**Prosthetic Aortic Valves: A Surgical and Bioengineering Approach**

Dimosthenis Mavrilas1, Efstratios Apostolakis2 and Petros Koutsoukos3

*1Mechanical Engineering & Aer/tics, University of Patras, 2School of Medicine, University of Ioannina, 3Chemical Engineering, University of Patras, Greece* 

#### **1. Introduction**

56 Aortic Valve Surgery

Tseng E, Lee C, Cameron DE et al. (1997). Aortic valve replacement in the elderly: risk

Urban M, Pirk J, Turek D & Netuka I. (2011). In patients with concomitant aortic and mitral

Valfrè C, Rizzoli G, Zussa C, et al. (2006). Clinical results of Hancock II versus Hancock Standard at long-term follow-up. *J Thorac Cardiovasc Surg* 132:595-601. Vidaillet H, Granada JF, Chyou PH, et al. (2002) A populationbased study of mortality among patients with atrial fibrillation or flutter. *Am J Med* 113:365–370. Walther T, Falk V, Langebartels G, Kruger M, et al. (1999). Prospectively randomized

regression of left ventricular hypertrophy. *Circulation* 100(suppl 19):II6 –10. Walther T, Simon P, Dewey T, et al. (2007). Transapical minimally invasive aortic valve implantation: multicenter experience. *Circulation* 116(suppl 11):I-240–I-245. Waszyrowski T, Kasprzak JD, Krzeminska-Pakula M, Dziatkowiak A & Zaslonka J. (1997).

mechanical prosthesis: 8-year follow-up study. *Clin Cardiol* 20:843– 848. Weiss BM, von Segesser LK, Alon E, et al. (1998). Outcome of cardiovasc surgery and

Webb JG, Pasupati S, Humphries K, et al. (2007). Percutaneous transarterial aortic valve

Westaby S, Horton M, Jin XY, et al. (2000). Survival advantage of stentless aortic

Wolf PA, Abbott RD & Kannel WB. (1991). Atrial fibrillation as an independent risk factor

Wos S, Jasinski M, Bachowski R, et al. (1996). Results of mechanical prosthetic valve replacement in active valcular endocarditis. *J Cardiovasc Surg (Torino)* 37:29-32. Yacoub M, Rasmi NR, Sundt TM, et al. (1995). Fourteen-year experience with homovital homografts for aortic valve replacement. *J Thorac Cardiovasc Surg* 110:186 –193. Ye J, Cheung A, Lichtenstein SV, et al. (2007). Six-month outcome of transapical

Yinon Y, Siu SC, Warshafsky C, et al. (2009). Silversides CK. Use of low molecular weight

Yum Kl, Miller DC, Moore KA, et al. (1995). Durability of the Hancock MO bioprostheses

double valve replacement? *Interact Cardiovasc Thorac Surg* 12:238-242. Vahasilta T, Saraste A, Kito V, et al. (2005). Malmberg M, Kiss J, Kentala E, Kallojoki M,

valve disease is aortic valve replacement with mitral valve repair superior to

Savunen T. Cardiomyocyte apoptosis after antegrade and retrograde cardioplegia.

evaluation of stentless versus conventional biological aortic valves: impact on early

Early and long-term outcome of aortic valve replacement with homograft vs

pregnancy: a systematic review of the period 1984–1996. *Am J Obstet Gynecol*

replacement in selected high-risk patients with aortic stenosis. *Circulation* 116:755–

transcatheter aortic valve implantation in the initial seven patients. *Eur J* 

heparin in pregnant women with mechanical heart valves. *Am J Cardiol*

compared with the standard aortic valve bioprostheses. *Ann Thorac Surg* 60

factors and long term results. Ann Surg 225:793– 804.

*Ann Thorac Surg* 80:2229–2234.

179:1643–1653.

bioprostheses. *Ann Thorac Surg* 70:785-791.

*Cardiothorac Surg* 31(1):16-21.

1;104(9):1259-1263.

Suppl:S221– S228.

for stroke: The Framingham Study. *Stroke* 22:983-988.

763.

The need for replacement of damaged or malfunctioning organs or tissues in the human body has led through intense and innovative research during the past century to the development of materials and devices which are compatible with living tissues. Materials' compatibility consists in their capability to be accepted by the body when implanted and in contact with other tissues and body fluids. These materials, known as biomaterials, are the fundamental tools for engineering implantable devices dedicated to function in a specific way that substitutes corresponding function of the tissues or organs replaced due to malfunction, in synergism with the surrounding biological environment. Bone fracture healing by the incorporation of plates was known as early as the beginning of the century. Implants to replace heart valves and hip joints have been reported in the early 60s (Park & Lakes, 1992). Among the problems recorded when the first implants were employed were corrosion, mechanical failure and rejection by the body. The latter remains the main problem in the development of novel biomaterials which can be used as implants. Biomaterials now play a major role in replacing or improving the function of every major body system (skeletal, circulatory, nervous, etc.). Commonly employed implants include orthopedic devices such as total knee and hip joint replacements, spinal implants, and bone fixators; cardiac implants such as artificial heart valves and pacemakers; soft tissue implants such as breast implants and injectable collagen for soft tissue augmentation; and dental implants to replace teeth/root systems and bony tissue in the oral cavity.

When a man-made material is placed in the human body, tissue reacts to the implant in a variety of ways depending on the material type and function. The mechanism of tissue attachment depends on the tissue response to the implant surface. In general, materials can be placed into three classes that represent the tissue response they elicit: inert, bioresorbable, and bioactive. **Inert materials** such as titanium and alumina (Al2O3) are nearly chemically inert in the body and exhibit minimal chemical interaction with adjacent tissue. A fibrous tissue capsule will normally form around inert implants. Tissue attachment with inert materials can be through tissue growth into surface irregularities, by bone cement, or by press fitting into a defect. This morphological fixation is not ideal for the long-term stability of permanent implants and often becomes a problem with orthopedic and dental implant applications. **Bioresorbable materials**, such as tricalcium phosphate and polylactic-

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 59

the leaflet). Leaflets are multilaminate composite tissue structures of 3 layers (Figure 2.): the ventricularis, composed of elastin- rich fibers aligned in a radial direction, perpendicular to the leaflet margin, the fibrosa, on the aortic side of the leaflet, comprising primarily fibroblasts and collagen fibers arranged circumferentially, parallel to the leaflet margin and the spongiosa, a layer of loose connective tissue at the base of the leaflet, between the fibrosa and ventricularis, composed of fibroblasts, mesenchymal cells, and a glycosaminoglycan (GAG) rich organic matrix. This composite tissue structure provides tensile strength and pliability to the leaflets for decades of repetitive motion per minute (Freeman & Otto, 2005).

Fig. 3. Schematic representation of the three aortic valve commissures in natural and open

The commissures (figure 3) form tall, peaked spaces between the attachments of neighbouring cusps, and reach the so-called aortic sino-tubular junction. The latter is a ridge, called also "supraortic ridge" that separates the sinus and tubular portions of the ascending aorta (Malouf et al., 2008). The commissure between the right and non-coronary (or posterior) aortic leaflet overlies the membranous septum and corresponds to that laid between the anterior and septal leaflets of the tricuspid valve (Malouf et al., 2008). The commissure between the right and left aortic leaflet contacts its corresponding pulmonary ones and overlies the infundibular septum. Finally, the intervalvular fibrosa, at the commissure between the left and non-coronary aortic leaflet, fuses the aortic valve to the

During left ventricular systole, the systolic pressure inside rises exceeding the aortic pressure and the aortic valve is passively opened. Blood, ejected by the left ventricle (LV) pushes the aortic cusps upward and away from the centre of the aortic lumen. During this phase of cardiac cycle, the major opening diameter of the aortic valve is about equal to that of the ascending aorta at the level of sino-tubular junction (Stewart et al., 1998). In fact, the relatively inaccurate measurements of LV and intraaortic pressure during left ventricular catheterization show that there is no clinically significant gradient across the normal aortic valve. This measurement is curious because it suggests that blood does not travel down a pressure gradient during ejection. If flow out of the LV and through the outflow tract and the aortic valve was simply frictional, then a pressure gradient should needed between the LV and ascending aorta to obtain blood flow, according to the low of the physics: Q(flow)=Pressure Gradient/Resistance. However, this assumption does not hold when considering pulsatile blood flow in large vessels such as the aorta, since the inertia and

The aortic surface of each leaflet is rougher than its ventricular surface.

configuration (acknowledged art work by G. Athanassiou)

anterior mitral leaflet (Edwards, 1996; Malouf et al., 2008).

