(Circulation. 1995;92:1627-1633.)
© 1995 American Heart Association, Inc.
Articles |
Endocardial Carbon-Braid Electrodes
A New Concept for Lower Defibrillation Thresholds
Presented in part at the 66th Scientific Sessions of the American Heart Association, Atlanta, Ga, November 8-11, 1993, and published in abstract form (Circulation. 1993;88[pt 2]:I-593).
From the I. Medizinische Klinik, Technische Universität München, Germany (E.U.A., P.C.F., E.M.); the Departments of Medicine and Pathology, Duke University Medical Center, Durham, NC; the Engineering Research Center for Emerging Cardiovascular Technologies and the Department of Biomedical Engineering of the School of Engineering, Duke University, Durham, NC.
Correspondence to Raymond E. Ideker, MD, PhD, Cardiac Rhythm Management Laboratory, Volker Hall G78A, 1670 University Blvd, Birmingham, AL 35294-0019.
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Abstract |
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Background In the treatment of patients with life-threatening ventricular arrhythmia, transvenous implantable cardioverter/defibrillators provide significant advantages over devices requiring a thoracotomy. This study tested the hypothesis that a new carbon-fiber electrode, designed at the Technische Universität in Munich, Germany, has a lower defibrillation threshold (DFT) than standard transvenous defibrillation electrodes.
Methods and Results In 8 mongrel dogs (weight, 25.2±0.8 kg; heart weight, 192±19 g), we examined the efficacy and electrical characteristics of a right ventricular endocardial carbon prototype defibrillation electrode (9.5F, 4.4-cm2 surface) compared with a standard CPI 0062 Endotak electrode and a Medtronic 6966 Transvene endocardial right ventricular defibrillation electrode. The new electrode consists of 24 braided, tubular carbon filaments, each containing 1000 highly isotropic carbon fibers of 7-µm diameter, yielding a theoretical electrical surface of 480 cm2. The DFTs were determined in random order between each of the three right ventricular electrodes and a subcutaneous wire array anode placed on the left thorax. A standard step-down/-up DFT protocol of 20-V shock steps was applied. Two different biphasic waveforms with a 1-ms delay between phases were tested: 3.2-ms first phase/2.0-ms second phase, and 6.0-ms first phase/6.0-ms second phase. For the 3.2/2.0-ms waveform, we found a significantly lower DFT for the carbon lead (4.96±1.58 J) compared with the CPI 0062 (6.93±1.67 J) and the Medtronic 6966 (7.49±0.99 J) leads. For the 6.0/6.0-ms waveform, the DFT for the carbon electrode (5.97±2.09 J) was significantly lower than for the Medtronic 6966 lead (8.55±1.93 J) but not for the CPI 0062 lead (6.30±1.41 J). The impedance with carbon was lower than with the other two leads for the 6.0/6.0-ms waveform but not for the 3.2/2.0-ms waveform. For the carbon electrode, the 3.2/2.0-ms waveform had a lower DFT than the 6.0/6.0-ms waveform.
Conclusions The present canine study found a lower DFT for a new carbon electrode compared with DFTs for endocardial defibrillation electrodes made of standard metal. Further long-term animal studies and clinical studies are needed to determine whether carbon materials and braided-lead technology are practical and beneficial in patients.
Key Words: pacemakers • death, sudden • defibrillation
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Introduction |
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Implantable cardioverter/defibrillators (ICDs) are effective in reducing the rate of cardiac sudden death, in most studies down to 1% and 5% at 1 and 3 years, respectively.1 2 3 4 5 6 7 Ongoing randomized trials will determine whether ICDs decrease mortality from all causes.8 Biphasic shocks, transvenous leads, pectoral implantation of defibrillators, and electrode configurations using the defibrillator case as one pole ("active can" configuration) are all improving implant technology.9 10 11 12 13 14 15 16 17 18
Improvements in lead systems could increase sensing sensitivity, lower defibrillation and pacing thresholds, and, in conjunction with advances in capacitor and battery technology, may reduce the size of future devices. While past research has focused on the size and shape of defibrillation electrodes and on shock strength, duration, waveform, and polarity,19 20 21 22 23 less attention has been paid to the electrode material used for transvenous leads.24
In cardiac pacing, carbon has been shown to enhance the efficiency of energy transfer from the pacing tip to the adjacent myocardium.25 26 27 28 29 The primary beneficial properties of carbon are low thrombogenicity,30 low impedance, and virtually no polarization.31 In analogy to pacing, where the introduction of carbon and of low-polarization electrodes has substantially contributed to modern 25-g pacemakers, we hypothesized that defibrillation electrodes made of nonmetallic materials with low polarization and enhanced total surface structure may improve the energy transfer and sensing characteristics of defibrillation electrodes, thus contributing to a reduction in defibrillation energy requirements.
