Artificial Heart: Technology, History, and Future Prospects

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Background

A healthy heart operates as a dual‑pumped organ, with two atria and two ventricles. The right atrium receives deoxygenated blood from the body and contracts into the right ventricle, which then propels it to the lungs. The left atrium receives oxygenated blood from the lungs and pushes it into the left ventricle, the chamber that distributes blood to the rest of the body. Each cardiac cycle begins with atrial contraction, followed by ventricular contraction, ensuring efficient blood circulation.

Congestive heart failure (CHF) is the progressive loss of the heart’s pumping capacity and remains a leading cause of mortality worldwide. The American Heart Association reports that approximately five million Americans live with CHF, and over 400,000 new cases are diagnosed annually. Roughly half of these patients die within five years of diagnosis. In 1998, the U.S. health system spent roughly $95 billion treating heart disease.

While pharmacologic therapy and surgical interventions can alleviate symptoms, a heart transplant remains the definitive cure for end‑stage heart failure. In 1998, about 7,700 patients were listed for transplant, yet only 30 % received an organ. This shortage spurred the development of artificial hearts and ventricular assist devices (VADs) as bridge‑to‑transplant or destination therapy options.

An ideal artificial heart must deliver continuous, reliable blood circulation and oxygenation for extended periods—typically at least 100,000 beats per 24 hours—without lubrication or routine maintenance, while adapting its pumping rate to the patient’s activity level and minimizing infection or thrombosis risks.

The two principal categories of circulatory support devices are the heart‑lung machine and the mechanical heart. The heart‑lung machine, comprising an oxygenator and a pump, is employed during open‑heart procedures and is limited to a few hours of use due to blood hemolysis over time. Mechanical hearts, on the other hand, reduce the workload of a failing native heart and can function as temporary or permanent support. LVADs, for example, assist the left ventricle by delivering a portion of the normal cardiac output. Currently, around 4,000 LVADs are implanted worldwide, and the U.S. market for these devices is estimated at $12 billion annually.

History

Since the late 19th century, scientists have pursued mechanical solutions to replace or supplement cardiac function. It took nearly a century before the first human heart‑lung machine, developed by John H. Gibbon Jr., was successfully used in 1953. Four years later, a plastic artificial heart was implanted in a canine model.

The National Heart Institute (now the National Heart, Lung, and Blood Institute) formalized an artificial heart program in 1964, culminating in the first total artificial heart implantation in a human in 1969.

In the 1970s, the focus shifted toward left ventricular assist devices (LVADs) and the use of biocompatible materials. The first successful LVAD implantation occurred in 1970, and subsequent decades saw the introduction of polyurethane and plastic pumps with extended lifespans. Regulatory tightening by the FDA in the 1980s increased development costs, leading to consolidation in the field.

Dr. Robert Jarvik, a pioneer in the field, introduced the Jarvik‑7 artificial heart in 1983. This aluminum‑plastic device, driven by a large external compressor, was implanted in Barney Clark in 1985. Clark survived 112 days, and subsequent recipients lived up to 20 months, demonstrating the feasibility of implantable support.

The desire for fully implantable, battery‑powered devices led to NIH funding in 1988 and the development of an electric LVAD by 1991. In 1999, Charlie Chappis became the first patient discharged home with an implanted LVAD, marking a milestone toward long‑term therapy.

Raw Materials

Artificial hearts and LVADs are constructed from a combination of metals, polymers, ceramics, and biological components. The primary metallic alloy used is titanium‑aluminum‑vanadium, chosen for its biocompatibility and mechanical strength. Titanium parts are precision‑machined and receive a microsphere coating that promotes endothelialization, creating a natural blood‑compatible lining.

The pump diaphragm is fabricated from a textured polyurethane that encourages blood‑cell adhesion, reducing hemolysis. Polyester grafts attach the device to the aorta, while pig heart valves serve as biologic replacements for native valves. Motor components are typically made from titanium or high‑strength ceramics.

