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Oxygenator anatomy and function
Oxygenator Anatomy and Function J u l i e A. Wegner, PhD, CCP The natural lung is the organ responsible for oxygen and carbon dioxide exchange between the blood and the outside environment. This function is accomplished by the large surface area and high permeability of the gas exchange interface, the alveolar-capillary membrane. These same features are fundamental to the design of an artificial lung, or oxygenator. Additional lung-like features essential to the design of an ideal oxygenator include the ability to achieve balanced oxygen and carbon dioxide exchange with minimal blood damage and blood activation. The purpose of this
review is to present the past and current developments of the oxygenator designs in terms of the structural and functional features of the natural lung as well as the limitations in the ability to mimic the features of the lung because of the lack of appropriate technology.
HE NATURAL lung is the organ responsible for oxygen and carbon dioxide exchange between the blood and the outside environment. The challenge of replacing the function of the natural lung with an artificial lung, or oxygenator, requires a thorough understanding of the structural and functional features that make the natural lung an efficient gas exchange organ. The oxygenator must be able to provide efficient and balanced gas exchange under extreme conditions of hemodilution and temperature with minimal blood damage and blood activation. The purpose of this article is to present the past and current developments of the oxygenator designs in terms of the structural and functional features of the natural lung as well as the limitations in the ability to mimic the features of the lung because of the lack of appropriate technology. Brief reviews of lung anatomy and function and the development of oxygenators are presented. In addition, the factors that limit the gas exchange efficiency and long-term utilization of the present-day membrane oxygenators are presented, as well as possible solutions to improvements in oxygenator design.
membrane, interstitial space, capillary basement membrane, and capillary endothelium. The average thickness of the alveolarcapillary interface is 0.5 lain and is highly permeable to both 02 and CO2. The large surface area and low resistance to gas diffusion of the alveolar-capillary interface are the two main factors responsible for the high efficiency of gas exchange in the lung. Gas exchange at the alveolar-capillary interface occurs through passive diffusion, as described by Ficks Law: Vgas (Volume/Unit Time) = (A/T) × D × (P1 - P2), where A equals surface area, T equals interface thickness, D equals diffusion coefficient, and P1 - P2 is the pressure gradient across the interface (ie, PAO2 - Pc o , where PAO2 is the alveolar oxygen tension and Pc o is the cat~illary oxygen tension). The diffusion coefficient is d~pendent on the solubility of the gas, temperature, and permeability of the interface. Differences in the diffusion coefficient, pressure gradient, and mechanism of transport in the blood between of 02 and CO2 suggest that different factors will limit 02 and CO2 exchange at a gas exchange interface, such as the alveolar-capillary membrane. Based on the components of the Fick equation, it can be inferred that the exchange of a given gas (Vgas) can be increased by increasing the surface area available for gas exchange, decreasing the thickness of the interface, increasing the diffusion coefficient, or increasing the pressure gradient. Under normal conditions, the interface thickness (T) and diffusion coefficient (D) are constant, and thus do not contribute to significant changes in gas exchange capacity. At rest, the entire alveolar-capillary surface area is not available for gas exchange because of mismatching in ventilation and perfusion. In other words, capillary blood flow is not always in contact with ventilated alveoli and vice versa. However, the gas exchange surface area can be rapidly expanded by alterations in the distribution of gas flow and blood flow in the lung. Thus, the natural lung has the ability to regulate gas exchange by adjusting the surface area available for gas exchange. Gas exchange can also be influenced by the individual pressure gradients across the interface. The gradients are dependent on both the O2 and CO2 content of the blood in the capillaries (ie, venous admixture) and the composition of the gas in the alveoli. Under conditions of room air and normal
T
STRUCTURE AND FUNCTION OF THE NATURAL LUNG
The lung is an efficient gas exchange organ, in part because of the large surface area generated by the branching network of the air conduction and blood-carrying conduits. The airways, or bronchi, sequentially branch into bronchioles, which in turn branch into alveolar ducts and finally individual alveoli. The spherical design of the alveoli creates a large surface area for gas exchange. Gas flow into and out of the alveoli is regulated at the level of the bronchioles, which are lined with smooth muscle. A similar branching pattern exists for the pulmonary blood vessels, which branch from the pulmonary arteries to the arterioles and finally into the capillaries that vascularize the alveoli. The pulmonary capillaries are 0.5 to 1.0 mm in length and 3 to 7 bun in diameter. Because at least 95% of oxygen (02) is transported in the blood as oxyhemoglobin (HbO2), the small diameter of the pulmonary capillaries ensures that the red blood cells are in constant contact with the endothelial cell surface, and thus can participate in gas exchange (ie, a high surfacearea- to-volume ratio). Regulation of blood distribution in the pulmonary vasculature occurs primarily at the arteriolar level. The convergence of the airway and blood vessel networks creates the gas exchange interface, the alveolar-capillary membrane. There are an estimated 300 million alveoli in the lungs, and the total alveolar surface area of the lung ranges from 60 to 80 m2. The alveolar-capillary interface consists of several layers, including the alveolar epithelium, alveolar basement
Copyright© 1997by W.B. Saunders Company KEY WORDS: lung, gas exchange, oxygenator, biocompatibility, membrane oxygenator, blood flow dynamics
From the University Medical Center, Tucson, AZ Address reprint requests to Julie A. Wegner, PhD, CCP, University Medical Center, 1501 N Campbell Ave, Room 4402, Tucson, AZ 85724. Copyright © 1997 by W.B. Saunders Company 1053-0770/97/1103-000353.00/0
Journal of Cardiothoracic and Vascular Anesthesia, Vol 11, No 3 (May), 1997: pp 275-281
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ventilation, the pressure gradient for 0 2 is approximately 40 mmHg, and approximately 6 mmHg for CO2. The gradients for 02 and CO2 can be modified by alterations in the metabolic rate of the individual as well as the composition of the inspired gas (O2) and the ventilation rate (CO2). Another important feature of the natural lung is that 02 exchange and CO2 exchange are metabolically balanced. Although 02 and CO2 exchange are metabolically coupled, under artificial conditions such as artificial ventilation, they can be modulated independently. A final factor that can influence the volume of gas exchange across the alveolar-capillary interface is the exposure time of the blood at the interface. At rest the contact time is approximately 0.75 seconds; however, because of the efficiency of gas exchange in the lung, gas transfer normally takes place within 0.25 seconds. Thus, the contact time is normally not a limiting factor in gas exchange in the healthy lung, even under conditions of maximal exercise. Another important feature of the natural lung is the continuous gas exchange performance over the lifetime of the individual. Gas exchange performance is maintained at both the alveolar and capillary level. The alveolar wall is composed of a layer of epithelial cells and the epithelial basement membrane. Cells within the alveolar epithelial layer secrete a phospholipid substance called surfactant, which prevents the collapse of alveoli during ventilation. Also found within the alveolar wall are macrophages. Alveolar macrophages are highly phagocytic cells that remove debris from the alveolar spaces. Thus, both surfactant and alveolar macrophages maintain the surface of the alveolar interface. The pulmonary capillaries are lined with the endothelial cells. A primary function of the vascular endothelium is to keep the blood in a fluid state through regulation of hemostasis and vascular tone. Thus, the vascular endothelium maintains blood flow to the capillaries and ensures contact of blood at the alveolar interface. The vascular endothelium also serves as a barrier between the blood and subendothelial surfaces. Exposure of blood to the subendothelial surfaces results in activation of hemostasis and inflammation. Thus, the vascular endothelial layer plays an important role in ensuring biocompatibility within the vasculature. DESIGN AND DEVELOPMENT OF THE ARTIFICIAL LUNG, THE OXYGENATOR
