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Monitoring Radio-Frequency Thermal Ablation with Ultrasound by Low Frequency Acoustic Emissions – In Vitro and In Vivo Study
Ultrasound in Med. & Biol., Vol. 37, No. 5, pp. 755–767, 2011 Ó 2011 Published by Elsevier Inc. on behalf of World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/$ - see front matter
doi:10.1016/j.ultrasmedbio.2010.11.008
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Original Contribution MONITORING RADIO-FREQUENCY THERMAL ABLATION WITH ULTRASOUND BY LOW FREQUENCY ACOUSTIC EMISSIONS – IN VITRO AND IN VIVO STUDY ITAI WINKLER and DAN ADAM Biomedical Engineering, Technion City, Haifa, Israel (Received 11 June 2010; revised 5 November 2010; in final form 12 November 2010)
Abstract—The object of this study was to evaluate the monitoring of thermal ablation therapy by measuring the nonlinear response to ultrasound insonation at the region being treated. Previous reports have shown that during tissue heating, microbubbles are formed. Under the application of ultrasound, these microbubbles may be driven into nonlinear motion that produces acoustic emissions at sub-harmonic frequencies and a general increase of emissions at low frequencies. These low frequency emissions may be used to monitor ablation surgery. In this study, a modified commercial ultrasound system was used for transmitting ultrasound pulses and for recording raw RF-lines from a scan plane in porcine (in vitro) and rabbit (in vivo) livers during radio-frequency ablation (RFA). The transmission pulse was 15 cycles in length at 4 MHz (in vitro) and 3.6 MHz (in vivo). Thermocouples were used for monitoring temperatures during the RFA treatment. In the in vitro experiments, recorded RF signals (A-lines) were segmented, and the total energy was measured at two different frequency bands: at a low frequency band (LFB) of 1–2.5 MHz and at the transmission frequency band (TFB) of 3.5–4.5 MHz. The mean energy at the LFB and at the TFB increased substantially in areas adjacent to the RF needle. These energies also changed abruptly at higher temperatures, thus, producing great variance in the received energy. Mean energies in areas distant from RF needle showed little change and variation during treatment. It was also shown that a 3 dB increase of energy at the low frequency band was typically obtained in regions in which temperature was above 53.3 ± 5 C. Thus, this may help in evaluating regions undergoing hyperthermia. In the in vivo experiments, an imaging algorithm based on measuring the LFB energy was used. The algorithm performs a moving average of the LFB energies measured at segments within the scan plane.Results show that a colored region is formed on the image and that it is similar in size to a measurement of the lesion from gross pathology, with a correlation coefficient of 0.743. (E-mail: [email protected]) Ó 2011 Published by Elsevier Inc. on behalf of World Federation for Ultrasound in Medicine & Biology. Key Words: Radio-frequency ablation (RFA), Treatment monitoring, Acoustic emissions, Ultrasound, Signal processing.
ultrasound (HIFU), have been suggested and introduced into clinics. RFA is considered the most commonly used approach for liver tumor ablation. The RFA procedure includes the insertion of a radiofrequency (RF) electrode (RF needle) into the tumor, usually aided by ultrasound imaging guidance. After the placement of the RF needle at the target location, a strong alternating electric current is delivered to the tumor. The treatment goal is to cause immediate cell death by heat to the tumor cells and their adjacent marginal area to ensure treatment success but with minimal damage to nearby healthy tissue. The electrical current causes agitation of ions that causes frictional heat, leading to quick temperature elevation and tissue ablation around the RF needle and in adjacent areas (Daniels et al. 2007; Gazelle et al. 2000). Lesion size depends on the RF output power and duration and it
INTRODUCTION Treatment of primary and secondary malignant hepatic tumors, considered to be one of the most common malignant tumors worldwide, is limited. Chemotherapy and radiation therapy are in many cases ineffective against primary and secondary hepatic tumors, thus, resection surgery is a potential solution. However, most patients are not candidates for this surgery for various reasons (Buscarini et al. 2005; Chiou et al. 2007). In recent years, minimally invasive ablation approaches, such as radio-frequency ablation (RFA) and laser ablation, or noninvasive approaches such as high-intensity focused
Address correspondence to: Itai Winkler, Biomedical Engineering, 32000 Technion City, Haifa 32000, Israel. E-mail: Itai. [email protected] 755
