Understanding the Principles of Continuous-Flow Rotary Left Ventricular Assist Devices

Introduction

Mechanical circulatory support with durable left ventricular assist devices (LVADs) has become a mainstay therapy for treatment of advanced heart failure refractory to optimal medical management and has now surpassed cardiac transplantation as the most frequently utilized surgical therapy for treatment of end-stage heart failure. Important progress in the field has occurred with the introduction of continuous-flow (CF) rotary pump technology that has supplanted older pulsatile displacement pumps. CF rotary pump technology has afforded patients improved survival and reduced occurrence of major adverse events owing to the smaller size of these pumps, quieter operation, improved power efficiency, and greater reliability and durability. CF rotary pumps have significantly different flow characteristics compared to pulsatile pumps, and CF rotary pumps in clinical use have differing device designs that influence the flow properties of the pump and its interaction with the native heart. There are several important features of the CF rotary pump technology that permit characterization of the pumps based on axial versus centrifugal design, bearing design, hydrodynamic properties, and mode of pump speed operation. Despite the marked advantages of this technology, numerous challenges in the management of patients remain, related in part to limitations in the ability of CF rotary pumps to adapt to marked changes in preload and afterload conditions, reduced pulsatility, and lack of algorithms that provide physiologic input into pump speed control. A thorough understanding of the theory of operation and design features of the pumps is necessary to provide care for patients supported with this technology.

Historical perspective

Despite the controversies surrounding the issues of maintaining pulsatile or nonpulsatile circulation, the field of mechanical circulatory support (MCS) has significantly evolved from the use of volume displacement pulsatile pumps to CF rotary blood pumps. CF rotary pumps now dominate the field of MCS, with nearly all LVAD implants in the United States representing CF rotary pump technology. This trend has continued for both short-term (i.e., nonimplantable temporary devices) and implantable, durable MCS device designs. CF rotary pumps offer several advantages over pulsatile, volume displacement pumps. These advantages include (1) a smaller pump size (due to the elimination of the displacement chamber required for a pulsatile pump), (2) fewer moving parts (limited to a single internal rotating impeller) resulting in greater durability and reliability, (3) limited blood contacting surfaces, (4) lower noise, and (5) reduced energy requirements that have translated into significantly improved clinical outcomes. The small size of implantable CF rotary pumps has facilitated the expansion of the therapy into children or patients with small body habitus and has promoted applications including implantable durable biventricular support, total artificial heart support, and minimally invasive surgical approaches.

General pump design

A CF rotary pump consists of blood inlet and outlet ports and a single internal rotating element, that is, rotor or impeller, that is suspended within a pump housing that propels blood forward by spinning the impeller at high speeds, imparting significant kinetic energy to the blood that overcomes outflow resistance to the pump. The spinning of the impeller is accomplished by sequentially actuating an electrical current and creating a magnetic field that is coupled to the internal magnets within the impeller. CF rotary pumps in clinical use today are powered by a brushless DC motor and currently require an external power source (most often provided by rechargeable batteries or an AC power cord) that is transmitted to the internal pump through a percutaneous cable or driveline. In the near future, wireless energy transfer systems should eliminate the need for the percutaneous lead to power implantable pumps.

CF rotary pump design: axial versus centrifugal pumps

There are essentially two types of CF rotary pumps in clinical use today: centrifugal flow pumps and axial flow pumps ( Fig. 7.1 ). The primary difference between centrifugal flow and axial flow pumps lies in the design of their rotating element, or impeller. In a centrifugal pump, the outlet path is positioned 90 degrees relative to the axis of rotation or centerline of the impeller and the rotating element acts as a spinning disk with blades that capture fluid and propel the blood from impeller blades to the outflow cannula along a tangential course. In an axial flow pump, the inlet and outlet blood paths are positioned parallel relative to the axis of rotation or centerline of the impeller and the rotating element operates like a propeller in a pipe that pushes fluid forward. In both cases, blood exits opposite to the direction of thrust generated by the pump motor.

Fig. 7.1, (A) Diagram of the axial-flow pump. Blood enters at the inlet end of the rotor and is driven along the axis of rotation or centerline of the rotor to the out-flow end of the pump. (B). Diagram of a centrifugal blood pump where blood enters the pump inlet cannula along the axis of rotation or centerline of the impeller and is driven outward, tangentially to the outlet of the pump.

Spinning of the impeller at high rotational speeds transfers significant kinetic energy to the blood and the blood leaves the pump at a higher pressure and velocity than at its entry. The rate at which a pump adds energy to a fluid is :

W = \frac{d V}{d t} \times Δ P = Q \times Δ P

$W = \frac{d V}{d t} \times Δ P = Q \times Δ P$

where $Q = \frac{d V}{d t}$ $Q = \frac{d V}{d t}$ is the flow and ∆ P is the difference in pressure between the inlet and outlet orifices of the pump. The efficiency of a pump is defined as the ratio of the power output to the required power input:

η_{p u m p} = \frac{Q \times Δ P}{W_{i n p u t}}

$η_{p u m p} = \frac{Q \times Δ P}{W_{i n p u t}}$

Centrifugal devices generally have greater hydraulic efficiency at energy transfer and provide CF at rotational speeds that are much slower, approximately 2000 to 6000 rpm compared with 8000 to 15,000 rpm for pumps with axial flow designs. This generally means that a lower shear stress is exerted on the blood elements and that there is a lower risk of hemolysis. The hydraulic efficiency, defined previously as the ratio of power imparted to the fluid divided by the power input to the impeller, is related to the ability of a pump to transport fluid with minimal power loss over the blood flow path. Hydraulic efficiency is an important, but not sole, determinant of overall LVAD system efficiency. LVAD systems require power supplies and controllers that have unique methods of operating the pump. It is the sum of the efficiency of the motors and controllers and the hydraulic efficiency that contribute to the overall system efficiency.

Bearing design/impeller suspension

CF rotary pump designs can be further distinguished by the mechanism of impeller suspension or levitation and include use of (1) mechanical bearings ( Fig. 7.2 ), (2) hydrodynamic bearings (fluid forces), (3) hydrodynamic bearings working in synergy with magnetic suspension ( Figs. 7.3 and 7.4 ), and (4) variations of active and/or passive magnetic suspension ( Fig. 7.4 ).