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RECENT ADVANCES IN POLYMERIC HEART VALVES RESEARCH

Recent Advances in Polymeric Heart Valves Research

Yee Han, Kuan1, Prasad Dasi, Lakshmi2, Ajit P., Yoganathan 3, Hwa Liang, Leo1*

1Division of Bioengineering, National University of Singapore

2Department of Mechanical Engineering, Colorado State University

3Wallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology

*Correspondence author:

Leo Hwa Liang, Division of Bioengineering, National University of Singapore, Singapore; Tel: +65-6516-5608, Fax: +65-6872-3069, email address: [email protected]


Abstract

Heart valve replacement is fast becoming a routine surgery worldwide and heart valve

prostheses are today considered among the most widely used cardiovascular devices.

Mechanical and bioprostheses have been the traditional choices to the replacement surgeries.

However, such valves continue to exposed patients to risks, including thrombosis, infection

and limited valve durability. In recent years, advances in polymer science give rise to an

important and new class of artificial heart valve made predominantly of polyurethane-based

materials, which show improved biocompatibility and biostability. These polymeric heart

valves have demonstrated excellent hemodynamic performance and good durability with

excellent fatigue stress resistance. Advancements in the designs and manufacturing methods

also suggested improved in the durability of polymeric heart valves. Animal studies with

these valves have also shown good biocompatibility with minimal calcification of the valve

leaflets. With these promising progresses, polymeric heart valves could be a viable alternative

in the heart valve replacement surgeries in the near future.

Keywords: Polyurethane, Hemodynamic, Reynolds shear stress, Bioprostheses, Calcification.


Recent Advances in Polymeric Heart Valves Research

Existing commercial prosthetic heart valve designs can be classified into two groups:

bioprosthetic, or tissue valves, and mechanical heart valves. The former is made from a

combination of synthetic materials and chemically treated animal tissue mainly porcine in

origin, whereas the latter is manufactured entirely from synthetic materials. In the 1960s, Dr

Charles Hufnagel performed one of the earliest successful mechanical heart valve surgeries

where six of the eight patients who received a caged ball heart valve survived the operation

(Hufnagel, Gillespie, Conrad, Mercier, & Evangelist, 1966). However, current artificial heart

valves are still exposed to risks of thrombosis and thromboembolism, tissue overgrowth,

infection, anti-coagulant related hemorrhage, and valve failure due to material fatigue or

chemical change (Black, & Drury, 1994; Yoganathan, 2000; Starr, Fessler, Grunkemeier, &

He, 2002; Korossis, Fisher, & Ingham, 2000).

One major drawbacks associated with the implantation of mechanical heart valves is

the need for daily chronic anticoagulation therapy to reduce the risk of thrombosis and

thromboembolic complications. These patients are expose to an increased risk of bleeding,

infection, and/or autoimmune response (Walker, & Yoganathan, 1992). Blood flow through

mechanical prostheses can lead to high turbulent stresses that may damage and/or activate

blood elements and initiate platelet aggregation. Platelet aggregation can lead to thrombus

formation with disastrous consequences for the patient. Thrombi may even detach from the

valve and become lodged in a downstream blood vessel, thus reducing or even cutting off the

blood supply to vital tissues.

An alternative to mechanical valves is using tissue valves which utilize the concept of

a trileaflet configuration with one central orifice that mimics the design of the native valve.

These valves have a lower potential for blood element damage than their mechanical

counterparts. Nevertheless, the mechanical properties of tissue valves appear to degrade more


rapidly, and they are prone to calcification (Black, & Drury, 1994). Often, implanted tissue

valves do not last for more than ten years, and reoperation is necessary.

The polyurethane (PU) trileaflet polymeric heart valve is the latest development in

prosthetic heart valve research (Hyde, Chinn, & Phillips, 1999). The design is based on the

natural aortic valve and is inherently appealing from a hemodynamic viewpoint. Although

this particular valve design is still in a developmental stage, preliminary studies have shown

excellent forward flow hemodynamic properties equivalent to that of a tissue heart valve and

promise a durability comparable to that of a mechanical heart valve (Bernacca, Mackay,

Wilkinson, & Wheatley, 1995; Jansen et al., 1991). However, recent animal trials involving

polymeric valves have reported problems mainly related to tearing of the leaflets and

thrombus formation occuring along the stent region of the valve (Wheatley et al., 2001). In

addition, results of long-term in vivo evaluation have suggested that calcification could be a

limiting factor to long-term function of polymeric valves (Bodnar, Frater, & eds, 1991). The

leaflets and basal attachments, such as the commissural region of the polymer valves, have

experienced extrinsic calcification associated with surface microthrombi that appear to be

independent of structural defects suggesting that the flow characteristics inside the polymeric

valve may be a contributor to the observed blood clots.

This review will discuss the recent development of polymeric heart valves,

particularly on its choice of materials. Much attention will also focus on the in vitro

hemodynamics characterization these polymeric heart valves. The review will also discuss

the in vivo experiments involving these polymeric heart valves and their susceptibility to

calcifications.

