Prosthetic heart valves: Catering for the few

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Biomaterials 29 (2008) 385–406

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Leading Opinion

Prosthetic heart valves: Catering for the few$

Peter Zilla�, Johan Brink, Paul Human, Deon Bezuidenhout

Christian Barnard Department of Cardiothoracic Surgery, University of Cape Town Medical School and Groote Schuur Hospital,

Anzio Road, 7925 Observatory, Cape Town, South Africa

Received 2 August 2007; accepted 23 September 2007

Available online 24 October 2007

Abstract

Prosthetic heart valves epitomize both the triumphant advance of cardiac surgery in its early days and its stagnation into a

retrospective, exclusive first world discipline of late. Fifty-two years after the first diseased heart valve was replaced in a patient,

prostheses largely represent the concepts of the 1960s with many of their design-inherent complications. While the sophisticated medical

systems of the developed world may be able to cope with sub-optimal replacements, these valves are poorly suited to the developing

world (where the overwhelming majority of potential valve recipients reside), due to differences in age profiles and socio-economic

circumstances. Therefore, it is the latter group which suffered most from the sluggish pace of developments. While it previously took less

than 7 years for mechanical heart valves to develop from the first commercially available ball-in-cage valve to the tilting pyrolytic-carbon

disc valve, and another 10 years to arrive at the all-carbon bi-leaflet design, only small incremental improvements have been achieved

since 1977. Similarly, bioprosthetic valves saw their last major break-through development in the late 1960s when formalin fixation was

replaced by glutaraldehyde cross linking. Since then, poorly understood so-called ‘anti-calcification’ treatments were added and the

homograft concept rediscovered under the catch-phrase ‘stentless’. Still, tissue valves continue to degenerate fast in younger patients,

making them unsuitable for developing countries. Yet, catheter-delivered prostheses almost exclusively use bioprosthetic tissue, thereby

reducing one of the most promising developments for patients of the developing world into a fringe product for the few first world

recipients. With tissue-engineered valves aiming at the narrow niche of congenital malformations and synthetic flexible leaflet valves

being in their fifth decade of low-key development, heart valve prostheses seem to be destined to remain an unsatisfying and exclusive

first world solution for a long time to come.

r 2007 Elsevier Ltd. All rights reserved.

Keywords: Heart valve; Bioprosthesis; Calcification; Immune response; Degradation; Thrombogenicity

1. Introduction

The emergence of heart surgery was closely linked to thetechnological advances of the post-second world warperiod. Pioneering spirit and innovative academic mindsbroke new ground on an almost daily basis. Once theheart–lung machine was invented, it did not take long until

e front matter r 2007 Elsevier Ltd. All rights reserved.

omaterials.2007.09.033

ote: Leading Opinion: This paper is one of a newly

of scientific articles that provide evidence-based scientific

pical and important issues in biomaterials science. They

tures of an invited editorial but are based on scientific

e features of a review paper, without attempting to be

These papers have been commissioned by Editor-in-Chief

r factual, scientific content by referees.

ing author. Tel.: +2721 406 64 76; fax: +27 21 448 59 35.

ess: [email protected] (P. Zilla).

prosthetic heart valves were implanted into patients [1,2].By the second half of the 1970s, these crude initialprototypes had already evolved into the valve prostheseswe know today: bioprosthetic tissue (BPT) valves had longundergone the quantum leap from unstable formalintreatment [3] to the far superior glutaraldehyde fixation[4] and mechanical prostheses had moved from the originalball-and-cage valve [5] to tilting disc [6] and even bi-leafletdesigns [7]. Against this background it is disappointingthat, more than three decades later, commercial productsare still based on the largely unchanged concepts of the1970s. Glutaraldehyde-fixed tissue valves have been onlymarginally improved by eliminating design flaws andadding so-called ‘anti-calcification’ treatments. Mechanicalvalves are still mostly of the bi-leaflet design andimprovements were restricted to better-polished pyrolytic

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ARTICLE IN PRESSP. Zilla et al. / Biomaterials 29 (2008) 385–406386

carbon coatings or some minor reductions in turbulences inthe crucial hinge area. In contrast to the pioneering years,courage and incentive towards a new developmentalquantum leap have been mostly absent from commercialproduct strategies.

2. Population needs and obstacles

Unfortunately, socioeconomic circumstances still limitaccess to valve surgery to a fraction of the patients whoactually need it. While the bulk of the estimated 275,000 [8]to 370,000 [9] annual valve replacements benefit predomi-nantly the older patients of the First World, developingcountries with their much higher incidence of rheumaticfever [10] often have no access to heart surgery. Related toEuropean service levels, for instance, cardiac surgery isonly available to 8.1% of the Chinese and 6.9% of theIndian population [11]. However, given the fast growingeconomies of some of these countries, it is predictable thatthey will soon have a high demand for prosthetic heartvalves that are affordable and that address the specificneeds of their mainly young patients. Since performancedemands, complications and degeneration distinctly differbetween age groups, the largely geriatric experience profilefrom developed regions [12] does not allow generalizedconclusions as to ‘‘what is the best replacement heartvalve?’’ for recipients in threshold countries. The often-deadly complications associated with mechanical valveprostheses accumulate over patient years. In developing

Fig. 1. Typical age-distribution of patients undergoing heart valve replacemen

recipients in a First World population are predominantly in the age group of 6

from 20 to 70 years in a Developing Country such as South Africa (blue line)

Groote Schuur Hospital (University of Cape Town) shows, a significant prop

countries these complications are multiplied by a lack ofcompliance with anti-coagulation due to infrastructuraland educational deficits [13,14]. Together with the youngerage of patients (Fig. 1) the individual likelihood ofexperiencing a thrombotic or thromboembolic disaster inthe course of a life-time is therefore distinctly higher thanin first-world patients [13,14]. Unfortunately, the alter-native use of contemporary tissue valves is equally over-shadowed in these patients due to the fast degeneration ofbioprostheses in young recipients.Taking all these factors into account, it is obvious that:

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The overwhelming majority of potential recipients ofprosthetic heart valves reside in developing countries.The fact that 85% of all open-heart procedures areperformed in countries representing 11% of theworld population emphasizes how unrepresentative thecurrent heart valve recipients are with regard to globalneeds [15].
� The fact that the economies of many of these threshold
countries are growing fast makes it predictable thatcardiac surgery will be increasingly accessible to thosemillions who are in need of a valve replacement. Sinceimproved socio-economic circumstance will only have adelayed impact on reducing the incidence of rheumaticfever, a huge demand for valve replacements will exist inthese ‘newly developed’ countries for a relatively longtime.
� The unattractive profit margins for ‘first-world-pros-
theses’ in developing countries as well as the relatively

the First World and in a Developing Country. While prosthetic valve

9 years (red line) they are broadly disseminated over an age-spectrum

s the age distribution of 2000 consecutive heart valve recipients at the

on of patients is even younger than 20 years.

ARTICLE IN PRESSP. Zilla et al. / Biomaterials 29 (2008) 385–406 387

small market in developed countries, together with theexceptionally long ‘bench-to-profit’ delays for heartvalve prostheses have hampered major research invest-ments in this field during the past three decades.
� Inasmuch as locally manufactured prosthetic heart
valves will help to partially overcome the affordabilityfactor for threshold countries, they will be replicas ofvalves which were designed to address the specificcircumstances of first world patients.
� Therefore, the discrepancy between the world’s needs
and industry’s commitment will continue to widen formany more years: an ever increasing number of youngpatients from threshold countries in need of a newconcept of a truly long-lasting valve prosthesis on theone side as opposed to the regardless pursuit ofcontinuously stepped-up sale-efforts of non-innovative,long-established products of yesterday on the other.A good example of the latter is the ‘sales-speak’ of‘actual’ freedom from structural valve failure as opposedto ‘actuarial’. While the actuarial median failure time ofbioprosthetic valves in elderly first-world patients isapproximately 14 years [16], the ‘actual’ failure-rateafter the same period of time as reported by industry-supported research is 8% in AVRs and 11% in MVRs[17]! As a damage-control measure, the Journal of HeartValve Disease announced in an Editorial its intention toreject any manuscripts in the future that report ‘actual’rather than ‘actuarial’ data [18].