**1.2 Physiology of the normal aortic valve** 

polyglycolic acid copolymers, are designed to be slowly degraded under the biologicalbiochemical action of the living organism in bioresorbable products (like water, ions of electrolytes or CO2), replaced by living tissue (such as bone or soft tissues in tissue engineering) or liberating drugs, as in drug-delivery applications. **Bioactive materials** bond to surrounding tissues (like bone or soft tissues) through a time-dependent, kinetic modification of the surface triggered by their contact and function after implantation with parts of living organism. In particular, ion-exchange reactions or ion incorporation into the crystal lattice between the bioactive implant and the surrounding body fluids results in the formation of a biologically active interface layer on the implant surface responsible for the relatively strong interfacial bonding.

#### **1.1 Anatomy of the normal aortic valve**

The aortic valve (figure 1) is composed of three components: the annulus, the cusps or leaflets and the commissures. The annulus of the valve, in contrast to this of atrioventricular valves, does not located at the same level. Here, the annulus consists of ventriculo-arterial junction and is oriented in a curvilinear, semilunar fashion. It consists of three almost semicircular dense fibrous collagen structures forming three scallops, the whole encircling the ventriculo-aortic junction like a coronet. The three aortic leaflets are folds of endocardium with a central lamina fibrosa, which is locally thickened. Each leaflet is attached to the aortic wall (its upper part) and to the left ventricle (its lowest part or nadir of

Fig. 1. Anatomy of aortic valve construct. a. View from aorta & b. from left ventricle. c: View from aorta, showing leaflet junction, sinus of Valsalva and coronary ostium. Pictures are from a porcine aortic valve, of similar anatomy with human aortic valve

Fig. 2. Histological sections (a and b) and SEM microphotograph (c) of valvular leaflet tissue, demonstrating its multilaminate, 3D fiber reinforced composite structure, make it suitable for the complicated leaflet movements during valve function. F: Fibrosa, S: Spongiosa & V: Ventricularis

polyglycolic acid copolymers, are designed to be slowly degraded under the biologicalbiochemical action of the living organism in bioresorbable products (like water, ions of electrolytes or CO2), replaced by living tissue (such as bone or soft tissues in tissue engineering) or liberating drugs, as in drug-delivery applications. **Bioactive materials** bond to surrounding tissues (like bone or soft tissues) through a time-dependent, kinetic modification of the surface triggered by their contact and function after implantation with parts of living organism. In particular, ion-exchange reactions or ion incorporation into the crystal lattice between the bioactive implant and the surrounding body fluids results in the formation of a biologically active interface layer on the implant surface responsible for the

The aortic valve (figure 1) is composed of three components: the annulus, the cusps or leaflets and the commissures. The annulus of the valve, in contrast to this of atrioventricular valves, does not located at the same level. Here, the annulus consists of ventriculo-arterial junction and is oriented in a curvilinear, semilunar fashion. It consists of three almost semicircular dense fibrous collagen structures forming three scallops, the whole encircling the ventriculo-aortic junction like a coronet. The three aortic leaflets are folds of endocardium with a central lamina fibrosa, which is locally thickened. Each leaflet is attached to the aortic wall (its upper part) and to the left ventricle (its lowest part or nadir of

Fig. 1. Anatomy of aortic valve construct. a. View from aorta & b. from left ventricle. c: View from aorta, showing leaflet junction, sinus of Valsalva and coronary ostium. Pictures are

b c

V

Fig. 2. Histological sections (a and b) and SEM microphotograph (c) of valvular leaflet tissue, demonstrating its multilaminate, 3D fiber reinforced composite structure, make it suitable for the complicated leaflet movements during valve function. F: Fibrosa, S: Spongiosa & V:

S

F

from a porcine aortic valve, of similar anatomy with human aortic valve

a b c

relatively strong interfacial bonding.

Ventricularis

a

**1.1 Anatomy of the normal aortic valve** 

the leaflet). Leaflets are multilaminate composite tissue structures of 3 layers (Figure 2.): the ventricularis, composed of elastin- rich fibers aligned in a radial direction, perpendicular to the leaflet margin, the fibrosa, on the aortic side of the leaflet, comprising primarily fibroblasts and collagen fibers arranged circumferentially, parallel to the leaflet margin and the spongiosa, a layer of loose connective tissue at the base of the leaflet, between the fibrosa and ventricularis, composed of fibroblasts, mesenchymal cells, and a glycosaminoglycan (GAG) rich organic matrix. This composite tissue structure provides tensile strength and pliability to the leaflets for decades of repetitive motion per minute (Freeman & Otto, 2005). The aortic surface of each leaflet is rougher than its ventricular surface.

Fig. 3. Schematic representation of the three aortic valve commissures in natural and open configuration (acknowledged art work by G. Athanassiou)

The commissures (figure 3) form tall, peaked spaces between the attachments of neighbouring cusps, and reach the so-called aortic sino-tubular junction. The latter is a ridge, called also "supraortic ridge" that separates the sinus and tubular portions of the ascending aorta (Malouf et al., 2008). The commissure between the right and non-coronary (or posterior) aortic leaflet overlies the membranous septum and corresponds to that laid between the anterior and septal leaflets of the tricuspid valve (Malouf et al., 2008). The commissure between the right and left aortic leaflet contacts its corresponding pulmonary ones and overlies the infundibular septum. Finally, the intervalvular fibrosa, at the commissure between the left and non-coronary aortic leaflet, fuses the aortic valve to the anterior mitral leaflet (Edwards, 1996; Malouf et al., 2008).

#### **1.2 Physiology of the normal aortic valve**

During left ventricular systole, the systolic pressure inside rises exceeding the aortic pressure and the aortic valve is passively opened. Blood, ejected by the left ventricle (LV) pushes the aortic cusps upward and away from the centre of the aortic lumen. During this phase of cardiac cycle, the major opening diameter of the aortic valve is about equal to that of the ascending aorta at the level of sino-tubular junction (Stewart et al., 1998). In fact, the relatively inaccurate measurements of LV and intraaortic pressure during left ventricular catheterization show that there is no clinically significant gradient across the normal aortic valve. This measurement is curious because it suggests that blood does not travel down a pressure gradient during ejection. If flow out of the LV and through the outflow tract and the aortic valve was simply frictional, then a pressure gradient should needed between the LV and ascending aorta to obtain blood flow, according to the low of the physics: Q(flow)=Pressure Gradient/Resistance. However, this assumption does not hold when considering pulsatile blood flow in large vessels such as the aorta, since the inertia and