Introduction of Polymeric Heart Valves

Pioneering works in polymeric heart valve replacement in human was first reported in


1958 (Roe, Owsley, & Boudoures, 1958). This is followed by a single trileaflet silicone

rubber (SR) aortic valve implanted in several patients between 1960 and 1962 (Roe, Kelly,

Myers, & Moore, 1966). The clinical trials were discontinued due to the high occurrence of

thromboembolic events. The choice of material is an important factor in the development of

polymeric heart valve. The chosen polymer should have acceptable characteristics with

regard to biocompatibility, hemocompatibility, anti-thrombogenicity, resistance to

degradation and calcification (Daebritz et al., 2004; Gallocher, Aguirre, Kasyanov, Pinchuk,

& Schoephoerster, 2006). Since then, several synthetic polymers have been tested as leaflet

materials, including silicone, polyofelin rubber, Silastic, SR-impregnated Dacron mesh,

polytetrafluoroethylene (PTPE) and PU.

In 1960s, silicone was considered a choice due to its great flexibility and

biocompatibility. However, in clinical investigations, formation of thrombosis and thickening

of valve leaflets were found. Silicone suffered from structural failure and short durability, and

was abandoned as a possible material replacement (Roe, Kelly, Myers, & Moore, 1966).

Another study also showed the inadequate durability of silicone and polyofeline rubbers

while in vivo (Cheeta, & Llyod, 1980). PTFE or Teflon valves were considered as a

polymeric heart valve material attributing to their good hemodynamic properties. However,

its low resistance to thromboembolism and stiffening of the leaflets prevented PTFE from

further development (Nistal, 1990). Calcification was also found on PTFE valves (Imamura,

& Kaye, 1977).

In 1982, a breakthrough came when Wisman and colleagues invented a trileaflet valve

fabricated from segmented polyurethane (SPU) (Wisman et al., 1982). The large central flow

orifice of the leaflet valve is similar to tissue valves, which reduces turbulence and blood

trauma as a result of lower transvalvular pressure drop and better hydraulic efficiency. The

valve demonstrated high flexure endurance, great strength and inherent nonthrombogenic


characteristics.

Choice of Materials

PU has been widely used in biomedical application and it exhibits several favourable

properties due to its two-phased microstructure consisting of hard crystalline segments and

soft elastomeric segments. The hard segment confers structural strength while the soft

segment determines the flexibility and elasticity of the material.

Polyurethanes elastomers are copolymers containing blocks of low molecular weight

polyesters or polyethers linked together by a urethane group. These elastomers are based on

three monomers: (1) an isocyanate source, (2) a macroglycol, and (3) a chain extender, or

curative. The isocyanate can either be aromatic or aliphatic. One example of aromatic

isocyanate is the methylene diphenyl diisocyanate (MDI), which is a very reactive isocyanate.

Usually, it does not require a catalyst for synthesis if a diamine chain extender is used. An

example of an aliphatic isocyanate is the cyclo-aliphatic diisocyanate H12MDI. The aromatic

isocyanate containing polymers have a tendency to yellow upon exposure to ultraviolet

radiation, and some aromatic polyurethanes, such as those made from methylene dianiline, or

MDA, are known carcinogens and display immediate cell toxicity. However, elastomers

produced from less toxic aliphatic isocyanates have inferior physical properties and are

significantly more expensive (Szycher, 1988).

The reaction of a diisocynate with a short chain diol or diamine (‘chain extenders’)

forms the hard segments. Commonly used extenders for medical use are the 1,4 butane diol

or ethylene diamine. On the other hand, the reaction of the diisocynate with higher molecular

weight polyols, such as polyether, polyesther or polycarbonate, forms the soft segments.

Medical grade segmented polyurethane (SPU) have been applied successfully in various

cardiovascular devices, including artificial hearts, ventricle assist devices and blood pumps


(Bernacca, Mackay, Gulbransen, Donn, & Wheatley, 1997; Bernacca, Straub, & Wheatley,

2002).

In recent years, remarkable advancement in polymer synthesis has resulted in a more

bio-stable PU. Recent studies have yielded a wider range of SPUs for polymeric heart valve

leaflets (Alferiev et al., 2001; Bernacca, Mackay, Gulbransen, Donn, & Wheatley, 1997;

Seifalian et al., 2003; Wheatley et al., 2001). The main types of PU which have been

developed and tested are polyester urethane, polyether urethane (PEU) and polycarbonate

urethane (PCU).

PEU were found to be unsuitable for long-term implants due to the hydrolytic

degradation. Such polymers are vulnerable to oxidative degradation and suffers

environmental stress cracking and metal ion oxidation under in vivo conditions (Stokes,

McVenes, & Anderson, 1995). PCU were tested and proven to have a higher oxidative

stability (Salacinski, Odlyha, Hamilton, & Seifalian, 2002). When compared to PEU, the

biodegradation of PCU was significantly lower and limited to a thinner surface layer

(Christenson, Anderson, & Hiltner, 2004).

In order to improve the biostability of PU, the chemical structures have been

modified. Bernacca and colleagues developed a polyurethane valve which was chain-

extended with ethylene diamine (PEUE). Good phase separation of polyurethane soft and

hard segments with good rubbery characteristics helped to explain the excellent PEUE valve

durability (Bernacca, Mackay, & Wheatley, 1996).

