A biomaterial is a natural or artificial material that remains in contact with one or more internal components of the human body for the purpose of replacing organ function or treating an abnormal condition. Biocompatibility refers to the effect of a specific biomaterial on exposed host tissues, whereas hemocompatibility refers to the specific effects of a biomaterial or circulatory support system on blood components, coagulation cascade, and the tendency for thrombus formation. Successful blood pump design requires special knowledge of microlevel interactions between blood elements and the contact surface as well as macrolevel considerations that include choice of prosthetic valves (if required), inflow and outflow port design, and blood flow pathways within the pump. The ideal biocompatible surface for blood is functioning endothelium, but the creation of a functioning endothelial layer on a bioprosthetic surface remains elusive. Minimizing thrombogenicity requires avoidance of highly thrombogenic biomaterials, specific features of pump design, and pharmacologic inhibition of the coagulation cascade. The requirement for blood-exposed pump components that minimize thrombogenicity has limited compatible materials for blood-exposed surfaces to titanium, polymers (primarily polyurethanes), silicone, graphite, and pyrolytic carbon.

A fundamental concept for the understanding of blood–pump surface interaction is the process of protein adsorption to biomaterial surfaces. Following exposure of pump surfaces to circulating blood in vivo, a protein layer develops that covers the biomaterial surface. The make-up of this protein layer is determined by the protein composition of the patient’s blood, the chemical composition of the biomaterial surface (more specifically, surface charge and hydrophobicity), and surface topography (rough vs. smooth surface, porous vs. nonporous). Concentration of proteins in blood, net protein charge relative to the biomaterial surface, distribution of charges on the protein surface, and ability of the protein to undergo conformational changes all contribute to the propensity for a given protein to adsorb to the pump surface. ^, Protein interactions with the biomaterial vary over time and are therefore dynamic. The change in composition of proteins that adsorb to the pump surface over time is termed the Vroman effect . The specific details of these protein-surface interactions contribute directly to the likelihood of pump thrombogenicity, because these proteins are biologically active and can initiate platelet adhesion and activation and trigger coagulation cascades.

Both smooth and rough surface designs have been used successfully in pump design. The textured titanium surface of the HeartMate series of LVADs stimulates the formation of a thin, stable coagulum that, although counterintuitive, has proven effective in minimizing development of pump thrombus. ^{, ,}

Application of computer simulations called computational fluid dynamics (CFD) analyses has greatly facilitated the ability to predict the effects of shear stresses in the pump flow pathway and areas of relative stasis on platelet activation and thrombus formation.

Contact between the pump surfaces and specific plasma proteins, including factor XII (Hageman factor), prekallikrein, and factor XI, can initiate the coagulation cascade via the intrinsic clotting system (also called the contact system ) particularly in areas of relative blood stagnation ( Fig. 20.1 ).

The coagulation cascade. Intrinsic clotting system (also called contact system ) is shown on left; extrinsic clotting system (cellular injury) is shown on right. Red arrows indicate inhibitory influences on coagulation cascades. F, Factor; HMWK, high-molecular-weight kininogen; PK, prekallikrein.

(From Holman WL, Teitel ER, Itescu S. Biologic barriers to mechanical circulatory support. In: Frazier OH, Kirklin JK, eds. ISHLT Monograph Series: Mechanical Circulatory Support . Philadelphia: Elsevier; 2006:9.)

The additional critical component of thrombus formation is platelet adhesion, aggregation, and activation . Normal endothelium is antithrombogenic, in part related to active biochemical reactions involving nitric oxide and prostacyclins. In areas of low shear rates, fibronectin functions as the adhesive system for platelets, whereas von Willebrand factor (VWF; factor VIII) is the active adhesive protein for platelets at high shear rates, particularly in combination with fibrinogen.

Adhesion of platelets to vascular subendothelium is facilitated by VWF, which forms a bridge between collagen fibrils in the vessel wall and platelet receptors. Following contact activation, platelets undergo changes in shape in which they display spreading pseudopods and release the contents of their α granules (which contain fibrinogen, fibronectin, thrombospondin, VWF, β-thromboglobulin, platelet factor IV, and platelet-derived growth factor), which stimulate thrombin generation. Platelet aggregation is stimulated by the release of adenosine diphosphate (ADP). Under laminar flow conditions, the released granule contents cannot accumulate, but when flow is nonlaminar, platelets are more likely to aggregate and accumulate. As newly recruited platelets release their granular contents, thrombin formation is facilitated by the generation of fibrinogen/fibrin bridges. Increasing thrombin generation accelerates the formation of thromboxane A ₂ and the release of ADP, which further promotes conversion of fibrinogen to fibrin and platelet activation and aggregation. Damaged red cells from shear stress–induced hemolysis also release ADP, further perpetuating platelet activation.

Preserved vascular endothelial cell function is critical for prevention of intravascular thrombus. Synthesis and release of prostacyclins inhibit platelet activation. Thrombomodulin, an endothelial cell product, neutralizes the procoagulant properties of thrombin and activates protein C, a potent anticoagulant that destroys factors Va and VIIIa. Antithrombin III binds to the endothelial cell plasma membrane and inactivates thrombin.

Specific technologic barriers challenging successful MCS include development of corrosion-resistant materials with minimal toxicity and a high level of structural integrity, management of specific blood-contacting surfaces to minimize thrombogenicity and damage to blood elements, blood pump design, and methods to store energy.

Energy to generate flow from circulatory assist pumps requires conversion of either electrical or pneumatic energy (compressed gas) into kinetic energy (energy of motion). Pulsatile or volume-displacement pumps are usually driven by electric motors that either transfer power directly to a pusher plate mechanism or compress gas or liquid in order to transfer energy to the blood sack or pusher plate. CF (rotary) pumps, currently used in clinical practice, use electric motors to transmit kinetic energy to the blood.

The principles of Starling’s law also apply to circulatory pumps, in that the pump must respond to higher inflow into the pump by increasing output. As in the natural heart, this balance is maintained in pulsatile pumps by variations in stroke volume or pump rate and to some degree with CF pumps by changing the pressure difference across the pump (referred to as ΔP ) by changes in preload and afterload.

