As the left ventricle begins to fail, cardiac output falls and intracardiac filling pressures rise. The main goals of MCS are to decompress the failing ventricle and augment systemic perfusion [6]. Mechanical unloading of the left ventricle leads to a decrease in the severity of mitral regurgitation, less pulmonary congestion, and a reduction in pulmonary arterial hypertension, all of which, in turn, can result in improved right ventricular (RV) function. Partial support pumps provide several liters of flow to augment the reduced native ventricular contribution to the total output, whereas full support pumps provide upwards of 6 to 7 L of flow with the native heart contributing little to the total output. Restoration of forward flow and the normalization of filling pressures also reduces neurohormonal activation, with attendant benefits on cardiorenal function; as a result, temporary VAD support may allow reverse ventricular remodeling and sufficient recovery of ventricular function to permit explantation in selected patients [7].

The hemodynamic benefits of mechanical circulatory support with a LV assist device (LVAD) are also associated with favorable structural changes within the cardiac myocytes and extracellular matrix. In studies of isolated human cardiac myocytes, LVAD support increased the magnitude of contraction, shortened the time of peak contraction, and reduced the time to 50% relaxation. In addition, responses to beta-adrenergic stimulation were greater in isolated myocytes after LVAD support. This suggests that mechanical unloading might reverse the downregulation of beta-adrenergic receptors and improve cardiac responsiveness to inotropic stimulation [8,9,10,11,12,13,14,15]. In vivo, mechanical unloading with an LVAD is known to be associated with alteration of gene and protein expression within the cardiac myocyte [16,17], a reduction in nuclear size and DNA content, and a reduction in fibrosis and collagen content within the cardiac extracellular matrix [10,11].

VADs are typically implanted in parallel with the native right- or left-sided circulation. For long-term LVADs, the pump inflow is from a cannula placed directly into the LV apex and the pump outflow is a cannula that is anastomosed to the
ascending aorta just distal to the aortic valve. Pulsatile systems typically have valves in the inflow and outflow cannulae, whereas continuous flow devices do not. For percutaneous systems, the pumps may be placed across the aortic valve and into the left ventricle or into the left atrium via transcatheter puncture of the interatrial septum. RVADs typically have inflow from the right atrium rather than the RV apex as RV apical cannulation typically provides less reliable flow. The venous blood may also be accessed from the cavae or femoral veins. Outflow is directed to the main pulmonary artery just distal to the pulmonic valve through either direct or transvenous cannulation.

Early-generation VADs are volume displacement pumps which fill and empty asynchronously with the cardiac cycle creating pulsatile arterial flow. The pulsatile pumps mostly fill passively or have limited ability to augment their filling; thus, the beat-to-beat filling of the pump depends partially on the cardiac cycle. Although most of the blood volume entering the left ventricle is diverted into the pump, the left ventricle does occasionally fill enough to eject and contribute to the total cardiac output. In settings of hypovolemia, the pump will fill less quickly and thus the pump rate will slow down, the converse is true in the setting of hypervolemia, thus maintaining a relatively constant state of decompression of the left ventricle.

In contrast to pulsatile flow pumps, continuous flow pumps have a continuously rotating impeller which produces forward flow. The left ventricle is continuously and actively unloaded and therefore the left ventricle rarely can fill to the point where it can eject blood during systole. Thus, the patient has little pulsatile contribution from their native ventricles and hence has little to no pulse pressure, but rather have a mean blood pressure. Patients supported with continuous flow LVADs therefore require Doppler ultrasound to assess their blood pressure. Continuous flow pumps are generally one of two major types: axial or centrifugal flow. Axial flow pumps have the impeller rotating in the same plane as the blood flow, whereas centrifugal pumps accelerate the blood perpendicularly to the axis of inflow. They typically have only one moving part (the impeller) which is magnetically or hydrodynamically suspended resulting in little wear over time. Given the size and wear considerations, among others, continuous flow pumps are now the pump of choice for long-term support. The internal and external components of representative pulsatile and continuous flow pumps are as seen in Figure 45.1.