Another nanocomposite PU polymer, PCU and polyhedral oligomeric silsesquioxanes

(POSS) was developed and studies have shown that the polymer possesses greater

thromboresistance than PTFE (Kannan et al., 2006). In vivo study conducted using the POSS-

PCU composite found minimal inflammation and capsular formation in a sheep model

compared to control. No degradation was found even after 36 months of post-implantation in


the sheep model (Kannan et al., 2007). The increased fibrinogen adsorption on POSS-PCU,

and large contact-angle hysteresis suggested that POSS-PCU inhibits inflammation by

adsorbing and inactivating fibrinogen on its surface. This composite has enhanced interfacial

biocompatibility and better biological stability as compared with conventional silicone

biomaterials.

In a study by Kidane and his team, it was found that POSS-PCU nanocomposite

possess excellent mechanical strength, good surface properties and resistance to platelet

adhesion. It has also shown comparable hardness and tear strength to commercially available

products such as Estane, Chronoflex C and Elasteon (Kidane et al., 2009). The inclusion of

nanoscale POSS particles into the polymer matrix was suggested as the contributing factor to

the increment in tensile strength and hardness of POSS-PCU. The hydrophobicity factor of

the polymer helps resist the adsorption of blood proteins, which is necessary in preventing

thrombosis and calcification on the heart valve surfaces.

In 2004, Jiang and colleagues developed and demonstrated a novel one-piece,

tricuspid valve made entirely from polyvinyl alcohol cryogel (PVA-C). The mechanical

properties of the valve were successfully demonstrated with reduction in flexural stresses of

the leaflets. The stress concentrations at the commissural areas where the leaflets are attached

to the stent were also avoided (Jiang, Campbell, Boughner, Wan, & Quantz, 2004). However

further studies are required to evaluate the potential of PVA-C valve as a superior alternative

to the current valves.

More recently, a novel polyolefin poly(styrene-b-isobutylene-b-styrene) (SIBS) was

being proposed as a viable polymer for use as valve due to its oxidation resistance. SIBS also

showed suitable hemocompatibility and mechanical durability, with over 350 million cycles

evaluated in vitro in the early stages. The material was comparable to PU in platelet

deposition tests, as well as tensile and fatigue properties after reinforcement of polypropylene


fibres. Platelet activation due to the polymer valve was not found to be significantly different

from either a St. Jude bileaflet valve or a St. Jude Tissue valve. (Gallocher, Aguirre,

Kasyanov, Pinchuk, & Schoephoerster, 2006). Such encouraging traits open the possibility of

SIBS as a potential material choice for further development in polymeric heart valve.

Polymeric Heart Valve Design

There are several essential requirements of a heart valve design. The valve should fit

nicely into the host anatomy with minimal thrombogenicity and damage to blood cells. The

valve leaflets should also ensure minimum resistance to the forward flow and open at a

minimum systolic trans-valvular pressure difference. This promote smooth washout and

hence are less susceptible to trauma-promotion (causing thrombus, atherosclerosis and

hemolysis) (Bellhouse, & Reid, 1969). There should also be appropriate sealing in the close

position to minimize backward flow. The stress peaks in valve components should be low to

ensure durability and minimal change to the geometric features such as creep (Mercer,

Benedicty, & Bahnson, 1973). Figure 1 shows an example of a valve design.

Attempts to manufacture leaflet valves from polymers started in the 1950s and as

early as the inception of the first mechanical heart valve prostheses. However, only recently

the first realistic solution was developed. The major hurdles for the success of the polymeric

valves pertained to valve design and its chemical composition to eliminate thrombogenicity,

calcification, mechanical stress, and durability. The recent design of polyurethane valve is the

stent-supported bioprosthetic heart valve. Here the polymeric leaflets are mounted onto

flexible stents that lead to a circular orifice during forward flow. However, the leaflets, stent,

and the sewing ring in polymeric valves are made of a multilayered cohesively bonded single

material and are available at various degrees of hardness instead of each leaflet being sewn

separately onto the stents as in the bioprosthetic valve (Daebritz et al., 2004). The leaflets at


relax state is semi-opened to minimize membrane stresses during diastole and systole. Stent

posts are flexible to ensure tight leaflet closure during diastole. All current bioprosthetic heart

valves are tri-leaflet, however polymeric valves have also been designed with a bileaflet

configuration specifically for mitral position implantation (Daebritz et al., 2004). See Figure

2 for the three stent aortic design and two stent mitral design of the polymeric heart valves

Many designs have been introduced over the years of heart valve development such as

the single-leaflet silicone-covered Dacron cusps implanted by Hufnagel in the 1970s

(Hufnagel, 1977) and asymmetric bileaflet mitral valve with a large anterior and small

posterior cusp (Sachweh, & Daebritz, 2006). However, the most common design is the

trileaflet design and the early basic trileaflet aortic valve design was used in the 1970s (Mori

et al., 1973). There is another variation such as a three-leaflet J-3 valve, manufactured in a

medium open position with almost flat leaflets. This valve showed superior hydrodynamic

performance with minimum pressure drop and very low energy losses compared with other

commercially available valves. Besides that, the downstream flow field of the valve indicated

low shear stress when observed with Laser-Doppler-Anemometry. The J-3 valve

demonstrated great durability tests, with lifetime equivalent up to 17 years (Jansen et al.,

1991).

