21 Radiofrequency Ablation:
Principles and Techniques
Riccardo Lencioni, Dania Cioni, Jacopo Lera, Erika Rocchi, Clotilde Della Pina, and
Laura Crocetti
R. Lencioni, MD; D. Cioni, MD; J. Lera, MD; E. Rocchi, MD;
C. Della Pina, MD; L. Crocetti, MD
Division of Diagnostic and Interventional Radiology, Depart- ment of Oncology, Transplants and Advanced Technologies in Medicine, University of Pisa, Via Roma 67, 56126 Pisa, Italy
CONTENTS
21.1 Introduction 307
21.2 Radiofrequency: How it Works 307 21.2.1 Basic Principles 307
21.2.2 Mechanisms of Energy Deposition 308 21.3 The Bio-heat Equation: How to
Increase Tumor Destruction 310
21.4 Technical Clues for Clinical Application 311 21.4.1 Imaging 311
21.4.2 Anesthesiology Care 312 21.4.3 Ablation Protocols 312 References 314
21.1
Introduction
Image-guided radiofrequency (RF) ablation is a minimally invasive procedure that has emerged as the most powerful percutaneous technique for tumor destruction and is nowadays established as the primary ablative modality at most institutions (Goldberg 2002; Lencioni et al. 2003; Shibata et al. 2002). In fact, recent improvements in RF technology have permitted the creation of in vivo spherical ablation zones exceeding 5 cm in diam- eter with a single probe insertion, thus substan- tially increasing the potential of the technique in clinical application (Berber et al. 2004; Lencioni et al. 2004a). A thorough understanding of the ba- sic principles, mechanisms of energy deposition, modulation of tissue physiologic characteristics to increase tumor destruction, and technical clues in clinical application is essential for optimal use of RF ablation.
21.2
Radiofrequency: How it Works 21.2.1
Basic Principles
The goal of RF ablation is to induce thermal injury to the tissue through electromagnetic energy de- position. The term RF ablation applies to coagula- tion induced by all electromagnetic energy sources with frequencies less than 900 kHz, although most devices function in the range of 375–500 kHz. The term RF refers not to the emitted wave but rather to the alternating electric current that oscillates in this frequency range. In monopolar RF ablation, the patient is part of a closed-loop circuit that includes an RF generator, an electrode needle, and a large dis- persive electrode (ground pads). An alternating elec- tric field is created within the tissue of the patient.
Because of the relatively high electrical resistance of tissue in comparison with the metal electrodes, there is marked agitation of the ions present in the target tissue that surrounds the electrode, since the tissue ions attempt to follow the changes in direc- tion of the alternating electric current. The agitation results in frictional heat around the electrode. The discrepancy between the small surface area of the needle electrode and the large area of the ground pads causes the generated heat to be focused and concentrated around the needle electrode (Gazelle et al. 2000; Rhim et al. 2001).
The thermal damage caused by RF heating is de-
pendent on both the tissue temperature achieved
and the duration of heating. Heating of tissue at
50–55°C for 4–6 min produces irreversible cellular
damage. At temperatures between 60°C and 100°C
near immediate protein coagulation is induced,
with irreversible damage to mitochondrial and cy-
tosolic enzymes as well as nucleic acid-histone pro-
tein complexes (Goldberg et al. 2000a). Cells expe-
riencing this extent of thermal damage most often,
but not always, undergo coagulative necrosis over
the course of several days. In fact, the zone of co-
agulation, while predominantly comprising coagu- lative necrosis, often lacks the classic well-defined histologic appearance of coagulative necrosis in the acute postablation period or even within some zones of adequately ablated tissue for many months after ablation. Indeed, in many cases, specialized stains are required to confirm that cellular death has been achieved after thermal ablation (Goldberg et al.
2000a).
For this reason and because many tumors un- dergo central necrosis without ablation therapy, the term “coagulation” is preferred over the use of “ne- crosis” alone, because it denotes that the ablation in- tervention is actively leading to tumor destruction.
The more generalized term “coagulation” is pre- ferred over the term “coagulative necrosis” because the latter term has a well-defined meaning in the pa- thology literature, including the absence of visible nuclei within the dead cells (Goldberg et al. 2003).