The major cause of hemolysis in blood pumps is rapid acceleration or deceleration of red cells through the pump, which can induce red cell membrane fracture. In general, pump-induced hemolysis is considered acceptable if the plasma free hemoglobin is maintained at less than 19 mg/dL. The rate of pressure increases and flow-channel velocities are maintained at levels designed to avoid high shear stresses within the blood pathways within an MCSD.

Proper application of fluid dynamics is critical to minimize thrombus formation. Because blood stasis, particularly flow cessation promotes clot formation, stationary vortex flow must be avoided because the central stagnant portion of the vortex can become a nidus for thrombus formation.

Power sources and alarms must provide reliability and durability backed up by software programs designed to activate appropriate alarm systems when deviations from normal function occur. Approximately 1.6 watts of power are needed to pump 6 L/min at 120 mmHg. Power in excess of 1.6 watts is both wasted and converted to heat that must safely dissipate within the body.

MCS designs have evolved considerably over time. Early technology utilized pulsatile (volume displacement) designs that were replaced by more modern CF (rotary) pumps, which include axial flow and centrifugal pump designs. Despite the controversies surrounding the issues of maintaining pulsatile or nonpulsatile circulation, the field of MCS has evolved from the use of volume displacement pulsatile pumps to CF rotary pumps. The evolution to CF rotary pumps has been associated with major improvements in device reliability resulting in major improvements in patient survival and reduction in adverse events. Rotary pumps designs included both axial and centrifugal pumps, with rotary pumps with centrifugal flow design and with total magnetic levitation of the internal rotor demonstrating greater safety compared to other rotary pump designs. ^, Currently, the use of pulsatile pumps and early generation of rotary pumps with axial flow design has been completely abandoned in adult patients, and these pumps are of historical interest, having been completely replaced by contemporary rotary pumps with centrifugal flow and total magnetic levitation of the internal rotor.

A CF rotary pump consists of blood inlet and outlet ports and a single internal rotating element (i.e., 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 future, wireless energy transfer systems should eliminate the need for the percutaneous lead to power implantable pumps.

There are essentially two types of CF rotary pumps in clinical use today: a centrifugal flow pump and axial flow pump ( Fig. 20.2 ). The primary difference between centrifugal flow and axial flow pumps lies in the design of their rotating element. 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.

(A) 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. (B) Diagram of an 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 outflow end of the pump.

(From Mehra MR, Naka Y, Uriel N, et al. A fully magnetically levitated circulatory pump for advanced heart failure. N Eng J Med. 2017;376:440-450.)

The rate at which a pump adds energy to a fluid is :

where Q = dV dt is the flow and ∆ P is difference in pressure between inlet and outlet orifices of the pump. The efficiency of a pump is defined as the ratio of the useful power output to the required power input:

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 earlier 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.

CF rotary pump designs can be further distinguished by the mechanism of impeller suspension or levitation and include use of: (1) mechanical bearings ( Fig. 20.3 ); (2) hydrodynamic bearings (fluid forces); (3) hydrodynamic bearings working in synergy with magnetic suspension ( Fig. 20.4 ); or (4) variations of active and/or passive magnetic suspension ( Fig. 20.5 ).

Diagram of a continuous flow rotary pump (axial pump) using a mechanical bearing or pivot design to suspend the internal rotor. Insert at lower right-hand corner of the figure demonstrates a “ball and socket” mechanical bearing design. The surfaces of the bearing are interspersed with a thin film of fluid to reduce friction and wear of the bearing.

(Reproduced with permission of Abbott. © 2025 Abbott. All rights reserved.)

Picture of a continuous flow rotary pump (HVAD; Medtronic Inc., Minneapolis, MN) utilizing a combination of magnetic and hydrodynamic suspension of the internal rotor. The insert in the upper left-hand corner of the picture details the hydrodynamic surface of the impeller that utilizes fluid forces to oppose magnetic forces generated from opposing magnets within the center post and impeller.

(With permission from Medtronic Inc., Minneapolis, MN.)

Diagram of a continuous flow rotary pump with total magnetic levitation of the internal impeller.

(Reproduced with permission of Abbott Laboratories. © 2025 Abbott. All rights reserved.)

A mechanical pivot design or contact bearing utilizes mechanical bearings on spherical surfaces rotating in sockets (see Fig. 20.3 ). These bearings are made of ceramic components with low friction coefficient (i.e., ceramic or ruby) and have long durability attributed to, in part, boundary lubrication from blood immersion. ^, Although simplistic in design, a major limitation of this mechanism of impeller support is that the point of contact of the mechanical bearings is a point of friction and heat generation that represent potential sites for fibrin deposition with thrombus formation. Washing of these points by blood or fluid is necessary to provide adequate dissipation of heat. In addition, the concentration of hydrodynamic loads on these bearings, especially at stress concentration points, makes these contact points theoretically susceptible to wear and fatigue.

Noncontact bearing designs incorporate either hydrodynamic or magnetic forces or a combination of both to achieve suspension of the internal impeller without the use of mechanical contact supports. ^{, ,} Without the need for mechanical contacts, these suspension systems do not have points of friction and theoretically have an infinite or limitless lifespan. These levitation systems must control for six degrees of freedom of the impeller: three rotational movements (rotation around “x”, “y” and “z” axes) and three translational movements (displacement without rotation). Rotation around the axis of rotation or centerline of the pump is achieved by electromagnetic coupling of the motor stator and internal magnet of the impeller.

Hydrodynamic bearings utilize pressure generated from a fluid film that acts to separate the rotating surfaces of the impeller from the stationary base of the pump (see Fig. 20.4 ). ^, However, in order to generate a hydrodynamic force, the surface of the impeller must be moving and when the impeller is stationary (e.g., at pump startup or stop, impeller and stationary surfaces, such as the pump housing, will come into contact that could result in damage to the surfaces of the impeller or pump housing, resulting in potential niduses for thrombus formation). Contact can also occur if rotating impeller speeds are not sufficient to overcome outflow pressures resulting in impeller contact with the housing or, in the absence of contact, cause a significant decrease in the distance between impeller and pump housing resulting in increased shear and friction of blood elements passing between impeller and pump housing. Importantly, hydrodynamic forces dissipate quickly as a function of distance, thus this limits the distances between the stationary base (pump housing) and rotating impeller. As a result, the load-bearing fluid film is prone to higher shear stresses, that theoretically can result in more damage to blood elements with hemolysis. Hydrodynamic suspension systems do not utilize position sensors resulting in a less complicated electronic design that enhances the ability to miniaturize pump sizes.

