Blood pressure pulse is an intrinsic aspect of the human cardiovascular system. Even in diseased states, pulse pressure (PP) magnitude is largely maintained. The cardiac cycle is required for ventricular filling, repolarization and relaxation of cardiac myocytes, and ventricular ejection. The physiologic benefits of PP in the circulatory system are not well understood. Immediately after leaving the ventricle, blood flow pulsatility is attenuated by arterial compliance. It is widely believed that PP is negligible at the capillaries. This calls into question the role of pulsatility for oxygen and nutrient delivery, which is one of the primary functions of the cardiovascular system.
Recent advances in mechanical circulatory support devices have brought the debate over the importance of pulsatility into the clinical realm. New generation ventricular assist devices (VADs) provide continuous outflow at the benefit of smaller size, simpler implantation, and improved durability (Figure 1). Although these features have promoted the acceptance of continuous flow VAD (CVAD) therapy for heart failure (HF) patients, there are also concerns over potential risk(s) of sustained, “nonpulsatile” (or diminished pulsatility) circulatory support. This review will present potential differences between continuous and pulsatile VAD therapy, with a focus on clinical implications and emerging CVAD technology.
Pulsatility and Mechanical Circulatory Devices
Traditionally, pulsatility has been described by arterial PP or pulsatility index (PI). Pulse pressure is the difference between the maximum systolic and minimum diastolic aortic pressure.1–3 Pulsatility index is the difference between maximum and minimum blood flow velocity, normalized to the average velocity (Figure 2).3,4 Although these measures of pulsatility are sufficient for diagnosing most cardiovascular conditions, they do not adequately quantify the dynamic energies associated with blood flow.1,5,6 The measures of energy equivalent pressure (EEP) and surplus hemodynamic energy (SHE) are better able to quantify the hemodynamic energies in pulsatile pressure and flow waveforms. Energy equivalent pressure is defined by,
in which, Q is the instantaneous blood flow, P is the instantaneous pressure, and t is time.7 Surplus hemodynamic energy is the difference between EEP and mean arterial pressure (MAP) defined as,
Pulsatile flow VAD (PVAD) is able to supplement EEP and SHE, even beyond that of the native ventricle. Undar et al.1 demonstrated that PVAD can generate significantly greater EEP by producing surplus energies that differ by >100%, even when PP are identical. Other VAD outflow features can influence hemodynamic energy, as described by Ising et al.,8 including VAD pump rate (synchronous or asynchronous), systolic:diastolic phase ratio (pulse width), and time shift with respect to native cardiac cycle (copulsation or counterpulsation). Continuous flow VADs do not augment pulsatility under standard operation but rather diminish native PP. A recent review discusses metrics for quantifying pulsatility, benefits of EEP and SHE, and their application for optimizing CVAD operation and patient outcomes.9 With emerging CVAD technologies, there is an opportunity to enhance outflow feature(s) using pump speed (rpm) modulation algorithms.
Role of a Pulse
Regardless of how clinicians and engineers quantify pulsatility, biological cells and organs are innately adapted in detecting dynamic changes of pressures and flow. At the cellular level, pulsatility can be characterized by cyclic shear stress, cyclic strain, and PP.10,11 These mechanical forces constantly induce a variety of cellular signaling pathways, and this process is termed mechanotransduction. A well-described mechanotransduction system is the vascular endothelium. Differences in blood flow patterns, such as laminar, pulsatile, or oscillatory, can have drastic effects on endothelial cell regulation of apoptosis, angiogenesis, atherosclerosis, vascular remodeling, and systemic blood pressure. Endothelial cell function is dependent on the mean, maximum/minimum, and frequency of mechanical stimulation.12–15
Pulsatile flow (PF) produces greater endothelial responses than CF, which may be due, in part, to the synergistic effects of cyclic stretch and shear stress. Nakata et al.13 demonstrated that PF augments the wall shear stress as a function of maximum flow rate. This implies that PF can generate greater wall shear stresses than CF even if mean flow rates are equal. Pulse frequency may also influence vascular response, as endothelial-dependent vascular relaxation was maximized with a 4.2 Hz pulse frequency in rabbit aorta preparations.12 Nakano et al.16 demonstrated that systemic vascular resistance was reduced by increasing PP or pulse rate in a in vivo canine model.
The scope of flow-dependent responses mediated through the vascular endothelium is vast, spanning from gene induction to blood pressure regulation. The molecular signaling involved with endothelial mechanotransduction is thoroughly reviewed by Li et al.17 To focus more specifically on physiologic-level responses, significant changes in vascular function and remodeling have been observed as early as 1–3 days after initiating altered hemodynamic profiles.18–21 Recent findings demonstrated that reduced vessel stretch in porcine carotid arteries diminished endothelial-dependent vascular relaxation21,22 and constriction.21,23 The potential mechanism of reduced vasorelaxation was identified to be attenuated enzymatic nitric oxide production. Impaired vasoconstriction was associated with increased proliferation, increased apoptosis, and reduced contractile protein expression in vascular smooth muscle cells. Other potential consequences of non-PF are elevated oxidative stress, inflammation,24 and arterial and myocardial remodeling.19,23,25 In a study of prolonged nonpulsatile left heart bypass, the aorta of goats became significantly thinner and the proportion of low contractility smooth muscle cell type was increased.26 Nishinaka et al.27 followed up these findings by measuring chronic changes in systemic vascular resistance in vivo. The vascular sensitivity to phenylephrine was reduced after 6 weeks of nonpulsatile cardiopulmonary bypass (CPB) support, but with pulsatile CPB the vascular sensitivity was maintained. In these studies, the nonpulsatile flow animals still experienced an average PP of 12 mm Hg.26,27 Considering these findings, it seems likely that differences between CVAD and PVAD support extend beyond hemodynamics and may also impact cellular function and remodeling.
Pulsatility may also play an important role in the microcirculation and end-organ perfusion. It has been hypothesized that PP is required to open capillary lumens and to enable blood flow through the capillary beds.28–30 Orime et al.30 and Sezai et al.31 demonstrated that after myocardial infarction was induced in pigs, PF CPB support was able to restore microcirculation of end organs back to pre-infarct conditions. Continuous flow support was not as effective in recovering microcirculation of the kidney, liver, stomach, and skin. Perfusion of epicardial and endocardial tissue, carotid artery, and the brain (white and gray matter) was not significantly different between CF and PF support.30,31 In an acute study using a doxorubicin-induced HF canine model, only perfusion of the heart, brain, and kidney was examined.32 These animals were supported by an LVAD operated in PF or CF mode. As with the previous studies, the perfusion of heart, brain, and kidney were not significantly affected by flow mode. Consequently, the authors drew the opposite conclusion, likely due to their limited organ selection, that flow mode does not impact end-organ perfusion.
The role of pulsatility in maintaining microcirculation is debated, largely on the premise that pulse (native or device-generated) is significantly diminished at the capillary level regardless. A complicating factor is that measurement of microcirculatory blood flow may compromise waveform morphology. Several groups have demonstrated that pulse does exist at the capillary level and that microcirculatory flow patterns do vary between PF and CF support.28,33,34 Dobsák et al.35 found that CF resulted in vasoconstriction of the venules, as measured by vessel diameter. This finding is consistent with the earlier discussion of vascular mechanotransduction, but this observation was only statistically valid for the first several hours of flow.
