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Defibrillator shocks ranging in intensity from three to nine times current threshold were delivered to four isolated, metabolically supported, beating canine hearts. The shocks produced an immediate, current-dependent depression of left ventricular isovolumic systolic pressure. This depression was transient, reproducible, and was followed by a transient overshoot in ventricular systolic pressure. Then 1 mg propranolol hydrochloride in 1 ml H2O was injected into the coronary arterial supply of the isolated heart, and the shocks were repeated. The magnitude of the immediate cardiac depression after shock was unchanged; however, the time required for full recovery of left ventricular systolic pressure to a pre-shock control value was prolonged. In addition, the transient overshoot in ventricular systolic pressure seen in the untreated state was absent. These results are consistent with the hypothesis that defibrillatory shocks produce a direct activation of cardiac adrenergic nerves, which aid in recovery of ventricular contractility following defibrillator shocks.

A promising therapy for ventricular fibrillation--a life-threatening cardiac arrhythmia--is implantation of an automatic defibrillator. A critical component of such a device is the system that detects the presence of ventricular fibrillation. Automatic systems for detecting ventricular fibrillation have been tested with arrhythmias produced by electric shocks in normal canine hearts, but have not been tested with spontaneous arrhythmias in hyperirritable hearts. We have developed an animal model to create arrhythmias without electrical stimulation and have used it to test our automatic defibrillator. This model permits evaluation of both reliability to diagnose VF and reliability to reject other tachyarrhythmias.

Therapy for severe chronic lung disease currently includes the administration of supplemental oxygen to prevent breathlessness and tissue hypoxia. Although effective, this therapy is unnecessarily costly, because oxygen is administered to the patient during expiration as well as inspiration. To eliminate this inefficiency, a delivery system that senses the inspiratory effort and delivers oxygen to the patient only during inspiration was developed. The 11 X 5 X 8-cm flow control unit attaches easily to a portable oxygen supply. The components of the system have an expected life of five years, and the 9-V battery provides power for about one month of use. Manual controls permit accommodation to the respiratory pattern of the patient. Preliminary evaluation of the system showed that its effectiveness in producing tissue oxygenation is similar to that of continuous oxygen systems. The system has potential applications in ambulatory oxygen therapy and in other clinical settings to improve the cost/benefit ratio of oxygen treatment.

A transient, dose-dependent cardiac depression was produced by defibrillator shocks in an isolated, working canine heart preparation perfused with oxygenated arterial blood from a support dog. Accompanying this depression was an efflux of potassium (K+), forced out of the myocardial cells by the passage of defibrillating current. The transient increase in extracellular K+ concentration was recorded graphically in the venous outflow. It was found that 5-msec rectangular wave shocks, from three to ten times defibrillatory current threshold, released dose-related pulses of K+. It is concluded that because K+ is a myocardial depressant, at least part of the myocardial depression after defibrillation is caused by the release of K+ from the myocardial cells.

Because the automatic implantable defibrillator may be used in patients having ischemic heart disease, it is important to know whether myocardial ischemia changes the threshold for ventricular defibrillation under experimental conditions simulating automatic internal defibrillation. We determined changes in ventricular defibrillation threshold following coronary occlusion, using an electrode catheter designed for use with an automatic implantable defibrillator. Acute myocardial ischemia was produced without thoracotomy in 10 dogs (experimental group) by embolization with a plastic bead injected via a catheter into the left coronary artery. A control group of 4 dogs had only saline injected into the artery. Defibrillation threshold was measured at 15-min intervals from 1 hour before embolization to 2 hours after embolization. In the control group, voltage, current, energy, and impedance were unchanged after injection of saline into the coronary artery, and india ink perfusion revealed no ischemic areas. In the experimental group postembolization threshold current and energy were significantly higher than preembolization values: 0.47 vs 0.40 A/kg and 1.01 vs 0.80 J/kg, respectively (p less than 0.01). The magnitude of the peak change in threshold current after embolization was positively correlated (r = 0.79) with the size of the ischemic zone, determined by weighing unstained areas after india ink perfusion. Defibrillation threshold for a catheter electrode configuration increases for at least 2 hours following onset of acute myocardial ischemia. This finding must be accounted for in the design and use of an automatic implantable defibrillator.

A method is developed for determing actual values of circuit elements in a damped sine wave (Lown waveform) defibrillator solely from measurements of the output, using two or more power resistors and a storage oscilloscope. If a defibrillator containing capacitance C, inductance L, and internal resistance RI, is discharged into increasing 5- to 100-ohm resistive loads R, it is shown for underdamped output waveforms that â = RI/2L + R/2L and ĉ = CRI + CR, where â = pi[t2tan(pit1/t2)], ĉ = 2â/[â2 + (pi/t2)2], t1 = time from onset to peak, and t2 = time from onset to first zero crossing of the output waveform on the oscilloscope trace. Linear plots of â vs R are constructed for seven defibrillators, and values of RI and L computed as intercept/slope and 1/(2 slope) respectively. C is given by the slope of a linear plot of ĉ vs R. Delivered energy is accurately predicted as stored energy X R/(RI + R).

