Plasma separator

30 April 14, 1986 C&EN

Artificial blood vessel

Blood oxygenator

Artificial kidney

Total artificial heart

I Artificial heart valve

SPECIAL REPORT

Biomaterials in

Artificial Organs Helen E. Kambic, Shun Murabayashi, Yukihiko Nose, Cleveland Clinic Foundation

The ability to replace damaged organs and blood ves­sels totally or in part has improved the quality of life and increased the lifetimes of many persons. The foun­dation for this success is improved biomaterials. Bio­materials are substances or combinations of substances that can be used in a system that treats, augments, or replaces any tissue, organ, or function of the body.

A widening variety of biomaterials is being used today outside the traditional areas of plastic and recon­structive surgery, dentistry, and bone or muscle repair. Artificial organs, for example, are expected to play a greater role in preventive medicine, especially in the early prevention of end-stage organ failure.

For several decades, bodily parts have been replaced or repaired by direct substitution with natural tissue or selected synthetic materials. At the beginning of the 20th century, devices and materials were focused on life-threatening situations, such as those involving the need for blood transfusions. Largely by modifying the composition and properties of plastics, chemists, bio­chemists, engineers, and others have developed to­day's biomedical devices.

Cardiovascular materials The development of devices for use in the cardiovas­

cular system has been hampered by the shortage of materials of sufficient strength and durability for long-term implantation. Other problems with materials for such use include blood-clot formation and mineraliza­tion on the material surface, mechanical failure, hemo­

lysis (destruction of red blood cells), material degrada­tion, and the activation of blood coagulation factors.

Cardiac devices fall into three general groups. One of these is transient devices, such as temporary left ventricular assist devices (LVAD) or the intra-aortic balloon pump, which assist the pumping of blood by a damaged heart for a few days to two weeks. Another group is interim or temporary LVADs that may be used for up to a month, to support a patient awaiting a heart transplant. A third group is permanent devices such as the artificial heart, which support the total circulation.

For blood pump applications, one important consid­eration is the fatigue life of the ventricular sacs or diaphragm used to pump the blood. In a pump beating 100 times a minute, the diaphragm flexes more than 50 million times a year. The diaphragm or sac also re­quires elastomers that undergo cyclic deformation and flexing while in contact with blood. Poly ether-type polyurethane elastomers are the most versatile materi­als that meet these needs, partly because they can be solution cast, injection molded, or extruded into a vari­ety of shapes and configurations. They also are more durable and less hydrolyzable than polyester-type ure-thanes.

Polyurethane's surface structure, characterized by different organic functional groups (hydrocarbon, amine, ether, amide, ester, carbamate), affects its bio­logical performance. Polyurethanes are highly inho-mogeneous materials, thus they may contain a variety of low-molecular-weight compounds that tend to mi­grate to the surface. These compounds may alter the material's surface properties by leaching out from the polyurethane into the biological media, thereby changing its surface structure.

April 14, 1986 C&EN 31

Special Report

These changes are important in biomedical applica­tions in which the surface of the polymer comes in direct contact with blood or tissue and, therefore, influ­ences blood compatibility. "Blood compatibility" re­fers to a variety of phenomena that occur at the blood/ material interface. Thrombus (clot) formation at a sur­face, for example, is influenced by the character of the surface. Blood clots can be thrown off from the surface of a device or develop within the blood vessels by coagulation factors activated by interactions on the sur­face.

Therefore, the surface properties and structure of a material, hydrodynamic factors (blood flow, shear stress), changes in blood coagulability after exposure to a material, and the processes for making a device are all related to thrombus formation. The surface chemistry and surface structure of many biomaterials are well defined. However, once used in constructing a device, their blood reactivity may be altered significantly by surface-structure changes induced by the extraction of compounds into the blood or tissue.

Blood flow within a device directly affects thrombus formation. At low rates of blood flow, prolonged blood contact with the surface activates many coagulation factors. At high flow rates, blood experiences a high shear rate, leading to hemolysis and activation of coag­ulation factors. The configuration of a device also de­termines its potential for clot formation. Blood stagna­tion areas and undesirable changes in blood-flow pat­terns can occur if a device is poorly designed. Thrombi can form as a result of the high shear rates caused by the formation of vortices where flow patterns change or in blood stagnation areas.

Changes in the level of such blood constituents as white cells, platelets, complement, and fibrin also in­fluence clot formation, as does such blood-flow dy­

namics as pulse rate and pressure. In addition, contami­nants, leachables, temperature, humidity, method of sterilization, and design affect the potential for clots forming on a material.

The segmented polyether-type polyurethanes are noncrosslinked polyurethanes whose physical proper­ties depend on the presence of soft and hard segments in the polymer chain. Hard segments are formed by urethane and urea linkages, whereas soft segments are created by polyether residues formed by reactions of diisocyanate with polyether macroglycols.

Segmented polyurethanes can be either aromatic or aliphatic. The aromatic polymers are based on methy-lenebis(4-phenylisocyanate) and either poly(tetra-methylene glycol) or poly(propylene glycol). The ali­phatic analog of these phenylisocyanate compounds is methylenebis(4-cyclohexylisocyanate).

The ratio of hard to soft segments determines the hardness, ultimate stress, ultimate strain (deforma­tion), and elongation of the resulting polymer.

In addition to the polyurethanes, Hexsyn, a sulfur-vulcanized polyolefin rubber made by Goodyear, is being evaluated for use in blood pumps. Hexsyn is polymerized from a mixture of 95% 1-hexene, 3% 4-methyl-l,4-hexadiene, and 2% 5-methyl-l,4-hexa-diene.

Evaluating the polymers Ideally, materials would be evaluated by actually

testing them for long periods of time in blood pumps either in the laboratory or in animals. Because of both the excessive time and cost involved, however, this is not feasible for testing new materials. In 1980, the National Heart, Lung & Blood Institute of the National Institutes of Health funded three projects to evaluate the fatigue properties of elastomers and their applica­bility for blood pumps.

Robert W. Penn and coworkers at the National Bu­reau of Standards evaluated Kontron's Cardiothane, a block copolymer of 90% polyether urethane and 10% polydimethylsiloxane. John L. Kardos and associates at Washington University tested Ethicon's Biomer, a seg­mented aromatic polyether-based polyurethane. And Carl R. McMillan, then at Monsanto Research Corp., evaluated Goody ear's Hexsyn.

All three of these groups have been involved in developing short-term fatigue tests that will predict long-term in-vitro performance and indicate the fa­tigue life of materials for potential use in circulatory assist devices. These studies will be of value in investi­gating all materials used in artificial organs for which fatigue life, decomposition products, and elastomer degradation in blood, water, and gas are of critical importance.

Accelerated fatigue tests and static mechanical/phys­iological tests have confirmed that the durability of materials for cardiovascular devices is influenced by the testing methods selected. On the basis of such pre­dictive tests, Biomer, Hexsyn, and Cardiothane should be considered for most cardiovascular applications.

The advantages of Biomer include its excellent flex life, extreme toughness, good fatigue resistance, ability

A technician assembles a left ventricular heart-assist device at Thoratec Laboratories in Berkeley, Calif.