Debate Over VAD-Generated Pulse
Arterial PP is an intrinsic part of human physiology and organ function. Still, advances and trends in VAD technology have provoked doubts as to whether a “physiologic” PP is needed. Some of the hindrances toward a scientific consensus are definition of CF (nonpulsatile versus diminished pulsatility), lack of standardized metrics to quantify pulsatility, and clinical study limitations (i.e., sample size, demographics, and length of LVAD support).2,36,37 Many of the early reports investigating PF and CF support were with CPB for short duration. Ji and Undar38 completed an extensive review of “pulsatile” and nonpulsatile flow CPB with adult and pediatric patients, in which pulsatile support improved end-organ function and blood flow and decreased systemic inflammation. Voss et al.3 attempted to optimize pulse waveform morphology with CPB. However, the pulsatile waveforms did not improve end-organ perfusion and induced blood trauma, leading the authors to speculate that PF was not required during CPB support.
There are two limitations associated with extrapolating CPB findings to clinical LVAD therapy. First, CPB trials are significantly shorter (hours) compared with long-term VAD therapy. The issue of extended nonpulsatile CPB support and “indwelling mechanical hearts” was first addressed by Mandelbaum and Burns in 1965.39 Jett40 hypothesized that the relative benefits of PF may be trivialized by systemic adaptation to chronic nonpulsatile flow. In contrast to the acute CPB results,30,31,38 reports indicated that CVAD and PVAD have equivalent pre- and post-transplant mortality.41,42 Letsou et al.43 reported that the renal and hepatic function was preserved or improved in bridge-to-transplant (BTT) patients (52–514 days) by the Jarvik 2000 CVAD. This patient population was primarily indicated for BTT and post-transplant survival (6–7 months), which may still be too short to extrapolate over predicted destination therapy (DT) implant periods (3–5 years).
Second, the requirement for aortic cross-clamping during CPB results in a nonpulsatile flow. With CVAD therapy, the native ventricle may continue to contract and eject through the aortic valve or the inflow cannula, resulting in a “diminished pulse,” as reported by Letsou et al.43 with Jarvik 2000 CVAD support. A similar observation was reported with the DeBakey CVAD, in which the flow profiles in the middle cerebral arteries of six New York Heart Association class IV HF patients were nearly physiologic.4 In all patients, the mean VAD flow was ≥3.0 L/min and the LV was adequately volume unloaded. The PI in all these patients steadily increased 2–3 months after implant. Continuous and non-PFs are often considered synonymous, which make interpretation of clinical finding ambiguous.
Concerns with CVAD
The HeartMate II (Thoratec, Pleasanton, CA) LV AD was approved by the US Food and Drug Administration (FDA) in 2008 for BTT and in 2010 for DT. The success of the HeartMate II has propelled the development of new CVAD technologies, including the approval of the (HeartWare, Miami Lakes, FL) centrifugal blood pump for BTT in November 2012. Aortic insufficiency (AI), gastrointestinal (GI) bleeding, and right HF events have been reported with CVAD that were not previously seen with PVAD, which has resurrected the debate over the potential need for “pulsatility” or phasic volume unloading. Left ventricular unloading by mechanical support may directly influence ventricular remodeling and inflammatory response, but the consequences of continuous unloading compared with cyclic unloading are not fully understood.
Differences in ventricular unloading between CVAD and PVAD have been reported,44,45 in which CVAD support provided a greater reduction in LV end-diastolic volume resulting in a leftward and downward shift of pressure-volume (P-V) loop. These findings suggest that CVAD may reduce left atrial pressure and LV end-diastolic pressure by at least 50% more than PVAD. Continuous flow VAD may also increase mean aortic pressure and mean diastolic aortic pressure over baseline and PVAD. These studies also suggest that coronary perfusion decreases with increasing CVAD support.45
A major limitation of in vivo LVAD testing has been the lack of a universally accepted large animal HF model. Bartoli et al.46 developed a chronic ischemic HF bovine model (coronary microembolization) and compared acute in vivo PVAD and CVAD responses. Continuous flow VAD support provided greater LV unloading than PVAD support, characterized by lower LV end-diastolic and end-systolic volumes, reduced LV end-diastolic and end-systolic pressures, and increased diastolic aortic pressure. Differences in LV unloading between CVAD and PVAD in clinical trials have been less pronounced. Garcia et al.47 investigated LV unloading between PVAD (HeartMate XVE, Thoratec) and CVAD (HeartMate II, Thoratec). Although the devices were operated at a fixed rate (rpm), the speed of the HeartMate II was selected to provide PI >4 to avoid suction events. After 1 month of support, CVAD patients had larger reduction in end-diastolic (94–77 ml decrease) and end-systolic (83–77 ml decrease) ventricular volumes than the PVAD patients, which was associated with a proportional drop in ventricular diameter. Pulmonary vascular resistance was lower with PVAD, but reductions in filling pressure were statistically indiscernible between CVAD and PVAD.
Two other clinical studies that reported reduced ventricular volumes with VAD support concluded that PVAD produced greater unloading than CVAD.48,49 Klotz et al.48 demonstrated that end-diastolic and end-systolic LV volumes were comparable after at least 30 days of CVAD or PVAD support. However, PVAD patients had larger pre-implant LV volumes, which produced a greater relative decrease than CVAD group. The LVAD outflow for the CVAD group was also significantly less than that of the PVAD patients (3.6 ± 0.9 L/min vs. 5.1 ± 1.0 L/min), making comparisons hard to interpret. Thohan et al.49 reported that PVAD support provided greater LV unloading than CVAD support, as assessed by changes in end-diastolic diameter, end-diastolic and end-systolic volume, and LV mass. However, CVAD and PVAD were able to equally reduce cellular indicators of HF: tumor necrosis factor (TNF)-α, collagen content, and myocyte size. In this study, the average level of support for each VAD type was not presented.
Letsou et al.50 addressed the issue of unequal levels of LVAD support by examining LV mechanical unloading under equivalent PVAD and CVAD outflow rates (3.39 and 3.36 L/min, respectively). The authors reported that synchronous PVAD provided greater LV unloading in HF porcine model (ligated left anterior descending coronary arteries) as quantified by cardiac output, MAP, and filling pressure (left atrial pressure), but morphological or geometric analysis of the ventricle was not reported.
Left ventricular unloading may initiate reverse remodeling and recovery of the heart.51–53 A study investigating the cytoskeletal-extracellular matrix linker protein, dystrophin, in HF patients supported by CVAD or PVAD found that remodeling was similar between pump types, with slightly greater recovery in PVAD patients.53 Although the results of LV unloading by CVAD and PVAD are quite similar and elicit opposing conclusions, molecular indicators may identify more distinct differences.
Overall, the clinical studies suggest that PVAD provide greater LV unloading. However, in one study, the CVAD delivered 30% less flow support than the PVAD48 and another study did not report the mean VAD flows.49 Continuous flow VAD are often run at a lower level of support than PVAD to avoid septum shift, eliminate suction events, maintain moderate pulsatility, and allow for periodic opening of the aortic valve.47,54 Thus, observations of augmented unloading with PVAD may be due to a higher level of support. The experimental studies, in which VAD flow settings were better matched between devices, demonstrated more LV unloading with CVAD support. This implies that LV unloading is more dependent on operational settings than device type.
Pulsatility and Ventricular Recovery
Left ventricular unloading and remodeling with LVAD support has enabled spontaneous myocardial recovery in a small cohort of HF patients55–57 that has led to successful VAD weaning and removal.58–61 In long-term follow-up of LVAD patients that recovered cardiac function and had their device removed, survival rates were equivalent to transplant patients.62 Although the level of LV unloading appears to be equivalent in CVAD and PVAD patients, it has been reported that there is a lower rate of myocardial recovery with device removal in CVAD patients.54,63 In a recent study, myocardial recovery in 144 PVAD and 243 CVAD patients were evaluated.54 Criteria for sustained myocardial recovery were sinus rhythm, minimal mitral regurgitation, LV ejection fraction greater than 45%, and LV end-diastolic diameter of <55 mm in at least four consecutive tests. Under these criteria, myocardial recovery was three times more likely for PVAD than CVAD. In agreement with this study, Kato et al.64 concluded that PVADs are more effective for recovering LV function as evidenced by improved ejection fraction, dP/dtmax, and mitral E/E′, and lower levels of brain natriuretic peptide and markers of extracellular matrix remodeling.
