Intrinsic Investigations of Physiological Interaction for Pulsatile Rotary Blood Pumps

1. Introduction

A number of devices exist to support the circulation of a heart failure patient, with the main distinguishing feature relating to their mode of outflow, i.e. pulsatile or continuous (1). Although a nearly physiological pulsatility in the middle cerebral arteries was developed in patients with continuous axial rotary blood pumps (RBPs) support, the debate continues regarding the benefit of maintaining flow pulsatility (2).

Pulsatile RBPs have the advantage of producing physiological pressure and flow within the cardiovascular system. This may be beneficial to improve microcirculation and prevent complications in the gastrointestinal circulation (3). However, the excessively large size and unreliable mechanical function of such devices far out way the benefits that pulsatile flow may provide. Rotary pumps overcome these issues by offering smaller and more reliable devices. However, the reduced pulse operation may be harmful to the vasculature and/or end organ perfusion. For example, Nishimura et al. (1998) found that prolonged non-pulsatile RPBs reduced aortic wall thickness and decreased the volume ratio of smooth muscle cells (4). Therefore, pulsing an RBP speed to induce pulse pressure with a rotary pump is an attractive control strategy. However, the benefits relating to the timing and operation of this mode are unclear (4).

The ability to control the speed of a continuous flow or Rotary Blood Pump (RBP) may enable the production of a simulated pulse, however to date circulatory support is conventionally provided at a constant mean rotational speed, which, despite the sensitivity to left ventricular and aortic pressure, invariably reduces pulse pressure (5, 6). Additionally, the insensitivity of the device to the altering hemodynamic environment within which it is operating can predispose the left ventricle to collapse (7), or regurgitant pump flow (8). Vandenberghe et al. performed a simulation of synchronous and asynchronous ventricular assist with phase shifting. They found that the synchronous pulsing could maximize stroke volume and ventricular pressure, while asynchronous pulsing led to highly unphysiological hemodynamic (9). However, the best timing of the pulses to favor interaction between heart and RBP was not yet decided, unless an accurate triggering of pulses was achieved. On the other hand, Jansen et al. created a simulation model to investigate the effect of pulsatility phase timing on the LVSW. However, this study was limited by the computational environment within which it was conducted (10).

The purpose of this study was to investigate the physiological interaction for RBPs in relation to cardiac output. The parameters of left and right ventricular for volume and pressure were used in order to find the most appropriate timing strategy to minimise the potential for ventricular suck down, enhance aortic pulse pressure, and to improve the degree of ventricle washout or unloading.

1. MethodsCardiovascular system (CVS) model description

An electrical equivalent circuit analogue of the cardiovascular system (CVS) model in conjunction with RBP, is used to conduct this study (Fig. 1). In the pump model, the parameters of pump rotational speed, pump differential pressure, inlet pressure, and outlet pressure were adjusted to study the steady-state and heart failure (HF) measurements. This model consists of four modelled compartments, including pulmonary circulation, heart, systemic circulation, and rotary pump. The model was validated using a dog’s data (11).

Fig.1. Electrical equivalent circuit analog of CVS - LVAD interaction. Rin: inlet cannulae resistances; Rout: outlet cannulae resistances; Lin: inlet cannulae instances; Lout: outlet cannulae instances; Rsuc: suction resistance; thor;1 & Pthor;2: intrathoracic pressures

• Parameters adjustment procedure

In the beginning, the rotational speed of RBP was adjusted until the pump outlet flow reached 5L/min. At pump differential pressure 90 mmHg, which was obtained when inlet pressure and outlet pressure equalled 10mmHg and 100 mmHg respectively. Steady-state pilot simulations were performed to find the pressure head-flow relationship of the RBPs at the speed. The inlet pressure was increased from zero mmHg, and the resulting outlet pressure and pump outlet flow were recorded. After that, the inlet pressure was again increased from zero mmHg. However, the outlet resistance valve was also adjusted as per the cannulae response in order to maintain the outlet pressure at 100 mmHg. The corresponding pump outlet flow was then recorded.