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varies greatly between different individuals and different tissue types. Blood perfusion, tissue properties and anisotropy make it difficult to predict for each case the resultant lesion size and shape. Ultrasound imaging is the most common technique for guiding the insertion of the RF needle. (Buscarini et al. 2005; Chiou et al. 2007; Rhim and Dodd 1999; Tateishi et al. 2005). However, B-mode imaging does not provide sufficient visualization of the lesion formation and treatment progress (Buscarini et al. 2005; Chiou et al. 2007; Rhim and Dodd 1999; Varghese et al. 2001). Eventually, the lack of sufficient real-time monitoring ability leads to a high recurrence rate, greater than 33% (Varghese et al. 2001), and repeated surgeries. During the clinical ablative process, numerous microbubbles are formed because usually, the heating is fast and temperatures reach boiling temperatures. These microbubbles are strong reflectors and, thus, produce strong echoes that enable their visualization in regular B-mode imaging. This appearance of new reflectors in the treated area is currently used to assess the heated region. However, this visualization technique is poor for assessing the true damage caused to the tumor and the completion of the treatment (Gazelle et al. 2000). Many studies have attempted to use ultrasound (US) thermometry to image temperature changes. The most common approaches are based on measuring echo shifts (Maass-Moreno and Damianou 1996; Maass-Moreno et al. 1996; Simon et al. 1998; Souchon et al. 2005). Other approaches include measuring frequency shifts (Amini et al. 2005; Gaitini et al. 2005; Seip and Ebbini 1995) and changes in backscattered energy (Arthur et al. 2005). Echo shifts result from two main mechanisms: changes in the speed of sound (SOS) and tissue expansion. However, the SOS does not change monotonically with temperature (Techvipoo et al. 2004). Moreover, the natural variation between individuals creates the need for specific calibration curves for each treatment (Arthur et al. 2005). Another problem with echo shift and frequency shift methods is their sensitivity to movements and bubble formation. Both methods are based on tracking echo or frequency displacements by some correlation technique. Therefore, both techniques are very sensitive to tissue motions and bubble formation that lead to severe decorrelation between measurements. Therefore, the echo shift and frequency shift methods work well for small temperature elevations but fail for high temperature elevations typical in ablation surgeries (Gaitini et al. 2005; Miller et al. 2002). Other studies have directly addressed the problem of treatment monitoring, e.g., using elastography (Fahey et al. 2004; Kolokythas et al. 2008; Varghese et al. 2001) or the acoustic nonlinearity parameter (b/a) (Liu
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et al. 2008). Elastography also suffers from the decorrelation of signals due to bubble formation. Therefore, elastography measurements are usually made after a cool down period, when most microbubbles have been dissolved and the lesion becomes stiffer than its surrounding. During heating, even at small temperature elevations (at the early stages of the ablation treatment), microbubbles are formed as the temperature rises (Mclaughlan et al. 2007; Sanghvi et al. 1995). These microbubbles can be detected and distinguished from the surrounding tissue by their typical emissions, caused when the microbubbles are insonified by US (Biagi et al 2005; Leighton 1997; Mast et al. 2008). A free bubble in liquid media, insonified by US with sufficient intensity and duration (number of cycles), may emit different subharmonics (Biagi et al. 2005; Leighton 1997) as well as other low frequency emissions (Leighton 1997; Mast et al. 2008). In this study, an attempt was made to exploit these emissions at low frequencies for the assessment of the treatment progress. A modified commercial US system was used for transmitting and receiving US waves, so as to interrogate a scan plane during RFA. From the acquired RF signals, low frequency band (LFB) energies were evaluated at different regions within the scan plane. Thermocouple measurements were used as control measurements. MATERIALS AND METHODS Experiment set-up In vitro. Fresh whole porcine livers were purchased from a local slaughter house less than 48 h post mortem and were used on the same day of purchase. Each liver was brought to room temperature (19–23 C) and two to three samples, no smaller than 7 3 4 3 2 cm3, were cut from it. The samples were visually checked to be certain they contained a minimum amount of large blood vessels. The samples were held by a specially designed sample holder, which has several holes in the front to allow insertion of the RF needle and two needle thermocouples (Fig. 1). The sample holder has a hollow floor to allow contact between the tissue samples and a metal plate that serves as the dispersive electrode. The metal plate was connected in series to a power resistor, to achieve total impedance of around 120 U. The RF ablation system used with an internally cooled needle (cool-tip RF, Vallylab; Tyco Healthcare Group LP, Boulder, CO, USA) theoretically produces a lesion of a cylindrical shape (Lobik et al. 2005). Two needle thermocouples, with a diameter of 0.25 mm (Type T; Omega Engineering, Stamford, CT, USA) and placed inside 10 cm syringe needles, were inserted into
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Fig. 1. (a) Experimental set-up. The liver samples were fixed by the sample holder and were in direct contact with the metal plate under it, serving as the dispersive electrode. The radio-frequency (RF) needle with the two needle thermocouples were placed through designated holes in front of the sample holder and perpendicular to US beam direction. (b) Cross-section after a typical radio-frequency ablation (RFA) treatment.