The efficiency of gas exchange in the natural lung is mainly attributable to the large surface area generated by the airway and circulatory networks and the low resistance to diffusion. These same features are essential to the design of an efficient artificial lung. Other necessary features of an ideal oxygenator include minimal trauma to blood, thromboresistance, minimal reaction with blood components, minimal generation of gaseous microemboli, ability to maintain performance over long periods, low prime volume, consistent physical properties, reliability, ease of use, and low cost. This section briefly reviews the development of oxygenators relative to both the essential design elements and the advances made in gas exchange technology. The primary challenge in the design of artificial lungs was the creation of adequate surface area for gas exchange. Direct gas-blood interfaces were first developed because of the lack of a highly permeable, solid material for an interface between the blood and gas phases. The first oxygenators developed were the
JULIE A. WEGNER
film oxygenators. These oxygenators used rotating cylinders to spread blood in a continuously renewed thin film with gas flowing over the top. The surface area generated was large; however, the actual volume of blood in contact with the gas-blood interface was small compared with the total volume of blood circulating in the device. In addition, blood streaming occurred within the thin blood film, which further decreased the red blood cell interaction with the exchange interface. Thus, large diffusion distances were generated, which limited the rate of both 02 and CO2 gas transfer, and decreased gas exchange efficiency. Because of the greater solubility of CO2 in blood, the rate of 02 transfer was limited to a greater degree than CO2, resulting in an uncoupling of 02 and CO2 gas exchange. A large prime volume was also required to obtain sufficient gas exchange for extracorporeal support. Although the direct gasblood interface eliminated a potential resistance to diffusion, the gas-blood interface was found to cause protein denaturation and blood damage. Even though the gas-blood interface caused progressive deterioration of the blood components, an acceptable alternative was not available that could generate the large gas exchange surface area required for extracorporeal support. The next generation of oxygenators, the bubble oxygenators, were designed to increase gas exchange efficiency by increasing the contact between the blood phase and the gas phase by dispersing gas bubbles of 100% 02 into venous blood. Gas exchange occurred in the resultant foam because of the intimate mixing of blood and oxygen. The improved gas exchange efficiency resulted in a significant reduction in the prime volume. Bubble size was found to influence 02 and CO2 gas exchange because of the differences in solubility between 02 and CO2. Small bubbles with a high surface-area-to-volume ratio favored 02 transfer, whereas larger bubbles favored CO2 transfer. The number of bubbles was determined by gas flow rate. Although gas exchange efficiency was significantly increased in bubble oxygenators, balanced 0 2 a n d CO 2 was difficult to achieve because of the dependence of 02 and CO2 gas exchange on both gas bubble size and gas-to-blood-flow ratios. Balanced 02 and CO2 transfer had to be compromised to achieve adequate gas exchange efficiency. Another consequence of the generation of gas bubbles in the blood was the requirement of a defoaming area and filtration unit within the oxygenator to remove gaseous emboli from the bloodstream before entering the arterial reservoir. The addition of defoaming surfaces into the oxygenator design increased the exposure of blood to foreign surfaces, without the benefit of significant gas exchange. Although the bubble oxygenator created a large surface area for gas exchange without the need of a thin blood path, the primary limitation to gas exchange was still the blood phase. The blood phase resistance could be decreased by increasing the gas flow (ie, increasing the number of bubbles). However, improvements in gas exchange efficiency were limited by the ability to remove the gaseous microemboli generated by the gas exchange process, and by the increase in blood damage attributable to increased exposure of blood to gaseous bubbles. The clinical significance of the impact of direct exposure of blood to a gas interface is controversial and may not be clinically significant in routine cases of cardiopulmonary bypass. However, because of the progressive nature of blood
OXYGENATOR ANATOMY
damage and protein denaturation at the gas-blood interface, the significance may increase in more critically ill patients and in situations requiring long-term extracorporeal circulatory support. The observation that the direct gas-blood interface caused significantblood damage led to th e development of oxygenators that interposed highly permeable materials between the gas phase and blood phase. The development of liquid fluorocarbons, in which both 02 and CO2 were highly soluble and were compatible with blood, led to the production of a liquid-liquid oxygenator. ~,2 The design concept of the liquid-liquid oxygenator was the same as the bubble oxygenator, except that the gaseous oxygen bubbles were encapsulated in fluorocarbon, which eliminated the direct gas-blood interface. Similar to the bubble oxygenator, increases in gas exchange efficiency were at the expense of increased generation of gaseous microemboli and an uncoupling of 02 and CO2 exchange. Although adequate gas exchange was achieved with a liquid-liquid oxygenator in a dog model of veno-veno bypass,2 the maintenance of oxygenator performance and the biocompatibility of the oxygenator were not studied. The advancements in polymer technology led to the production of thin, semipermeable membrane materials that could be used as a solid interface, or membrane, between the gas and blood phases in oxygenators. The initial challenge in membrane technology was to generate a consistent and reliable membrane with high permeability to both 02 and CO2. The first membrane solid polymer materials included polyethylene, polypropylene, and ethylcellulose. The permeability of CO2 in these membrane materials was only five times greater than the permeability of 02. Thus, CO 2 transfer was membrane-limited, whereas 02 transfer was blood-phase limited. A second challenge to membrane oxygenator development was to design an artificial circulation that would ensure adequate distribution of blood flow and maximize blood contact with the membrane. The first membrane oxygenators were flat-sheet, sandwich-type designs in which blood flowed between two membrane sheets with gas flow on the opposite side of the membrane. The inability to design appropriate blood flow pathways to achieve even blood distribution and narrow blood paths resulted in a significant resistance to 02 transfer in the blood phase. The combination of low membrane permeability to CO2 and a high resistance to 02 transfer in the blood phase required both a large surface area and prime volume to achieve adequate and balanced gas exchange. Thus, the prime volume required by first-generation membrane oxygenators was greater than that of bubble oxygenators. However, although protein adsorption and denaturation occurred at the membrane surface, in contrast to the direct gas-blood interface, the thin protein layer remained in place, resulting in less blood damage and blood activation. 3,4 Thus, membrane oxygenators were considered to be more biocompatible than bubble oxygenators. The next generation of membranes were silicone-based with increased permeability to CO2, thus improving the balance of 02 and CO2 transfer, gas exchange efficiency, and decreasing prime volume. However, CO2 transfer was still membranelimited because it was technically difficult to make thin membranes without inconsistencies of membrane properties.