32 April 14, 1986 C&EN

A scientist at Battelle-Columbus Laboratories observes a sheep from which blood continuously is removed and then returned to the animal. The external blood comes in contact with a number of experimental polymers used in artificial organs to determine how the surfaces of these materials interact chemically with various components of the blood

to be autoclaved, and high resistance to creep (dimen­sional changes after prolonged exposure to stress). One disadvantage is its relatively high permeability to flu­ids such as water.

The advantages of Hexsyn include its low permeabil­ity to body fluids and low cost of raw materials. Its disadvantages include a lower resistance to creep (com­pared with Biomer) and low tear strength.

Cardiothane performs well in fatigue tests under a variety of conditions. One of its disadvantages, as with some other polyurethanes, is a high permeability to water, a low resistance to creep, high cost, and the necessity that it be film-cast in multiple layers under rigid controls.

The Pellethane 2363 series of aromatic polyether-based polyurethanes, made by Upjohn, are tough and resistant to fatigue, with a high flex life. However, they have a considerable tendency to creep under load and cannot withstand autoclave temperatures without dis­tortion.

Not one of these materials has yet fulfilled all of the requirements for blood compatibility and mechanical durability. But careful consideration of their properties and extensive testing have led us to conclude that Biomer, Hexsyn, and Cardiothane have the best poten­tial for use in chronic implantable devices, especially those with flexing components.

For chronically implanted blood pumps, toughness and flex life of the material are important bulk proper­ties. Surface properties include the blood compatibility of the materials in terms of platelet attachment and the potential for clot formation.

Although the bulk and surface properties of materi­als are different, they depend on two major variables: molecular structure and molecular weight. These vari­ables can be manipulated to optimize surface proper­ties. But the final product may be less than optimum with respect to its bulk physical properties.

To overcome this problem, surface treatments, coat­ings, and polymer blends or laminates are used to form a surface that is independent of the underlying base or structural polymer. Nearly two decades ago, efforts were begun to identify or synthesize artificial surfaces that interact safely with blood. Two approaches have been employed. One relies on the use of nonthrombo-genic smooth surfaces; the other uses polymers with rough or textured surfaces to promote the deposition of a biologically compatible layer called the pseudo-neointima.

The use of smooth surfaces has the advantages of ease of fabrication, minimal thrombosis, and minimal deposition of the blood protein fibrin. Improper poly­mer synthesis and quality control procedures can limit the flex life of these polymers, as can the introduction of microscopic defects produced during fabrication. These factors can lead to polymer degradation and to blood cell and protein deposition. Treatment of the patient with an anticoagulant, such as sodium warfarin (Coumadin), and with antiplatelet drugs, such as aspi­rin and dipyridamole (Persantine), is usually required after surgery.

Treatments to create rough or textured surfaces de­velop in time a biologically derived and highly flexible blood-contacting layer. The major disadvantages of this approach are the necessity to control the pseudo-neointimal layer thickness and to provide the patient with anticoagulant drugs. Moreover, the development on the polymer surface of calcium deposits from the blood can lead to the loss of the material's flexibility.

During the early 1960s, many devices were fabricat­ed with what investigators thought were smooth sur­faces. Smoothness was desired to eliminate stagnant areas in the pumps and to reduce blood clotting. These devices, however, were poorly designed. There were seams within them, and the flow patterns were poor. In a short time, despite anticoagulant therapy, clot forma-

April 14, 1986 C&EN 33

Soecial Report

Polymerization reactions lead to formation of aromatic polyurethanes Pellethane, Biomer

/ / \N OCN-</ \ } - C H 2 - ^ V> NCO + H—0-hCH2CH2CH2CH2 —θ4-Η Jn

Methylenebis(4-phenylisocyanate)

Η Ο

Poly(tetramethylene glycol)

Ο Η 0 = C = N - ( ^ \ V - C H 2 ^ y - N - C - o | - C H 2 C H 2 C H 2 C H 2 — O - f C - N ^ f / - C H 2 - ^ N ^ - N ^ C ^ O

Isocyanate-terminated prepolymer

Η Ο - CH2CH2CH2CH2-OH

j „ . Butanediol

Ο Η Ο Η . > H O τ - . O H Η Ο

4·(ΟΗ2)4-0-0-Ν-<^ ^ C H 2 - ^ y-N-C-0-UcH2CH2CH2CH2— O-Lc-N-^f γ~^^ζ\ V N - C - O H C H 2 ) 4 4 -

Pellethane (Upjohn product)

Isocyanate-terminated prepolymer

Η Ο Η

H2N-CH2CH2 -NH2

Η Ο 4 - ( C H 2 ) 2 - N - C - N - ^ ν ~ Ο Η 2 Η ^ V - N - C - o 4 - C H , C H , C H 9 C H , — O - J - C - N - ^ ^ - C H ,

Ethylenediamine

Ο Η Ί l"

\ / * \ / Biomer (Ethicon product)

Η Ο Η // V N - C - N - Î C H , ) ,

Polymerization reactions yield the aliphatic polyurethane Tecoflex ο Dicylohexylmethanediisocyanate

OCN- -CH, NCO + H~0-J~CH2CH2CH2CH2—O-

Poly(tetramethylene glycol)

Η + HO-CH 2CH 2CH 2CH 2 -OH

1,4-Butanediol

I » Ϊ Γ \ ^ Λ V fi Γ 4(CH2)4-0-C~NHf VCH 2 - / \ - N - C - o 4 —

• -,Ο Η Ί H ι

CH2CH2CH2CH2—0-4-C - Ν CHC

Tecoflex (Thermedics product)

H I

N -

0 II

-c-- 0 - "(CH2)4J

tion, with deposition of the clots, was a common prob­lem. Eventually, many of the surfaces cracked and de­formed, leading to complete failure of the device. More recently, the development of better blood-pump con­trol systems, improved pump design, and the absence of flow obstructions or seams have provided more ac­ceptable results.

One type of material for smooth-surface cardiac pros­thesis is so-called "biolized" gelatin. Researchers in the Cleveland Clinic Foundation's department of artificial organs have developed a smooth gelatin protein (bio­lized) surface for the blood-contacting areas of their blood pumps.

Biolization is a process for making biocompatible

and blood-compatible materials by either chemical or thermal treatment of natural tissue or protein. This is done by extracting proteins or other biological compo­nents from natural tissue and blending them with wa­ter-soluble synthetic polymers or coating them on the textured surfaces of polymers. This concept offers the versatility of selecting specific polymers, such as the polyolefin rubbers, polyether-based polyurethanes, or silicones, depending on their physical properties, for a given application. Hence, scientists are not restricted to using only natural tissue or a narrow range of avail­able medical-grade polymers.

The smooth glutaraldehyde crosslinked gelatin coat­ings on blood pumps show no evidence of pseudo-

34 April 14, 1986 C&EN

neointimal formation because this surface, after con­tact with blood, is covered by a thin layer of protein-aceous material and is essentially free of cellular and fibrin deposits. With such smooth coatings, pseudo-neointimal formation is undesirable because these sur­faces provide no firm anchorage for cellular deposits. Should these deposits become detached, they could lead to serious complications, such as a stroke.

