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.
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