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The typical membrane oxygenator designs were of the flat-sheet and spiral-coil types. The development of a microporous membrane, which contained small pores (0.5 to 1.0 ~m) and was highly permeable to both CO2 and 02, led to the next major advancement in membrane oxygenator technology. The existence of pores significantly increased the permeability of the membrane without having to decrease the thickness or reliability of the membrane. Furthermore, the elimination of the membrane resistance to CO2 exchange improved the balance between 02 and CO2 exchange, such that the control of 02 and CO2 could be accomplished independently by mechanisms similar to that used in the natural lung (ie, manipulation of sweep rate [ventilation], blood flow, and gas composition). Another result of the increase permeability of microporous membranes was a decrease in prime volume. Thus, gas exchange efficiency was significantly improved with the introduction of microporous membranes. The biocompatibilityproperties of the microporous membrane were similar to those of the solid membranes. Under normal operating conditions, the hydrophobic nature of the membrane surface did not allow blood to enter the pores, and thus did not allow a direct gas-blood interface at the pore. The most common microporous membrane oxygenator design is the hollow-fiber type in which the membrane is formed into fibers that are bundled or woven together. The fibers are 200 to 250 pm in diameter, 10 to 15 cm long, and the membrane thickness is 25 to 50 pm. The fiber configuration increases the amount of blood in contact with the gas exchange surface, and thus prime volume requirements. The blood flow through fiber bundle is either on the outside of the fibers (extraluminal flow [ELF]) or through the inside of the fibers (intraluminal flow [ILF]). The primary advantages of the ELF compared with the ILF design include a greater gas exchange surface, decreased prime volume requirement, and decreased resistance to blood flow. The ELF design is the most common type of hollow-fiber membrane oxygenator. Although the existence of micropores in the membrane significantly increased the gas exchange of membrane oxygenators, long-term use results in the progressive wetting of the surface, plasma leakage through the pores, and subsequent deterioration of membrane performance. The possible mechanisms of the progressive wetting to the membrane surface include water vapor condensation,5 ultrafiltration caused by changes in the transmembrane pressures, 6 or alteration of membrane surface tension caused by adsorption of phospholipids.7,8 The current membrane oxygenators mimic the natural lung by providing a highly permeable blood-membrane interface for gas exchange; however, the efficiency of the membrane oxygenator is two to three times less than the efficiency of the natural lung at rest, and about eight times less than the natural lung under conditions of maximal exercise. 9 The primary limiting factor to efficient gas exchange in membrane oxygenators appears to be blood phase resistance to both 02 and CO2 diffusion. A second factor that has limited the use of microporous membranes in situations of long-term extracorporeal support is the progressive decrease in gas exchange function. The mechanisms responsible for the deterioration of perfor-
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mance are unclear, but the biocompatibility of the device is thought to play a role. BLOOD PHASE RESISTANCE TO GAS EXCHANGE
With the development of highly permeable microporous membranes, the primary limitation to gas exchange is gas diffusion in the blood phase. The resistance to diffusion in the blood phase is a consequence of the anatomy of blood circulation in oxygenators. Typically, blood enters the oxygenating unit through a single conduit, which empties into an expanded chamber. From the inflow chamber, blood flow simultaneously divides into a large number of blood paths or functional channels fie. the "capillary" area). The dimensions of the functional channels are l0 to 15 cm in length and 150 to 250 #In in diameter. The blood paths simultaneously converge into the outflow chamber and then a single outflow conduit. The blood circulation in oxygenators is in contrast to that found in the natural lung, which is an extensive branching network of arteries, arterioles, capillaries, venules, and veins. The dimensions of the pulmonary capillaries are 0.5 to 1.0 mm in length and 3 to 7 pin in diameter. One consequence of the artificial oxygenator circulation is the blood damage associated with the conduits, surfaces, and blood flow patterns, which generate a spectrum of shear forces and uneven blood distribution. Computational fluid dynamics have been employed to guide the development of flow pattern geometry to minimize regions of high shear and stasis, and to promote uniform blood flow through the oxygenator unit.l° A second consequence of the current inability to mimic the pulmonary microcirculation is the size of the blood path required to permit a low resistance to blood flow through the oxygenatQr and uniform blood distribution. The width of the blood path in membrane oxygenators is approximately 25 times the width of the blood path found in the pulmonary capillaries. Thus. the blood actually participating in gas exchange (ie. the blood in contact with the membrane surface) is small compared with the total volume of blood in the path. In addition, boundary layers are generated in thick blood paths. Boundary layers exist in the native circulation, but not at the level of the capillaries where gas exchange takes place. A boundary layer is a cell-free layer of plasma adjacent to a solid surface, which by reasons of local fluid dynamics is not renewed at the same rate as blood in the core of the blood path. The lack of red blood cells in the boundary layer increases the diffusion distance to the membrane, thus increasing the resistance to diffusion. 9-1M3 The thickness of the boundary layer is a function of the width and length of the blood path. hematocrit, velocity of blood flow, and the shear stress or frictional forces at the membrane surface. The boundary layer begins to develop as blood flows into a conduit. As flow continues along a surface, the boundary layer increases m thickness until the boundary layers meet at the core of the blood path. ie, when flow is fully developed. Thus. gas exchange will be more efficient at the inlet than at the outlet of the oxygenator. The length of time required to achieve a given saturation of hemoglobin also increases with the thickness of the boundary layer. Thus, a longer exposure time is required, suggesting that gas transfer efficiency will decrease at high blood flows. Possible solutions for decreasing the blood phase limitation
of gas transfer would be to increase the surface-area-to-volume ratio and decrease the thickness of the boundary layer. Decreasing the width of the blood path would increase the surface-areato-volume ratio. A decrease in the diameter of the flow path would also decrease the maximum thickness of the boundary layer. However, such a reduction in blood path width would increase the resistance to blood flow and pressure drop across the oxygenator, and thus require a greater pressure for blood flow. The alterations in flow dynamics caused by a reduction in blood path diameter also might compromise blood distribution to all of the fibers within the oxygenator. Thus, the inflow and outflow ports would have to be designed to maximize blood distribution and minimize shear stress and stasis. A reduction in blood flow path would als0 increase the contact of blood with foreign surfaces, thus increasing hemolysis and blood activation. 14 Thus, an increase in the amount of blood participating in gas exchange could potentially compromise the hemocompatibility and biocompatibility of the oxygenator. Another method to decrease the boundary layer effect is to modify the laminar flow characteristics of blood flow by increasing the shear rate and turbulence of the blood path by the introduction of secondary flOWS.9'11'15'16 Secondary flow is any type of fluid motion, steady or periodic, in which the fluid is moving in a direction different from that of the primary flowJ 2 In other words, the introduction of secondary flows into the flow path results in greater blood mixing, which increases the number Of cells in the boundary layer. Incorporation of secondary flows in the oxygenator blood flow patterns has been shown to reduce the surface area required for gas exchange by 30% to 50%.1719 The introduction of secondary flows can either be passive or active. Passive secondary flows do not require the addition of energy to the blood path. Methods of achieving passive secondary flows include putting obstacles in the blood path, membrane indulations, membrane texturing, helical flow systems, or alterations in fiber geometry relative to blood flow. 17 Active secondary flows are achieved by adding energy to the blood as it flows through the oxygenator. Examples of active secondary flows include an enclosed rotating disc, inner cylinder rotating axial flow, and pulsed flow vortex shedding. 17 Active secondary flows can also be achieved by the conduct of perfusion, such as the use of pulsatile blood flow. The gas exchange efficiency of membrane oxygenators was increased by the use of pulsatile blood flow. 9,20,21 However, in one study pulsatile flow was found to have no influence on the gas exchange efficiency of a membrane oxygenator, 22 which suggests that the effect of pulsatile flow may be dependent on the blood fl0w characteristics of the membrane oxygenator used in the extracorporeal circuit. Th e incorporation of secondary flows introduces turbulence into the blood flow path. Compared with laminar blood flow, turbulent blood flow requires greater pressure to generate a given blood flow. Thus, a greater pressure force is generated. In addition, secondary flows, or turbulence, generate greater shear rates. Shear rates of 1,000 to 2,000/s were found to be required to influence gas transfer efficiency of membrane oxygenators. 9J1,15,16 The shear rates generated in membrane oxygenators ranged from 480 to 2,100/s.14 Although increases in shear rate can potentially augment gas transfer, an increase in the