Magnetic forces may be passive without the consumption of power (permanent magnet) or active (induction of magnetic field with electricity) in design (see Fig. 20.5 ). ^, Passive magnets utilize rare earth magnets such as neodymium boron-iron magnets within the rotor and along the housing to achieve suspension of the impeller. Permanent magnet bearings allow for a larger gap between the static motor armature and the rotor, while eliminating the need for mechanical bearings, lubrication, sealing or purging fluid. However, the magnetic forces are dependent on the instantaneous position of the rotor (i.e, magnetic forces are greater when magnets—impeller and pump housing—are at closer distances). Magnetic bearings are often used in conjunction with hydrodynamic or electromagnetic bearings. Control of any axis can be active or passive. In active mode, the repelling force is adjusted by changing the supplied current to the electromagnets based on the instantaneous position or other feedback. In a passive mode, the supplied current to the electromagnets is constant. The electromagnetic force adjusts automatically based on the position itself (i.e., as the impeller approaches the magnet in the pump housing, the repelling forces increase). However, electromagnetic systems are complex, requiring position sensors, electromagnets and extra conductors, connector pins, and electronics to execute the dedicated position control algorithm. If, for example, there is an electrical contact failure in a connector pin, or momentary instability encountered in the control algorithm, pump failure can occur. To make the system fail-safe, electromagnetic bearing elements frequently utilize hydrodynamic bearings as a fail-safe, which require no control. Distances between impeller and pump housing maintained by magnetic fields are a function of the strength of the magnetic field and hence, a function of the size of the magnet. Thus, overall pump size and degree of miniaturization are a function of the size of the magnet needed to maintain the desired gap between impeller and pump housing. Overall efficiencies of the pump are also determined, in part, by magnet strength. The major advantages of magnetic bearings are that distances between impeller and pump housing are generally larger and the degree of shear stresses on blood elements is less compared to mechanical contact bearings or hydrodynamic suspension systems.

Durable pulsatile VAD designs have all been abandoned for use in the adult population and are of historical interest only. Previous implantable, durable pulsatile VADs designed for intracorporeal use included the HeartMate VE (Abbott Labs, Chicago, IL) and its updated version, the HeartMate XVE (Abbott Labs, Chicago, IL) device, the Novacor (WorldHeart, Ottawa Canada) device, and IVAD was unique in that the device had the capability to support both the right ventricle and left ventricle or either ventricle, separately. Durable VADs with paracorporeal design included the Thoratec pVAD (Abbott Labs, Chicago, IL) and Berlin Heart (Berlin Heart, Berlin, Germany).

Numerous CF LVADs have been or are currently in clinical use, including such devices as the HeartMate II LVAS (Abbott Laboratories, Chicago, IL), Medtronic HVAD (Medtronic, Inc., Minneapolis, MN), EVAHEART®2 (EVAHEART, Bellaire, TX), Jarvik 2000 (Jarvik, Inc., New York, NY), BrioVAD (BrioHealth Solutions, Burlington, MA), Corheart 6 VAS (Shenzen Medical Technology, China), and others. The current CF device with the widest global application is the HeartMate 3 (Abbott Laboratories), which will be described in detail in this section.

The HeartMate 3 LVAD (Abbott) is a compact intrapericardial-positioned CF rotary pump with centrifugal design. ^, The unique feature of the HeartMate 3 is the full magnetic levitation of the internal rotor. The design differs from previous devices due to active control of rotation and levitation of the rotor that permits impellor movement farther away from the outer housing creating larger gaps for blood flow within the device. The larger gaps may reduce blood component trauma and improve hemocompatibility of the pump. The HeartMate 3 completed clinical evaluation in the Multicenter Study of MagLev Technology in Patients Undergoing Mechanical Circulatory Support Therapy with HeartMate 3 (MOMENTUM 3) clinical study. The results of the MOMENTUM 3 clinical study demonstrated superiority of the HeartMate 3 over the HeartMate II and demonstrated the best overall survival ever reported in a clinical study of durable MCS ( Fig. 20.10 ). ^,

The HeartMate 3 LVAD.

(Reproduced with permission of Abbott. © 2025 Abbott. All rights reserved.)

The major advantage of the TAH over VADs is the provision for biventricular support with an implantable device. The major disadvantage of the TAH is the necessity of removing the native heart, which removes the option of recovery. The second major disadvantage of current iterations is the space requirement for the device, such that only patients with a body surface area of about 1.7 m ² or greater are suitable.

The only TAH device currently available for clinical use is the SynCardia TAH-t (SynCardia Systems Inc., Tucson, Ariz.). The SynCardia TAH-t replaces the native ventricles and all four valves in the orthotopic position.

Syncardia total artificial heart.

The SynCardia TAH-t is a pneumatic pulsatile pump in which a rigid spherical outer housing around each artificial ventricle supports a seamless blood-contacting segmented polyurethane diaphragm, two intermediate diaphragms, and an air diaphragm ( Fig. 20.11 ). Two Medtronic 27-mm inflow valves and two Medtronic 25-mm outflow valves provide unidirectional flow. The full ejection volume of each ventricle is 70 mL per beat, and the TAH typically generates a cardiac output of 7 to 8 L/min. The atrial components are sewn to the native atrial cuffs, and the atrioventricular and ventricular outflow graft connections are via “quick snap” connectors. The external console consists of one primary and one secondary pneumatic driver, air tanks, transport batteries, an alarm, and a computer monitoring system. Current portable pneumatic drivers and consoles allow out-of-hospital living with full ambulation.

The SynCardia TAH-t.

(From Copeland JG, Smith RG, Arabia FA, et al. Cardiac replacement with a total artificial heart as a bridge to transplantation. N Engl J Med . 2004;351(9):859-867.)