In 1994, Leat and his team designed a leaflet heart valve, termed ‘alpharabola’, which

has a continuous radius of curvature. The smallest radius is at the center of the leaflet at the

free edge and it demonstrated good opening characteristics (Leat, & Fisher, 1994). An

important understanding is that that the leaflet geometry can reduce the maximal tensile stress

up to 60% of those of the same spherical valve using the same mesh density (Clift, & Fisher,

1996). Another valve design with three thin polyurethane leaflets suspended from the inside

of a flexible polyurethane frame was also developed. The closed-leaflet geometry is elliptical

in the radial direction and hyperbolic in the circumferential direction. The valves showed


promising durability tests, with all six valves exceeded the lifetime equivalent of 10 years.

Nevertheless, one valve failed due to a leaflet tear after reaching 12 years cycling (Mackay,

Wheatley, Bernacca, Fisher, & Hindle, 1996).

Another interesting design is a stentless and fiber-reinforced synthetic valve as shown

in Figure 3. This design mimics stress-reduction and may decrease the tears and perforation

in the leaflets. The prototype also demonstrated low resistance to forward flow. However, the

manufacturing procedure to produce a stentless valve is complicated due to the irregular

shape of the mold to keep the fiber manually in place during winding procedure (Cacciola,

Peters, & Baaijens, 2000).

Leaflet Thickness

Thin SPU leaflets between 100-150 µm perform well in terms of trans-valvular

pressure gradients and energy losses. This leaflet thickness may achieve good hydrodynamic

flows but may compromise the durability. Although the leaflets can be manufactured as thin

as 100 µm to be flexible and durable, there are concerns at thickness less than 150 µm

(Mackay, Wheatley, Bernacca, Fisher, & Hindle, 1996). In order to reduce the imparted

bending stress, a valve that undergoes cyclical flexion must be manufactured from low elastic

modulus material.

SPU is widely used because it can flex easily. This helps to minimize blood flow

disturbance and resistance. The ease of flexion criterion must be counterbalanced by the

ability to withstand the high reverse pressures and mechanical stresses present when the valve

opens and closes in a normal physiological cycle.

Bernacca and his team examined the hydrodynamic behavior of biostable

polyurethane valves varying Young's modulus from 5 to 63.6 MPa and a mean leaflet

thickness from 48 to 238 µm. They demonstrated that valve hydrodynamic function is not


significantly affected by the material modulus with little difference in the pressure gradient.

However, the increasing leaflet thickness is produces a significant rise in pressure gradient.

Energy loss during forward flow is proportional to the behaviour of the pressure gradient.

Increased in leaflet thickness is associated with the decrease in valve closing efficiency

(Bernacca, O'Connor, Williams, & Wheatley, 2002).

In another study by Bernacca and his colleagues, leaflet thickness was varied from 60

to 200 µm and it was also found that the thickness of a polyurethane valve is critical to its

function. Durability of the valve was directly related to leaflet thickness, with good durability

achieved with median leaflet thickness of approximately 150 µm. The polymeric heart valves

exceeded 800 million cycles (Bernacca, Mackay, Gulbransen, Donn, & Wheatley, 1997). In a

more recent in vitro fatigue testing, the durability of PCU valves with leaflet thickness

varying between 100 and 300 µm was between 600 million (15.8 years) to 1 billion cycles

(26 years) of normal human function (Daebritz et al., 2003).

Fabrication processes

Today, all polymeric heart valves are made mainly out of polyurethane, with a

combination of solution casting and injection molding. The stents or frames are injection

molded and typically have a thickness of approximately 3 mm. The polyurethane frames are

then molded onto steel formers of ellipto-hyperbolic leaflet shape (Mackay, Wheatley,

Bernacca, Fisher, & Hindle, 1996) and dipped into a concentrated polyurethane solution so

that they coat the whole valve to form the leaflets. The polymer valve is then allowed to dry

while hanging free edge downward. The leaflet edge is later cut and trimmed by precision a

laser cutting tool. The thickness of the leaflet ranges from 80 to 300 µm (Bernacca, Mackay,

Gulbransen, Donn, & Wheatley, 1997; Daebritz et al., 2003; Daebritz et al., 2004; Mackay,

Wheatley, Bernacca, Fisher, & Hindle, 1996). Some polyurethane valves contain a stiffening


ring of radio-opaque MRI compatible titanium alloy to facilitate radiographic imaging.

There are several types of fabrication techniques for polymeric heart valves; (1) Dip

casting, (2) Film fabrication, and (3) Cavity moulding. The fabrication of the polymeric valve

usually involves coating a semi-rigid stent made of poly-ether-ether-ketone (PEEK) in

polyurethane (Wisman et al., 1982). Traditionally, SPU valves have been manufactured by

dip-coating, which involves the use of a specifically designed mandrel. It involved multiple

dips in a weak polymer solution (Jansen et al., 1991). The major disadvantage for this method

is that the leaflet thickness distribution cannot be control accurately.

In film fabrication, pre-cast SPU film is solvent-bonded to the valve frame and

thermally formed to the leaflet shape. This method allows greater control over the desired

geometry of the valve (Leat, & Fisher, 1995). However, this method would also yield a valve

that was inferior in term of durability due to an inconsistent leaflet-frame interface.

