The use of short-term mechanical circulatory support today is a common procedure when pharmacologic treatment fails in patients with cardiogenic shock after myocardial infarction or heart failure after cardiotomy. The most widely used device in these situations is the intraaortic balloon pump, which was introduced in 1968. However, in recent years, there have been efforts made to develop more powerful percutaneous assist devices. Examples are the Hemopump and the percutaneous cardiopulmonary bypass. 1–4
One of the problems with percutaneous systems is that their size must correspond with the place they are going to be introduced to in the circulatory system. The femoral artery is often accessed for insertion, and the size of the device is therefore limited to the size of the vessel.
Most existing percutaneous devices depend on a residual function of the left ventricle. In more advanced heart failure, a left ventricular assist device requiring major surgery is needed such as Heartmate, Thoratec, or others. 5–8
We are currently developing a percutaneous circulatory support device and have earlier reported the effect of the device in bench tests and in animal studies. 9 The device is a short-term circulatory support system aimed to be used from 1 day up to a week. However, the pump has been tested several weeks in vitro and up to 4 days in vivo, and no cable break has occurred.
The device is based on a flexible catheter with the proximal end connected to a drive unit and a pump head at the distal end. The pump head consists of a foldable propeller surrounded by foldable filaments. With an umbrella-like mechanism in the drive unit, the propeller unfolds, and the filaments form a cage around the propeller to protect the wall of the aorta. After use, the mechanism disengages the propeller and cage and closes the pump head. This mechanism is based on a dual catheter where the inner catheter (shaft) is longitudinally displaced against the outer catheter.
The rotation of the propeller is transmitted from the drive unit by way of a rotating wire inside the catheter, and a console monitors the rotational speed (rpm). The speed can be adjusted from 0 to 15,000 rpm according to the hemodynamic demands, although a higher speed is possible. However, increasing the rotational speed may result in a clinically significant hemolysis with the current propeller design. The console also controls the vital functions of the device such as the internal purge pressure and the power consumption.
The advantage of the construction of the pump head is the small size (4.6 mm) during the insertion into the aorta by way of the femoral artery, but, when unfolded, a propeller, 15 mm in diameter, is revealed surrounded by the cage, 21 to 22 mm in diameter (Figure 1).
The basic principle for the function of the catheter is that rotation of a propeller creates a pressure drop in front of the propeller and a pressure rise behind it. The difference between these pressures is the pressure gradient; this has been characterized earlier. 10
With the pump placed in the proximal part of the descending aorta, a pressure drop in front of the pump will lead to an afterload reduction for the left ventricle but may also lead to a reduced perfusion pressure in the coronary arteries and the arteries supplying the upper extremities and the head. Behind the pump, the pressure rise will improve the perfusion of the internal organs, including the kidneys.
The pressure drop in front of the propeller may mimic the beneficial effect of drug-induced vasodilation in heart failure. However, an arterial pressure reduction below a critical level could lead to critical perfusion disturbances, with organ failure as a consequence. 11,12
The current pump is aimed at combining afterload reduction and increased perfusion pressure for the internal organs. Beneficial results have been shown in a previous study performed in healthy pigs with the pump head placed in the ascending aorta. 13 However, a safe placement in the ascending aorta by way of the femoral artery may be technically difficult to accomplish. The aim of the current study was therefore to examine the effect on afterload reduction in an acute model with left ventricular failure with the pump positioned in the proximal part of the descending aorta.
Materials and Methods
Anesthesia and Operative Management
Seven domestic calves with a mean weight of 84 (range 70–95) kg were used in the experiment. The local ethical committee approved the study, and the animals received human care in compliance with the Guide for the Care and Use of Laboratory Animals (NIH publication 85-23, revised 1985).