OXYGE NATOR ANATOMY
shear-induced effects on the blood als0 occurs. The magnitude of hemolysis was found to increase linearly with increases in shear rate. 14 In addition, the adhesion patterns of platelets and neutrophils were also found to be dependent on shear rate. 23 Thus, both the gas exchange efficiencY and the biocompatibility of blood are affected by the shear rates generated in membrane oxygenators. With the current membrane materials available, the improvement in gas exchange efficiency would have to be at the expense of a reduction in biocompatibility. However, shear-induced effects on the blood introduced by secondary flows may be balanced by the decrease in surface area required for gas transfer because of greater efficiency. Another development that may decrease the blood phase resistance to gas transfer, without a change in oxygenator design, is the use of perfluorocarbons (PFC) as a blood substitute rather than crystalloid or colloid solutions.24 The solubility of 02 is increased in PFCs compared with blood or crystalloid/colloid solutions, which decreases the resistance to diffusion and the requirement of red blood cells in contact with the membrane surface. Animal studies have shown adequate gas transfer ability and stability of PFCs during extracorporeal circulation.25,26 Studies have also shown that efficiency of O2 transfer was greatest for hollow-fiber membrane oxygenators compared with solid membrane and bubble oxygenatorsY ,27 Thus, the use of blood substitutes during extracorporeal circulation may increase the efficiency of gas transfer without the introduction of shear stresses or increases in surface area. Issues that still need to be addressed regarding blood substitutes are the long-term performance characteristics of membrane oxygen~ ators, exposed PFC emulsions, and the biocompatibility of PFCs. OXYGENATOR BIOCOMPATIBILITY
Biocompatibility and hemocompatibility are complex problems that encompass shear stresses, surface area-to-volume ratio, geometry of blood flow, type and composition of the foreign surface, duration of exposure, and patient status. Furthermore, the biocompatibility characteristics of a given material may be different under different flow conditions, or in different oxygenator designs. A systemic inflammatory response is induced in all patients exposed to extracorporeal circulation, part because of the activation of blood components that come in contact with foreign surfaces. However, the severitY of the systemic inflammatory response is highly variable, and only in a minority of patients does the response result in clinically significant adverse events.2S,29The complexity of biocompatibility implies that no single alteration in oxygenator design will effectively improve the biocompatibility of the entire system. The first interaction of blood with an artificial surface results in the instantaneous adsorption of protein to the surface with subsequent protein denaturation or blood component adhesion. 14 In oxygenators without a solid surface, such as bubble oxygenators, the denatured proteins are continually released back into the circulation, where they can exhibit unfavorable biological properties, such as expression of vasoactivity, alteration of coagulation factors, and initiation of red blood cell agglutination with subsequent alterations in microcirculatory blood flow. 3,3° In contrast, the adsorption of protein to a solid surface, such as the membrane interface, results in a thin protein
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layer that remains in place and subsequently alters the reactivity of the surface such that further protein denaturation and blood damage are moderated. 3,4,21The type of protein adsorbed on the membrane surface also influences the biocompatibilitY of a device. Membrane materials that preferentially adsorb albumin have been shown to have a reduced thrombogenic tendency compared with membranes with low albumin adsorption. 17,31 Precoating the extracorporeal circuit with fresh frozen plasma has been shown to decrease complement and platelet activation because of the composition of the protein layer deposited and receptor sites available on the membrane surface before initiation of cardiopulmonary bypass. ~2 Coating the surface of membranes with polymeric phospholipids, to mimic the lipid surface of the endothelium, may be another method to alter the amount and type of protein adsorbed, and to improve biocompatibility.33 Because the biocompatibility of a device depends in part on the degree of blood activation when blood comes in contact with a foreign surface, attenuation of contact activation should improve biocompatibility. One approach to increasing the biocompatibility of a device is to make the surface more endothelial-like. On e primary function of the endothelium is to maintain blood in a fluid and flowing state. The surface of the vascular endothelium is antithromb0genic because of the presence of several proteins, such as heparan sulfate. The heparan sulfate protein functions similarly to heparin sulfate to prevent thrombin formation. Thus, one approach to increasing biocompatibility is to coat the membrane and other extracorporeal surfaces with heparin, either by a covalently or ionically bonding processes. The heparin coating creates an antit!arombotic surface that is both flow-dependent and antithrombin IIIdependent. Most studies Comparing heparin-coated circuits with non-heparin-coated circuits have shown a decrease in complement activation,33-35 a decrease in platelet activation,35 and a decrease in neutrophil activation36 with heparin coating. Overall, the conclusions from these studies are that heparin coating improves biocompatibility. However, the significance of the improvement Of biocompatibility with respect to clinical outcome is still uncertain. In addition to the exPression of antithrombogenic proteins, endothelial cells also express surface proteins, or release substances that influence the activation of platelets, comPlement, and neutrophils. Nitric oxide, a potent vasodilator and platelet inhibitor, is released by endothelial ceils. The addition of nitric oxide to the sweep gas has been shown to significantly reduce platelet activation during cardiopulmonary bypass. 3v Other attempts to mimic the endothelial cell environment include the use of DNA technology to incorporate proteins such as complement receptor 1, a cell surface inhibitor of complement, into membrane surfaces to prevent local complement activation.3s Other future membrane Coatings might include monoclonal antibodies against the adhesion proteins exPressed on platelets and neutrophils.
CONCLUSIONS
The gas exchange efficiency of current membrane oxygenators appears to be sufficient for routine cardiopulmonary bypass. However, further improvements will be required so that
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all patients, including children and small adults, can benefit from the small prime volumes associated with highly efficient oxygenators. Advances in gas exchange efficiency will require a thorough understanding of the factors limiting gas diffusion in the blood phase. An appreciation of how improvements in gas exchange efficiency will impact on the biocompatibility of the device will also be required. The research and development of the intravascular oxygenator may help lead the way to advancements in extracorporeal oxygenator designs. The ability to maintain gas exchange performance for long duration is the primary limitation of current hollow-fiber membrane oxygenators. Knowledge of the contributing factor(s) causing deterioration of membrane performance is lacking. The cause is probably multifactorial and associated with the magnitude of the inflammatory response induced by exposure to extracorporeal circulation. In addition, the deterioration of patient status because of long-term exposure to nonpulsatile blood flow and anticoagulation must be taken into account. Few
studies have examined the nature of the inflammatory response during long-term support. Thus, a better understanding of the components of the inflammatory response and their interactions will be required before improvements in long-term performance can be made. The magnitude of the inflammatory response induced by exposure to extracorporeal circulation is related to the biocompatibility of the circuit or device. Even though the inflammatory response is induced in all patients, the magnitude of the response varies widely. Thus, a clear picture of the impact of biocompatibility on i~atient outcome is lacking and has slowed down advancements in this area. The importance of biocompatibility will be greatest in critically ill patients as well as patients younger than 2 years and older than 80 years. Again, a better understanding of the interaction of patient status on the magnitude of the inflammatory response will result in further improvements in the biocompatibility of membrane oxygenators.
REFERENCES
1. Howlett S, Dundas D, Sabiston DC: Fluid fluorocarbon as oxygenator in experimental extracorporeal circulation. Arch Surg 91:643-645, 1965 2. Pitzele S, Sze S, Dobell ARC: Gas exchange across a liquid-liquid interface. Trans Am Soc Artif Intern Organs 18:375-383, 1972 3. Peirce EC: Is the blood-gas interface of clinical importance? Ann Thorac Surg 17:526-529, 1974 4. Tamari Y, Aledort L, Puszkin E, et al: Functional changes in platelets during extracorporeal circulation. Ann Thorac Surg 19:639647, 1975 5. Palder SB, Shaheen KW, Whittlesey GC, et al: Prolonged extracorporeal membrane oxygenation in sheep with a hollow-fiber oxygenator and a centrifugal pump. Trans Am Soc Artif Intern Organs. 34:820-822, 1988 6. Tamari Y, Tortolani AJ, Maquine M, et al: The effect of high pressure on microporous membrane oxygenator failure. Artif Organs 15:15-22, 1991 7. Montoya JR Shanley CJ, Merz SI; et al: Plasma leakage through microporous membranes: Role of phosph01ipids. ASAIO J 38:M399405, 1992 8. Butler BO, Kurnsz M, Conti VR: Postbypass surface tension and its role in bubble filtration. Perfusion 1:187-191, 1986 9. SpeathEE: Blood oxygenation in extracorporealdevices: Theoretical considerations. CRC Crit Rev Bioeng 1:383-417, 1973 10. Goodin MS, Thor EJ, Haworth WS: Use of computational fluid dynamics in the design of the Avecor Affinity oxygenator. Perfusion 9i217-222, 1994 11. Colton CK: Fundamentals of gas transport in blood, in Zapol WM, Qvist J (eds): Artificial Lungs for Acute Respiratory Failure: Theory and Practice. New York, NY, Academic, 1976, pp 3-42 12. Galletti PM, Colton CK: Artificial lungs and blood-gas exchange devices, in Bronzino JD (ed): The Biomedical Engineering Handbook. Boca Raton, FL, CRC, 1995, pp 1879-1897 13. Marx TI, Snyder WE, St John AD, et al: Diffusion of oxygen into a film of whole blood. J Appl Physiol 15:1123-1129, i960 14. Solen KA, Whiffen JD, Lightfoot EN: The effect of shear, specific surface, and air interface on the development of blood emboli and hemolysis. J Biomed Mater Res 12:381-399, 1978 15. Keller KH: Effect of fluid shear on mass transport in flowing blood. Fed Proc 30:1591-1599, 1971 16. Ebert RC, Dengle SK, Curtis RM: Mathematical and experimen-