Conventional CPB is recommended for most implants. In the setting of prior cardiac surgery, percutaneous catheters should be placed in the femoral artery and vein, which facilitates rapid initiation of CPB in the event of bleeding or acute cardiac decompensation during sternotomy (see Chapter 5 ).

A pulmonary artery catheter is routinely advised to assess the response of pulmonary artery pressure to pharmacologic interventions and device implantation. Continuous measurement of pulmonary artery and central venous (or right atrial) pressures provides critical information in determining the need for inhaled nitric oxide and/or biventricular support to treat right HF (see Chapter 4 ).

Intraoperative transesophageal echocardiography (TEE) is necessary during MCS implants for assessment of biventricular function, evaluation of aortic regurgitation and presence of a patent foramen ovale, identification of intracardiac and aortic air bubbles during implantation and subsequent de-airing maneuvers, and assessment of inflow cannula position within the left ventricle.

Prior to sternotomy, formal discussions occur with anesthesia and echocardiography colleagues regarding the pulmonary artery and central venous pressures (CVP) and interventions to optimize right ventricular (RV) function, the presence and severity of aortic insufficiency and any need to address the aortic valve surgically, the presence or absence of intraventricular thrombus, and the presence of a patent foramen ovale (which if present has to be surgically closed to prevent postimplant right to left shunting with arterial desaturation).

The most common approach to implantation involves a median sternotomy incision ( Table 20.1 ). The pericardium can be incised to the left of the midline to create a pericardial “flap” to ensure adequate coverage of the outflow graft during closure. Once incised longitudinally to the diaphragm, a “T” is created by incising the pericardium toward the right and then the left. At the apex, the pericardium is scored laterally several centimeters above the left phrenic nerve so as not to constrict the device in the pericardium. It is important to ensure appropriate incision of the pericardium to obtain wide exposure of the apex for a device positioned within the pericardium or to guide creation of a small preperitoneal pocket if needed to prevent compression of the right ventricle that could impair right ventricular function ( Fig. 20.12 ). Even for devices intended for intrapericardial placement, creation of a small preperitoneal space at the apex of the left ventricle can facilitate proper alignment of the inflow cannula. Insufficient leftward dissection will result in inappropriate positioning of the inflow cannula (particularly with a markedly enlarged heart and leftward displacement of the apex) such that the inflow portion points against that lateral wall rather than into the central portion of the LV cavity, producing chronic partial inflow obstruction.

Patient and surgical preparation. Patients with advanced HF are preoperatively maintained with inotrope infusions to optimize organ function. To minimize infectious risks, a new central venous pulmonary artery (PA) catheter should be inserted no more than 24 hours preoperatively. The routine use of PA catheters for LVAD implantation may assist with perioperative therapies including pulmonary vasodilator therapy with nitric oxide as well as inotropic and vasoconstrictor support. The general preparation is similar to other cardiac surgical procedures with the patient prepared and draped in the standard sterile fashion. Routine broad-spectrum antibiotics are administered intravenously. In addition to the PA catheter, an arterial line and a Foley catheter are inserted. A transesophageal echocardiogram (TEE) is used to assess for a patent foramen ovale, valvular dysfunction, and intracardiac thrombi before implantation and is used to optimize pump speed while weaning from cardiopulmonary bypass following implantation. A standard median sternotomy is performed and the pericardium is incised; this incision is carried to the left, along the pericardial reflection to the apex.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg . 2013;18:230-8, with permission from Elsevier). Ao = aorta.

Assembly of the pump and components are performed at this time specific to each device type. The tunneling of the percutaneous lead is performed before administration of heparin to prevent bleeding in the rectus sheath. This process is common to all devices designed with an external power source. The tunnel begins at the inferior aspect of the incision at the posterior rectus SHEATH 2 cm lateral to the midline and passes through the rectus muscle, exits the anterior rectus sheath, and continues in the subcutaneous tissue to exit the skin near the costal margin at the anterior axillary line ( Fig. 20.13 ). It is important not to place the exit site posteriorly because this can interfere with the patient’s ability to provide self-care with the exit site. The cloth-covered sections of the driveline are completely buried underneath the skin so that the interface at the skin exit site is with the smooth surface of the driveline. This technique significantly reduces the risk of infection at the skin exit site of the percutaneous lead. A fixation point for the percutaneous lead can be created approximately 10 cm beyond the exit site with a 3-0 Prolene (Ethicon, Somerville, NJ) suture to distribute any tugging to this point rather than the exit site to prevent trauma to the exit site in the early postoperative period when the percutaneous lead is potentially vulnerable to dislodgement. A 5-0 subcuticular Prolene suture can be used to create a purse string around the exit site to encourage apposition of the skin around the exit site of the lead. While the traditional approach for percutaneous lead exit site placement during LVAD implantation is the right upper quadrant of the abdomen, alternative sites have been explored to mitigate complications such as infection and driveline trauma. Some centers utilize the left upper quadrant or subcostal areas, aiming to allow for a shorter, more direct route to the pump, potentially reducing tension and driveline wear. Additionally, a tailored approach, considering individual patient anatomy, body habitus, and risk factors for driveline complications, is increasingly advocated in contemporary practice.

The percutaneous lead is tunneled through the right rectus sheath and exits just below the costal margin at the anterior axillary line.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg. 2013;18:230-8, with permission from Elsevier).

CPB is established with ascending aorta and right atrial cannulation (unless emergent femoral cannulation is required). Selective cannulation of the superior and inferior vena cavae with snares would be required for venous drainage if a concomitant procedure was planned on the right side (i.e., closure of patent foramen ovale or tricuspid valve repair). After initiation of CPB, the heart is maintained in a continuously beating state without cardioplegic arrest unless additional cardiac procedures are needed such as aortic valve repair or replacement. Suction is maintained on the aortic needle vent and continuous surveillance by TEE is important for detection of any air in the ascending aorta.