Other considerations are diminished coronary blood flow (CoF) with CVAD or different pump management protocols.47,54,65 Ootaki et al.65 reported significantly reduced coronary flow with increasing CVAD support in a porcine model, suggesting that lower cardiac perfusion with CVAD may impair myocardial recovery. Diminished CoF may be proportional to reduced left ventricular (LV) workload, which is a typical effect of CF support. It has further been suggested that higher degree of support and asynchronous pulsatile pumping provided by PVAD may stimulate myocardial recovery by cyclically resting and reloading, or “training,” the ventricle.55
Another concern of CVAD support in terms of VAD weaning is cardiac disuse atrophy.46,66 In a bovine HF model, Bartoli et al.46 reported significantly reduced ventricular stroke volume (SV) and cardiac metabolic demands (rate-pressure product) with CVAD, concluding that the native workload of the heart was lower with CVAD support. Increased myocardial stiffness has been observed in patients supported by the HeartMate VE.51 Although PVAD support improved LV and right ventricular (RV) passive P-V curves, ventricular masses, and matrix metalloproteinase (MMP)-1 and MMP-9 expression, there was also an associated worsening of myocardial stiffness, total collagen mass, collagen cross-linking, and myocardial tissue levels of angiotensin I and II. Myocardial force generation was equal between the LV free wall sections for normal, dilated cardiomyopathy and PVAD-assisted patients.51 This result was also reported in a controlled experiment with isolated rat hearts subjected to prolonged unloading.67 Despite cardiac atrophy, contractile function was preserved as determined by maximal developed pressure, ventricular contractility (dP/dtmax), and ventricular relaxation (dP/dtmin). Consequently, ventricular stiffness and atrophy may not be sufficient indicators of diminished myocardial function.
Ventricular assist device beat rate and VAD SV reduction weaning protocols with PVAD were tested in a bovine model (Figure 3).55 Slaughter et al.55 concluded that SV reduction provided a better weaning strategy (stable volume reloading) than beat rate reduction that created transient mechanical reloading due to asynchronous filling and emptying cycles. Continuous flow VAD may be approximated to have an infinite beat rate, in which the level of support may be controlled by the motor speed (rpm). Hence, CVAD can achieve steady mechanical reloading by decreasing the VAD rpm (Figure 3). However, determination as to whether non-PF can induce significant recovery is unclear.
As a result of the uncertainties about long-term CF support, contradictory data regarding physiologic responses to CF and PF, and potential for developing clinical myocardial recovery therapy, there is a renewed interest in VAD-generated pulsatile outflow. When considering that CVAD provides comparable or improved survival rates, quality of life, and adverse event rates, while also improving device reliability, ease of implantation, and ease of operation, there is no discussion of returning to the PVAD model.36,68–73 There is growing speculation that with constant CVAD speed (rpm), the native heart may provide sufficient pulsatility to meet physiologic needs and minimize the risk of adverse events.4,36,74 In addition, there is an ongoing work to incorporate pulsatility into current and emerging CVAD models.
Clinical Comparisons of Pulsatile and Continuous Flow
Clinical Studies of CVAD Versus PVAD
A summary of reported clinical findings with CVAD and PVAD support in HF patients is presented in Figure 4. One concern associated with extended CVAD support was that diminished pulsatility would not provide sufficient end-organ perfusion. As previously stated, the Jarvik 2000 CVAD maintained renal and hepatic function over a 6 month period.43 These HF patients were candidates for BTT and had normal or mild hepatic or renal dysfunction before implantation. This same group of investigators later compared end-organ function between CVAD (Jarvik 2000 and HeartMate II) and PVAD (HeartMate XVE) over a span of 15 months.75 Up through 12 months, the markers of hepatic and renal function were nearly identical between CVAD and PVAD patients. At 15 months, patients with CVAD support did demonstrate over two times more blood urea nitrogen and serum glutamic oxaloacetic transaminase than PVAD support. During device support (6–15 months), LDH was consistently greater with CVAD support compared with PVAD. However, these results were not statistically different due to data deviation. Thus, the authors concluded that PVAD and CVAD provide adequate end-organ perfusion over prolonged time periods. Kamdar et al.76 measured similar end-points over 3 months of support for axial CVAD, centrifugal CVAD, and PVAD. Indicators of renal and hepatic function were maintained or improved for all three devices. Continuous flow VADs were operated to enable the aortic valve to open at a ratio of 1:3 beats. Consequently, PP was not significantly reduced by CF support but was nearly doubled with PF support (baseline to 3 months: centrifugal, 31.5–31.7 mm Hg; axial, 35.1–30.5 mm Hg; pulsatile, 30.8–55.8 mm Hg), which is consistent with other reports.4,74
The importance of pulsatility in the inflammatory process may also be considered. It was found that the common inflammatory marker, TNF-α, was not significantly different between the MicroMed DeBakey CVAD and the Novacor PVAD.77 The same was true for the inflammatory markers polynuclear leukocyte elastase and anaphylatoxin C3a. Conversely, interleukin 6 and anaphylatoxin C5a were significantly elevated with CVAD support. The authors only speculated at the mechanism of selective activation for these two inflammatory molecules, offering the opinion that the response was less likely due to lower PP and more likely the molecular association with the coagulation system or VAD-induced blood trauma.
Currently, two pumps are approved by the FDA for DT: the HeartMate XVE PVAD and the HeartMate II CVAD. A few studies have directly compared the physiologic response(s) with each of these devices. Changes in hemodynamics and exercise capacity were measured in a small cohort of patient after 3 months of HeartMate XVE (n = 16) or HeartMate II (n = 18) support.78 Improvements in cardiac output and exercise capacity were similar between both patient groups. However, pulsatile support resulted in greater LV volume unloading. In many HeartMate II patients, the aortic valve did not open, and in three patients the aortic valve opened intermittently. The average PP (approximated from the group averaged arterial systolic pressures and diastolic pressures) with HeartMate II support was 21 mm Hg, less than one half that of the pulsatile XVE. Garcia et al.47 also compared ventricular unloading delivered by HeartMate XVE and HeartMate II and found the pumps to be equally effective. The only noteworthy difference between the pulsatile and continuous device groups was that the HeartMate XVE produced a greater decrease in pulmonary vascular resistance.
A recent study of patient data from Organ Procurement and Transplantation Network/United Network for Organ Sharing Thoracic Registry database was used to evaluate post-transplant outcomes of HeartMate XVE or HeartMate II recipients.79 The average time on the transplant waiting list was 6–7 months for both VAD patient groups. The overwhelming finding was that both VAD types resulted in similar 1 and 3 year survivals. The data showed that after 3 years post-transplant, the HeartMate II survival remained stable, whereas HeartMate XVE survival dropped rapidly. The HeartMate II also demonstrated less risk of early graft rejection and less infection. This is in opposition to an earlier study that reported higher post-transplant rejection rate and severity with Micromed DeBakey CVAD.42 However, pre- and post-transplant survival rates between CVAD and PVAD support were also reported to be similar. Nativi et al.71 examined >2,000 transplant patients bridged with VAD over an 8 year period from the International Society for Heart and Lung Transplantation registry. They compared post-transplant survival of patients with PVAD, CVAD, and without VAD assistance during the years 2004–2008. The PVAD data were also compared with PVAD patients from 2000 to 2004, which better controls for technological and surgical advances over time. Post-transplant survival of PVAD patients has significantly improved in the modern era. In addition, survival rates of patients with CVAD, with PVAD, and without VAD assistance are not significantly different from each other.