The blood-contacting surface of the blood pump is a seamless gelatin layer coated both on a textured Hex-syn rubber diaphragm and on a pump housing covered with a textured polyurethane, Avcomat 610. This sur­face is blood compatible, hydrophilic, durable, and cell nonadhesive. Blood pumps treated in this way have been removed free of clots from calves implanted with a total artificial heart or a left ventricular assist device even more than 10 months after implantation.

One textured surface used to promote pseudoneoin-tima formation is the flocked surface produced by at­taching polyester fibrils to a polyurethane substrate, such as Tecoflex, an aliphatic polyether-based polyure­thane developed by Michael Szycher of Thermedics in Waltham, Mass. When the material is exposed to blood, plasma proteins adsorbed on the polyester fibrils at­tract platelets that adhere and accumulate in the inter­stices of the flocked matrix. Fibrin deposits between the platelets attract and retain both red and white blood cells. After the fibrils are covered, specialized cells called fibroblasts colonize on this lining and be­gin to synthesize collagen. In time, the collagen surface replaces the fibrin to produce a stable pseudoneointi-mal lining. The resulting integrated textured surface is used in the clinical left ventricular assist devices made by Thermedics.

Textured surfaces also can be formed on metals by a sphere-sintering process, which uses fused titanium spheres that have been sifted to yield powders ranging from 45 to 225 μτη in diameter. These spheres, applied to the nonmoving parts of blood pumps, are coated with pseudoneointimal layers that are securely an­chored to the powdered metal surfaces.

Expanding on this texturization concept, several lab­oratories have developed cell-seeding techniques. Cul­tured human endothelial cells are used to accelerate the development of a thin, organized interface on a polyester fiber or on an expanded polytetraf luoroethy-lene surface.

Endothelial cells, it should be pointed out, line the interior of all natural vessels. The ideal surface is one with an endothelial cell lining (intima) that most close­ly approximates the lining of the natural vessel. An endothelial-like surface is especially important in situ­ations involving low blood flow, because it can reduce clotting on the surface. This lining also can regenerate and repair itself.

Some scientists consider the formation of pseudo­neointimal linings composed of native collagen and endothelial cells highly desirable because it represents an end stage in the intravascular healing process. Cell-seeding techniques have been applied successfully to both blood pumps and vascular grafts.

In 1980, NIH's National Heart, Lung & Blood Insti-

At Cleveland Clinic Foundation, a textured coating of polyurethane is applied to a blood-pump housing

tute established goals and criteria for developing heart devices and supportive techniques in an effort to im­prove the treatment of heart disease. These research directives focused on the development of temporary and permanent (two- to five-year) LVADs and the total artificial heart. Circulatory support using various forms of mechanical ventricular assist is now an accept­ed mode of therapy.

LVADs mechanically support the failing heart and

Photomicrograph shows the poly ester-fibril-textured surface used as the blood-contacting surface for the diaphragm of a Thermedics left ventricular assist device

April 14, 1986 C&EN 35

Special Report

At Thoratec Labs, a polyurethane is poured on a mandrel to make a diaphragm for a left ventricular assist device

circulation to decrease cardiac work, increase the blood supply to vital organs, and increase the oxygen supply to the myocardium (muscular layer of the heart wall). When these devices are used, the natural heart remains intact within the chest. Either implanted or placed ex­ternally, they are attached to the natural heart by means of a tubular connector. Blood from the pump flows through another connector attached to the aorta.

William Pierce and coworkers at Pennsylvania State University developed one version of the diaphragm ventricular assist pump, called the Pierce/Donachy Ventricular Assist Device, now commercially available through Thoratec Laboratories in Berkeley, Calif. The unit contains an air-driven, seamless pumping cham­ber made of smooth Biomer in a rigid polysulfone cas­ing. Bjôrk-Shiley tilting-disk valves are used at the inlet and outlet positions. The pump receives blood from the tip of the left ventricle or from the left atrium by means of a segmented polyurethane tube and ejects this blood at the rate of 6.5 L per minute into the descending aorta through another tube. Before it was used clinically in 1984, the pump was evaluated in 12 calves for an average of 70 days. In clinical trials at Penn State, this pump has been used in 32 patients.

Since 1973, Thermedics has been working on im­plantable, electrically powered left heart devices, which will soon be available commercially. Titanium alloy is used for the pump housing because of its excel­lent strength-to-weight ratio and corrosion resistance. Tecoflex polyurethane is used for the diaphragm, which has a textured polyester flocked surface at the blood interface.

Nosé and coworkers at the Cleveland Clinic Founda­tion have developed an LVAD for both temporary and permanent use in humans. The pump uses a pusher plate made of titanium and a biolized, gelatin-coated Hexsyn diaphragm. The device can pump 9.6 L per minute at 120 beats per minute.

This pump is being adapted for use with two energy conversion systems, one electrical, the other thermal.

The electrohydraulic system is being developed by Nimbus Inc., of Rancho Cordova, Calif. With this de­sign, the patient would wear the main power source (an external battery) incorporated into a vest and a belt pack. A secondary power source would be provided by an implanted battery, which, acting as a backup system, could operate for about 30 minutes if power fails from the external source. Thus, as proposed, the patient could be completely free of external components for a limited time.

With this assist device, electrical power and signals must be transmitted across the intact skin by the induc­tive coupling of an external primary coil and an im­planted secondary coil. This integrated blood pump and electrohydraulic system are now being evaluated in calves at the Cleveland Clinic Foundation.

Nimbus also has developed a fully implantable ther­mal energy converter for left ventricular assist devices. The basic concept of the thermal power system in­volves the conversion of heat to pneumatic power and the subsequent use of this power to actuate and control the blood pump. The Cleveland Clinic Foundation's biolized LVAD will use this type of actuation system.

Peer M. Portner and coworkers at Novacor Medical Corp. in Berkeley, Calif., have developed an integrated LVAD that uses a seamless polyurethane sac symmetri­cally actuated by dual pusher plates. This electrome­chanical actuation system uses a solenoid energy con-

A brushless torque motor (bottom left) operates this left ventricular assist system (top) developed by Thermedics, Electrical commutator has been removed (bottom right). System includes an implantable battery package (upper right) to permit operation if main power source fails

36 April 14, 1986 C&EN

vertor to energize two opposite pusher plates within the pump. The Novacor group also has developed a belt worn around the waist that accommodates a bat­tery pack and electrical controls for energy transmis­sion across the skin for operating its implantable left ventricular assist system.

A pneumatically activated partial mechanical heart has been developed by Adrian Kantrowitz and col­leagues at Sinai Hospital in Detroit. The device is cast of segmented polyester polyurethane and is flocked with polyester fibrils to establish a pseudointima. The device's 30-cc pumping chamber is implanted in the body to replace a portion of the wall of the descending thoracic aorta.

Inflated in time with and out of phase with the natu­ral heart, it augments central aortic pressure during diastole (expansion of the heart). In addition, it reduces the input impedance at the root of the aorta at the beginning of systole (contraction of the heart). These actions, together with their effects on coronary artery flow and left ventricular pressures, wall tension, and oxygen requirements, reduce the external work of the heart while improving cardiac metabolism and increas­ing blood perfusion to the vital organs.

Clinical studies carried out since the early 1970s have shown that the Kantrowitz partial mechanical heart can alleviate intractable congestive heart failure. Such a partial mechanical heart appears to provide minimal alteration to hemostasis (arrest of bleeding) during car­diac assistance and, when turned off, behaves like a passive vascular graft.

