where Coxy = O2 concentration in blood (mlO2/Lblood), tHb = hemoglobin concentration (g/dl), %SAT = hemoglobin O2 saturation (%), Qb = blood flow rate (L/min), Coxy-post = postoxygenator O2 concentration (mlO2/Lblood), and Coxy-pre = preoxygenator O2 concentration (mlO2/Lblood). The CO2 transfer rate (VCO2, ml/min) was calculated as given in Equation 8:
where ΔCO2 = % CO2 in the outlet gas sample and F = sweep gas flow (L/min). The error bars denote the standard error of mean for each measurement.
The CFD results, as shown in Figure 2, show a well-distributed blood flow profile through the device with minimal regions of low blood flow. The results further demonstrate the formation of transient flow vortices through the M-Lung housing. These vortices are dampened by the membrane fiber bundle, with the magnitude of dampening dependent on the permeability of the membrane fiber bundle (Figure 2B).
A comparison of the flow through the M-Lung obtained using simulations and optical flow visualization is shown in Figure 3B. The PIV results demonstrate similar flow patterns to those obtained computationally; after the fluid passes through the gate of the concentric divider.
In in vitro studies, the M-Lung prototype (Figure 6A), comprising a fiber bundle surface area of 0.28 m2 and priming volume of 47 ml, demonstrated a rated flow of 2.0 L/min (Figure 6B). Further, the device was able to remove 200 ml/min of CO2 at rated flow using a sweep gas of 16.0 L/min (Figure 6C). The transit time at rated flow is 1.8 s (47 cc volume at 2,000 cc/min).
The blood-side pressure drop across the device comprising a fiber bundle of porosity = 0.59 was 49 and 106 mm Hg at blood flow rates of 1.0 and 2.0 L/min, respectively (Figure 7). This pressure drop was within the specifications designed for 1 L/min blood flow and pressure drop 60 mm Hg. The highest shear forces generated within the housing are 30 dyne/cm2.
The permeability and porosity of the fiber bundle can be modulated in the fabrication process by varying the number of fibers/cm or the spacing between each layer of fibers. The relationship between fiber bundle permeability and simulated M-Lung pressure drop, priming volume, and fiber surface area is shown in Figure 8, A–C. The simulations demonstrate that looser wrapping (or fewer fibers per mat) would result in increased porosity and permeability accompanied by significant reduction in the pressure (Figure 8A). This would result in larger calculated void volume (Figure 8B) and lower surface area (Figure 8C) resulting in less O2 transfer per square meter.
The key to maximize O2 transfer per surface area in membrane lungs is mixing, and thereby disrupting the diffusion boundary layer, through an increase in vorticity and secondary flows.2 Flow through circular paths will induce secondary flows to create mixing, yet a long circular spiral has high resistance. Concentric circles connected by gates created secondary flows without increasing resistance. Intense vortices are created within the blood path (Figure 2), but these vortices are dampened by the fiber bundle. Significant mixing also occurs as blood passes through the gates. The net result of the mixing on oxygenation manifested as efficiency of oxygenation per surface area and as rated flow. In vitro studies demonstrate highly efficient gas exchange within the M-Lung; the prototype maintained an outlet O2 saturation of ≥95% for blood flow rates up to 2.0 L/min with a fiber bundle surface area of 0.28 m2 and priming volume of 47 ml. The rated flow of the prototype was twice the goal in our specifications. Several similar size devices are compared in Table 2. The O2 transfer efficiency (mlO2/m2/min) of the M-Lung prototype is 357 mlO2/m2/min. The O2 transfer efficiency of similar commercial devices with predominantly straight flow path is 128–204 mlO2/m2/min. The experimental Pedipump lung has a single circular chamber, and blood is driven by an integral centrifugal pump.22 The O2 transfer efficiency is 400 cc/m2/min.
