DOI 10.1515/bmt-2013-0026 Biomed Tech 2014; 59(2): 125–133
Georg Wagnera,*, Peter Schlansteina, Sandra Fiehe, Tim Kaufmann, Rüdger Kopp, Ralf Bensberg, Thomas Schmitz-Rode, Ulrich Steinseifer and Jutta Arens
A novel approach in extracorporeal circulation: individual, integrated, and interactive heart-lung assist (I3-Assist) Abstract: Extracorporeal life support (ECLS) is a wellestablished technique for the treatment of different cardiac and pulmonary diseases, e.g., congenital heart disease and acute respiratory distress syndrome. Additionally, severely ill patients who cannot be weaned from the heart-lung machine directly after surgery have to be put on ECLS for further therapy. Although both systems include identical components, a seamless transition is not possible yet. The adaption of the circuit to the patients’ size and demand is limited owing to the components available. The project I³-Assist aims at a novel concept for extracorporeal circulation. To better match the patient’s therapeutic demand of support, an individual number of one-size oxygenators and heat exchangers will be combined. A seamless transition between cardiopulmonary bypass and ECLS will be possible as well as the exchange of components during therapy to enhance circuit maintenance throughout long-term support. Until today, a novel oxygenator and heat exchanger along with a simplified manufacturing protocol have been established. The first layouts of the unit to allow the spill- and bubble-free connection and disconnection of modules as well as improved cannulas and a rotational pump are investigated using computational fluid dynamics. Tests were performed according to current guidelines in vitro and in vivo. The test results show the feasibility and potential of the concept. Keywords: ECLS; HLM; individual therapy; modular design; normative testing.
These authors contributed equally to this article. *Corresponding author: Dipl.-Ing. Georg Wagner, Department of Cardiovascular Engineering, Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, Pauwelsstraße 20, 52074 Aachen, Germany, Phone: +49 241 80 89 888, Fax: +49 241 80 82 144, E-mail: [email protected] Peter Schlanstein, Sandra Fiehe, Tim Kaufmann, Thomas Schmitz-Rode, Ulrich Steinseifer and Jutta Arens: Department of Cardiovascular Engineering, Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, 52074 Aachen, Germany a
Rüdger Kopp and Ralf Bensberg: Department of Intensive Care Medicine, University Hospital Aachen, RWTH Aachen University, 52074 Aachen, Germany
Background In extracorporeal circulation (ECC), two groups are distinguished [1, 7, 8]: 1. For short-term support of several hours during cardiac surgery, a heart-lung machine (HLM) or cardiopulmonary bypass (CPB) is used. The first use of a HLM in an adult patient was reported by John Gibbon in 1953 [19]. The need for blood transfusions was an issue right from the start of ECC, owing to infections and the amount of blood needed [40]. Although Beall et al. [9] reported successful open heart surgery without excessive use of blood transfusions in 1967, hemodilution remained unfeasible for small patients. During the next decades, blood-free surgery in neonatal and pediatric patients was achieved by using hypothermia and minimized circuits with reduced tubing length and diameter [16, 20, 28, 30, 42]. To further decrease the need of blood transfusions, vacuum-assisted venous drainage has been established, but results in undesirable side effects such as suction events or venous collapse [14, 22, 32, 33, 36, 44, 45]. 2. For long-term support from several hours up to several weeks, an extracorporeal life support (ECLS; i.e., heart and lung support) or an extracorporeal lung assist can be used. Even in the 1950s, researchers thought of a continuing support to allow for recovery of the heart and lung. It took another 20 years until the first ECLS was used successfully in 1972 [23]. Regarding ECLS, circuit-related complications become more likely the longer the support lasts [11, 18, 21, 31]. Although the exchange of components is difficult [31], preparations of the circuit can be made, e.g., a second line containing a primed oxygenator in case of
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126 G. Wagner et al.: Individual, integrated, and interactive assist mechanical failure of the oxygenator [1]. The critical parameters regarding the mortality rate of patients on ECLS are, among others, the comorbidities [6, 11, 18, 31], the time needed to initiate the circuit [1, 6, 10, 25, 29, 39, 41, 43], and the duration of support [21, 43]. The time needed to start the system can be reduced by the experience of the user and the design of the system [1, 11, 31, 41, 43], including the possibility of transport on a cart [31] or in a car [11] to get the ECC to the patient. Whereas HLM is used in the operating room to enable cardiac surgery, ECLS is used in a wide range of bridging techniques. Bridge-to-recovery includes patients with possible reversible indications such as myocarditis or cardiogenic shock during or after surgery; bridge-to-bridge comprises all patients waiting for a permanent assist device, while bridge-to-transplant consists of patients awaiting organ transplantation [1, 6, 11, 21, 31, 37, 41]. A HLM usually consists of a venous reservoir, a pump, an oxygenator with integrated heat exchanger, and the tubing. An ECLS circuit contains a pump, an oxygenator, and tubing. Although the type of pump and oxygenator are usually different in both systems, essentially identical components for the most parts are applied. However, there is no system yet available that allows switching between the two modes in order to reduce the trauma to the patients, e.g., in those patients who cannot be weaned from HLM after surgery [1, 6]. The priming volume, owing to the dimensions of the components (pump, oxygenator, reservoir, tubing), causes hemodilution [16, 30, 38]. The surface area in contact with blood, resulting mostly from the membrane but also from the dimensions of the components, causes an inflammatory response and the activation of clotting
[16, 18, 37]. Clotting is treated with anticoagulants applied as a coating of the components before use or injected directly into the circuit [1, 6, 16, 18, 37]. At the cannulation site, identified as the major contributor to the clotting activation, this anticoagulation can lead to bleeding complications [37, 38, 43]. Although a wide range of oxygenators, reservoirs, and other circuit components are available on the market, they can be divided into the three general size ranges (pediatric, adolescent, and adult) according to their performance regarding maximum blood flow and gas exchange rates. In order not to operate the ECC beyond its range, the next larger components are chosen rather than the small device. Thereby, the priming volume and surface area of the ECC are exceeding the therapeutic need in most cases as the continuous body weight range from neonates to adults has to be covered with the discrete sizes available. This results in the circuit exceeding the actual therapeutic demand, thus increasing priming volume and artificial surface area and, thereby, the rate of complications. Numerous publications report a reduced length of tubing in order to create a smaller circuit; it appears to be common practice to reduce the circuit-related side effects [30, 33, 34, 38, 44]. To address these issues, the project I³-Assist aims at developing a highly integrated and modular extracorporeal system that can be adapted to the changing individual therapy of a patient. The key aspect of the I3-Assist system is the modularity. This is being achieved by fragmentation of the functional units, heat exchanger and oxygenator, into small modules with one size only (see schematic drawings in Figure 1). An optimized gas and heat exchange surface area and thus a minimized priming volume of the system are therefore realized. During the assembly of the system, the modules can be combined by the user to
Figure 1 Conceptual idea of the project. ECLS, extracorporeal life support; HLM, heart-lung machine; Oxy, oxygenator; HE, heat exchanger.
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accomplish the suitable exchange rate, e.g., one oxygenator module for a neonate, five oxygenator modules for a small adult, or eight oxygenator modules for a tall adult. All components will be assembled before the surgery by the perfusionist on a connecting unit. The oxygenator and heat exchanger modules of the HLM/ECLS system will be exchangeable under operating conditions. Thus, defective components can be exchanged easily. Additionally, an immediate and seamless transition between operation modes can be carried out and the system can be modified according to changing individual needs during surgery and therapy. Owing to the highly integrated design, the system can be placed near the operating table and can be used for inter- and intrahospital transport. This article gives an overview of the status quo of the ongoing research project and further steps planned.
Methods The project is carried out in a partnership of four parties: –– The Department of Intensive Care Medicine of the University Hospital RWTH Aachen as a clinical partner and provider of the users’ perspective; –– The qcmed GmbH (Aachen, Germany) as a consultant to implement a quality management (QM) system that suits all partners’ needs and enables an easy technology transfer between the partners; –– A company with manifold experiences in the manufacturing and distribution of HLM and ECLS devices; and –– The Department of Cardiovascular Engineering (CVE), Institute of Applied Medical Engineering (AME), University Hospital RWTH Aachen as a research facility. Primarily at the start, the project highly profited from the vicinity of the partners’ location, as strategy and engineering workshops as well as risk analysis meetings were frequent. The communication plan and the assignment of tasks had to be defined to put all competences to optimal use; the short traveling distance was of great benefit. During the requirement specification process, a basic draft listing all functional units was developed. These functional units are (in the order of blood flow): cannulas, reservoir, pump, heat exchanger, oxygenator, and a connecting unit, like a mounting plate to link all units. The research, as well as the construction and development, of these components were divided between the company and the CVE/AME.