Another study of post-transplant responses focused specifically on pulmonary hemodynamics.80 Data for baseline, HeartMate II support (135 day average support time point), and 1 month post-transplant were presented. Continuous flow VAD support produced significant decreases in systolic and diastolic pulmonary artery pressures. In addition, pulmonary vascular resistance decreased from 3.6 to 2.1 Woods units, even with CF support. Patients with or without severe pulmonary hypertension responded equally to CVAD therapy. The pulmonary hemodynamics and total cardiac output remained stable at 1 month post-transplant.
Together, these studies suggest that CVAD are equally as successful as PVAD for bridging patients to transplant and post-transplant survival (Figure 4). It is also appears that CVAD and PVAD provide similar hemodynamic pressures, flows, and ventricular unloading. However, investigation of ventricular recovery and inflammation response indicate that there may be important differences between PF and CF support. Evaluation of noncardiac tissue responses to long-term CVAD support have been limited, specifically vascular tissue that is closely linked to the altered hemodynamics and neurohormonal pathways.
Clinical Events with CVAD
Despite similarities in hemodynamic parameters and patient survival, CVAD usage is associated with several clinical complications that are less common for PVAD support. Common clinical risks associated with CVAD support are AI or fusion, arteriovenous malformation (AVM) and bleeding events, thrombosis, right HF, and pulmonary hypertension.41,81–84 These events were less common with PVAD, which have other known risks such as infection or device malfunction.85
Aortic insufficiency and fusion.
Aortic insufficiency has been commonly reported with CVAD support.86 Diminished aortic PP in CVAD outflow produces a constant and elevated transvalvular pressure gradient. A biomechanical study of aortic valve leaflets during CVAD support indicates that average strain is increased due to augmented minimum systolic strain. Elevated pressure gradient and strain may lead to AI, reduced valve opening, and aortic root or valve leaflet remodeling.87 Patients with CVAD (HeartMate II) were two times more likely to develop AI than patients with PVAD (HeartMate XVE; 14.3% vs. 6.0%, respectively).88 Aortic insufficiency was more common in patients when the aortic valve did not open. An additional study compared AI frequency for multiple CVAD types with that of a PVAD and also determined CVAD and aortic valve opening to be risk factors for AI.89 Thus, the suggested operation of CVAD that permit occasional aortic valve opening offers protection against the development of AI.
A complication of CVAD support that is seemingly contradictory to AI is aortic valve fusion. The increased transvalvular pressure gradient and strain associated with AI can also stimulate aortic valve leaflet remodeling, potentially leading to valve fusion. A retrospective evaluation of samples from HeartMate II BTT patients found that eight of nine patients had evidence of commissural fusion of the aortic valve leaflets.90 Fusion has also been reported with PVAD support as a result of organized thrombosis at the cusps of the aortic valve.91 As would be more common with CVAD, the pumping algorithm in this study kept the aortic valve permanently closed. The authors suggested an automatic venting cycle to clear thrombi from the aortic valve area.92 These data suggest that the critical difference between CF and PF support with respect to AI and aortic valve fusion is the occurrence of aortic valve opening, which is protective for both disorders.
Bleeding events and arteriovenous malformations.
One of the most commonly reported CVAD complications has been GI bleeding. Letsou et al.93 first reported an increased GI bleeding with CVAD therapy in 2005. Gastrointestinal bleeding was observed in 14% of Jarvik 2000 CVAD patients and 15% of HeartMate II patients.41,94 Crow et al.95 and Stern et al.96 performed comparative studies of bleeding events between CVAD and PVAD. Both concluded that CVAD patients had higher incidence of GI bleeding than PVAD patients. Crow et al.95 noticed trends of lower body mass index and longer support duration for CVAD patients with bleeding events compared with those without bleeding events. Stern et al.96 identified the patient age and the use of aspirin preoperatively as risks of GI bleeding. The large majority of the PVAD pumps used for comparison were the HeartMate XVE, which does not require anticoagulation therapy. This biases direct comparison and could imply anticoagulation treatment as a predominant risk of GI bleeding in VAD patients.
Letsou et al.93 determined the cause of bleeding in all their patients to be AVMs. In a much larger study with HeartMate II patients (n = 172), 19% were identified to have GI bleeding and one-third of these patients had associated AVM.97 The association of CVAD and GI bleeding has been correlated to the association of GI bleeding with aortic stenosis, known as Heyde’s syndrome.98 Both aortic stenosis and CVAD therapy result in a diminished PP, which may promote the formation of AVMs.93 Another similarity is the elevated shear stress, through either the calcified aorta or the VAD impeller region. High shear stress is hypothesized to disrupt the pro-coagulation molecule, von Willebrand factor (vWF), as characterized by the loss of high-molecular-weight vWF multimers. Reduced high-molecular-weight vWF multimers have been observed in HeartMate II patients.99 The anecdotal evidence linking CVAD and vWF has generated significant clinical interest. However, it appears that “acquired von Willebrand” disease occurs almost ubiquitously in CVAD patients while bleeding events occur in a much smaller population, which suggests that vWF may not be a sufficient marker for bleeding events.
Another CVAD complication of primary concern is pump thrombosis. Pulsatile flow and CF VAD activate the coagulation system through multiple influences, such as device materials, surface texture, blood-device contact time, and potential hemolysis. High operational speeds achieved by CVAD may increase the risk of hemolysis due to turbulent flow and high shear stresses.100,101 A standard treatment with VAD implant includes anticoagulation therapy. Unfortunately, anticoagulation therapy can increase the risk of bleeding, and the risk of thromboembolism remains elevated.81 The HeartMate II pump has demonstrated a low thrombosis risk as a result of a redesign after the European clinical trials.81,102 This redesign included textured surface features, which was adapted from the HeartMate XVE design. John et al.103 reported one thromboembolic event and one suspected CVAD pump thrombus in 45 patients supported by HeartMate II device using a heparin-based anticoagulation regimen. A modified anticoagulation regime suggests that intravenous heparin administration may not be needed.104 This study included over 400 HeartMate II patients that were divided between three different anticoagulation protocols. Heparin did not provide a significant reduction in thrombosis, ischemic stroke, or hemorrhagic stroke. In addition, between days 3 and 30 post-implant bleeding events requiring transfusion were significantly reduced compared with anticoagulation with heparin. Low occurrence of pump thrombus has also been reported with the Jarvik 2000. This is speculated to be a result of high-velocity blood stream through the pump providing a continuous “washing effect,” which is a unique feature of the Jarvik 2000 CVAD.94,105 With PVAD support, the cyclic filling/emptying of pump volume and frequent opening of the aortic valve provide regular washing of the pump and ventricle. Flow modulation may provide an alternate mechanism for “washing” CVAD, for which control algorithms may be designed with programmed speed (rpm) oscillations.92
Right heart failure and pulmonary hypertension.