The CFD results show a well-distributed blood flow profile through the M-Lung with minimal regions of low blood flow. The results also demonstrate the formation of transient flow vortices through the M-Lung housing, suggesting that the gated design promotes passive secondary flow mixing. This mixing is dampened by the membrane fiber bundle, with the magnitude of dampening dependent on the permeability of the membrane fiber bundle. The simulations demonstrate that an increase in secondary flow mixing and a significant reduction in the pressure drop across the M-Lung could be achieved by increasing the fiber bundle permeability. The specific pattern of secondary flows generated depends on a variety of factors, including inlet flow rate and pulsatility and specific configuration of dividers and gates. Because the inlet flow rate and pulsatility are dependent on the perfusion system (pump or native circulation), variations of this design will incorporate a variety of divider and gate configurations to meet different specifications.
Despite its low priming volume and efficient gas exchange, the M-Lung maintains a sufficiently low blood-side pressure drop across the relevant range of blood flow rates. At a blood flow rate of 1.0 L/min, the M-Lung had a blood-side pressure drop <50 mm Hg, which is within the range supported by a systemic arteriovenous pressure gradient for the intended application of CO2 removal. At its rated flow of 2.0 L/min, the M-Lung prototype had a pressure drop of 106 mm Hg, which is acceptable but very high for pump perfusion and too high for pulmonary artery perfusion.
The pressure drop can be regulated to any desired level by changing the density of fibers (porosity or permeability) and the size of the device, which we will do for other prototypes depending on the driving pressure source. These relationships simulated by CFD are shown in Figure 8.
We did not test hemolysis, thrombogenicity, or durability because the housing components were fabricated by 3D printing which creates a rough surface. The next prototypes will be injection molded or machined, creating a very smooth surface for durability and blood damage testing.
The prototype was able to remove 200 ml/min of CO2 at the rated flow and gas-to-blood flow ratio of 16:1. The CO2 clearance is largely dependent on the sweep gas flow rate-to-blood flow rate ratio and can be further increased by increasing the sweep gas flow. The relatively short gas exchange fibers allow the M-Lung to maintain a sufficiently low gas side pressure drop at higher sweep flows and minimize water accumulation in gas phase enhancing the CO2 removal efficiency/capacity of the device. Testing of the durability of CO2 clearance awaits prolonged animal trials.
This M-Lung design meets the functional requirements for providing CO2 removal for adults and total respiratory support for infants and small children (with a pump). In addition to sufficient gas transfer, it comprises a low priming volume, allowing for minimal hemodilution and rapid transit time. The rapid transit time at rated flow and the absence of corners where blood can stagnate should minimize thrombogenicity. Thrombogenicity and anticoagulant requirements will be assessed in chronic animal trials. Hemolysis testing will be done in vitro and in vivo using devices with smooth blood contact surfaces.
We intentionally began these studies with a small membrane lung design. Given its compact size, controllable resistance, and potential for low thrombogenicity, this size M-Lung has potential for providing long-term ambulatory respiratory support, in arteriovenous mode for clearance of CO2 for patients with end-stage chronic obstructive pulmonary disease, or (at higher porosity) pulmonary artery to left atrium mode for total lung support in children. The design could be scaled up to achieve total adult respiratory support (i.e., a rated flow of 5 L/min). Scaling up to adult size devices will be done with a return to CFD and PIV to optimize the properties of a device with a 5 L/min rated flow, pressure drop <120 mm Hg at rated flow, and CO2 transfer two times O2 transfer at rated flow. This would require 0.6–0.8 m2 of gas transfer surface, in contrast with current contemporary adult devices, which comprise 1.5–1.8 m2.
In conclusion, using CFD and optical flow visualization methods, a unique hollow fiber membrane lung has been designed based on concentric circular blood flow paths connected by gates. In prototype testing, flow patterns and pressure drop across the M-Lung device matched those predicted by CFD. The prototype M-lung has a rated flow of 2 L/min with a fiber surface area of 0.28 m2, and the oxygenation efficiency is 357 ml/min/m2. The CO2 clearance of the M-lung is 200 ml/min at the rated blood flow. Future work will evaluate the performance of the M-Lung in in vivo to further assess its biocompatibility and chronic performance.
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artificial lung; circular flow; mixing; oxygenationCopyright © 2017 by the American Society for Artificial Internal Organs