For the time being, we need to accept that the limited
market for valve prostheses in today’s world of high profitswith mass-selling low-tech products such as coronary stentswill continue to perpetuate a situation whereby heart valveresearch is too cumbersome and unattractive for the mainplayers to sufficiently invest in truly disruptive technologies.This is opposed by a flurry of innovative ideas byUniversity-based entrepreneurs and start-up companies.However, the particularly long and costly lead-time betweenproduct development and marketing as well as the moderateprofit margins will unfortunately prevent many of theseinnovations from becoming a mainstream product. Untilthen, the hype around the endovascular delivery of yester-day’s concepts may ironically and unintentionally turn intoa success story for the huge number of potential recipients ofheart valve prostheses in threshold countries: In contrast tothe heavily calcified heart valve pathology in the first world,young patients with rheumatic valves almost naturally offerthemselves as ideal recipients of endovascularly or transa-pically placed prostheses as their diseased native valves aremostly delicate and non-calcified. With the incentive of apotential mass market, durability issues in young patientsmay eventually be confronted.
3. Mechanical heart valves

The groundbreaking initial years in the development ofprosthetic heart valves are tightly linked to mechanical

prostheses. A ‘mechanical’ problem such as a leaking valveof a ‘pump’ naturally seemed amenable to an ‘engineering’solution. Without this ability to exploit concomitantdevelopments in engineering and material sciences, pre-vious pioneers were doomed to fail. Ten years after the firstknown operation on a heart valve by Harvey Cushing atthe Peter Bent Brigham hospital in Boston in 1913, ElliottCarr Cutler and the cardiologist Samuel A. Levineattempted a surgical valvotomy in a young female patientwith mitral valve stenosis. This operation, hailed as amilestone by the British Medical Journal, had a subsequentmortality of 90% and was soon abandoned.In stark contrast to this rather gloomy era, the heady era

of replacement valves was the decade from the early 1950sto the early 1960s, when break-through developments inscience and engineering were vigorously applied to clinicalneeds. Insight into the physiology of hypothermia allowedFloyd Lewis of Minnesota to gain open access to an atriumin 1952 followed by Charles Hufnagel’s implantation of acaged-ball heart valve into the descending aorta of tenpatients (six survived the operation) in the same year. By1953, electrical engineering had provided the pacemaker,and the miniaturization of industrial technologies hadallowed John Gibbon Jr. of Philadelphia to introduce theheart lung machine, thereby for the first time allowingunlimited open access for the surgical replacement of aheart valve on July 22, 1955 in Sheffield, UK. The epitomesof this pioneering spirit were the use of synthetic materialsfor the first mitral valve replacement by Nina Braunwald in1960 and the joint effort of the surgeon Albert Starr andthe electrical engineer Miles ‘‘Lowell’’ Edwards—founderof the medical device company ‘American Edwards

Laboratories’ in 1950—to improve the ball-and-cage valveto a level where it could reliably be used as a replacementheart valve (Fig. 2). The first human implant of this valveon September 21, 1960 concluded the most vibrant decadeof development in this field. During the subsequent 15years major insights into fluid dynamics, material sciencesand engineering were still carried over into the improve-ment of replacement heart valves, but they were no matchfor the quantum leap of the preceeding era. The two crucialelements of these post-pioneer years were the adaptation oftechnologies that allowed the production of tilting discs orsemi-discs (Fig. 2) as occluders and the utilization ofmaterials developed in the course of atomic energyresearch. The scientist J.C. Bokros from the ‘General

Atomic Company’ had previously investigated pyrolyticcarbon materials for the coating of nuclear fuel particles.Since this material’s atomic microstructure was found toincrease resistance to cracking and distinctly loweredthrombogenicity, it was introduced into heart valveprostheses. For more than three decades all commercialmechanical heart valves have either been using it forsurface coating (by depositing gaseous hydrocarbons—usually methane—onto a heated graphite substrate attemperatures of between 1800 and 2300 1C) or for entireleaflets and housings of the valves (made from 100%

ARTICLE IN PRESS

Fig. 2. The development of mechanical heart valves experienced an

exponential slow-down over it’s more than 50-year history: while it took

only 7 years from the first commercially available ball-in-cage valve to the

era of pyrolytic carbon-coated tilting disc valves and 10 years from the

latter to the first all-carbon bileaflet design, progress of the past 30 years

was incremental at most with leaflets being either more parallel aligned or

mildly curved and hinge mechanisms slightly less thrombogenic.

P. Zilla et al. / Biomaterials 29 (2008) 385–406388

pyrolytic carbon using fluidized bed processes). Overall,mechanical heart valve developments were largely con-cluded in 1977 with the introduction of an all-pyrolyticcarbon bileaflet valve (St. Jude Medical/SJM). Whatfollowed was attention to engineering detail rather thangrand new concepts. Not surprisingly, the one aspect ofengineering which represented the monopoly componentof production, pyrolytic carbon, became the centre ofcorporate interest and ‘merry-go-round’ acquisitions: first,Dr. Bokros of the ‘General Atomic Company’ helped to setup their medical division under the name ‘‘Carbomedics’’.At one stage, ‘‘Carbomedics’’ manufactured 17 heartvalves for 16 different companies! Subsequently, Dr.Bokros split off ‘‘Carbon Implants’’. When Medtronicbought it in 1992 to obtain their ‘‘Parallel Valve’’, Bokrossplit off the ‘‘Medical Carbon Research Institute’’ (MCRI)in 1994, which designed the On-X valve and later changedthe name MCRI to On-X. At the same time, ‘‘Carbome-dics’’ (Austin, TX) itself was bought by ‘‘Sulzer Medica’’ ofSwitzerland in the late 1980s, and then in turn by ‘‘Sorin’’,the medical division of Fiat (the Italian automobilemanufacturer), which eventually transferred the produc-tion from Austin to an FDA approved site in Milan, Italy.The incestuous nature of corporate acquisitions in thisfield is further highlighted by another pioneer ending inItaly: After Pfizer took over Shiley, the very durableBjork-Shiley tilting disc valve was re-designed, leading toone of the biggest disasters in the history of medicalimplants when many of the convexo-concave Bjork-Shileyvalves developed mechanical failure [19–21]. EventuallyPfizer sold Shiley to Sorin, but Sorin chose to exclude therights to the failing ‘convexo-concave’ valve from thispurchase.

3.1. Mechanical durability

Out of 86,000 convexo-concave Bjork-Shiley tilting-discvalves, 619 experienced a fracture of the outflow strut,which led to the patient’s death in two-thirds of the cases.On a smaller scale, 46 leaflet escapes were reported out of200,000 implants of the Edwards–Duromedics bileafletvalve [22]. As unsettling as these failures were, and still are,for the many patients living with one of these prostheses(e.g. 23,000 Bjork-Shiley valves in North America alone),they provided crucial insight into key elements of the threemajor challenges of engineering, namely closing load,material fatigue and cavitation.As far as the first is concerned, the major load on

mechanical heart valves generally arises from transvalvularpressure generated at and after valve closure leading toboth ‘impact wear’ and ‘friction wear’. Impact wear usuallyoccurs between occluder and ring in tilting discs and thehinge regions in bileaflets. At the same time, friction wearoccurs between occluder and struts in tilting discs andbetween leaflet pivots and hinge cavities in bileaflets. In thecase of the Bjork-Shiley valve, the damage occurred on theoutlet strut, causing bewilderment as to how the closingload could damage the outlet strut. In 1984, Shiley itselfprovided the answer by calling the discovery of a transient(o0.5ms) occlusion impact on the outlet strut tip due toover-rotation of the disc ‘bimodal closure phenomenon’.This force was ten times higher than the actual openingimpact, causing bending stresses that exceeded the strutwire’s fatigue endurance limit. In Duromedic’s case, wherethe leaflet’s seat was created through a lip or shelf on thehousing (rather than a cavity in the housing like in all othermodels), erosion contributed to the breakages. All theseinsights were addressed by modern strut- and hinge-designs, some of them (Bjork-Shiley) by slightly increasingthe outlet strut clearance, some of them (On-X) by evenclaiming a ‘two-point landing system’ that reduces the loadimpact.The second most important cause of failure, material

fatigue as a consequence of the polycrystalline character-istics of metals, was successfully addressed by a gradualshift from metal alloys to pyrolytic carbon, a material thatbelongs to the ‘turbostatic carbons’ and thus has a similarstructure to graphite. Initially, strength and wear resistanceof pyrolytic carbon were increased by adding up to 8% ofits weight as silicon, co-deposited from silicon carbides, butrecent improvements by the same group, which has beendriving pyrolytic carbon development for the past 40 years,led to what they call ‘pure carbon’. Allegedly, theelimination of the silicon component resulted in a 50%stronger material. The switch to carbon was first pioneeredin the occluder discs when pyrolytic carbon-coated graphitereplaced the original Delrin disc in Bjork-Shiley valves in1968 (first implanted in 1971). Since then, all manufac-turers have been using this carbon-combination for theproduction of their mono/bileaflet discs except for MCRI(On-X) which introduced leaflets made of solid, silicon-free