the tissue through the sample holder holes, 2.5 mm and 5 mm away from the RF needle (Fig. 1). The RF needle and the two thermocouples were oriented along a line parallel to the ultrasound transducer’s surface and perpendicular to the imaged plane. The temperature was recorded every 1 s during the whole experiment by software (Virtual Bench Logger; National Instruments, Austin, TX, USA). A Vivid III US imaging system with a S5 phased array probe (GE Healthcare, Giles, UK) was used for transmitting and receiving the US beams at the selected scan plane. The transmission frequency was set to be 4 MHz with bursts of 15 cycles in length. The transducer’s focus was set to a depth similar to the depth of the RF needle or at most 1 cm below it. Output intensity was set to produce a mechanical index of MI 5 0.6. The transducer was held fixed during the entire experiment. The raw US analog signals (RF lines) were acquired using custom software, designed for acquisition of the Vivid III analog signals before they are processed by the imaging system. Because the current produced by the RF generator disrupts the US data acquisition, a switch box was designed to turn off the ablation current for brief periods during data acquisition. The switch box was controlled by the custom data acquisition software. The current was turned off every 5 s for 0.2 s, during which the RF signals from the whole scan plane, as well as the thermocouple measurements, were acquired. The resultant duty cycle of the RFA treatment was 96%. Twelve acquisitions (frames), consisting of all the RF lines comprising a scan plane, were made as reference, before the ablation was started, and 24 were made after the ablation current was terminated. The RF needle is 15 cm long and has an exposed tip of 3 cm in length. The tips of the thermocouples were aligned with the center of the exposed part of the RF needle. Insertion of the RF needle and the thermocouples was
done under regular B-mode imaging, to confirm the correct placement of the instruments. The US imaging was also used to make sure that the scan plane and the chosen target for ablation consisted of a minimal amount of large scatterers. The metal plate located under the tissue sample was clearly visualized in the B-mode image and served as an indicator that the US beams were not blocked from reaching the region-of-interest (ROI). The RF generator output was manually set to 16 watts. The duration of each RF treatment was around 15 min, which was the time needed to achieve a desired temperature of 80–90 C at thermocouple 1 (closer to the RF needle). In vivo. The in vivo experiments were carried out using rabbits (NZW, 3.5–4.5 kg). The protocol for the experiments was approved by the Technion committee for the supervision of animal experiments. Rabbits were anesthetized with a mixture of ketamine (35 mg/kg), xylazine (2 mg/kg), acepromazine (0.3 mg/kg) and butorphanol (0.5 mg/kg) (SC). An additional dose of 3 mg/kg ketamine was injected (IV) every 10 min. Vital signs and anesthesia strength were checked frequently. The liver was exposed by an abdominal section. The liver was pulled outward slightly and two pads were inserted beneath the liver to expose a larger surface area to allow better placement of the instruments (US transducer, RF needle and thermocouples). The US transducer, RF needle and thermocouples were fixed during the entire procedure. As detailed in the in vitro section, a B-mode image was acquired before the treatment started to identify the locations of the RF needle and the thermocouples within the tissue. RF current was administered for 5–10 min with power output of 20–40 Watts. Fifteen rabbits were included in the study. Each rabbit was subjected to a single RFA treatment.