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 61

aortic ridge increased and the surface area and mass of the aortic leaflets also increased (Hall & Julian, 1989). In the 6th and 7th decades the fibrosa of the leaflet begins to calcify, first at the point of attachment to the aortic wall, which is where maximum flexion occurs. This calcification may gradually extend throughout the valve, limiting the valve's opening. Occasionally, the calcified leaflets may develop local ulcerations and thrombus formations (Otto et al., 1999; Schwartz & Zipes, 2005). Of course, in the cases of the rheumatic disease the course of chronic inflammatory disease produce the above-mentioned changes (calcification, thickening of the leaflets, fusion of commissures, local ulcerations, sub-

In normal adults, the area of the aortic valve orifice is 2.6 to 3.5 cm2. Experimental studies have suggested that the aortic orifice must be reduced to approximately one quarter of this in order to diminish significantly the cardiac output. A reduction in this area to 1 cm2 is associated with a rise in left ventricular systolic pressure and a pressure drop across the aortic valve (Hall & Julian, 1989). Aortic stenosis is generally considered to be critical when the systolic pressure difference across the valve exceeds 50 mmHg in the presence of a normal cardiac output or if the effective aortic orifice is less than 0.4 cm2 (Hall & Julian, 1989). According to Rahimtoola, the aortic valve area has to be reduced by about 50% of normal before a measurable gradient can be demonstrated (Rahimtoola, 2004). When a pressure gradient develops between LV and aorta LV pressure is increased, ventricular wall stress increased contributing to development of myocardial hypertrophy and the LV function impairs. A diastolic dysfunction is caused of a combination of impaired myocardial relaxation during diastolic phase and increased myocardial stiffness (Hess et al., 1993). Patients with severe LV hypertrophy may exhibit LV diastolic dysfunction, which consequently may produce the syndrome of clinical heart failure with symptoms of paroxysmal nocturnal dyspnea, orthopnea or even pulmonary oedema, even if the systolic

Prospective studies on the rate of hemodynamic progression in patients diagnosed with aortic stenosis documented an increasing rate of aortic jet velocity, in average 0.3 m/s per year, with an increase in mean trans-aortic pressure gradient of 7 mm Hg per year and a decrease in aortic valve area of 0.1 cm2 per year (Brener et al., 1995; Faggiano et al., 1996; Freeman & Otto, 2005). During this later course, for the cases of symptomatic aortic stenosis or of the asymptomatic with significant (> 50 mmHg) trans-valvular mean gradient, surgical

There are four options for the management of the severe calcific aortic stenosis: the balloon aortic valvuloplasty, the "open" aortic valve commissurotomy, the percutaneous aortic

At **balloon aortic valvuloplasty** (BAV), a ballon-catheter is introduced after a femoral artery puncture and retrograde till the left ventricle (Diethrich, 1993; Smedira et al., 1993). Inflation of the balloon within the aortic orifice can stretch the calcified annulus, fracture calcified areas and dissect the fused commissures. Disadvantages of the method, as the risk of stroke and increase of pre-existent valve regurgitation (Cormier & Vahanian, 1992) are controversial to increase of effective orifice area. An overall 65% survival and 40% free of death or reoperation over 1-year survival has been reported (Davidson et al., 1990). However, no beneficial effect on long-term clinical outcome demonstrated due to significant residual obstruction from leaflet thickening and annular calcification, resulted in severe re-stenosis

endothelial atherosclerotic plaques, etc.) much earlier.

LV function is normal (Rahimtoola, 2004).

**1.4 The surgical management of the stenotic aortic valve** 

valve implantation, and the "classic" aortic valve replacement.

management is indicated.

momentum of the blood ejected from the LV is much more important than resistance. In fact, it has been experimentally shown that at the first 40% of the ejection phase, during blood acceleration, small pressure gradients (about 10 mmHg) are observed between LV and aorta (Hall & Julian, 1989). This small gradient persists for about 45% of the ejection after which the gradient reverses as forward blood flow is decelerated.

The fibrous wall of the sinuses of Valsalva (figure 1) at the nadirs of annular scallops is not extensible in contrast to their upper parts (at the level of commissures) where it produces its biggest increase of aortic radius, about 16% in the peak of systole, due to the fibro-elastic composition of the aortic wall (Williams et al., 1989). During this phase the commissures move apart, making the fully open orifice triangular, the free margins of the aortic leaflets becoming almost straight lines between commissures. However, they do not flatten against the sinus wall, which is an important factor for the subsequent valve closure (Williams et al., 1989). Most of the blood ejected during systole is directed to the ascending aorta while a small volume enters into the sinuses of Valsalva. Valve geometry in the sinus region produces vortical blood flow at systole, which helps to coronary perfusion, maintain the triangular "mid-position" of the leaflets and probably initiate their re-approximation at the end of systole. During diastole, the three aortic leaflets fall passively towards the centre of the aortic lumen and, under the pressure of the supravalvular blood column they hermetically contact each other along lines of coaptation. Therefore during diastole the three normal aortic leaflets, such as the yacht-sails, support the entire intraaortic blood column and prevent its partial regurgitation into the left ventricle. Experiments have shown that only 4% of blood ejected during systole regurgitates through the centre of the valve during diastole. In the absence of sinuses of Valsalva the regurgitant blood may be increased up to 23% (Williams et al., 1989).

#### **1.3 The diseased stenotic aortic valve**

Normal aortic valve histology and anatomy may be changed under pathologic conditions with corresponding alterations in its normal physiological function. Age-related changes in fibromuscular skeleton of the heart include myxomatous degeneration and collagen infiltration, called aortic valve sclerosis. This sclerosis is observed in as many as 30% of elderly people, namely in 25% of people 65 to 74 years of age and in 48% of people older than 84 years (Freeman & Otto, 2005; Otto et al., 1999; Stewart et al., 1997). Histopathologic studies of aortic sclerosis show focal subendothelial plaquelike lesions on the aortic side of the leaflet that extend to the adjacent fibrosa layer. Similarities to atherosclerosis are present in these lesions, with prominent accumulation of "atherogenic" lipoproteins, including LDL and lipoprotein(a), evidence of LDL oxidation, an inflammatory cell infiltrate and microscopic calcification (Olsson et al., 1999; Otto et al., 1999; Wallby et al, 2002). The initiation of these lesions is possibly due to increased mechanical or decreased shear stress, similar to that seen in early atherosclerotic lesions (Freeman & Otto, 2005). Of note, these changes are more prominent on the aortic surface of the leaflets where the mechanical stress of the aortic valve is highest, especially in the flexion area near the attachment to the aortic root. Shear stress across the endothelium of the non-coronary cusp is lower than the left and right coronary cusps because of the absence of diastolic coronary flow, which likely explains why the non-coronary cusp is often the first cusp affected (Freeman & Otto, 2005).

Other age-related changes become in the valve, in the aortic wall, as well as in the myocardium. The central nodules on the cusps and the closure lines become more prominent. At the same time the Valsalva's sinuses are stretched, the diameter of the supra-

momentum of the blood ejected from the LV is much more important than resistance. In fact, it has been experimentally shown that at the first 40% of the ejection phase, during blood acceleration, small pressure gradients (about 10 mmHg) are observed between LV and aorta (Hall & Julian, 1989). This small gradient persists for about 45% of the ejection

The fibrous wall of the sinuses of Valsalva (figure 1) at the nadirs of annular scallops is not extensible in contrast to their upper parts (at the level of commissures) where it produces its biggest increase of aortic radius, about 16% in the peak of systole, due to the fibro-elastic composition of the aortic wall (Williams et al., 1989). During this phase the commissures move apart, making the fully open orifice triangular, the free margins of the aortic leaflets becoming almost straight lines between commissures. However, they do not flatten against the sinus wall, which is an important factor for the subsequent valve closure (Williams et al., 1989). Most of the blood ejected during systole is directed to the ascending aorta while a small volume enters into the sinuses of Valsalva. Valve geometry in the sinus region produces vortical blood flow at systole, which helps to coronary perfusion, maintain the triangular "mid-position" of the leaflets and probably initiate their re-approximation at the end of systole. During diastole, the three aortic leaflets fall passively towards the centre of the aortic lumen and, under the pressure of the supravalvular blood column they hermetically contact each other along lines of coaptation. Therefore during diastole the three normal aortic leaflets, such as the yacht-sails, support the entire intraaortic blood column and prevent its partial regurgitation into the left ventricle. Experiments have shown that only 4% of blood ejected during systole regurgitates through the centre of the valve during diastole. In the absence of sinuses of Valsalva the regurgitant blood may be increased up to