pyrolytic carbon in 1996. In contrast to leaflets, thehousings and struts were made out of metals until wellinto the 1980s. Initial attempts at overcoming the fatigue ofmetals involved either using alloys or particularly strongpure metals such as titanium. ‘Starr-Edwards’ ball-and-cage valves were the first to replace the original stainlesssteel by a cobalt–chromium–molybdenium–nickel-alloy, amaterial that was later also used for the Bjork-Shileyvalves. The introduction of titanium to mechanical heartvalves also took place during the days of ball-in-cage valveswhen pure titanium was used for the Smeloff-Cutter valveas early as in 1965 [23]. Later, tilting-disc designs such asthe Medtronic Hall (1977), Sorin Monoleaflet (1977),Lilleihei-Kastner (1978) and Omniscience (1978) valvemachined their housings out of a solid block of titanium.Eventually, the introduction of the bileaflet design in 1977coincided with the advent of the ‘all carbon’ valves whereboth leaflets and housing were made of pyrolytic carbon.This bileaflet concept has continuously dominated themarket for the past 30 years: St. Jude Medical (SJM; since1977; 800,000 sold); Duromedics (1982–1988; 200,000 sold[24]; reintroduced in 1990); Carbomedics (introduced in1986 in France, 1993 in USA; 500,000 sold by 2004);Advanced Tissue Science (ATS; since 1992; 135,000 sold);Medtronic (since 1999, after ‘Parallel’ valve was boughtfrom ‘Carbon Implants in 1992) and On-X (since 1996outside USA, 2001 USA; 50,000 sold). Yet, Titanium is stillpart of the housing in some designs (e.g. Lockring andLockwires in ATS and the housing body in Sorin Bicarbonwhich later became the Edwards Mira in 1997). Overallit appears as if material science at least has provided uswith hardly improvable, optimal materials for mechanicalheart valves.

The third most important mode of failure is ‘cavitation’.This well-known phenomenon is defined by the rapidformation of vaporous microbubbles in a fluid due to alocal reduction of pressure below vapour pressure—similarto boiling [25]. When the pressure conditions for ‘cavita-tion’ are present, bubbles will start to form and grow.When the pressure recovers, the formed bubbles willimplode, producing pressure and thermal shock wavesthat can impinge on solid surfaces. Pressure oscillations,flow decelerations, tip vortices and ‘squeeze jets’ are allcapable of inducing cavitation, the last one being the mostcontributive factor. Squeeze jets are formed when the valveis closing and the blood between the occluder and the valvehousing is ‘squeezed’ out to create a high-speed jet. This inturn creates vortices that can lead to additional cavitation.Damages found on failed Duromedics valves, for instance,were associated with confined areas of erosion and pittingon leaflets and housing [26] caused by cavitation.

Overall, the disasters of the 1970s and 1980s seem tohave helped eliminate mechanical failure. The fact that nota single failure was reported on later designs (e.g. after halfa million Carbomedics implants) indicates that combinedefforts were successful. Part of these efforts is therequirements for much more extended animal tests and

clinical trials before obtaining the European CE Mark andAmerican FDA approval. Unfortunately, this rigorous pre-market testing added significantly to the research anddevelopment costs of the prostheses. The resulting costincreases of mechanical heart valves added anotherobstacle to the widespread use of these prostheses indeveloping countries.

3.2. Weighing life and death

The Bjork-Shilry ‘‘convexo-concave debacle’’ did notonly help to identify the weak points of mechanical heartvalve engineering but also taught us to deal calmly withpotential failure by balancing the risk of valve fractureagainst the risk of surgery to replace the valve. The decisionof both the University of Alabama and the Mayo Clinic toremove every concavo–convex valve may not hold retro-spective scrutiny. An average re-operation mortality of 8%may well outweigh the actually occurred 0.9% incidence ofstrut breakage, even if risk-stratification takes the threeparticularly prone groups into account which were retro-spectively identified [27]: (a) 33mm mitral valves, forinstance, were 23 times more likely to fracture than21–25mm aortic valves as were (b) valves with moreflexible outlet struts, as determined by hook deflection andload deflection tests during manufacture and (c) valvesproduced by one particular welder.

3.3. Biological limits of mechanical solutions

Given the fact that both the ‘fluid’ flowing through thevalves and the ‘pump’ into which these valves areincorporated consist of living cells, problems are likely tobe augmented by cascades of biological events. Fluid shearforces and turbulences, for instance, do not only lead toenergy losses but to an activation of the platelet andcoagulation system which has the potential to immobilizethe entire valve or embolize into vital organs. Similarly,small valvular orifices do not only lead to higher pressuresin the outlet chamber but to a remodelling process of thecontractile tissue that may eventually lead to irreversiblepump failure.

3.3.1. Effective orifice area (EOA)

Overall, it appears as if minimizing the energy requiredto eject the blood through the valve has been moresuccessful than eliminating the activation of the coagula-tion cascade. The ‘effective orifice area’ (EOA) is a meansof expressing the degree to which a prosthesis impedesblood flow through the valve and thus increases the energyloss. In contrast to the energy efficient central flow ofnatural heart valves, caged-ball valves completely blockedthe central flow. While the first tilting disc valve (Melrose)in 1964 introduced some degree of central flow, the majorbreak-through occurred in 1977 with the first bi-leafletvalve. The magnitude of this development is best expressedby the increase of EOA from 1.5 to 2.1 cm2 in tilting disc


valves to 2.4–3.2 cm2 in bileaflet designs of the same outerdiameter. Again, improvements after 1977 were incremen-tal rather than disruptive. By moving the sewing ring supraannularly, for instance, additional increases of the innerdiameter were possible (e.g. 17% On-X; Carbomedics; etc).Furthermore, by successively aligning the leaflets parallelwith the blood stream, unobstructed central flow becamealmost a reality. Again, most of the progress was alreadymade before 1978 when the tilting disc opening-angle of 601of the first Bjork-Shiley valve was increased to 751 (inaortic; 701 in mitral valves) in the Medtronic Hall in 1977and to 801 in the Lilleihei-Kaster valve (later Omniscience)in 1978. The introduction of the bileaflet design onlyincreased this angle to 851 (SJM 1977), with some makeseven having a lower opening angle than tilting discs(Duromedics 771/731 in aortic/mitral valves, respectively,and Carbomedics 801/781). Even with a 901 opening angle(Medtronic Parallel; On-X), however, the actual travel arcthrough which the leaflets swing is/was mostly less than 601(e.g. 551/601 SJM; 531/551 Carbomedics; Duromedics581/621 and 501 Medtronic Parallel), resulting in a steeperclosing angle and thus carrying the potential of a higherregurgitation volume.

3.3.2. Thromboembolic risks

Mechanical heart valves remain vulnerable to thrombusformation due to high shear stress (one of the strongestplatelet activators), flow separation/stagnation (sinceVirchow’s days the dreaded cause of coagulation) andblood damage (through the release of pro-coagulantsubstances). As a consequence, anti-coagulation therapyhas been a hitherto sine qua non. In general, thromboem-bolism (e.g. stroke) is more pronounced in mitral valvereplacements, and within this group, more so in patientswith large left atria and in those in atrial fibrillation and/orcardiac failure [28,29].

In tilting-discs, flow separation occurs behind the valvestruts and distal to the leaflet edge as a result of acombination of high velocity and stagnant flows. Of alldesigns, the largest turbulent stresses of about 150 Pa werefound behind tilting disc valves in the minor flow regionparallel to the tilt axis [30]. The bileaflet models have highstress during forward flow and leakage regurgitation aswell as adjacent stagnant flow in the hinge area. As it turnsout, the hinge area is the most critical part of bileaflets andis where the thrombus formation usually commences.

Prosthetic valve-induced clot formation-triggered bytissue factor exposure, platelet activation and contactactivation by foreign materials—occurs in three steps:initiation, amplification and propagation. Tissue factor(TF) release usually occurs through cell rupture with bloodtrauma playing a central role. High stresses during leakageflow in aortic valves and forward flow in mitral valves leadto blood damage releasing both TF and the platelet-activating ADP. Plasma factor VII binds to TF, setting offa chain reaction which activates factor Xa and Va whichbind to each other to produce thrombin which in turn

activates platelets and factor VIII. In parallel, plateletactivation is triggered when shear stresses reach a levelhigher than 60–80 dyn/cm2, followed by von Willebrandfactor binding to the platelet receptor GPIb and calciumrelease. GP IIb–IIIa facilitates platelet–platelet adhesion.Last, but not least, contact activation begins when factorXII binds to the non-endothelialized artificial surfaces ofprostheses. This in turn activates prekallikrein and high-molecular weight kininogen. Although ‘back-wash’ designshave aimed at reducing the thrombus formation in thehinge area [31–35], mechanical prostheses still require anti-coagulation and the promise of freedom from thromboem-bolism is still elusive and remains the Achilles heel of allmechanical prostheses, even in the first world. Althoughthe incidence of thromboembolism has changed from the4.5% per patient year [36–41] in the days of the ball-in-cagevalves, it is still 0.6–2.5% in Omnicarbon tilting discvalves [42,43]; 1.3–4.2% in Medtronic Hall valves[13,32,33,37,44–47]; 1.1–3.7% in SJM valves [14,48–54];0.6–4.3% in Carbomedics valves [55,56]; 0.0–3.3% in ATSvalves [57,58] and 1.8% in On-X valves [59]. The mostlethal and debilitating forms of thromboembolism, how-ever, is thrombotic prosthetic obstruction. While thepreviously reported estimates for the tilting disc valveswere between 0.1% and 1.2% per patient year in the firstworld, there is general recognition that this became anextremely rare complication in the era of bileaflet valvesunless anti-coagulation had been stopped. Therefore, giventhe unreliable degree of anti-coagulation in developingcountries, it is not surprising that thrombotic valveobstruction is still a feared occurrence outside thedeveloped world. Unfortunately, the true number ofpatients dying as a consequence of valve clotting will oftenremain unrecognized in these geographic regions due to thelack of centralized death registries and cultural resistancesto autopsies.