An online tool enabled access to the latest versions of strategic documents, such as URS, DRS, and RIA, for all partners at any time. For details concerning the quality management within this project, please be referred to the article “Implementation of Quality Management in Early Stages of Research and Development Projects at a University” within this journal [17]. The general specifications were defined by all parties. The overall system dimensions should not exceed 0.15 m³, whereas each disposable component should be able to fit in a standard hospital bin. On the basis of input from the user, the minimal circuit containing one oxygenator to support one infant should allow for a nominal blood flow of 800 ml/min. The pump should provide a flow range from 0 up to 8000 ml/min and be able to create a pressure of up to 650 mm Hg at maximum flow. As CAD software ProEngineer Wildfire 5.0 (Parametric Technology Cooperation, Needham, MA, USA) was used by the CVE/AME, the company used different software. As the different CAD systems were well established in each party, we refrained from establishing a corporate system. For file sharing between research, development, and manufacturing, the platform-independent file formats IGES and STEP were used. We encountered no problems during the exchange and further processing of the files emanating from the different software solutions used. The test parameters for oxygenators defined in DIN EN 12022 [13] and ISO 7199 [24] allow for a wide range of input settings. Depending on the chosen setting within this range, the performance of the same oxygenator may vary in an even wider range, due to the non-linear behavior of oxygen binding. Blood is used as test fluid for performance and hemolysis tests, which differs on a daily basis even within the same species and when collected from more than one specimen. The quality is highly dependent on proper handling during collection and transport. Therefore, performing comparable hemolysis test for several design variants requires performing the test on all designs on the same day and with the same blood. To reduce the need for blood, facilitate the handling, and create results that are reliably comparable with each other, blood replacement fluids are used frequently at our institute when applicable (e.g., for hydraulic tests).
Results An initial feasibility study was performed using a Medos Hilite hard-shell reservoir for children (Medos Medizintechnik AG, Stolberg, Germany) as fluid capacity, Hilite
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128 G. Wagner et al.: Individual, integrated, and interactive assist 800 LT oxygenators, and a Medos deltastream DP3 blood pump with DC driving unit. A HAES mixture (Fresenius Kabi Deutschland GmbH, Bad Homburg, Germany) with a dynamic viscosity of 3.63 mPas at 20°C was used as blood substitute. This study was conducted to show the feasibility of a passive flow distribution using three parallel lines, as shown in Figure 2. One design for both the dividing and the collection unit were defined as spherical sectors. The setup enabled the quantification of the flow distri bution within the three lines and a study of the flow distri bution with one line being obstructed. At the nominal flow, the oxygenators developed a mean pressure loss of 33 ± 2 mm Hg per line. The 1/4-in tubing within each line created a pressure loss equal to 60% of the oxygenator. Each line was equally exposed to the same flow within a range of ± 3%. The results suggested that a uniform passive flow distribution was nearly independent of the relative arrangement of the lines. The flow resistance of heat exchangers and oxygenators supported the uniformity by reducing the influence of fluid-wall interactions, which increase with longer flow paths.
Construction and design The manufacturing of those laboratory models that needed conventional machining was done in the CVE/ AME’s workshop. Immersion processes were performed by the company involved. Early testing of concepts and variations was elementary during the design process. Therefore, rapid prototyping was used frequently especially at the project start and was performed in the CVE/ AME’s workshop using Full Cure 720 material (Objet
Figure 2 Initial feasibility study of passive flow deviation.
Geometries Ltd., Rehovot, Israel). This allowed fast comparisons between different design options. As advantageous as a three-dimensional (3D) printer may be, there are a few limitations regarding the system we used. The material used is not suitable for contact with blood; it macerates in contact with water and shows plastic flow even at temperatures below 50°C. Although outer and inner structures can be created without the limitations of conventional machining, the partly required postprocessing of oriented parts is restricting the creative scope for development. Additionally, the costs are still considerably high; in some cases, conventional manufacturing is less expensive. However, it is beneficial that it takes only a couple of hours to print a component.
Cannulas The key aspects regarding the cannula development were minimizing the trauma at the cannulation sites, reducing risk of suction events at the withdrawal cannula, and an improving hemodynamic performance along with an optimized flow pattern at the return cannula. A jet is usually produced at the outflow site. The jet’s impact may loosen plaques, damage the aortic wall, and lead to low perfusion of the upper extremities [26]. Design studies on the return cannula were performed, investigating the effects of outer and inner structures on the expansion of the jet. The hemolytic potential of the requisite design elements has to be investigated closely. After hydraulic tests with several new design options, hemolysis test were performed with two different designs. In vitro testing was performed according to available standards and guidelines (such as ISO 7199 [24] or ASTM F 1830 [5]). Comprehensive tests regarding the hydraulic performance were conducted using the laboratory models with both blood substitute and porcine blood. Porcine blood was collected from an abattoir in a 5-l canister that was prepared with 75,000 IU heparin-sodium, 8 ml of 50% glucose solution, and 60 ml of 0.9% saline solution. A hemolysis test for one new cannula design with corresponding control circuits was performed. A 24-Fr Medos arterial cannula (Medos Medizintechnik AG, Stolberg, Germany) with an angled tip was used as a reference device. A circuit without cannula was used as a reference circuit. Each circuit contained the same kind and length of tubing, reservoir, and roller pump. The occlusion and rotational speed of the roller pumps were calibrated to ensure comparable settings. The results showed a higher hemolysis rate in one design; further investigations were performed using computational fluid dynamics (CFD)
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simulations as more design options could be investigated and compared more easily in a shorter period. Design features cannot be displayed to not affect the patent process. CFD simulations of blood flow deviation in the aortic arch dependent on different cannula designs were performed using a method that was established and experimentally validated in previous studies [26]. 3D models of different cannula tip designs were incorporated in a CAD model of the human aortic arch, including the three supraaortal vessels. The model was meshed with unstructured tetrahedral elements. Even though mesh independency was already achieved at 100,000 elements, a fine mesh density of 2.7 million elements was chosen to provide a higher spatial resolution of the final results. Furthermore, five layers of prismatic elements were generated to resolve boundary layers and to more accurately calculate wall shear stress. Flow rates of 2000, 4000, and 6000 ml/min were simulated using a non-Newtonian blood model with a hydrodynamic viscosity threshold value of 3.6 mPas for high shear rates. Arterial outlets were modeled as static pressure openings at 55 mm Hg. Additionally, a pressure drop due to systemic resistance was modeled using the Darcy-Weißbach equation. 1 ∆P = L ρU n2 2 In this equation, ρ is the fluid density and Un is the normal velocity at each boundary. L is defined as a dimensionless loss coefficient and set to 50 based on a sensitivity study [27]. Investigations of the cannula comparing the current design and the new designs were conducted. The results for a flow rate of 6 l/min and a constant aortic pressure
showed an increase of 9.5% in shear rate for new tip #1, whereas both others performed similar. The maximum pressure on the aortic wall was reduced by 4.9% in new tip #2 and by 22.8% in new tip #1 compared to the standard tip. A comparison of new tip #1 and the conventional tip is shown in Figure 3; the improved stress level on the aortic wall is achieved by a smoothed flow pattern with good expansion of the jet. Owing to the new fluid path, the pressure loss over the cannula is slightly increased.
Reservoir At first, the objectives of the new reservoir were to combine the benefits of a hard-shell reservoir with those of a soft venous blood bag. The hard shell allows a defined position and for venous drainage when sealed airtight. The bag’s ability to inhibit blood-air contact, a major trigger of coagulation, is advantageous. In vitro tests with the first prototypes showed that the available elastic materials were not suitable for the required performance; the elastic deformation was exceeded, disabling the return to the initial shape. Therefore, a reassessment of the requirements was necessary. A survey performed among perfusionists revealed a bias toward hard-shell reservoirs. A new design is currently being developed.
Pump A rotational pump is advantageous compared with a roller pump, especially for long-term applications, as the hemolytic rate is considerably lower [1] and no electrostatic charge is developed. Regarding the transport in and between hospitals, roller pumps are considerably heavier
Figure 3 Pressure loss and shear stresses induced to the aortic wall: comparison of a current tip (left) and new design #1 (right).
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than rotational pumps. Additionally, the priming volume is higher because of the tubing length required inside the roller pump [2, 38]. The wide range of flow rate results from the modularity of the system, as the performance should be good for low and for high flow rates. The first design, generated with the support of CFD simulation, is currently being improved. The hydraulic performance matched the expectations; however, once a larger amount of air had entered the pump chamber, the hydraulic performance dropped drastically. Furthermore, the air was removed very slowly, thereby reducing the blood flow for a longer period.
