The Cox-Maze (CM) procedure was first performed in 1987 by Dr James L. Cox.1 After several iterations, the CM III procedure was introduced and became the criterion standard operation for the treatment of atrial fibrillation (AF).2 Despite its success, this traditional “cut-and-sew” CM procedure was not widely adopted because of its technical complexity and required time on cardiopulmonary bypass (CPB).
The development of surgical ablation technologies has simplified the surgical treatment of AF. With the advent of ablation technology, it has become possible to replace most of the standard lesions of the original CM procedure with lines of ablation, making the procedure simpler and faster to perform.3 Multiple energy sources, including radiofrequency (RF), microwave, cryoablation, and high-intensity focused ultrasound, have been used to create lesions with varying degrees of success.4–8 Our group has demonstrated the ability of bipolar RF clamps to reliably create transmural lesions in chronic animal models and, as such, has predominantly used this technology in the clinical setting.9–13
The CM IV (CMIV) procedure, which combines bipolar RF ablation and cryoablation to replace most of the traditional atrial incisions of the CM lesion set, was introduced by our group14 in 2002. The use of ablation technology simplified the procedure, shortened operative duration, allowed for minimal-access techniques, and maintained the success of the CM III, with 89% freedom from AF and 78% freedom from AF off of antiarrhythmic medications at 1 year.3 However, despite its many improvements, the CMIV still requires CPB because the bipolar clamps require an arrested heart to create the left atrial lesion set.
The introduction of epicardial ablation technology attempts to address these issues by allowing linear lines of ablation to be placed through minimal-access techniques on the epicardial surface of the beating heart. Unfortunately, the inability of these devices to reliably create transmural lesions on the beating heart has limited their clinical application.7 Unlike the bipolar RF clamps, these technologies have been ineffective for a number of reasons. One major problem is that the endocardial surface is protected from thermal injury because of the heat sink of the circulating cavitary blood pool.15 Other issues that have been implicated in the shortcomings of epicardial ablation technology include the anatomic variability of pathologic atria, poor tissue-device contact, and the deficiencies of the particular technology.
To overcome these shortcomings, a novel epicardial RF ablation system (Cobra Fusion; Estech, San Ramon, CA USA) was developed (Fig. 1). Unlike previous devices, this technology delivers both monopolar and bipolar energy via a temperature-controlled ablation algorithm in an attempt to eliminate the heat sink of the endocavitary blood pool from an epicardial approach. The present study was designed to test the ability of this novel RF ablation device to create contiguous linear lines of ablation on the beating heart in an acute porcine study using an epicardial approach.
This study was approved by the Washington University School of Medicine institutional review board for animal research. All animals received humane care in compliance with the Guide for the Care and Use of Laboratory Animals published by the National Institutes of Health (publication no. 85–23, revised 1985).
The Cobra Fusion Surgical Ablation System (Estech, San Ramon, CA USA) consists of two suction-assisted epicardial RF ablation catheters that deliver both monopolar and bipolar energy. Two ablation systems are available: a six-electrode 150-mm ablation catheter designed to position the device around the pulmonary veins and a 50-mm ablation catheter, which contains two 50-W electrodes that terminate 5 mm from the end of the suction chamber. In this study, the 50-mm catheter was used because the 150-mm device does not conform to the anatomic constraints of the porcine atria. The ablation system also includes an Electrosurgical Unit, which provides power to the ablation catheter via a cable permanently attached to the handle, and a vacuum source controlled via a three-way stopcock that attaches to the ablation catheter via accessory tubing. Once the ablation catheter is affixed to the tissue, suction is maintained at −500 mm Hg throughout the ablation algorithm.
Unlike other commercially available RF ablation devices that use impedance algorithms, this device uses a temperature-controlled heating algorithm that uses two thermocouples per electrode, sampled at 20-millisecond intervals with real-time feedback to power, to attain and maintain the target temperature16 of 70°C. Further, using temperature-controlled ablation prevents char formation on the tissue-electrode interface, which can limit lesion depth. The settings used in this investigation were identical to those recommended clinically by the manufacturer and included two separate 60-second tissue ablations, the first of which consisted of bipolar RF energy directed from the active RF electrodes toward the integrated indifferent electrodes on the other side of the device (Fig. 2). Without disrupting the suction, the indifferent electrode pads are plugged into the receptacles on the Electrosurgical Unit, and the device is then switched to monopolar mode by selecting the corresponding button on the device handle. Monopolar RF energy is subsequently directed from the active RF electrodes on the device toward the indifferent electrode pads on the subject’s back.