3.3.3. Coumadin-related haemorrhage

The necessity for anti-coagulation with Coumadin(Warfarin) that minimizes, but never completely eliminatesthe risk of thromboembolism also exposes patients to theadditional risk of major haemorrhagic complications. Themost lethal and debilitating of these is intracranialhaemorrhage, which occurs at a rate of between 0.8%and 3.7% per year [60]. Although INR requirements havecome down from 3.5–4.5 in ball-in-cage-valves to 2.5–3.5,and anti-coagulation-caused haemorrhages are signifi-cantly fewer being reported by various authors [61,62],the numbers are still worrying. While they were reported tobe between 0.6% and 3.7% per patient year in StarrEdwards valves [38,40], they were 0.0–2.7% in Omnicarbonvalves [63–65], 1.4–3.2% in Medtronic Hall valves [47,66];0.3–2.8% in SJM valves [53,54], 0.0–2.8% in Carbomedicsvalves [53,67], 0.0–4.9% in ATS valves [53] and 0.0% inOn-X valves [59]. Naturally, these complications occur at adistinctly higher rate in the developing world where theability to ensure good anti-coagulation control is very


limited and often impossible due to poor socio-economiccircumstances [13,14,68–70]. Typically, half of thesepatients presenting with obstructive valve dysfunction havean INR of 1.4 or less at the time of failure [71].

3.3.4. Pannus overgrowth

Excess tissue growth across the sewing ring can leadeither to a narrowing of the orifice or to leafletimmobilization. Although widely underestimated, it islikely the primary cause of obstructive prosthetic valvefailure in patients on correct anti-coagulation therapy.Pure pannus overgrowth is estimated to be behindobstructive valve failure in between 31% [72] and 53%[73] of cases. Even in geographic regions where the INRcontrol is sub-optimal, the reason for obstructive valvefailure is still overwhelmingly pannus rather than thrombusrelated [71]. In the aortic position, pannus formation ismainly on the inflow side [74] while in the mitral position itoccurs both on the atrial and the ventricular side [71].In order to prevent tissue ingrowth into the clearance ofleaflets, one contemporary valve design introduced a longerhousing cylinder with the goal of creating an ingrowthbarrier.

Although pannus formation is a complex event, inflam-mation certainly contributes to a sustained growth signalfor the hyperplastic tissue. One of the sources of growthstimulation is the chronic foreign body reaction against thesewing ring material. The dominant cells in this reactionare macrophages and giant cells which are known to secretean extended range of powerful cytokines and growthfactors, including IL-1, PDGF, bFGF, etc. [75].

3.4. Way forward for mechanical valves

Modern investigative technologies (microvascular flowvisualization, computational fluid dynamics modelling,laser Doppler velocimetry and anemometry measurements)have contributed to the latest bileaflet hinge pocket designwith the goal of further reducing turbulences and pro-coagulant ‘pockets’. The two most recent bileaflet valves onthe market, the On-X and the Medtronic Advantage valve,have been designed and evaluated using these modalities[76]. This has already resulted in one study reportinglowered rates of valve thrombosis and thromboembolismin a third world population [77]. Whether attention todetail such as lower turbulence levels and better polishedmaterials may eventually allow the safe implantation ofmechanical valves into the young patients of developingcountries, regardless of their poor anti-coagulation control,will need to be elucidated in large, dedicated studies. Onesuch study pursued by SJM has just been approved.Another reduced anti-coagulation trial by On-X, usinglowered anti-coagulation targets in the mitral position andonly Aspirin (without Coumadin) in the aortic position,has also been approved. In the meantime, an incrementallowering of the INR-bar may be on the cards for manyof the current prostheses. Most recent guidelines for the

anti-coagulation of mechanical heart valves for the firsttime ventured into the ‘grey zone’ of INRsp2.2 [78].

4. Tissue valves

Bioprosthetic heart valves were the answer to thethromboembolic complications of mechanical valves. Sincecommercial tissue quantities could be only obtained fromanimal sources, antigen masking through cross-linking wasan integral part of the concept from the beginning.In general, few improvements have been added since the

introduction of glutaraldehyde fixation in the late 1960s [4]and its translation into commercial valves (Hancock 1972;Carpentier Edwards Porcine 1970/1976). Other than theelimination of design flaws such as ill-placed suspensionstitches (Ionescu Shiley 1976) [79] or over-zealous efforts toreduce the valve profile (BioImplant) [80], poorly under-stood ‘anti-calcification’ treatments were added [81–83]. Atthe same time, market-driven preferences for either bovinepericardial or porcine aortic valves cultivated a pseudo-sense of progress amongst surgeons while in fact theymerely reflected the circumstances of the time rather thanscientific insight. While in previous eras clinical failures dueto design flaws affected the surgeon’s preferences forbovine or porcine products, prion and virus concerns aswell as tissue availabilities and production advantagesdetermine today’s marketing thrusts towards ‘bovine’ or‘porcine’ valves. Last, but not least, the much-hailed ‘new’concept of ‘stentless-valves’ (SJM Toronto SPV 1991;Edwards Prima 1991; Medtronic Freestyle 1991; BioCor1994), which emulated the homograft concept of the 1960sgrossly overstated the biomechanic advantages. Slowlyemerging mid-term results confirm the largely unfulfilledexpectations [84]. The fact that stentless roots do notcalcify more than homograft roots [85] is less a reflection ofthe success of stentless valves than the degree of degenera-tion occurring in homografts.

4.1. Insight into failure modes

The absence of living tissue, as well as the necessity tomask antigenicity, represents the system-inherent problemsof contemporary tissue valves. Instead of recognizing thesecore issues and addressing them from the beginning,however, there was a widely prevailing perception thatimmunogenicity non-vital tissue was acceptable as long asthe right ‘anti-calcification’ treatment was found. On thebasis of these two assumptions, bioprosthetic heart valveresearch focused almost exclusively on ‘down-stream’problems such as the intrinsic potential of glutaraldehydeto trigger calcification although it was well knownthat calcification affects less than half of failed tissuevalves while tears, as a consequence of inflammation,collagen degradation and a lack of repair mechanisms werethe predominant modes of failure [86,87]. Eventually, therole of inflammation was revisited and a link betweenremnant-immunogenicity and tissue degeneration was

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Fig. 4. Support for a possible role of graft-specific IgG in the calcification

of bioprosthetic tissue provided by atomic absorption spectroscopy

analysis of calcium levels in subdermally implanted glutaraldehyde-fixed

porcine aortic wall tissue in the rabbit. An almost three-times higher level

of calcification was found in bioprosthetic tissue that was exposed to

serum containing graft-specific antibodies Ref. [91] with permission.


established [88], but recognition and implementation areslow.

4.1.1. Remnant tissue immunogenicity

It has long been known that the low-dose glutaraldehydetreatment used for the fixation of bioprosthetic heart valvesreduces immunogenicity but does not abolish it (Fig. 3)[89]. The low-grade fixation used in commercial valvepreparations fails to significantly alter membrane-boundreceptors or structural glycoproteins [4]. Therefore, thetissue continues to elicit both cellular and humoral immuneresponses [90]. By increasing crosslink density, tissueantigenicity could be reduced [88,89,91] and the humoralantibody-response mitigated. Most importantly, by de-monstrating a link between immune response and calcificdegeneration, a key aspect of bioprosthetic degenerationcould be clearly traced to insufficiently masked immuno-genicity [88,92–95] (Fig. 4). After identifying porcinefibronectin as a major persisting antigen (Fig. 3) [96], themain challenge ahead will be to identify more tissueantigens and elucidate the exact mechanisms throughwhich antibody binding facilitates degeneration.