Oxygenator Two types of oxygenators can be distinguished by their manufacturing process: those that contain a coiled fiber bundle and those that contain a layered fiber bundle. In our experiments, we found that a coiled oxygenator usually shows a higher pressure loss, as the flow is deflected more often to ensure a counter flow of blood versus gas and water. A layered oxygenator is usually easier to prime, owing to the less complicated structure, but shows less uniform flow deviation when used at lower flow ranges. Another distinction can be made by the hollow fibers used; fibers used in HLM are not plasmaproof over periods of more than 12 h [15, 35]; those fibers used in ECLS last up to 7 days. This results in the limited working time of the oxygenators. A third type, silicone membranes, were used until recently, especially in the United States, but are now being replaced successively because of their comparatively high pressure loss and lower gas transfer rates. Regardless of the inner structure, the gas exchange rate within an oxygenator should be ≥ 55 ml/min of oxygen and ≥ 50 ml/min of carbon dioxide per liter blood at all times. The development of the one-size oxygenator within this project was driven by the therapeutic need of the smallest patients. The requested nominal blood flow of 800 ml/min results in a gas exchange area of approximately 0.35 m2, as shown in a comparison of available oxygenators in Figure 4, extracted from publicly available data sheets (e.g., Medos Medizintechnik AG, Stolberg, Germany; Sorin AG, Milan, Italy; Maquet GmbH, Rastatt, Germany; Terumo Corporation, Tokyo, Japan). The first rapid prototyped models of the oxygenator modules consisted of a coiled fiber bundle and simple housing parts. The device was slim and tall, with the outer diameter of the fiber bundle being 56 mm and the length being 60 mm. The geometry arose partly from
Surface area for gas exchange in m2
130 G. Wagner et al.: Individual, integrated, and interactive assist
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Max. blood flow in l/min Maquet safe micro
Sorin lilliput D901
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Figure 4 Comparison of available oxygenators regarding surface area and blood flow.
the comparative sample we used, the Medos Hilite 800 LT (Medos Medizintechnik AG, Stolberg, Germany), and partly from the preceding research projects at the CVE/ AME, the NeonatOx and the MiniHLM [3, 4]. The overall dimensions of the whole system defined in URS and DRS compelled a revision of this concept. The current patent pending version contains the same surface area but it is more compact. It contains a layered bundle with an outer diameter of 50 mm and a height of 30 mm. A special focus was an easy manufacturing process, on the one hand to enable the assembling in our laboratory, and on the other hand to reduce production costs later on. The latter will reduce costs for the user as well, regarding procurement prices and inventory management. The number of housing parts could be reduced to three, two of which are identical for heat exchanger and oxygenator. The potting, a process by which fibers and housing are joined using a glue and the separation of blood and fluid is ensured, was performed following a simple protocol. Silicone is used as potting material by the CVE so that the custommade housings can be reused. A custom-made centrifuge, including controller, was developed and built for the new potting process of oxygenator and heat exchanger modules. Simulations of the oxygenator were performed regarding the pressure loss, the fluid path, and stagnation areas. An isotropic model as well as an anisotropic model, to account for the direction of the fibers, was used to embed the permeability of the bundle. Different inlet geometries for the latest oxygenator were investigated, supporting an optimized flow distribution. As there is no powerful simulation tool yet to calculate the efficacy of construction elements and design changes regarding the gas exchange
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capability, detailed tests leading to design and test iterations were expected and necessary. Regarding in vitro testing, a new test setup was implemented to conduct long-term performance tests up to 12 h on the oxygenators’ gas exchange efficiency. A custommade double-reservoir allows for connecting two circuits in a cross-flow manner. One is the test circuit for the oxygenator, the other one is a deoxygenation circuit, an oxygen consumer, and carbon dioxide producer, which is used to maintain venous parameters for the blood entering the test circuit. All tests were run according to the newly developed test records, which enabled a sound comparison of different test objects. Five oxygenators of the latest design showed an average performance factor of 96% of the required oxygen transfer and 107% of the required CO2 removal. To further optimize the design of oxygenators, better simulation tools are needed. The flow patterns inside a fiber bundle cannot be described with numerical methods yet, at least not on a large scale. The current approach of an anisotropic model needs further improvement to support the design process. In addition, a more reliable prediction of the hemolytic potential using simulations would be helpful. Thus far, this can only be done qualitatively by comparison of different designs. Accurate calculation of the normalized index of hemolysis (NIH) is not yet possible but would be a powerful tool to support the design process. Regarding the design of the oxygenator, the fibers are the limiting part in current designs. The way of processing required, the gas exchange efficiency, and the longterm durability are constraints in new developments. New materials would greatly contribute to further progress toward long-term stable and further miniaturized oxygenators.
Heat exchanger Commercially available oxygenators usually contain a heat exchanger; the modularization allows separate modules for both. The heat exchanger and oxygenator should share identical housing parts as many as possible to minimize production costs. Additionally, their performances have to match each other. Therefore, the heat exchangers’ development was postponed until construction and design of the oxygenator was completed. The manufacturing protocol is identical to the oxygenators’, except for the fibers used. The heat exchanger performance factor in vitro was R = 0.8, calculated according to ISO 7199 [24]:
R=
TBlood ,out -TBlood ,in TWater ,in -TBlood ,in
Our research examining publically available data sheets showed that commercially devices have factors between 0.63 and 0.8. The modularization of the heat exchanger and oxygenator creates a high grade of flexibility regarding the adaption of the circuit during therapy. The surface area of materials in contact with the patients’ blood can be reduced to a minimum at all times, but the priming volume of the system is increased by the number of modules. The youngest patients could most likely benefit more from a further decreased priming volume. Therefore, the separation of heat exchanger and oxygenator should be thoroughly surveyed.
Connecting unit All fluids are channeled by the connecting unit, which should allow for attachment and removal of the modules. Connecting modules should be possible without leakage of blood, water, or gas into or out of the circuit. Furthermore, a hemocompatible flow without stagnation areas, swirls, or interfering details must be provided. Additionally, priming of the blood paths and general handling should be easy when operating all fluid channels of a module at the same time. The channel system, replacing the tubing used to date, is designed with all fluid paths separated and all necessary connections. Simulations will be used to investigate the flow distribution within the channels. By now, we have designed a laboratory sample of a connector fulfilling the requirements outlined above, which is currently being manufactured in the workshop and subsequently will be thoroughly tested.
In vivo testing The in vivo testing was authorized by the government of the state of North Rhine-Westphalia, Germany. All animals received medical care according to the German guidelines for laboratory animal care, which are in line with the National Institutes of Health guidelines. On the basis of the test records for the in vitro tests, protocols for in vivo tests have been developed. A novel animal model was developed by the clinical partners in order to mimic the transition between HLM and ECLS use of the devices during
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132 G. Wagner et al.: Individual, integrated, and interactive assist surgery and subsequent intensive care. The novel protocol was validated before the final device testing using commercially available components in two acute (short-term) tests in a pig model. Reversible hypoxia and hypercapnia were induced by mechanical ventilation with decreased oxygen percentage and increased level of carbon dioxide. This first validation of the novel animal testing protocol showed good results; the animals were hydrodynamically stable and could be supported according to the varying levels of hypoxia as the blood gas levels were maintained within a normal range by the circuit.
Discussion and future prospects Today, users are preparing their circuits to exchange components more easily in case mechanical failure of a device occurs during operation [1, 31]. The demand to extend the use of ECLS by adding new functions to it, e.g., a possibility to run renal replacement therapy in parallel or by offering customized support solutions, is expressed in literature. The hopes are that such systems could save lives and costs while sparing resources at the same time [11, 12, 21, 30, 37]. The chance to develop a whole new ECC with all its components and search for integrative potential enabled new ways in the construction and design of these components, which would not have been seen in separate research of each component. The close cooperation of the partners and their will to combine their diverse expertise was beneficial to establish a fast and productive communication. This was notably effective during the start of the project. The project is still ongoing. The feasibility of the concept was already shown in vitro as well as in vivo as most of the components are available as laboratory
models. Some components required more research and development as expected during planning, e.g., the connecting unit, particularly the connector itself. Further tests of the performance and hemolytic behavior of the components are going to be performed at the CVE/AME. In vivo tests are planned to validate the overall design of the system. The transfer into industrially produced prototypes will be the next step after the in vivo validation of the design in long-term-use. A potential field of research could be the level of oxygen saturation created by the ECC. Clinicians reported weaning from ECLS when a saturation of 65% could be established. If the required performance is able to be reduced from the current demand of a saturation of 100%, the dimensions of the device, the surface area, the priming volume, and the durability of the fibers could be improved. This would probably reduce the demand on anticoagulants as well. Acknowledgments: This project was performed in collaboration between the industrial partner, the QM consultant qcmed GmbH, the Department of Intensive Care Medicine of the University Hospital RWTH Aachen, and the Department of Cardiovascular Engineering (CVE), Institute of Applied Medical Engineering (AME), University Hospital RWTH Aachen. The project is co-funded by the European Union (ERDF – European Regional Development Fund – Investing in your future) and the German federal state North Rhine-Westphalia (NRW), under the operational program Ziel2 “Regional Competitiveness and Employment” 2007–2013 (EFRE). Received March 28, 2013; accepted November 6, 2013; online first December 11, 2013
References [1] Allen S, Holena D, McCunn M, Kohl B, Sarani B. A review of the fundamental principles and evidence base in the use of extracorporeal membrane oxygenation (ECMO) in critically ill adult patients. J Intensive Care Med 2011; 26: 13–26. [2] Ando M, Takahashi Y, Suzuki N. Open heart surgery for small children without homologous blood transfusion by using remote pump head system. Ann Thorac Surg 2004; 78: 1717–1722. [3] Arens J, Schoberer M, Lohr A, et al. NeonatOx: a pumpless extracorporeal lung support for premature neonates. Artif Organs 2011; 35: 997–1001. [4] Arens J, Schnoering H, Pfennig M, et al. The Aachen MiniHLM – a miniaturized heart-lung machine for neonates with an integrated rotary blood pump. Artif Organs 2010; 34: 707–713.