The rapid and drastic unloading of the LV by CVAD may contribute to RV volume overload, deviated RV size, and septal shift, which may adversely impact RV contractility. A clinical study comparing HeartMate II and HeartMate XVE demonstrated similar incidence of right HF with CVAD and PVAD support (41% vs. 35%, respectively).41,106 If RVAD support is required following LVAD implant, then the 1 year survival rate drops from 79% to 59%.107 Furthermore, based on early studies indicating that CVAD could not unload the LV as completely as PVAD, the ability of CVAD to relieve pulmonary hypertension associated with HF was questioned.48,49 It has also been shown that CVAD can reduce pulmonary pressures as effectively as PVAD, yet pulmonary vascular resistance was significantly higher with CVAD.47,80
Emerging CVAD Technologies
As a result of the uncertainties with using CVAD for DT and the persistence of CVAD-associated clinical risks, research to generate pulsatility with CVAD models has been ongoing. Although the ability of a CVAD to produce PF is primarily dependent on motor dynamics, centrifugal CVAD possess flatter head curves (ΔP vs. flow) and less tendency for suction events than axial CVAD.108 Hence, centrifugal CVAD output larger flow pulsatility in response to LV pressure changes and can be modulated over a wider flow range. Accordingly, most PF development for CVAD has been conducted with centrifugal pumps. Lim et al.109 presented a proof-of-concept study that coupled a CF centrifugal pump with an intra-aortic balloon pump (IABP) in cardiac arrest pigs and quantified SHE, EEP, and PP. Integrating CVAD with IABP produced a fivefold increase in PP compared with CVAD alone. Importantly, SHE was 100 times greater when CVAD flow was augmented with IABP (20,219.8 vs. 133.2 erg/cm3). In an effort to generate pulsatility with standard CVAD, speed modulation algorithms are being developed. Bearnson et al.110 demonstrated that centrifugal CVAD can produce physiologic PP via speed modulation using a trapezoidal profile. Power losses due to pump acceleration in the pulsatile mode were offset by power conservation during deceleration. This is an important observation concerning VAD battery life and durability.
Early LVAD modulation strategies favored asynchronous control.110,111 Later work implemented sinusoidal and synchronous flow modulation strategies.50,112,113 Vandenberghe et al.114 varied the timing of VAD control modulation, but kept the flow synchronous and did not vary speed amplitude or pulse width. Overall, these studies demonstrated that pulsatility can be created by CVAD flow modulation, yet comparison is difficult because each study only tested certain parameters.
Ising et al.8 performed a computer simulation study of flow modulation parameters and the impact on cardiovascular hemodynamics and coronary perfusion (Table 1). Over 150 different combinations of pump speed amplitude, pump rate, pulse width, and time shift with respect to ventricular beat were tested. Synchronous modulation provided the greatest reduction in LV external work (LVEW), but asynchronous modulation more dramatically increased EEP and SHE. Shifting beat timing had a strong impact on CoF, with counterpulsation significantly improving myocardial supply-demand ratios (CoF/LVEW).8 These results were confirmed by Wang et al.,115 who investigated many of the same parameters to modulate PediVAS centrifugal pump flow in a pediatric mock circulatory model of CPB. Future studies should aim to reproduce the work of Ising et al.8 and Wang et al.115 in adult mock circulatory loops and in vivo animal experiments.
A CVAD that currently uses pump speed modulation is the HeartWare HVAD. The HVAD modulates speed through a Lavare cycle (±200 rpm, 3 second cycle once per minute), which allows intermittent opening of the aortic valve for washing the aortic root.92 Although the HVAD speed change function is intended as an antithrombotic precaution, this speed modulation capability is being further developed for inducing and optimizing pulsatility.116,117 Furthermore, the development the HeartMate III (Thoratec) has included preclinical tests with a pulse mode.118,119 The HeartMate III successfully produced near-physiologic PP and flow waveforms, both in an in vivo sheep model and mock circulation loops. Farrar et al.119 reported a significant increase in EEP and a fourfold increase in dP/dt. However, current CVAD models can only generate PP of 20–30 mm Hg due to technological limitations (maximum flow ~10 L/min). It still needs to be determined whether pulsatility achievable with CVAD flow modulation is sufficient to normalize vascular responses and allow for myocardial recovery.
Advances in LVAD technology using rotary pumps have raised new clinical and scientific questions as to the importance and significance of pulsatile blood pressure and flow. Positive patient outcomes and quality of life with the CF HeartMate II (Thoratec) have drastically shifted the clinical landscape from PVAD to CVAD use in HF therapy. To gain widespread clinical acceptance of emerging CVAD therapies, clinical data from a large patient population with support up to 5 years that demonstrates efficacy, safety, and reliability are needed. There is compelling evidence to suggest that CVAD support for BTT is comparable (or superior) with PVAD. However, differences in molecular responses, specific clinical complications, and lower occurrence of VAD weaning and myocardial recovery have been reported with CVAD, which makes extrapolation for long-term support (DT or ventricular recovery) a potential cause for concern. In anticipation of this unmet need, flow modulation control strategies are being developed to generate a pulse using CVAD. Consequently, future studies focused on elucidating the physiologic responses to varying levels of CVAD-produced pulsatility are warranted.
1. Undar A, Masai T, Frazier OH, Fraser CD Jr. Pulsatile and nonpulsatile flows can be quantified in terms of energy equivalent pressure during cardiopulmonary bypass for direct comparisons. ASAIO J. 1999;45:610–614
2. Undar A, Rosenberg G, Myers JL. Major factors in the controversy of pulsatile versus nonpulsatile flow during acute and chronic cardiac support. Asaio J. 2005;51:173–175
3. Voss B, Krane M, Jung C, et al. Cardiopulmonary bypass with physiological flow and pressure curves: Pulse is unnecessary! Eur J Cardiothorac Surg. 2010;37:223–232
4. Potapov EV, Loebe M, Nasseri BA, et al. Pulsatile flow in patients with a novel nonpulsatile implantable ventricular assist device. Circulation. 2000;102(19 suppl 3):III183–III187
5. Undar A, Rosenberg G, Myers JL. Part 1: Principles of research on pulsatile and nonpulsatile perfusion during chronic support. Asaio J. 2005;51:303–304
6. Guan Y, Karkhanis T, Wang S, et al. Physiologic benefits of pulsatile perfusion during mechanical circulatory support for the treatment of acute and chronic heart failure in adults. Artif Organs. 2010;34:529–536
7. Shepard RB, Simpson DC, Sharp JF. Energy equivalent pressure. Arch Surg. 1966;93:730–740