tal methods for design and evaluation of membrane oxygenators. Artif Organs 2:19-34, 1978 17. Galletti PM: Cardiopulmonary bypass: A historical perspective. Artif Organs 17:675-686, 1993 18. Drinker PA, Lehr JL: Engineering aspects of ECMO technology. Artif Organs 2:6-1 !, 1978 19. Gaylor JDS, Hickey S, Bell G, et al: Membrane oxygenators: Influence of design on performance. Perfusion 9:173-180, 1994 20. Bodell BR, Head JM, Head LR, et al: A capillary membrane oxygenator. J Thorac Cardiovasc Surg 46:639-650, 1963 21. Peirce EC: Tile theory and function of the membrane lung. Mount Sinai J Med 40:119-134, 1973 22. Gourlay T, Taylor KM: Pulsatile flow and membrane oxygenators. Perfusion 9:189-196, 1994 23. Jessen M, Kulkarni P, Constantinescu A, et al: Influence of blood flow and surface characteristics on cell adhesion on membrane oxygenator circuit surfaces. ASAIO J 41:63A, 1995 (suppl) 24. Ferguson ER, Clymer JJ, Spruell RD, et al: Perfluorocarb0n oxygen transport: A comparative study of four oxygenator designs. ASAIO J 40:M649-M653, 1994 25. R°us° JA' Engelman RM' Anisim°wicz L' et al: A c°mparis°n °f blood and Fluosol-DA for cardiopulmonary bypass. J Cardiovasc Surg 26:447-453, 1985 26. Usuba A, Motoki R, Ogata Y, et al: Effect and safety of liposome-encapsulated hemoglobin Net Red Cells (NRCs) as a perfusate for total c~irdiopulmonary bypass. Art Cells Blood Subs Immob Biotech 23:337-346, i995 27. Holman WL, McGiffin DC, Vicente WVA, et al: Use of current generation perfluorocarbon emulsions in cardiac surgery. Art Cells Blood Subs Immob Biotech 22:979-990, 1994 28. Taylor KM: The systemic inflammatory response syndrome after cardiac operations. Ann Thorac Surg 61:1607-1608, 1996 29. Cremer J, Martin M, Redl H, et al: Systemic inflammatory response syndrome after cardiac operations. Ann Thorac Surg 61:17141720, 1996 30. Lee WH, Hairston P: Structural effects on blood proteins at the gas-blood interface. Fed Proc 30:1615-1620; 1971 31. Knight PM, Lyman DJ: Evaluation of the thrombogenicity of selected microporous oxygenator membranes. Artif Organs 9:28-36, 1985 32. Collett JR, Plunkett PF: Effect of serum precoat on complement
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activation in membrane oxygenators. Proc Am Soc Extra-Corporeal Tech, 1990, pp 57-61 33. Gu YJ, van Oeveren W, Akkerman C, et al: Heparin-coated circuits reduce the inflammatory response to cardiopulmonary bypass. Ann Thorac Surg 55:917-922, 1993 34. Videm V, Svennevig JL, Fosse E, et al: Reduced complement activation with heparin-coated oxygenator and tubings in coronary bypass operations. J Thorac Cardiovasc Surg t03:806-813, 1992 35. Fukutomi M, Kobayashi S, Niwaya K, et al: Changes in platelet, granulocyte, and complement activation during cardiopulmonary
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bypass using heparin-coated equipment. Artif Organs 20:767-776, 1996 36. Thelin S, Bagge L, Hultman J, et al: Heparin-coated cardiopulmonary bypass circuits reduce blood cell trauma: Experiments in the pig. Eur J Cardiothorac Surg 5:486-491, 1991 37. Sly MK, Prager MD, Eberhart RC, et al: Inhibition of surfaceinduced platelet activation by nitric oxide. ASAIO J 41:M394-M398, 1995 38. Hornick P, George A: Blood contact activation: Pathophysiological effects and therapeutic approaches. Perfusion 11:3-19, 1996
review is to present the past and current developments of the oxygenator designs in terms of the structural and functional features of the natural lung as well as the limitations in the ability to mimic the features of the lung because of the lack of appropriate technology.
HE NATURAL lung is the organ responsible for oxygen and carbon dioxide exchange between the blood and the outside environment. The challenge of replacing the function of the natural lung with an artificial lung, or oxygenator, requires a thorough understanding of the structural and functional features that make the natural lung an efficient gas exchange organ. The oxygenator must be able to provide efficient and balanced gas exchange under extreme conditions of hemodilution and temperature with minimal blood damage and blood activation. The purpose of this article is to present the past and current developments of the oxygenator designs in terms of the structural and functional features of the natural lung as well as the limitations in the ability to mimic the features of the lung because of the lack of appropriate technology. Brief reviews of lung anatomy and function and the development of oxygenators are presented. In addition, the factors that limit the gas exchange efficiency and long-term utilization of the present-day membrane oxygenators are presented, as well as possible solutions to improvements in oxygenator design.
membrane, interstitial space, capillary basement membrane, and capillary endothelium. The average thickness of the alveolarcapillary interface is 0.5 lain and is highly permeable to both 02 and CO2. The large surface area and low resistance to gas diffusion of the alveolar-capillary interface are the two main factors responsible for the high efficiency of gas exchange in the lung. Gas exchange at the alveolar-capillary interface occurs through passive diffusion, as described by Ficks Law: Vgas (Volume/Unit Time) = (A/T) × D × (P1 - P2), where A equals surface area, T equals interface thickness, D equals diffusion coefficient, and P1 - P2 is the pressure gradient across the interface (ie, PAO2 - Pc o , where PAO2 is the alveolar oxygen tension and Pc o is the cat~illary oxygen tension). The diffusion coefficient is d~pendent on the solubility of the gas, temperature, and permeability of the interface. Differences in the diffusion coefficient, pressure gradient, and mechanism of transport in the blood between of 02 and CO2 suggest that different factors will limit 02 and CO2 exchange at a gas exchange interface, such as the alveolar-capillary membrane. Based on the components of the Fick equation, it can be inferred that the exchange of a given gas (Vgas) can be increased by increasing the surface area available for gas exchange, decreasing the thickness of the interface, increasing the diffusion coefficient, or increasing the pressure gradient. Under normal conditions, the interface thickness (T) and diffusion coefficient (D) are constant, and thus do not contribute to significant changes in gas exchange capacity. At rest, the entire alveolar-capillary surface area is not available for gas exchange because of mismatching in ventilation and perfusion. In other words, capillary blood flow is not always in contact with ventilated alveoli and vice versa. However, the gas exchange surface area can be rapidly expanded by alterations in the distribution of gas flow and blood flow in the lung. Thus, the natural lung has the ability to regulate gas exchange by adjusting the surface area available for gas exchange. Gas exchange can also be influenced by the individual pressure gradients across the interface. The gradients are dependent on both the O2 and CO2 content of the blood in the capillaries (ie, venous admixture) and the composition of the gas in the alveoli. Under conditions of room air and normal
T
STRUCTURE AND FUNCTION OF THE NATURAL LUNG
The lung is an efficient gas exchange organ, in part because of the large surface area generated by the branching network of the air conduction and blood-carrying conduits. The airways, or bronchi, sequentially branch into bronchioles, which in turn branch into alveolar ducts and finally individual alveoli. The spherical design of the alveoli creates a large surface area for gas exchange. Gas flow into and out of the alveoli is regulated at the level of the bronchioles, which are lined with smooth muscle. A similar branching pattern exists for the pulmonary blood vessels, which branch from the pulmonary arteries to the arterioles and finally into the capillaries that vascularize the alveoli. The pulmonary capillaries are 0.5 to 1.0 mm in length and 3 to 7 bun in diameter. Because at least 95% of oxygen (02) is transported in the blood as oxyhemoglobin (HbO2), the small diameter of the pulmonary capillaries ensures that the red blood cells are in constant contact with the endothelial cell surface, and thus can participate in gas exchange (ie, a high surfacearea- to-volume ratio). Regulation of blood distribution in the pulmonary vasculature occurs primarily at the arteriolar level. The convergence of the airway and blood vessel networks creates the gas exchange interface, the alveolar-capillary membrane. There are an estimated 300 million alveoli in the lungs, and the total alveolar surface area of the lung ranges from 60 to 80 m2. The alveolar-capillary interface consists of several layers, including the alveolar epithelium, alveolar basement
Copyright© 1997by W.B. Saunders Company KEY WORDS: lung, gas exchange, oxygenator, biocompatibility, membrane oxygenator, blood flow dynamics
From the University Medical Center, Tucson, AZ Address reprint requests to Julie A. Wegner, PhD, CCP, University Medical Center, 1501 N Campbell Ave, Room 4402, Tucson, AZ 85724. Copyright © 1997 by W.B. Saunders Company 1053-0770/97/1103-000353.00/0
Journal of Cardiothoracic and Vascular Anesthesia, Vol 11, No 3 (May), 1997: pp 275-281
