THE IDEAL VALVE

Many acceptable substitutes exist today for the replacement of diseased human heart valves. The ideal valvular prosthesis as described by Harken ¹ remains the Holy Grail of cardiac surgery. According to Harken, the ideal valve would be durable with a longevity approaching that of a native valve. Thrombogenicity would be nonexistent, and there would be no need for supplemental anticoagulation. In addition, the ideal replacement valve would have no inherent gradient, thus allowing unimpeded outflow. It also would be easily implanted and readily available. Finally, growth commensurate with that of the recipient would be possible.

HISTORY

The first human heart valve operation was a digital valvotomy of a stenotic aortic valve performed by Tuffier in 1914. Cutler, Souttar, Brock, Swan, and Harken refined valvotomies and commissurotomies in the ensuing decades. The need for a replacement valve arose out of the quest for an effective treatment of valvular insufficiency. In 1950, Hufnagel developed a ball valve, designed to be placed in the descending thoracic aorta. He was the first to implant a prosthetic valve in a human when he implanted his valve in the descending thoracic aorta of a patient with severe aortic insufficiency.² In 1956, Gordon Murray implanted an aortic homograft in the descending thoracic aorta of a patient with severe aortic regurgitation.³ With the introduction of cardiopulmonary bypass, open valve replacement became a possibility. In 1960, Braunwald and Harken successfully performed mitral and aortic valve replacements with valves made of polyurethane.⁴ In 1961, Albert Starr, a surgeon, and Lowell Edwards, an engineer, developed a caged ball valve with which they achieved long-term survival.⁵ Caged ball valves worked well in many patients, but their high profile made implantation difficult in patients with small ventricular chambers and small aortic roots. High inherent gradients, along with a less favorable thromboembolic profile, made caged ball valves less desirable as new prostheses were introduced.

Several investigators designed disc valves, which functioned by having a disc pivot into an open or closed position as dictated by flow across the valve. The first of these tilting disc valves was the Wada hingeless valve introduced by the Japanese surgeon Jura Wada in 1966.⁶ The Lillehei-Kaster valve, a hingeless valve with a freely rotating pivoting disc retained by struts, was introduced in 1967 by C. Walton Lillehei and Robert L. Kaster. Viking Björk, working with Shiley Laboratories, developed a similar version of a hingeless pivoting disc valve. Although the hemodynamic profile was improved compared with the caged ball valves, these early pivoting disc valves were subject to occasional thrombosis. The 60-degree convexoconcave Björk-Shiley disc valve was prone to catastrophic structural failure secondary to fracture of its welded struts, allowing escape of the occluder disc.⁷ Seeking to improve on the problems of durability and thrombogenicity with these initial pivoting disc valves, Karl-Victor Hall, along with Woien and Kaster and the Medtronic Corporation, introduced the Medtronic-Hall valve in 1977.

The St. Jude Medical bileaflet mechanical prosthesis was also introduced in 1977.⁸ It has gone through several refinements and is currently the most commonly used mechanical valve prosthesis.

Independently, in 1969, Carpentier and Hancock developed the first porcine xenografts.⁹ Ionescu developed the first glutaraldehyde-preserved bovine pericardial valve in 1971.¹⁰

Limited long-term durability secondary to leaflet calcification and subsequent perforation plagued early bioprostheses. Significant progress has been made with the latest generation bioprostheses by state-of-the-art fixation and preservation methods to achieve extended durability. Stentless xenografts were introduced in 1986 as a way to counteract the inherent gradient found with stented bioprostheses.¹¹

Ross reported the use of aortic homografts in 1962 and subsequently developed the Ross procedure, which uses a pulmonary autograft for aortic valve replacement.¹²

HEMODYNAMIC ASSESSMENT OF CARDIAC VALVES

The normal valve area is 3.0 to 4.0 cm ² for the aortic valve and 4.0 to 5.0 cm² for the mitral valve. Stented bioprostheses and mechanical prostheses have inherent gradients because the actual valve is supported by a stent and housing material. The effective orifice area (EOA) refers to the true cross-sectional area of the prosthetic valve orifice through which blood must flow. The “gold standard” for measuring valve area is the Gorlin formula, which uses data derived from catheterization and applies hydraulic formulas for fixed orifice systems to heart valves.¹³ The formula is as follows:

where A is the cross-sectional area of the valve (cm²), Q is the flow across the valve (mL/sec), and Δp is the pressure gradient across the valve (mm Hg); 44.3 is a derived coefficient.