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After the procedure and a cool down period, the rabbits were euthanized (Thiopental, 650 mL/kg, IV). The ablated areas were incised for visual examination of the treatment outcome and for lesion measurements. The in vivo set-up differed from the in vitro phase for certain parameters. The transmission frequency was 3.6 MHz with bursts of 15 cycles in length and with an intensity of MI 5 0.6. Insertions of the RF needle and the thermocouples were made through a custom built holder with the same spacing between the RF needle and the thermocouples as was described in the in vitro section. Data processing and analysis In vitro. The acquired RF lines were sampled at 20 MHz (by the Vivid III) and were filtered with two band pass filters (BPF) with pass bands in the following range: low frequency band (LFB) 5 1–2.5 MHz, transmission frequency band (TFB) 5 3.5–4.5 MHz. The LFB was measured to study the behavior of the low frequency emissions from different areas within the imaging plane during RFA. The higher frequency band (TFB), which includes the transmission frequency, was measured to study the changes of the tissue echogenicity during RFA. The US transmit pulse had a bandwidth of 0.5 MHz. The filtered signals were uniformly segmented to 64 sample windows, which correspond to 2.5 mm in length, without overlaps. LFB and TFB energies in each window were calculated using the L2 norm:
Eðf1 ; f2 Þ5
ð f2 f1
jXðf Þj df z 2
N 1 X n50
jx½n hBPF j
2
(1)
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where Eðf1 ; f2 Þ denotes the energy in the band specified by f1 ; f2 , Xðf Þ is the Fourier transform of a time signal, x½n is the sampled signal and hBPF is the filter impulse response. Because the US probe used here was a phased array probe, energy calculations made from the acquired A-lines are given in a polar coordinate systemðr; qÞ, where r represents the distance from the center of the transducer center and q represents the angle of the beam in relation to the normal to the transducer’s surface. The energy measurements were first transformed to a Cartesian coordinate system using a weighted average of the two nearest energy measurements in the polar representation. Weights were assigned according to the distances as presented in eqn (2): Eðx; yÞ5W1 E1 ðr; qÞ1W2 E2 ðr; qÞ d2 W1 5 ; d1 1d2
W2 5
d1 d1 1d2
(2) (3)
where E1 ; E2 represent the energies of the neighboring locations in polar coordinates, and d1 ; d2 are the Euclidean distances between the vertex in the Cartesian grid to the closest polar neighbors. After obtaining the LFB and TFB energy maps for the entire scan plane, data was analyzed in three ways, as follows. (1) The average LFB energy was calculated in two windows of sizes 1.6 mm 3 2.5 mm (lateral x axial), as described in Figure 2. Each of the two window centers is positioned at the same distance from the RF needle as that of the thermocouples and at the
Fig. 2. Experiment set-up (cross-section view). The radio-frequency (RF) needle and the two thermocouples separated by 2.5 mm spacing on a line parallel to the transducers surface. Two regions (windows) are centered in the same distances from the RF needle as the thermocouples and on the same line parallel to the transducers surface.
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same depth. When the average LFB energy within a window increased by 3 dB, the temperature at that moment, as measured by the corresponding thermocouple, was registered. This enabled us to correlate between the LFB energy and the temperature, assuming that the temperature at the thermocouple sites is similar to the temperature at the symmetrical points on the other side of the RF needle. This assumption is likely valid because liver tissue is relatively homogenous, as evident by the circular symmetry of the resultant lesion. (2) The average LFB and TFB energies were calculated for seven adjacent windows, as described in Figure 3. All windows were on the same line formed by the RF needle and the thermocouples and at the same depth. All windows were non-overlapping (Fig. 3). The window size was 2 mm 3 2.5 mm (lateral 3 axial), thus enabling a reasonable spatial resolution. The axial size of the window was limited by the number of samples used for the energy calculation in each window. Average energies for these windows were calculated over 50 consecutive frames from the moment thermocouple 1 reached 60 C in each of the samples. (3) Two large windows within the scan plane of 5 mm 3 5 mm were selected. The first window was a large region within the ablated area. The second window was centered at a distant location from the RF needle, assumed not to be affected by the treatment (Fig. 4). The average LFB and TFB energies and energy variances were calculated for the two regions and for seven consecutive frames from three different start-
Fig. 3. Illustration of the location of the windows where the low frequency band (LFB) and transmission frequency band (TFB) energies were measured. In each of the windows numbered 1–7, the mean LFB and TFB energies were calculated.