Our study aimed to determine the defibrillation threshold (DFT) with a new carbon-braid right ventricular (RV) defibrillation electrode in 8 dogs and to further our understanding of the contribution of the electrode material to the overall performance of the transvenous shocking lead. The new carbon electrode was compared with two conventional platinum-iridium transvenous RV defibrillation electrodes, the CPI 0062 Endotak and the Medtronic 6966 Transvene. Two different biphasic waveforms, 3.2/2.0 ms and 6.0/6.0 ms, were explored.20
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Methods |
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The present study was approved by the Institutional Animal Care and Use Committee at Duke University. It conformed to the guidelines of the American Heart Association on research animal use, adopted November 11, 1984.
Carbon Electrodes
In conjunction with the Electronic
Engineering Department at the
Technische Universität München, we conducted a search for
appropriate carbon materials to be used as defibrillation electrodes.
On the basis of our in vitro studies,32 we selected
electrically conductive carbon fibers manufactured by BASF. Each fiber
consists of 1000 nearly isotropic filaments of 7 µm diameter. Each
filament comprises 1000 molecular chains of 10 nm diameter; each chain
is twisted to form a helical structure. The excellent conductivity of
this carbon material is achieved by a special manufacturing process to
make conduction nearly isotropic.
The material used has a high tensile
modulus (23.9x103
kg/mm2, or three times the tension of
steel).33 Additionally, it shows a very smooth surface
(Fig 1
) and high flexibility.34
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Conduction is further enhanced with a braid configuration of 24 carbon filaments and 8 platinum-iridium filaments.35 36 37 The interwoven platinum-iridium strands consist of nine 25-µm-thick individual platinum-iridium single fibers twisted together to form one strand of 85 µm thickness.
Calculations have shown that the theoretical electrical surface (480 cm2) of all fibers together amounts to 109 times the outer geometric surface of this lead (4.4 cm2). The prototypes tested in this study were manufactured and tested according to our specifications (B. Braun Cardiovascular). Preliminary mechanical bench testing showed that no fractures or damage to the electrode could be detected after 7.5 million cycles of a flex-fatigue test in which the fibers were bent from +90° to -90° with a radius of curvature of 6 cm at a frequency of 1.2 Hz.
Animal Preparation
In 8 mongrel dogs, anesthesia was induced
with
intravenous pentobarbital (30 to 35 mg/kg body wt) and
maintained with a continuous infusion of pentobarbital at a rate of
approximately 0.05 mg/kg per minute.38 Succinylcholine (1
mg/kg) was also given intravenously at the time of
anesthesia induction. Supplemental doses of succinylcholine
(0.25 to 0.5 mg/kg) were given as needed to maintain muscle relaxation.
The animals were intubated with a cuffed endotracheal tube and
ventilated with room air and oxygen through a Harvard respirator
(Harvard Apparatus Co). A peripheral
intravenous line was inserted, and normal saline was
continuously infused. A femoral arterial line was placed
for hemodynamic monitoring as well as for
arterial blood gas analysis and electrolyte
measurements. Normal metabolic status was maintained
throughout the study by taking blood samples every 30 to 60 minutes and
correcting any abnormal values. ECG leads were applied for continuous
ECG monitoring of lead II. For delivery of an external rescue shock,
cutaneous defibrillation patches (R2 Medical Systems Inc) were placed.
Body temperature was measured and maintained at 35°C to 37°C with a
thermal mattress and heat lamp. At the end of the study, euthanasia was
induced with an injection of potassium chloride. The heart was then
removed, weighed, and macroscopically examined for carbon-fiber
particles.
Electrode Configurations
Three different endocardial
electrodes were positioned in random
order in the RV: (1) a 9.6F CPI 0062 Endotak defibrillation electrode
with a 3.7-cm-long RV coil and a pacing electrode tip 6 mm distal from
the RV electrode; (2) a 10.5F Medtronic 6966 Transvene electrode with a
5-cm-long defibrillation electrode 25 mm proximal to the pacing tip;
and (3) a 9.5F carbon-platinum-iridium prototype electrode with a
5-cm-long defibrillation electrode 15 mm proximal to the tip (Fig
2).
