Technologies Oxygenation Pumping Of Blood In Assisted Perfusion Biology Essay
The cardiovascular system is composed of the heart, blood, and blood vessels. This system functions in the delivery of oxygen, nutrient molecules, and hormones and the removal of carbon dioxide, ammonia and other metabolic wastes. The heart pumps oxygen-rich blood through arteries. Veins carry blood from various parts of the body back to the heart. Capillaries are the points of exchange between the blood and surrounding tissues. The cardiovascular system is one of the most important systems in the body as none of our cells and tissues can function without adequate oxygen and thus blood supply.
The essential goals of cardiopulmonary bypass (CPB) are to maintain adequate circulation and respiration by diverting blood flow to an extracorporeal circuit and that functionally replaces the heart and lung to facilitate surgery of the heart and great vessels.
The essential components of the CPB system remains the same despite the different machine designs. The main components of the artificial circuit include the cannulae, venous reservoir, oxygenator, heat exchanger; pump itself, and the tubing. Venous blood is drained by gravity into the reservoir via a cannula placed in the right atrium or a large vein, pumped through the oxygenator and the heat exchanger, then returned into the patient’s arterial system via a cannula in the aorta or other large artery. Transit though the oxygenator reduces the partial pressure of carbon dioxide in the blood and raises oxygen content.
Major advances in CPB have been achieved over the years. Oxygenators are more efficient, the materials used are more biocompatible, the pumps are designed to improve fluid dynamics and the circuits are therefore safer.
HISTORY OF CPB
Open heart surgery and CPB developed in incremental steps by many investigators who were often in tangential or even irrelevant fields. A collaboration of physiologists, chemists, physicists, engineers and physicians were all driven by a common goal of diverting the circulation through synthetic devices that could effectively sustain patients who otherwise would be succumb to treatable lesions of the heart, lungs or other vital organs (Stammers 1997).
The concept of perfusion of blood and the use of mechanical devices for oxygenation was initially shown by La Gallois in 1812, where extracorporeal circulation was feasible with experimentation that involved the injection of blood into the carotid arties of decapitated rabbits. La Gallois hypothesised that it might be possible to replace the function of the heart with an artificial pump that could be used to perfuse the body (Stammers 1997).
The idea of the roller pump was introduced in 1855 by Porter and Bradley. This device was used with the idea of using rollers that exerted a force on a tube in a rotational manner to drive blood through the tube in one direction. This pump was used for transfusion of blood directly from donor to recipient.
Some studies were carried out with perfusion of muscles and organs in the years to follow, but not until it was pointed out by Brown-Sequard in 1858, that the success of such perfusions depended on the use of oxygenated blood. Brown-Sequard arterialised desaturated blood by whipping it into a frothy foam and perfused it into the served arms of decapitated criminals. He made the observation of temporary disappearance of the rigor mortis of muscles. Since then it was definite that blood was essential for successful perfusion.
The first report of artificial oxygenator was made in 1869 by Ludwig and Schmidt, who oxygenated defibrinated blood by shaking it with gas in a balloon. This was one of the first reports to use defibrinated blood, caused by the process of whipping which also showed the importance of anticoagulation in facilitating extracorporeal flow (Stammers 1997).
The first development of a heart-lung machine and earliest use of bubble and film-type oxygenators were by Von Frey and Gruber as they described a blood pump in 1885 in which gas exchange occurred as blood flowed into a thin film over the inner surface of a slanted rotating cylinder where it was exposed to oxygen. In 1895 Jacobi passed blood through an excised animal's lung that was aerated by artificial respiration.
In 1916, Jay McLean discovered that “courin” which is extracted from an ox’s heart had anticoagulation properties. This anticoagulant was named “heparin” by professor Howell. Heparin was extracted in large quantities in 1933 by Charles and Scott from beef lungs, which was just in time for the Gibbons’ studies. Chargaff and Olson then reported in 1939 that protamine neutralised heparin in vitro (Edmunds Jr 2002).