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ing points. The starting points were chosen to be the moments thermocouple 1 reached temperatures of 50, 60 and 70 C. All data processing was performed using MATLAB (Mathworks, Natick, MA, USA). In vivo. The data measured during the in vivo experiments were processed similarly to that obtained during the in vitro experiments and they were also processed to provide images of the calculated parameters. The formation of image sequences that allow monitoring of the treatment first requires the computation of the LFB energy from all segments within the scan plane or in a predefined ROI. The energy band that was studied, 0.5–0.7 MHz, was somewhat different than during the in vitro studies. This band produced better results due to a better rejection of natural scatterers. To confirm that the energy in this frequency range was due to microbubbles, a single element transducer with a center frequency of 0.5 MHz (V318; Panametrics, Waltham, MA, USA) was used to passively record signals during the experiment. For each of the Vivid III recordings, the control system issued a trigger to a computer scope (Compuscope CS14100; Gage Applied Inc., Lachine, Canada) that recorded 540 K samples at a sampling rate of 25 MHz. The first 120 K samples were recorded before the Vivid III received the trigger and the rest were recorded after the trigger. Energy in the 0.5–0.7 MHz band was calculated from two different segments in the recorded signal. The first segment was before the Vivid III was activated and the second was after it was activated, i.e., during ultrasound transmission. The segments were 110 K samples in length. At each time step, an image was created by calculating the average of the last 10 frames. Because the ablated regions were expected to produce higher LFB energies than untreated regions and natural scatterers are severely attenuated due to the low frequency range used for the LFB energy calculations, the resultant image is expected to produce a contrast between treated and untreated areas. After calculating all the pixel values that comprise an image, spatial filtering was performed by a median filter (a square mask of 5 3 5 pixels, lateral 3 axial), followed by a spatial averaging filter (of 5 3 3 pixels, lateral 3 axial), to improve the image and reduce artifacts, and by gamma correction ðg51=2Þ. The median filter was used to remove high energy measurements from the images. These high energy measurements are caused by the wide energy distribution. Removal of these high energies is desired for creating a tighter energy range before assigning colors according to the energy levels. The averaging filter was implemented to smooth the image and produce a visually clearer result.
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Fig. 4. The two selected windows: each window was of a size of 5 3 5 mm. The ablation window corresponds to an area that is subjected to ablation. The reference window is at a distant location from the radio-frequency (RF) needle and is not subjected to ablation.
RESULTS In vitro Energy elevations of 3 dB in the LFB were measured from the graphs of energy vs. time for the two windows and the temperatures associated with these 3 dB increases were found from the thermocouple recordings to be 53.3 6 5 C (mean 6 std, n 5 36). At higher temperatures, rapid changes in the LFB energy could be observed, which may be explained by rapid formation and implosion or dissolution of microbubbles. Figure 5 presents averaged results of the LFB and TFB energies, recorded from the moment that thermocouple 1 reached 60 C, as a function of the distance from the RF needle, calculated over 18 experiments. All energies were normalized by the energy measured before the beginning of the heating within a rectangular
window whose center was the RF needle, averaged over the first 10 recorded frames. Mean energies are presented in arbitrary units. The bars represent the standard error of the measurements. It can be seen in Figure 5 that both the LFB energy and TFB energy decrease as the distance from the RF needle increases. Figure 6 depicts representative recordings of the LFB and TFB energies measured at different single ‘‘pixels’’, i.e., energy calculated from a single RF segment, taken from different locations within the windows illustrated in Figure 4 during RFA treatment. Energies are displayed in dB relative to the mean noise level, which was measured in a large region of the image before the heating started. The energy measured in the ablated areas tended to vary abruptly from frame to frame, thus, producing
Fig. 5. Mean energy measured during 50 frames from the moment thermocouple 1 reached 60 C. (a) Mean low frequency band (TFB) energy. (b) Mean transmission frequency band (LFB) energy. (average 6 standard error).