4.1.2. Inflammatory degradation

In the vast majority of explanted tissue valves, inflam-matory cells are either found to cover the surfaces [97] orfocally infiltrate into the tissue [98]. Since the initial signs ofxenograft rejection in the form of polymorphnuclearinfiltrates [99,100] give way to a more macrophage- andforeign body giant cell-dominated phenomenon, a percep-tion of harmlessness ensued. Given the destructive poten-tial of giant cells to erode even synthetic materials,however, it becomes obvious that these ‘war-formations’

Fig. 3. Demonstration of the need for increased crosslink density to

mitigate residual immunogenicity of glutaraldehyde fixed porcine aortic

wall tissue. Commercially applied concentrations of the dialdehyde (up to

1.0%) failed to avoid an IgG response to porcine bioprosthetic tissue—in

particular the glycoprotein fibronectin—in a rabbit subdermal implant

model. Only glutaraldehyde concentrations in excess of 1.0% (here 3.0%

ideally with L-lysine extension) were able to quench the response

(Densitometry plot of a porcine Fibronectin Western blot using 1:100

dilutions of rabbit sera) Ref. [91] with permission.

of macrophages are serious culprits of destruction. Inbovine pericardial heart valve prostheses, for instance,macrophages are regularly found invading and focallydegrading the prosthetic collagen [90,101] (Fig. 5). Simi-larly, clinically implanted porcine valves also showinflammatory cells inside the BPT [97]. Conspicuously,dense infiltrates of inflammatory cells are found in valveswhich had failed due to tissue tearing [100,102]. Macro-phage-mediated tissue degradation is further evidenced bythe direct proof of collagen phagocytosis [99,103]. Inclinical series, as many as 82% of failed valves showedsigns of collagen phagocytosis by macrophages in trans-mission electron microscopy [104] (Fig. 6).

4.1.3. Mechanical damage/lack of repair

The consequences of BPT being crosslinked and non-vital—and thus deprived of repair mechanisms—are mostobvious when it comes to mechanical wear and tear: in themechanically more stressed mitral position, as many as75% of failed porcine prostheses show a rupture of a freecusp edge and 43% of the cusp belly [105]. In vital, nativevalves both a macromechanical and a micromechanicalprinciple reduce the impact of mechanical load. Macro-mechanically, the distinct dilatation of the annulus duringsystole results in the triangular stretching of cusp tissue,thereby avoiding acute bending at the commissures beyond601 [106]. Micromechanically, a sophisticated tri-layeredultrastructural architecture ascertains optimal stress reduc-tion whereby the lamina spongiosa provides a ‘sliding gap’between the relatively smooth lamina ventricularis on theoutflow side and a folded lamina fibrosa on the inflow side.Bioprosthetic heart valves—stented or stentless—defyboth these principles of valve mechanics: by either

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Fig. 5. Torn leaflet of a pericardial prosthesis leading to acute mitral

regurgitation 56 months after implantation. All retrieved tissue samples

showed significant inflammatory infiltration into the depth of the leaflet.

Although the inflammation was macrophage-dominated [(a) CD 68/Ham

56; 40� ] there was also significant involvement of granulocytes

dominated [(b) neutrophil elastase; 20� ].

Fig. 6. Typical transmission electron micrograph of a macrophage

infiltrating into the depth of a porcine bioprosthetic heart valve. The

phagocytotic vesicles containing remnants of collagen are evidence that

these macrophages are no ‘innocent bystanders’ but rather active players

in the degeneration of bioprosthetic tissue.

P. Zilla et al. / Biomaterials 29 (2008) 385–406 393

stent-mounting, or through glutaraldehyde fixation, annu-lar dilatation during systole is either eliminated ordramatically diminished. As a consequence, the bendingangle becomes acute and reverse curvature and bucklingensues. By eliminating the ‘sliding gap’ between theventricularis and the fibrosa as a result of crosslinking,a major means of stress distribution is lost.

4.1.4. Pannus overgrowth

Inasmuch as similar triggers initiate the tissue over-growth of the sewing ring as in mechanical valves, theprocess is augmented by the inflammatory potential of theBPT. Amongst other factors, the extent of this inflamma-tion is a reflection of the degree of insufficient immunemasking of the BPT. A clear correlation between pannus

overgrowth onto leaflet tissue and cross-link density hasbeen shown [96].

4.2. Scope for improvements

4.2.1. Anti-calcification treatments

Although the unifocal attention to intrinsic tissuecalcification was disproportionate, it remains an integralcomponent of bioprosthetic degeneration and as suchneeds to be part of improvement strategies. Tissuecalcification is a multifactorial process, of which insuffi-cient immune masking is only one component. Glutar-aldehyde is another factor, as this dialdehyde is known tointrinsically elicit calcification. Initially, the glutaralde-hyde-based crosslinks themselves were suspected of beingpro-calcific [107]. However, when extraction methods forthe unbound derivates of glutaraldehyde were developed, itbecame clear that the intrinsic potential of glutaraldehydeto contribute to the calcification process rests in the freedialdehydes and their polymerization products rather thanthe crosslinks. By extracting unbound glutaraldehyde, twomajor improvements were achieved: tissue became non-toxic allowing even endothelial cells to grow on it [108] andcalcification could be distinctly mitigated, even if muchhigher concentrations of glutaraldehyde were used forfixation [109–111]. The flurry of approaches marketablytermed ‘‘anti calcification treatments’’ often representedvariations on glutaraldehyde detoxification. Substancessuch as diphosphonates, for instance, are believed to acttwofold: aldehyde stabilization occurs through binding viaSchiff’s bases and subsequent reduction by NaBH4, whilerestriction of crystal growth is thought to occur throughdirect diphosphonate binding to developing hydroxyapa-tite nucleation sites [109]. The precise mechanism of theanticalcification effect of a-amino-oleic-acid (AOAs)—the


anticalcification substance incorporated into Medtronic’sMosaic and Freestyle valves—is not fully understoodeither, but it is thought that it also binds with its twoamino groups to free aldehyde groups. Residual aldehydeis pivotal for the binding of AOA [112].

4.2.2. Influence of crosslinking on degeneration

Cross-linking appears to have a direct and an indirecteffect on tissue calcification: while the density of cross-linksindirectly acts through different degrees of immunemasking [88,89,91], there is an additional direct effectthrough the chemistry of the cross-linking process itself.One empirical approach was therefore to experiment withalternative crosslinking procedures such as Photofix[113,114] or alternative crosslinking agents such as Epoxycompounds [115], Carbodiimide [116], Diglycidyl [117],Reuterin [118], Genipin [119] and many others. Analternative approach to this empirical permutation ofcrosslinking agents emerged in the wake of tissue- andbio-engineered valves. The realization that engineeringprinciples can be applied to hitherto empirical processes[120] led to a more fundamental and mechanistic analysisof the ‘engineerable’ elements of cross-linking. Afteranchoring the cross-links to the carboxyl rather than theamino groups, for instance, it could be demonstrated thatunblocked amino groups in the tissue have a strong pro-calcific effect [121] and vice versa, the length andhydrophobicity of the blocking agent determine the degreeof suppression. Once such rational ‘engineering’approaches allowed the almost complete elimination oftissue calcification [122], other consequences of tissuefixation, such as stiffness, became amenable to a similarapproach [123].

4.2.3. Cellular antigen removal (decellularization processes)

Cell extraction has been tried since the 1950s [124]. Fromthe beginning, the primary goal was to extract cellularimmunogenicity and thereby eventually manage to avoidcrosslinking. The misconception of this approach was thatit exclusively targeted cell surface molecules (MHC-II inthe case of allografts and the Gal-a1,4-Gal xenoantigen inheterografts), implying that the extracellular matrix is non-immunogenic. Apart from the fact that extracellular matrixhas been shown to be immunogenic [96], the decellulariza-tion process itself holds a pro-inflammatory potential.Apart from the uncertainty with which solubilized antigensand remnant detergents are eluded from the tissue, thedecellularization process itself may even liberate pro-inflammatory substances. Arachidonic acid, for instance,a major component of the membranes of cellular orga-nelles, may potentially not be removed by aqueous buffersand thereby play a role in neutrophil chemotaxis throughleukotriene B4, a product of the arachidonic acid pathway.The latter may partly explain the massive polymorph-nuclear response to the clinically implanted ‘‘Syner-GraftTM’’ valves [125,126] that were used as rightventricular outflow tract replacement in Ross procedures.