The total artificial heart The Jarvik-5 and Jarvik-7 pneumatic total artificial

hearts were developed by Robert K. Jarvik under the direction of Willem Kolff at the University of Utah. They are the culmination of Kolf f's pioneering efforts in artificial heart development since 1957. The Jarvik-7 heart is smaller than the Jarvik-5 and can fit within the human chest.

The two ventricles of the Jarvik-7 heart are made of polyurethane, supported on aluminum bases. A four-layer Biomer diaphragm has graphite dispersed be­tween the layers to act as a lubricant. The resulting flexibility permits normal blood inflow pressures to depress the diaphragm and fill the ventricle during diastole. The four-layered pumping diaphragm can withstand the left pump's high pumping pressure (more than 200 mm Hg) and is designed so that stress points that can cause defects in the Biomer are elimat-ed. Polycarbonate rings support the tilting-disk Bjôrk-Shiley heart valves.

While the patient is on heart-lung bypass, the natu­ral ventricles are removed. Polyurethane cuffs are then sutured to the two remaining atria and to the two large arteries that connect with the heart. The device's right and left ventricles are snapped into the cuffs and con­nected to the external pneumatic drive system by ap­propriate drive lines.

The first human recipient of a Jarvik-7 artificial heart was a 61-year-old dentist, Barney B. Clark. The opera­tion performed on Dec. 2,1982, by William C. DeVries

Adrian Kantrowitz at Sinai Hospital in Detroit devised this partial mechanical heart It is made of segmented polyurethane and is coated with polyester fibrils for use in patients with congestive heart failure

at the University of Utah Medical Center, succeeded in saving Clark from immediate death. After receiving the mechanical heart, he lived for 112 days before dy­ing of circulatory collapse.

William J. Schroeder, a 52-year-old diabetic, had a Jarvik-7 heart implanted by DeVries at the Humana Hospital Audubon in Louisville Nov. 25, 1984. While recovering from the surgery, he suffered three strokes on Dec. 13, resulting in permanent speech damage. Despite these problems, Schroeder has lived with the artificial heart for more than a year and remains the world's longest-living recipient.

Since then, other patients have received artificial hearts of various types at the University of Arizona Medical Center in Phoenix, Penn State's Milton S. Her-shey Medical Center in Hershey, Presbyterian Univer­sity Hospital in Pittsburgh, Karolinska Institute in Stockholm, and elsewhere. The first woman recipient of an artificial heart, Mary Lund, received a small Jar­vik-7 on Dec. 18,1985, at Abbott Northwestern Hospi­tal in Minneapolis. In some cases, the devices were used only until the patient could receive a transplanted human heart.

Problems with artificial hearts Infection is the leading cause of disease and death in

long-term surviving animals (such as calves) with total artificial hearts. Infection rates of 25 to 50% have been reported in animals that have had artificial heart im­plants. The organisms most often detected in blood culture have been Pseudomonas aeruginosa, Enterobacter cloacae, Escherichia coli, and Enterococcus, all of which are intestinal species.

The predominance of gram-negative bacteria (74%) suggests that the contamination is derived from the animal itself. The various pressure lines, pressure transducers, and monitoring lines are the most likely avenues of bacterial invasion from the outside. Al­though many antibiotics, such as penicillins, are avail­able for combating bacteria, the eradication of infec-

April 14, 1986 C&EN 37

Special Report

tion remains a formidable task once microorganisms gain a foothold into the interstices of the various pros­thetic materials and devices.

In recent years, calcification or mineralization of blood pump surfaces also has emerged as a significant problem. Calcification of artificial heart valves has been observed frequently in humans and occurs very rapidly with valves implanted in children. About 5% of these valves malfunction for this reason within six years of implantation. Valves made from animal tissue apparently are subject to early calcification (within 30 days) in young calves.

Such mineral deposits cause premature mechanical failure of cardiac devices. The binding of serum calci­um and phosphates and the deposition of yet-un­known components may form nucleation sites for the growth of the deposits. Dennis L. Coleman and his group at the University of Utah have proposed that, initially, lipids or phospholipids in the blood are ad­sorbed onto polymers via adsorption of lipoprotein complexes. Phosphate groups on the phospholipids could attract and bind calcium, thus starting the calcifi­cation process.

Jane B. Lian of Children's Hospital in Boston has evaluated the calcific deposits formed on assist devices that are made from Biomer and contain diaphragms coated with flocked polyester fibril. Calcification on polyester-flocked surfaces starts at the polymer-pseu-doneointimal interface, rather than at the blood-con­tacting surface. Studies using transmission electron mi­croscopy show that much cellular debris collects on the polyester fibrils and within the fibrin pseudoneointi-mal matrix of the flexing portions of the diaphragm. Needles of hydroxyapatite have been found in associa­tion with cell-wall membranes and other cell debris. X-ray diffraction has revealed the presence of a poorly

At Symbion Inc., the diaphragm of a Jarvik-7 heart is cast in four layers of flexible Biomer polyurethane. Each layer is lubricated with graphite

crystalline hydroxyapatite form of calcium phosphate. Gel electrophoresis indicates that two calcium-binding amino acids, O-phosphoserine and O-carboxyglutamic acid (GLA), are present in significant concentrations that correlate with the calcium concentration of the sample. The GLA-containing protein osteocalcin, which is synthesized in bone and circulates in the blood, also accumulates in the calcified deposits.

Hiroaki Harasaki of the Cleveland Clinic Founda­tion has found in studies with calves that calcium phos­phate is deposited in injured, degenerated, or dead tissue. Dead, degenerated blood cells in this tissue be­come the nuclei for the precipitation and crystalliza­tion of calcium salt. This type of deposition occurs easily on surfaces with imperfections or contaminants, such as dust or air bubbles.

The cause of calcification is not yet known. But re­gardless of the materials used or the species and ages of the animals involved, the common phenomenon at the blood-material interface appears to be calcification. This phenomenon occurs especially in the moving part of the blood pump, where perforation of the dia­phragm can cause pump failure.

Heart valves Since the early 1960s, scientists have developed

nearly 50 different heart valves. Today, the ones most commonly used include mechanical prostheses and tis­sue valves (bioprostheses).

Early devices to replace the individual valve leaflets and eventually carry out the entire valve function were made with polyamides, poly(ethylene terephthalate), polytetrafluoroethylene, silicone rubber, polymethyl methacrylate, or stainless steel. Valves made from pericardium, vein, or atrium tissue provided good short-term results, but eventually malfunctioned and caused blood clotting. Coatings made of such sub­stances as graphite, heparin, and benzalkonium chlo­ride (Winthrop's Zephiran chloride) were used to re­duce blood clotting, but clinically they also proved unsatisfactory.

Nearly 75,000 of these prosthetic heart valves are implanted annually throughout the world, about 30,000 in the U.S. alone. Although the desirability of replacing defective heart valves has been clearly estab­lished, much uncertainty exists about the optimal tim­ing for surgery, the proper prosthesis to use, and the ultimate prognosis for the individual patient.