Monitoring radio-frequency thermal ablation d I. WINKLER and D. ADAM
Fig. 6. Representative recordings of energy measured from four segments during a typical experiment. Each pixel represents a small area in the scan plane. (a) Low frequency band (LFB) energies measured from two nonablated segments. (b) LFB energies measured from two ablated segments. (c) Transmission frequency band (TFB) energies measured from two nonablated segments. (d) TFB energies measured from two ablated segments. (e) Temperature measured by the two thermocouples during the experiment.
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a high variance of energy at both frequency bands compared to the base level variance. In the non-heated areas, the LFB and TFB energies changed slowly and were usually lower than the measured energies in the heated areas. In the non-heated regions, rapid changes in the energy were recorded from time to time, but in general, the energy graphs were smoother. In Figure 6, some jumps can be seen to occur sporadically at the time interval 400–800 s. The cause of these occasional rapid changes is not clear but we speculate that it may result from some type of treatment induced movements. The LFB energy measured in the non-heated areas was typically smoother than the TFB energy recorded from those regions. These points are depicted in Figure 7. Figure 7 presents energies (average 6 standard error) calculated from two large regions (windows) within the scan plane at different temperatures. Results are for 18 measurements; the bars represent the standard error of the measurements. The average normalized LFB and TFB energies measured in the window within the heated region increased as the temperature increased. Significant increases of LFB energy can be seen in ablated areas when heating from 50 to 70 C (p , 0.05). A significant increase of the TFB energy was also observed at these temperatures (p , 0.04). Both energies showed similar behavior. The variance of the LFB and TFB energies, measured at segments within the heated region and the reference one, also both increased with temperature. A significant increase in the mean LFB and TFB variances was found between the base level of 22 to 70 C (p , 0.02 for the LFB and p , 0.01 for the TFB).
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This was probably due to large bubbles being trapped in the tissue. The diameter of the region that changed its color was measured and compared with the manual pathology measurements made after a cool down period, performed after the animal was euthanized. The manual measurement of the lesion was taken along a line parallel to the liver lobe surface (the surface that is distal from the US transducer) and through the point of the RF needle location. This liver lobe surface was in contact with the surface of the rabbit’s stomach during the experiment. Because the rabbit stomach surface was visible from the B-mode image taken before the initiation of ablation, the orientation of that surface in the algorithm images could be found. Thus, the image measurement was performed along a line passing through the RF needle and with the same direction as the stomach surface. The boundaries for the lesion measurements were found by including all pixel values in the range of 5% to 100% along the line of measurement and were taken from the frame that presented the largest measurement along this line. The comparison between manual and image measurements is presented in Figure 10. As can be seen in Figure 10, a correlation coefficient of 0.743 was found between the lesion diameter, as obtained by the manual measurement, and the region in which the color changed, as measured from the images of the algorithm. The error calculated from these 14 experiments is –0.3 6 1.4 [mm] (average 6 std) or 3% 6 13.7%. Two measurements were removed as outliers (see Discussion section). DISCUSSION
In vivo Figure 8 presents the energy recorded within the 0.5–0.7 MHz band as a function of time (frame), which increased with time from the moment the ablation procedure was initiated and decreased once the ablation was terminated. This indicates that the emissions in the low frequency band were caused by the treatment. Some residual energy, caused by the RF generator, was also present, but this value was constant as can be seen from the graphs. An example of the algorithm output, together with the temperature measurements and a B-mode image acquired before the initiation of the ablation, are presented in Figure 9. As can be seen, change in color that correspond to an increase of the LFB energies occur in the region associated with the probe location. The region in which color changed increased in size as long as the ablation current was on and it started to shrink once the ablation current was stopped. Even after a cool down period, some areas still produced increased LFB energy.
In vitro In vitro tissue samples had large bubbles and air cavities trapped in the blood vessels. These bubbles appear as strong echoes in the B-mode image and produce artifacts in the experiments. When the air cavity is large enough it can cause shadowing and block the US waves. To decrease the probability of having trapped air, the liver samples were taken from the distal part of each liver lobe, where smaller blood vessels are present. Tissue samples were also gently massaged to remove the trapped air from inside the large vessels. Despite these efforts, six experiments were excluded due to strong echoes in areas near the RF needle. The liver samples used in the in vitro experiments were not degassed. From our experience, degassing could only be partially performed and may damage the sample. The degassing process may also change the ability to form microbubbles within the tissue due to the decreased gas concentration in the liquids. In previous papers, degassing was applied only for short durations (Mclaughlan et al.