However, regardless of whether the inflammatory responseis co-triggered by matrix immunogenicity and byproductsof the extraction process or only by matrix immunogeni-city: the latter will need to be addressed. The most obviousway of dealing with it is an additional cross-linking step.Interpreting contemporary research programs, however, itseems unlikely that the cross-linking of decellularized BPTwill be eagerly embraced. The reason for this reluctance liesin the fact that fixation would conflict with one of theunfulfilled hopes behind decellularization. This unfulfilledhope is the misconception that—once implanted—an‘ideal’, ‘natural’ scaffold such as a decellularized matrixwould be eagerly populated by host cells. Whereas someevidence appears to support the ingrowth potential ofacellular xenogenic matrices, at least in the sheep [127–129]and rat [130], the debate is skewed by the multitude ofdecellularization methodologies in use, the wide range ofextraction times employed and the questionable animalmodels used to validate the process. Our own experiences(unpublished) with exhaustively washed decellularizedporcine aortic valves have confirmed a failure to repopulatein rat, rabbit and primate models, with any ingrowthpotential confounded by a massive granulocytic response.This is supported by the findings of others confirming theability of decellularized tissue to attract mononuclear cellsand granulocytes [131,132]. The last blow to the concept oftissue ‘vitalization’ came with the clinical implantationof SynerGraftTM Valves, where no evidence of matrixrepopulation was seen even after a year of implantation.Stock et al. [133] even suggested that, in an inert acellularmatrix devoid of MMPs and their inhibitors (TIMPs),repopulation by cells from adjacent tissue is biologicallyunlikely [133].

4.3. Way forward for tissue valves

The mantra of bioprosthetic research must be therecognition of its system-inherent limitations such asmatrix immunogenicity and perpetual non-vitality. Inas-much as the removal of the xenogenic cells reduces theimmune burden, the continual immunogenicity of theextracellular matrix will always require masking. Similarly,believing that the extraction of cells will leave voids behindwhich are inviting spaces for ingrowing host cells means toignore the decades-old knowledge that even the best ofdecellularized heart valves do not get repopulated:Although homografts are non-crosslinked they remainlargely acellular even after years of implantation [134].However, if our goal is less ambitious and we just want tosignificantly extend the longevity of a non-vital bio-prosthesis, engineered crosslinking already offers a solu-tion, potentially augmented by combining the process witha preceeding immune burden reduction such as decellular-ization. A more recent discovery may even allow one toachieve immune masking without cross-linking: by fillingnon-crosslinked tissue with a hydrogel, antigens andproteins seem to have been spatially rendered inaccessible

ARTICLE IN PRESS

Fig. 7. A novel approach to abrogate bioprosthetic tissue calcification and degradation involves the filling of the tissue with hydrogel rather than cross-

linking the amino or carboxyl groups of the tissue as conventionally done in tissue fixation. The ramified blue lattice of the hydrogel filling the space

between two yellow collagen strands schematically shows how a ‘space filler’ has replaced cross-links. The tissue structure is well preserved after filling with

20% acrylamide (+ bis-acrylamide crosslinker; UV initiation) (top left; Goldner; original magnification: 40� ) and calcification (60 days rat

subcutaneous) is well below that seen after standard glutaraldehyde fixation (bottom right) [135].


by cells, enzymes or immune molecules. It was hypothe-sized that it was this inaccessibility that prevented thein vivo degradation, inflammatory infiltration and calcifica-tion of hydrogel-infiltrated tissue (Fig. 7) [135].

5. Polymeric flexible-leaflet valves

The promise of synthetic heart valves was to haveprostheses available that have the durability of mechanicalvalves and the haemocompatibility of tissue valves. Inreality, most of the polymeric valves were the opposite,combining the durability of tissue valves with thethrombogenicity of mechanical valves. Therefore, yearsafter the first flexible leaflet polymeric heart valves wereimplanted into patients [2,136], they have yet to reach aperformance level which makes them clinically acceptablebeyond the short-term use in artificial hearts (e.g.Abiomed, Berlin-Heart, Medos [137]).

5.1. Valve design

Similar to synthetic vascular grafts, material and designaspects stood in the foreground during the 1960s and 1970swhile biocompatibility was added as a third pillar from the1980s onwards. Initially, most researchers used the basictrileaflet (Fig. 8a) aortic design for their prototype valves[138–141]. Advances in fluid-dynamic measurements and

numerical modelling subsequently allowed for optimiza-tion of flow dynamics and leaflet stress. Variations rangedfrom hemi-cylindrical cusps [142] to leaflets in half-openposition [143,144]; from variable-curvatures [145] toelliptical and hyperbolic shapes [146] and from a conicalbase with spherical upper part [147] to high profile cusps(Fig. 8b) [137,148]. Notable exceptions to the tricuspiddesign include the very first bileaflet mitral valve implantedby Braunwald [2], the single-leaflet silicone-covered Dacroncusps implanted by Hufnagel in the 1970s [149] and anasymmetric bileaflet mitral valve [137,150] comprising akidney-shaped stent-ring containing a large anterior andsmall posterior cusp, thereby mimicking mitral flowconditions [148] (Fig. 8c). Overall, design and materialimprovements led to a gradually increased durability(up to 900 million cycles) and fatigue resistance nowequals that of bioprostheses [151]. There still seems to behigh variability in cycle life not only between differentvalve designs and different studies [137,146,148,150]but also between similar valves within one and the samestudy [151].

5.2. Leaflet materials

Since 1959 [137,138], the overwhelming majority ofpolymer valves were made of polyurethanes (PUs). Theseuser-friendly materials have continuously been favoured

ARTICLE IN PRESS

Fig. 8. Polymeric heart valves: (a) a frame machined from polyetheretherketone (PEEK) and coated with a thin layer of leaflet polyurethane. Leaflets of a

commercially available polyetherurethane suitable for animal implantation (Estane 58315, BF Goodrich, Westerlo-Oevel, Belgium) were dip-coated onto

the frame. This valve design has achieved durabilities in excess of 400 million cycles (10.5 years) during in vitro fatigue testing [162]; (b) and (c):

polycarbonate urethane (PCU) tri-leaflet and bi-leaflet valves intended for the aortic and mitral positions. These particular designs achieved in vitro

durabilities of up to 600 million (15.8 years) and 1 billion (26 years) cycles, respectively [137].


for blood contacting applications in spite of their initiallypoor long-term chemical stability [152]. However, new-generation materials such as Elasteon [153] seem to havelargely overcome the problem of bio-degradation. Simi-larly, improved fabrication technologies [145,154] resultedin even and constant leaflet thickness thereby improvingfatigue resistance and durability [151,154,155]. SegmentedPUs are generally flexible and durable at approx 100 mmleaflet thickness, although there was some concern that aleaflet thickness of less than 150 mm may not be sufficient[156]. Overall, the choice of thickness and materialmodulus is a trade-off between tensile strength andimparted bending stresses. The difficulty of this choice isreflected in the broad range of leaflet thicknesses(60–400 mm) reported [137,147,148,150,151,155].

The second most widely used polymer family wasSilicone rubbers [140]. Both the ball of the first successfulmechanical valve and the leaflets of one of the firstclinically implanted polymeric heart valves [141] weremade from this material. Disappointing results [136,157],however, led to silicones eventually being abandoned.

In general, most of the biomaterials of the 1960s and1970s were either tried for vascular grafts or synthetic heartvalves. Expanded Teflon was one of the materials whichwas used for both applications [139,158,159]. Mostrecently, a novel polyolefin, poly(styrene-b-isobutylene-b-styrene) (SIBS), with excellent chemical stability, hasbeen proposed as alternative to the traditional polymers[160,161]. The material was found to match PU in plateletdeposition tests, and was comparable with PU in terms oftensile and fatigue properties, albeit only after reinforce-ment with polypropylene fibres [160,161]. Altogether, thepromising result with both the latest generation ofbiostable urethanes [162] and innovative new materialssuch as fibre-reinforced SIBS [160] for the first timeindicate that after 50 years of failure a new door mayeventually have opened a small gap.

5.3. In vivo performance

Initial results of the 1960’s with synthetic flexible leafletvalves were catastrophic [136,163]. During the subsequentdecade improvements were only moderate [149] or short-lived. An example of the latter was the re-discovery ofePTFE for a tricuspid heart valve prosthesis [139]. Afterpromising initial results, its leaflets were soon found to beprone to stiffening, free-edge inversion and calcification[159] confirming a previous clinical trial that had led to ahigh mortality rate, associated with leaflet rupture andthickening almost 15 years earlier [158].Until the mid to late 1980s, calcification [142,164,165]

had almost inevitably led to leaflet immobilization, ruptureand perforation [166]. This seemingly insurmountablelimitation briefly resulted in an almost complete disconti-nuation of polymer valve programs. Subsequent materialimprovements during the 1990s [146,151,167] allowed there-commencement of some of these programs culminatingin a mitral implant study [162] where no valve relateddeaths occurred. Furthermore, bioprosthetic anticalcifica-tion treatments of the 1980s and early 1990s wereincorporated in an attempt to modify the leaflet surfacestowards more ‘biocompatibility’. Covalently bound bipho-sphonate [168,169], the extraction of low molecular weightcomponents [167] from methanol-extracted polyetherurethane (PEU) and the modification of a polyurethanevalve with polyethylene oxide and sulfonate groups(PU-PEO-SO3) all showed reduced polymer calcification.The latter also showed a decrease in thrombogenicity andcrack formation [170]. Similarly, heparin, taurine andaminosilane modifications equally resulted in extendeddurability [171]. Yet, even the latest generations of polymervalves continue to experience some degree of extrinsiccalcification [137,148,150] indicating that the goal of along-lasting synthetic heart valve may remain elusive aslong as no cellular components such as endothelial cells are


an integral part of the concept. The first successful attemptsto grow endothelial cells on polyurethanes and PU-siliconerubber co-polymers were made in the 1980s [172–174].A promising current approach to endothelialization in-volved the pre-seeding of ECs onto cholesterol-modifiedpolyurethane cusps. This modification resulted in increasedcollagen synthesis and cell retention in vitro andin vivo [175,176].