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ventilation, the pressure gradient for 0 2 is approximately 40 mmHg, and approximately 6 mmHg for CO2. The gradients for 02 and CO2 can be modified by alterations in the metabolic rate of the individual as well as the composition of the inspired gas (O2) and the ventilation rate (CO2). Another important feature of the natural lung is that 02 exchange and CO2 exchange are metabolically balanced. Although 02 and CO2 exchange are metabolically coupled, under artificial conditions such as artificial ventilation, they can be modulated independently. A final factor that can influence the volume of gas exchange across the alveolar-capillary interface is the exposure time of the blood at the interface. At rest the contact time is approximately 0.75 seconds; however, because of the efficiency of gas exchange in the lung, gas transfer normally takes place within 0.25 seconds. Thus, the contact time is normally not a limiting factor in gas exchange in the healthy lung, even under conditions of maximal exercise. Another important feature of the natural lung is the continuous gas exchange performance over the lifetime of the individual. Gas exchange performance is maintained at both the alveolar and capillary level. The alveolar wall is composed of a layer of epithelial cells and the epithelial basement membrane. Cells within the alveolar epithelial layer secrete a phospholipid substance called surfactant, which prevents the collapse of alveoli during ventilation. Also found within the alveolar wall are macrophages. Alveolar macrophages are highly phagocytic cells that remove debris from the alveolar spaces. Thus, both surfactant and alveolar macrophages maintain the surface of the alveolar interface. The pulmonary capillaries are lined with the endothelial cells. A primary function of the vascular endothelium is to keep the blood in a fluid state through regulation of hemostasis and vascular tone. Thus, the vascular endothelium maintains blood flow to the capillaries and ensures contact of blood at the alveolar interface. The vascular endothelium also serves as a barrier between the blood and subendothelial surfaces. Exposure of blood to the subendothelial surfaces results in activation of hemostasis and inflammation. Thus, the vascular endothelial layer plays an important role in ensuring biocompatibility within the vasculature. DESIGN AND DEVELOPMENT OF THE ARTIFICIAL LUNG, THE OXYGENATOR
The efficiency of gas exchange in the natural lung is mainly attributable to the large surface area generated by the airway and circulatory networks and the low resistance to diffusion. These same features are essential to the design of an efficient artificial lung. Other necessary features of an ideal oxygenator include minimal trauma to blood, thromboresistance, minimal reaction with blood components, minimal generation of gaseous microemboli, ability to maintain performance over long periods, low prime volume, consistent physical properties, reliability, ease of use, and low cost. This section briefly reviews the development of oxygenators relative to both the essential design elements and the advances made in gas exchange technology. The primary challenge in the design of artificial lungs was the creation of adequate surface area for gas exchange. Direct gas-blood interfaces were first developed because of the lack of a highly permeable, solid material for an interface between the blood and gas phases. The first oxygenators developed were the
JULIE A. WEGNER
film oxygenators. These oxygenators used rotating cylinders to spread blood in a continuously renewed thin film with gas flowing over the top. The surface area generated was large; however, the actual volume of blood in contact with the gas-blood interface was small compared with the total volume of blood circulating in the device. In addition, blood streaming occurred within the thin blood film, which further decreased the red blood cell interaction with the exchange interface. Thus, large diffusion distances were generated, which limited the rate of both 02 and CO2 gas transfer, and decreased gas exchange efficiency. Because of the greater solubility of CO2 in blood, the rate of 02 transfer was limited to a greater degree than CO2, resulting in an uncoupling of 02 and CO2 gas exchange. A large prime volume was also required to obtain sufficient gas exchange for extracorporeal support. Although the direct gasblood interface eliminated a potential resistance to diffusion, the gas-blood interface was found to cause protein denaturation and blood damage. Even though the gas-blood interface caused progressive deterioration of the blood components, an acceptable alternative was not available that could generate the large gas exchange surface area required for extracorporeal support. The next generation of oxygenators, the bubble oxygenators, were designed to increase gas exchange efficiency by increasing the contact between the blood phase and the gas phase by dispersing gas bubbles of 100% 02 into venous blood. Gas exchange occurred in the resultant foam because of the intimate mixing of blood and oxygen. The improved gas exchange efficiency resulted in a significant reduction in the prime volume. Bubble size was found to influence 02 and CO2 gas exchange because of the differences in solubility between 02 and CO2. Small bubbles with a high surface-area-to-volume ratio favored 02 transfer, whereas larger bubbles favored CO2 transfer. The number of bubbles was determined by gas flow rate. Although gas exchange efficiency was significantly increased in bubble oxygenators, balanced 0 2 a n d CO 2 was difficult to achieve because of the dependence of 02 and CO2 gas exchange on both gas bubble size and gas-to-blood-flow ratios. Balanced 02 and CO2 transfer had to be compromised to achieve adequate gas exchange efficiency. Another consequence of the generation of gas bubbles in the blood was the requirement of a defoaming area and filtration unit within the oxygenator to remove gaseous emboli from the bloodstream before entering the arterial reservoir. The addition of defoaming surfaces into the oxygenator design increased the exposure of blood to foreign surfaces, without the benefit of significant gas exchange. Although the bubble oxygenator created a large surface area for gas exchange without the need of a thin blood path, the primary limitation to gas exchange was still the blood phase. The blood phase resistance could be decreased by increasing the gas flow (ie, increasing the number of bubbles). However, improvements in gas exchange efficiency were limited by the ability to remove the gaseous microemboli generated by the gas exchange process, and by the increase in blood damage attributable to increased exposure of blood to gaseous bubbles. The clinical significance of the impact of direct exposure of blood to a gas interface is controversial and may not be clinically significant in routine cases of cardiopulmonary bypass. However, because of the progressive nature of blood
OXYGENATOR ANATOMY
damage and protein denaturation at the gas-blood interface, the significance may increase in more critically ill patients and in situations requiring long-term extracorporeal circulatory support. The observation that the direct gas-blood interface caused significantblood damage led to th e development of oxygenators that interposed highly permeable materials between the gas phase and blood phase. The development of liquid fluorocarbons, in which both 02 and CO2 were highly soluble and were compatible with blood, led to the production of a liquid-liquid oxygenator. ~,2 The design concept of the liquid-liquid oxygenator was the same as the bubble oxygenator, except that the gaseous oxygen bubbles were encapsulated in fluorocarbon, which eliminated the direct gas-blood interface. Similar to the bubble oxygenator, increases in gas exchange efficiency were at the expense of increased generation of gaseous microemboli and an uncoupling of 02 and CO2 exchange. Although adequate gas exchange was achieved with a liquid-liquid oxygenator in a dog model of veno-veno bypass,2 the maintenance of oxygenator performance and the biocompatibility of the oxygenator were not studied. The advancements in polymer technology led to the production of thin, semipermeable membrane materials that could be used as a solid interface, or membrane, between the gas and blood phases in oxygenators. The initial challenge in membrane technology was to generate a consistent and reliable membrane with high permeability to both 02 and CO2. The first membrane solid polymer materials included polyethylene, polypropylene, and ethylcellulose. The permeability of CO2 in these membrane materials was only five times greater than the permeability of 02. Thus, CO 2 transfer was membrane-limited, whereas 02 transfer was blood-phase limited. A second challenge to membrane oxygenator development was to design an artificial circulation that would ensure adequate distribution of blood flow and maximize blood contact with the membrane. The first membrane oxygenators were flat-sheet, sandwich-type designs in which blood flowed between two membrane sheets with gas flow on the opposite side of the membrane. The inability to design appropriate blood flow pathways to achieve even blood distribution and narrow blood paths resulted in a significant resistance to 02 transfer in the blood phase. The combination of low membrane permeability to CO2 and a high resistance to 02 transfer in the blood phase required both a large surface area and prime volume to achieve adequate and balanced gas exchange. Thus, the prime volume required by first-generation membrane oxygenators was greater than that of bubble oxygenators. However, although protein adsorption and denaturation occurred at the membrane surface, in contrast to the direct gas-blood interface, the thin protein layer remained in place, resulting in less blood damage and blood activation. 3,4 Thus, membrane oxygenators were considered to be more biocompatible than bubble oxygenators. The next generation of membranes were silicone-based with increased permeability to CO2, thus improving the balance of 02 and CO2 transfer, gas exchange efficiency, and decreasing prime volume. However, CO2 transfer was still membranelimited because it was technically difficult to make thin membranes without inconsistencies of membrane properties.