In 1915 Hooker described an oxygenator in which blood was dropped on a rotating horizontal disc. The centrifugal force spread the blood out to a thin film, and it was then thrown off the disc to the inside of a cylinder, where it fell to the bottom and was collected and transfused. Bayliss in 1928 then increased the efficiency of this device by adding more than one disc. This adaptation increased the surface area, enabling perfusion of large subjects.
Coinciding Hooker in 1915, Richards and Drinker reported on an oxygenator for isolated organ perfusion. Their device used silk screens over which blood was poured. The blood ran down the screens which exposed two sides to oxygen, giving twice the efficiency of the solid disc oxygenator. The blood also ran down the screens in a turbulent fashion which caused mixing and more even oxygenation of the deeper layer of the blood film (Edmunds Jr 2002).
In 1926, Professors S. S. Brukhonenko and S. Tchetchuline in Russia designed a machine that used an excised lung from a donor animal as an oxygenator and two mechanically actuated blood pumps. Their machine was used initially to perfuse isolated organs but later was used to perfuse entire animals.
Dale and Schuster described and produced a prototype of the pumping mechanism for extracorporeal circulation in 1928. The device was a valved pump that delivered a pulsatile wave, which was created to mimic that of the normal circulation. This later became the predominant device used for extracorporeal perfusion (sigmamotor pump) (Stammers 1997).
Daly and Thorpe then utilised the idea of Hooker and added three of the horizontal discs with a concentric arrangement in 1933.
In 1931 John Gibbon hypothesised the idea that a heart-lung machine would temporarily replace the function of a patient’s heart and lungs to permit safe entry and exit to the patient’s great vessels of the heart. John Gibbon probably contributed more to the success of the development of the heart-lung machine than anyone else.
Debakey modified the twin roller pump that had been described in 1855 by Porter and Bradley. In 1934 Debakey’s pump was capable of outputs of up to 5L/min and became the standard by which all positive displacement pumps used to day for extracorporeal flow have been modelled (Stammers 1997).
In line with Debakey’s pump in 1934, Gibbon began his work to build the first heart-lung machine from rubber, glass, homemade valves, a rotating cylinder, finger cots and assorted laboratory air pumps and paraphernalia. He cannulated the jugular vein and femoral artery of an ether-anesthetised cat and bypassed its circuitry with the rotating cylinder oxygenator that he invented and two finger cot pumps. Since then, Gibbon worked to perfect his heart-lung machine (Edmunds Jr 2002).
In 1937, Gibbon reported the first successful demonstration that life could be maintained by an artificial heart and lung and that the native heart and lungs could resume function.
Gibbon's resumed his work at the shortly after WWII, and in that time, other groups also worked on heart-lung machines. In 1950, Bigelow and his colleagues introduced hypothermia for correction of heart defects. From there a series of surgeries were carried out in the human heart. The major problem associated with the development of early blood gas exchange devices was that the total delivery of oxygen and removal of carbon dioxide could only provide partial support of the circulation. In 1951, Dennis and his associates combined the features of both Gibbon’s and Bjork’s rotating disk type oxygenator, replacing the system with rotating screens that were immersed into a reservoir of blood. Using this design, Dennis performed the first total cardiopulmonary bypass in a 6 yea old patient with an incomplete atrioventricular canal (Stammers 1997, Edmunds Jr 2002, Gravlee, Davis et al. 2007).
Using the hypothermic technique, Lewis closed an ASD in the world’s first successful operation within the open human heart under direct vision. Lewis used blood inflow stasis and hypothermia by surface cooling.
Gibbons work then followed by the use of a stationary vertical screen oxygenator designed by IBM to perform the first successful bypass in man and closure of an ASD in 1953 (Stammers 1997).
The idea of “cross-circulation” was proposed and used by Lillehei in 1955, where the arterial and venous circulation of a mother and child were connected by tubing in series. The mother’s heart and lungs maintained the circulatory and respiratory functions of both, whilst surgeons operated on the child’s heart.
