Centrifugal Pump Clinic

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Thispaper describes the single passage CFD model based to analyze the flow in blade passages of a centrifugal pump. The model consists of the flow passage between two impeller blades and the spaces in the inlet eye as well as in the volute. The incompressible Navier-Stokes equations in the conservation form are solved by a finite difference method. The code is designed to investigate the velocity and pressure distributions and intended to investigate how the pump design fluid flow through the rotor as well as the pump performance. An early part of the paper investigates the behavior of the model as well as validity of the assumptions made in the model. Then, applications to a rotodynamic heart pump are presented.

N2 - This paper describes the single passage CFD model based to analyze the flow in blade passages of a centrifugal pump. The model consists of the flow passage between two impeller blades and the spaces in the inlet eye as well as in the volute. The incompressible Navier-Stokes equations in the conservation form are solved by a finite difference method. The code is designed to investigate the velocity and pressure distributions and intended to investigate how the pump design fluid flow through the rotor as well as the pump performance. An early part of the paper investigates the behavior of the model as well as validity of the assumptions made in the model. Then, applications to a rotodynamic heart pump are presented.

AB - This paper describes the single passage CFD model based to analyze the flow in blade passages of a centrifugal pump. The model consists of the flow passage between two impeller blades and the spaces in the inlet eye as well as in the volute. The incompressible Navier-Stokes equations in the conservation form are solved by a finite difference method. The code is designed to investigate the velocity and pressure distributions and intended to investigate how the pump design fluid flow through the rotor as well as the pump performance. An early part of the paper investigates the behavior of the model as well as validity of the assumptions made in the model. Then, applications to a rotodynamic heart pump are presented.

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The innovation of coronary-bypass surgery made Cleveland Clinic a world-recognized center in cardiac surgery by the mid-1970s. Floyd Loop, MD, who became Chairman of Thoracic and Cardiovascular Surgery in 1975, continued the development of coronary surgery and greatly enhanced the academic profile of Cleveland Clinic. Dr. Loop began to recruit physicians nationwide, including Toby Cosgrove, MD, from Massachusetts General Hospital. Also during the 1970s, Cleveland Clinic began attracting increasing numbers of international patients, particularly from the Middle East, including Crown Prince Khalid of Saudi Arabia, who came to Cleveland Clinic for coronary artery bypass surgery in 1972.

In 1971, Floyd Loop, MD, identifies the mammary artery as the preferred conduit for bypass surgery. As a result, bypass surgery undergoes one of its most important refinements, as the mammary artery produces better results than using a vein from the leg.

Thomas Meaney, MD, Chair of the Division of Radiology from the 1960s to the 1980s, helped Cleveland Clinic acquire a CT scanner in 1972, which at that time, was one of only four in existence. This was strictly a neruoimager. Ten months later, Cleveland Clinic acquired a CT scanner for body imaging. Dr. Meany urged Cleveland Clinic to buy a pioneering MRI machine in 1983, and Cleveland Clinic radiologists led the way in developing the technology.

In 1979, Leonard Golding, MD, successfully installs a left ventricular assist device (LVAD) with a centrifugal pump, used to bridge patients until a donor heart becomes available or, increasingly, as a long-term alternative to transplantation.

Recently three different neonatal extracorporeal membrane oxygenation (ECMO) circuits have been employed in our clinic. These circuits were compared for clotting and bleeding complications. Initially, we used an ECMO circuit containing a roller pump and venous bladder without severe complications. Manufacturing of circuit components was discontinued, necessitating the replacement of this circuit by a circuit with a centrifugal pump with 3/8 inch inlet and outlet. Acute increase of oxygenator resistance requiring emergency changeout became unexpectedly a regularly occurring complication. The increase in resistance was suspected to be caused by oxygenator clotting, although oxygenator function was preserved. To prevent this complication, we changed to a levitating centrifugal pump with 1/4 inch inlet and outlet, after which no oxygenator malfunction has been observed. Macroscopic and electron microscopic analysis demonstrates that small clots are formed within the circuit, presumably in or near the centrifugal pump, which are transported to the oxygenator and clog up the hollow fiber layer at the inlet side, barely penetrating the oxygenator beyond this first layer. Our results suggest that low blood velocities accompanied with recirculation of blood within or near the centrifugal pump and/or heat generation within the pump could contribute to the formation of these clots.