Anesthesia was induced with a dose of 1,000 mg of ketamine (50 mg/ml Ketalar, Parke Davis, Morris Plains, NJ) given intramuscularly. Thiopental sodium 25 mg/ml (Pentothal, Abbott Laboratories, North Chicago, IL) at a dose of 10 mg/kg and atropine (Kabi Pharmacia, Uppsala, Sweden) at a dose of 0.5 mg were given intravenously before tracheostomy (no. 8.5 tube). Anesthesia and muscular relaxation were maintained with a continuous infusion of fentanyl at a dose of 0.1 mg/hour (50 μg/ml, Leptanal, Janssen Cilag, Bucks, UK) and pancuronium bromide (2 mg/ml, Pavulon, Organon Teknika, Boxtel, The Netherlands) given at a rate of 0.1 mg/kg per hour. The animals were ventilated with a Siemens Servoventilator 900 (Siemens-Elema AB, Stockholm, Sweden). A volume controlled, pressure regulated ventilation of 10 L/min (20 breaths per min; positive end-expiratory pressure 8 cm H2O; inspired oxygen fraction, 0.5) with 1% to 2% Isoflurane (Abbott Scandinavia, Stockholm, Sweden) was used.
A Foley catheter was inserted into the urinary bladder by way of a suprapubic cystostomy. A right neck incision was made and an arterial line introduced into the right carotid artery and connected to a pressure transducer (Ohmeda 1DT 1Rose, Murray Hill, NJ) to monitor the proximal arterial pressure. A right femoral incision was made to expose the femoral artery and an arterial line established and connected to a pressure transducer to monitor the distal arterial pressure. A Doppler flow transducer (Transonic T 201, Ithaca, NY) was connected to the same artery to measure the femoral artery flow. The left carotid artery was exposed by means of a left neck incision and a flow transducer connected to the artery to measure the carotid flow (Transonic T201). A left-sided lumbar incision was made, and a Doppler probe (Transonic T201) was connected to the left renal vein to measure the renal blood flow. A sternotomy was performed, the heart exposed, and the pulmonary trunk dissected free. A flow transducer (Transonic) was connected to the pulmonary artery to enable the continuous measurement of the cardiac output (CO) before the induction of heart failure. Then, the proximal left anterior descending artery was found and equipped with a flow transducer (Transonic) to measure the flow.
The left atrial pressure (LAP) was measured with a catheter through the left atrial appendage, and the left ventricular pressure (LVP) was measured by way of a line through the ventricular wall. The right heart was similarly equipped, and the pressure in the pulmonary artery was measured with a line into the vessel. The catheter pump was introduced by way of a cut down in the left femoral artery and was placed in the upper part of the descending aorta.
Left Ventricular Failure
We used a failure model, which has been described earlier. 14 A biventricular inotropic depression was obtained with a combination of a beta-blocker and a calcium antagonist. Metoprolol (17.3 ± 4.8 mg) was given intravenously as a bolus dose to block the beta-receptor system (Seloken 1 mg/ml Hässle, Göteborg, Sweden), and 7.5 ± 3.5 mg of verapamil (Isoptin, 2.5 mg/ml Knoll, Germany) was given as a bolus dose followed by a continuous infusion of 14.3 ± 5.3 mg during the test.
A centrifugal pump was applied to bypass the right ventricle (Rota-Flow, Jostra GmbH, Hechingen, Germany) and to enable an appropriate filling of the left ventricle. The in-line to the centrifugal pump was a two stage catheter inserted into the right atrial appendage and into the inferior caval vein. The outflow line was inserted into the pulmonary trunk, and a tourniquet was tightened around the pulmonary artery with the outflow cannula inside during the tests to prevent competitive flow through the pulmonary valve. The speed of the centrifugal pump was increased to cause a volume overload in the left ventricle and a backward congestion with a LAP of more than 20 mm Hg. A VVI pacemaker was used to control the heart rhythm and increase the LAP (Figure 2).