8. Ising M, Warren S, Sobieski M, Slaughter M, Koenig S, Giridharan G. Flow modulation algorithms for continuous flow left ventricular assist devices to increase vascular pulsatility: A computer simulation study. Cardiovasc Eng Technol. 2011;2:90–100
9. Soucy KG, Koenig SC, Giridharan GA, Sobieski MA, Slaughter MS. Defining pulsatility during continuous flow ventricular assist device support. J Heart Lung Transplant. 2013;32:581–587
10. Ali MH, Schumacker PT. Endothelial responses to mechanical stress: Where is the mechanosensor? Crit Care Med. 2002;30(5 suppl):S198–S206
11. Davies PF, Barbee KA, Volin MV, et al. Spatial relationships in early signaling events of flow-mediated endothelial mechanotransduction. Annu Rev Physiol. 1997;59:527–549
12. Hutcheson IR, Griffith TM. Release of endothelium-derived relaxing factor is modulated both by frequency and amplitude of pulsatile flow. Am J Physiol. 1991;261(1 pt 2):H257–H262
13. Nakata M, Tatsumi E, Tsukiya T, et al. Augmentative effect of pulsatility on the wall shear stress in tube flow. Artif Organs. 1999;23:727–731
14. Busse R, Fleming I. Pulsatile stretch and shear stress: Physical stimuli determining the production of endothelium-derived relaxing factors. J Vasc Res. 1998;35:73–84
15. Dimmeler S, Fleming I, Fisslthaler B, Hermann C, Busse R, Zeiher AM. Activation of nitric oxide synthase in endothelial cells by Akt-dependent phosphorylation. Nature. 1999;399:601–605
16. Nakano T, Tominaga R, Morita S, et al. Impacts of pulsatile systemic circulation on endothelium-derived nitric oxide release in anesthetized dogs. Ann Thorac Surg. 2001;72:156–162
17. Li YS, Haga JH, Chien S. Molecular basis of the effects of shear stress on vascular endothelial cells. J Biomech. 2005;38:1949–1971
18. Liu SQ, Fung YC. Relationship between hypertension, hypertrophy, and opening angle of zero-stress state of arteries following aortic constriction. J Biomech Eng. 1989;111:325–335
19. Montorzi G, Silacci P, Zulliger M, Stergiopulos N. Functional, mechanical and geometrical adaptation of the arterial wall of a non-axisymmetric artery in vitro. J Hypertens. 2004;22:339–347
20. Zulliger MA, Montorzi G, Stergiopulos N. Biomechanical adaptation of porcine carotid vascular smooth muscle to hypo and hypertension in vitro. J Biomech. 2002;35:757–765
21. Thacher T, Gambillara V, Da Silva R, Montorzi G, Stergiopulos N, Silacci P. Oscillatory shear stress and reduced compliance impair vascular functions. Clin Hemorheol Microcirc. 2007;37:121–130
22. Thacher T, Gambillara V, da Silva RF, Silacci P, Stergiopulos N. Reduced cyclic stretch, endothelial dysfunction, and oxidative stress: an ex vivo model. Cardiovasc Pathol. 2010;19:e91–e98
23. Gambillara V, Thacher T, Silacci P, Stergiopulos N. Effects of reduced cyclic stretch on vascular smooth muscle cell function of pig carotids perfused ex vivo. Am J Hypertens. 2008;21:425–431
24. Pinaud F, Loufrani L, Toutain B, et al. In vitro
protection of vascular function from oxidative stress and inflammation by pulsatility in resistance arteries. J Thorac Cardiovasc Surg. 2011;142:1254–1262
25. Abbruzzese TA, Guzman RJ, Martin RL, Yee C, Zarins CK, Dalman RL. Matrix metalloproteinase inhibition limits arterial enlargements in a rodent arteriovenous fistula model. Surgery. 1998;124:328–334; discussion 334
26. Nishimura T, Tatsumi E, Takaichi S, et al. Prolonged nonpulsatile left heart bypass with reduced systemic pulse pressure causes morphological changes in the aortic wall. Artif Organs. 1998;22:405–410
27. Nishinaka T, Tatsumi E, Nishimura T, et al. Change in vasoconstrictive function during prolonged nonpulsatile left heart bypass. Artif Organs. 2001;25:371–375
28. Baba A, Dobsák P, Saito I, et al. Microcirculation of the bulbar conjunctiva in the goat implanted with a total artificial heart: Effects of pulsatile and nonpulsatile flow. ASAIO J. 2004;50:321–327
29. Ji B, Undar A. Comparison of perfusion modes on microcirculation during acute and chronic cardiac support: Is there a difference? Perfusion. 2007;22:115–119
30. Orime Y, Shiono M, Nakata K, et al. The role of pulsatility in end-organ microcirculation after cardiogenic shock. ASAIO J. 1996;42:M724–M729
31. Sezai A, Shiono M, Orime Y, et al. Major organ function under mechanical support: Comparative studies of pulsatile and nonpulsatile circulation. Artif Organs. 1999;23:280–285
32. Eya K, Tuzun E, Conger J, et al. Effect of pump flow mode of novel left ventricular assist device upon end organ perfusion in dogs with doxorubicin induced heart failure. ASAIO J. 2005;51:41–49
33. Konishi H, Sohara Y, Endo S, Misawa Y, Fuse K. Pulmonary microcirculation during pulsatile and non-pulsatile perfusion. ASAIO J. 1997;43:M657–M659
34. Lee JJ, Tyml K, Menkis AH, Novick RJ, Mckenzie FN. Evaluation of pulsatile and nonpulsatile flow in capillaries of goat skeletal muscle using intravital microscopy. Microvasc Res. 1994;48:316–327
35. Dobsák P, Novakova M, Baba A, et al. Influence of flow design on microcirculation in conditions of undulation pump-left ventricle assist device testing. Artif Organs. 2006;30:478–487
36. Slaughter MS. Long-term continuous flow left ventricular assist device support and end-organ function: Prospects for destination therapy. J Card Surg. 2010;25:490–494
37. Hickey PR, Buckley MJ, Philbin DM. Pulsatile and nonpulsatile cardiopulmonary bypass: Review of a counterproductive controversy. Ann Thorac Surg. 1983;36:720–737
38. Ji B, Undar A. An evaluation of the benefits of pulsatile versus nonpulsatile perfusion during cardiopulmonary bypass procedures in pediatric and adult cardiac patients. ASAIO J. 2006;52:357–361
39. Mandelbaum I, Burns WH. Pulsatile and nonpulsatile blood flow. JAMA. 1965;191:657–660
40. Jett GK. Physiology of nonpulsatile circulation: Acute versus chronic support. ASAIO J. 1999;45:119–122
41. John R, Kamdar F, Liao K, Colvin-Adams M, Boyle A, Joyce L. Improved survival and decreasing incidence of adverse events with the HeartMate II left ventricular assist device as bridge-to-transplant therapy. Ann Thorac Surg. 2008;86:1227–1234; discussion 1234
42. Klotz S, Stypmann J, Welp H, et al. Does continuous flow left ventricular assist device technology have a positive impact on outcome pretransplant and posttransplant? Ann Thorac Surg. 2006;82:1774–1778
43. Letsou GV, Myers TJ, Gregoric ID, et al. Continuous axial-flow left ventricular assist device (Jarvik 2000) maintains kidney and liver perfusion for up to 6 months. Ann Thorac Surg. 2003;76:1167–1170
44. Giridharan GA, Ewert DL, Pantalos GM, et al. Left ventricular and myocardial perfusion responses to volume unloading and afterload reduction in a computer simulation. ASAIO J. 2004;50:512–518
45. Koenig SC, Pantalos GM, Gillars KJ, Ewert DL, Litwak KN, Etoch SW. Hemodynamic and pressure-volume responses to continuous and pulsatile ventricular assist in an adult mock circulation. ASAIO J. 2004;50:15–24
46. Bartoli CR, Giridharan GA, Litwak KN, et al. Hemodynamic responses to continuous versus pulsatile mechanical unloading of the failing left ventricle. ASAIO J. 2010;56:410–416
47. Garcia S, Kandar F, Boyle A, et al. Effects of pulsatile- and continuous-flow left ventricular assist devices on left ventricular unloading. J Heart Lung Transplant. 2008;27:261–267