[5] ASTM. ASTM F 1830:1997 – standard practice for selection of blood for in vitro evaluation of blood pumps, 1997. [6] Bakhtiary F, Keller H, Dogan S, et al. Venoarterial extracorporeal membrane oxygenation for treatment of cardiogenic shock: clinical experiences in 45 adult patients. J Thorac Cardiovasc Surg 2008; 135: 382–388. [7] Bartlett RH. Extracorporeal life support in the management of severe respiratory failure. Clin Chest Med 2000; 21: 555–561. [8] Bartlett RH, Roloff DW, Cornell RG, Andrews AF, Dillon PW, Zwischenberger JB. Extracorporeal circulation in neonatal respiratory failure: a prospective randomized study. Pediatrics 1985; 76: 479–487.
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G. Wagner et al.: Individual, integrated, and interactive assist 133 [9] Beall A Jr, Yow E Jr, Bloodwell RD, Hallman GL, Cooley DA. Open heart surgery without blood transfusion. Arch Surg 1967; 94: 567–570. [10] Chen Y-S, Yu H-Y, Huang S-C, et al. Experience and result of extracorporeal membrane oxygenation in treating fulminant myocarditis with shock: what mechanical support should be considered first? J Heart Lung Transplant 2005; 24: 81–87. [11] Combes A, Leprince P, Luyt C-E, et al. Outcomes and long-term quality-of-life of patients supported by extracorporeal membrane oxygenation for refractory cardiogenic shock. Crit Care Med 2008; 36: 1404–1411. [12] Cruz D, Bellomo R, Kellum JA, de Cal M, Ronco C. The future of extracorporeal support. Crit Care Med 2008; 36: S243–S252. [13] DIN EN 12022:1999 – Blutgasaustauscher. DIN Deutsches Institut für Normung e.V., Berlin: Beuth Verlag, März 1999. [14] Durandy Y. Blood transfusion in pediatric cardiac surgery. Artif Organs 2010; 34: 1057–1061. [15] Eash HJ, Jones HM, Hattler BG, Federspiel WJ. Evaluation of plasma resistant hollow fiber membranes for artificial lungs. ASAIO J 2004; 50: 491–497. [16] Edmunds LH. Advances in the heart-lung machine after John and Mary Gibbon. Ann Thorac Surg 2003; 76: S2220–S2223. [17] Fiehe S, Wagner G, Schlanstein P, et al. Implementation of quality management in early stages of research and development projects at a university. Biomed Eng/Biomed Technik 2014; 59: 135–145. [18] Fleming GM, Gurney JG, Donohue JE, Remenapp RT, Annich GM. Mechanical component failures in 28,171 neonatal and pediatric extracorporeal membrane oxygenation courses from 1987 to 2006. Pediatr Crit Care Med 2009; 10: 439–444. [19] Gibbon J Jr. Application of a mechanical heart and lung apparatus to cardiac surgery. Minn Med 1954; 37: 171–185; passim. [20] Golab HD, Takkenberg JJ, Bogers AJJC. Specific requirements for bloodless cardiopulmonary bypass in neonates and infants; a review. Perfusion 2010; 25: 237–243. [21] Haines NM, Rycus PT, Zwischenberger JB, Bartlett RH, Undar A. Extracorporeal life support registry report 2008: neonatal and pediatric cardiac cases. ASAIO J 2009; 55: 111–116. [22] Hickey E, Karamlou T, You J, Ungerleider R. Effects of circuit miniaturization in reducing inflammatory response to infant cardiopulmonary bypass by elimination of allogeneic blood products. Ann Thorac Surg 2006; 81: 2367–2372. [23] Hill JD, O’Brien TG, Murray JJ, et al. Prolonged extracorporeal oxygenation for acute post-traumatic respiratory failure (shock-lung syndrome): use of the Bramson membrane lung. N Engl J Med 1972; 286: 629–634. [24] ISO 7199:2009 – Cardiovascular implants and artificial organs – blood-gas exchangers (oxygenators). ISO International Organization for Standardization, Geneva, April 2009. [25] Kato S, Morimoto S, Hiramitsu S, Nomura M, Ito T, Hishida H. Use of percutaneous cardiopulmonary support of patients with fulminant myocarditis and cardiogenic shock for improving prognosis. Am J Cardiol 1999; 83: 623–5, A10. [26] Kaufmann TAS, Hormes M, Laumen M, et al. Flow distribution during cardiopulmonary bypass in dependency on the outflow cannula positioning. Artif Organs 2009; 33: 988–992. [27] Kaufmann TAS, Wong KC, Schmitz-Rode T, Steinseifer U. Mimicking of cerebral autoregulation by flow-dependent cerebrovascular resistance: a feasibility study. Artif Organs 2012; 36: E97–101.
[28] Kawaguchi A, Bergsland J, Subramanian S. Total bloodless open heart surgery in the pediatric age group. Circulation 1984; 70: I30–I37. [29] Kawahito K, Murata S, Yasu T, et al. Usefulness of extracorporeal membrane oxygenation for treatment of fulminant myocarditis and circulatory collapse. Am J Cardiol 1998; 82: 910–911. [30] Kotani Y, Honjo O, Nakakura M, et al. Impact of miniaturization of cardiopulmonary bypass circuit on blood transfusion requirement in neonatal open-heart surgery. ASAIO J 2007; 53: 662–665. [31] Lan C, Tsai P-R, Chen Y-S, Ko W-J. Prognostic factors for adult patients receiving extracorporeal membrane oxygenation as mechanical circulatory support – a 14-year experience at a medical center. Artif Organs 2010; 34: E59–E64. [32] Lau CL, Posther KE, Stephenson GR, et al. Mini-circuit cardiopulmonary bypass with vacuum assisted venous drainage: feasibility of an asanguineous prime in the neonate. Perfusion 1999; 14: 389–396. [33] Merkle F, Boettcher W, Schulz F, Koster A, Huebler M, Hetzer R. Perfusion technique for nonhaemic cardiopulmonary bypass prime in neonates and infants under 6 kg body weight. Perfusion 2004; 19: 229–237. [34] Miyaji K, Kohira S, Miyamoto T, et al. Pediatric cardiac surgery without homologous blood transfusion, using a miniaturized bypass system in infants with lower body weight. J Thorac Cardiovasc Surg 2007; 134: 284–289. [35] Montoya JP, Shanley CJ, Merz SI, Bartlett RH. Plasma leakage through microporous membranes: role of phospholipids. ASAIO J 1992; 38: M399–M405. [36] Nakanishi K, Shichijo T, Shinkawa Y, et al. Usefulness of vacuum-assisted cardiopulmonary bypass circuit for pediatric open-heart surgery in reducing homologous blood transfusion. Eur J Cardiothorac Surg 2001; 20: 233–238. [37] Ota K. Advances in artificial lungs. J Artif Organs 2010; 13: 13–16. [38] Rahe-Meyer N, Solomon C, Tokuno M-L, et al. Comparative assessment of coagulation changes induced by two different types of heart-lung machine. Artif Organs 2010; 34: 3–12. [39] Reiss N, El Banayosy A, Posival H, Morshuis M, Minami K, Körfer R. Management of acute fulminant myocarditis using circulatory support systems. Artif Organs 1996; 20: 964–970. [40] Roche JK, Stengle JM. Open-heart surgery and the demand for blood. J Am Med Assoc 1973; 225: 1516–1521. [41] Schaible T, Hermle D, Loersch F, Demirakca S, Reinshagen K, Varnholt V. A 20-year experience on neonatal extracorporeal membrane oxygenation in a referral center. Intensive Care Med 2010; 36: 1229–1234. [42] Stein JI, Gombotz H, Rigler B, Metzler H, Suppan C, Beitzke A. Open heart surgery in children of Jehovah’s Witnesses: extreme hemodilution on cardiopulmonary bypass. Pediatr Cardiol 1991; 12: 170–174. [43] Ugaki S, Kasahara S, Kotani Y, et al. Extracorporeal membrane oxygenation following Norwood stage 1 procedures at a single institution. Artif Organs 2010; 34,: 898–903. [44] Ungerleider RM, Shen I. Optimizing response of the neonate and infant to cardiopulmonary bypass. Semin Thorac Cardiovasc Surg Pediatr Card Surg Annu 2003; 6: 140–146. [45] Willcox TW. Vacuum-assisted venous drainage: to air or not to air, that is the question. Has the bubble burst? J Extra Corpor Technol 2002; 34: 24–28.