Six domestic pigs weighing between 70 and 90 kg were used for this study. Each animal was premedicated with intramuscular ketamine (2.2 mg/kg), telazole (4.4 mg/kg), and xylazine (2.2 mg/kg); intubated; and anesthetized with 2% to 4% inhaled isoflurane. The animals were maintained on a volume-limited respirator, and end-tidal carbon dioxide, body temperature, heart rate, respiratory rate, blood pressure, and oxygen saturation were continuously monitored. Indifferent electrodes for use with monopolar ablation were placed on the posterior aspect of the torso. Arterial blood gases, blood counts, and electrolytes were monitored every 30 minutes and as needed throughout the procedure via catheters in the femoral artery and vein.
A median sternotomy was performed, and the heart was supported in a pericardial cradle. The azygous vein was isolated and ligated, the ascending aorta was exposed, and flow probes were placed around the superior and inferior vena cavae for continuous monitoring of cardiac output. A baseline activated clotting time was measured before a heparin bolus was administered, and heparin was redosed as needed to maintain an activated clotting time of greater than 350 throughout the procedure. Amiodarone was administered via bolus and subsequent constant infusion (10–12 mg/kg and 0.5 mg/kg per hour, respectively) for episodes of stable AF. Direct current cardioversion was performed for episodes of unstable AF. Phenylephrine was administered via bolus and/or constant infusion (100 μg and 1 μg/kg per minute, respectively) to maintain systolic blood pressure in the acceptable range throughout the procedure.
The experimental lesion set (Fig. 3) was carried out using the Cobra Fusion 50-mm Surgical Ablation System (Estech, San Ramon, CA USA), as described above. Five standard lesions were created in the following locations in every animal: anterior left atrial appendage, posterior left atrial appendage, left atrial free wall, posterior right atrial appendage, and superior vena cava. In several animals, the anatomy was such that the superior vena cava lesion extended on to the superior portion of the inferior vena cava. According to the specific anatomy of the animal, an additional lesion was variably created either on the posterior right atrial appendage or, if space allowed, on the anterior right atrial appendage. Two 60-second tissue ablations were performed at each site, as described previously. The goal temperature during both ablations was 70°C, and suction was maintained at −500 mm Hg throughout the ablation time. The maximum power obtained, the maximum and minimum temperatures obtained, and the presence of any temperature discrepancy of greater than 5°C between electrodes as measured by the ablation system were recorded for each ablation.
After the ablations were completed, the ends of each lesion were marked on the epicardial surface with 4–0 Prolene suture for later identification. The aorta was subsequently cross clamped, and 60 mL of 2% 2, 3, 5-triphenyl-tetrazolium chloride solution was perfused through the coronary circulation via infusion into the root of the aorta. The animal was then euthanized with antegrade infusion of 3 M of potassium chloride. Postoperatively, the heart was excised en bloc and soaked in 1% 2, 3, 5-triphenyl-tetrazolium chloride solution for 6 to 12 hours at 4°C to stain any viable myocardium.
After staining, the heart was evaluated macroscopically for evidence of charring, tissue rupture, and blood coagulation. The individual lesions were excised, and gross digital photography of the epicardial and endocardial surfaces was carried out in proximity to a caliper for later calibration. Large intracavitary trabeculae were divided before analysis. The lesions were subsequently cross-sectioned at 5-mm intervals in a plane perpendicular to the long axis of the ablation line. The individual cross sections were again digitally photographed in proximity to a caliper for later calibration. Lesion depth, lesion width, and tissue thickness were determined with a commercial software (Adobe Photoshop, San Jose, CA USA).17 Tissue thickness was measured at the site of transmurality or at the site of deepest lesion penetration in the case of nontransmural lesions. Measurements of lesion depth and width included all of the unstained area, starting at the border of the pink halo region surrounding each lesion. Both the transmurality of each individual tissue cross section and the transmurality of each 5-cm lesion created by the device were examined.
Continuous and categorical variables are expressed as mean ± SD and as number and percentage, respectively, unless otherwise specified. Comparisons were performed using a two-tailed Student t test for normally distributed continuous variables and the Fisher exact test for categorical variables. Univariate logistic regression was performed on clinically relevant variables including tissue depth, body temperature, cardiac output, heart rate, systolic blood pressure, diastolic blood pressure, oxygen saturation, the maximum and minimum temperatures reached in both the bipolar and monopolar settings, and the maximum power reached in both the bipolar and monopolar settings to determine independent predictors of lesion transmurality. Data analyses were performed using Systat 13 (Systat Software, Inc, San Jose, CA USA). P values less than 0.05 were considered statistically significant.