Normal aortic valve histology and anatomy may be changed under pathologic conditions with corresponding alterations in its normal physiological function. Age-related changes in fibromuscular skeleton of the heart include myxomatous degeneration and collagen infiltration, called aortic valve sclerosis. This sclerosis is observed in as many as 30% of elderly people, namely in 25% of people 65 to 74 years of age and in 48% of people older than 84 years (Freeman & Otto, 2005; Otto et al., 1999; Stewart et al., 1997). Histopathologic studies of aortic sclerosis show focal subendothelial plaquelike lesions on the aortic side of the leaflet that extend to the adjacent fibrosa layer. Similarities to atherosclerosis are present in these lesions, with prominent accumulation of "atherogenic" lipoproteins, including LDL and lipoprotein(a), evidence of LDL oxidation, an inflammatory cell infiltrate and microscopic calcification (Olsson et al., 1999; Otto et al., 1999; Wallby et al, 2002). The initiation of these lesions is possibly due to increased mechanical or decreased shear stress, similar to that seen in early atherosclerotic lesions (Freeman & Otto, 2005). Of note, these changes are more prominent on the aortic surface of the leaflets where the mechanical stress of the aortic valve is highest, especially in the flexion area near the attachment to the aortic root. Shear stress across the endothelium of the non-coronary cusp is lower than the left and right coronary cusps because of the absence of diastolic coronary flow, which likely explains

why the non-coronary cusp is often the first cusp affected (Freeman & Otto, 2005).

Other age-related changes become in the valve, in the aortic wall, as well as in the myocardium. The central nodules on the cusps and the closure lines become more prominent. At the same time the Valsalva's sinuses are stretched, the diameter of the supra-

after which the gradient reverses as forward blood flow is decelerated.

23% (Williams et al., 1989).

**1.3 The diseased stenotic aortic valve** 

aortic ridge increased and the surface area and mass of the aortic leaflets also increased (Hall & Julian, 1989). In the 6th and 7th decades the fibrosa of the leaflet begins to calcify, first at the point of attachment to the aortic wall, which is where maximum flexion occurs. This calcification may gradually extend throughout the valve, limiting the valve's opening. Occasionally, the calcified leaflets may develop local ulcerations and thrombus formations (Otto et al., 1999; Schwartz & Zipes, 2005). Of course, in the cases of the rheumatic disease the course of chronic inflammatory disease produce the above-mentioned changes (calcification, thickening of the leaflets, fusion of commissures, local ulcerations, subendothelial atherosclerotic plaques, etc.) much earlier.

In normal adults, the area of the aortic valve orifice is 2.6 to 3.5 cm2. Experimental studies have suggested that the aortic orifice must be reduced to approximately one quarter of this in order to diminish significantly the cardiac output. A reduction in this area to 1 cm2 is associated with a rise in left ventricular systolic pressure and a pressure drop across the aortic valve (Hall & Julian, 1989). Aortic stenosis is generally considered to be critical when the systolic pressure difference across the valve exceeds 50 mmHg in the presence of a normal cardiac output or if the effective aortic orifice is less than 0.4 cm2 (Hall & Julian, 1989). According to Rahimtoola, the aortic valve area has to be reduced by about 50% of normal before a measurable gradient can be demonstrated (Rahimtoola, 2004). When a pressure gradient develops between LV and aorta LV pressure is increased, ventricular wall stress increased contributing to development of myocardial hypertrophy and the LV function impairs. A diastolic dysfunction is caused of a combination of impaired myocardial relaxation during diastolic phase and increased myocardial stiffness (Hess et al., 1993). Patients with severe LV hypertrophy may exhibit LV diastolic dysfunction, which consequently may produce the syndrome of clinical heart failure with symptoms of paroxysmal nocturnal dyspnea, orthopnea or even pulmonary oedema, even if the systolic LV function is normal (Rahimtoola, 2004).

Prospective studies on the rate of hemodynamic progression in patients diagnosed with aortic stenosis documented an increasing rate of aortic jet velocity, in average 0.3 m/s per year, with an increase in mean trans-aortic pressure gradient of 7 mm Hg per year and a decrease in aortic valve area of 0.1 cm2 per year (Brener et al., 1995; Faggiano et al., 1996; Freeman & Otto, 2005). During this later course, for the cases of symptomatic aortic stenosis or of the asymptomatic with significant (> 50 mmHg) trans-valvular mean gradient, surgical management is indicated.

#### **1.4 The surgical management of the stenotic aortic valve**

There are four options for the management of the severe calcific aortic stenosis: the balloon aortic valvuloplasty, the "open" aortic valve commissurotomy, the percutaneous aortic valve implantation, and the "classic" aortic valve replacement.

At **balloon aortic valvuloplasty** (BAV), a ballon-catheter is introduced after a femoral artery puncture and retrograde till the left ventricle (Diethrich, 1993; Smedira et al., 1993). Inflation of the balloon within the aortic orifice can stretch the calcified annulus, fracture calcified areas and dissect the fused commissures. Disadvantages of the method, as the risk of stroke and increase of pre-existent valve regurgitation (Cormier & Vahanian, 1992) are controversial to increase of effective orifice area. An overall 65% survival and 40% free of death or reoperation over 1-year survival has been reported (Davidson et al., 1990). However, no beneficial effect on long-term clinical outcome demonstrated due to significant residual obstruction from leaflet thickening and annular calcification, resulted in severe re-stenosis

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 63

mechanical point of view, cardiac valves work as common one-way valves of hydraulic systems. Despite their simple working principle, the design and orthotopic surgical implantation of cardiac valves was not possible till the development of extracorporeal blood circulation devices, introduced by Gibbon (Gibbon, 1954), by which blood circulation was maintained during open heart surgical procedures. Two main types of prosthetic heart valves are available: Mechanical (MHV) and biological or bioprosthetic (BHV) heart valves. MHVs in general composed of two main parts: One non-moving, which is sutured properly in the anatomic region of the failed, surgically excised, natural valve, in the interior region of which a second, moving part is included, passively guided by pressure difference changes between the inlet and outlet regions around it. The main difference is on the type of the moving part (occluder): The ball type and the disk type valves. For each of the two basic constructions many different designs and materials were used and a great branch of technology was developed. In a near parallel approach, BHVs, based on mimicking the design and function of natural heart valves were developed and applied. Although they simply were made of animal derived valves of proper size and structure, like porcine aortic valves, after biochemical treatment for removing antigenic factors, different designs have been introduced nowadays using combinations of artificial and natural derived