DeWall made a huge advance in 1955 by developing a bubble oxygenator with a unique method for removing bubbles from the freshly oxygenated blood. However, bubble oxygenators produced a constant stream of microscopic bubbles that escaped filters and bubble traps. Soon after in 1963, Bodwell explored the use of tubular capillary membrane oxygenators and the first commercially available membrane oxygenator composed of silicone rubber sheets with inlet and outlet gas tubes was designed by Bramson in 1965. Lamdon and his colleagues and Bramson’s group developed clinically used membrane oxygenators using a stack design in 1969 (Stammers 1997, Edmunds Jr 2002, Gravlee, Davis et al. 2007).
Mr Ben Lipps design introduced the hollow-fibre oxygenator design in 1970. He developed a method to make miles of 200micrometre diameter hollow polycarbonate siloxane fibres with a discrete wall thickness which were bundled together, potted both ends in Silastic, cut off the potted ends to expose the fibres and placed whole into a jacket into which oxygen was admitted (Stammers 1997, Edmunds Jr 2002, Gravlee, Davis et al. 2007).
In the 1970s several engineers collaborated in the design of centrifugal pumps. This device moved blood by the process of centrifugation through the creation of a vortex in a constrained plastic housing. The centrifugal force is created without the action of any occlusive mechanisms so that the blood trauma normally created by positive displacement pumps could theoretically be reduced. Centrifugal pump use continued to rise as they are considered user friendly (Stammers 1997).
Principle of gas exchange in the lungs
Inhaled air travels through the airways to the alveoli. Blood is pumped out of the heart through the pulmonary arteries to a network of capillaries that surround the alveoli. The oxygen of the inhaled air diffuses out of the alveoli into the blood while carbon dioxide in the blood moves into the alveoli to be exhaled. The alveoli are the final branchings of the respiratory tree and act as the primary gas exchange units of the lung. The gas-blood barrier between the alveolar space and the pulmonary capillaries is extremely thin, allowing for rapid gas exchange. To reach the blood, oxygen must diffuse through the alveolar epithelium, a thin interstitial space, and the capillary endothelium; CO2 follows the reverse course to reach the alveoli. The oxygen-rich blood is then returned to the heart through the pulmonary veins (Martini & Nath 2009, Silverthorn 2001).
Gas exchange in the Lungs serves two functions; the first is that it replenishes the blood’s oxygen supply in the pulmonary capillaries, second is that it removes carbon dioxide from the pulmonary capillaries. Gas exchange occurs between the air in the alveoli, through the respiratory membrane, to the red blood cells in the blood of the pulmonary capillaries. Carbon dioxide’s membrane solubility is 20 times greater than that of oxygen, so CO2 can diffuse across the respiratory membrane much more rapidly (Martini & Nath 2009, Silverthorn 2001).
Under most circumstances, diffusion distance, surface area, and membrane thickness are constants and are maximised to facilitate diffusion. The most important factor for gas exchange in the alveoli is the concentration gradient where the flow of gases occurs from high partial pressure regions to low partial pressure regions, and this rule governs the exchange of oxygen and carbon dioxide in the alveoli (Martini & Nath 2009, Silverthorn 2001).
Pumps and Oxygenators
In the development of CPB, there were two basic functions that had to be addressed: methods to oxygenate the blood and the means to pump the blood through a device and back to the patient. Today many different oxygenators and pumps are available commercially, enabling perfusionists and surgeons to choose the one that best fills their requirements. However, the pioneers of CPB had to build everything they needed.
A pump is required to produce blood flow. Currently, roller and centrifugal pump designs are the standard of care. Both modern designs can provide pulsatile (pulsed, as from a heartbeat) or non-pulsatile blood flow to the systemic circulation.
The idea of the roller pump was introduced in 1855 by Porter and Bradley. In the beginning designs with up to three rollers had been proposed. But the twin roller pump has now become the standard (Gravlee, Davis et al. 2007).
Many of the early experiments used a pump developed by Dale and Schuster in 1928, that caused blood within it to move forward. Another early machine that was used clinically was the sigmamotor pump. This device had occlusive fingers that rhythmically compressed the tubing to propel the blood in a forward direction. These pumps produced a pulsatile flow and were thought at the time to be more physiologic than the roller pump. However, because of their complexity, the difficulty with synchronisation at high flow rates, and the extensive hemolysis they caused, the pulsatile device lost favour over time, and the roller pump became the standard.