Jiang et al showed a cavity mould consists of a static assembly and a moving

assembly can be used to fabricate the whole valve structure. Hot polymer is poured into the

mould and placed in a water bath. The cyclic freezing and thawing process at – 20°C and +

20°C forms a thin film of polymer (Jiang, Campbell, Boughner, Wan, & Quantz, 2004). The

valve manufacturing method has shown to significantly influence the performance of the

valve leaflets. Valves manufactured using a dip-casting method has shown up to 31%

reduction in open leaflet bending strains compared to valves manufactured from solvent-cast

sheets of polyurethane (Corden, David, & Fisher, 1995).

The longer durability of dip cast valves compared to thermally formed film fabricated

leaflets was also verified by Leat and Fisher, although no significant difference was found in

the hydrodynamic function of the valves between the two prescribed methods. In their

comparison study, one method is by thermally forming the leaflets from film and bonded to

the frame while the other method is by dip casting the leaflets on to the frame using s-shaped


mandrel. A simple and reproducible fabrication method using a combination of solution

casting and injection moulding was used by Mackay et al in their polyurethane heart valve

development (Mackay, Wheatley, Bernacca, Fisher, & Hindle, 1996). These valves were

tested and found to exhibit pressure gradients similar to those for a bioprosthesic valve with

considerably lower levels of regurgitation and leakage than bileaflet MHVs or bioprosthesic

valves.

In vitro studies of polymeric heart valves

In vitro studies have been conducted to determine the performance of polymeric heart

valves. Many of these valves have been tested in vitro under pulsatile flow conditions for the

hydrodynamic functions. The levels of regurgitation and leakage are considerably less than

the mechanical and porcine bioprosthetic valves (Mackay, Wheatley, Bernacca, Fisher, &

Hindle, 1996).

It has been shown that polymeric heart valves generate a centralized flow with a

relatively flat profile similar to that of bioprostheses and native heart valves in vitro studies.

The flow patterns of four different trileaflet polyurethane heart valves were compared to the

Hancock porcine heart valve and Ionescu-Shiley 3 pericardial valve by measuring the down-

stream velocity distributions using Laser-Doppler-Anemometry (Herold et al., 1987). It was

observed that the velocity distributions at peak systole are very similar between the

polyurethane valve and Hancock prosthesis. Both flow fields can be characterized as a

triangular jet-flow, with the recirculation areas located near the right aortic wall downstream

of the stent post and in the opposite sinus cavity. However, flat plug type flow was observed

for the Ionescu-Shiley 3 valve. The flow is positive with recirculation in the aortic sinus

cavity and one small transient recirculating vortex near the right aortic wall. Overall, the flat

flow profile of polyurethane valves is similar to those of the native and bioprostheses valves


with vortex formation observed behind the leaflets in the sinus cavity.

De’Souza and colleagues characterized the velocity and turbulence waveforms distal

to several trileaflet polymeric heart valves with varying degrees of stenosis (D'Souza,

Butterfield, & Fisher, 2003). They observed that the shape and orientation of the spatial

velocity profile in a polyurethane heart valve depicts the effect of the constriction caused by

the valve leaflet in a similar manner to that observed in bioprosthetic valves.

In another study, the flow dynamics of two polyurethane valves are investigated; one

with leaflets bound to the edge of the supporting stents (PUI) and the other with leaflets

mounted outside the stents (PUII) (Chandran et al., 1989). The PUII valves have larger inlet

orifice diameter with a smaller tissue annulus diameter, which provide more unobstructed

area for blood flow. Such valve design was also found to reduce the maximum turbulent

shear stresses distal to the valves, which appears more advantageous. Furthermore, the

turbulent shear stress in the PUII valve was half that in the PUI valve showing that the

increase in orifice diameter can help reduce the turbulent shear stresses and corresponding

damage to blood elements. When compared to tissue valves, polymeric valves have larger

turbulent shear stresses. The orifice diameters in polymer valves are usually smaller than

those in bioprostheses, and as a result, polyurethane valves have larger observed velocity

magnitudes and larger velocity gradients than the bioprostheses.

Recent studies have also shown that the trileaflet design is more efficient than double

or quadruple arrangements (Cheeta, & Llyod, 1980; Jansen et al., 1991; Leat, & Fisher, 1995;

Mackay, Wheatley, Bernacca, Fisher, & Hindle, 1996), and that the maximal stresses in these

trileaflet configurations were found near the commissural attachment, which is consistent

with other synthetic and bioprosthetic valves (Leat, & Fisher, 1995). Corden and team

calculated the bending stresses and strains at the free leaflet edge by measuring the curvature

and the study showed that the highest curvatures occurred at the commissural regions when


the valves are fully open (Corden, David, & Fisher, 1995).