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Georg Wagnera,*, Peter Schlansteina, Sandra Fiehe, Tim Kaufmann, Rüdger Kopp, Ralf Bensberg, Thomas Schmitz-Rode, Ulrich Steinseifer and Jutta Arens
A novel approach in extracorporeal circulation: individual, integrated, and interactive heart-lung assist (I3-Assist) Abstract: Extracorporeal life support (ECLS) is a wellestablished technique for the treatment of different cardiac and pulmonary diseases, e.g., congenital heart disease and acute respiratory distress syndrome. Additionally, severely ill patients who cannot be weaned from the heart-lung machine directly after surgery have to be put on ECLS for further therapy. Although both systems include identical components, a seamless transition is not possible yet. The adaption of the circuit to the patients’ size and demand is limited owing to the components available. The project I³-Assist aims at a novel concept for extracorporeal circulation. To better match the patient’s therapeutic demand of support, an individual number of one-size oxygenators and heat exchangers will be combined. A seamless transition between cardiopulmonary bypass and ECLS will be possible as well as the exchange of components during therapy to enhance circuit maintenance throughout long-term support. Until today, a novel oxygenator and heat exchanger along with a simplified manufacturing protocol have been established. The first layouts of the unit to allow the spill- and bubble-free connection and disconnection of modules as well as improved cannulas and a rotational pump are investigated using computational fluid dynamics. Tests were performed according to current guidelines in vitro and in vivo. The test results show the feasibility and potential of the concept. Keywords: ECLS; HLM; individual therapy; modular design; normative testing.
These authors contributed equally to this article. *Corresponding author: Dipl.-Ing. Georg Wagner, Department of Cardiovascular Engineering, Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, Pauwelsstraße 20, 52074 Aachen, Germany, Phone: +49 241 80 89 888, Fax: +49 241 80 82 144, E-mail: [email protected] Peter Schlanstein, Sandra Fiehe, Tim Kaufmann, Thomas Schmitz-Rode, Ulrich Steinseifer and Jutta Arens: Department of Cardiovascular Engineering, Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, 52074 Aachen, Germany a
Rüdger Kopp and Ralf Bensberg: Department of Intensive Care Medicine, University Hospital Aachen, RWTH Aachen University, 52074 Aachen, Germany
Background In extracorporeal circulation (ECC), two groups are distinguished [1, 7, 8]: 1. For short-term support of several hours during cardiac surgery, a heart-lung machine (HLM) or cardiopulmonary bypass (CPB) is used. The first use of a HLM in an adult patient was reported by John Gibbon in 1953 [19]. The need for blood transfusions was an issue right from the start of ECC, owing to infections and the amount of blood needed [40]. Although Beall et al. [9] reported successful open heart surgery without excessive use of blood transfusions in 1967, hemodilution remained unfeasible for small patients. During the next decades, blood-free surgery in neonatal and pediatric patients was achieved by using hypothermia and minimized circuits with reduced tubing length and diameter [16, 20, 28, 30, 42]. To further decrease the need of blood transfusions, vacuum-assisted venous drainage has been established, but results in undesirable side effects such as suction events or venous collapse [14, 22, 32, 33, 36, 44, 45]. 2. For long-term support from several hours up to several weeks, an extracorporeal life support (ECLS; i.e., heart and lung support) or an extracorporeal lung assist can be used. Even in the 1950s, researchers thought of a continuing support to allow for recovery of the heart and lung. It took another 20 years until the first ECLS was used successfully in 1972 [23]. Regarding ECLS, circuit-related complications become more likely the longer the support lasts [11, 18, 21, 31]. Although the exchange of components is difficult [31], preparations of the circuit can be made, e.g., a second line containing a primed oxygenator in case of
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126 G. Wagner et al.: Individual, integrated, and interactive assist mechanical failure of the oxygenator [1]. The critical parameters regarding the mortality rate of patients on ECLS are, among others, the comorbidities [6, 11, 18, 31], the time needed to initiate the circuit [1, 6, 10, 25, 29, 39, 41, 43], and the duration of support [21, 43]. The time needed to start the system can be reduced by the experience of the user and the design of the system [1, 11, 31, 41, 43], including the possibility of transport on a cart [31] or in a car [11] to get the ECC to the patient. Whereas HLM is used in the operating room to enable cardiac surgery, ECLS is used in a wide range of bridging techniques. Bridge-to-recovery includes patients with possible reversible indications such as myocarditis or cardiogenic shock during or after surgery; bridge-to-bridge comprises all patients waiting for a permanent assist device, while bridge-to-transplant consists of patients awaiting organ transplantation [1, 6, 11, 21, 31, 37, 41]. A HLM usually consists of a venous reservoir, a pump, an oxygenator with integrated heat exchanger, and the tubing. An ECLS circuit contains a pump, an oxygenator, and tubing. Although the type of pump and oxygenator are usually different in both systems, essentially identical components for the most parts are applied. However, there is no system yet available that allows switching between the two modes in order to reduce the trauma to the patients, e.g., in those patients who cannot be weaned from HLM after surgery [1, 6]. The priming volume, owing to the dimensions of the components (pump, oxygenator, reservoir, tubing), causes hemodilution [16, 30, 38]. The surface area in contact with blood, resulting mostly from the membrane but also from the dimensions of the components, causes an inflammatory response and the activation of clotting
[16, 18, 37]. Clotting is treated with anticoagulants applied as a coating of the components before use or injected directly into the circuit [1, 6, 16, 18, 37]. At the cannulation site, identified as the major contributor to the clotting activation, this anticoagulation can lead to bleeding complications [37, 38, 43]. Although a wide range of oxygenators, reservoirs, and other circuit components are available on the market, they can be divided into the three general size ranges (pediatric, adolescent, and adult) according to their performance regarding maximum blood flow and gas exchange rates. In order not to operate the ECC beyond its range, the next larger components are chosen rather than the small device. Thereby, the priming volume and surface area of the ECC are exceeding the therapeutic need in most cases as the continuous body weight range from neonates to adults has to be covered with the discrete sizes available. This results in the circuit exceeding the actual therapeutic demand, thus increasing priming volume and artificial surface area and, thereby, the rate of complications. Numerous publications report a reduced length of tubing in order to create a smaller circuit; it appears to be common practice to reduce the circuit-related side effects [30, 33, 34, 38, 44]. To address these issues, the project I³-Assist aims at developing a highly integrated and modular extracorporeal system that can be adapted to the changing individual therapy of a patient. The key aspect of the I3-Assist system is the modularity. This is being achieved by fragmentation of the functional units, heat exchanger and oxygenator, into small modules with one size only (see schematic drawings in Figure 1). An optimized gas and heat exchange surface area and thus a minimized priming volume of the system are therefore realized. During the assembly of the system, the modules can be combined by the user to
Figure 1 Conceptual idea of the project. ECLS, extracorporeal life support; HLM, heart-lung machine; Oxy, oxygenator; HE, heat exchanger.
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accomplish the suitable exchange rate, e.g., one oxygenator module for a neonate, five oxygenator modules for a small adult, or eight oxygenator modules for a tall adult. All components will be assembled before the surgery by the perfusionist on a connecting unit. The oxygenator and heat exchanger modules of the HLM/ECLS system will be exchangeable under operating conditions. Thus, defective components can be exchanged easily. Additionally, an immediate and seamless transition between operation modes can be carried out and the system can be modified according to changing individual needs during surgery and therapy. Owing to the highly integrated design, the system can be placed near the operating table and can be used for inter- and intrahospital transport. This article gives an overview of the status quo of the ongoing research project and further steps planned.
Methods The project is carried out in a partnership of four parties: –– The Department of Intensive Care Medicine of the University Hospital RWTH Aachen as a clinical partner and provider of the users’ perspective; –– The qcmed GmbH (Aachen, Germany) as a consultant to implement a quality management (QM) system that suits all partners’ needs and enables an easy technology transfer between the partners; –– A company with manifold experiences in the manufacturing and distribution of HLM and ECLS devices; and –– The Department of Cardiovascular Engineering (CVE), Institute of Applied Medical Engineering (AME), University Hospital RWTH Aachen as a research facility. Primarily at the start, the project highly profited from the vicinity of the partners’ location, as strategy and engineering workshops as well as risk analysis meetings were frequent. The communication plan and the assignment of tasks had to be defined to put all competences to optimal use; the short traveling distance was of great benefit. During the requirement specification process, a basic draft listing all functional units was developed. These functional units are (in the order of blood flow): cannulas, reservoir, pump, heat exchanger, oxygenator, and a connecting unit, like a mounting plate to link all units. The research, as well as the construction and development, of these components were divided between the company and the CVE/AME.