Different surgical operations were approached before extracorporeal circulation by implantation of artificial valve designs in peripheral vessels, like that of Hufnagel & Harvey in 1952 (Hufnagel & Harvey, 1953), when an aortic valve, made by a combination of biologically inert materials (a lucite® tube-like design with a mobile spherical poppet inside) was implanted in the descending thoracic aorta of a patient with a significant aortic insufficiency. However, it was at end of 50s when a caged ball type MHV was introduced. This design, in its final appearance, consists of a metallic ring with a soft material in its perimeter for stable suturing on surrounding soft tissues without blood leakage through the suturing line. A cage design, usually a three metal struts, is welded in the ring into which a ball made of silicon or other polymeric material is moved from the closed position, where it is pushed in touch with the ring, to the open once where it is attached the top of the cage (figure 4a). Starr and co-workers began in 1960 (their first report appeared in 1963) (Starr et al., 1963) to implant the caged ball aortic prosthesis in the orthotopic position with many of these prostheses remained well functioning for up to 40 years (Shiono et al., 2005). Major problems with these initial so-called "ball valves" were the compromised hemodynamic performance (small effective orifice area, big size of the sewing ring, turbulent flow) and the thromboembolic complications (high-grade haemolysis, thrombosis). For the last reason, all these valves required intense anticoagulation therapy

To find a solution in these problems, after a substitution of ball with a disk type occluder (caged disk valves) to achieve less moving mass and reducing the valve height inside the aortic root, the second-generation of prosthetic valves, the so-called "tilting-disk valve", was developed in 1968 (Emery et al., 2008) (figure 4b). Tilting disk valves were the result of evolution in MHV technology towards reduction of whole volume, occluder mass and surface area projected vertical to the blood flow axis, maximizing of opening angle of the disk and designed the disk shape so as to approach a near physiological central flow velocity profile. Haemocompatibily was also improved by a minimization of the blood

biomaterials.

(Ezekowitz, 2002).

**2.1 Evolution of mechanical heart valve prostheses** 

typically occurred within months (Bonow et al., 1998; Cormier & Vahanian, 1992; Freeman & Otto, 2005; Smedira et al., 1993).

The **"open" aortic valvulotomy**, performed rarely, usually during another open heart operation, is based on the -by a scalpel- commissurotomy of the fused commissure (-s). Because of the excessive calcification and rigidity of the leaflets, a central postoperative insufficiency is anticipated. The main indication is the case of congenital aortic stenosis with one or two congenitally fused commissures. In fact, in young patients, if the valve is pliable, mobile, and free of calcification, simple commissurotomy may be feasible. The operative mortality in these cases does not exceed 1% (Rahimtoola, 2004).

The impetus for the development of **percutaneous or transcatheter aortic valve replacement** (PAVI or TAVI) lies in the need for an intervention that is more durable than balloon aortic valvuloplasty and that can be used in patients who are too risk for the "classic" aortic valve replacement. The basic concept is based on the use of an outer expandable stent (scaffold) to resist the rigidity of the calcified aortic annulus and native leaflets (Davidson & Baim, 2008). In the inner surface of this stent three appropriately prepared pericardial or porcine leaflets are fixed constituting –after full expansion of the stent- a well functioning valvular prosthesis. The first implantation in the human being was done in France since 1992 by Cribier et al. (Cribier et al., 2002). The introduction of the catheter bearing the valved-stent requires direct femoral or iliac artery access, while in a few cases with stenotic iliac arteries the catheter is introduced through the apex of the left ventricle (transapical introduction), after a left anterior thoracotomy. The results of this method after more than 10 years of application are encouraged. Procedural mortality is 2-3%, one-month survival about 88% and the 1-year is ranged 65 to 78% (Bosmans et al., 2011).

Finally, **the aortic valve replacement** is the "classical" surgical treatment of calcific aortic stenosis, especially for the elderly. During this method, the three calcified leaflets of the valve are resected, as well as the calcific deposits of the annulus. Then, sutures are passed circumferentially through the annulus and the sewing ring of the prosthesis. Finally, the sutures are tied down in the native annulus. Aortic valve replacement by using 3rdgeneration prosthetic valves –mechanical or biological- obtains excellent early and late outcomes with low mortality and morbidity. Recent surgical series report operative mortality rates for aortic valve replacement as low as 1%, increasing to 9% in higher-risk patients. Long-term survival after valve replacement is 80% at 3 years, with an age-corrected survival postoperatively that is nearly normalized (Freeman & Otto, 2005; Rahimtoola, 2010). Significant postoperative morbidity, such as thromboembolism, hemorrhagic complications from anticoagulation, prosthetic valve dysfunction, and endocarditis, are rare and occur at a rate of 2% to 3% per year (Rahimtoola, 2010).

## **2. Aortic valve prostheses: The parallel evolution of mechanical & bioprosthetic valves**

Cardiac valve prostheses are devices designed and constructed properly to assure unidirectional blood flow. Like the natural cardiac valves, they work passively, due to pressure difference across their structure, with parts able to be moved between two positions: open position, when blood circulates, as in the case of aortic valves from left ventricle to the aorta during myocardial contraction (systolic phase) and closed position, when blood circulation stops (during the diastolic phase for the aortic valve). From a

typically occurred within months (Bonow et al., 1998; Cormier & Vahanian, 1992; Freeman

The **"open" aortic valvulotomy**, performed rarely, usually during another open heart operation, is based on the -by a scalpel- commissurotomy of the fused commissure (-s). Because of the excessive calcification and rigidity of the leaflets, a central postoperative insufficiency is anticipated. The main indication is the case of congenital aortic stenosis with one or two congenitally fused commissures. In fact, in young patients, if the valve is pliable, mobile, and free of calcification, simple commissurotomy may be feasible. The operative

The impetus for the development of **percutaneous or transcatheter aortic valve replacement** (PAVI or TAVI) lies in the need for an intervention that is more durable than balloon aortic valvuloplasty and that can be used in patients who are too risk for the "classic" aortic valve replacement. The basic concept is based on the use of an outer expandable stent (scaffold) to resist the rigidity of the calcified aortic annulus and native leaflets (Davidson & Baim, 2008). In the inner surface of this stent three appropriately prepared pericardial or porcine leaflets are fixed constituting –after full expansion of the stent- a well functioning valvular prosthesis. The first implantation in the human being was done in France since 1992 by Cribier et al. (Cribier et al., 2002). The introduction of the catheter bearing the valved-stent requires direct femoral or iliac artery access, while in a few cases with stenotic iliac arteries the catheter is introduced through the apex of the left ventricle (transapical introduction), after a left anterior thoracotomy. The results of this method after more than 10 years of application are encouraged. Procedural mortality is 2-3%, one-month survival about 88% and the 1-year is ranged 65 to 78% (Bosmans et al.,

Finally, **the aortic valve replacement** is the "classical" surgical treatment of calcific aortic stenosis, especially for the elderly. During this method, the three calcified leaflets of the valve are resected, as well as the calcific deposits of the annulus. Then, sutures are passed circumferentially through the annulus and the sewing ring of the prosthesis. Finally, the sutures are tied down in the native annulus. Aortic valve replacement by using 3rdgeneration prosthetic valves –mechanical or biological- obtains excellent early and late outcomes with low mortality and morbidity. Recent surgical series report operative mortality rates for aortic valve replacement as low as 1%, increasing to 9% in higher-risk patients. Long-term survival after valve replacement is 80% at 3 years, with an age-corrected survival postoperatively that is nearly normalized (Freeman & Otto, 2005; Rahimtoola, 2010). Significant postoperative morbidity, such as thromboembolism, hemorrhagic complications from anticoagulation, prosthetic valve dysfunction, and endocarditis, are rare

mortality in these cases does not exceed 1% (Rahimtoola, 2004).

and occur at a rate of 2% to 3% per year (Rahimtoola, 2010).

**2. Aortic valve prostheses: The parallel evolution of mechanical &** 

Cardiac valve prostheses are devices designed and constructed properly to assure unidirectional blood flow. Like the natural cardiac valves, they work passively, due to pressure difference across their structure, with parts able to be moved between two positions: open position, when blood circulates, as in the case of aortic valves from left ventricle to the aorta during myocardial contraction (systolic phase) and closed position, when blood circulation stops (during the diastolic phase for the aortic valve). From a

& Otto, 2005; Smedira et al., 1993).