In a recent study, explanted livers of cirrhotic pa- tients who had RF ablation of their hepatocellular carcinomas were examined. A detailed histopatho- logic analysis showed that, unlike classic tissue necrosis, the treated lesions all showed “thermal fixation,” with preserved tissue architecture and microscopic cellular detail. The cellular staining characteristics faded with time, but the treated tis- sue became brittle, resisted tissue breakdown, and generated a minimal wound healing response. At the periphery of the lesion, the fibrous septa of the cirrhotic liver and vascular structures appeared to demarcate or limit progression of the ablation front.
A narrow hypocellular fibrous boundary with a fo- cal “foreign body” giant cell-type reaction devel- oped around the edge of the ablation zone (Coad et al. 2003).
At 110°C, tissue vaporizes and carbonizes.
These processes usually retard optimal ablation due to a resultant decrease in energy transmission (Goldberg et al. 1996). For adequate destruction of tumor tissue, the entire target volume must be sub- jected to cytotoxic temperatures. Thus, an essential objective of ablative therapy is achievement and maintenance of a 50–100°C temperature through- out the entire target volume for at least 4–6 min.
However, the relatively slow thermal conduction from the electrode surface through the tissues may increase the duration of application to 10–30 min.
On the other hand, the tissue temperature should not be increased over these values to avoid carbon- ization around the tip of the electrode due to exces- sive heating (Gazelle et al. 2000; Goldberg et al.
1996; Rhim et al. 2001).
21.2.2
Mechanisms of Energy Deposition
In the early experiences with RF treatment, a major limitation of the technique was the small volume of ablation created by conventional monopolar elec- trodes. These devices were capable of producing cylindrical ablation zones not greater than 1.6 cm in the short axis (Goldberg et al. 1995). Therefore, multiple electrode insertions were necessary to treat all but the smallest lesions. Subsequently, several strategies for increasing the ablation zone achieved with RF treatment have been used.
Heat efficacy is defined as the difference between the amount of heat produced and the amount of heat lost. Therefore, effective ablation can be achieved by optimizing heat production and minimizing heat loss within the area to be ablated. The relationship between these factors has been characterized as the
“bio-heat equation.” The “bio-heat” equation gov- erning RF-induced heat transfer through tissue has been previously described by Pennes (1948), with this equation simplified to a first approximation by Goldberg et al. (2000b) as follows:
Coagulation =
energy deposited × local tissue interactions – heat lost Heat production is correlated with the intensity and duration of the RF energy deposited. On the other hand, heat conduction or diffusion is usually explained as a factor of heat loss in regard to the electrode tip. Heat is lost mainly through convec- tion by means of blood circulation (Patterson et al. 1998).
Therefore, most investigators devoted their atten- tion to strategies that increase the energy deposited into the tissues and several corporations have man- ufactured new RF ablation devices based on tech- nologic advances that increase heating efficacy. To accomplish this increase, the RF output of all com- mercially available generators has been increased to 150–200 W, which may potentially increase the in- tensity of the RF current deposited at the tissue.
Major progress was achieved with the introduc- tion of modified electrode needles, including in- ternally cooled electrodes and multitined expand- able electrodes with multiple retractable prongs on the tip (Berber et al. 2004; Lencioni et al.
1998; Lencioni et al. 2001; Lencioni et al. 2004b;
Solbiati et al. 1997). These techniques enabled a
substantial and reproducible enlargement of the ab-
lation zone produced with a single needle insertion,
and prompted the start of clinical application of RF ablation (Fig. 21.1).
Internally cooled electrodes (Radionics, Tyco Healthcare Group, Burlington, MA) consist of dual- lumen electrodes with an exposed active tip of vari- able length. Internal cooling is obtained by con- tinuous perfusion with chilled saline and is aimed at preventing overheating of tissues nearest to the electrode to minimize carbonization and gas forma- tion around the tip. The tip contains a thermocouple for recording the temperature of the adjacent tissue.
To increase the size of the ablation, the company placed three of the cooled electrodes in a parallel triangular cluster with a common hub. This device produces a significantly larger ablation than does a single cooled electrode (Goldberg et al. 1998a).
Pulsing of RF energy (i.e. alternation of very high RF current for several seconds followed by minimal RF deposition for a defined period) has also been described as a method that allows overall increased current deposition (Goldberg et al. 1999).
Multitined expandable electrodes have an ac- tive surface which can be substantially expanded by prongs deployed from the tip. The number of prongs and the length of their deployment varies according to the device and to the desired volume of ablation.