The apical cannula for current devices is designed to insert directly into the LV apex. The pump is connected to the left ventricle via an inflow collar or apical cuff. The apex of the heart is exposed by lifting the apex of the heart and placing several sponges underneath the diaphragmatic base. The apex of the LV is identified and marked. We generally identify the apical dimple by marking the course of the left anterior descending coronary artery, drawing a mark along the acute margin of the RV, and projecting this mark onto the obtuse margin. The apical dimple is typically located just anteriorly and laterally to the confluence of these two marks. A stab incision is placed into the apex, a 16F Foley catheter is inserted into the LV, the balloon is inflated, and gentle retraction is placed on the catheter to lift the apex. The LV is vented through the catheter to prevent LV distension and facilitate sewing of the apical cuff to the myocardium. The apical cuff is sewn to the LV apex before coring the LV apex (ie, “sew-then-core” method). The Foley catheter is passed through the apical connector before insertion in the LV ( Fig. 20.14 ).

The Foley catheter is inserted into the apex of the left ventricle. The apical cuff of the LVAD is sewn to the apex of the left ventricle using 4 sutures of 2-0 Ethibond (Ethicon, Somerville, NJ) with pledgets, placed in an interrupted, horizontal mattress fashion. A running suture of 4-0 or 3-0 Prolene around the cuff is then performed for hemostasis (not shown).

(Reprinted from Stulak JM, Abou El Ela A, Pagani FD. Implantation of a Durable Left Ventricular Assist Device: How I Teach It. Ann Thorac Surg . 2017;103:1687-92, with permission from Elsevier.)

The apical cuff is first attached to the LV apex using four stitches of 2-0 Ethibond suture with pledgets placed in an interrupted, horizontal mattress fashion at four quadrants of the apical cuff. The suture is passed through the sewing ring of the apical cuff and the apical cuff is then secured to the apex of the LV. Next, a 4-0 or 3-0 Prolene suture is then continuously stitched in a running fashion around the apical cuff ( Fig. 20.14 ) to ensure hemostasis. Alternatively, the apical cuff can be secured to the apex using 12, 2-0 Ethibond sutures with pledgets placed into the myocardium and circumferentially around the perimeter of the apical cuff in an interrupted, horizontal mattress fashion. The suture is then passed through the sewing ring of the apical cuff and the apical cuff is secured to the apex of the LV ( Fig. 20.15 ).

Alternatively, the apical cuff can be secured to the apex using 12, 2-0 Ethibond sutures with pledgets placed into the myocardium in an interrupted, horizontal mattress fashion. The suture is then passed through the sewing ring of the apical cuff and the apical cuff is secured to the apex of the LV.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg . 2013;18:230-8, with permission from Elsevier).

Next, using a coring knife specifically designed for each specific LVAD brand, a core of LV muscle is removed through the center of the apical cuff ( Fig. 20.16 ). The inside of the left ventricle is carefully inspected to ensure the absence of LV thrombosis or large trabeculae that may obstruct the orifice of the inlet cannula ( Fig. 20.17 ). The heart is then gently filled to evacuate air in the left ventricle. The apical inflow cannula is inserted and secured during continuous TEE surveillance of the ascending aorta for air. With the heart and pump filled with blood, the operating table is rotated away from the surgeon, and the pump is de-aired to allow residual air in the pump to collect in the outflow graft and be eliminated. When no further residual intracardiac air is identified by TEE, the outflow graft is measured to an appropriate length, trimmed, and clamped proximally. An end-to-side anastomosis is then constructed to the exteriorized right lateral aspect of the ascending aorta, usually with continuous 4-0 or 5-0 polypropylene suture ( Fig. 20.18 )

Next, using a coring knife specifically designed for each specific LVAD brand, a core of LV muscle is removed through the center of the apical cuff.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg . 2013;18:230-8, with permission from Elsevier).

(A) At this point, CO ₂ can be infused into the operative field to facilitate deairing of the LV. (B) The inflow cannula is then inserted into the apical cuff and the device is positioned with the outflow graft and driveline parallel to the diaphragm. The outflow graft is positioned with a gentle curve around the perimeter of the acute margin of the right ventricle following the atrioventricular groove to the ascending aorta. Additional deairing is accomplished by passively filling the heart and pump and elevating the apex and gently shaking the ventricle. (C) The outflow graft is then distended, clamped, and trimmed to proper length.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg . 2013;18:230-8, with permission from Elsevier).

Outflow graft anastomosis. It is crucial to determine the appropriate length of the outflow graft as excess length may lead to kinking of the graft. Conversely, a short graft may cause added tension to the anastomosis and lead to bleeding or obstruction. The outflow graft should lie along the atrioventricular groove of the right side. The outflow graft should be trimmed at a 45 degree angle to prevent kinking at the aortic anastomosis and create laminar flow within the ascending aorta. (A) The anastomosis is performed along the proximal anterolateral ascending aorta. A partial-occlusion clamp is placed without dissecting the plane between the aorta and PA. The aorta is incised, and the anastomosis is sewn with a 4-0 polypropylene suture. The anastomosis is deaired. The partial-occlusion clamp is then removed to assess hemostasis. A vascular clamp should remain on the outflow graft while the device is off to avoid retrograde flow into the aorta. (B) A deairing cannula is placed in the ascending left ventricle between the outflow graft and aortic cannula. It is important to place the graft as proximal as possible to allow for future cannulation and aortic anastomosis at the time of heart transplantation.

(Reprinted from Romano MA, Haft J, Pagani FD, HeartWare HVAD: principles and techniques for implantation, Oper Tech Thorac Cardiovasc Surg . 2013;18:230-8, with permission from Elsevier).

Appropriate inotropic support is initiated to optimize RV function, with or without the addition of inhaled pulmonary vasodilator such as nitric oxide or prostaglandin. The driveline and its connecting tubing are passed off the surgical field and connected to the pump console. Weaning from CPB is a critical phase of the operation. Data from multiple sources, CPB, LVAD device, TEE, hemodynamic monitor, and visual inspection of the heart are synthesized in real time to facilitate optimal weaning. CPB is slowly weaned to one-half total flow to facilitate native LV ejection and initial air removal steps. Air removal and filling are guided by TEE findings and hemodynamics as well as visual inspection of the right ventricle. After initial removal of air, the LVAD is actuated at its lowest rpm speed setting, and the clamp on the outflow graft is removed, permitting forward device flow. Air removal is continued, CPB is weaned to one-quarter flow, and rpm speeds are increased slowly. After confirmation of air removal, the patient is weaned from CPB, and further adjustments to rpm speed are made with TEE guidance and hemodynamic assessments. Factors to consider include device flow, hemodynamics, LV septal wall position, atrioventricular opening, and degree of mitral regurgitation. It is important to prioritize septal position by TEE to ensure rightward deviation as the priority in assessment of optimal LVAD speeds while obtaining satisfactory device flows. After this, transient aortic valve opening at least every third beat can be attempted with changes in device speed to avoid stasis in the aortic root and ensure adequate LV filling, followed by further speed changes to reduce the degree of mitral regurgitation, if any.