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Fig. 7. (a) and (b) Low frequency band (LFB) and transmission frequency band (TFB) energy ratio at 50, 60, and 70 C. Ratio is calculated in relations to mean energy at the same region prior to RFA treatment. (c) and (d) Averaged energy variance of in the selected region. All values in these graphs are calculated from 18 experiments (average 6 standard error).
2007; Souchon et al. 2005). Other researchers simply suspended the tissue in a partially degassed buffer (Mast et al. 2008). The purpose of the first set of experiments was to find a typical temperature where a 3 dB increase in the measured LFB energy was observed. It was shown that such an increase usually occurs when temperatures of 53.3 6 5 C are reached. This result indicates that at these temperatures microbubbles were already formed, which increased the back scattered energy. This comparison is made under the assumption that the temperatures on both sides of the RF needle are symmetrically distributed. It would have been more accurate to measure the energy right at the thermocouples, however, echoes and reverber-
ations from the thermocouples and the RF needle disrupt the energy measurements. Moreover, the insertion of the thermocouples into the tissue may produce cavities filled with air, further affecting energy measurements from this area. Therefore, all energy measurements were made at regions that did not include the thermocouples but were at symmetrical distances on the other side of the RF needle and with a safety margin of 2 mm from it, assuming that liver tissues are relatively homogenous and that there is no asymmetric perfusion involved. The lesions that were created by the ablation were measured after each experiment and were found to be symmetrical on both sides of the RF needle, thus, supporting the assumption of a symmetrical heating profile.
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Fig. 8. (a) In vivo experimental set-up. The imaging probe can be seen on top with the Panametrics V318 probe that is located on the side. (b) (top) Energy calculated from the 0.5–0.7MHz band vs. frame as was recorded by the passive transducer during transmission (green) and before transmission (blue). (bottom) A subtraction of the two.
The second set of measurements was made to help establish a relationship between the distance from the RF needle and the average LFB energy. A general increase of the mean LFB energy was observed as the distance from the RF needle decreased. To combine results from different experiments, normalization had to be performed; the energy values were normalized by the energy reflected from the RF needle itself. This was found to work well when there was no obstruction of the acoustic path between the US transducer and the RF needle, but when a large air cavity blocked part of the transmitted waves, the normalization produced erroneous values. To exclude calculations based on improper normalization, it was assumed that in the last window (No. 7) the LFB energies should have comparable values because the region inside window 7 was not exposed to the ablation. Three measurements were excluded based on this criterion. Calculations of the mean energy ratios and their variances were based on measurements performed over a relatively large region. The variance of the LFB and TFB energies, as described in the graphs of Figure 7, shows a substantial increase at temperatures over 60 C. A similar result was obtained by Gertner et al. (1997), who showed that a large variability of the backscatter coefficient occurs at temperatures above 60 C. The variance of the TFB energy in the reference window, which was not subjected to the heating, increased as well. The reason for this increase is probably due to motion that was induced by changes in the SOS and thermal expan-
sion. This result exemplifies the advantage of using the LFB energy over the TFB energy because the latter is more sensitive to natural scatterers. When treating a large region with high temperatures, the large amount of bubbles that are formed may create an acoustic barrier that prevents the ability to measure the LFB energy from areas beyond this barrier. Such acoustic barriers were clearly visualized in the in vitro experiments, where large lesions were produced (1–2 cm in diameter). When the ablation process ends and the tissue starts to cool down, the microbubbles created during the ablation dissolve quickly, leading to improved acoustic penetration. Large bubbles dissolve much slower and lead to the creation of the hyperechoic regions typically seen after ablation treatments. Echoes from these large bubbles appear as strong and sparse speckles in B-mode imaging, but they cannot really help in estimating lesion size and shape (Rhim and Dodd 1999). The mechanical index (MI) used during the experiments was 0.6. Thus, US induced cavitation is assumed to be only a minor contributor to the formation of microbubbles in the tissue. The thermal index (TI) used during US exposures was high (TI 5 3.0) due to the long transmission of 15 cycles. However, the duty cycle (4%) of the US transmission that was used throughout the experiments was low, thus, preventing significant tissue heating due to the US transmission. Even though all experiments were performed under the same parameters for the RFA instrument (output
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Fig. 9. Measurements performed during the treatment. (a)–(i) Display of the processed measurements performed during one of the ablation treatments. Images were improved using a median filter followed by an averaging filter and gamma correction (g 5 1/2). The dark red dot marks the location of the radio-frequency (RF) needle. (j) B-mode image of the tissue before the experiment. The RF needle is visualized as the bright dot near the middle of the image. (k) Section of the ablated liver after the procedure. (l) Temperature recorded during this experiment.
