature physiological biomineral in bone (1,75).

**5. Conclusion** 

generation biologic valve for patients at about 60 years of age derive improved life expectancy and event-free expectancy regardless of the need for concomitant coronary artery surgery (Birkmeyer et al., 2000). Of course, for special patient groups the indications should be changed. Patients who do need a long-term anticoagulation such as those with chronic atrial fibrillation, intracardiac thrombus, history of thromboembolic events, hypercoagulable state or low ejection fraction, should receive a mechanical valve regardless of age. In the contrary, patients with contraindication to anticoagulants, with bleeding disorders, women of child-bearing age, should receive a biologic valve (Bonow et al., 2006). Patients with chronic renal failure have a higher risk of earlier bioprosthetic valve degeneration, and also an increased incidence of anticoagulation-related complications. For that reason, the current ACC/AHA guidelines (2006) do not recommend the routine use of a mechanical prosthesis (Bonow et al., 2006).

## **6. The future: Tissue engineering (TE)**

Artificial devices designed and manufactured for mid and long term implantation in patients has to satisfy quality criteria of biocompatibility and function during all implantation time expected. This time period is varied, from a few months, for temporary used prostheses, like some orthopaedic fixation plaques and screws, to long life function, as in the case of prosthetic heart valves. Despite of their evolution and future trends, even if medical technology could make implants satisfying that criteria, prosthetic devices made of, in its best, biologically inert biomaterials cannot meet a serious clinical problem: they cannot follow changes in patient's body from the time of implantation to end of their expected life. In other words, they cannot grow up and remodelled with patient. Tissue engineering is a recent technological approach in the construction of artificial implants that can be gradually remodelled into the patient in real living tissue and organs, following regeneration and auto repair capabilities similar to that of the other natural patient body components (Kretlow & Mikos, 2008; Zilla et al., 2008). Attempts for the construction of TE implants are spread to different tissues and organs, like dermal parches, cartilage, bone and cardiovascular implants and TE or hybrid organs like pancreas or liver. Design and construction of cardiovascular TE implants, like heart valves and blood vessels, is still a challenge because of numerous worldwide needs and the severity of possible failure.

As a general rule TE valves composed of two groups of biomaterials. One group is composing the scaffold, a structure having the morphology of natural heart valves, usually a biodegradable flexible composite synthetic membrane of a polymeric fibber network embedded in amorphous organic matrix. Different structure of valve parts like valve ring, wall stent and leaflets give to the synthetic valve mechanical strength and flexibility identical to function like natural heart valves. However, scaffolds have a temporary role as, in addition and in parallel with their normal physiological function as heart valves, they may have the ability to support cell adhesion in their structure. Different cells, like fibroblasts, smooth muscle cells and endothelia may adhere, proliferate, stimulated and function into the scaffold valve structure to produce different tissue components that will synergy to compose suitable valvular tissue. Some of these valve cells are transported into scaffold material prior implantation, followed by in vitro cell culture and parallel mechanical valve function in special designed devices, bioreactors. A hybrid structure of synthetic and living biomaterials is made in vitro, which by implantation in the living organism is expected to continue remodelling into real living tissue and organ. As the final

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 79

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Understanding of cell-biomaterial-biomechanics interaction needs a multidisciplinary synergism in order to result in successful TE valve, avoiding possible future undesirable side effects, like valve failure or carcinogenesis.

#### **7. References**


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**Part 3** 

**Aortic Root Replacement** 