An online tool enabled access to the latest versions of strategic documents, such as URS, DRS, and RIA, for all partners at any time. For details concerning the quality management within this project, please be referred to the article “Implementation of Quality Management in Early Stages of Research and Development Projects at a University” within this journal [17]. The general specifications were defined by all parties. The overall system dimensions should not exceed 0.15 m³, whereas each disposable component should be able to fit in a standard hospital bin. On the basis of input from the user, the minimal circuit containing one oxygenator to support one infant should allow for a nominal blood flow of 800 ml/min. The pump should provide a flow range from 0 up to 8000 ml/min and be able to create a pressure of up to 650 mm Hg at maximum flow. As CAD software ProEngineer Wildfire 5.0 (Parametric Technology Cooperation, Needham, MA, USA) was used by the CVE/AME, the company used different software. As the different CAD systems were well established in each party, we refrained from establishing a corporate system. For file sharing between research, development, and manufacturing, the platform-independent file formats IGES and STEP were used. We encountered no problems during the exchange and further processing of the files emanating from the different software solutions used. The test parameters for oxygenators defined in DIN EN 12022 [13] and ISO 7199 [24] allow for a wide range of input settings. Depending on the chosen setting within this range, the performance of the same oxygenator may vary in an even wider range, due to the non-linear behavior of oxygen binding. Blood is used as test fluid for performance and hemolysis tests, which differs on a daily basis even within the same species and when collected from more than one specimen. The quality is highly dependent on proper handling during collection and transport. Therefore, performing comparable hemolysis test for several design variants requires performing the test on all designs on the same day and with the same blood. To reduce the need for blood, facilitate the handling, and create results that are reliably comparable with each other, blood replacement fluids are used frequently at our institute when applicable (e.g., for hydraulic tests).
Results An initial feasibility study was performed using a Medos Hilite hard-shell reservoir for children (Medos Medizintechnik AG, Stolberg, Germany) as fluid capacity, Hilite
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128 G. Wagner et al.: Individual, integrated, and interactive assist 800 LT oxygenators, and a Medos deltastream DP3 blood pump with DC driving unit. A HAES mixture (Fresenius Kabi Deutschland GmbH, Bad Homburg, Germany) with a dynamic viscosity of 3.63 mPas at 20°C was used as blood substitute. This study was conducted to show the feasibility of a passive flow distribution using three parallel lines, as shown in Figure 2. One design for both the dividing and the collection unit were defined as spherical sectors. The setup enabled the quantification of the flow distri bution within the three lines and a study of the flow distri bution with one line being obstructed. At the nominal flow, the oxygenators developed a mean pressure loss of 33 ± 2 mm Hg per line. The 1/4-in tubing within each line created a pressure loss equal to 60% of the oxygenator. Each line was equally exposed to the same flow within a range of ± 3%. The results suggested that a uniform passive flow distribution was nearly independent of the relative arrangement of the lines. The flow resistance of heat exchangers and oxygenators supported the uniformity by reducing the influence of fluid-wall interactions, which increase with longer flow paths.
Construction and design The manufacturing of those laboratory models that needed conventional machining was done in the CVE/ AME’s workshop. Immersion processes were performed by the company involved. Early testing of concepts and variations was elementary during the design process. Therefore, rapid prototyping was used frequently especially at the project start and was performed in the CVE/ AME’s workshop using Full Cure 720 material (Objet
Figure 2 Initial feasibility study of passive flow deviation.
Geometries Ltd., Rehovot, Israel). This allowed fast comparisons between different design options. As advantageous as a three-dimensional (3D) printer may be, there are a few limitations regarding the system we used. The material used is not suitable for contact with blood; it macerates in contact with water and shows plastic flow even at temperatures below 50°C. Although outer and inner structures can be created without the limitations of conventional machining, the partly required postprocessing of oriented parts is restricting the creative scope for development. Additionally, the costs are still considerably high; in some cases, conventional manufacturing is less expensive. However, it is beneficial that it takes only a couple of hours to print a component.
Cannulas The key aspects regarding the cannula development were minimizing the trauma at the cannulation sites, reducing risk of suction events at the withdrawal cannula, and an improving hemodynamic performance along with an optimized flow pattern at the return cannula. A jet is usually produced at the outflow site. The jet’s impact may loosen plaques, damage the aortic wall, and lead to low perfusion of the upper extremities [26]. Design studies on the return cannula were performed, investigating the effects of outer and inner structures on the expansion of the jet. The hemolytic potential of the requisite design elements has to be investigated closely. After hydraulic tests with several new design options, hemolysis test were performed with two different designs. In vitro testing was performed according to available standards and guidelines (such as ISO 7199 [24] or ASTM F 1830 [5]). Comprehensive tests regarding the hydraulic performance were conducted using the laboratory models with both blood substitute and porcine blood. Porcine blood was collected from an abattoir in a 5-l canister that was prepared with 75,000 IU heparin-sodium, 8 ml of 50% glucose solution, and 60 ml of 0.9% saline solution. A hemolysis test for one new cannula design with corresponding control circuits was performed. A 24-Fr Medos arterial cannula (Medos Medizintechnik AG, Stolberg, Germany) with an angled tip was used as a reference device. A circuit without cannula was used as a reference circuit. Each circuit contained the same kind and length of tubing, reservoir, and roller pump. The occlusion and rotational speed of the roller pumps were calibrated to ensure comparable settings. The results showed a higher hemolysis rate in one design; further investigations were performed using computational fluid dynamics (CFD)
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simulations as more design options could be investigated and compared more easily in a shorter period. Design features cannot be displayed to not affect the patent process. CFD simulations of blood flow deviation in the aortic arch dependent on different cannula designs were performed using a method that was established and experimentally validated in previous studies [26]. 3D models of different cannula tip designs were incorporated in a CAD model of the human aortic arch, including the three supraaortal vessels. The model was meshed with unstructured tetrahedral elements. Even though mesh independency was already achieved at 100,000 elements, a fine mesh density of 2.7 million elements was chosen to provide a higher spatial resolution of the final results. Furthermore, five layers of prismatic elements were generated to resolve boundary layers and to more accurately calculate wall shear stress. Flow rates of 2000, 4000, and 6000 ml/min were simulated using a non-Newtonian blood model with a hydrodynamic viscosity threshold value of 3.6 mPas for high shear rates. Arterial outlets were modeled as static pressure openings at 55 mm Hg. Additionally, a pressure drop due to systemic resistance was modeled using the Darcy-Weißbach equation. 1 ∆P = L ρU n2 2 In this equation, ρ is the fluid density and Un is the normal velocity at each boundary. L is defined as a dimensionless loss coefficient and set to 50 based on a sensitivity study [27]. Investigations of the cannula comparing the current design and the new designs were conducted. The results for a flow rate of 6 l/min and a constant aortic pressure
showed an increase of 9.5% in shear rate for new tip #1, whereas both others performed similar. The maximum pressure on the aortic wall was reduced by 4.9% in new tip #2 and by 22.8% in new tip #1 compared to the standard tip. A comparison of new tip #1 and the conventional tip is shown in Figure 3; the improved stress level on the aortic wall is achieved by a smoothed flow pattern with good expansion of the jet. Owing to the new fluid path, the pressure loss over the cannula is slightly increased.
Reservoir At first, the objectives of the new reservoir were to combine the benefits of a hard-shell reservoir with those of a soft venous blood bag. The hard shell allows a defined position and for venous drainage when sealed airtight. The bag’s ability to inhibit blood-air contact, a major trigger of coagulation, is advantageous. In vitro tests with the first prototypes showed that the available elastic materials were not suitable for the required performance; the elastic deformation was exceeded, disabling the return to the initial shape. Therefore, a reassessment of the requirements was necessary. A survey performed among perfusionists revealed a bias toward hard-shell reservoirs. A new design is currently being developed.
Pump A rotational pump is advantageous compared with a roller pump, especially for long-term applications, as the hemolytic rate is considerably lower [1] and no electrostatic charge is developed. Regarding the transport in and between hospitals, roller pumps are considerably heavier
Figure 3 Pressure loss and shear stresses induced to the aortic wall: comparison of a current tip (left) and new design #1 (right).
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than rotational pumps. Additionally, the priming volume is higher because of the tubing length required inside the roller pump [2, 38]. The wide range of flow rate results from the modularity of the system, as the performance should be good for low and for high flow rates. The first design, generated with the support of CFD simulation, is currently being improved. The hydraulic performance matched the expectations; however, once a larger amount of air had entered the pump chamber, the hydraulic performance dropped drastically. Furthermore, the air was removed very slowly, thereby reducing the blood flow for a longer period.