Calculations of the Heart Work
To avoid a beat to beat comparison between the different animals with regular and irregular pulse curves, the calculation of the heart work was based on the ventricular work performed in 1 minute. The calculation of the left ventricular heart work (left ventricular hydraulic performance [LVHP]) is expressed in watts, according to LVHP = (MAP − LAP) × CO × 0.0136 × 9.81/60, where MAP is mean arterial pressure in the upper aorta, LAP is the left atrial pressure, CO is the cardiac output, 0.0136 is the density of mercury, 9.81 is the gravitational constant, and 60 represents seconds per minute.
Measurements, Data Collection, and Statistics
After induction of left ventricular failure, the hemodynamic values were recorded for the last 3 minutes of the stabilizing period before the pump was started, and the mean values of this period were calculated (baseline). Then, the pump was run for 6 minutes at 14,000 rpm, the hemodynamic values were recorded for the last 3 minutes of this period, and mean values were calculated and compared with the baseline values (14,000 rpm). The pump was then abruptly stopped to observe any rebound phenomenon during the first 30 seconds, and the hemodynamic values were recorded (30 seconds).
Data were collected digitally at 200 Hz, and mean values were calculated every 5 seconds. The data are expressed as mean values with one standard deviation. To compare the difference in the failure model between the baseline values and with the pump running, a paired Student’s t-test was applied. p < 0.05 was considered statistically significant, and the analyses were performed with commercial software (Microsoft Excel and Lotus 1-2-3).
The hemodynamic values are presented in Tables 1 to 3. There was a 9 ± 6 mm drop in mean pressure in the carotid artery from 80 ± 11 to 71 ± 13 mm Hg (p = 0.008) and a 17 ± 11 mm mean pressure rise from 78 ± 10 to 95 ± 20 mm Hg (p = 0.005) in the femoral artery at 14,000 rpm.
The LAP increased from 5 ± 4 mm Hg to 25 ± 3 mm Hg (p < 0.001) after induction of failure. With the pump on, the pressure dropped to 20 ± 5 mm Hg (p = 0.008). The left ventricular systolic pressure dropped 9 ± 4 mm Hg from 109 ± 17 to 100 ± 19 mm Hg (p = 0.004), and the mean left ventricular pressure dropped 8 ± 3 mm Hg from 70 ± 16 to 62 ± 7 mm Hg. Left ventricular performance did not change with the pump on but increased immediately after stopping the pump from 0.40 ± 0.13 to 0.45 ± 0.14 W (p = 0.002). The carotid flow decreased 15% with the pump on from 283 ± 150 to 240 ± 119 ml/min (p = 0.029) and returned to baseline values after pump stop. The renal, femoral, and coronary flow did not change significantly with the pump on, but coronary flow increased from 39 ± 19 to 44 ± 21 ml 30 seconds after stopping the pump (p = 0.01).
A pressure gradient was observed in all animals when the pump was started. Of the total pressure gradient of 24 mm Hg, about one third presented as a pressure drop in front of and two thirds as a pressure rise behind the propeller. The pressure gradient developed immediately when the pump was started and disappeared within seconds when the pump was abruptly stopped (Figure 3).
Although the pump was placed in the descending aorta about 20 to 30 cm away from the aortic valve, the resulting pressure drop in front of the propeller was propagated backward by way of the left ventricle and left atrium and to the pulmonary artery. A reduced afterload was measured both as a decreased mean carotid arterial pressure and a reduction of the left ventricular systolic pressure during pump operation. Of special interest was the influence on the pressure in the left atrium, indicating a reduced preload and a reduced central congestion. If LAP was lowered, the left ventricular end-diastolic pressure and left ventricular volume was reduced likewise, leading to reduced left ventricular wall stress (Figure 4).
No significant pressure changes in the right ventricle were observed because the flow bypassed the ventricle through the centrifugal pump. The ventricle worked against a closed pulmonary artery because of the tourniquet around the vessel. The continuous flow from the pump also explains the small pressure fluctuations in the pulmonary artery.
