Percutaneous Vascular Access

History of Aortic Counterpulsation

The concept of arterial counterpulsation was introduced in 1961 by Clauss and coworkers and involved use of an external “ventricular” chamber that filled with blood from a catheter in the iliac artery and was subsequently compressed by a piston. Compression of the “ventricle” was synchronized to either the QRS complex of an electrocardiogram (ECG) or the impulse of a pacemaker, so that a counter pulse of blood was delivered into the arterial system during diastole. It was demonstrated in dogs that cardiac stroke work and LV end-systolic pressures could be substantially reduced with the use of a counterpulsation into the aorta. The following year, Moulopaulus and associates adapted the model to create an intra-aortic balloon pump (IABP) that could provide a similar counterpulsation without the need for blood reservoirs. The investigators used a balloon that was rapidly inflated and deflated with carbon dioxide during native diastole. The IABP was subsequently adapted and described for clinical use by Kantrowitz and colleagues in 1968.

The original polyurethane balloon measured 1.8 cm in diameter by 14.8 cm in length when inflated (helium was used because its low density allows rapid delivery to and from the balloon) and displaced 32 mL of blood. There is little difference in the modern IABP and that originally described, other than the availability of different-sized balloons (30- to 50-mL balloons) and subtle differences in the materials used to make the catheters. The extracorporeal components of the IABP now include an electronically controlled pump with a solenoid valve in continuity with a pressurized helium source. The valve controls the flow of helium into and out of the balloon at intervals timed to either pressure changes on an arterial transducer, ECG signals (i.e., the QRS complex), or a ventricular pacer signal. This timing of balloon inflation and deflation is critical to attain optimal physiologic benefit of the cardiac support.

The physiologic rationale for the efficacy of the IABP is that balloon deflation provides a rapid, synchronized reduction in impedance (afterload) during isovolemic LV contraction. This is followed by a rapid, synchronized increase in aortic pressure during isovolemic LV relaxation (diastolic augmentation) caused by balloon inflation. In combination, these events achieve two important goals. First, LV systolic unloading directly reduces stroke work, which in turn reduces myocardial oxygen consumption during the cardiac cycle. Second, diastolic augmentation raises arterial blood pressure and provides better coronary arterial perfusion during diastole, yielding increased oxygen delivery to the myocardium. The IABP does not directly move or redistribute blood flow; however, peak diastolic coronary flow velocity can be increased as much as 87% with IABP augmentation and peak diastolic flow velocity by as much as 117%. Since introduction into clinical use in 1968, the IABP has remained an important adjunct to supporting patients in cardiogenic shock. Myocardial recovery is promoted by the reduction of cardiac work and the simultaneous increase in myocardial oxygen supply. However, therapeutic success is dependent on the patient having a minimum degree of LV function that, in combination with IABP support, facilitates an adequate cardiac output to sustain end-organ function. When this minimal cardiac output is not met, alternative mechanical cardiac assistance must be considered.

History of Mechanical Assist Devices

The need for effective mechanical cardiac assist devices became apparent in the 1950s during the development of CPB for open-heart surgery. Initial attempts with prolonged postoperative CPB demonstrated that the bypass circuit was damaging to both end-organ function and blood constituents after several hours of use. The first attempt at isolated extracorporeal LV support was with a simple roller pump in 1962. Subsequently, femoral venous-to-femoral arterial CPB was successfully used by Spencer and colleagues in four patients with postcardiotomy cardiac failure.

Simultaneous to Spencer and colleagues' work with extracorporeal systems, DeBakey designed the first intracorporeal LV assist device (LVAD), the DeBakey blood pump. This device consisted of a Dacron-reinforced silicone rubber tube with an inner chamber of blood from the left atrium that was connected to the descending thoracic aorta. Pressurized air was instilled into the outer chamber by an external pneumatic controller to compress the inner blood chamber, timed to the R wave of the QRS complex. Blood flow was directed from the left atrium to the descending aorta with the use of ball valves at both the inflow and outflow ends of the device. The DeBakey blood pump was first used in a patient who died 4 days postoperatively of neurologic complications. A remodeled extracorporeal version was subsequently used for postcardiotomy failure in a 37-year-old woman after aortic and mitral valve replacements. The device was needed for 10 days, but the patient survived.

By 1972, investigators at the Texas Heart Institute had developed a pneumatically driven LVAD designed to be implanted in the abdomen. This device had a blood chamber compressed by pulses of air delivered into the pump by a percutaneous driveline. Modern devices have chamber compression that is electrically powered via percutaneous drivelines. Paracorporeal, pneumatically driven devices were a parallel development. Paramount to the evolution of these devices was the sponsorship of the Artificial Heart Program of the National Heart, Lung, and Blood Institute, which was chartered in 1964.

By the 1960s, continuous flow, as compared to pulsatile, pumps were under development. Over the subsequent 15 years, centrifugal pumps were perfected and introduced into clinical use. These pumps work on the principle of a forced, constrained vortex devised from three magnetic cones. They have been shown to be useful in a variety of clinical settings where short-term mechanical support is needed and an IABP is inadequate. Several types of small, axial-flow or rotary pumps have also been developed, including some that allow for percutaneous deployment. These are generally constructed of a magnetically suspended impeller that rotates at extremely fast rates (25,000 to 35,000 rpm). The axial rotary pump technology has some potential advantages over pulsatile devices; they are quite small with few moving parts and do not require a compliance chamber. The latest generation of rotary pump technology utilizes fully magnetically levitated rotors that completely eliminate the need for seals or bearings. This technology reduces the risk of damage to blood elements and may lead to lower rates of thromboembolism.

Counterpulsation/Intra-aortic Balloon Pump