Oxygenator Two types of oxygenators can be distinguished by their manufacturing process: those that contain a coiled fiber bundle and those that contain a layered fiber bundle. In our experiments, we found that a coiled oxygenator usually shows a higher pressure loss, as the flow is deflected more often to ensure a counter flow of blood versus gas and water. A layered oxygenator is usually easier to prime, owing to the less complicated structure, but shows less uniform flow deviation when used at lower flow ranges. Another distinction can be made by the hollow fibers used; fibers used in HLM are not plasmaproof over periods of more than 12 h [15, 35]; those fibers used in ECLS last up to 7 days. This results in the limited working time of the oxygenators. A third type, silicone membranes, were used until recently, especially in the United States, but are now being replaced successively because of their comparatively high pressure loss and lower gas transfer rates. Regardless of the inner structure, the gas exchange rate within an oxygenator should be ≥ 55 ml/min of oxygen and ≥ 50 ml/min of carbon dioxide per liter blood at all times. The development of the one-size oxygenator within this project was driven by the therapeutic need of the smallest patients. The requested nominal blood flow of 800 ml/min results in a gas exchange area of approximately 0.35 m2, as shown in a comparison of available oxygenators in Figure 4, extracted from publicly available data sheets (e.g., Medos Medizintechnik AG, Stolberg, Germany; Sorin AG, Milan, Italy; Maquet GmbH, Rastatt, Germany; Terumo Corporation, Tokyo, Japan). The first rapid prototyped models of the oxygenator modules consisted of a coiled fiber bundle and simple housing parts. The device was slim and tall, with the outer diameter of the fiber bundle being 56 mm and the length being 60 mm. The geometry arose partly from
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Figure 4 Comparison of available oxygenators regarding surface area and blood flow.
the comparative sample we used, the Medos Hilite 800 LT (Medos Medizintechnik AG, Stolberg, Germany), and partly from the preceding research projects at the CVE/ AME, the NeonatOx and the MiniHLM [3, 4]. The overall dimensions of the whole system defined in URS and DRS compelled a revision of this concept. The current patent pending version contains the same surface area but it is more compact. It contains a layered bundle with an outer diameter of 50 mm and a height of 30 mm. A special focus was an easy manufacturing process, on the one hand to enable the assembling in our laboratory, and on the other hand to reduce production costs later on. The latter will reduce costs for the user as well, regarding procurement prices and inventory management. The number of housing parts could be reduced to three, two of which are identical for heat exchanger and oxygenator. The potting, a process by which fibers and housing are joined using a glue and the separation of blood and fluid is ensured, was performed following a simple protocol. Silicone is used as potting material by the CVE so that the custommade housings can be reused. A custom-made centrifuge, including controller, was developed and built for the new potting process of oxygenator and heat exchanger modules. Simulations of the oxygenator were performed regarding the pressure loss, the fluid path, and stagnation areas. An isotropic model as well as an anisotropic model, to account for the direction of the fibers, was used to embed the permeability of the bundle. Different inlet geometries for the latest oxygenator were investigated, supporting an optimized flow distribution. As there is no powerful simulation tool yet to calculate the efficacy of construction elements and design changes regarding the gas exchange
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capability, detailed tests leading to design and test iterations were expected and necessary. Regarding in vitro testing, a new test setup was implemented to conduct long-term performance tests up to 12 h on the oxygenators’ gas exchange efficiency. A custommade double-reservoir allows for connecting two circuits in a cross-flow manner. One is the test circuit for the oxygenator, the other one is a deoxygenation circuit, an oxygen consumer, and carbon dioxide producer, which is used to maintain venous parameters for the blood entering the test circuit. All tests were run according to the newly developed test records, which enabled a sound comparison of different test objects. Five oxygenators of the latest design showed an average performance factor of 96% of the required oxygen transfer and 107% of the required CO2 removal. To further optimize the design of oxygenators, better simulation tools are needed. The flow patterns inside a fiber bundle cannot be described with numerical methods yet, at least not on a large scale. The current approach of an anisotropic model needs further improvement to support the design process. In addition, a more reliable prediction of the hemolytic potential using simulations would be helpful. Thus far, this can only be done qualitatively by comparison of different designs. Accurate calculation of the normalized index of hemolysis (NIH) is not yet possible but would be a powerful tool to support the design process. Regarding the design of the oxygenator, the fibers are the limiting part in current designs. The way of processing required, the gas exchange efficiency, and the longterm durability are constraints in new developments. New materials would greatly contribute to further progress toward long-term stable and further miniaturized oxygenators.
Heat exchanger Commercially available oxygenators usually contain a heat exchanger; the modularization allows separate modules for both. The heat exchanger and oxygenator should share identical housing parts as many as possible to minimize production costs. Additionally, their performances have to match each other. Therefore, the heat exchangers’ development was postponed until construction and design of the oxygenator was completed. The manufacturing protocol is identical to the oxygenators’, except for the fibers used. The heat exchanger performance factor in vitro was R = 0.8, calculated according to ISO 7199 [24]:
R=
TBlood ,out -TBlood ,in TWater ,in -TBlood ,in
Our research examining publically available data sheets showed that commercially devices have factors between 0.63 and 0.8. The modularization of the heat exchanger and oxygenator creates a high grade of flexibility regarding the adaption of the circuit during therapy. The surface area of materials in contact with the patients’ blood can be reduced to a minimum at all times, but the priming volume of the system is increased by the number of modules. The youngest patients could most likely benefit more from a further decreased priming volume. Therefore, the separation of heat exchanger and oxygenator should be thoroughly surveyed.
Connecting unit All fluids are channeled by the connecting unit, which should allow for attachment and removal of the modules. Connecting modules should be possible without leakage of blood, water, or gas into or out of the circuit. Furthermore, a hemocompatible flow without stagnation areas, swirls, or interfering details must be provided. Additionally, priming of the blood paths and general handling should be easy when operating all fluid channels of a module at the same time. The channel system, replacing the tubing used to date, is designed with all fluid paths separated and all necessary connections. Simulations will be used to investigate the flow distribution within the channels. By now, we have designed a laboratory sample of a connector fulfilling the requirements outlined above, which is currently being manufactured in the workshop and subsequently will be thoroughly tested.
In vivo testing The in vivo testing was authorized by the government of the state of North Rhine-Westphalia, Germany. All animals received medical care according to the German guidelines for laboratory animal care, which are in line with the National Institutes of Health guidelines. On the basis of the test records for the in vitro tests, protocols for in vivo tests have been developed. A novel animal model was developed by the clinical partners in order to mimic the transition between HLM and ECLS use of the devices during
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132 G. Wagner et al.: Individual, integrated, and interactive assist surgery and subsequent intensive care. The novel protocol was validated before the final device testing using commercially available components in two acute (short-term) tests in a pig model. Reversible hypoxia and hypercapnia were induced by mechanical ventilation with decreased oxygen percentage and increased level of carbon dioxide. This first validation of the novel animal testing protocol showed good results; the animals were hydrodynamically stable and could be supported according to the varying levels of hypoxia as the blood gas levels were maintained within a normal range by the circuit.
Discussion and future prospects Today, users are preparing their circuits to exchange components more easily in case mechanical failure of a device occurs during operation [1, 31]. The demand to extend the use of ECLS by adding new functions to it, e.g., a possibility to run renal replacement therapy in parallel or by offering customized support solutions, is expressed in literature. The hopes are that such systems could save lives and costs while sparing resources at the same time [11, 12, 21, 30, 37]. The chance to develop a whole new ECC with all its components and search for integrative potential enabled new ways in the construction and design of these components, which would not have been seen in separate research of each component. The close cooperation of the partners and their will to combine their diverse expertise was beneficial to establish a fast and productive communication. This was notably effective during the start of the project. The project is still ongoing. The feasibility of the concept was already shown in vitro as well as in vivo as most of the components are available as laboratory
models. Some components required more research and development as expected during planning, e.g., the connecting unit, particularly the connector itself. Further tests of the performance and hemolytic behavior of the components are going to be performed at the CVE/AME. In vivo tests are planned to validate the overall design of the system. The transfer into industrially produced prototypes will be the next step after the in vivo validation of the design in long-term-use. A potential field of research could be the level of oxygen saturation created by the ECC. Clinicians reported weaning from ECLS when a saturation of 65% could be established. If the required performance is able to be reduced from the current demand of a saturation of 100%, the dimensions of the device, the surface area, the priming volume, and the durability of the fibers could be improved. This would probably reduce the demand on anticoagulants as well. Acknowledgments: This project was performed in collaboration between the industrial partner, the QM consultant qcmed GmbH, the Department of Intensive Care Medicine of the University Hospital RWTH Aachen, and the Department of Cardiovascular Engineering (CVE), Institute of Applied Medical Engineering (AME), University Hospital RWTH Aachen. The project is co-funded by the European Union (ERDF – European Regional Development Fund – Investing in your future) and the German federal state North Rhine-Westphalia (NRW), under the operational program Ziel2 “Regional Competitiveness and Employment” 2007–2013 (EFRE). Received March 28, 2013; accepted November 6, 2013; online first December 11, 2013
References [1] Allen S, Holena D, McCunn M, Kohl B, Sarani B. A review of the fundamental principles and evidence base in the use of extracorporeal membrane oxygenation (ECMO) in critically ill adult patients. J Intensive Care Med 2011; 26: 13–26. [2] Ando M, Takahashi Y, Suzuki N. Open heart surgery for small children without homologous blood transfusion by using remote pump head system. Ann Thorac Surg 2004; 78: 1717–1722. [3] Arens J, Schoberer M, Lohr A, et al. NeonatOx: a pumpless extracorporeal lung support for premature neonates. Artif Organs 2011; 35: 997–1001. [4] Arens J, Schnoering H, Pfennig M, et al. The Aachen MiniHLM – a miniaturized heart-lung machine for neonates with an integrated rotary blood pump. Artif Organs 2010; 34: 707–713.