48. Klotz S, Deng MC, Stypmann J, et al. Left ventricular pressure and volume unloading during pulsatile versus nonpulsatile left ventricular assist device support. Ann Thorac Surg. 2004;77:143–149; discussion 149
49. Thohan V, Stetson SJ, Nagueh SF, et al. Cellular and hemodynamics responses of failing myocardium to continuous flow mechanical circulatory support using the DeBakey-Noon left ventricular assist device: A comparative analysis with pulsatile-type devices. J Heart Lung Transplant. 2005;24:566–575
50. Letsou GV, Pate TD, Gohean JR, et al. Improved left ventricular unloading and circulatory support with synchronized pulsatile left ventricular assistance compared with continuous-flow left ventricular assistance in an acute porcine left ventricular failure model. J Thorac Cardiovasc Surg. 2010;140:1181–1188
51. Klotz S, Foronjy RF, Dickstein ML, et al. Mechanical unloading during left ventricular assist device support increases left ventricular collagen cross-linking and myocardial stiffness. Circulation. 2005;112:364–374
52. Scheinin SA, Capek P, Radovancevic B, Duncan JM, McAllister HA Jr, Frazier OH. The effect of prolonged left ventricular support on myocardial histopathology in patients with end-stage cardiomyopathy. ASAIO J. 1992;38:M271–M274
53. Vatta M, Stetson SJ, Jimenez S, et al. Molecular normalization of dystrophin in the failing left and right ventricle of patients treated with either pulsatile or continuous flow-type ventricular assist devices. J Am Coll Cardiol. 2004;43:811–817
54. Krabatsch T, Schweiger M, Dandel M, et al. Is bridge to recovery more likely with pulsatile left ventricular assist devices than with nonpulsatile-flow systems? Ann Thorac Surg. 2011;91:1335–1340
55. Slaughter MS, Sobieski MA, Koenig SC, Pappas PS, Tatooles AJ, Silver MA. Left ventricular assist device weaning: Hemodynamic response and relationship to stroke volume and rate reduction protocols. ASAIO J. 2006;52:228–233
56. Levin HR, Oz MC, Chen JM, Packer M, Rose EA, Burkhoff D. Reversal of chronic ventricular dilation in patients with end-stage cardiomyopathy by prolonged mechanical unloading. Circulation. 1995;91:2717–2720
57. Frazier OH, Benedict CR, Radovancevic B, et al. Improved left ventricular function after chronic left ventricular unloading. Ann Thorac Surg. 1996;62:675–681; discussion 681
58. Birks EJ, Tansley PD, Hardy J, et al. Left ventricular assist device and drug therapy for the reversal of heart failure. N Engl J Med. 2006;355:1873–1884
59. Birks EJ, Yacoub MH, Banner NR, Khaghani A. The role of bridge to transplantation: Should LVAD patients be transplanted? Curr Opin Cardiol. 2004;19:148–153
60. Slaughter MS, Silver MA, Farrar DJ, Tatooles AJ, Pappas PS. A new method of monitoring recovery and weaning the Thoratec left ventricular assist device. Ann Thorac Surg. 2001;71:215–218
61. Nakatani T, Sasako Y, Kobayashi J, et al. Recovery of cardiac function by long-term left ventricular support in patients with end-stage cardiomyopathy. ASAIO J. 1998;44:M516–M520
62. Farrar DJ, Holman WR, McBride LR, et al. Long-term follow-up of Thoratec ventricular assist device bridge-to-recovery patients successfully removed from support after recovery of ventricular function. J Heart Lung Transplant. 2002;21:516–521
63. Morshuis M, El-Banayosy A, Arusoglu L, et al. European experience of DuraHeart magnetically levitated centrifugal left ventricular assist system. Eur J Cardiothorac Surg. 2009;35:1020–1027; discussion 1027
64. Kato TS, Chokshi A, Singh P, et al. Effects of continuous-flow versus pulsatile-flow left ventricular assist devices on myocardial unloading and remodeling. Circ Heart Fail. 2011;4:546–553
65. Ootaki Y, Kamohara K, Akiyama M, et al. Phasic coronary blood flow pattern during a continuous flow left ventricular assist support. Eur J Cardiothorac Surg. 2005;28:711–716
66. Kinoshita M, Takano H, Taenaka Y, et al. Cardiac disuse atrophy during LVAD pumping. ASAIO Trans. 1988;34:208–212
67. Brinks H, Tevaearai H, Mühlfeld C, et al. Contractile function is preserved in unloaded hearts despite atrophic remodeling. JThorac Cardiovasc Surg. 2009;137:742–746
68. Lahpor J, Khaghani A, Hetzer R, et al. European results with a continuous-flow ventricular assist device for advanced heart-failure patients. Eur J Cardiothorac Surg. 2010;37:357–361
69. Strueber M, Birks E, Jansz P, O’Driscoll G, Wieselthaler G. Clinical results of the international HeartWare® LVAS bridge to transplant trial. J Heart Lung Transplant. 2009;28
70. Pagani FD, Miller LW, Russell SD, et al.HeartMate II Investigators. Extended mechanical circulatory support with a continuous-flow rotary left ventricular assist device. J Am Coll Cardiol. 2009;54:312–321
71. Nativi JN, Drakos SG, Kucheryavaya AY, et al. Changing outcomes in patients bridged to heart transplantation with continuous- versus pulsatile-flow ventricular assist devices: An analysis of the registry of the International Society for Heart and Lung Transplantation. J Heart Lung Transplant. 2011;30:854–861
72. Rogers JG, Aaronson KD, Boyle AJ, et al.HeartMate II Investigators. Continuous flow left ventricular assist device improves functional capacity and quality of life of advanced heart failure patients. J Am Coll Cardiol. 2010;55:1826–1834
73. Slaughter MS, Rogers JG, Milano CA, et al.HeartMate II Investigators. Advanced heart failure treated with continuous-flow left ventricular assist device. N Engl J Med. 2009;361:2241–2251
74. Jahanmir S, Hunsberger AZ, Heshmat H, et al. Performance characterization of a rotary centrifugal left ventricular assist device with magnetic suspension. Artif Organs. 2008;32:366–375
75. Radovancevic B, Vrtovec B, de Kort E, Radovancevic R, Gregoric ID, Frazier OH. End-organ function in patients on long-term circulatory support with continuous- or pulsatile-flow assist devices. J Heart Lung Transplant. 2007;26:815–818
76. Kamdar F, Boyle A, Liao K, Colvin-Adams M, Joyce L, John R. Effects of centrifugal, axial, and pulsatile left ventricular assist device support on end-organ function in heart failure patients. JHeart Lung Transplant. 2009;28:352–359
77. Loebe M, Koster A, Sänger S, et al. Inflammatory response after implantation of a left ventricular assist device: Comparison between the axial flow MicroMed DeBakey VAD and the pulsatile Novacor device. ASAIO J. 2001;47:272–274
78. Haft J, Armstrong W, Dyke DB, et al. Hemodynamic and exercise performance with pulsatile and continuous-flow left ventricular assist devices. Circulation. 2007;116(11 suppl):I8–15
79. Ventura PA, Alharethi R, Budge D, et al. Differential impact on post-transplant outcomes between pulsatile- and continuous-flow left ventricular assist devices. Clin Transplant. 2011;25:E390–E395
80. John R, Liao K, Kamdar F, Eckman P, Boyle A, Colvin-Adams M. Effects on pre- and posttransplant pulmonary hemodynamics in patients with continuous-flow left ventricular assist devices. J Thorac Cardiovasc Surg. 2010;140:447–452
81. Caccamo M, Eckman P, John R. Current state of ventricular assist devices. Curr Heart Fail Rep. 2011;8:91–98