The basic roller pump consists of two rollers, 180 degrees apart, that rotate in a circle through a half circular raceway. A length of flexible tubing between 1/4 and 5/8 inch inner diameter is placed between the rollers and the raceway. The rollers rotating in a circular movement compress the tubing against the raceway, squeezing the blood ahead of the rollers. The rollers are set to almost completely occlude the tubing, and operate essentially as a positive displacement pump, each passage of a roller through the raceway pumping the entire volume of the fluid contained in the tubing segment between the rollers. As a positive displacement pump, high positive pressures can be generated at the pump outlet and high suction (negative) pressures can be generated at the pump inlet. Roller pumps are typically driven by a constant speed motor which draws blood at a substantially constant rate (Morgan, Codispoti et al. 1998, Mulholland, Shelton et al. 2005).
Occlusive roller pumps positively displace blood through the tubing using a peristaltic motion. As the rollers roll the blood through the tubing, the tubing is intermittently occluded where positive and negative pressures are generated on either side of the point of occlusion. The direction of the blood flow can be altered by the direction of pump head rotation. Displacement roller pumps are relatively independent of circuit resistance and hydrostatic pressure, and output depends on the number of rotations of the pump head and the internal diameter of the tubing used. Pulsatile or non-pulsatile (laminar) flow can both be achieved by positive displacement pumps, which both have their advantages and disadvantages.
Pulsatile flow, being more physiologically correct than constant flow, may have a beneficial impact on the efficacy of the extracorporeal perfusion. This can result in improved patient outcomes following cardiac bypass surgery. Pulsatile flow is important for cerebral oxygenation and autoregulation, and for other tissue perfusion and capillary blood flow. Pulsatile flow also stimulates the endothelial cells that line normal blood vessels, causing them to elongate and secrete local factors (Nitric oxide and prostaglandins) into the vessel wall (intramural release) and into the blood stream (intraluminal release). These factors maintain vascular tone (vessel relaxation), inhibit clot formation on the vessel inner surface (platelet adhesion and aggregation), inhibit monocyte adherence and chemotaxis, and inhibit smooth muscle cell migration and proliferation (Gosh, Falter et al. 2009, Mulholland, Shelton et al. 2005).
An acceptable way of creating pulsatile flow is to vary the speed of the pump in a cyclical manner. This is easily accomplished electronically by the pump controller. However, the inertia of the spinning elements of the pump tends to render the resulting waveform more sinusoidal than the natural heartbeat waveform and forces the wave period to be longer than the natural period. In addition, the components of the bypass circuit downstream of the pump, such as the oxygenator and arterial filter, also damp the pulses due to their volumetric holdup (Gosh, Falter et al. 2009).
Non-pulsatile flow in comparison is known to have a detrimental effect on cell metabolism and organ function. The shear stress generated by the roller pumps during pulsatile perfusion however, leads to increased hemolysis. Another disadvantage of roller pumps is the occurrence of cavitation, which is the formation and collapse of gas bubbles due to the creation of pockets of low pressure by precipitous change in mechanical forces due to sudden occlusion of the inflow to the pump (Gosh, Falter et al. 2009).
Nonocclusive roller pumps were also used. The Sausse (Rhone-Poulenc) pump stretches a distensible silicon tubing of an ovoid or elliptical cross section and shape memory compliance longitudinally around pin rollers mounted 120 degrees apart on a rotating wheel, the tubing being held in place below the wheel by connectors retained in a notched fixed base. This tubing, herein called a "header" tubing, is not compressed against a raceway (as for a roller pump), but is held in tension across the rollers, restricting the lumen of the header tubing across the rollers. This segments the header tubing into portions defined by leading and trailing adjacent rollers. The rotation of the wheel moves fluid captured between adjacent rollers in the direction of the rotation. The material and thickness of the wall of the header tubing are selected so the tubing between the rollers will expand or collapse as a function of pump inlet pressure (available venous return). Collapse of the tube will restrict the flow rate of the liquid as a function of the pump inlet pressure. When the obstruction is released, blood flows downstream propelled by the increased stroke volume of the distended header tubing. The header tubing stretched over the rollers therefore functions as a built-in capacitance reservoir, eliminating the need for the reservoirs that are required for roller and centrifugal pumps (Gravlee, Davis et al. 2007, Morgan, Codispoti et al. 1998).