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The typical membrane oxygenator designs were of the flat-sheet and spiral-coil types. The development of a microporous membrane, which contained small pores (0.5 to 1.0 ~m) and was highly permeable to both CO2 and 02, led to the next major advancement in membrane oxygenator technology. The existence of pores significantly increased the permeability of the membrane without having to decrease the thickness or reliability of the membrane. Furthermore, the elimination of the membrane resistance to CO2 exchange improved the balance between 02 and CO2 exchange, such that the control of 02 and CO2 could be accomplished independently by mechanisms similar to that used in the natural lung (ie, manipulation of sweep rate [ventilation], blood flow, and gas composition). Another result of the increase permeability of microporous membranes was a decrease in prime volume. Thus, gas exchange efficiency was significantly improved with the introduction of microporous membranes. The biocompatibilityproperties of the microporous membrane were similar to those of the solid membranes. Under normal operating conditions, the hydrophobic nature of the membrane surface did not allow blood to enter the pores, and thus did not allow a direct gas-blood interface at the pore. The most common microporous membrane oxygenator design is the hollow-fiber type in which the membrane is formed into fibers that are bundled or woven together. The fibers are 200 to 250 pm in diameter, 10 to 15 cm long, and the membrane thickness is 25 to 50 pm. The fiber configuration increases the amount of blood in contact with the gas exchange surface, and thus prime volume requirements. The blood flow through fiber bundle is either on the outside of the fibers (extraluminal flow [ELF]) or through the inside of the fibers (intraluminal flow [ILF]). The primary advantages of the ELF compared with the ILF design include a greater gas exchange surface, decreased prime volume requirement, and decreased resistance to blood flow. The ELF design is the most common type of hollow-fiber membrane oxygenator. Although the existence of micropores in the membrane significantly increased the gas exchange of membrane oxygenators, long-term use results in the progressive wetting of the surface, plasma leakage through the pores, and subsequent deterioration of membrane performance. The possible mechanisms of the progressive wetting to the membrane surface include water vapor condensation,5 ultrafiltration caused by changes in the transmembrane pressures, 6 or alteration of membrane surface tension caused by adsorption of phospholipids.7,8 The current membrane oxygenators mimic the natural lung by providing a highly permeable blood-membrane interface for gas exchange; however, the efficiency of the membrane oxygenator is two to three times less than the efficiency of the natural lung at rest, and about eight times less than the natural lung under conditions of maximal exercise. 9 The primary limiting factor to efficient gas exchange in membrane oxygenators appears to be blood phase resistance to both 02 and CO2 diffusion. A second factor that has limited the use of microporous membranes in situations of long-term extracorporeal support is the progressive decrease in gas exchange function. The mechanisms responsible for the deterioration of perfor-
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mance are unclear, but the biocompatibility of the device is thought to play a role. BLOOD PHASE RESISTANCE TO GAS EXCHANGE
With the development of highly permeable microporous membranes, the primary limitation to gas exchange is gas diffusion in the blood phase. The resistance to diffusion in the blood phase is a consequence of the anatomy of blood circulation in oxygenators. Typically, blood enters the oxygenating unit through a single conduit, which empties into an expanded chamber. From the inflow chamber, blood flow simultaneously divides into a large number of blood paths or functional channels fie. the "capillary" area). The dimensions of the functional channels are l0 to 15 cm in length and 150 to 250 #In in diameter. The blood paths simultaneously converge into the outflow chamber and then a single outflow conduit. The blood circulation in oxygenators is in contrast to that found in the natural lung, which is an extensive branching network of arteries, arterioles, capillaries, venules, and veins. The dimensions of the pulmonary capillaries are 0.5 to 1.0 mm in length and 3 to 7 pin in diameter. One consequence of the artificial oxygenator circulation is the blood damage associated with the conduits, surfaces, and blood flow patterns, which generate a spectrum of shear forces and uneven blood distribution. Computational fluid dynamics have been employed to guide the development of flow pattern geometry to minimize regions of high shear and stasis, and to promote uniform blood flow through the oxygenator unit.l° A second consequence of the current inability to mimic the pulmonary microcirculation is the size of the blood path required to permit a low resistance to blood flow through the oxygenatQr and uniform blood distribution. The width of the blood path in membrane oxygenators is approximately 25 times the width of the blood path found in the pulmonary capillaries. Thus. the blood actually participating in gas exchange (ie. the blood in contact with the membrane surface) is small compared with the total volume of blood in the path. In addition, boundary layers are generated in thick blood paths. Boundary layers exist in the native circulation, but not at the level of the capillaries where gas exchange takes place. A boundary layer is a cell-free layer of plasma adjacent to a solid surface, which by reasons of local fluid dynamics is not renewed at the same rate as blood in the core of the blood path. The lack of red blood cells in the boundary layer increases the diffusion distance to the membrane, thus increasing the resistance to diffusion. 9-1M3 The thickness of the boundary layer is a function of the width and length of the blood path. hematocrit, velocity of blood flow, and the shear stress or frictional forces at the membrane surface. The boundary layer begins to develop as blood flows into a conduit. As flow continues along a surface, the boundary layer increases m thickness until the boundary layers meet at the core of the blood path. ie, when flow is fully developed. Thus. gas exchange will be more efficient at the inlet than at the outlet of the oxygenator. The length of time required to achieve a given saturation of hemoglobin also increases with the thickness of the boundary layer. Thus, a longer exposure time is required, suggesting that gas transfer efficiency will decrease at high blood flows. Possible solutions for decreasing the blood phase limitation
of gas transfer would be to increase the surface-area-to-volume ratio and decrease the thickness of the boundary layer. Decreasing the width of the blood path would increase the surface-areato-volume ratio. A decrease in the diameter of the flow path would also decrease the maximum thickness of the boundary layer. However, such a reduction in blood path width would increase the resistance to blood flow and pressure drop across the oxygenator, and thus require a greater pressure for blood flow. The alterations in flow dynamics caused by a reduction in blood path diameter also might compromise blood distribution to all of the fibers within the oxygenator. Thus, the inflow and outflow ports would have to be designed to maximize blood distribution and minimize shear stress and stasis. A reduction in blood flow path would als0 increase the contact of blood with foreign surfaces, thus increasing hemolysis and blood activation. 14 Thus, an increase in the amount of blood participating in gas exchange could potentially compromise the hemocompatibility and biocompatibility of the oxygenator. Another method to decrease the boundary layer effect is to modify the laminar flow characteristics of blood flow by increasing the shear rate and turbulence of the blood path by the introduction of secondary flOWS.9'11'15'16 Secondary flow is any type of fluid motion, steady or periodic, in which the fluid is moving in a direction different from that of the primary flowJ 2 In other words, the introduction of secondary flows into the flow path results in greater blood mixing, which increases the number Of cells in the boundary layer. Incorporation of secondary flows in the oxygenator blood flow patterns has been shown to reduce the surface area required for gas exchange by 30% to 50%.1719 The introduction of secondary flows can either be passive or active. Passive secondary flows do not require the addition of energy to the blood path. Methods of achieving passive secondary flows include putting obstacles in the blood path, membrane indulations, membrane texturing, helical flow systems, or alterations in fiber geometry relative to blood flow. 17 Active secondary flows are achieved by adding energy to the blood as it flows through the oxygenator. Examples of active secondary flows include an enclosed rotating disc, inner cylinder rotating axial flow, and pulsed flow vortex shedding. 17 Active secondary flows can also be achieved by the conduct of perfusion, such as the use of pulsatile blood flow. The gas exchange efficiency of membrane oxygenators was increased by the use of pulsatile blood flow. 9,20,21 However, in one study pulsatile flow was found to have no influence on the gas exchange efficiency of a membrane oxygenator, 22 which suggests that the effect of pulsatile flow may be dependent on the blood fl0w characteristics of the membrane oxygenator used in the extracorporeal circuit. Th e incorporation of secondary flows introduces turbulence into the blood flow path. Compared with laminar blood flow, turbulent blood flow requires greater pressure to generate a given blood flow. Thus, a greater pressure force is generated. In addition, secondary flows, or turbulence, generate greater shear rates. Shear rates of 1,000 to 2,000/s were found to be required to influence gas transfer efficiency of membrane oxygenators. 9J1,15,16 The shear rates generated in membrane oxygenators ranged from 480 to 2,100/s.14 Although increases in shear rate can potentially augment gas transfer, an increase in the