The continuity equation uses echo Doppler–acquired data to calculate EOA. Both a standard and a simplified equation exist. The standard equation is as follows:

where SV is the stroke volume (derived from flow measurement) and VTI is the velocity-time integral across the prosthesis (also measured by Doppler).

The simplified equation is EOA = Qp/Vp, where Qp is peak flow (mL/sec) and Vp (mL/sec × cm²) is the peak flow velocity across the valve.

The continuity equation has been shown to be accurate for bioprostheses.¹⁴ When it is applied to bileaflet valves, it has been shown to underestimate valve area because of localized high-velocity jets.¹⁵

Although valve manufacturers have often used the geometric cross-sectional area of prostheses to describe their size, the parameter that is clinically important is the functional orifice of the valve, the EOA. In general, EOA is proportional to valve size for a given type of prosthesis, and the EOA is generally always less than the cross-sectional area of a patient’s valve anulus. Accordingly, some degree of obstruction is always imparted by prostheses, which can be particularly important in the case of the aortic valve. Many surgeons and cardiologists believe that particularly for aortic valve replacement, the EOA of an implanted aortic prosthesis must be matched to the size of the patient to provide sufficient gradient relief. This is particularly the case during exercise. Although the concept of patient-prosthesis mismatch makes intuitive sense, its impact on patient morbidity and mortality in the short or long term remains unclear and is the subject of numerous studies. Some authors have suggested that patient-prosthesis mismatch can be avoided when the ratio of EOA to patient body surface area exceeds 0.85 cm²/m². The manufacturer’s labeled valve size does not delineate the true internal or external diameter of a given valve and often does not represent the true EOA. Indeed, labeling of the valve sizes is arbitrary, and as such it cannot be used to compare values among different manufacturers and is not a true indicator of EOA.¹⁶

MORBIDITY AND MORTALITY GUIDELINES FOR VALVE OPERATIONS

Clinical studies are crucial for determining outcomes after cardiac valve operations, and precise definitions of outcomes are critical in comparing prostheses. To address this need, the councils of the Society of Thoracic Surgeons and the American Association for Thoracic Surgery formulated the Ad Hoc Liaison Committee for Standardizing Definitions of Prosthetic Heart Valve Morbidity. The initial report of this committee was issued in 1988,¹⁷ with an update following in 1996.¹⁸ The report strictly defines types of morbidity and mortality that can occur after valvular operations. An understanding of these definitions is crucial for interpreting studies dealing with valvular prostheses.

The guidelines distinguish two types of mortality: hospital mortality and 30-day mortality. Hospital mortality refers to death occurring at any time before discharge during a patient’s initial hospital stay. Thirty-day mortality, also referred to as operative mortality, is death that occurs at any time or place within 30 days of operation. There are several precise definitions of valve-related morbidity. Structural valve deterioration refers to “any change in function (a decrease of one New York Heart Association functional class or more) of an operated valve resulting from an intrinsic abnormality of the valve that causes stenosis or regurgitation.”¹⁸ It includes “changes intrinsic to the valve, such as wear, fracture, poppet escape, calcification, leaflet tear, stent creep, and suture line disruption of components … of an operated valve.”¹⁸ Thrombotic or infectious causes of valve dysfunction are not included.