5.4. Way forward for polymer valves

Over the course of half a century, polymeric valvedesigns have been improved regarding flow characteristicsand stress reduction. During the same period of time,degradation-resistant materials were developed, processingmethods evolved that allow reproducible high-qualitymanufacturing, and thrombogenicity and calcification havebeen partly addressed through surface modifications. Yet,without the integration of living cells into these con-structs—either through blood borne fallout healing [177] oractive incorporation through seeding—it seems unlikelythat the full potential of polymeric valves will ever come tofruition.

6. Tissue-engineered valves

Tissue-engineering technologies providing living auto-logous heart valves with the capacity of regeneration andgrowth have shown promising experimental results andinitial human applications have been reported [178–181].Resorption of the porous leaflet scaffold initially led toovershooting fibrosis [182,183] but fine-tuning of theresorption process through the use of poly-glycolic-acid(PGA)/poly-4-hydroxybutyrate (P4HB) instead of poly-lactic acid (PLA) co-polymers led to living heart valveswith a tri-layered structure and many features of a nativeleaflet [181]. The ability to visualize the vitality and cellularactivity of heart valve cells in vivo is a backbone of follow-

Fig. 9. Novel multimodality molecular imaging techniques enable the monito

and disease. The long-axis view (a) shows the aortic root and arch, followed by

caused by the uptake of nanoparticles by activated cells. Colour-coded signal i

(modified from [217] with permission/mouse model/VCAM-1).

up requirements of future clinical trials with tissue-engineered valves (Fig. 9).While the majority of researchers pursue a concept

whereby a scaffold is seeded in vitro with autologous cells,an alternative approach aims at the in vivo recruitment ofcells [184] through the incorporation of biological signalsinto a degradable scaffold. The latter is supported by theobservation that circulating endothelial progenitor cells(ECP) led to the spontaneous endothelialization of theblood surface of clinically implanted ventricular assistdevices [185,186]. Similarly, ECPs have been shown tohome to stents coated with CD 34 antibodies (CD34 beingcell surface molecules of ECPs) [187].Although scientifically representing a quantum-leap

improvement to today’s heart valve prostheses, tissue-engineered heart valves still have a long way to go towardscompeting with contemporary heart valve prosthesesregarding safety, functionality, logistic feasibility andfinancial viability. A clearly unmet medical need andindication, however, is in paediatric applications due to theinherent growth-potential of tissue-engineered valves.Furthermore, recent studies have demonstrated the feasi-bility of using prenatal foetal cells—which can be obtainedduring pregnancy and used for the tissue-engineeringprocedure prior to birth—to provide living, autologousheart valves for early correction of congenital heart defects[178,188] (Fig. 10). Therefore, should tissue-engineeringtechnologies become an accepted clinical tool, it willmost likely happen through the ‘‘niche’’ of paediatricapplications.

7. Catheter-based valves

As attractive as the concept of a catheter-delivered heartvalve is, the emperor’s new clothes syndrome remains:while it potentially reduces the implantation trauma, itperpetuates the unsatisfying situation regarding availableprostheses. By using mainly conventional bioprosthetic

ring of cell activation and remodelling enzymes during valve development

a short axis view (b) which shows negative signal enhancement (darkening)

ntensities (c) show focused uptake of the nanoparticles in the commissures

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Fig. 10. Ideally, an autologous cardiovascular substitute would be used

for the repair of congenital cardiovascular malformations which has the

potential to grow and regenerate, thereby avoiding secondary damage to

the immature heart and re-operations with their associated mortality and

morbidity over a life-time. If used at or shortly after birth, the process of

engineering such an autologous construct would need to be initiated

prenatally. Conceptually, foetal stem cells could be harvested from extra-

embryonic foetal tissues such as chorionic villi, the umbilical cord or

amniotic fluid. After differentiation and proliferation in vitro, cells could

be seeded onto a biodegradable scaffold and conditioned in a bioreactor,

mimicking physiological conditions. The feasibility to fabricate foetal

tissues according to this concept has been recently demonstrated by the

in vitro fabrication of autologous living heart valve tissues based on

prenatal progenitor cells derived from umbilical cord [218], chorionic villi

[188] and amniotic fluid [219].


material, delivery-associated hype is not matched byprognostic expectations. This is particularly true for thelargely young patients in developing countries, whose non-calcified rheumatic valves would theoretically make themideal recipients for endovascularly placed prostheses.Given the rapid degeneration of all contemporary flexibleleaflet valves in young recipients, however, the very patientgroup that would dramatically benefit from transcatheterdelivery is the one that hardly qualifies for it.

7.1. Principle

The concept behind catheter-based valve delivery issimple: a valved stent is collapsed into a catheter, thecatheter tip is positioned at the site of the valve to bereplaced, and the stent is balloon-expanded to unfurl thenew valve. The execution of the concept, however, is verycomplex. In addition to the strenuous demands alreadyplaced on surgically implanted valves, and to the ingeniousstent and catheter designs required to deliver them,catheter-based valves must withstand being forced into a

tube at a fraction (approximately a third) of their expandeddiameter, and then be able to unfold into the precisegeometries required for function and durability. Althoughball-and-cage [189] and tilting disk concepts [190] wereproposed and evaluated, flexible-leaflet designs are prob-ably the most suitable candidates to fulfill the requiredcriteria, and it follows that transcatheter valve leaflets arepreferentially fabricated from BPT or flexible polymersused in more conventional designs. Although the first type(BPT) suffers from degenerative flexural failure, 50 years ofresearch have yet to produce a valve of the second type(polymeric) that approximates its success (as seen insections above). Hyperelastic metal leaflets comprise athird option, but relatively little is currently known aboutthe suitability of this emerging technology for valveapplications.

7.2. For few of the few?

If there was ever a development that was tailor-made forthe millions of patients in threshold countries who need avalve replacement, it was the catheter-based delivery ofheart valve prostheses. By offering patients with earlyregurgitation who have no access to open heart surgery butaccess to a modern C-arm visualization system a tool forthe placement of a functioning valve and thereby prevent-ing ventricular dilatation, morbidity patterns of entirepopulations could change. Yet, the available bio-materialsare so unsuitable for these patients that catheter-basedheart valves turned from a promise for millions into themost exclusive first world prostheses of all. Even worse,contemporary catheter-based concepts cater for two first-world fringe-indications: severely diseased patients who areinoperable due to co-morbidities [191] and pulmonaryplacements in patients with congenital heart disease inorder to obviate or delay the comorbidity associated withre-operation [192,193]. Even if indications may eventuallybe extended to the majority of first world patients, themarket is small. Given this small potential market, it isastonishing how all the major players rushed to the scene.Since the first transcatheter implants in animals occurred in1992, with Andersen et al. and Pavcnik et al. delivering astented porcine valve and a ball-and-cage valve (withinflatable ball), respectively [189,194], there are reported tobe more than 20 companies currently working on thistechnology [192]. A Pubmed search for ‘‘percutaneousheart valve’’ receives over 1500 hits, which include morethan 150 reviews!Since the first human catheter-based replacement of a

pulmonary valve in pediatric patients by Bonhoeffer et al.[195] and an aortic valve by Cribier et al. [196], eventsfollow hot on the heels of one another. Bonhoeffer hassince treated more than 120 patients with pulmonaryinsufficiencies using a porcine jugular vein valve in aplatinum stent [197] (now the Medtronic Melodys valve)(Fig. 11), while the Edwards Sapiens (equine pericardialtricuspid valve in a stainless steel stent) [198] (Fig. 11), and

ARTICLE IN PRESS

Fig. 11. Two examples of percutaneously delivered heart valves, the

Edwards Lifesciences ‘‘Sapien’’ aortic valve prosthesis (left) and the

Medtronic ‘‘Melody’’ transcatheter pulmonary valve prosthesis (right).

Although both models use unconventional animal sources (equine

pericardial for the Edwards valve and bovine jugular for the Medtronic

valve), the debut of two major commercial players highlights how

‘conventional’ the actual valves are—even if the delivery is catheter

based. Reprinted from [76] with permission of Universal Medical Press,

Inc.