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A bipolar pacing catheter for induction of
fibrillation and backup
pacing was positioned through a left jugular incision into the apex of
the RV and left in place throughout the trial. One of the three leads
was randomly chosen to be placed under fluoroscopic guidance through a
right jugular incision at the RV apex. The bipolar pacing catheter was
used as a reference to place each subsequent electrode in a position as
close to the previous electrodes as possible within the RV. A silver
wire array (each wire 18 cm long) with a surface area of 6
cm239 40 41 was tunneled
subcutaneously in the
left lateral chest wall (Fig 3). The same three
endocardial electrodes were used in all eight canine trials.
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Defibrillation Protocol and Data Acquisition
Defibrillation
success was determined by a threshold
value.42 43 44 Eight animals underwent
defibrillation trials.
After the subcutaneous wire electrode was placed, the three endocardial
electrodes were randomized following a randomization table and the
first electrode was positioned in the RV; the DFT was determined for
the 6.0/6.0-ms biphasic waveform and then for the 3.2/2.0-ms waveform.
Thereafter, the second and third endocardial leads were positioned and
tested in the same manner.
Defibrillation testing was performed by following a modified up/down protocol45 starting with a leading-edge voltage of 400 V. The initial step size was 20 V. If the first shock failed, shocks were incremented in 20-V steps until a successful defibrillation occurred. If the first shock succeeded, shocks were decremented in 20-V steps until a shock failed. The DFT was defined as the lowest voltage required to defibrillate.
Ventricular fibrillation was induced by 60-Hz alternating current through the RV apex pacing electrode with the return electrode on the dog's chest wall. Fibrillation was allowed to continue for 10 seconds46 47 before defibrillation was attempted. A failed shock was followed by a rescue shock of higher voltage delivered between the electrodes. If the rescue shock failed, it was followed by external defibrillation given by a Life-Pak 8 defibrillator (Physio-Control Corp) via the cutaneous patches.
A minimum of 4 minutes elapsed between each fibrillation-defibrillation attempt. Fibrillation was not reinitiated until blood pressure and heart rate returned to normal. The defibrillation electrodes were connected to an external defibrillator (model HVS-02, Ventritex, Inc). The defibrillator delivered a single-capacitor biphasic shock from a 150-µF capacitor bank. The truncated exponential biphasic shock utilized a second phase of opposite polarity to the first with the second phase leading-edge voltage equal to the first phase trailing-edge voltage rounded to the nearest 10 V. The biphasic duration was set at 3.2/2.0 ms or 6.0/6.0 ms with a 1-ms delay between phases. During the first phase, the endocardial electrode was the cathode and the subcutaneous wire array was the anode. The actual current and voltage waveforms were digitized at 20 kHz and recorded by a Data Precision 6100 waveform analyzer. Signal analysis software within the analyzer was used to obtain impedance and energy measurements. The data were transferred to a Sun workstation and to a Macintosh computer for further analysis.
Documentation
The carbon-fiber electrode was microscopically
examined and
photodocumented after each trial.
Statistical Analysis
Data are expressed as mean±SD for
energy, voltage, current, and
impedance for all DFT determinations. Statistical significance was
assessed by ANOVA (Super ANOVA, Abacus Concepts, Inc) for multiple
comparisons and by the Friedman and Wilcoxon tests. A value of
P
.05 was considered significant.
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Results |
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A total of 48 DFT determinations was performed in 8 dogs. The Table
summarizes the results with the three different RV
electrodes and the two waveforms regarding total energy, impedance,
current, and voltage of the first phase at the DFT.
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The lowest absolute total energy at DFT was found with the
carbon-platinum-iridium prototype electrode and the short waveform
(4.96 J). This energy was 28% lower than that found with the CPI
Endotak 0062 lead and the short waveform (6.93 J) and 34% lower
compared with the Medtronic Transvene 6966 lead (7.49 J) and the short
waveform (P.05).
Fig 4A depicts the individual results obtained with the
3.2/2.0-ms biphasic waveform regarding energy, impedance, current, and
voltage at DFT. The results for each dog and the mean values associated
with each of the three different electrodes are shown. Total energy
with the carbon prototype electrode is significantly lower compared
with the CPI 0062 and Medtronic 6966 electrodes. The same holds true
for first-phase current and voltage at DFT.