Leo and colleagues performed Laser Doppler Velocimetry (LDV) and Particle

Imaging Velocimetry (PIV) studies of three Aortech polymeric heart valves investigating the

effect of commissural design and leaflet thickness on the hemodynamic performance of the

heart valves. These studies indicated that commisural design and leaflet thickness can

influence the thrombogenic potential of tri-leaflet polymeric valves. The studies, as seen in

Figure 4 and 5, identified the following regions of high shear stress and high velocity flow in

the three polymeric heart valves: 1) the leakage jet inside the valve during diastole; 2) the

trailing edge of the valve; 3) the vortex ring surrounding the forward jet during the

acceleration phase; 4) the edge of the central orifice jet, which extended from the inside of

the valve into the distal part of the flow chamber; and 5) the distal region of the flow chamber

where turbulence mixing occurred during systole. Low flow velocities were observed in the

following regions of the three polymeric heart valves: 1) the split flow inside the valve during

diastole; 2) a vortex inside the sinus region; and 3) flow inside the recirculation zone along

the chamber wall during diastole. Additional low flow regions were observed in prototype A:

1) the inflow stent region during late systole; and 2) the closed commissural region during

systole (Leo, 2005).

Using flow visualization and LDV, several studies have demonstrated that trileaflet

bioprostheses produced jet-type flow fields like those observed in the polymeric heart valves

(Woo, Williams, & Yoganathan, 1983; Woo, Williams, & Yoganathan, 1983; Yoganathan et

al., 1983; Yoganathan et al., 1986; Yoganathan, Woo, & Sung, 1986). In general, the

pericardial and Aortech polymeric heart valve designs created flatter velocity profiles than

the porcine and Abiomed valve designs. All three Aortech polymeric heart valves produced

jets with flat velocity profiles at peak systole, identical to those observed in the Ionescus-

Shirley pericardial, Carpentier-Edwards 2650 porcine, Hancock pericardial, and Carpentier-


Edwards pericardial valves. In contrast, the velocity profiles of the central jets in the

Carpentier-Edwards 2625 porcine, Hancock modified-orifice porcine, Hancock II porcine,

and the Abiomed valves were more parabolic in shape.

The diameter of the central jets in the tissue bioprostheses at peak systole is typically

larger than that in the polymeric heart valves. The bioprosthetic valves studied have larger

tissue annular diameters compared with the polymeric prostheses (see Table 1). For a given

cardiac output, valves with larger tissue annular diameters typically have larger jet diameter

compared to those with smaller tissue annular diameters. Central orifice jet diameters

between 15 - 25 mm were observed in the bioprostheses at peak systole with the smallest

occurring in Carpentier-Edwards 2625 and Hancock standard-orifice porcine valves, and the

largest in the Hancock pericardial and Carpentier-Edwards pericardial valves(Yoganathan et

al., 1986). The jet diameters for the Aortech polymeric heart valves were comparable to that

of the Abiomed trileaflet polymeric valve prosthesis, which had a jet diameter of 14 mm

(Woo, Williams, & Yoganathan, 1983; Woo, Williams, & Yoganathan, 1983). The Aortech

prototype A had a jet diameter of 13 mm, while both Aortech prototypes B and C a jet

diameter of 16 mm.

Flow separation was observed at the edges of the valve leaflets and around the jet.

Flow reattachment occurred when the central orifice jet diverged and impinged on the

chamber wall downstream of the valve. The location of reattachment points varies from valve

to valve. Under the same cardiac output, valves with larger orifice sizes tend to have

reattachment points closer to the trailing edge of the valve leaflet. For instance, the central

orifice jets of the Abiomed 25 mm and 21 mm valves were observed to reattach 40 mm and

60 mm downstream from the valves, respectively (Woo, Williams, & Yoganathan, 1983). A

common observation pertaining to the reattachment points of the central orifice downstream

of the valve was made in the flow visualization studies as well as in current studies.


Reattachment of the orifice jet occurred at different downstream locations in the upper and

lower parts of the flow chamber. This observation is attributed to the asymmetric nature of

the central orifice jet. In the Abiomed 25 mm and 21 mm valves, the central orifice jet

reattached approximately 75 mm and 40 mm downstream of the sewing ring in the upper and

lower parts of the flow chamber, respectively (Woo, Williams, & Yoganathan, 1983). For the

tissue bioprostheses, the jet typically reattached in the upper part of the flow chamber at a

location between 55 – 90 mm downstream from the valve sewing ring, and at a location

between 35 – 50 mm in the lower part of the flow chamber. These findings are consistent

with those made for the three Aortech polymeric heart valves; the central jet in the three

prototypes valves reattached at a location between 45 - 55 mm downstream from the valve

sewing ring in the lower part of the flow chamber. For both prototypes A and C, the jet

reattachment point in the upper part of the flow chamber seem to occur beyond the

measurement plane, which only extended 77.5 mm downstream from the valve sewing ring.

The exception to the jet reattachment location in the upper part of the flow chamber was

observed in both the Hancock pericardial and the Aortech prototype B valves. The central

orifice jets in both valves were observed to reattach closer to the valve at approximately 55

mm downstream from the valve sewing ring in the upper part of the flow chamber.

The leaflet profile and flow patterns of the three Aortech polymeric heart valves at

peak systole was similar to that seen in the pericardial heart valve where a small recirculation

flow was observed at the edge of the leaflet during systole as illustrated in Figure 6 (Black,

Cochrane, Lawford, Reul, & Yoganathan, 1991). The shape of the open leaflet may explain

the formation of a vena contracta in the central orifice jet. This narrowing of the central

orifice jet immediately downstream of the valve was due to the opening configuration of the

valve during systole where the shape of the leaflet was such that the flow area at the leaflet’s

trailing edge was more constricted than that inside the valve.