2011).

**bioprosthetic valves** 

mechanical point of view, cardiac valves work as common one-way valves of hydraulic systems. Despite their simple working principle, the design and orthotopic surgical implantation of cardiac valves was not possible till the development of extracorporeal blood circulation devices, introduced by Gibbon (Gibbon, 1954), by which blood circulation was maintained during open heart surgical procedures. Two main types of prosthetic heart valves are available: Mechanical (MHV) and biological or bioprosthetic (BHV) heart valves. MHVs in general composed of two main parts: One non-moving, which is sutured properly in the anatomic region of the failed, surgically excised, natural valve, in the interior region of which a second, moving part is included, passively guided by pressure difference changes between the inlet and outlet regions around it. The main difference is on the type of the moving part (occluder): The ball type and the disk type valves. For each of the two basic constructions many different designs and materials were used and a great branch of technology was developed. In a near parallel approach, BHVs, based on mimicking the design and function of natural heart valves were developed and applied. Although they simply were made of animal derived valves of proper size and structure, like porcine aortic valves, after biochemical treatment for removing antigenic factors, different designs have been introduced nowadays using combinations of artificial and natural derived biomaterials.

#### **2.1 Evolution of mechanical heart valve prostheses**

Different surgical operations were approached before extracorporeal circulation by implantation of artificial valve designs in peripheral vessels, like that of Hufnagel & Harvey in 1952 (Hufnagel & Harvey, 1953), when an aortic valve, made by a combination of biologically inert materials (a lucite® tube-like design with a mobile spherical poppet inside) was implanted in the descending thoracic aorta of a patient with a significant aortic insufficiency. However, it was at end of 50s when a caged ball type MHV was introduced. This design, in its final appearance, consists of a metallic ring with a soft material in its perimeter for stable suturing on surrounding soft tissues without blood leakage through the suturing line. A cage design, usually a three metal struts, is welded in the ring into which a ball made of silicon or other polymeric material is moved from the closed position, where it is pushed in touch with the ring, to the open once where it is attached the top of the cage (figure 4a). Starr and co-workers began in 1960 (their first report appeared in 1963) (Starr et al., 1963) to implant the caged ball aortic prosthesis in the orthotopic position with many of these prostheses remained well functioning for up to 40 years (Shiono et al., 2005). Major problems with these initial so-called "ball valves" were the compromised hemodynamic performance (small effective orifice area, big size of the sewing ring, turbulent flow) and the thromboembolic complications (high-grade haemolysis, thrombosis). For the last reason, all these valves required intense anticoagulation therapy (Ezekowitz, 2002).

To find a solution in these problems, after a substitution of ball with a disk type occluder (caged disk valves) to achieve less moving mass and reducing the valve height inside the aortic root, the second-generation of prosthetic valves, the so-called "tilting-disk valve", was developed in 1968 (Emery et al., 2008) (figure 4b). Tilting disk valves were the result of evolution in MHV technology towards reduction of whole volume, occluder mass and surface area projected vertical to the blood flow axis, maximizing of opening angle of the disk and designed the disk shape so as to approach a near physiological central flow velocity profile. Haemocompatibily was also improved by a minimization of the blood

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 65

al., 1962). The first orthotopic insertion of homograft valve was performed in 1962 by Barratt-Boyes (Barratt-Boyes, 1964). The introduction of other biological valves began in 1967 when Senning used pieces of fascia lata of the patient for replacement of the diseased

Xenograft or heterograft valves, animal derived heart valves or BHVs made from different animal derived tissues are alternatives, offering the advantage of been prepared much prior the operation, available in different sizes and designs. BHVs include a variety of heart valve replacement using as substitutes heart valves of different orientations and technologies. Among different types, porcine aortic and bovine pericardial xenograft BHVs has been established as valve substitutes. Porcine aortic valves, after a treatment for removing excess fatty and aortic wall tissue and part of septal myocardial tissue from valve leaflets are imposed in biochemical preparation, aimed in removing antigenic factors (valve cells) and stabilize the remaining acellular valve tissue against enzyme reactions by different chemical compounds. Formaldehyde was first used for porcine valve fixation (Angel, 1972), substituted later by glutaraldehyde because of its ability for double-edge cross linking of collagen molecules (Woodroof, 1972), resulted in better longevity of valves. The stabilized tissue valve is sutured in specially designed frames composed of aortic ring and three commissures. A metallic or polymeric frame is used as a skeleton, covered with biocompatible textiles (like Teflon® or Dacron®) onto which the valve tissue is sutured with

Improvement of BHV function was achieved with the use of different membranous soft tissues for the construction of valve leaflets and suturing them in similar artificial frames like that of the porcine valves. Percutaneous tissue was used for that scope; however, pericardial tissue from different animals was finally used alternatively. Pericardial tissue is a big membrane enclosing the heart. Its histology is similar to heart valve leaflets with respect of its composition of collagen and elastin fiber networks in different layers inside an amorhous organic matrix of glycosaminoglycans (GAGs), proteoglycanes (PGs), and other proteins. Extracellular water solution of electrolytes and soluble proteins compose 65-70% of the tissue mass weight. Cells, like fibroblast, epithelial, muscle cells and other types are present. Despite these similarities with other soft tissues, including heart valve leaflets, fibber structure of bovine pericardium is quite different. Fibber orientation in valve leaflets is specified for supporting their motion and strengthening mechanical stress developed during valve function. The different anatomic position and function of pericardial tissue resulted in a different fibber orientation, varied across its surface. For this reason special attention is given in selection criteria of specific regions from the whole pericardial membrane for better suitability to function as heart valve leaflets (Simionescu et al., 1993). The benefits from the use of pericardial tissue (especially bovine ones, which is the standard selection last years) instead of porcine valves were better opening area of the valve, coaptation of the valve leaflets and flexibility in design of valve anatomic configuration. However, similar problems with porcine BHVs, tissue deterioration and

The first xenograft valve, a stent-mounted porcine aortic valve was implanted by Binet et al. in 1965 (Binet et al., 1965), while the glutaraldehyde-preserved stent-mounted porcine valves were introduced by Carpentier et al., in 1967 (Carpentier et al., 1969). Over the past 40 years, advances in tissue fixation (bovine pericardium and porcine aortic valves) methodologies and chemical treatments to prevent calcification, have yielded improvements

aortic valve (Senning, 1967; Ionescu & Ross, 1969).

permanent sutures).

calcification still remain.

in the longevity of bioprostheses.

contact area, material coating with biocompatible compounds (like pyrolitic carbon) and appropriate disk morphology for smoother blood flow around it. The most usable models were the Björk-Shiley and the Medronic Hall type. The main problems with these valves were the rare rupture of metal strut supporting disk movement and subsequent embolisation by the disk, non-axial flow, even in the models of Björk-Shiley with disk opening angle of 72°. Due to the excessive turbulent flow through the two orifices of the valve (a small and a bigger), a high-grade of haemolysis in patients was reported. However, good long-term results characterized that type of MHV (Oxenham et al., 2003).