In the event of severe RV failure early post-bypass with a CF pump, left atrial or pulmonary artery diastolic pressure should be carefully observed for rapid fall, and pump speed should then be rapidly decreased while observing for aortic air by TEE. If air is seen in the ascending aorta, the outflow graft must be clamped and CPB rapidly reestablished when heparinization is adequate.

Prior to sternal closure, the outflow graft and apical cannulation site should be covered with native pericardium to facilitate safe sternal reentry and dissection at the time of cardiac transplantation.

The anterolateral thoracotomy approach (nonsternotomy) is an alternative approach to median sternotomy for durable LVAD implantation with reported lower blood transfusions, less incidence of severe RV dysfunction, and shorter time to extubation. The most common surgical approach for exposure for placement of the inflow cannula is a lateral thoracotomy usually in the fifth or sixth intercostal space. Access for anastomosis of the outflow graft to the ascending aortic can be obtained with either an upper hemisternotomy approach or right anterior thoracotomy at the level of the second or third intercostal space ( Fig. 20.19 ). Institution of CPB can be central or peripheral. One of the touted key benefits of the lateral thoracotomy approach is preservation of RV geometry by leaving most of the pericardium intact and not lifting the heart. The downside is the learning curve and potentially longer operative times. A group of patients who could benefit the most from this approach are those with previous cardiac surgery operations. The upper sternotomy or right anterior thoracotomy avoids the need for complete redo-sternotomy. The more important benefit is avoiding the need for complete dissection around the heart and the need to lift the heart to core the LV apex. The Implantation of the HeartMate 3 in Subjects With Heart Failure Using Surgical Techniques Other Than Full Median Sternotomy (SWIFT) study is a prospective, postmarket, nonblinded, single-arm study of the HeartMate 3 that evaluated durable LV implant strategies other than median sternotomy. The SWIFT study demonstrated no differences in length of hospital stay, blood product utilization, adverse events, functional status, or health-related quality of life between sternotomy and thoracotomy approaches. The study validated the noninferiority of the nonsternotomy surgical technique to the median sternotomy technique.

Pump insertion by left thoracotomy with aortic access by (A) partial sternotomy or (B) right thoracotomy.

(Reprinted from Gosev I, Pham DT, Um JY, Anyanwu AC, Itoh A, Kotkar K, Takeda K, Naka Y, Peltz M, Silvestry SC, Couper G, Leacche M, Rao V, Sun B, Tedford RJ, Mokadam N, McNutt R, Crandall D, Mehra MR, Salerno CT. Ventricular assist device using a thoracotomy-based implant technique: Multi-Center Implantation of the HeartMate 3 in Subjects With Heart Failure Using Surgical Techniques Other Than Full Median Sternotomy (HM3 SWIFT). J Thorac Cardiovasc Surg . 2024;168(5):1474-1484, with permission from Elsevier.)

Tricuspid regurgitation is usually addressed at durable LVAD implant when severe due to the adverse effect on forward flow, RV overload, and hepatic congestion. Recent data suggest that although surgical correction of significant tricuspid regurgitation successfully reduces tricuspid regurgitation, correction of the tricuspid regurgitation does not reduce the incidence of right HF following durable LVAD implantation.

Aortic regurgitation should be addressed when a mild degree or more as aortic regurgitation will generally worsen over longer support times and aortic regurgitation adversely effects net forward cardiac output. Aortic regurgitation is especially important to consider in patients who may be supported longer than a year. Options for addressing important aortic regurgitation include aortic valve repair (i.e., Park stitch) ( Fig. 20.20 ) , aortic valve replacement, or patch closure of the aortic valve anulus. Aortic valve repair is the more common procedure when feasible and has similar durability to AV replacement with optimal patient selection. Mechanical valves in the aortic valve position are usually replaced with a bioprosthesis to avoid thrombotic complications associated with incomplete valve opening following durable LVAD implantation. Mechanical valves in the mitral valve position are not routinely replaced as there are currently no data to support routine explant and exchange with a bioprosthesis.

Park Stitch: Pledgeted 4-0 Prolene sutures are applied to approximate the fibrous nodules of Arantius to create a coaptation stitch.

(Reprinted from Park SJ, Liao KK, Segurola R, Madhu KP, Miller LW. Management of aortic insufficiency in patients with left ventricular assist devices: a simple coaptation stitch method (Park’s stitch). J Thorac Cardiovasc Surg . 2004;127(1):264-6, with permission from Elsevier.)

Controversy surrounds the issue of whether important mitral valve regurgitation should be addressed at the time of durable LVAD implantation. Recent data from the MOMENTUM 3 clinical study demonstrated that significant preoperative mitral regurgitation resolved from a preoperative incidence of 43.5% to a postoperative incidence of 6.2% with HeartMate 3 support and that mitral regurgitation present at baseline or follow-up did not affect late outcomes.

Closure of a patent foramen ovale or atrial septal defect is necessary at the time of durable LVAD implantation. LV unloading with MCS can reduce left atrial pressures below that of right atrial pressures, producing a significant right to left shunt resulting in significant systemic hypoxia and the risk of a paradoxical embolus. Shunting may not be immediately present as the patient is transitioned from CPB to LVAD support but may become evident later as right sided medical support (i.e., inotropes or inhaled pulmonary vasodilators) is weaned in the intensive care unit.

The technique for implantation of the SynCardia TAH-t has been described in detail by Copeland and colleagues (see Fig. 20.11 ). Before systemic heparinization, the arterial outflow conduits are appropriately preclotted, and drivelines are tunneled and brought out under the left costal margin. The left-sided ventricular driveline is positioned approximately 5 cm below the costal margin in the midclavicular line. The RV driveline is brought through the skin about 5 cm medial to the left driveline. Both ventricles are then covered with a surgical towel and placed on the left chest until implantation.