Table 1 compares the peak velocities and Reynolds shear stress values of previous

valve studies with those obtained in the current polymeric valve experiments. It can be seen

that the peak forward velocities obtained from previous studies were comparable to those

recorded in the current study. The peak Reynolds shear stress levels calculated downstream

of all valves during systole was typically between 1,000 – 2,900 dynes/cm2. The highest

Reynolds shear stress of 8,347 dynes/cm2 was recorded in Aortech prototype A, while both

the Abiomed 21 mm and the Carpentier-Edwards 27 mm 2625 porcine valves recorded

Reynolds shear stresses of 4,500 dynes/cm2. Reynolds shear stress studies performed with the

trileaflet bioprostheses revealed a Reynolds shear stress distribution identical to those from

the current polymeric valves. Elevated Reynolds shear stress levels were typically observed

along the edge of the central orifice jet and were spread out over a wider area farther

downstream from the valve as the energy of the central jet becomes dissipated. All of the

aortic valves studied produced Reynolds shear stress in excess of 200 dynes/cm2 during the

majority portion of systole. It is therefore clear that the elevated Reynolds shear stress levels

could lead to sub-lethal and/or lethal damage to the blood elements (Leo, 2005).

Calcification in polymeric heart valves

Calcification remains a major cause of clinical failure in polymeric heart valves

(Akutsu, Dreyer, & Kolff, 1959). Calcification occurs when nodular deposits of calcium

phosphates or other calcium-containing compounds form on cardiovascular devices. There

are several causes of heart valve calcification; (1) Host factors, (2) Fixation conditions, and

(3) Mechanical effects: deposition of calcium phosphates occurs more readily at regions of

stress concentrations in cusps (Vyavahare et al., 1997).

The hemodynamic function of the polymeric heart valves will be affected when

undesired calcium phosphates deposit on the valve surface. This often leads to valve


obstruction and subsequently stiffening of leaflet cusps (Nistal et al., 1990). The most

common mineral deposit found on biomaterials is a poorly crystalline calcium phosphate

known as apatite, which is related to calcium hydroxyapatite, the mineral that is responsible

for the structural integrity in the bones. Calcification was observed. It was also suggested

that the tendency of calcification occurrence may be more related to the presence of small

molecular weight contaminants and oligomers than the bulk polymer itself (Fisher et al.,

1992).

Calcification may occur on regions exposed to high mechanical deformation such as

flexure points in the polymeric valve leaflets. Also calcification may occur on the implant

surface associated with attached tissue or cells or within structural components. Studies have

shown that the commissural region of the valve leaflets have the higher tendency of

calcification, mainly due to the high bending deformations and bending strains that occur in

this region (Corden, David, & Fisher, 1995). In 1987, Glasmacher et al investigated in vitro

calcification behaviour of polyurethane biomaterials but encounter problems in finding

suitable calcifying solution. They also noticed that high mechanical stresses due to cyclical

flexing and other elastomer deformations caused calcification deposits (Glasmacher, Reul,

Rau, & Wieland, 1987). The calcification modelling issue was addressed by Fisher et al by

examining a range of methods of studying the calcification process. Test protocols were

described are proven useful in the early calcification investigation (Fisher et al., 1992).

Calcification in polymeric heart valves can either be intrinsic or extrinsic. If damaged

tissue or cellular components become trapped in these surface defects, they may eventually

calcify leading to extrinsic calcification. However, intrinsic calcification of calcium deposits

formed within the leaflet surfaces have also been observed without any surface defects

(Bernacca, Mackay, Wilkinson, & Wheatley, 1995; Herold et al., 1987). Studies have shown

that surface defects that occur during the fabrication process can lead to extrinsic calcification


whereas material composition has more influence on the occurrence of intrinsic calcification

(Bernacca, Mackay, Wilkinson, & Wheatley, 1997; Bernacca, Straub, & Wheatley, 2002).

In another study, attempts were made to modify the surface of polyurethane valves in

order to resist calcification. Alferiev and colleagues showed that biophosphonate derivatized

polyurethanes improve calcification resistance in polymeric heart valves. Biophosphonates

resist calcification by disrupting the growth of the calcium phosphate crystal (Alferiev et al.,

2001).

In vivo valve performance on animal models

Pre-clinical assessment for polymeric heart valves includes in vivo testing in animal

models. Animal models remain the best means for evaluating anticalcification strategies in

polymeric heart valves. Typical animal studies involve both subdermal (rat) and circulatory

(sheep or calf) implants. Sheep is considered less thrombogenic than human and is widely use

to investigate thromboembolic complications. Meanwhile, a growing calf model represents an

extreme calcification model, and more thrombogenic than sheep, due to their fast growth and

extreme hemodynamic workload (Schoen, Levy, Hilbert, & Bianco, 1992).

In animal studies performed in the 1950s, polyurethane trileaflet valves were

implanted in dogs. However, most dogs only survived for a few hours post operation

(Akutsu, Dreyer, & Kolff, 1959). Following that, studies on calves by comparing various

polyurethane valves with two Ionescus-Shirley bovine pericardial valves demonstrated that

polyurethane valves have similar advantages to bioprostheses in term of thrombus formation

(Herold et al., 1987). However, all explanted valves were calcified and thrombosed, likely

due to the severe hemodynamic and surface alterations by the calcium deposits on the

upstream and downstream sides of the leaflets.