Technology for local heat therapy of cancer is evolving rapidly at a number of technologically diverse and geographically scattered institutions and companies. No single technology is superior to others in all applications, and no single company, laboratory, or research group has all the answers. An ideal system would provide focused heating at depth in a predictable fashion, with little probability of generating undesired hot spots in normal tissues and little interference with monitoring equipment. Existing systems approximate this ideal to different degrees, depending on the anatomy and geometry of the tumor and its surrounding tissues. In the foregoing discussion the important problem of measuring temperatures in tumors and normal tissues has been slighted. At the present time, all thermometry is necessarily invasive, and there are limitations to the number of points at which temperatures can be measured utilizing percutaneously placed catheters as conduits for thermometers. However, further advances in the art, the science, and the technology of local heat therapy are likely to be forthcoming in the next few years from a diverse community of investigators and young companies who are following an interesting variety of approaches. Continued research and development in the spirit of constructive, rather than destructive, competition will certainly advance the field substantially--much to the benefit of patients. At present, however, clinical engineers should realize that hyperthermia therapy for cancer is still experimental. Despite the flurry of commercial activity, considerable caution should be exercised in the purchase and use of hyperthermia equipment.

Successful cancer therapy must selectively destroy tumor tissue while sparing the host's normal tissues. Local heat treatment can have such a selective effect because abnormalities in tumor blood vessels supply less oxygen to heat-stressed tumor cells and are less efficient in cooling tumor tissue by blood perfusion.

Local hyperthermia therapy for cancer can produce selective heating of solid tumors on the basis of known physical laws. If energy is deposited in the general region of the tumor, temperature tends to develop in the tumor higher than that in surrounding normal tissues. The goal of therapy is to achieve cytotoxic temperature elevations in the tumor for an adequate period of time, without damaging nearby normal tissues. Several modalities exist for local heat treatment, of which radiofrequency and ultrasound offer the most promise for controlled, localized heating at depth. A paucity of blood flow in the tumor compared to that in adjacent normal tissues can enhance selective tumor heating considerably. The tumor types that have reduced flow in their central regions are especially vulnerable to heat therapy, both because they can be heated more efficiently and because hypoxic and acidotic tumor tissues are more susceptible to damage by heat. This effect is more pronounced in larger tumors, which have smaller surface-to-volume ratios and so lose heat less rapidly by thermal diffusion. Selective heat treatment of larger tumor masses with low blood perfusion, therefore, is physically practical and rational therapy. Vigorous research efforts are now underway at many centers to optimize this approach.

Our defibrillation-detecting system uses both ECG and right ventricular impedance change (delta Z). We studied the effect of catheter electrode spacing on delta Z in 10 dogs with body weights of 10 to ,5 kg and heart weights of 67 to 220 g. Impedance to 20-kHz 100--muA square waves was measured between two 1-cm-long electrodes mounted on a No. 12F catheter and wedged into the right ventricular apex. Catheters with spacings of 5 to 25 mm between the electrodes were tested during sinus rhyhm and ventricular fibrillation. During sinus rhythm the mean beat-to-beat delta Z was 23 +/- 4 omega using the 5-mm spacing. Wider spacing gave smaller, and hence less desirable, delta Z. Some delta Z signals were recorded during ventricular fibrillation, and 5-mm spacing was more sensitive to these than wider spacing. This resulted in some prolongation of the time between onset of fibrillation and application of the defibrillation shock, but should decrease false positive diagnosis of fibrillation. No clear relationship was observed between delta Z and body weight or heart weight. We conclude that the 5-mm spacing is best for detection of pumping by the catheter-impedance method in hearts of this weight range.

The feasibility of combining local heat treatment with whole-body hypothermia in an effort to improve therapeutic gain was assessed. Superficial, nonperfused phantom tumors were fashioned in eight anesthetized mongrel dogs by transplantation of the spleen from the abdomen to a subcutaneous site on the hindlimb. After pretreatment of the animal with the vasodilator hydralazine (0.5 mg/kg, IV) to enhance normal tissue perfusion, the spleen implant was heated with a 2450-MHz microwave diathermy apparatus, first with the animal's core body temperature in the normal range (39 degrees C) and then after the animal had been packed in ice to reduce core temperature to 30 degrees C. Applied power density and temperatures in both the phantom tumor and underlying muscle tissue were recorded during brief interruptions of diathermy until steady-state temperatures had been achieved. Under normothermic conditions with time-averaged applied power of 0.038 W/ml to phantom tumor and 0.014 W/ml to underlying muscle, tumor temperature rose to 45.9 +/- 1.8 degree C, while muscle temperature remained at 40.5 +/- 0.7 degree C. During whole-body hypothermia applied power could be increased to 0.114 W/ml in phantom tumor and to 0.025 W/ml in muscle. Muscle temperature rose only to 33.8 +/- 1.6 degree C, while that of the nonperfused phantom tumor rose to 53.6 +/- 4.3 degrees C with systemic hypothermia.(ABSTRACT TRUNCATED AT 250 WORDS)