Corvalves (bovine pericardial valve in a Nitinol (NiTi)self-expanding stent) [76,199,200] are arguably the mostcommonly used aortic prostheses. Other valves include the3F (Enables and Entratas) (bioprosthetic leaflets in self-expanding NiTi stents) and Sadra Medical Lotuss

(pericardium/NiTi) aortic valves, the Shelhighs (porcinepulmonic/NiTi) pulmonary valve, the Aortexs valve, andthe Palmaz-Baileys valve composed entirely of Nitinol[76]. A first assessment of the clinical experience with morethan 120 transvenous placements of pulmonary valvesyears after correction of congenital malformations, wasfavourable in comparison to surgical pulmonary valvereplacement in a historical cohort from the same institution[201]. The median hospital stay was 2 days (compared to 7in the surgical group). There was no mortality and the earlymorbidity was 5.8% (vs. 8.5% in the surgical group).Percutaneous aortic valve replacement in humans was firstperformed as a femoral transvenous, transseptal procedurewith antegrade access to the aortic annulus [198]. In theinitial attempt in 26 high-risk patients (turned down forconventional aortic valve surgery due to high risk), 22deployments were successful and 4 failed (2 patients couldnot tolerate the guidewire across the mitral valve and in 2the valve migrated). Complications were high with 27%early (o30days) and 41% late deaths. The survivors havereturned to a normal life with mean transvalvular gradientsbeing 11mmHg and paravalvular leak grade 0–1 in sevenand grade 2 in four patients. Subsequently, Webb et al.[202] from Vancouver reported a femoral transarterialprocedure with retrograde access to the aortic annulus.Their initial results in 18 patients showed successfuldeployment in 14 (the 2 failures were due to iliac arterialinjury in two and prosthetic valve embolization in anothertwo patients). There were 2 post-procedural deaths and onesurgical aortic valve replacement for failure—in the 14patients with successful deployment paravalvular leakagevaried from none to moderate. Because of the technicaldifficulty in deploying the device, Ye, Webb and colleagues

from Vancouver performed a hybrid procedure utilized atransapical access to deploy the valve antegradely underradiographic control [203]. There were no intraproceduralmortalities or morbidities but one patient died on day 12 ofpneumonia and 2 died later on days 51 and 85, respectivelyof non-cardiac causes; the remaining 4 have done well andcompleted 6-month follow-up—only one has more thantrivial or mild aortic incompetence and none have hadvalve-related complications. A larger series of 30 very highrisk patients (average age 8275.1 years) from Leipzig usinga similar method of transapically introduced valves wasmore favourable with successful deployment of the devicein 29 of the 30 patients, experiencing 3 deaths [204].Extracorporeal circulatory support was only used in theinitial 43% patients and thereafter the procedures weresuccessfully performed without circulatory support. Over-all, this initial experience confirms transcatheter placementof heart valves as a promising approach for fringe groups.

7.3. Design/procedural challenges

In addition to securing seating—in first world patientsmostly on top of heavy calcium deposits—and possibleinterference with other cardiac structures, access and valveplacement are important considerations in valve develop-ment [193]. In the pulmonary position, access is typicallygained via the femoral vein [195], although transventricularpulmonary placement [205] has been advised in senescentpatients when the RVOT is large and the valve and cathetersize precludes traditional percutaneous placement.In the aortic position, the initial antegrade transseptal

procedure via the femoral vein [206] was technicallydemanding encountering difficulties in accurately seatingand deploying the device via this tortuous route (via anatrial septostomy, through the mitral valve and having tonegotiate an acute curvature in the left ventricle). There-fore, the transseptal approach was subsequently aban-doned for retrograde arterial placement [202,207]. Thisprocedure was also later extended to mitral valves[208,209]. Although better tolerated by patients, it had itsown challenges—especially in older patients—in requiringa large caliber femoral artery capable of receiving a 24F(8mm) catheter (commonly used for currently preferred26mm valves). Moreover, the long catheter route alsomade accurate deployment of the device difficult. This ledYe et al. [203] from Vancouver to utilize a similar cathetermounted valve to implant an aortic valve via the apex ofthe left ventricle antegradely via a small left anteriorthoracotomy and without the use of cardiopulmonarybypass. Although not strictly complying with the conven-tional interpretation of ‘endovascular’ placement, transa-pical placement via endoscopy or small thoracotomy[203,210,211] allows the catheter-mounted valve conduitto be introduced via the left ventricular apex more directly.The shorter, more rigid and larger deployment systemallows more accurate and secure prosthetic valve deploy-ment in the aortic annulus. This is done via radiographic

ARTICLE IN PRESS

Fig. 12. Image of the ‘‘ValveXchange’’ transapically exchangeable

bioprosthetic heart valve. Rather than being a permanent valve, the

exchangeable valve is a two-piece device with leaflets that can be replaced

after having worn out. In the main image, the exchange process is shown

through the apex of the heart. A special trocar locks onto the ‘‘docking

station’’ (the sewing cuff and valve stent) to stabilize the heart and valve,

and a valve removal tool is inserted. The stent posts are grasped, the valve

lifted from the docking station, collapsed and pulled out through the

trocar. A new valve is immediately inserted and the procedure is done

completely off-pump. The inset image shows a retrograde approach, in

which a similar valve holder is passed from the outflow aspect to lock on

to the docking station. In this approach, the valve is passed over the shaft

of the valve holder during the exchange ([212], with permission).


and echocardiographic control. It also allows a larger sizevalve to be deployed.

7.4. What is in the pipeline?

An innovative way of accepting the shortcomings ofcontemporary tissue valves while reducing the stress,morbidity and eventually higher mortality of the re-operations is the concept of the transapical ‘renewal’ of adegenerated, surgically implanted valve. One system(ValveXchange, Inc.) consists of a two-piece valve andthe associated tools for replacing an exchangeable leafletcomponent without requiring cardiopulmonary bypass(Fig. 12) [212]. Another one is a valve by Sadra Medical,that also contains a re-positionable anchor, and allows forvalve replacement [213].

Other interesting developments include suture-less clo-sure after transapical access [214], a low profile valve thatcan be introduced with 11–16F catheter (used in the firstretrograde placement in humans) [207], and polymericvalves by Sochman (folding silicone disk in a Nitinol) [215]and Attmann (novel low profile PU valve in Nitinol stent24mm diameter that can be crimped down to 14F catheter)[216]. Paniagua et al. [207] have developed a low profile,metal-stented, tissue valve that can be introduced with a11–16F catheter).

8. Where to with replacement heart valves?

Contemporary heart valve prostheses are extremely fine-tuned to a first world situation. When using a mechanicalvalve we replace one disease by another with turning youngto middle-aged patients into haemophiliacs with theadditional risk of clotting up their valve or having astroke. Only the sophisticated medical system of thedeveloped world can cope with this challenge without asignificant level of morbidity and mortality. In the samefirst world population, tissue valves currently represent anequally fine-tuned compromise for the aged, but areotherwise almost as unsatisfying as in previous decades.With no disruptive technology on the commercial horizon,these valves will remain unacceptable for the youngpatients of threshold countries while getting more trouble-some for developed countries. The reason for the latter isthe increasing challenge for globally operating companiesto source internationally acceptable animal products in thewake of prions and viruses. Moreover, if developmentalsteps were particularly difficult to bring from bench tomarketing before, recently tightened European regulationswill not act as an encouragement for manufacturers, sincemost of the tissue valves were previously developed in theUnited States but clinically tested in Europe.

Tissue-engineered valves will undoubtedly eventuallyrepresent the future. However, judging by the snail-paceat which tissue engineering has developed since it becamevogue in the late 1970s, its clinical application in the field ofreplacement heart valves—even in the narrow paediatric

niche—may take at least another generation. In view of thesmall market, it may yet remain in the domain ofUniversities and start-up companies.Given the current low incentive for the major companies

to develop a catheter-based heart valve that does not needanticoagulation but nevertheless has a long durability inyoung patients, one needs to ask where the niches may bewithin existing technologies that may benefit the hugeuntapped number of potential recipients in thresholdcountries. Judging by the technological abilities of coun-tries such as India, China or Brazil, first world develop-ments could be (and already are!) easily and affordablyreproduced. One of the main obstacles, however, wouldbe licensing. Although hardly any of the developmentsthat have shown great promise—from engineered cross-linking to polymer filling, from heparin bonding to


decellularization—may eventually be implemented in thefirst world, research was often financed and patented by itsmajor commercial players. Therefore, generous licensingprograms could not only make valve replacements anacceptable treatment for many times the numbers ofpatients who currently benefit from it, but also providethe overdue feedback on performance patterns in youngrecipients. With patients from threshold countries beinglargely young and thus ideal for endovascularly placedvalves due to the lack of calcification, endovascularlyimplanted tissue valves pioneering improvements alreadylying in the drawers of companies, may create a win–winsituation which overcomes the 50 year long stalemate in thedevelopment of commercially available heart valve pros-theses.

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Prosthetic heart valves: Catering for the few

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