The commercially available devices were developed to monitor the ablation process so that high-tem- perature coagulation may occur without exceeding a 110°C maximum temperature threshold. One de- vice (RITA Medical Systems, Mountain View, CA) relies on direct temperature measurement. This kind of electrode, in fact, is made by an insulated outer cannula that houses nine curved electrodes of various lengths, which deploy out from the trocar tip. Five of the electrodes are hollow and contain thermocouples in their tips that are used to measure the tissue temperature. Probe-tip temperatures, tis- sue impedance, and wattage are displayed on the RF generator and are graphically recorded by dedicated software. Maximum power output of the RF genera- tor, amount of electrode array deployment from the trocar, and duration of the effective time of the abla- tion (time at target temperature) depend on the de- sired volume of ablation. In fact, the generator runs by an automated program and maintains the target temperature throughout the procedure. At the end of the procedure, the coagulation of the needle track can be done after retraction of the hooks with the aim of preventing any tumor cell dissemination (Lencioni et al. 2001; Lencioni et al. 2004a).
Another manufacturer (Boston Scientific, Natick, MA) produces an RF ablation device that relies on
Fig. 21.1a–d. Current RF equipment: a RF generator Model 1500 X, with a maximum output of 200 W and b Model 90 StarBurst XL multitined expandable electrode with nine prongs (RITA Medical Systems, Mountain View, CA).
c Expandable electrodes for 2.0-4.0-cm ablation Zones (Boston Scientific, Natick, MA). d single or cluster internally Cool-tip RF electrodes (Radionics, Tyco Healthcare Group,
b
a
electrical measurement of tissue impedance rather than on tissue temperature. The electrode is made by an insulated 14-gauge outer needle that houses ten retractable curved electrodes. The electrodes are manufactured in different lengths. In application, the tip of the needle is advanced to the target tissue and the curved electrodes are deployed to full ex- tension. The generator is switched on and energy is administered until a rapid rise in impedance occurs.
The impedance of the tissue increases as the tissue desiccates. It is assumed that an ablation is success- ful if the device impedes out (Rhim et al. 2001).
In addition to the previously described devices, several other designs for RF electrodes were recently developed and some of them are currently under- going clinical investigations. These include bipolar devices with two active electrode applicators placed in proximity to achieve contiguous coagulation and
d c
perfusion electrodes that have small apertures at the tip of the prongs that allow fluids (i.e. normal or hypertonic saline) to be infused into the tissue during the ablation procedure (Fig. 21.2) (Burdio et al. 2003; Kettenbach et al. 2003; Miao et al. 2001;
Mulier et al. 2003).
lesion to be treated. This has led some investiga- tors to study modifications of the underlying tumor characteristics in an attempt to improve RF thermal ablation. These modifications can be divided, on the basis of the bio-heat equation, as (a) strategies that permit an increase in the overall deposition of energy through an alteration in tissue electrical con- ductivity, (b) strategies that improve heat retention within the tissue, and (c) strategies that decrease the tolerance of tumor tissue to heat (Goldberg et al. 2000b).
For a given RF current, power deposition is strongly dependent on local electrical conductivity.
Intratumoral injection of saline solution prior to or during the application of RF current alters tissue conductivity and thereby allows greater deposition of RF current and increased tissue heating and coagu- lation (Goldberg et al. 2001a; Livraghi et al. 1997).