Rotary or centrifugal pumps have three basic designs being, axial, diagonal and radial which all have different qualities. The centrifugal pump is composed of cone with impellers encased in an outer plastic housing. The cone spins as a result of the magnetic force that is generated when the pump is activated. The spinning cone creates a negative pressure that sucks blood into the inlet, creating a vortex by the rotating impellers. The energy created in the cone creates pressure and blood is then forced out of the outlet. As a safety feature, this pump disengages when air bubbles are introduced. The centrifugal force draws blood into the center of the device. Blood is propelled and released to the outflow tract tangential to the pump housing. Rotational speed determines the amount of blood flow, which is measured by a flowmeter placed adjacent to the pump housing. If rotational frequency is too low, blood may flow in the wrong direction since the system is non-occlusive in nature (Gravlee, Davis et al. 2007, Gosh, Falter et al. 2009). Centrifugal pumps may produce less hemolysis and platelet activation than roller pumps, but this does not correlate with any difference in clinical outcome. These pumps are more expensive as the pump head is of single use and may be prone to heat generation and clot formation on the rotating surfaces in contact with the blood. They are therefore used for complex surgery of prolonged duration which the damage to blood associated with roller pumps may be theoretically disadvantageous (Gosh, Falter et al. 2009).
Many different paths were likewise taken to accomplish oxygenation of the blood. For a short period, cross-circulation was used in 1954 in which the arterial and venous circulation of a mother and child were connected by tubing in series. The mother’s heart and lungs maintained the circulatory and respiratory functions of both, whilst surgeons operated on the child’s heart. Other surgeons used human donor lungs incorporated into the bypass system. Both of these approaches were abandoned early because of their high risk and complexity in a clinical setting. Others turned to mechanical methods, which were adapted and refined to the state we have today. There were many problems with early mechanical oxygenators as they were large and cumbersome, with complex assemblies that had to be dismantled, cleaned, sterilised and reassembled each time a procedure was performed. Oxygenators such as the rotating horizontal disc and silk screens over which blood was poured to the use of the vertical screen oxygenator by Gibbon and the bubble oxygenator by DeWall which made a huge advance in 1955 with a unique method for removing bubbles from the freshly oxygenated blood.
In general, three types of oxygenators has been used and further developed over the years with two different approaches, one involving direct contact between blood and oxygen, and others in which a membrane was placed between the blood and oxygen.
In Film-type oxygenators where direct blood interaction occurred, various techniques were employed to produce a thin blood film, and gas exchange takes place on the surface of the exposed blood film. Because there is no mechanical introduction of the gas into the blood, the blood trauma caused by this pump oxygenator is generally less than that of other oxygenators that were developed later on. A large surface area is necessary for adequate gas exchange therefore requiring a high priming volume of this type of device (Iwahashi, Yuri et al. 2004).
Although many different types of bubble oxygenators have been developed, typically all bubble oxygenators have three distinctly identifiable sections: an oxygenating section, a defoaming section, and an arterial reservoir. Direct blood interaction with the bubble oxygenator therefore also occurred.
Oxygen and other gases are introduced into the oxygenating section through small tubes or a porous member. Each small tube or porous member creates small bubbles which are dispersed in the blood.
Oxygen and carbon dioxide are exchanged across the boundary layer of the blood and gas bubbles. In most devices, the majority of the oxygenation takes place in this section. However, in some devices the blood has a short residence time in the oxygenating section; thus, oxygenation continues to occur as the blood passes through the subsequent defoaming section (Iwahashi, Yuri et al. 2004, Haworth 2003).