A total of 38 lesions and 270 tissue cross sections were examined. There was no charring of the epicardial surface and no clot identified on the endocardial surface for any ablation. There were no atrial wall perforations identified intraoperatively or on subsequent photograph analysis.
Representative photographs demonstrating the epicardial and endocardial surfaces of an intact lesion, as well as cross sections of tissue found to have both transmural and nontransmural ablations, are demonstrated in Figure 4.
The maximum power and the maximum and minimum temperatures obtained for each ablation were not significantly different between the transmural and nontransmural lesions (Table 1). The lesion dimensions were also recorded for each ablation (Table 1). As expected, the mean lesion depth was greater for the transmural (3.1 ± 1.3 mm) versus nontransmural lesions (2.2 ± 1.5 mm, P = 0.02); however, the lesion width was also greater for the transmural (10.6 ± 2.9 mm) versus nontransmural lesions (7.8 ± 2.8 mm, P = 0.001). The mean ± SD tissue depth of the transmural cross sections was 3.1 ± 1.3 mm, whereas the mean ± SD tissue depth of the nontransmural cross sections was 3.4 ± 2.1 mm (P = 0.57, Table 1).
The efficacy of this device is depicted graphically via a scatter plot (Fig. 5) of lesion depth versus tissue thickness. Transmurality was documented in 94% (254/270) of all cross sections, and 68% (26/38) of all lesions were transmural along the entire ablation line. There was a significant difference in lesion transmurality based on where the cross section was located within a lesion, with middle sections reaching 96% (187/195) transmurality, and end sections being 89% (67/75) transmural (P = 0.04).
No significant difference in lesion transmurality was identified on the basis of the anatomic location of the lesion itself on the heart (Table 2, P = 0.45 for the distribution).
Of 270 total cross sections, 16 were identified as being nontransmural. Of the cross sections located at the end of the ablation line, 11% (8/75) were found to be nontransmural, whereas only 4% (8/195) of the cross sections located within the line of ablation were found to be nontransmural (P = 0.04). Logistic regression analysis demonstrated that failure of the device to create transmural lesions was associated with only low body temperature (P = 0.015). Failure to create transmural lesions was not associated with cardiac output (P = 0.54).
In the surgical treatment of AF, the ability to consistently create transmural lesions is critical to preventing AF recurrence. Although previous studies of ablation technologies have demonstrated success in consistently creating transmural lesions with the use of CPB, epicardial ablation devices have historically had difficulty in creating transmural lesions on the beating heart.7 A previous epicardial ablation device, for example, was tested in our laboratory and demonstrated 67% transmurality when tested on the beating heart in an acute animal model.17 In a long-term study with this device, only 8% of the lesions were transmural along their entire length, and no lesion was able to create lasting conduction block.18 Microwave epicardial ablation devices were also tested on the beating heart in acute animals, creating transmural lesions only 36% and 38% of the time.17,19 The poor efficacy of these devices is one of the factors that have contributed to the often disappointing results with minimally invasive, off-pump surgical approaches for AF.20 Although bipolar RF clamping devices have been shown to reliably create transmural lesions around the pulmonary veins from a completely epicardial approach, completion of the full CMIV lesion set requires that one jaw of the clamp be inserted into the atrium, risking air embolism if performed without CPB. This has limited these clamps to performing only pulmonary vein isolation on the beating heart, which has had poor efficacy, particularly in patients with longstanding persistent AF.21 These results highlight the need for a novel epicardial device for the completion of the CMIV lesion set on the beating heart and for the refinement of minimally invasive techniques.
This study investigated the ability of a novel RF ablation system, the Estech Cobra Fusion, to reliably create transmural lesions on the beating heart from an epicardial approach in an acute porcine model. Our results show that this device allows for epicardial ablation on the beating heart without the use of CPB, with a success rate that is better than that of most previous devices that we have tested.13,22 This device was able to create transmural ablation in 94% of the cross sections tested, with 68% of the lesions being transmural across their entire length, without significant discrepancies in performance based on anatomic location of the lesion on the heart, cardiac output, or the thickness of the target tissue.