Problems associated with valve prosthesis have not been completely eliminated, even though this research has been in progress for almost 25 years. Among the many possible complications with heart valves are thromboembolism, red blood cell destruction, valve failure due to material degradation or fatigue, tearing of suture lines, tissue overgrowth from the natural blood vessel, and back flow caused by improper valve closing.

Caged-ball, caged-disk, and tilting-disk heart valves are the types most widely used today. The original Smeloff-Cutter and Starr-Edwards caged-ball valves contained a silicone rubber ball. Wear problems and materials degeneration in these heart valves resulted

38 April 14, 1986 C&EN

Artificial hearts have noticeably different designs Jarvik-7 device Pennsylvania State University device

Bjôrk-Shiley valves

Four-layer Biomer diaphragm

Biomer diaphragm

Plastic valves

Cleveland Clinic Foundation device

Tissue valves

Aluminum base

Drive lines to pneumatic drive system

Gelatin-coated polyolefin diaphragm

Pusher plate-.

Polysulfone casing

Drive lines

Titanium housing

Actuator assembly

Drive lines

in ball-diameter changes, ball cracking, and embolism (the obstruction of a blood vessel by, for example, a clot or air bubble). Degradation of the silicone rubber balls caused by ball swelling, cracking, and lipid absorption became more and more serious as more valves were used for longer periods.

Modifications of the basic design led to the use of Stellite 21 (a cobalt-chromium-tungsten alloy) for the ball and the use of alloys of cobalt, chromium, molyb­denum, or nickel to replace the stainless steel cage. Other changes included a cloth covering for the valve body or strut and an increased orifice-to-ball ratio. In­corporation of these changes resulted in the current 6120 Starr-Edwards valves.

Bjôrk-Shiley and Lillehei-Kaster tilting-disk valves were introduced in 1969. The Bjôrk-Shiley valve con­sists of a Stellite 21 cage, with a disk that pivots around an eccentric axis in the plane of the valve orifice. The original disk, made of acetal resin (Du Pont's Delrin), has been replaced by a disk of graphite-covered pyro-lytic carbon for increased blood compatibility. This type of valve is used in the Jarvik-7 artificial heart.

The Lillehei-Kaster valve has a pivoting-disk design, with the disk inclined at an 18° angle in the closed position to facilitate rapid opening and closing. Pyro-lite, a carbon-silicon alloy, is used for the disk, with a

titanium metal frame. The improved disk is composed entirely of pyrolytic carbon formed over a graphite substrate.

Another adaptation of an all-pyrolytic-carbon valve is the bileaflet St. Jude valve, used clinically since 1982. This valve consists of two leaflets that rotate within the orifice housing. This design provides a superior high-flow orifice, with the leaflets opening to an angle of 85° to the plane of the orifice ring.

Bioprosthetic valves are grafts of heart valves from deceased humans, pigs, calves, and other animals or constructed from nonvalvular human and animal tis­sue. Tissue-fabricated valves duplicate the flow of nat­ural valves, and most can be used without anticoagula­tion treatment. Valves made from pig aortic valves perform satisfactorily in most adult patients for up to 10 years.

Alain Carpentier of the Laboratoire d'Étude des Pros­theses Cardiaques at Paris' Hospital Broussais, a major advocate of natural-tissue valves, implanted the first one in 1965. Carpentier used glutaraldehyde to treat porcine valve collagen tissue and thus prevent both immunological reactions and the long-term denatur­ation of collagen. Such denaturation can cause damage to natural tissue.

Components of the valve, such as the cells, soluble

These artificial heart valves are made in a variety of configurations and use different materials of construction

April 14, 1986 C&EN 39

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Special Report

This x-ray shows a silicone caged-ball valve that was placed in the mitral position in the heart

proteins, mucopolysaccharides, and structural glyco­proteins that are antigenic are denatured by oxidation, using sodium metaperiodate. The oxidation produces free aldehyde groups that react with the free amino groups of adjacent molecules of collagen to form inter-molecular crosslinkages. Denaturation of the collagen is prevented by these spontaneous linkages and by the formation of glutaraldehyde crosslinkages. Glutaralde-hyde is a dialdehyde that can crosslink two molecules of collagen. The resulting five-carbon link produces the irreversible intermolecular crosslinking of two rel­atively distant amino groups.

The Hancock porcine bioprosthesis, made by Han­cock Laboratories in Anaheim, Calif., by a stabilized glutaraldehyde process, has been used clinically since 1970. These valves, after preservation with glutaralde­hyde, are sutured to a poly(ethylene terephthalate)-covered flexible polypropylene stent (a molded sup­port onto which the valve leaflets are sutured into place).

The Ionescu-Shiley pericardial valve made by Shiley Laboratories of Irvine, Calif., is sutured onto a poly­e thylene terephthalate)-covered titanium stent.

Nonvalvular tissues that have gained clinical accep­tance for heart valves are the dura mater (a membrane that envelops the brain), obtained from human cadav­ers, and bovine pericardium (a membrane that encloses the heart). Dura mater leaflets are mounted on a poly­e thylene terephthalate) velour-covered stent.

The incidence of clotting with natural-tissue valves is greatly reduced, compared with that with synthetic mechanical valves, but not altogether eliminated. Problems such as infection, calcification of valve leaf­lets, and immunological reactions have led to valve failure. These problems underscore the need for suit­able tissue procurement, processing, sterilization, and preservation. Nevertheless, the overwhelming evi­dence is that natural tissues, when properly treated, are superior to any synthetic materials because of their low tendency to produce blood clots.

Vascular grafts A frequently used surgical method for treating pa­

tients with atherosclerosis involves either the replace­ment or bypass of clogged blood vessels. To date, how­ever, no synthetic material has satisfied all the func­tional requirements for replacing or bypassing blood vessels.

Polyester fiber, such as Du Pont's Dacron, remains the preferred textile material for arterial replacements. According to Lester Sauvage, director of the cardiovas­cular reconstruction division of the Bob Hope Interna-

A technician at Cleveland Clinic Foundation inspects a trileaflet heart valve made of dura mater tissue (a membrane surrounding the brain). Valve is seen in closeup above

40 April 14, 1986 C&EN

tional Heart Research Institute in Seattle, "The best arterial prosthesis is one that has the potential to be completely incorporated by the body. More than any other material. . . Dacron has stood the test of time. It is light, durable, has excellent surgical handling charac­teristics, and can be produced in a wide variety of configurations to satisfy almost any surgical require­ment."

The first attempts at blood-vessel substitution in hu­mans were made using arteries taken from cadavers. Once the blood vessel was sterilized and sutured in place, it performed well. In time, however, calcifica­tion and a loss of elasticity resulted in clot deposition.

Today, the human saphenous vein, a nonessential vein taken from the patient's leg, is used for most coronary artery bypass surgery and remains the stan­dard by which other arterial grafts are compared. Coro­nary bypass surgery involves removing the saphenous vein from the leg and using it to bypass a partial or total blockage of one or more of the coronary arteries. The bypass is made by suturing one end into the aorta and the other into the coronary artery beyond the point of occlusion. Between 150,000 and 200,000 of these proce­dures are done annually in the U.S.

