# Bladeless Heart Pump Design: Modeling and Numerical Analysis
## 1 Introduction  
Cardiovascular diseases, including heart failure, remain one of the major health challenges in developed countries [1]. Heart pumps had become unrivaled tools in situations of heart failure without adequate donor heart availability [2]. While a total artificial heart (TAH) replaces the heart, ventricular assist devices (VAD) support a failing heart. Recently, heart pumps (VAD or TAH) are used to bridge the time to heart transplantation, as a bridge to recovery or as a permanent solution [3, 4]. Based on the type of flow, heart pumps can be divided into two categories: pulsatile and continuous flow pumps. The pulsatile flow heart pump is a positive displacement volumetric pump, while the continuous flow pump is a rotary pump [5]. Years of research and implementation have shown that pumps with continuous flow output cause less blood damage and have superior properties [6, 7, 8]. In terms of engineering aspects, a centrifugal pump is superior to an axial flow device: higher efficiency, less power consumption, and less blood damage [9, 10]. To this day, multiple clinical data showed major complications with infections, thrombosis, hemolysis, and hemorrhagic events [11, 12, 13]. Reliability, lifetime, and the type of bearing also have a significant impact on pump hemocompatibility [14, 15]. The influence of the pump design on blood can be significant, hence hemocompatibility should be a priority when designing the heart pump. Hemocompatibility depends on multiple categories representing specific blood functions: thrombosis, coagulation, platelet function, hematology, and immunology (in accordance with ISO 10993-4). Therefore, minimizing the amount of mechanical damage to blood cells is important for hemocompatibility [16]. Blood is a suspension of elastic particulate cells in a liquid plasma which can be modeled as a multiphase non-Newtonian fluid with time dependent viscosity. In blood rheology, many different blood models are developed [17; 18]. To date, there is no best commonly accepted model, and no model can fully express the complex blood rheology and dependence on many parameters. Therefore, at this stage of our research, more attention is paid to the design of the pump with a simplified blood model. According to [18], if the fluid is not exposed to high shear stresses, the difference between the non-Newtonian and Newtonian models is minimal. The incompressible Newtonian approximation for blood was employed with a fluid density of \(1050\,\mathrm{kg/m^{3}}\) and a viscosity of \(0.0035\,\mathrm{kg/(m\cdot s)}\)[19]. The basic concept of a bladeless heart pump consisting of several flat rotating disks is presented in Figure 1[20; 21]. To prevent fluid from circulating continuously in the space between the two disks, a flow separator is inserted in the space between the inlet and outlet pipes. It forces the fluid to stop rotating and directs it to the outlet pipe. The uniqueness of this pump is that the mechanical energy is transferred to the fluid by rotation of flat disks without blades to ensure good hemocompatibility of the pump. The construction principles of the Tesla pump have been known for many years [22; 23]. However, construction of the bladeless heart pump presented in this paper is significantly different. The Tesla pump has a fluid inlet axial to the disk, while the outlet is tangential to the disk rotation. The disks are fixed to each other with solid links due to the geometric design. In contrast, the input to the presented pump is tangential rather than axial, which makes the first essential difference. Another important difference is that due to the geometric design of the tangential input, it is possible to directly connect the disks to the shaft, so the disks are not interconnected. The main benefit of the new design is that there is no negative impact of the disk links on the fluid, precisely because of the direct connection of the disks to the shaft [19].