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shear-induced effects on the blood als0 occurs. The magnitude of hemolysis was found to increase linearly with increases in shear rate. 14 In addition, the adhesion patterns of platelets and neutrophils were also found to be dependent on shear rate. 23 Thus, both the gas exchange efficiencY and the biocompatibility of blood are affected by the shear rates generated in membrane oxygenators. With the current membrane materials available, the improvement in gas exchange efficiency would have to be at the expense of a reduction in biocompatibility. However, shear-induced effects on the blood introduced by secondary flows may be balanced by the decrease in surface area required for gas transfer because of greater efficiency. Another development that may decrease the blood phase resistance to gas transfer, without a change in oxygenator design, is the use of perfluorocarbons (PFC) as a blood substitute rather than crystalloid or colloid solutions.24 The solubility of 02 is increased in PFCs compared with blood or crystalloid/colloid solutions, which decreases the resistance to diffusion and the requirement of red blood cells in contact with the membrane surface. Animal studies have shown adequate gas transfer ability and stability of PFCs during extracorporeal circulation.25,26 Studies have also shown that efficiency of O2 transfer was greatest for hollow-fiber membrane oxygenators compared with solid membrane and bubble oxygenatorsY ,27 Thus, the use of blood substitutes during extracorporeal circulation may increase the efficiency of gas transfer without the introduction of shear stresses or increases in surface area. Issues that still need to be addressed regarding blood substitutes are the long-term performance characteristics of membrane oxygen~ ators, exposed PFC emulsions, and the biocompatibility of PFCs. OXYGENATOR BIOCOMPATIBILITY
Biocompatibility and hemocompatibility are complex problems that encompass shear stresses, surface area-to-volume ratio, geometry of blood flow, type and composition of the foreign surface, duration of exposure, and patient status. Furthermore, the biocompatibility characteristics of a given material may be different under different flow conditions, or in different oxygenator designs. A systemic inflammatory response is induced in all patients exposed to extracorporeal circulation, part because of the activation of blood components that come in contact with foreign surfaces. However, the severitY of the systemic inflammatory response is highly variable, and only in a minority of patients does the response result in clinically significant adverse events.2S,29The complexity of biocompatibility implies that no single alteration in oxygenator design will effectively improve the biocompatibility of the entire system. The first interaction of blood with an artificial surface results in the instantaneous adsorption of protein to the surface with subsequent protein denaturation or blood component adhesion. 14 In oxygenators without a solid surface, such as bubble oxygenators, the denatured proteins are continually released back into the circulation, where they can exhibit unfavorable biological properties, such as expression of vasoactivity, alteration of coagulation factors, and initiation of red blood cell agglutination with subsequent alterations in microcirculatory blood flow. 3,3° In contrast, the adsorption of protein to a solid surface, such as the membrane interface, results in a thin protein
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layer that remains in place and subsequently alters the reactivity of the surface such that further protein denaturation and blood damage are moderated. 3,4,21The type of protein adsorbed on the membrane surface also influences the biocompatibilitY of a device. Membrane materials that preferentially adsorb albumin have been shown to have a reduced thrombogenic tendency compared with membranes with low albumin adsorption. 17,31 Precoating the extracorporeal circuit with fresh frozen plasma has been shown to decrease complement and platelet activation because of the composition of the protein layer deposited and receptor sites available on the membrane surface before initiation of cardiopulmonary bypass. ~2 Coating the surface of membranes with polymeric phospholipids, to mimic the lipid surface of the endothelium, may be another method to alter the amount and type of protein adsorbed, and to improve biocompatibility.33 Because the biocompatibility of a device depends in part on the degree of blood activation when blood comes in contact with a foreign surface, attenuation of contact activation should improve biocompatibility. One approach to increasing the biocompatibility of a device is to make the surface more endothelial-like. On e primary function of the endothelium is to maintain blood in a fluid and flowing state. The surface of the vascular endothelium is antithromb0genic because of the presence of several proteins, such as heparan sulfate. The heparan sulfate protein functions similarly to heparin sulfate to prevent thrombin formation. Thus, one approach to increasing biocompatibility is to coat the membrane and other extracorporeal surfaces with heparin, either by a covalently or ionically bonding processes. The heparin coating creates an antit!arombotic surface that is both flow-dependent and antithrombin IIIdependent. Most studies Comparing heparin-coated circuits with non-heparin-coated circuits have shown a decrease in complement activation,33-35 a decrease in platelet activation,35 and a decrease in neutrophil activation36 with heparin coating. Overall, the conclusions from these studies are that heparin coating improves biocompatibility. However, the significance of the improvement Of biocompatibility with respect to clinical outcome is still uncertain. In addition to the exPression of antithrombogenic proteins, endothelial cells also express surface proteins, or release substances that influence the activation of platelets, comPlement, and neutrophils. Nitric oxide, a potent vasodilator and platelet inhibitor, is released by endothelial ceils. The addition of nitric oxide to the sweep gas has been shown to significantly reduce platelet activation during cardiopulmonary bypass. 3v Other attempts to mimic the endothelial cell environment include the use of DNA technology to incorporate proteins such as complement receptor 1, a cell surface inhibitor of complement, into membrane surfaces to prevent local complement activation.3s Other future membrane Coatings might include monoclonal antibodies against the adhesion proteins exPressed on platelets and neutrophils.
CONCLUSIONS
The gas exchange efficiency of current membrane oxygenators appears to be sufficient for routine cardiopulmonary bypass. However, further improvements will be required so that
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all patients, including children and small adults, can benefit from the small prime volumes associated with highly efficient oxygenators. Advances in gas exchange efficiency will require a thorough understanding of the factors limiting gas diffusion in the blood phase. An appreciation of how improvements in gas exchange efficiency will impact on the biocompatibility of the device will also be required. The research and development of the intravascular oxygenator may help lead the way to advancements in extracorporeal oxygenator designs. The ability to maintain gas exchange performance for long duration is the primary limitation of current hollow-fiber membrane oxygenators. Knowledge of the contributing factor(s) causing deterioration of membrane performance is lacking. The cause is probably multifactorial and associated with the magnitude of the inflammatory response induced by exposure to extracorporeal circulation. In addition, the deterioration of patient status because of long-term exposure to nonpulsatile blood flow and anticoagulation must be taken into account. Few
studies have examined the nature of the inflammatory response during long-term support. Thus, a better understanding of the components of the inflammatory response and their interactions will be required before improvements in long-term performance can be made. The magnitude of the inflammatory response induced by exposure to extracorporeal circulation is related to the biocompatibility of the circuit or device. Even though the inflammatory response is induced in all patients, the magnitude of the response varies widely. Thus, a clear picture of the impact of biocompatibility on i~atient outcome is lacking and has slowed down advancements in this area. The importance of biocompatibility will be greatest in critically ill patients as well as patients younger than 2 years and older than 80 years. Again, a better understanding of the interaction of patient status on the magnitude of the inflammatory response will result in further improvements in the biocompatibility of membrane oxygenators.
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