Experimental findings demonstrated that ablative temperatures can be generated farther from an RF electrode by increasing tissue electrical conductiv- ity with NaCl solution injection. However, because both volume and concentration of saline solution influence tissue heating and the coagulation diam- eter in a nonlinear fashion, optimal parameters for injection of saline solution must be determined for each type of RF apparatus used and for the different types of tumors and tissues to be treated. In an ex- perimental study performed in a controlled system of agar phantoms (Lobo et al. 2004), excessive in- crease of tissue electrical conductivity obtained with very high saline concentrations decreased the extent of heating. Increased electrical conductivity, in fact, has competing effects on RF ablation: it enables in- creased energy deposition and greater heating, but it also increases the energy required to heat a given volume of tissue. If this amount of energy cannot be delivered, that is, it is beyond the maximum genera- tor output, then less actual heating, and thus less coagulation, is achieved (Lobo et al. 2004). A major concern for the clinical application of percutaneous injection strategies is the possibility of determining a nonuniform alteration of tissue electrical conduc- tivity because of the difficulty of achieving uniform fluid diffusion and distribution. Irregularly shaped areas of coagulation have been observed with RF performed during simultaneous saline injection (Livraghi et al. 1997). Also, saline-enhanced abla- tion may potentially lead to distortions of the abla- tion zone shape because of the spread of fluid outside the target, along paths of least resistance (Ahmed et al. 2002; Boehm et al. 2002). Goldberg et al. (2001a) have observed this phenomenon in tissue samples in
Fig. 21.2. The RITA model 100 (Starburst Xli) has five electrodes with infusion sites for hypertonic saline at the end. They alternate with four metal prongs with thermocouples at their tip to monitor the tissue temperature. Two subtypes exist to create RF lesions of up to 5 cm or 7 cm
21.3
The Bio-heat Equation: How to Increase Tumor Destruction
Despite technologic advances and electrode modi-
fications that have effectively increased RF energy
deposition and tissue heating, inadequate coagula-
tion can represent a clinical problem in some cir-
cumstances. Specifically, there are multiple and
often tissue specific limitations which may cause
heterogeneity of heat deposition throughout a given
which large volumes of saline (25 ml) were injected.
These results suggested that infusion of very small volumes of fluid could increase the ablation volume without risks of undesired leakage.
Perfusion-mediated tissue cooling reduces the extent of coagulation produced by thermal ablation (Goldberg et al. 1998b). Modeling of the bio-heat equation shows that for a given tissue and power deposition, the effects of tissue blood flow predomi- nate. RF-induced coagulation is also more limited and variable in vivo than ex vivo. Coagulation in vivo is often shaped by vasculature in the vicinity of the ablation. Experiments in which hepatic per- fusion is altered by mechanical or pharmacologic means during RF ablation of normal liver tissue and tumors show that blood flow is largely respon- sible for this reduction in observed coagulation (Goldberg et al. 1998c; Rossi et al. 2000). Several strategies for reducing blood flow during ablation therapy have been proposed. Total portal inflow oc- clusion (Pringle maneuver) has been used at open laparotomy and at laparoscopy. Angiographic bal- loon occlusion of the hepatic artery can be used but proved useful for hypervascularized tumors only (Rossi et al. 2000). de Baere et al. (2002) published promising results showing that temporary hepatic vein or portal branch occlusion during RF ablation can facilitate the treatment of large tumors (>3.5 cm in maximum diameter) or tumors in contact with the walls of large vessels. However, RF ablation per- formed during vascular occlusion can increase the risks associated with the treatment. In one study, portal vein thrombosis was significantly more fre- quent after RF ablation performed during a Pringle maneuver in patients with liver cirrhosis (de Baere et al. 2003). Pharmacologic modulation of blood flow and antiangiogenesis therapy are theoretically possible but should currently be considered experi- mental (Goldberg et al. 1998c).
As far as decreasing tissue resistance to heat, based upon the well-documented relationship between the effects of some chemotherapeutic agents and hyper- thermia, recent studies began to explore the poten- tiation of effects that can be achieved with a combi- nation of chemotherapy and RF ablation (Goldberg et al. 2001b; Goldberg et al. 2002). These advances have already raised substantial clinical interest be- cause preliminary results from a randomized study using combined RF/liposomal doxorubicin therapy in patients with primary and secondary liver tumors demonstrated significant increases in tumor necro- sis compared with RF ablation alone (Goldberg et al. 2002). Ahmed et al. (2003) provided insight into
the factors that can potentially improve the outcome of this combination therapy. Validation in tumor models of larger size and different histological types will be a necessary next step to confirm the utility of this combination therapy before its adoption in a wider clinical setting. Further research exploring all of the factors that influence intratumoral liposome delivery, including liposome size, charge, circula- tion time, and composition, is ongoing. Other com- bination therapies, such as combined treatment of RF ablation and acid acetic injection, are currently under investigation (Lee et al. 2004).