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Fig 4B
shows the results for the 6.0/6.0-ms waveform. Here,
a
significant difference in total energy is found between the carbon and
Medtronic 6966 electrodes. With this longer waveform, a significantly
lower impedance is associated with carbon compared with the CPI 0062
and the Medtronic 6966 electrodes.
Fig 5 compares the two waveforms when used together with
the three different leads. For carbon, the energy requirements at DFT
were significantly lower with the 3.2/2.0-ms waveform compared with the
longer 6.0/6.0-ms waveform (4.96 J versus 5.97 J, P.05).
This represents a reduction of 17%; however, no significant
difference regarding the influence of waveform was found for the CPI
0062 and Medtronic 6966 electrodes. The waveform of the shock had
virtually no influence on peak voltage, peak current, and impedance
with the carbon electrode.
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Discussion |
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For patients requiring defibrillation energies above or near the maximum the ICD can deliver, efforts have been made to reduce DFTs by changing lead systems, lead configurations, patch position and polarity, defibrillation waveform,15 19 20 and location of an RV electrode.17 But these efforts are not always successful. Consequently, about 1% to 9% of patients48 49 do not receive the device.
Therefore, ways to reduce the DFT are helpful not only in reducing the size and physical characteristics of the defibrillator but also in increasing the number of patients who may benefit from a simple transvenous system. In addition, lower thresholds may allow greater device longevity and would give a higher safety margin, which could be important in those patients who develop high thresholds after implantation.1 48
Previous findings also demonstrated that decreasing the defibrillation shock voltage produces less risk of cardiac damage.50 The risk of refibrillation from high voltage gradient areas51 also decreases with lower energies. Additionally, a device that delivers less energy would require less time to charge, thus reducing the period that a cardiac patient may be without blood flow. Lower DFTs could also permit more shocks before the battery becomes depleted.
The positive experience with carbon in cardiac pacing prompted us to develop a defibrillation electrode based on carbon fibers.34 Our rationale was that by increasing the total internal electrical surface of the coil used for defibrillation in relation to the outer dimensions of the electrode, the transmission of the shock may be facilitated so that impedance is lowered. Calculations of the theoretically possible total electrical surface of the carbon braid (24 strands of carbon filaments each containing 1000 fibers of 7 µm diameter, each braided with a gear of 45%) yielded an increase of total internal surface of all fibers in relation to the outer geometric surface of the lead by a factor of more than 100.37 This number is a theoretical upper limit; the actual effective increase in surface area is probably smaller, but our animal results suggest an increased electrical surface available for current transmission during shock delivery similar to our in vitro findings (references 32 and 37 and F. Evans et al, unpublished data, 1995). In addition, it has been shown that carbon has low thrombogenicity and may induce less tissue reaction than conventional metallic materials52 that may have a long-term effect on impedance.
The current density and potential gradient fields throughout the ventricles are thought to be of prime importance in determining the success or failure of defibrillation. Lowering impedance by changing electrode size or location may increase the DFT if it creates a less efficient shock field.50 53 54 55 Therefore, all possible care was taken to ensure the same placement for each of the three electrodes within each individual dog. The pacing catheter in the RV used for the induction of fibrillation was used as a marker for correct positioning of the three defibrillation electrodes. In addition, we designed the carbon defibrillation electrode to be of similar outer diameter as the two standard metallic defibrillation electrodes to avoid systemic errors caused by different electrode sizes. Nevertheless, the CPI 0062 electrode coil is shorter (3.7 cm long) than the Medtronic 6966 (5 cm long) and the carbon (5 cm long) leads, which may represent a limitation of this study. Both standard electrodes were selected on the basis of their widespread clinical application.12 14 17 56
Another study57 has shown that increasing the electrode length in a superior vena cava location from 5 to 8 cm has reduced impedance by 7%, leading to a reduction of 10% in the voltage at DFT. It is assumed that both impedance and field configurations contributed to the percentage reduction in voltage DFT.
In addition to an increase in surface area, a reduction in polarization may also translate to lower impedances. With metallic electrodes, polarization voltages range between 2% and 8% of the total voltage applied with defibrillation (references 57 and 58 and F. Evans et al, unpublished data, 1995). In general, the lower the voltage is for defibrillation, the higher the degree of polarization voltage. Since the polarization voltage is ineffective for defibrillation, the energy requirements for defibrillation with polarizing electrodes are higher.