In this scenario, the pump was then configured to alter the rotational speed, with pulse peaks in systole or diastole. The pulse amplitude was 500 rpm, with the mean impel speed 2800 rpm. With the recorded, resulting hemodynamic measurements of left ventricular pressure and RBP outflow, the flow-preload relationship of the mixed-flow pump within a cardiac cycle is studied. Resulting hemodynamic (left ventricular and aortic pressure, pump outflow, and ventricular volumes) were recorded.

1. Results

Steady-state simulations were performed to characterize the pressure, volume - flow relationship of left and right ventricular at constant pump rotation speeds. The characteristic curves were achieved by varying resistance of the CVS circulation, i.e. head pressure. The flow measured in this condition was the mass flow rate through the aorta. Peak diastolic left ventricular volume was limited to 4.5 L/min, as shown in Fig.2. End-diastolic left ventricular pressure was eight mmHg, as observed in Fig. 3.

Fig.2. End-diastolic left ventricle volume vs Cardiac output

Fig.3. End-diastolic left ventricle pressure vs Cardiac output

For the right ventricular side, the associated rend diastolic right ventricular pressure compared with the cardiac output are presented in Fig. 4. The end-diastolic right ventricular volume and cardiac output are compared in Fig. 5. End-systole volume, stroke work, pulse pressure, and ejection fraction were influenced by the timing of pulsations, as summarized in Table 1.

Table 1. Summary of results of end systole volume, stroke work, pulse pressure and ejection fraction

	In Systole	Continuous	In Diastole
End Systole Volume	183.8mL	164mL	171.5mL
Stroke Work	0.19Joules	0.18Joules	0.16Joules
Ejection Fraction	19%	16%	14%
Pulse Pressure	21mmHg	10mmHg	7.5mmHg

Fig.4. End-diastolic right ventricle pressure vs Cardiac output

Fig.5. End-diastolic right ventricle volume vs Cardiac output

1. Discussion

Alteration of the pump speed to induce pulsatility in RBPs was found to influence end-diastolic left and right ventricular pressure and volume. Among the three situations, (1. Constant speed 2. Pulse peak in systole 3. Pulse peak in diastole) pulse peak in diastole gave the largest ESV and the smallest stroke work and aortic pulse pressure, as shown in Table 1.

These findings suggest that pulsing the RBP in diastole can create a potentially ideal environment for myocardial recovery, as the load on the heart is minimized. Furthermore, the reduced pulse pressure from continuous operational mode, whilst not intuitively beneficial for end-organ perfusion, may, in fact, benefit recovery. This is due to the maintenance of a higher diastolic aortic pressure higher, and thus, coronary blood flow is expected to increase, as seen in other counter-pulsation devices (12).

Placing the pulse peak in systole, on the other hand, created the largest pulse pressure, by effectively amplifying the remnant ventricular pressure. This may be beneficial in improving end-organ perfusion, particularly during exercise conditions. The resulting reduction in end-systolic volume, whilst predisposing the ventricle to suck down events, creates the added benefit of increased ejection fraction and thus ventricular washout (13).

The pulsatile output from an RBP might be beneficial for microcirculation However, the effect on coronary circulation should also be considered. With good aortic pressure autoregulation maintained between 60 and 200 mmHg, normal coronary blood flow is usually maintained despite changes to coronary perfusion pressure. However, due to the extravascular pressure, the endocardium is more susceptible to ischemia, especially at lower perfusion pressures (14).

Care must, therefore, be taken to maintain a diastolic pressure of greater than 60 mmHg. And the lowest trough gave by ‘pulse peak in systole’ could easily make diastolic aortic pressure lower than the minimum diastolic pressure for normal coronary blood flow (60 mmHg). Although ‘Pulse peaks in diastole’ maintained the largest diastolic aortic pressure, the timing of this pressure may not coincide with ventricular relaxation, due to the phase delay caused by the cannula (15).