The saphenous vein has become the method of choice in dealing with many arterial occlusions. It con­tains normal endothelium—a single layer of endothe­lial cells, which is the most nonthrombogenic surface known today. It can be obtained in most patients in segments long enough so that not only the anterior surface of the heart can be reached but also the posteri­or surface. This posterior surface is where most of the bypass procedures have to be constructed either to the branches of the circumflex coronary artery or to the branches of the right coronary artery.

The Surgikos Reinforced Artergraft, marketed by Johnson & Johnson, is made of selected bovine carotid artery treated with the enzyme ficin and crosslinked with dialdehyde starch. Its outer surface is reinforced with polyester open-mesh tubing that minimizes the possibility of an aneurysm (an abnormal dilation of the blood vessel wall).

Meadox Medicals of Oakland, N.J., has developed a glutaraldehyde-stabilized human umbilical cord vein called Biograft. The advantage of starting with a blood vessel of biological origin is that it is not likely to cause clotting. The human umbilical cord is a long, branch-free blood conduit having characteristic endothelial cells covering a membrane that provides the mechani­cal strength and elasticity needed for blood vessels. Initial clinical results from implanting this graft are encouraging.

In the past five years, expanded polytetrafluoroethyl-ene (PTFE) grafts have become increasingly popular as a synthetic nontextile graft for reconstructive proce­dures, such as above- or below-the-knee bypasses for limb salvage. In the late 1960s, W. L. Gore & Associates in Flagstaff, Ariz., developed a process for mechanical­ly stretching PTFE to produce a semiporous material that maintains its inert biological properties. Porosity makes the graft wall permeable to water, cells, and tissue fluids. The more porous a graft material, the

A worker at Bard Cardiosurgery in Billerica, Mass., inspects polyester fiber tubing used to replace large blood vessels. Previously, the tubing was tested for, among other things, porosity and mechanical properties

more elaborate the preclotting technique required by the surgeon prior to graft implantation. Preclotting renders the walls of the graft impervious to blood by the accumulation of a combined fibrin and thrombus layer.

Tissue ingrowth on PTFE and the eventual forma­tion of a pseudoneointima depend on the pore size. The tubular PTFE consists of solid nodes from which extend longitudinally oriented fibrils.

PTFE grafts do not have to be preclotted to prevent blood leakage. Upon exposure to blood, the internodal spacing permits cell penetration and protein adher­ence, followed by platelet aggregation, eventual clot formation, and the development of a stable protein layer and pseudoneointima.

Since their introduction in 1975, PTFE grafts have been used in more than 500,000 clinical procedures, mainly for arterial reconstructions in lower limbs, limb salvage, and for blood access in dialysis procedures.

Although the saphenous vein is the graft of choice for arterial reconstructions, in nearly a third of all pa­tients it is either unavailable because of prior surgery or is inadequate because of its small diameter, short usable length, previous disease (such as phlebitis), or other problems. The availability of PTFE grafts for low­er-limb bypasses has provided a vascular substitute for high-risk patients whose only alternative would be major amputation.

Blood oxygenators Many materials have been used for constructing

blood oxygenators, often referred to as heart-lung ma­chines. The initial materials, such as glass and stainless

April 14, 1986 C&EN 41

Special Report

steel, gradually have been replaced by plastics. The introduction of polycarbonate, polypropylene, polyvi­nyl chloride, and polyethylene have made it relatively easy to mass-produce oxygenators rather inexpensive­ly. With growing demand for oxygenators by the late 1950s, the concept of an oxygenator was shifted from a precision-made, expensive, large device to a simple, inexpensive, small disposable unit.

The basic construction of an oxygenator involves any one of several types of units employing a bubble-type, membrane-film-type, or hollow-fiber-type de­sign. Most oxygenators have been of the bubble type, involving the direct contact of blood with oxygen and carbon dioxide. Prolonged use of this system, however, results in protein denaturation and destruction of red blood cells, which restricts its use to a maximum of six hours.

The most important advance in oxygenator develop­ment was the introduction in the mid-1950s of the membrane-type oxygenator. Membrane units are a step closer to the normal physiological condition in which gas contact occurs indirectly via a gas-permeable mem­brane. Blood trauma is greatly reduced compared with that in direct-gas-contact units, although some interac­tion does occur between membrane and blood. Mem­brane materials have included PTFE, PVC, and cello­phane, but by 1955 silicone rubber and cellulose ace­

tate were used predominantly. Silicone rubber, however, is preferred because of its high gas perme­ability and blood compatibility.

General Electric's MEM 213, a silicone rubber/poly­carbonate copolymer, is highly permeable to both oxy­gen and carbon dioxide. Microporous membranes, such as North Star Research's ethylcellulose perfluoro-butyrate and Owens-Illinois' poly(alkyl sulfones), have been introduced in the past decade.

Coil, plate, and hollow-fiber membrane oxygenators are all safer and more effective than the bubble devices. Because they avoid direct contact of blood with gas, they reduce the risk of blood trauma, protein denatur­ation, and air embolism.

A new generation of capillary-membrane oxygen­ators using bundles of microporous polypropylene hollow fibers has been developed by Terumo Corp. of Japan. The device contains 60,000 straight hollow fi­bers that are arranged longitudinally.

The use of hollow fibers for gas exchange is not new; many prototypes employed polydimethylsiloxane/poly­carbonate copolymers or silicone rubber as the gas bar­rier. The limitations in gas-transfer rates, requiring a large surface area, impeded the development of hol­low-fiber oxygenators. The subsequent availability of microporous membranes, however, solved the limita­tions in gas exchange. The hollow-fiber design pro-

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High flow rates used in modern preparative LC can cause excessive pressure drops. Typically, at 300 mL/min a standard Rheodyne analytical valve can cause a pressure drop of 150 psi.

Tech Note #7 tells you how to use a flow rate vs. pressure drop graph and select a substitute Rheodyne valve that reduces pressure drop to a negligible 17 psi. Send for Tech Note #7. Please address Rheodyne, Inc., PO. Box 996, Cotati, California 94928 USA. Phone (707) 664-9050.

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duces less turbulence and resistance on the blood side, and the smooth surface helps decrease blood trauma. Terumo's device, marketed initially in Japan in the early 1980s, is the first hollow-fiber oxygenator to be offered commercially.

Use of carbon A unique carbon material, originally developed for

encapsulating nuclear reactor fuel, has properties sur­prisingly appropriate for biomedical applications. Pro­duced by CarboMedics of Austin, Tex., it has a less crystalline (turbostratic) carbon structure, rather than the more ordered crystalline arrangements of diamond (the cubic polymorph) or graphite (the hexagonal poly­morph).

The turbostratic structure is present in all carbons used medically. Turbostratic carbons have both cova-lent and van der Waals bonding. The material exhibits unusually large elasticity before fracturing. Its high strength and high fracture strain permit the design of multicomponent devices that, under cyclic loading, re* sist breakage. Carbons with the turbostratic structure include the LTI (low-temperature isotropic) pyrolytic carbon, ULTI (ultralow-temperature isotropic carbon), and glassy carbon.

LTI carbon is formed by pyrolyzing hydrocarbons at temperatures between 1000 and 2400 °C in a fluidized

bed. The carbon, deposited on substrates to a thickness of 1 mm, is chemically inert and biocompatible. It also resists wear and fatigue and apparently is unaffected by the cellular components of blood. This combination of properties makes it especially useful for construct­ing mechanical heart valves. According to Axel D. Haubold, vice president of CarboMedics, more than 530,000 mechanical heart valves have been made of this material since its introduction in 1969.