CPB is established and the aorta cross-clamped. The recipient heart is excised on the ventricular side of the atrioventricular groove, leaving a small rim of ventricular muscle. The great vessels are divided just above the aortic and pulmonary valves (see Fig. 22.21 ). The mitral and tricuspid valve leaflets are excised and the chordae trimmed. The coronary sinus ostium in the right atrium is oversewn to prevent backflow of blood to the AV groove.

Ventricular rim with AV (atrioventricular valves) valves and chordae excised. Great vessels are transected just above sinutubular junction.

(Redrawn from Copeland JG, Dowling RD, Tsau PH. Total artificial hearts. In: Frazier OH, Kirklin JK, eds. ISHLT Monograph Series: Mechanical Circulatory Support . Philadelphia: Elsevier; 2006:105.)

A “neopericardium” is constructed out of three 15- × 20-cm sheets of expanded polytetrafluoroethylene (ePTFE). The first sheet of ePTFE is sutured to the pericardial reflection as posteriorly as possible at the level of the superior vena cava, inferior vena cava, and pulmonary veins on the right side (without injuring the phrenic nerve); the second sheet is sutured anterior to the left pulmonary veins on the left side; and the third sheet is placed to cover the entire diaphragmatic reflection. The three ePTFE sheets are later folded over the artificial ventricles after completion of the implantation.

The outer rims of the left and right atrial cuffs, which consist of approximately 1 cm of ventricular muscle and fat in the AV groove, are buttressed circumferentially with 1-cm-wide strips of Teflon felt to reinforce the anastomoses between the atrial cuffs and the quick connects and provide maximal hemostasis. The atrial quick connects are trimmed to 6 mm from the connectors, turned inside out, and anastomosed to the atrial cuff (first left, then right), using 3-0 polypropylene. The suture line includes the felt strip–strengthened free ventricular wall and the interventricular septum, taking care to achieve a maximally hemostatic suture line. After completion of both anastomoses, the quick connects are everted back to normal configuration, and biological glue is applied to the suture lines (see Fig. 20.22 ).

Testing left atrial cuff for hemostasis after placement of atrial quick connector.

(Redrawn from Copeland JG, Dowling RD, Tsau PH. Total artificial hearts. In: Frazier OH, Kirklin JK, eds. ISHLT Monograph Series: Mechanical Circulatory Support . Philadelphia: Elsevier; 2006:105.)

The outflow conduits from each artificial ventricle are cut to an appropriate length, determined by briefly placing the artificial ventricles in the mediastinum. The entire length of the pulmonary artery is preserved and anastomosed to the conduit with 4-0 polypropylene. The aortic anastomosis is similarly constructed. All suture lines are tested for hemostasis using a plastic tester device with saline via a three-way stopcock (see Fig. 20.22 ).

The ventricular connections to the quick connects are facilitated by grasping each side of the quick connect with a large clamp while pushing in the device. To avoid conduit angulation, the LV outflow should be as close as possible to the native aorta. Orientation of the ventricle and the outflow mount is fixed by the atrial connection. Prior to completing the aortic connection, the pump and conduit are filled with saline. The right atrial inflow quick connect and the right arterial conduit are attached to the right artificial ventricle similarly. The right side is filled by partially releasing the inferior vena cava tape, completing device implantation (see Fig. 20.23 ).

Complete implantation of SynCardia TAH-t in orthotopic position.

(Redrawn from Copeland JG, Dowling RD, Tsau PH. Total artificial hearts. In: Frazier OH, Kirklin JK, eds. ISHLT Monograph Series: Mechanical Circulatory Support . Philadelphia: Elsevier; 2006:105.)

The patient is placed in steep Trendelenburg, and the aortic cross-clamp is removed with the ascending aorta vented. With the TAH pumping at 40 beats/min, vigorous de-airing is completed under TEE guidance. The TAH pumping rate is then increased and CPB discontinued while maintaining a CVP of 12 to 15 mmHg. Cardiac output from the TAH is usually about 7 to 8 L/min at this point. After obtaining meticulous hemostasis, the previously secured ePTFE “neopericardium” is closed over the TAH anteriorly. A rectangular piece of ePTFE is passed around the proximal aorta and secured. During chest closure, special attention is focused on device output and TEE monitoring of inferior vena caval and left pulmonary venous flows.

The mainstay of pump anticoagulation therapy involves blockade of the intrinsic contact coagulation cascade and platelet aggregation. (See further discussion under “ Biological Barriers ”.)

Although pulsatile flow can usually be detected by direct arterial line monitoring in the early postimplant period, ambulatory blood pressure may not be reliably measured with a standard blood pressure cuff or with automatic blood pressure devices. The normal Korotkoff sounds are typically absent with the minimal pulsation of CF pumps. The most reliable estimate of mean or systolic blood pressure (depending on degree of pulsatility) is use of a Doppler device over the radial artery while gradually deflating a standard blood pressure cuff. Monitoring outpatient blood pressure with CF pumps is important because of the occasional development of severe hypertension (possibly related to abnormal autonomic deregulation), which, if undetected and untreated, can result in a major intracerebral bleeding event.

When pump speed in a rotary pump exceeds the ability of the left ventricle to supply a continuous inflow of blood into the inlet cannula, the resultant rapid decrease in size of the ventricular chamber can result in cavitary collapse around the inlet cannula, which impedes entrance of blood and creates a “suction event.” The result is a sudden decrease in pump flow. Ventricular suction is remedied by decreasing pump speed. The HeartMate 3 device detects sudden changes in pump flow by a change in pulsatility index (a measure of flow pulse through the pump described by the relationship: pulse index = [maximum flow– minimum flow]/mean flow). Sudden drops in pulsatility index and a decrease in pump flow indicate a likely suction event. The HeartMate 3 device incorporates a suction detection algorithm in which pump speed automatically reduces to a set rpm if a sudden change in pump flow pulsatility is detected, after which the pump speed slowly returns to the set speed. Observation of this phenomenon indicates the need to reduce the set pump speed.

