Computational Fluid Dynamics-Based Design Optimization for an Implantable Miniature Maglev Pediatric Ventricular Assist Device

According to recently available statistics, approximately 36,000 new cases of congenital heart disease (CHD) occur each year [1]. Of these, several studies suggest that 9200, or 2.3 per 1000 live births, require invasive treatment or result in death in the first year of life [2]. The very limited options available to treat ventricular failure in these infants and young children have motivated us to develop the PediaFlow® ventricular assist system, which features a miniature rotodynamic blood pump having a magnetically levitated impeller and streamlined blood flow path. It is intended to be fully implantable, providing chronic (up to 6 months) circulatory support from birth to 2 years of age (3 kg to 15 kg body weight) with a nominal pressure rise of ∼100 mmHg for the flow range of 0.3 ∼ 2.3 L/min. By consideration of the hemodynamic requirements of this population [3], a nominal design point of 1.5 LPM with a target pressure rise of 100 mmHg was chosen for the design procedure. An important design requirement is the need for optimizing and miniaturizing the flow path to maximize hemodynamic performance while minimizing shear-stress induced blood trauma. A unique feature of magnetically levitated axial-flow blood pumps in general, and the PediaFlow® in particular, is a continuous annular fluid gap between rotor and housing. The dimensions of this gap are limited by the requirements for magnetic stiffness and motor efficiency, but must be sufficiently large to permit desired flow of blood and to avoid excessive shear stress and other flow disturbances. When shared with the impeller blades, a narrow annular flow gap generally necessitates greater rotational speeds to generate sufficient pressure rise and flow rate. However, the combination of small gap and high speed can be a formula for blood cell damage without sufficient optimization of flow path geometry including the blade profiles. Because of design tradeoffs such as these, which span across several subsystems of the PediaFlow® (electromagnetics, heat transfer, rotordynamics, etc.), our group has adopted a numerical, multidisciplinary approach to optimization. With regard to the flow path, we employed a CFD-based design optimization system developed by Wu et al. [4,5] that integrates a robust and flexible inverse blade design tool, automatic mesh generators, parameterized geometry models, and mathematical models of blood damage, integrated with commercial CFD packages. The PediaFlow® pump has evolved over four generations, denoted as PF1, PF2, PF3, and PF4. The PF3 evolved from its predecessor (PF2, [6]) by the realization that the rotor can operate above its rotordynamic critical rotational speed, which reduced the requirement for magnetic suspension stiffness. This, in turn, permitted a relatively larger annular gap; hence, smaller rotor diameter. Although PF3 demonstrated excellent in vivo biocompatibility over 72 days [7], adverse fluid–structure interaction caused an unstable operational range greater than 0.8 L/min and 18,000 revolutions per minute (rpm). This study describes the use of the aforementioned CFD-based design optimization tools for overcoming this limitation and thereby expanding the operating range and improving hydrodynamic performance in transitioning from the PF3 to a frozen PF4 design, as illustrated in Fig. ​Fig.11. Fig. 1 Two generations of the PediaFlow® ventricular assist device: PF3 and PF4. Cutaway (bottom) shows critical internal components of PF4.