Yet another aspect of polarization is its effect on sensing. The initial experience with the CPI Endotak 0062 electrode incorporating a 6-mm distance from tip to shocking coil has shown that redetection of fibrillation was compromised.53 A redesign of this electrode has solved the problem by increasing the tip-to-coil distance from 6 to 12 mm (CPI Endotak 0072). Two factors are considered to cause difficulties in redetecting fibrillation: (1) the high voltage gradients in the myocardium near the electrode that may decrease the electrical signal generated by this tissue,59 and (2) polarization voltages that make postshock sensing more difficult. With respect to both issues, nonmetallic electrodes may be beneficial.
Although a greater tip-to-coil distance seems beneficial for good sensing, a closer tip-to-coil distance may be beneficial for low DFTs. Animal studies41 60 have shown that sensing strategies and electrode location affect DFTs. One study41 found that moving the defibrillation coil away from the RV apex caused a significant increase in DFT voltage and energy for a total endocardial electrode system. A second study60 showed that the location of the shocking coil with respect to the distance from tip to coil (1.2 versus 2.0 cm) also had a significant impact on the DFT. DFTs with the 2.0-cm distance were significantly higher in energy (39%).
An explanation for this finding may be that the area of the lowest potential gradients and the earliest activations after unsuccessful shocks for most transvenous lead systems is the left ventricle (LV).55 61 Hence, the closer the defibrillation coil is to the septum and to the LV, the lower the DFT.51 55 62 It is thus desirable to have a defibrillation electrode that incorporates a close tip-to–shocking coil distance without compromising postshock sensing. The differences in DFT between the Medtronic 6966 and the other two electrodes may be explained in part by the longer distance of 25 mm between the tip and the defibrillation coil compared with 6 mm in the CPI 0062 lead and 15 mm in the carbon lead.
A major thrust in the design of transvenous defibrillation leads has
been to reduce their diameter. Smaller leads are easier to apply and
may have less risk for insulation defects and fractures when applied by
a puncture technique. On the other hand, a reduction of the coil and
lead diameter results in a smaller surface and shadow area for
defibrillation. Past modeling studies have shown that decreasing the
diameter of a 5-cm-long coil in blood from 3 to 2 mm increased its
impedance by 
9%.57 To compensate for this and still
have a small defibrillation lead diameter with a large electrical
surface for defibrillation, the concept of braided design seems to
provide a good solution.
As hypothesized, the carbon electrode had a lower impedance than the other two electrodes, but only for the longer 6.0/6.0-ms waveform. The impedance was not lower for the carbon electrode for the shorter 3.2/2.0-ms waveform. Intriguingly, the effect of the carbon electrode on lowering the DFT was greater for the short waveform than for the long waveform, even though the impedance was not significantly lower for the carbon electrode with the short waveform.
Recent studies have shown that a capacitance of 60 to 90 µF gives the best energy yield63 64 but with reduced capacitance, shorter pulse widths are needed to achieve a similar tilt of 50% to 70%, which has been described as most energy efficient.65 All possible care was taken to ensure the same placement for each of the three electrodes within each individual dog. Thus, shorter pulse waves seem to be more likely in the future on the basis of present developments in reducing the size in capacitance. The specific reasons that carbon electrodes performed significantly better with the short waveform need further investigation.
This study reports the potential benefit of improved material and surface structure for endocardial defibrillation electrodes. Despite the positive results seen with this trial, there are many unanswered questions associated with this new material and design that need to be answered, eg, long-term stability, biocompatibility, and thrombogenicity. The transfer of the results from animal data to human application must also be questioned. Further in vitro and in vivo studies are needed to confirm or reject the concept of carbon materials and of braided electrode structure for transvenous defibrillation electrodes.
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Acknowledgments |
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This study was supported in part by National Institutes of Health research grants HL-42760 and HL-44066, American Heart Association Grant NC-93-SA-05, and The National Science Foundation Engineering Research Center grant CDR-8622201. The authors wish to thank Ellen Dixon-Tulloch, Jenny Hagler Knight, Sharon Melnick, and Robert Walker for their technical assistance, Herman Naarman, PhD, from BASF for providing the carbon raw material, and Steve Huntley from Angeian, Inc of Plymouth, Minn, for assistance in the fabrication of the carbon electrode prototypes.
Received February 2, 1995; accepted February 28, 1995.
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