82. John R, Kamdar F, Eckman P, et al. Lessons learned from experience with over 100 consecutive HeartMate II left ventricular assist devices. Ann Thorac Surg. 2011;92:1593–1599; discussion 1599
83. Thompson LO, Loebe M, Noon GP. What price support? Ventricular assist device induced systemic response. ASAIO J. 2003;49:518–526
84. Frazier OH. Unforeseen consequences of therapy with continuous-flow pumps. Circ Heart Fail. 2010;3:647–649
85. Lietz K, Miller LW. Destination therapy: Current results and future promise. Semin Thorac Cardiovasc Surg. 2008;20:225–233
86. Zamarripa Garcia MA, Enriquez LA, Dembitsky W, May-Newman K. The effect of aortic valve incompetence on the hemodynamics of a continuous flow ventricular assist device in a mock circulation. ASAIO J. 2008;54:237–244
87. May-Newman K, Enriquez-Almaguer L, Posuwattanakul P, Dembitsky W. Biomechanics of the aortic valve in the continuous flow VAD-assisted heart. ASAIO J. 2010;56:301–308
88. Pak SW, Uriel N, Takayama H, et al. Prevalence of de novo aortic insufficiency during long-term support with left ventricular assist devices. J Heart Lung Transplant. 2010;29:1172–1176
89. Hatano M, Kinugawa K, Shiga T, et al. Less frequent opening of the aortic valve and a continuous flow pump are risk factors for postoperative onset of aortic insufficiency in patients with a left ventricular assist device. Circ J. 2011;75:1147–1155
90. Mudd JO, Cuda JD, Halushka M, Soderlund KA, Conte JV, Russell SD. Fusion of aortic valve commissures in patients supported by a continuous axial flow left ventricular assist device. JHeart Lung Transplant. 2008;27:1269–1274
91. Rose AG, Park SJ, Bank AJ, Miller LW. Partial aortic valve fusion induced by left ventricular assist device. Ann Thorac Surg. 2000;70:1270–1274
92. Larose JA, Tamez D, Ashenuga M, Reyes C. Design concepts and principle of operation of the HeartWare ventricular assist system. ASAIO J. 2010;56:285–289
93. Letsou GV, Shah N, Gregoric ID, Myers TJ, Delgado R, Frazier OH. Gastrointestinal bleeding from arteriovenous malformations in patients supported by the Jarvik 2000 axial-flow left ventricular assist device. J Heart Lung Transplant. 2005;24:105–109
94. John R. Current axial-flow devices—The HeartMate II and Jarvik 2000 left ventricular assist devices. Semin Thorac Cardiovasc Surg. 2008;20:264–272
95. Crow S, John R, Boyle A, et al. Gastrointestinal bleeding rates in recipients of nonpulsatile and pulsatile left ventricular assist devices. J Thorac Cardiovasc Surg. 2009;137:208–215
96. Stern DR, Kazam J, Edwards P, et al. Increased incidence of gastrointestinal bleeding following implantation of the HeartMate II LVAD. J Card Surg. 2010;25:352–356
97. Demirozu ZT, Radovancevic R, Hochman LF, et al. Arteriovenous malformation and gastrointestinal bleeding in patients with the HeartMate II left ventricular assist device. J Heart Lung Transplant. 2011;30:849–853
98. Heyde EC. Gastrointestinal bleeding in aortic stenosis [letter]. NEngl J Med. 1958;259:196
99. Uriel N, Pak SW, Jorde UP, et al. Acquired von Willebrand syndrome after continuous-flow mechanical device support contributes to a high prevalence of bleeding during long-term support and at the time of transplantation. J Am Coll Cardiol. 2010;56:1207–1213
100. Kameneva MV, Burgreen GW, Kono K, Repko B, Antaki JF, Umezu M. Effects of turbulent stresses upon mechanical hemolysis: Experimental and computational analysis. ASAIO J. 2004;50:418–423
101. Yasuda T, Funakubo A, Miyawaki F, Kawamura T, Higami T, Fukui Y. Influence of static pressure and shear rate on hemolysis of red blood cells. ASAIO J. 2001;47:351–353
102. Frazier OH, Delgado RM 3rd, Kar B, Patel V, Gregoric ID, Myers TJ. First clinical use of the redesigned HeartMate II left ventricular assist system in the United States: A case report. Tex Heart Inst J. 2004;31:157–159
103. John R, Kamdar F, Liao K, et al. Low thromboembolic risk for patients with the HeartMate II left ventricular assist device. JThorac Cardiovasc Surg. 2008;136:1318–1323
104. Slaughter MS, Naka Y, John R, et al. Post-operative heparin may not be required for transitioning patients with a HeartMate II left ventricular assist system to long-term warfarin therapy. JHeart Lung Transplant. 2010;29:616–624
105. Frazier OH, Myers TJ, Westaby S, Gregoric ID. Use of the Jarvik 2000 left ventricular assist system as a bridge to heart transplantation or as destination therapy for patients with chronic heart failure. Ann Surg. 2003;237:631–636
106. Patel ND, Weiss ES, Schaffer J, et al. Right heart dysfunction after left ventricular assist device implantation: A comparison of the pulsatile HeartMate I and axial-flow HeartMate II devices. Ann Thorac Surg. 2008;86:832–840; discussion 832
107. Kormos RL, Teuteberg JJ, Pagani FD, et al.HeartMate II Clinical Investigators. Right ventricular failure in patients with the HeartMate II continuous-flow left ventricular assist device: Incidence, risk factors, and effect on outcomes. J Thorac Cardiovasc Surg. 2010;139:1316–1324
108. Moazami N, Fukamachi K, Kobayashi M, et al. Axial and centrifugal continuous-flow rotary pumps: A translation from pump mechanics to clinical practice. J Heart Lung Transplant. 2013;32:1–11
109. Lim CH, Son HS, Fang YH, et al. Hemodynamic energy generated by a combined centrifugal pump with an intra-aortic balloon pump. Asaio J. 2006;52:592–594
110. Bearnson GB, Olsen DB, Khanwilkar PS, Long JW, Allaire PE, Maslen EH. Pulsatile operation of a centrifugal ventricular assist device with magnetic bearings. ASAIO J. 1996;42:M620–M624
111. Qian KX. Pulsatile impeller heart: A viable alternative to a problematic diaphragm heart. Med Eng Phys. 1996;18:57–66
112. Cox LG, Loerakker S, Rutten MC, de Mol BA, van de Vosse FN. Amathematical model to evaluate control strategies for mechanical circulatory support. Artif Organs. 2009;33:593–603
113. Shi Y, Lawford PV, Hose DR. Numerical modeling of hemodynamics with pulsatile impeller pump support. Ann Biomed Eng. 2010;38:2621–2634
114. Vandenberghe S, Segers P, Antaki JF, Meyns B, Verdonck PR. Hemodynamic modes of ventricular assist with a rotary blood pump: Continuous, pulsatile, and failure. ASAIO J. 2005;51:711–718
115. Wang S, Rider AR, Kunselman AR, Richardson JS, Dasse KA, Undar A. Effects of the pulsatile flow settings on pulsatile waveforms and hemodynamic energy in a pedivas centrifugal pump. ASAIO J. 2009;55:271–276
116. Ising MS, Sobieski MA, Giridharan GA, Koenig SC, Slaughter MS. Hemodynamic responses to continuous flow LVAD speed modulation in computer simulation and in-vitro mock loop studies.19th Congress of the International Society of Rotary Blood PumpsSeptember 9-10, 2011Louisville, KY
117. Koenig SC. Pulsatile control of continuous flow pumps.ASAIO 58th Annual ConferenceJune 14-16, 2012San Francisco, CA
118. Bourque K, Dague C, Farrar D, et al. In vivo
assessment of a rotary left ventricular assist device-induced artificial pulse in the proximal and distal aorta. Artif Organs. 2006;30:638–642
119. Farrar DJ, Bourque K, Dague CP, Cotter CJ, Poirier VL. Design features, developmental status, and experimental results with the HeartMate III centrifugal left ventricular assist system with a magnetically levitated rotor. ASAIO J. 2007;53:310–315