The 3rd generation of mechanical valves was appeared in 1977 with the introduction of the St. Jude Medical (SJM) bileaflet valve coated with pyrolitic carbon (Emery et al., 2008; Gott et al., 2003). Over the following decades, the dramatic step of bileaflet prostheses nearly obviated the use of all other kinds of mechanical valves in all over the world (Emery et al., 2008). In fact, the low-profile SJM valve demonstrated low rates of thomboembolism, low trans-valvular gradients, low grade of haemolysis and minimal valvular insufficiency (Chambers et al., 2005; Gott et al., 2003; Walther et al., 2000). Out of the SJM valve, several other 3rd generation models were introduced such as the ATS Medical Prosthesis, the Sulzer CarboMedics , the On-X prosthesis and the Sorin prosthetic valve, all of them with similar haemodynamic and clinical outcomes (Chambers et al., 2005; Walther et al., 2000). Since the introduction of this 3rd generation of the valves and till over 2.1 million of these models have been implanted all over the world. In the meantime, many useful changes in valve design have been made in the new models. There are many changes on the effective flow orifice area, on the shape of the leaflets (straight, convex or concave), on the pivot style, on the angle of orientation of the leaflets (from 72° to 90°), on the sewing-ring etc. (Chambers et al., 2005; Gelsomino et al., 2002; Gott et al., 2003; Walther et al., 2000).

Fig. 4. Three generations of mechanical heart valves: a. caged ball, b. tilting disk & c. bileaflet mechanical heart valves

#### **2.2 Evolution of bioprosthetic heart valves**

Heart valve transplantation is the substitution of the diseased heart valves with healthy living heart valves (valves transplanted from genetically similar donors-homologous valves), including auto transplantation (the substitution of the aortic valve with the pulmonary valve and the later with a prosthetic non-living valve – the Ross procedure). Murray in 1956 demonstrated that human aortic valves from cadavers could be used as a valve transplant in the descending thoracic aorta in patients with aortic insufficiency (Murray, 1956). Based on this research Kerwin and co-workers six years later reported their first clinical applications in patients, with one of them having 6-year follow-up (Kerwin et

contact area, material coating with biocompatible compounds (like pyrolitic carbon) and appropriate disk morphology for smoother blood flow around it. The most usable models were the Björk-Shiley and the Medronic Hall type. The main problems with these valves were the rare rupture of metal strut supporting disk movement and subsequent embolisation by the disk, non-axial flow, even in the models of Björk-Shiley with disk opening angle of 72°. Due to the excessive turbulent flow through the two orifices of the valve (a small and a bigger), a high-grade of haemolysis in patients was reported. However,

The 3rd generation of mechanical valves was appeared in 1977 with the introduction of the St. Jude Medical (SJM) bileaflet valve coated with pyrolitic carbon (Emery et al., 2008; Gott et al., 2003). Over the following decades, the dramatic step of bileaflet prostheses nearly obviated the use of all other kinds of mechanical valves in all over the world (Emery et al., 2008). In fact, the low-profile SJM valve demonstrated low rates of thomboembolism, low trans-valvular gradients, low grade of haemolysis and minimal valvular insufficiency (Chambers et al., 2005; Gott et al., 2003; Walther et al., 2000). Out of the SJM valve, several other 3rd generation models were introduced such as the ATS Medical Prosthesis, the Sulzer CarboMedics , the On-X prosthesis and the Sorin prosthetic valve, all of them with similar haemodynamic and clinical outcomes (Chambers et al., 2005; Walther et al., 2000). Since the introduction of this 3rd generation of the valves and till over 2.1 million of these models have been implanted all over the world. In the meantime, many useful changes in valve design have been made in the new models. There are many changes on the effective flow orifice area, on the shape of the leaflets (straight, convex or concave), on the pivot style, on the angle of orientation of the leaflets (from 72° to 90°), on the sewing-ring etc. (Chambers et al.,

Fig. 4. Three generations of mechanical heart valves: a. caged ball, b. tilting disk & c. bileaflet

Heart valve transplantation is the substitution of the diseased heart valves with healthy living heart valves (valves transplanted from genetically similar donors-homologous valves), including auto transplantation (the substitution of the aortic valve with the pulmonary valve and the later with a prosthetic non-living valve – the Ross procedure). Murray in 1956 demonstrated that human aortic valves from cadavers could be used as a valve transplant in the descending thoracic aorta in patients with aortic insufficiency (Murray, 1956). Based on this research Kerwin and co-workers six years later reported their first clinical applications in patients, with one of them having 6-year follow-up (Kerwin et

good long-term results characterized that type of MHV (Oxenham et al., 2003).

2005; Gelsomino et al., 2002; Gott et al., 2003; Walther et al., 2000).

mechanical heart valves

a

**2.2 Evolution of bioprosthetic heart valves** 

al., 1962). The first orthotopic insertion of homograft valve was performed in 1962 by Barratt-Boyes (Barratt-Boyes, 1964). The introduction of other biological valves began in 1967 when Senning used pieces of fascia lata of the patient for replacement of the diseased aortic valve (Senning, 1967; Ionescu & Ross, 1969).

Xenograft or heterograft valves, animal derived heart valves or BHVs made from different animal derived tissues are alternatives, offering the advantage of been prepared much prior the operation, available in different sizes and designs. BHVs include a variety of heart valve replacement using as substitutes heart valves of different orientations and technologies. Among different types, porcine aortic and bovine pericardial xenograft BHVs has been established as valve substitutes. Porcine aortic valves, after a treatment for removing excess fatty and aortic wall tissue and part of septal myocardial tissue from valve leaflets are imposed in biochemical preparation, aimed in removing antigenic factors (valve cells) and stabilize the remaining acellular valve tissue against enzyme reactions by different chemical compounds. Formaldehyde was first used for porcine valve fixation (Angel, 1972), substituted later by glutaraldehyde because of its ability for double-edge cross linking of collagen molecules (Woodroof, 1972), resulted in better longevity of valves. The stabilized tissue valve is sutured in specially designed frames composed of aortic ring and three commissures. A metallic or polymeric frame is used as a skeleton, covered with biocompatible textiles (like Teflon® or Dacron®) onto which the valve tissue is sutured with permanent sutures).

Improvement of BHV function was achieved with the use of different membranous soft tissues for the construction of valve leaflets and suturing them in similar artificial frames like that of the porcine valves. Percutaneous tissue was used for that scope; however, pericardial tissue from different animals was finally used alternatively. Pericardial tissue is a big membrane enclosing the heart. Its histology is similar to heart valve leaflets with respect of its composition of collagen and elastin fiber networks in different layers inside an amorhous organic matrix of glycosaminoglycans (GAGs), proteoglycanes (PGs), and other proteins. Extracellular water solution of electrolytes and soluble proteins compose 65-70% of the tissue mass weight. Cells, like fibroblast, epithelial, muscle cells and other types are present. Despite these similarities with other soft tissues, including heart valve leaflets, fibber structure of bovine pericardium is quite different. Fibber orientation in valve leaflets is specified for supporting their motion and strengthening mechanical stress developed during valve function. The different anatomic position and function of pericardial tissue resulted in a different fibber orientation, varied across its surface. For this reason special attention is given in selection criteria of specific regions from the whole pericardial membrane for better suitability to function as heart valve leaflets (Simionescu et al., 1993). The benefits from the use of pericardial tissue (especially bovine ones, which is the standard selection last years) instead of porcine valves were better opening area of the valve, coaptation of the valve leaflets and flexibility in design of valve anatomic configuration. However, similar problems with porcine BHVs, tissue deterioration and calcification still remain.

The first xenograft valve, a stent-mounted porcine aortic valve was implanted by Binet et al. in 1965 (Binet et al., 1965), while the glutaraldehyde-preserved stent-mounted porcine valves were introduced by Carpentier et al., in 1967 (Carpentier et al., 1969). Over the past 40 years, advances in tissue fixation (bovine pericardium and porcine aortic valves) methodologies and chemical treatments to prevent calcification, have yielded improvements in the longevity of bioprostheses.

