E A A F t w t a r n a t g t t d w a t e 1 w O w d w p o c t m w a a d a “ p e m D C 1 DITORIAL COMMENTARY scending waveforms: The ramp to the Holy Grail? dam Zivin, MD rom Seattle Cardiology and Swedish Heart and Vascular Institute, Seattle, Washington. r c t f t c t o t f r b d a a q n i i d m v o w t n i r e c a i b ( p t o c b w g In this issue of Heart Rhythm, Shorofsky et al1 present he first human data on the efficacy of an ascending ramp aveform for ventricular defibrillation. As the authors note, he theoretical advantage to defibrillation waveforms with n ascending first phase has been previously described and ecently verified in a guinea pig model.2,3 Muscle tissue has both resistive and reactive compo- ents, impeding an instantaneous change in cellular volt- ge. Thus, energy transfer and “charging” of cardiac issue are more efficient if the delivered energy is applied radually. Practical engineering concerns have limited he application of this theory. In contrast to a conven- ional, exponentially decaying waveform, which uses the ischarge properties of the storage capacitor to define the aveform shape, with present technology, generation of high-voltage ascending ramp requires the addition of he bulk of an inductive element to the circuit. Using a “dummy” hot can and external waveform gen- rator, a modest but statistically significant reduction of 8% in delivered energy at defibrillation threshold (DFT) as achieved with an optimized ascending ramp waveform. f interest, the benefit was seen in about half of the patients, ith one fourth actually having a higher DFT. The quest for lower DFTs has been a goal of implantable efibrillator design from the outset. The first step forward as the addition of a second reverse-polarity phase (a bi- hasic waveform).4 This was followed by the development f nonthoracotomy unipolar systems and variations in lead onfiguration.5,6 Defibrillation of cardiac muscle requires delivery of a hreshold current—the “rheobase”—and development of a inimum voltage gradient. The curve of current vs pulse idth for tissue stimulation is a decaying exponential curve pproaching rheobase at infinite pulse width. However, for pulse of given energy, the effective defibrillating current elivered into a reactive load (such as cardiac muscle) peaks t a pulse width about twice the time constant—the chronaxie”—of the tissue, a finding reproduced in the resent study. As long as this effective current is at least qual to the rheobase, this point on the curve identifies the ost efficient pulse width and therefore the point of lowest FT. As the pulse width increases further, delivered cur- Address reprint requests and correspondence: Dr. Adam Zivin, Seattle ardiology, 1229 Madison Street, Suite 1500, Seattle, Washington 98122. iE-mail address:
[email protected]. 547-5271/$ -see front matter © 2005 Heart Rhythm Society. All rights reserved ent, and therefore the probability of defibrillation, de- reases and energy is “wasted.”7 This is the rationale behind runcation of the waveform. For capacitive-discharge wave- orms, the most efficient energy transfer from device to issue occurs when the circuit time constant, the product of apacitor size and lead impedance, is in the range of the issue time constant (the chronaxie). For shock impedances f 50 �, this is approximately 60 �F.8 A capacitor smaller han this leads to higher DFT, whereas the DFT “penalty” or larger capacitance is relatively small within a broad ange and permits higher stored energy for a greater defi- rillation safety margin. Typical capacitor size in current evices is 110 to 120 �F. In the study by Shorofsky et al, waveforms were gener- ted by an external waveform generator. To reduce implant- ble cardioverter-defibrillator (ICD) size, the circuitry re- uired to generate this waveform would need to be at least o bulkier than the current “simple” capacitor circuits. If an nductor is required, this may be problematic. Second, even f mean DFTs could be reduced, one still would want a evice capable of delivering a shock with an adequate safety argin. At present, with available capacitor technology and oltage handling constraints, this would require a capacitor f about 100 �F. Therefore, to reduce ICD generator size hile maintaining a comfortable margin of safety may ac- ually depend more on capacitor, battery, and circuit tech- ology than on reduction of DFT per se. A potentially more ntriguing possibility is not whether device size can be educed by lowering DFTs but if, for any given delivered nergy, this waveform is more effective. For ICDs, this ould be valuable for dealing with patients with high DFTs nd allow greater safety margins to accommodate changes n patient substrate (drugs, infarcts, etc.). For external defi- rillators including automatic external defibrillators AEDs), could this translate into successful defibrillation of atients with longer “down times”? Other techniques for reducing DFT are under investiga- ion but, as with the ascending ramp waveform, at the cost f increased system complexity.9 Use of smaller, fully dis- harging capacitors may reduce circuit complexity and size, ut without any reduction in DFT.10 We may be approaching a point of diminishing returns ith respect to reducing device size, but the potential for reater efficacy from this and other areas of investigation s promising. . doi:10.1016/j.hrthm.2005.01.007 R 1 396 Heart Rhythm, Vol 2, No 4, April 2005 eferences 1. Shorofsky SR, Rashba E, Havel W, Belk P, DeGroot P, Swerdlow C, Gold MR. Improved defibrillation efficacy with an ascending ramp waveform in humans. Heart Rhythm 2005;2:388–394. 2. Hillsley RE, Walker RG, Swanson DK, Rollins DL, Wolf PD, Smith WM, Ideker RE. Is the second phase of a biphasic defibril- lation waveform the defibrillating phase? Pacing Clin Electro- physiol 1993;16:1401–1411. 3. Guan D, Malkin R. Analysis of the defibrillation efficacy for 5-ms waveforms. J Cardiovasc Electrophysiol 2004;15:447–454. 4. Dixon EG, Tang ASL, Wolf PD, Meador JT, Fine MJ, Calfee RV, Ideker RE. Improved defibrillation thresholds with large contoured epicardial electrodes and biphasic waveforms. Circulation 1987;76:1176–1184. 5. Bardy GH, Johnson G, Poole JE, Dolack GL, Kudenchuk PJ, Kelso D, Mitchell R, Mehra R, Hofer B. A simplified, single-lead unipolar transvenous cardioversion-defibrillation system. Circulation 1993; 88:543–547. 6. Bardy GH, Dolack GL, Kudenchuk PJ, Poole JE, Mehra R, Johnson G. Prospective, randomized comparison in humans of a unipolar defibril- lation system with that using an additional superior vena cava elec- trode. Circulation 1994;89:1090–1093. 7. Kroll MW, Lehmann MH, Tchou PJ. Defining the defibrillation dos- age. In: Kroll MW, Lehmann MH, Implantable Cardioverter Defibril- lator Therapy: The Engineering-Clinical Interface. Norwell, MA: Klu- wer Academic Publishers, 1996:63–88. 8. Swerdlow CD, Brewer JE, Kass RM, Kroll MW. Application of models of defibrillation to human defibrillation data: Implications for optimizing implantable defibrillator capacitance. Circulation 1997;96: 2813–2822. 9. Dosdall DJ, Rothe DE, Brandon TA, Sweeney JD: Effect of rapid biphasic shock subpulse switching on ventricular defibrillation thresh- olds. J Cardiovasc Electrophysiol 2004;15:802–808. 0. Yamanouchi Y, Fishler MG, Mowrey KA, Wilkoff BL, Mazgalev TN, Tchou PJ. New approach to biphasic waveforms for internal defibril- lation: fully discharging capacitors. J Cardiovasc Electrophysiol 2000; 11:907–912.