Nonstructural dysfunction includes nonthrombotic and noninfectious causes of valvular stenosis or regurgitation that are not intrinsic to the valve itself. “Examples … include entrapment by pannus, tissue, or suture; paravalvular leak; inappropriate sizing or positioning; residual leak or obstruction from valve implantation or repair; and clinically important hemolytic anemia.”¹⁸ Morbid events are often reported as the composite linearized rate, or the number of events divided by the number of patient-years of follow-up (events/patient-years). The composite linearized rate for nonstructural dysfunction of the commonly available mechanical valves is 0.2 to 0.8 (events/patient-years) for the aortic position and 0.3 to 1.4 (events/patient-years) for the mitral position.¹⁹

Valve thrombosis is defined as thrombus in or about the valve that is not associated with infection and that interferes with valve function or obstructs blood flow through it. The composite linearized rate of thrombosis of mechanical valves is 0 to 0.2 (events/patient-years) for the aortic position and 0.4 to 0.8 (events/patient-years) for the mitral position.¹⁹

Embolism refers to any embolic event not associated with endocarditis that occurs after the immediate perioperative period and after the emergence from anesthesia. Embolic events are further delineated into neurologic events and peripheral embolic events. The composite linearized rate of thromboembolism ranges from 1.4 to 2.5 (events/patient-years) for the aortic position and 1.8 to 3.6 (events/patient-years) for the mitral position.¹⁹

A bleeding event refers to any clinically significant bleed requiring hospitalization or transfusion or causing death. A patient does not have to be taking an anticoagulant to sustain a bleeding event. Composite linearized rates vary from 0.8 to 2.5 (events/patient-years) for the aortic position and 1.2 to 2.2 (events/patient-years) for the mitral position.¹⁹

Endocarditis involving an operated valve is designated operated valvular endocarditis. “Morbidity associated with active infection, such as valve-thrombosis, thrombotic embolus, bleeding event, or paravalvular leak, is included under this category and is not included in other categories of morbidity.”¹⁸ Composite linearized rates for prosthetic valve endocarditis range from 0.4 to 0.7 (events/patient-years) for both the aortic and mitral positions.¹⁹

Consequences of morbid events are also defined by the guidelines. A reoperation is any operation on “a previously operated valve.”¹⁸ The composite linearized rate for reoperation ranges from 0.3 to 1.8 (events/patient-years) for the aortic position and 0.6 to 1.6 (events/patient-years) for the mitral position.¹⁹

PROSTHETIC HEART VALVES

The following section describes prosthetic heart valves commonly in use in the United States. Valves are grouped by their mode of construction. Discussion is by valve type and manufacturer. Order in listing in no way delineates preference by the authors.

Mechanical Prostheses

Caged Ball Prostheses

Starr-Edwards Silastic Ball Valve Prosthesis

The Starr-Edwards ball valve prosthesis (Edwards Lifesciences Corp., Irvine, CA) was introduced in 1966. The model currently available is constructed of a cage made from a single piece of titanium that encloses a Silastic ball. Although the valve is very durable and has a notable history, indications for its use today are limited because its hemodynamic and thromboembolic profiles do not match more modern designs (Fig. 76-1).¹⁹

Figure 76–1 Starr-Edwards Silastic ball valve prosthesis.

(Courtesy of Edwards Lifesciences Corp.)

Tilting Disc Prostheses

Medtronic-Hall Mechanical Heart Valve

The Medtronic-Hall tilting disc valve (Medtronic, Inc., Minneapolis, MN) was introduced in 1977 (Fig. 76-2). The valve housing is constructed from one piece of titanium alloy with no introduced welds or bends. The round central disc is made from tungsten-impregnated graphite with a pyrolytic carbon coating and has a central hole that allows the disc to be retained by a curved central guide strut that is part of the housing. It was designed to be an improvement on previous tilting disc valves in terms of durability, hemodynamic performance, and reduced thrombogenicity. Valve washing was improved by a relatively larger minor orifice and a disc that lifts out of the housing and rotates with opening.²⁰ Loss of structural integrity has not been reported. The valve can be rotated after implantation and has a low inherent transvalvular gradient. It has a moderately high profile in the open position. Occluder impingement is possible because its position at the equator of the valve housing makes it susceptible to obstruction from retained valve elements, sutures cut too long, or pannus.