Most recent data suggests that implantation of a continuous-flow LVAD, as compared with a pulsatile-flow device, significantly improved the probability of survival free of stroke and reoperation for device repair or replacement at 2 years in patients with advanced heart failure in whom medical therapy had failed and who were ineligible for transplantation. In addition, the 2 year actuarial survival with an LVAD was significantly better with a continuous-flow device than with a
pulsatile-flow device. The continuous-flow LVAD was also associated with significant reductions in the frequency of adverse events and the rate of repeat hospitalization, as well as with an improved quality of life and functional capacity [19,20,21,22].

Extracorporeal membrane oxygenation (ECMO) can provide pulmonary or cardiopulmonary support for up to a week or more. Blood is withdrawn from the circulation via an inflow cannula to an extracorporeal continuous flow pump, an oxygenator, and then back to the patient through an outflow cannula. There are two basic types of ECMO: venovenous (VV) and venoarterial (VA). In VV ECMO, the blood is withdrawn from a large central or peripheral vein (jugular or femoral) and oxygenated blood is returned via another large vein. Thus, VV ECMO does not provide hemodynamic support, but rather pulmonary support. For VA ECMO the inflow is typically via the femoral vein and the outflow is typically through the femoral artery and thus provides both oxygenation and mechanical circulatory support. VA ECMO is most commonly used in the setting of severe shock in the setting of acute infarction, fulminant myocarditis or cardiac arrest or after a failure to wean from cardiopulmonary bypass. In the setting of a failure to wean from bypass, the intraoperative cannulation that was used for cardiopulmonary bypass can be attached to the ECMO circuit, rather than having new cannula placed peripherally. Outside the setting of the operating room, both VV and VA EMCO can be rapidly instituted even at the bedside, as either configuration can be achieved through peripheral access, but should only be performed by experienced personnel.

Placement of an intracorporeal pump requires a sternotomy and cardiopulmonary bypass. The degree of perioperative bleeding can be affected by preexisting coagulopathy, liver congestion, and prior sternotomies or other concomitant corrective surgeries at the time of MCS. Most current-generation devices, whether temporary or permanent, require anticoagulation with heparin after post-operative bleeding subsides and then chronically with warfarin and, depending on the center, an antiplatelet agent(s). Most pulsatile devices have mechanical prosthetic valves requiring an INR of 2.5 to 3.5, whereas some of the current generation continuous flow devices may only require an INR of 1.5 to 2 [23]. Thus, there is a risk of continued or new onset bleeding throughout the duration of support, but current devices have a risk of bleeding requiring transfusion of about 0.85 per patient year beyond 30 days, which is a substantial improvement in comparison to previous generation pulsatile devices [24,25]. However, the continuous flow pumps present a unique risk for gastrointestinal bleeding. The high shear stress on the blood from the impeller can cause destruction of large multimers of von Willebrand factor (vWF), which results in a picture of acquired von Willebrand’s disease [26,27]. Bleeding risk is mostly manifest from gastrointestinal arteriovenous malformation (AVMs), it is unknown if the loss of vWF multimers results in bleeding from pre-existing AVMs or the lack of pulsatile flow predisposes to the development of AVMs [27,28,29]. Although most patients have a demonstrable loss of vWF multimers, only a minority of patients develop bleeding. For those who are awaiting transplant, bleeding requiring transfusion carries the additional risk of sensitization to human leukocyte anitgen (HLA) antigens that may limit the pool of suitable donor organs [30].

Aside from the infection risks associated with surgery and indwelling lines postoperatively, there is the additional chronic risk associated with the presence of the VAD itself and the associated driveline or cannulae. However, sepsis from any source can result in seeding of interior of the VAD or its components, which may necessitate more urgent and higher risk transplant or even device replacement [31]. Vegetations on LVAD prosthetic valves may also be a source of thromboembolism [32].

Embolism may result from the pump due to inadequate anticoagulation, the cardiac chambers due to arrhythmias such as atrial fibrillation, or may arise from the native vasculature as a result of the patients’ preexisting vascular atherosclerosis. The overall incidence of ischemic stroke varies greatly with type of device, however with the current generation devices the rate is 0.09 per patient-year overall and 0.05 per patient-year after 30 days [33]. Maintenance of goal INR is critical to minimize the risk of thromboembolism.