21.4
Technical Clues for Clinical Application 21.4.1
Imaging
Imaging is used in five separate and distinct ways in RF ablation: planning, targeting, monitoring, controlling, and assessing treatment response (Goldberg et al. 2003). Imaging techniques, includ- ing ultrasound (US), CT, MR imaging, and more recently positron emission tomography (PET), are used to help determine whether patients are suitable candidates for RF ablation. Pre-treatment imaging planning must define tumor size and shape, num- ber, and location within the liver relative to blood and biliary vessels, as well as critical structures (i.e.
gallbladder, gastrointestinal tract) that might be at risk for injury during the ablation (Goldberg et al.
2003). Targeting refers to the placement of the RF electrode into the tumor, which can be achieved by using US, CT, or MR imaging. The guidance system is chosen largely on the basis of operator preference and local availability of dedicated equipment such as CT fluoroscopy or open MR systems. Monitoring is the term used to describe the process with which ablation effects are viewed during the procedure.
Important aspects to be monitored include how well
the tumor is being covered and whether any adjacent
normal structures are being affected at the same
time (Goldberg et al. 2003). While the transient
hyperechoic zone that is seen on US within and sur-
rounding a tumor during and immediately after RF
ablation can be used as a rough guide to the extent
of tumor destruction, MR is currently the only imag-
ing modality with validated techniques for real-time
temperature monitoring (Quesson et al. 2000). The
term “controlling” is used to describe the intrapro-
cedural tools and techniques that are used to control the treatment. To control an image-guided ablation procedure, the operator can utilize the image-based information obtained during monitoring or auto- mated systems that terminates the ablation at a criti- cal point in the procedure (Goldberg et al. 2003).
Finally, imaging is used to assess the outcome of the procedure. Contrast-enhanced US performed after the end of the procedure may allow an initial evalua- tion of treatment effects. Contrast-enhanced CT and MR imaging are recognized as the standard modali- ties to assess treatment outcome, although promis- ing initial results have been reported by using PET after RF ablation of liver metastases (Donckier et al. 2003). CT and MR images obtained after treat- ment show successful ablation as a non-enhancing area surrounded by an enhancing rim. The enhanc- ing rim appears a relatively concentric, symmetric, and uniform process in an area with smooth inner margins. This is a transient finding that represents a benign physiologic response to thermal injury (initially, reactive hyperemia; subsequently, fibro- sis and giant cell reaction) (Goldberg et al. 2000a).
Benign periablational enhancement needs to be dif- ferentiated from irregular peripheral enhancement due to residual tumor that occurs at the treatment margin. In contrast to benign periablational en- hancement, residual unablated tumor often grows in scattered, nodular, or eccentric patterns (Goldberg et al. 2003). Later follow-up imaging studies should be aimed at detecting the recurrence of the treated lesion (i.e. local tumor progression), the develop- ment of new hepatic lesions, or the emergence of extrahepatic metastases.
21.4.2
Anesthesiology Care
Patient candidates to RF ablation can have a me- dium-to-high anesthesiology risk. Most of them, in fact, have been rejected for surgery for associated diseases involving the cardiovascular system. There is no consensus on the best anesthesiology care for RF ablation. Local anesthesia does not produce ad- equate pain relief. Some centers use general anesthe- sia and endotracheal intubation. Others, including our own, prefer to perform liver RF under conscious sedation. The association of an hypnotic drug with an ultrashort half-life analgesic drug allows a mild sedation of the patient, who can cooperate with the operator and bear the pain induced by treatment. One possible protocol consists of administering a bolus
of ketorolac (0.5-0.8 mg/kg) followed by infusion of propofol (1-2 mg/kg/h) and remifentanil (0.1 µg/kg/
min). However, drug posology has to be modulated in relation to the individual patient compliance and to the different phases of the procedure. The infusion of the hypnotic drug can be varied between 0.5 and 2 mg/kg/h to achieve a patient sedation that preserves the ability to do easy actions. The infusion of remifen- tanil can be varied between 0.05 and 0.15 µg/kg/min to obtain an optimal analgesia. Attention has to given to avoiding bolus administration of remifentanil, as this may cause respiratory depression. The procedure is performed under standard cardiac, pressure, and oxygen monitoring. A careful post-treatment proto- col is to be recommended following RF ablation. The patient is kept under close medical observation and re-scanned with US 1–2 h after treatment to detect any bleeding. An overnight hospital stay is scheduled. In most of the cases, patients may be discharged the day after the procedure.