More recently, these carbons have been used in re­constructing small joints in the hand and foot because of the material's similar elasticity to that of bone, its excellent wear properties, and its outstanding biologi­cal compatibility. Healthy bone can grow up to and into the carbon implant without forming the thick con­nective tissue capsule normally produced.

Ultralow-temperature isotropic carbon and glassy carbons are formed when carbon vapor is deposited under vacuum as a film on a wide variety of polymer or metal substrates. This procedure results in highly smooth or granular surfaces.

One interesting use of carbon is in repairing tendons and ligaments with composites of carbon fibers and absorbable polymers. John R. Parsons and coworkers at New Jersey Medical School in Newark have developed a partially absorbable composite carbon material for repairing severely damaged tendons and ligaments.

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CIRCLE 12 ON READER SERVICE CARD April 14, 1986 C&EN 45

Special Report

The material consists of poly(lactic acid), a biodegrad­able polyester of lactic acid, reinforced with fine, uni-axially aligned carbon fibers. In the body, it undergoes de-esterification to lactic acid, which is then metabo­lized. The composite provides mechanical strength to the repair site as new tissue develops around the car­bon fiber "scaffold" while the polymer is degraded and absorbed by the body.

Blood purification The use of biomaterials to purify blood is an area of

growing research interest. Hemodialysis and plasma­pheresis, two of the several treatment methods for blood purification, have one feature in common: The blood must be circulated outside the body.

Hemodialysis has been used during the past two decades to treat nearly 200,000 kidney disease patients worldwide. Without an artificial kidney, patients with acute kidney failure die within seven days.

Natural kidneys are made up of millions of micro­scopic filters, called nephrons, that cleanse the blood of wastes, such as urea, uric acid, ammonia, creatinine, and excess sodium and chloride ions. Nearly 180 L of fluid are filtered daily, but only 1 to 2 L actually are passed out of the body as urine.

When a patient is on hemodialysis, blood from an artery in the arm or leg flows via a tube into the artifi­cial kidney, where it spreads across a porous polymer membrane. The other side of the membrane is bathed continuously in circulating dialysis fluid. Wastes from the blood diffuse through the membrane's pores into the dialysis fluid until equilibrium is reached. This dialysis fluid is changed frequently to accelerate the blood's purification.

Coil-type dialyzers use tubular membranes wound around a plastic mesh and encased in polycarbonate, polystyrene, or acrylonitrile-styrene housings. Plate-type dialyzers are composed of laminated membranes and grooved plastic plates in alternate layers, similar to a sandwich. The convenient hollow-fiber-type config­uration uses about 10,000 hollow-fiber membranes with inner diameters of 200 to 300 μιη. The dialysis fluid passes countercurrently on the outside of the fi­bers through which the blood flows.

Since the introduction of hemodialysis by Willem J. Kolff in 1943, regenerated cellulose has been used as the hemodialysis membrane. The extracorporeal cir­cuit traditionally has been made of PVC tubing. The semipermeable membranes used for hemodialyzers are made from cuprammonium-type cellulose. The hol­low-fiber hemodialyzer is rapidly becoming the most popular type.

In recent years, new synthetic hemodialysis mem­branes made of polymethyl methacrylate, polycarbon­ate, polysulfone, polyacrylonitrile, and poly(ethyl-vinyl alcohol) have been developed in Japan and Eu­rope. These membranes are more compatible with blood than the older types.

In hemodialysis, a phenomenon called transient leu­kopenia (a temporary reduction in the number of white blood cells in the bloodstream) sometimes occurs with­in the first 30 minutes of dialysis. Although frequency

Among the various types of kidney dialysis units available are the hollow-fiber type (foreground), the coil type (right), and the flat-membrane type

and intensity of its symptoms may vary from patient to patient, leukopenia almost always occurs on the first use of a cellulosic membrane. With subsequent reuse of dialyzers, its symptoms decrease or may even disap­pear.

In the future, researchers will be placing increased emphasis on the mechanisms of leukopenia in hemodi­alysis patients and on its relation to the acute and chronic complications associated with dialyzer reuse.

Therapeutic plasmapheresis The past three decades have seen increased interest

in extending the methods developed for treating kid­ney disease to other disorders. Attempts to create an artificial liver were started nearly 20 years ago. Howev­er, lack of proper filter membranes halted this research for 10 years.

Recent developments in membrane technology and in the continuous separation of plasma from whole blood have provided the incentive for developing and using such procedures to treat liver, blood, nerve, and rheumatic diseases. In this area, much emphasis has centered on the use of therapeutic plasmapheresis.

Plasmapheresis is the separation of plasma from the cellular components of blood, with the cells being re­turned to the donor or patient. The primary use of this procedure is in collecting normal plasma from healthy donors for use, for example, in patients undergoing surgery.

Plasmapheresis also is used directly in treating a variety of chronically ill patients either by plasma ex­change or plasma perfusion. In plasma exchange, a volume of plasma is removed from the patient and replaced with an equal volume of normal plasma or a plasma substitute. In plasma perfusion, the plasma is separated from its cellular components and is then treated, either by absorption or filtration, to remove its unwanted constituents before the treated plasma and cells are returned to the patient.

Plasma separation currently is used to treat an in­creasing number of autoimmune, hematologic, and metabolic diseases that are difficult to treat by conven-

46 April 14, 1986 C&EN

tional means. With an immunological disease such as rheumatoid arthritis, this technique is aimed at remov­ing pathogenic proteins from the body, including globulins, antibodies, complement components, and circulating immune complexes.

The removal of toxic blood constituents is the aim of almost all therapeutic applications of plasma exchange. In patients with liver disease, the procedure also can be used to supply depleted substances and to regulate the composition of blood. Gorig Brunner of the University of Hanover in West Germany reports that, in severe liver ailments such as fulminant hepatic failure and chronic liver disease, all functions of the liver are se­verely impaired. The liver's three main functions of synthesis, regulation, and detoxification can, to a cer­tain extent, be supported by plasma exchange.

The products normally made by the healthy liver show a sharp decrease when the liver's synthetic func­tion fails. The replacement of diseased plasma with fresh plasma from healthy donors supplies such liver-synthesized products as fibrin, albumin, and globulin and can be considered as a support of liver synthesis.

Detoxification involves the removal of plasma that contains endogenous toxins such as mercaptans, phe­nols, free fatty acids, amines, and ammonia. Also re­moved are such serum amino acids as tyrosine, phenyl­alanine, and methionine, which serve as precursors of these phenols and mercaptans. The imbalance between both the aromatic branched and unbranched amino acids and some lipids is partially corrected by plasma exchange and may be regarded as support of the liver's regulation function.

In 1984, an estimated 90,000 therapeutic plasmapher­esis procedures were carried out in the U.S. alone, with an additional 30,000 such procedures performed in Europe and Japan. If plas­ma exchange is successful in the wide variety of dis­eases being investigated, various disease states will be competing for this sup­ply. Diseases associated with deficiencies in clot­ting factors, which are present in plasma, would also be competing with those diseases associated with an excess of plasma components.