[5] ASTM. ASTM F 1830:1997 – standard practice for selection of blood for in vitro evaluation of blood pumps, 1997. [6] Bakhtiary F, Keller H, Dogan S, et al. Venoarterial extracorporeal membrane oxygenation for treatment of cardiogenic shock: clinical experiences in 45 adult patients. J Thorac Cardiovasc Surg 2008; 135: 382–388. [7] Bartlett RH. Extracorporeal life support in the management of severe respiratory failure. Clin Chest Med 2000; 21: 555–561. [8] Bartlett RH, Roloff DW, Cornell RG, Andrews AF, Dillon PW, Zwischenberger JB. Extracorporeal circulation in neonatal respiratory failure: a prospective randomized study. Pediatrics 1985; 76: 479–487.
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G. Wagner et al.: Individual, integrated, and interactive assist 133 [9] Beall A Jr, Yow E Jr, Bloodwell RD, Hallman GL, Cooley DA. Open heart surgery without blood transfusion. Arch Surg 1967; 94: 567–570. [10] Chen Y-S, Yu H-Y, Huang S-C, et al. Experience and result of extracorporeal membrane oxygenation in treating fulminant myocarditis with shock: what mechanical support should be considered first? J Heart Lung Transplant 2005; 24: 81–87. [11] Combes A, Leprince P, Luyt C-E, et al. Outcomes and long-term quality-of-life of patients supported by extracorporeal membrane oxygenation for refractory cardiogenic shock. Crit Care Med 2008; 36: 1404–1411. [12] Cruz D, Bellomo R, Kellum JA, de Cal M, Ronco C. The future of extracorporeal support. Crit Care Med 2008; 36: S243–S252. [13] DIN EN 12022:1999 – Blutgasaustauscher. DIN Deutsches Institut für Normung e.V., Berlin: Beuth Verlag, März 1999. [14] Durandy Y. Blood transfusion in pediatric cardiac surgery. Artif Organs 2010; 34: 1057–1061. [15] Eash HJ, Jones HM, Hattler BG, Federspiel WJ. Evaluation of plasma resistant hollow fiber membranes for artificial lungs. ASAIO J 2004; 50: 491–497. [16] Edmunds LH. Advances in the heart-lung machine after John and Mary Gibbon. Ann Thorac Surg 2003; 76: S2220–S2223. [17] Fiehe S, Wagner G, Schlanstein P, et al. Implementation of quality management in early stages of research and development projects at a university. Biomed Eng/Biomed Technik 2014; 59: 135–145. [18] Fleming GM, Gurney JG, Donohue JE, Remenapp RT, Annich GM. Mechanical component failures in 28,171 neonatal and pediatric extracorporeal membrane oxygenation courses from 1987 to 2006. Pediatr Crit Care Med 2009; 10: 439–444. [19] Gibbon J Jr. Application of a mechanical heart and lung apparatus to cardiac surgery. Minn Med 1954; 37: 171–185; passim. [20] Golab HD, Takkenberg JJ, Bogers AJJC. Specific requirements for bloodless cardiopulmonary bypass in neonates and infants; a review. Perfusion 2010; 25: 237–243. [21] Haines NM, Rycus PT, Zwischenberger JB, Bartlett RH, Undar A. Extracorporeal life support registry report 2008: neonatal and pediatric cardiac cases. ASAIO J 2009; 55: 111–116. [22] Hickey E, Karamlou T, You J, Ungerleider R. Effects of circuit miniaturization in reducing inflammatory response to infant cardiopulmonary bypass by elimination of allogeneic blood products. Ann Thorac Surg 2006; 81: 2367–2372. [23] Hill JD, O’Brien TG, Murray JJ, et al. Prolonged extracorporeal oxygenation for acute post-traumatic respiratory failure (shock-lung syndrome): use of the Bramson membrane lung. N Engl J Med 1972; 286: 629–634. [24] ISO 7199:2009 – Cardiovascular implants and artificial organs – blood-gas exchangers (oxygenators). ISO International Organization for Standardization, Geneva, April 2009. [25] Kato S, Morimoto S, Hiramitsu S, Nomura M, Ito T, Hishida H. Use of percutaneous cardiopulmonary support of patients with fulminant myocarditis and cardiogenic shock for improving prognosis. Am J Cardiol 1999; 83: 623–5, A10. [26] Kaufmann TAS, Hormes M, Laumen M, et al. Flow distribution during cardiopulmonary bypass in dependency on the outflow cannula positioning. Artif Organs 2009; 33: 988–992. [27] Kaufmann TAS, Wong KC, Schmitz-Rode T, Steinseifer U. Mimicking of cerebral autoregulation by flow-dependent cerebrovascular resistance: a feasibility study. Artif Organs 2012; 36: E97–101.
[28] Kawaguchi A, Bergsland J, Subramanian S. Total bloodless open heart surgery in the pediatric age group. Circulation 1984; 70: I30–I37. [29] Kawahito K, Murata S, Yasu T, et al. Usefulness of extracorporeal membrane oxygenation for treatment of fulminant myocarditis and circulatory collapse. Am J Cardiol 1998; 82: 910–911. [30] Kotani Y, Honjo O, Nakakura M, et al. Impact of miniaturization of cardiopulmonary bypass circuit on blood transfusion requirement in neonatal open-heart surgery. ASAIO J 2007; 53: 662–665. [31] Lan C, Tsai P-R, Chen Y-S, Ko W-J. Prognostic factors for adult patients receiving extracorporeal membrane oxygenation as mechanical circulatory support – a 14-year experience at a medical center. Artif Organs 2010; 34: E59–E64. [32] Lau CL, Posther KE, Stephenson GR, et al. Mini-circuit cardiopulmonary bypass with vacuum assisted venous drainage: feasibility of an asanguineous prime in the neonate. Perfusion 1999; 14: 389–396. [33] Merkle F, Boettcher W, Schulz F, Koster A, Huebler M, Hetzer R. Perfusion technique for nonhaemic cardiopulmonary bypass prime in neonates and infants under 6 kg body weight. Perfusion 2004; 19: 229–237. [34] Miyaji K, Kohira S, Miyamoto T, et al. Pediatric cardiac surgery without homologous blood transfusion, using a miniaturized bypass system in infants with lower body weight. J Thorac Cardiovasc Surg 2007; 134: 284–289. [35] Montoya JP, Shanley CJ, Merz SI, Bartlett RH. Plasma leakage through microporous membranes: role of phospholipids. ASAIO J 1992; 38: M399–M405. [36] Nakanishi K, Shichijo T, Shinkawa Y, et al. Usefulness of vacuum-assisted cardiopulmonary bypass circuit for pediatric open-heart surgery in reducing homologous blood transfusion. Eur J Cardiothorac Surg 2001; 20: 233–238. [37] Ota K. Advances in artificial lungs. J Artif Organs 2010; 13: 13–16. [38] Rahe-Meyer N, Solomon C, Tokuno M-L, et al. Comparative assessment of coagulation changes induced by two different types of heart-lung machine. Artif Organs 2010; 34: 3–12. [39] Reiss N, El Banayosy A, Posival H, Morshuis M, Minami K, Körfer R. Management of acute fulminant myocarditis using circulatory support systems. Artif Organs 1996; 20: 964–970. [40] Roche JK, Stengle JM. Open-heart surgery and the demand for blood. J Am Med Assoc 1973; 225: 1516–1521. [41] Schaible T, Hermle D, Loersch F, Demirakca S, Reinshagen K, Varnholt V. A 20-year experience on neonatal extracorporeal membrane oxygenation in a referral center. Intensive Care Med 2010; 36: 1229–1234. [42] Stein JI, Gombotz H, Rigler B, Metzler H, Suppan C, Beitzke A. Open heart surgery in children of Jehovah’s Witnesses: extreme hemodilution on cardiopulmonary bypass. Pediatr Cardiol 1991; 12: 170–174. [43] Ugaki S, Kasahara S, Kotani Y, et al. Extracorporeal membrane oxygenation following Norwood stage 1 procedures at a single institution. Artif Organs 2010; 34,: 898–903. [44] Ungerleider RM, Shen I. Optimizing response of the neonate and infant to cardiopulmonary bypass. Semin Thorac Cardiovasc Surg Pediatr Card Surg Annu 2003; 6: 140–146. [45] Willcox TW. Vacuum-assisted venous drainage: to air or not to air, that is the question. Has the bubble burst? J Extra Corpor Technol 2002; 34: 24–28.
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