If thrombus develops on the pump rotor, an abnormal increase in power is required to maintain impeller speed. The increase in power consumption causes an inaccurately high estimate of pump flow and a decrease in pulsatility index caused by a reduction in cyclic power consumption (because continuously high-power consumption is caused by the rotor drag induced by thrombus). In contrast, inlet obstruction (secondary to inlet cannula misalignment or thrombus on the inlet cannula orifice), or outflow obstruction (such as kinking of the outflow graft or anastomotic narrowing) produces a decrease in power consumption along with decreased pump flow and pulsatility index. The diagnosis of inflow or outflow obstruction can be aided by echocardiographic imaging, which may reveal a dilated left ventricle, which is not improved by increasing pump speed.

In contrast to pulsatile devices (in which a hand-pumping device is usually available), CF pumps carry no manual option to maintain forward pump flow in the event of pump stoppage. The physiologic sequelae of pump stoppage are discussed in “Continuous Flow Ventricular Assist Devices.”

When pump stoppage occurs, most commonly resulting from damage to the driveline, rapid decompensation and death are common secondary to a combination of severe reduction in cardiac output and a variable degree of “pump/aortic” insufficiency due to the absence of valves. If the patient is viable, emergency transport to an experienced MCS center is mandatory while cardiac output is maximized with inotropic support. Upon hospital arrival, a surgical team should be ready for emergency device exchange. An echocardiogram is necessary to evaluate for possible intracavitary LV thrombus. If no thrombus is identified and soldering of damaged electrical wiring is an option, full heparinization followed by device restart (with the small possibility of thromboembolic stroke) may provide a more favorable risk/benefit ratio than emergency reoperation for pump exchange.

Survival rates for patients with advanced HF undergoing implantation of a durable mechanical circulatory device have significantly improved over the years. Current data indicate a 1-year survival rate of approximately 86% for patients supported by contemporary continuous-flow durable LVADs ( Fig. 20.24 ). Over the past 18 years, 1-year survival rates with durable LVADs have increased by approximately 35% ( Fig. 20.25 ). ^{, ,} In the 2001 seminal publication of the Randomized Evaluation of Mechanical Assistance for the Treatment of Congestive Heart Failure (REMATCH) trial, treatment with an early generation pulsatile durable LVAD (HeartMate XVE; Abbott) demonstrated a 1-year survival of 52% compared to 25% for medical management. Subsequent technologic advances improved 1-year survival to 68% with the next generation of axial rotary pumps (HeartMate II; Abbott) and to 77% with centrifugal rotary pumps (HeartWare HVAD; Medtronic).

Kaplan-Meier survival estimates compared by patients with non–totally magnetically levitated devices (HeartMate II and HeartWare HVAD, Medtronic) in the previous era (2013-2017), non–totally magnetically levitated devices (HeartMate II, Abbott and HeartWare HVAD, Medtronic) in the contemporary era (2018-2022), compared to a totally magnetically levitated device (HeartMate 3, Abbott) in the contemporary era (2018-2022).

Historical survival of patients with advanced heart failure supported with durable LVADs from pivotal clinical trials. ^{, ,} The green and blue curves depict the survival associated with medical therapy and early pulsatile LVAD technology (HeartMate XVE), respectively. Following the REMATCH trial, pulsatile pumps were replaced with smaller CF rotary pumps (HeartMate II, Abbott, Chicago, IL and HeartWare HVAD translating into improved survival ( orange , yellow , purple , and black curves). The contemporary CF rotary pump with total magnetic levitation of the internal impellor (HeartMate 3, red curve) is associated with a 2-year survival of 83%, paralleling that of heart transplantation. (HeartMate II and HeartMate 3 are trademarks of Abbott Laboratories or its related companies.)

(Reproduced with permission of Abbott. © 2025 Abbott. All rights reserved.)

From 2018 to 2022, the Multi-center Study of MagLev Technology in Patients Undergoing MCS Therapy With HeartMate 3 (MOMENTUM 3) compared outcomes of patients receiving the HeartMate II to the contemporary HeartMate 3 device (Abbott). In MOMENTUM 3, the 1- and 5-year survival was 84% and 52%, respectively, for patients receiving the HeartMate 3 compared to 82% and 44%, respectively, for patients receiving the older HeartMate II. The HeartMate 3 arm of the MOMENTUM 3 achieved the best survival in a durable LVAD clinical trial to date. ^, “Real world” survival observed within The Society of Thoracic Surgeons Intermacs National Database (STS Intermacs) has validated survival achievements observed in MOMENTUM 3 (see Fig. 20.24 ). ^, Five year survival from STS Intermacs is approximately 64% for patients receiving a primary LVAD implant. For patients 50 years of age or less, survival at 5 years is approximately 75%. For patients 70 years of age or older, survival at 5 years is 53%.

Survival on durable LVAD support, regardless of device intent, far exceeds the associated mortality from medical management of American Heart Association Stage D HF. Given that approximately 50% of patients receiving a durable LVAD demonstrate some degree of physiologic signs of cardiogenic shock or impaired tissue perfusion (i.e., STS Intermacs Patient Profiles 1 and 2) and with another 36% deemed stable but inotrope dependent (i.e., STS Intermacs Patient Profile 3), these outcomes exceed those of medical management of cardiogenic shock alone (survival of approximately 50% to 70% at 30 days). These high mortalities from cardiogenic shock and ambulatory American Heart Association Stage D HF support consideration for durable LVAD therapy to improve survival.

A number of patient-related risk factors are associated with higher postimplant mortality. Important risk factors reported for mortality with rotary pumps include older patient age, female gender, higher body mass index, lower serum albumin as a marker of nutritional state, renal dysfunction requiring dialysis, higher serum bilirubin as a marker of right HF, and the need for concomitant cardiac operations.

The requirement for biventricular support has been consistently identified to confer a significantly worse prognosis. Survival at 1 year for patients requiring concomitant RVAD support at the time of durable LVAD support is approximately 59%. ^, For patients with biventricular HF who require a TAH, survival is approximately 50% at 1 year (see Fig. 20.26 ). Among patients who ultimately need RV support, mortality appears to increase if the decision for RVAD placement is delayed until after the primary LVAD implant.

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