A reduction in the arterial pressure facilitates the ejection of the left ventricle. In clinical practice, vasodilators are often used to accomplish the reduction in arterial pressure. Many studies have been performed with nitroprusside in left ventricular failure with high filling pressures. Only small drops in arterial pressure are normally required to observe significant hemodynamic effects on the left ventricular filling pressure and on CO. 15–19 Other agents like prostacyclin, chlorpromazine, or nesiritide without inotropic effects have also shown similar effects. 20–24 Patients in cardiogenic shock or with a low CO after heart surgery often benefit from vasodilation. However, critically low blood pressure limits the use of these drugs. A further pressure reduction would lead to an impaired perfusion of internal organs and could lead to organ failure.
Our results suggest that this problem can be overcome to some extent with the current support system because a significant pressure rise behind the pump was observed. The device thereby combines an afterload reduction with a maintained or increased perfusion pressure for the lower internal organs.
We also studied the distribution of flow with a fixed CO from the aorta to four flow regions. Only the carotid flow showed a significant change (−15%) and returned to baseline values immediately after stopping the pump. The fall in arterial pressure in the carotid artery was of a similar magnitude (−12%).
The most important part of the carotid flow in humans is directed to the brain. In humans, the brain is autoregulated to maintain a constant blood supply regardless of blood pressure variations. The autoregulation ensures a sufficient blood supply down to a MAP of 60 mm Hg, and, below this point, the flow is pressure dependent. 25 We examined calves in which the relative size of the brain compared with the head is smaller than in humans. It is therefore not possible to draw any conclusions about the brain perfusion. Because of the importance of sufficient brain perfusion, more specific flow studies need to be performed where the blood pressure is below the point of autoregulation.
The LAD flow did not change with the pump on but increased 12% when the pump was stopped. Even the left ventricular work, which remained the same with the pump on, had a similar rise.
The coronary flow depends on autoregulation in a complex way. The autoregulation controls the flow to the different parts of the myocardium and is intimately linked to the myocardial metabolic demand. Although the driving force is the difference between the aortic pressure and the pressure in the coronary sinus, the resistance depends not only on the vessel’s properties but also partly on the left ventricular filling pressure. 26
In this respect, an increased end-diastolic pressure may both reduce the coronary flow and increase the myocardial demand. A distended left ventricle has a higher wall stress according to the law of LaPlace. Wall stress is an important determinant of oxygen demand and may therefore explain the difference in LAD flow between pump on and pump off. Within seconds after the pump was stopped, the LAP increased in all the animals, indicating an increased left ventricular end-diastolic volume (Figures 5 and 6). 27,28 The regional flow behind the pump was measured in the right femoral artery and in the left renal vein. After induction of heart failure, there was neither a change in flow with the pump on nor after 30 seconds when it was stopped. The increased pressure behind the pump is expected to have caused a higher flow if the vessels had been passive conduits. Active autoregulation can explain the nearly constant flow in the kidney and in the femoral artery. This conclusion is supported by a parallel fall in flow in both the femoral artery and the renal vein 10 seconds after stopping the pump with a return to “pump-on” values after 30 seconds, despite the lower perfusion pressure (Figure 7).
The current study shows that the new device is able to reduce the afterload of the left ventricle in an acute heart failure model. The pressure unloading effect seems to be similar to the effect accomplished with vasodilating therapy. The most pronounced difference between these two therapeutic options relates to the difference in pressure in the lower part of the aorta: vasodilation decreases the pressure while the pump increases it. This is important in a case of low perfusion pressure for the internal organs.
Despite the large pressure difference within the aorta with the pump on, we did not observe any radical change in regional distribution of flow. Three of the four flow areas investigated exhibited autoregulation, and our results indicate that these mechanisms were still active with the pump on. Future studies are needed with the pump in conditions of low blood pressure leading to loss of autoregulation. The crucial question is how large the pressure reduction can be in the upper part of the aorta and still be tolerated by a patient.
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