Design

Designing an LVAD requires meticulous attention to fluid dynamics to ensure adequate flow without thrombosis, selection of truly biocompatible materials, and optimization of motor efficiency to limit heat generation. Device dimensions are kept minimal to reduce the risk of immune rejection; a typical LVAD weighs about 1.2 kg (2.4 lb) and occupies 660 ml (1.4 pints).

Robert Jarvik

Born May 11, 1946, in Midland, Michigan, Jarvik pursued architecture and mechanical drawing at Syracuse University before switching to pre‑medicine after his father's heart disease. He earned a B.A. in zoology in 1968, later obtaining an M.A. in occupational biomechanics from New York University in 1971. After receiving his M.D. in 1976, Jarvik joined the University of Utah’s Institute for Biomedical Engineering, where he collaborated with Willem Kolff on artificial heart research.

The Manufacturing Process

• Components are custom‑manufactured by specialized third‑party suppliers, including precision machine shops and printed‑circuit board producers. Porcine valves are sewn into polyester grafts at facilities dedicated to heart‑valve fabrication.

  Following component acquisition, the LVAD assembly undergoes rigorous testing to verify compliance with all specifications. Once cleared, the device is sterilized and packaged for distribution.

Forming the Polyurethane Parts

• Some manufacturers fabricate polyurethane diaphragms in‑house using a proprietary liquid solution applied to a ceramic mandrel in successive layers. Each layer is cured and dried until the target thickness is achieved. The finished part is then removed from the mandrel and inspected. Alternatively, third‑party producers may employ injection or vacuum molding combined with RF welding.

Assembly

• Assembly is conducted in a sterile cleanroom to prevent contamination. Up to 50 precision components are bonded with high‑temperature‑curable adhesives. Parallel sub‑assemblies—including motor housing, percutaneous tube, and diaphragm pusher plates—are completed, inspected, and finally integrated into the complete system. Grafts are attached separately during the final assembly step.

Testing

• Post‑assembly, each device undergoes mechanical testing that simulates physiological pressures. All electronic circuits are evaluated with diagnostic equipment to confirm proper operation.

Sterilization & Packaging

• After successful testing, the device is transferred to an external sterilization service. Each unit is sealed in a sterile plastic tray and then placed in a custom protective case for shipment, safeguarding it from contamination and damage.

Quality Control

Most components undergo pre‑inspection before shipment to the manufacturer, but dimensional checks with tolerances in the millionths of an inch are performed on critical parts. To meet FDA regulations, every component—including adhesives—is traceable via lot and serial numbers, enabling full traceability.

Byproducts / Waste

Scrap titanium is reclaimed, remelted, and recast, minimizing metal waste. Because most parts pass pre‑inspection, only a small quantity of defective components is discarded. Post‑implantation devices are returned to the manufacturer for analysis, providing data that drive iterative design improvements.

The Future

Over the next decade, a wave of innovations is poised to transform mechanical circulatory support. Researchers at Pennsylvania State University are developing an electromechanical heart that receives power transcutaneously via radio‑frequency energy, enabling patients to carry a battery during the day and recharge overnight. The prototype is slated for human trials by 2001.

Other groups are exploring continuous‑flow pumps, which eliminate the need for a mechanical pump stroke and reduce device size and energy consumption. In Australia, Micromedical Industries Limited is advancing a continuous‑flow rotary pump expected for human implantation in 2001. Ohio State University is working on a self‑regulating plastic pump the size of a hockey puck, designed for short‑term support until native cardiac function recovers.

Thermo Cardiosystems, Inc. is developing a bearing‑less LVAD featuring a magnetically suspended rotor, promising virtually unlimited durability. Both continuous‑flow centrifugal and rotary variants are being engineered for transcutaneous energy transfer, making them fully implantable.

With donor heart shortages, permanent artificial heart replacements are gaining traction. LVADs are being refined by inventors such as Robert Jarvik and surgeons like Michael DeBakey. Total artificial hearts are a joint effort between the Texas Heart Institute and Abiomed, Inc. In Japan, research teams are creating silicone‑ball‑valve total hearts powered by a ceramic‑bearing centrifugal pump.

Alternative therapies, such as shape‑modifying clamps that can improve native heart efficiency by up to 30 %, are also under investigation, offering minimally invasive options for patients with structural heart disease.