At least 10 companies worldwide have devel­oped membrane plasma-separator modules. These components can be used with a wide range of blood flows, can permit stable and reproducible plasma separations for three to five hours, and are avail­able as sterile, disposable units. Most employ a hol­

low-fiber design, with membranes made of polyvinyl alcohol, polymethyl methacrylate, polyethylene, poly­propylene, cellulose diacetate, or PVC. The hollow fi­bers are encased in polycarbonate or polystyrene con­tainers, with the end fibers sealed with polyurethane.

The hollow-fiber membrane systems presently be­ing used and investigated generate a plasma product that contains few, if any, platelets. Thus, losses of blood cells into the plasma are minimal. For the same devices, more than 95% of the high-molecular-weight solutes (immunoglobulins, rheumatoid factor, albumin, and other proteins) pass into the plasma.

Membrane systems have certain advantages; howev­er, the selection of the membrane, its design into a module, and the determination of operating conditions have yet to be fully optimized.

Another plasma treatment system is known as plas­ma cryofiltration. Developed by the Cleveland Clinic Foundation's department of artificial organs in 1980, it involves two stages of filtration. In the first module, a hollow-fiber membrane made of cellulose diacetate separates the plasma from the blood cells. All blood cells are removed when the hollow fibers have a maxi­mum pore size of 0.2 μπ\. In the second module, the plasma is cooled to 4°C and filtered through a mem­brane also made of cellulose diacetate.

This double-step filtration uses temperature as a variable in separating pathological macromolecules from plasma. These cooled macromolecules, called cryogels, are plasma proteins that precipitate and gel at temperatures near 4 °C. Although these undesirable molecules exist in nonpathological concentrations in some diseased states, such as rheumatoid arthritis, 10 to 100 times greater levels are not uncommon.

Surrounded by two physicians and a nurse, a rheumatoid arthritis patient at the Cleveland Clinic Foundation undergoes a plasmapheresis procedure. The Cryomax system used is made by Parker Hannifin of Irvine, Calif.

April 14, 1986 C&EN 47

Special Report

Cryogels formed and retained in the membrane fil­ter are subsequently discarded. The filtered plasma is then reunited with the blood cells, warmed to physio­logical temperature, and returned to the patient.

The cryofiltration system used clinically is marketed by Parker Hannifin Corp. of Irvine, Calif. Called Cryo-max, it uses a 0.5-square-meter cellulose diacetate hol­low-fiber capillary module as the plasma separator and a 0.65-square-meter cellulose diacetate hollow-fiber capillary module as the cryofilter. Both modules are made by Asahi Medical Co. of Tokyo. This system has been used clinically in treating patients with rheuma­toid arthritis and other diseases.

Artificial organs and their biomaterials provide valu­able solutions to some major medical problems. These artificial devices, constructed from either natural or synthetic polymers, replace natural organs in function and design. Obviously, no single biomaterial is appli­cable for every potential organ replacement.

In the next decade, artificial organs will play a much greater role in preventive medicine, specifically in the early prevention of end-stage organ failure. Continued progress will depend, however, on the development of newer and more refined biomaterials to meet challeng­ing clinical needs. D

Selected readings Altieri, F. D., Artif. Organs, 7, 5 (1983). Boretos, J. W., Eden, M., eds., "Contemporary Biomaterials,"

Noyes Publications, Park Ridge, N.J., 1984. Bruck, S. D., "Properties of Biomaterials in the Physiological

Environment," CRC Press, Boca Raton, Fla., 1980. Haubold, A. D., Shim, H. S., Bokros, J. C, "Carbon in Medical

Devices" in "Biocompatibility of Clinical Implant Materials," Vol. 2, Williams, D. F., éd., CRC Press, Boca Raton, Fla., 1981.

Joyce, L. D., DeVries, W. C, Hastings, W. L, Olsen, D. B., Jarvik, R. K., Kolff, W. J., Trans. Am. Soc. Artif. Intern. Organs, 29, 81 (1983).

Kambic, H. E., Murabayashi, S., et al., Cleve. Clin. Q., 51, 105 (1984).

Malchesky, P. S., Sueoka, Α., Matsubara, S., Wojcicki, J., Nosé, Y., "Membrane Plasma Separation" in "Plasmapheresis: Therapeutic Applications and New Techniques," Nosé, Y., Malchesky, P. S., Smith, J., Krakauer, R., eds., Raven Press, New York, 1983.

McMillin, C, Artif. Organs, 7, 78 (1983). Rubin, L. R., éd., "Biomaterials in Reconstructive Surgery," C.

V. Mosby Co., St. Louis, 1983. Stanley, J. C, éd., "Biologic and Synthetic Vascular Prosthe­

sis," Grune & Stratton, New York, 1982. Szycher, M., Bernhard, W. F., "Hemocompatible Interfaces in

Artificial Heart Devices" in "Biocompatible Polymers, Metals, and Composites," Szycher, M., éd., Technomic Publishing Co., Lancaster, Pa., 1983.

Reprints of this C&EN special report will be available at $5.00 per copy. For 10 or more copies, $3.00 per copy. Send requests to: Distribution, Room 210, American Chemical Soci­ety, 1155—16th St., N.W., Washington, D.C. 20036. On orders of $20 or less, please send check or money order with request.

Helen E. Kambic has been di­rector of the biomaterials pro­gram in the Cleveland Clinic Foundation's department of artificial organs since 1981. After receiving a B.S. in biology/ chemistry in 1967 from Saint Francis College in Pennsylva­nia and an M.S. in biology/ chemistry in 1969 from Ford-ham University, she joined the staff of the Kidney Dis­ease Institute of the New York State Department of Health in Albany as a research scientist. In 1972, she began working in the Cleveland Clinic Foundation's department of artificial organs, where she first served as a research assistant and later as a project leader in the biomaterials program. She has authored more than 70 publications dealing with bio­materials, artificial organs, and related subjects. She has been active in the International Society for Artificial Organs and the American Society for Artificial Internal Organs.

Shun Murabayashi, who was born in Japan, did all of his university work at the Uni­versity of Hokkaido's school of engineering. With the aid of a scholarship from the Jap­anese government, he carried out his advanced studies lead­ing to a Ph.D. in engineering, with a specialty in radiation chemistry, in 1978. That year, he became a research fellow in the department of artificial organs at the Cleveland Clinic Foundation. Since 1981, he has been scientific director of the department's biocompatibility program. For the past three years, he also has been an adjunct associate professor in the Institute for Biomedical Engineering Research at the Univer­sity of Akron.

Yukihiko Nosê has been head of the Cleveland Clinic Foun­dation's department of artifi­cial organs since 1967. He joined the department as head of the research laboratory in 1964. Born in Japan, he re­ceived his M.D. in 1957 and his Ph.D. in surgical science in 1962 from the University of Hokkaido. From 1962 to 1964, he was a surgical re­search associate at Maimoni-des Hospital in New York City. He has published more than 400 scientific papers on artificial organs and has authored or edited eight books on the subject. A member of the editorial boards of several scientific journals, he also has served as a consultant and study section member for the National Institutes of Health's programs dealing with artificial kidneys, bioengineering, and surgery.

48 April 14, 1986 C&EN