Polyurethane valves were assessed in vivo in a growing sheep model and showed


good comparison with bioprosthesis and other mechanical valves (Wheatley et al., 2000). The

hemodynamic performance is better than the bioprosthesis and lower thrombogenicity when

compared to mechanical valves. The team found that high intensity transient signals (HITS)

were more frequent in the mechanical valve group than in either the porcine or the

polyurethane valve groups. There is also a correlation between the HITS count and platelet

aggregation in the valves, suggesting that these phenomenons are related and that changes in

valve factors, such as reduction in flow orifice area, that tend to increase valve

thrombogenicity could also increase the number of HITS. However, three of the eight valves

showed serious thrombosis, which highlighted a concern as sheep are less thrombogenic than

humans.

In a research conducted by Daebritz and his team, PCU valve prostheses for mitral

position showed great durability and hemodynamic performance in vivo for 22 weeks using a

growing calf model. When implanted correctly, the valve showed minimal degradation of

material and pathological findings showed that animals with PCU valves did clinically better

than those with Perimount and Mosaic commercial valves (Daebritz et al., 2003). Minor

thrombus deposits were observed at the commissural regions of the SPU valves, but the

bioprostheses showed moderate to severe thickening and deformation of the leaflets. One

bioprosthesis in the mitral position was completely covered by a huge thrombus.

Future prospects of polymeric heart valves

Polymeric heart valve designs have improved significantly to address the flow

characteristics and stress reduction. Structural and surface modification of materials also

helped to developed better degradation-resistant materials. Manufacturing methods have also

been improved to allow reproducible polymeric heart valves. Surface modifications of

materials have helped to address the thrombogenicity and calcification issue but it still


remains a challenge in the development of polymeric heart valves to improve the durability.

The hemodynamic performance of the valve can be continued to improve by designing valves

with minimal stresses on the leaflets. Perhaps the biocompatibility and the overall

hemodynamic function and valve durability can be improved by increasing the capability of

the synthetic surface to attract endothelial cells (Ghanbari et al., 2009). Looking forward, the

full potential of polymeric heart valves may be realized through the integration of living cells

into the valve structures.


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Figure 1. An Aortech 23 mm tri-leaflet polymeric heart valve (Leo, 2005).

Figure 2. Computer aided design of the polymeric heart valves; (a) three stent aortic design,

(b) two stent mitral design (Daebritz et al., 2003).

�

Figure 3. Stentless valve prototype with the leaflets in an open configuration (Cacciola,

Peters, & Baaijens, 2000).

Commissure region �

Coaptation region

High central region

Stent

(a) (b)


Figure 4. Velocity fields at the center plane of an Aortech polymeric heart valve during peak

systole. Velocity central orifice jet of 2.3 m/s was observed issuing from the valve orifice

with flow separation occurring at the trailing edge of the leaflet. Plug flow profile was

evident at the immediate downstream of leaflet trailing edge and became more parabolic at

the distal end of the flow chamber (Leo, 2005).

Figure 5. Reynolds shear stress contour plot of the center plane of an Aortech polymeric

heart valve during peak systole. High Reynolds shear stress values coincided with regions of

high velocity gradient typically observed at the edge of the central orifice jet and at the distal

part of the flow chamber (Leo, 2005).

Recirculation zone

Vena contracta

Distal region of flow chamber

��

��


Figure 6. Comparison of flow patterns and leaflet configuration at peak systole among the

three heart valve prostheses: (a) porcine bioprosthesis, (b) pericardial bioprosthesis, and (c)

Aortech polymeric heart valve (Black, Cochrane, Lawford, Reul, & Yoganathan, 1991; Woo,

Williams, & Yoganathan, 1983)

(a)

(b)

(c)

Front view

Side view


Table 1 Comparison of the velocity magnitudes and Reynolds shear stresses from various trileaflet valve designs (Leo, 2005).

Effective annular orifice diameter (mm)

Maximum phase averaged velocity during systole (m/s)

Maximum phase averaged TSS during systole (dynes/cm2)

Carpentier-Edwards 2625 porcine valve (27 mm)

23.0 3.3 4,500

Hancock modified orifice porcine valve (25 mm)

21.8 3.0 2,900

Ionesuc-Shiley pericardial valve (27 mm)

23.4 2.3 2,500

Carpentier-Edwards 2650 porcine valve (27 mm)

25.0 2.0 2,000

Hancock II porcine valve (27 mm)

24.0 2.6 2,500

Hancock pericardial valve (27 mm)

23.3 1.8 2,100

Ionescuc-Shiley low profile pericardial valve (27 mm)

23.0 2.2 2,400

Carpentier-Edwards pericardial valve (27 mm)

25.7 1.8 1,000

Abiomed (21 mm) 18.6 3.7 4,500

Abiomed (25 mm) 22.8 2.2 2,200

Aortech prototype A (23 mm)

23.0 2.8

8,347

Aortech prototype B (23 mm)

23.0 2.5

1,725

Aortech prototype C (23 mm)

23.0 2.3 1,327

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