1. Conclusion

Physiological interaction for RBPs in terms of left and right ventricular for volume and pressure with cardiac output is investigated. Pump flow volume in diastole changes frequently, while pump flow volume in systole does not change during varying blood pressure or at varying pump speed. This indicates that diastole would be more demanding for a high preload sensitivity when compared to systole. Therefore, it is needed to reduce outflow with a further small reduction of preload, and thus left ventricular pressure (LVP). In order not to suck too much blood from the ventricle, and to prevent from the abnormal reduction of end systolic volume (ESV) or even ventricular collapse.

REFERENCES

1. Rogers JG, Pagani FD, Tatooles AJ, Bhat G, Slaughter MS, Birks EJ, Boyce SW, Najjar SS, Jeevanandam V, Anderson AS, Gregoric ID. Intrapericardial left ventricular assist device for advanced heart failure. New England Journal of Medicine. 2017 Feb 2;376(5):451-60.

1. Granegger M, Masetti M, Laohasurayodhin R, Schloeglhofer T, Zimpfer D, Schima H, Moscato F. Continuous monitoring of aortic valve opening in rotary blood pump patients. IEEE Transactions on Biomedical Engineering. 2016 Jun;63(6):1201-7.

1. Pirbodaghi T. Mathematical modeling of rotary blood pumps in a pulsatile in vitro flow environment. Artificial organs. 2017 Aug;41(8):710-6.

1. Nishimura T, Tatsumi E, Takaichi S, Taenaka Y, Wakisaka Y, Nakatani T, Masuzawa T, Takewa Y, Nakamura M, Endo S, Nakata M. Prolonged nonpulsatile left heart bypass with reduced systemic pulse pressure causes morphological changes in the aortic wall. Artificial organs. 1998 May;22(5):405-10.

1. Bakouri M. Evaluation of an advanced model reference sliding mode control method for cardiac assist device using a numerical model. IET systems biology. 2018 Apr 1;12(2):68-72.

1. Bakouri MA, Salamonsen RF, Savkin AV, AlOmari AH, Lim E, Lovell NH. A Sliding Mode‐Based S tarling‐Like Controller for Implantable Rotary Blood Pumps. Artificial organs. 2014 Jul;38(7):587-93.

1. Vollkron M, Voitl P, Ta J, Wieselthaler G, Schima H. Suction events during left ventricular support and ventricular arrhythmias. The Journal of heart and lung transplantation. 2007 Aug 1;26(8):819-25.

1. Iizuka K, Nishinaka T, Takewa Y, Yamazaki K, Tatsumi E. The influence of pump rotation speed on hemodynamics and myocardial oxygen metabolism in left ventricular assist device support with aortic valve regurgitation. Journal of Artificial Organs. 2017 Sep 1;20(3):194-9.

1. 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 Journal. 2005 Nov 1;51(6):711-8.

1. Jansen‐Park SH, Mahmood MN, Müller I, Turnhoff LK, Schmitz‐Rode T, Steinseifer U, Sonntag SJ. Effects of interaction between ventricular assist device assistance and autoregulated mock circulation including frank–starling mechanism and baroreflex. Artificial organs. 2016 Oct;40(10):981-91.

1. Bakouri MA, Savkin AV, Alomari AH. Nonlinear modelling and control of left ventricular assist device. Electronics Letters. 2015 Apr 2;51(8):613-5.

1. Peters W, Henrichsen H, Miller S, Parkin R, inventors; Sunshine Heart Co Pty Ltd, assignee. Intraluminal inflatable counter-pulsation heart assist device. United States patent application US 10/634,642. 2004 May 20.

1. Schlangen J, Fischer G, Steendijk P, Petko C, Scheewe J, Hart C, Hansen JH, Ahrend F, Rickers C, Kramer HH, Uebing A. Does left ventricular size impact on intrinsic right ventricular function in hypoplastic left heart syndrome?. International journal of cardiology. 2013 Aug 20;167(4):1305-10.

1. 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 Journal. 2005 Nov 1;51(6):711-8.

1. Gilbert JC, Glantz SA. Determinants of left ventricular filling and of the diastolic pressure-volume relation. Circulation Research. 1989 May;64(5):827-52.