21.4.3
Ablation Protocols
An important factor that affects the success of RF ablation is the ability to ablate all viable tumor tis- sue and an adequate tumor-free margin. The most important difference between surgical resection and RF ablation of hepatic tumors is the surgeon’s in- sistence on a 1-cm-wide tumor-free zone along the resection margin. To achieve rates of local tumor recurrence with RF ablation that are comparable to those obtained with hepatic resection, physicians should produce a 360°, 1-cm-thick tumor-free mar- gin around each tumor (Cady et al. 1998). This cuff is necessary to ensure that all microscopic inva- sions around the periphery of a tumor have been eradicated. Thus, the target diameter of an ablation must be ideally 2 cm larger than the diameter of the tumor that undergoes treatment (Fig. 21.3) (Dodd et al. 2000; Patterson et al. 1998). Eradication of a tumor can therefore be achieved with a single abla- tion if the diameter of the tumor is 2 cm less than the diameter of tissue ablated. Currently, the maximum diameter of the in vivo ablation sphere produced by RF is 5.5–5.6 cm (Berber et al. 2004). Therefore the tumor to be treated should not exceed 3.5 cm in longest axis to obtain a safety margin of 1 cm all around the lesion.
Despite technological advances and all the efforts
made to obtain larger zones of ablation, moderate
to high rates of local tumor recurrence have been
reported, especially in larger tumors (Livraghi et al. 2000; Solbiati et al. 1997). A possible reason for failures in the treatment of large tumors is the inability to determine the optimal number of abla- tions and the exact location of electrode placement needed to completely destroy tumors larger than the size of a single ablation zone. Thus, appropri- ate protocols to determine the correct number of RF ablations have been devised (Chen et al. 2004; Dodd et al. 2000). In one model, developed for treatment of tumors larger than 3.5 cm in diameter, the regu- lar prism and the regular polyhedron were used to generate mathematical models for determining the correct preoperative protocol. Number of ablations ranged from 4 to 12 for spherical tumors with diam- eters respectively of 4.0–4.3 cm and 5.7–6.5 cm. In some cases, however, fewer electrodes were placed than were calculated mathematically because of the particular characteristics of the tumors (Chen et al. 2004). In this study, local recurrence rate did not exceed 24% in a series of 121 tumors with a mean diameter of 4.7 ±0.9 cm (range 3.6–7.0 cm) (Chen et al. 2004).
In addition to size, tumor location is one of the important factors that influence the likely outcome of therapy. Treatment of lesions adjacent to the gall- bladder or to the hepatic hilum is at risk of thermal injury of the biliary tract. Nevertheless, in experi- enced hands, RF ablation of tumors adjacent to the gallbladder was shown to be feasible although asso- ciated in most cases with self-limited iatrogenic cho- lecystitis (Chopra et al. 2003). In contrast, treatment of lesions located in the vicinity of hepatic vessels is possible, since flowing blood usually “refrigerates”
the vascular wall, protecting it from thermal injury:
in these cases, however, the risk of incomplete ab- lation of the neoplastic tissue adjacent to the vessel may increase because of the heat loss caused by the vessel itself. Lesions located along the surface of the liver can be considered for RF ablation, although their treatment requires experienced hands and may be associated with a higher risk of complica- tions. Percutaneous treatment of superficial lesions that are adjacent to any part of the gastrointestinal tract must be avoided because of the risk of thermal injury of the gastric or bowel wall (Buscarini and
Fig. 21.3a–d. Schematic model of an RF ablation protocol. a,b A 360°, 1-cm-thick tumor-free margin around each tumor must be achieved for successful ablation. Thus, the target diameter (D*) of an ablation must be ideally 2 cm larger than the diameter (D) of the tumor that undergoes treatment. c,d The multitined electrode needle is placed in the tumor and deployed gradually to encompass the tumor itself and 1 cm of tissue all around it
a
b
c d
Buscarini 2004; Livraghi et al. 2003). The co- lon appears to be at greater risk than the stomach or small bowel for thermally mediated perforation (Rhim et al. 2004). Gastric complications are rare, likely owing to the relatively greater wall thickness of the stomach or the rarity of surgical adhesions along the gastrohepatic ligament. The mobility of the small bowel may also provide the bowel with greater protection compared with the relatively fixed colon.
The potential risk of thermal damage to adjacent structures should be weighed against benefits on a case-by-case basis. A laparoscopy approach can also be considered in such instances, as the bowel may be lifted away from the tumor (Rhim et al. 2004).
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