Prosthetic Aortic Valves: A Surgical and Bioengineering Approach 67

• Non physiological geometry hemodynamic • Chronic anticoagulation therapy

• Risk of thromboembolism • Regular medical examinations

• Undesirable host reactions

• Tissue deterioration • Calcification

• Tissue deterioration • Calcification • Minimal availability • Undesirable immunologic reactions

*Valve type Advantages Disadvantages* 

• Easy implantation • Variety of size-designs • Availability

hemodynamic • Minimize anticoagulation

hemodynamic • Physiological remodeling • Minimal anticoagulation therapy

therapy • Availability

A mechanical valve prosthesis is recommended to patients having valve re-operation, regardless of the nature of the first operation, as the risk of the second operation is significantly higher (Gelsomino et al., 2002; Potter et al., 2005). The most debated age for making decision according the selection of type of prosthesis is the decade between 60 and 70 years. The final decision is dependent on other parameters, which have to be taken into account, like the existence of atrial fibrillation, chronic renal failure, cerebrovascular disease, history of gastrointestinal bleeding, contraindication to oral anticoagulants, etc. (Emery et al., 2002). In 1962, D. Harken summarized the 10 important characteristics that an ideal heartc valve

Table 1. Comparison of different types of heart valve substitutes

*Living grafts* • Ideal anatomy -

*Bioprosthetic* • Physiological anatomy -

*Mechanical* • Longevity

2. Be chemically inert and not damage blood elements.

10. Be technically practical to insert (Harken et al., 1962).

**3. Biomechanics of heart valve prostheses** 

5. Remain closed during the appropriate phase of the cardiac cycle.

7. Be inserted in a physiological site, generally the normal anatomical site.

After a near 50 years of evolution today's mechanical or biological valve do not satisfy all

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

3. Offer no resistance to physiological flows. 4. Close promptly (less than 0.05 second).

6. Lasting physical and geometric features.

8. Be capable of permanent fixation.

must satisfy:

1. Not propagate emboli.

9. Not annoy the patient.

those requirements.

Like the mechanical valves, the development of biologic valves is characterized by the appearance of first-, second-, and third-generation prostheses. **First-generation** bioprostheses were preserved with high-pressure fixation. They include the Medtronic Hancock Standard, and the Carpentier-Edwards Standard valves, both porcine prostheses. **Second-generation** prostheses are treated with low-, or zero-pressure fixation. Pericardial prostheses include the Carpentier-Edwards Perimount, and the Pericarbon Sorin prostheses. Porcine prostheses include the Medtronic Hancock II, the Medtronic Intact, and the Carpentier-Edwards Supraannular prostheses. In the third-generation prostheses were included all valves with zero-, or low-pressure fixation, decellularization of animal tissues and simultaneous antimineralization processes (e.g. a-amino oleic acid) to reduce material fatigue and calcification. In these models the stents have become gradually thinner and flexible, the profile much lower and the effective orifice area larger. Porcine prostheses include the Medtronic Mosaic, and the St. Jude Medical Epic valve. Pericardial prostheses include the Carpentier-Edwards Magna and the Mitroflow Pericardial valve.

Fig. 5. Pericardial bioprosthetic heart valve explanted due to severe calcification. A: aortic side, b: ventricular side, c: SEM micrograph of the same valve demonstrating calcific crystals deposited implemented with leaflet tissue fibber network

#### **2.3 Comparison of MHV with BHV replacement**

Aortic valve replacement by using a mechanical prosthetic valve is not indicated for all patients suffered from aortic valve stenosis. Generally, the two main advantages of the mechanical valves are the absence of degeneration (long-term rigidity) and the larger effective flow orifice, both contributing to better long-term outcomes. In fact, according to many studies followed over time frames, freedom from all-valve related events and from the risk of reoperation were improved in patients with mechanical valve prostheses as compared to those with biologic prostheses (Ezekowitz, 2002; Gott et al., 2003). On the other hand, the main disadvantage of mechanical valves is the obligatory need for long-life anticoagulation (Ezekowitz, 2002). The use of porcine BHVs resulted in a better function in patient circulatory system, improving failed valve insufficiency due to stenosis or regurgitation of diseased natural valves. A central blood flow with a near physiological velocity profile, low pressure gradient across the valve, near physiological leaflet movements and no need for long term anticoagulation therapy were the benefits of their use. However, porcine BHVs longevity remains limited, mainly because of calcification of valve leaflets. Calcific crystal deposits are gradually developed in valve leaflets caused stiffening and incompetence in their moving ability and valve dysfunction due to stenosis or regurgitation. BHV calcification is more often in younger patients, for which MHVs are the gold standard in heart valve replacement.

Like the mechanical valves, the development of biologic valves is characterized by the appearance of first-, second-, and third-generation prostheses. **First-generation** bioprostheses were preserved with high-pressure fixation. They include the Medtronic Hancock Standard, and the Carpentier-Edwards Standard valves, both porcine prostheses. **Second-generation** prostheses are treated with low-, or zero-pressure fixation. Pericardial prostheses include the Carpentier-Edwards Perimount, and the Pericarbon Sorin prostheses. Porcine prostheses include the Medtronic Hancock II, the Medtronic Intact, and the Carpentier-Edwards Supraannular prostheses. In the third-generation prostheses were included all valves with zero-, or low-pressure fixation, decellularization of animal tissues and simultaneous antimineralization processes (e.g. a-amino oleic acid) to reduce material fatigue and calcification. In these models the stents have become gradually thinner and flexible, the profile much lower and the effective orifice area larger. Porcine prostheses include the Medtronic Mosaic, and the St. Jude Medical Epic valve. Pericardial prostheses include the

Fig. 5. Pericardial bioprosthetic heart valve explanted due to severe calcification. A: aortic side, b: ventricular side, c: SEM micrograph of the same valve demonstrating calcific crystals

Aortic valve replacement by using a mechanical prosthetic valve is not indicated for all patients suffered from aortic valve stenosis. Generally, the two main advantages of the mechanical valves are the absence of degeneration (long-term rigidity) and the larger effective flow orifice, both contributing to better long-term outcomes. In fact, according to many studies followed over time frames, freedom from all-valve related events and from the risk of reoperation were improved in patients with mechanical valve prostheses as compared to those with biologic prostheses (Ezekowitz, 2002; Gott et al., 2003). On the other hand, the main disadvantage of mechanical valves is the obligatory need for long-life anticoagulation (Ezekowitz, 2002). The use of porcine BHVs resulted in a better function in patient circulatory system, improving failed valve insufficiency due to stenosis or regurgitation of diseased natural valves. A central blood flow with a near physiological velocity profile, low pressure gradient across the valve, near physiological leaflet movements and no need for long term anticoagulation therapy were the benefits of their use. However, porcine BHVs longevity remains limited, mainly because of calcification of valve leaflets. Calcific crystal deposits are gradually developed in valve leaflets caused stiffening and incompetence in their moving ability and valve dysfunction due to stenosis or regurgitation. BHV calcification is more often in younger patients, for which MHVs are the

Carpentier-Edwards Magna and the Mitroflow Pericardial valve.

a b c

deposited implemented with leaflet tissue fibber network

**2.3 Comparison of MHV with BHV replacement** 

gold standard in heart valve replacement.


Table 1. Comparison of different types of heart valve substitutes

A mechanical valve prosthesis is recommended to patients having valve re-operation, regardless of the nature of the first operation, as the risk of the second operation is significantly higher (Gelsomino et al., 2002; Potter et al., 2005). The most debated age for making decision according the selection of type of prosthesis is the decade between 60 and 70 years. The final decision is dependent on other parameters, which have to be taken into account, like the existence of atrial fibrillation, chronic renal failure, cerebrovascular disease, history of gastrointestinal bleeding, contraindication to oral anticoagulants, etc. (Emery et al., 2002).

In 1962, D. Harken summarized the 10 important characteristics that an ideal heartc valve must satisfy:


After a near 50 years of evolution today's mechanical or biological valve do not satisfy all those requirements.