The improved success of this device compared with other epicardial linear devices may be due to several features that distinguish it from other devices tested in our laboratory. Instead of using one mode of energy, the Cobra Fusion combines both bipolar and monopolar energy. Furthermore, the suction assistance offered by this device helps to ensure contiguous linear tissue-device contact, which has been difficult for handheld epicardial devices without a mechanism such as clamping. Similarly, whereas clamping devices have addressed the heat sink of the endocavitary blood pool by isolating the target tissues inside the jaws of the clamp, epicardial ablation devices have typically used unidirectional energy transfer in direct apposition to circulating endocavitary blood at 37°C. This design excludes the endocavitary blood by isolating the target tissue inside the suction chamber of the ablation catheter. To ensure that suction is maintained throughout the ablation time, the device also features an alarm to alert the user to a loss of contact.
Unlike other epicardial devices tested by our laboratory, adverse events such as atrial perforation or charring were not observed in any of the lesions tested. This is credited to the heating algorithm of this device, which uses temperature feedback control to determine the amount of energy delivered to target tissues. This mechanism delivers less energy to thinner tissues, which reach the target temperature of 70°C more rapidly than thicker tissues, while achieving good rates of lesion transmurality with a single application.
Although the overall transmurality achieved by this device was one of the highest of any epicardial ablation device tested by our laboratory, 32% of the lesions were not transmural along their entire length. This still falls short of the desired clinical performance and of the results obtained when using bipolar clamps.13 Several factors were identified that may improve the performance of this device. Nontransmural cross sections were significantly associated with low body temperature of the animal. This could be related to suboptimal performance of the device’s heating algorithm at lower tissue temperatures that thus require additional energy to achieve a target temperature of 50°C, which has been shown to produce myocardial cell death. This problem could be avoided clinically by carefully maintaining the patient at normothermia during a procedure. However, we did not specifically test the effect of body temperature on the performance of the device, and thus, further investigation is needed to clarify this relationship. In addition, there was a significant difference in transmurality based on the location of the cross section within a lesion, with end sections being more often nontransmural. This is likely because of the fact that the electrodes terminate 5 mm from the end of the suction chamber, further highlighting the importance of overlapping ablation lines in the clinical setting because even a 1-mm gap is sufficient for conduction of both sinus and fibrillatory wave fronts.15
There are several weaknesses of this study. The present study was performed on normal healthy porcine atria. Although several steps were taken to minimize the impact of anatomic differences between the porcine atria used in this study and human atria, important differences still exist. The pathologic human atrial wall has been shown to differ significantly from healthy atrial myocardium in thickness, fibrosis, and amount of epicardial fat, all of which could impact energy delivery and affect lesion transmurality.23,24 Thus, surgeons need to be aware of these differences when using this device in diseased human hearts.
In addition, the present study is an acute animal study, and it is possible that transmurality could be different both in a chronic animal model and in the clinical population. However, unlike other studies that have used conduction block to demonstrate transmurality in the acute setting, which is known to be unreliable, our study used vital staining.18 Because a lesion was judged to be transmural only if the tissue did not take up any of the vital stain, our assessment represents a conservative estimation of the efficacy of this device in this model. Furthermore, previous studies in our laboratory have shown a good correlation between acute and chronic histology using this methodology.
The experimental design of this study was not meant to replicate the full CMIV lesion set. Our experimental lesion set was designed to test the device over a range of atrial tissues and was further refined on the basis of the anatomic considerations of the porcine model. Specific lesions of the CMIV lesion set that were not tested include the box lesion around the pulmonary veins, the lesion to the mitral annulus, and the lesion to the tricuspid annulus. Further clinical investigation is needed to determine the efficacy of this device in these areas.
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This is an acute animal model evaluating the efficacy of the Fusion device in producing linear, transmural lesions from an epicardial approach. The histologic assessments support that the device is generally effective in producing the desired effect. However, there are several serious limitations to this study that should temper the enthusiasm of the surgeon considering using this device in the clinical setting. First and foremost, there is no correlation between tissue thickness and transmurality of the lesion. In addition, the 3-mm thickness of the porcine atrial tissue tested is often much thinner than the often pathologic human atrial tissue encountered in clinical practice. It would be helpful if future animal studies use a chronic model with assessment of electrical isolation because that is really the final arbiter in judging the success of a lesion. Finally, although animal studies are valuable in evaluating the efficacy of ablative power sources for AF, decisions regarding the possible use of these devices in one’s practice should be predicated on sound clinical results reported in peer-reviewed journals and meetings. Once again, Dr Damiano and his colleagues have made a valuable contribution in our understanding of the performance of a power source for surgical ablation.
Harold G. Roberts, MD, is the guest editor.