Ed H. Edwards is the vice president for HBE Engineering of Three Rivers, Michigan, a supplier of centrifugal pump minimum flow valves and orifices. He has worked in the pump and related equipment field since 1972 and has a B.S. from Western Michigan University.

The development of ventricular assist devices (VADs) over the past 5 decades as therapy for advanced heart failure (HF) has been extraordinary. Since the original VAD design by Michael DeBakey in the early 1960s, numerous devices for mechanical circulatory support have been engineered, assessed in preclinical studies, applied to human patients in large multicenter clinical trials, and now, select devices are Food and Drug Administration-approved therapy for advanced HF patients. This review highlights select examples of durable VADs from the engineering aspect of design and conception to experimental studies and clinical application underscoring the remarkable progression of such technology to now becoming the standard of care for many advanced HF patients.

The MCS era began with the founding of the Artificial Heart Program in 1964 at the National Institutes of Health. Michael DeBakey was the first to develop the original rudimentary LVAD prototypes and, in 1966, reported the first successful use of the LVAD in a young woman unable to be weaned from cardiopulmonary bypass. Since that time, the era of MCS ushered in the development of various support devices that entered laboratory and preclinical testing to improve on the initial LVAD designs. In 1994, the Food and Drug Administration (FDA) approved the pneumatically driven LVAD as bridge to transplant (BTT) and, in 1998, approved the self-contained, vented electric LVAD devices for the same indication. These first-generation LVADs have been shown to improve end-organ function, optimize hemodynamics, augment functional capacity and improve the quality of life (QOL) of end-stage HF patients awaiting heart transplantation with an acceptably low side effect profile [4, 5, 6, 7, 8, 9]. Figure 1 shows a representative first-generation LVAD with the device implanted in the abdominal cavity below the diaphragm, with the inflow cannula positioned in the left ventricular apex drawing blood from the left ventricle and the outflow cannula anastomosed to the ascending aorta [10]. There is a driveline cable that exits the skin and is attached to an external controller that is powered by portable batteries. This first-generation device is now essentially obsolete.

The landmark REMATCH trial (Randomized Evaluation of Mechanical Assistance for the Treatment of Congestive Heart Failure) assessed the outcomes of end-stage HF patients who were ineligible for heart transplantation that were randomized to medical management or implantation of the HeartMate vented electric LVAD (Thoratec Corporation, Pleasanton, Calif., USA) [10]. The investigators demonstrated a marked reduction of 48% in the risk of death from any cause in the group randomized to the LVAD arm with a 1-year survival rate of 52% in the device group compared to 25% in the medical therapy group. However, despite these results, this technology was not initially widely embraced by the medical community due to concerns regarding the large size of the pump, limited durability of the device and adverse events such as infections related to the large percutaneous driveline.

Over the past 15 years, the second generation of pumps utilizing continuous-flow (CF) technology has been developed with only a single moving part, the rotor, thus significantly reducing the size of the pump and providing more durable support with less mechanical wear compared to the first-generation LVADs. Recently, third-generation pumps without mechanical bearings have been developed that suspend the impeller with magnetic or hydrodynamic suspension systems. The CF rotary pump design consists of axial (mainly second-generation) or centrifugal (mainly third-generation) pump platforms. The main differences between axial and centrifugal pump performance may best be illustrated by analyzing the differences in pump head curves of the two designs. The pump head curve describes the relationship between the flow generated by an axial or centrifugal VAD and the difference in pressure across the inlet and outlet of the pump (pump delta P) that is generated as the VAD and is placed in between the left ventricle and systemic circulation. There is a separate pump head curve that may be generated with different pump speeds. With axial flow pumps, there is a steep and inverse linear relationship between flow and pump delta P. For example, flow increases as pump delta P decreases and vice versa. If the pump delta P is 40 mm Hg in ventricular systole and 80 mm Hg in diastole, for axial flow pumps, this produces a less pulsatile waveform ranging from 3 to 7 l/min during the cardiac cycle (fig. 2) [11].

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