As oxygen is bubbled through the blood in a bubble oxygenator, a certain amount of foaming necessarily occurs. This foam, and any entrapped air bubbles, must be removed from the blood before it is reinjected into the patient; otherwise, the entrapped air bubbles can form an embolus which can severely injure or kill the patient. Defoaming is generally accomplished by passing the blood over a material having a large surface area which has been treated with a defoaming agent (Stammers 1997).
The arterial reservoir provides an area where the defoamed blood is collected before reinjection into the patient. The reservoir acts as a safety feature in helping avoid accidental pumping of air into the blood lines. Should the blood supply to the oxygenator be accidently stopped, the reservoir must contain sufficient blood to allow the perfusionist to stop the output from the oxygenator before air enters the arterial line (Iwahashi, Yuri et al. 2004, Haworth 2003).
The simple design of the bubble oxygenator relied on a hydrostatic pressure head from the patient to the mixing chamber connected by the venous line to the right atrium or vena cava. Bubble oxygenators also have the heat exchanger downstream from the bubble chamber but proximal to the blood pump so that heat exchange occurs simultaneous with gas exchange. Blood trauma induced by bubble oxygenators is high compared to any other type of oxygenator. It was used for short-duration bypass procedures because it is inexpensive and easy to use (Gravlee, Davis et al. 2007).
Bubble oxygenators produced a constant stream of microscopic bubbles that escaped filters and bubble traps. Incorporated filters in bubble oxygenators helped reduced the number of particulate emboli, but gaseous, fat and fibrin emboli, platelet and leukocyte aggregates and red cell debris still entered the arterial perfusion line even after introducing arterial line filters.
Despite the success of bubble oxygenators in expanding the role of cardiac surgery, there was a growing concern that the direct contact of blood and gas caused significant derangements in the formed elements of blood, particularly when CPB lasted more than 1 hour (Stammers 1997). Work on membrane oxygenators was already in progress before the problem of the microembolic problems of the bubble oxygenator was discovered. Over the years many type of membrane oxygenators were invented including the spiral coil membrane oxygenator and the hollow-fibre membrane oxygenator (Edmunds Jr 2002).
Two general classes of application are foreseen for membrane oxygenators. One being to replace the existing blood-gas contact and the second was to provide periods of prolonged extracorporeal support, lasting days or possibly weeks. Membrane oxygenators separated the blood-gas interface and mimicked the endogenous alveolar capillary membrane features (Drinker 1972).
One of the improvements that increased the efficiency of gas exchange in membranes was the discovery of laminar blood flow, which created boundary layers along the membrane material. The layers in contact with the membrane material caused velocities to be slow and hence increased gas exchange. Thus, reducing blood film thickness and disrupting the laminar flow characteristics became apparent with the introduction of spacers, screens and sheets to promote gentle mixing of the layers of blood to optimise gas exchange (Stammers 1997).
The theory of membrane oxygenators is that they have the same effect as the patient’s biological lung where the laws of diffusion take place (diffusion of oxygen and carbon dioxide occurs from high to low concentrations). It is however not possible to produce a blood film in the membrane oxygenator as thin as that existing in natural lung tissue. Thus, it is always important to maintain effective mixing of the blood in the thicker blood film used in these oxygenators (Drinker 1972).
If effective blood mixing can be achieved, red cells that contain low levels of oxygen are constantly brought into close proximity to the gas exchange surface of the membrane oxygenator. The oxygen transfer from the surfaces of the gas exchange membrane to the plasma and then to red cells is thus effectively maintained because the distance of gas transfer would always be kept to a minimum. If the membrane is too thick, thus the distance between the gas and the blood is increased, the efficiency of the oxygenator will be poor (Iwahashi, Yuri et al. 2004).
There are two types of membrane oxygenators described as microporous and nonporous membranes. The plate-type nonporous gas permeable membrane consisted of supported sheet membranes. The also nonporous coil type membrane oxygenators were first used by using polyethylene membranes where multiple coils were used for effective gas exchange. This later was developed to the use of silicone rubber envelope reinforced with nylon knit wrapped around a central core and pure humidified oxygen was passed through it under a negative pressure. Blood flowed across the flat tubing parallel to the axis of the cylinder (Iwahashi, Yuri et al. 2004).
Silicone membranes dipped in a nylon mesh were then introduced as they had a higher permeability to oxygen and carbon dioxide. Silicone membranes required special priming techniques to ensure that all air was displaced from the system before perfusion of the patient was started. The biological compatibility of silicone made this material especially attractive for long term perfusion. The design of silicone membrane oxygenators in 1963 with a spiral-wound sheet configuration is still the standard to date for long-term extracorporeal perfusion.
Microporous membranes had microscopic pores throughout the membrane material. There are two major types of configurations of the membrane material being the hollow fibre type and the folded sheets with hollow fibres making up the predominance of devices. Historically there were two options of blood flow through a hollow fibre oxygenator: through the fibre or around it and the blood may flow either perpendicular to the fibre bundle or in the direction of the fibres (Gravlee, Davis et al. 2007).
The first hollow-fibre oxygenator with a microporous membrane was developed in 1981 by Terumo Corporation. Blood flow in this type of oxygenator was through the hollow fibres. The fibres were coated with a thin silicone film which did not completely prevent plasma leakage resulting from insufficient coating. The coating was then laid in thicker layers to prevent leakage (Iwahashi, Yuri et al. 2004). This was later abandoned due to high transmembrane pressure, activation of platelets and increased hemolysis.
A new design of microporous membrane oxygenators was then developed where the flow arrangements were reversed and the blood flow is now outside of the hollow fibres and gas flow is down the fibre lumen. This is a significantly more efficient arrangement for gas transfer, so that smaller membrane area will provide adequate gas transfer capability, and it is also substantially noncompliant so prime volume does not change with back pressure on the oxygenator. The blood flow outside the fibre is in a counter current direction to the gas flow inside the fibre which confers the advantage of optimised gas gradients during the dwell time (Gravlee, Davis et al. 2007).
Membrane oxygenators produced far fewer gaseous emboli than bubble oxygenators and also produce fewer particulate emboli as well as reducing the postoperative neurologic dysfunction. Long term support of a patient became known as Extracorporeal Membrane Oxygenation (ECMO) where silicone membrane oxygenators were suitable. That is because plasma is eventually able to wet a microporous membrane, a phenomenon called plasma breakthrough, and when that happens gas transfer rates through the membrane are substantially reduced. A solid silicone membrane does not suffer from this effect and so may be used for support of a patient for several days without significant decline in performance. Membrane oxygenators made extracorporeal support for acute respiratory distress possible and provided enough time for a patient’s heart to recover (Edmunds Jr 2002, Haworth 2003).
Extracorporeal circulation is known to elicit a profound inflammatory response by activating several protein-mediated systems involved in cellular immunity and protection. Extracorporeal circulation stimulates complement system, coagulation cascade as well as the release of powerful substances. Efforts to ameliorate these toxic effects have resulted in the coating of the circuit of part of the circuit with biocompatible coatings to prevent or reduce such effects.
The essential components of the CPB system remain the same despite the different machine designs. Major advances in CPB have been achieved over the years. In the development of CPB, there were two basic functions that had to be addressed: methods to oxygenate the blood and the means to pump the blood through a device and back to the patient. Today many different oxygenators and pumps are available commercially, enabling perfusionists and surgeons to choose the one that best fills their requirements. However, the pioneers of CPB had to build everything they needed.
Pulsatile and nonpulsatile roller pumps with an occlusive and non-occlusive method of pumping blood were designed. Centrifugal pumps were then introduced which reduced hemolysis of the blood and were more efficient. Film type, oxygen type and later membrane type oxygenators were developed to increase the efficiency and function of the extracorporeal system close to that of biological heart and lungs. Oxygenators are now more efficient, the materials used are more biocompatible, the pumps are designed to improve fluid dynamics and the circuits are therefore safer.
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