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Fig. 10. A comparison between manual and algorithmic lesion diameters from 13 measurements.
power, duration), the thermocouple recordings showed that different liver samples were heated at different rates and reached different maximal temperatures. Lesion diameters measured after a cool down period also varied significantly. This is probably due to the natural variability of the tissue samples, and may help explain some of the high variance that was observed in the in vitro experiments. However, the circular symmetry was observed in all of the measurements regardless of the produced lesion size. The turn-off periods, required for the acquisition of the US data, resulted in a duty cycle of 96%, which did not allow time for a significant cooling of the tissue. This is demonstrated in the two examples of temperature vs. time graphs depicted in Figures 6 and 9. It is assumed, therefore, that because the temperature remains elevated, additional microbubbles are formed and the existing ones mostly do not dissolve. Moreover, the US transmission that was applied during that period may have resulted in an increase in size of the microbubbles, due to the rectified diffusion mechanism. In this study spectrums were not included due to the usage of 64 samples segments, which leads to frequency resolution of 312.5 KHz (at 20 MHz sampling frequency). At this resolution it is difficult to interpret the spectrums and to obtain any conclusion regarding the source of the low frequency emissions.
In vivo A linear correlation coefficient of 0.743 was obtained between the manual measurements of the diameter of the lesion and measurements of the size of the region that underwent color change. This suggests that there is a correlation between the ablated areas and areas in which bubble formation occurred. Thus, detec-
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tion of tissue regions in which bubbles are being generated can be used to estimate lesion size. Because two different measurements types of the same parameter were compared, the optimal fit was forced to intercept the origin to appreciate the deviation of the slope from unity. It is clear, however, that the optimal fit may not necessarily intercept the origin. The cross-section produced by the sectioning of the liver, which was carried out after the experiment, probably did not always coincide with the scanned plane because it was performed manually. As the lesions showed great variation in shape, it is possible that even a millimeter shift between the planes could introduce large errors. We believe that the two outliers obtained in these measurements are due to lack of agreement between the scan plane and the sectioned plane. The algorithm performs temporal averaging of the LFB energy measurements, thus there is an inherent sensitivity to movements. However, if a higher US frame rate is used, the effect of movements may become negligible. To produce subharmonic responses and other low frequency emissions, which are produced by nonlinear bubble oscillations, several US cycles should be used (Biagi et al. 2005; Leighton 1994). Here, the transmission wave used consisted of 15 cycles to sufficiently excite the microbubbles. Increasing the number of cycles would increase the low frequency content of the backscattered signal. However, this long transmission wave leads to a degraded axial resolution. To efficiently produce subharmonic emissions and other low frequency emissions, the transmission frequency should be related to the bubble sizes in the media. Previous studies showed the ability of liver tissues to produce low frequency emissions during heating, when insonified with frequencies at the 3–4 MHz range (Mast et al. 2008, Mclaughlan et al. 2007). This frequency range imposes a limitation on the selection of the transducer, so that both LFB and TFB should be included in the transducer frequency band. It would be better to select the LFB as far from the TFB as possible to obtain a better rejection of natural scatterers. The bandwidth of the ultrasound probe limits this choice, as preserving a sufficiently high signal to noise ratio (SNR) is also required. Therefore, there is a tradeoff between the need to select the LFB as low as possible and maintaining a reasonable SNR for the measurements. In general, naturally occurring scatterers are not expected to produce any LFB emissions. However, due to the fact that the pulse has finite duration it contains low frequencies in it. If these low frequencies are close to the transmission frequency and are included in the transducer bandwidth, they are barely attenuated and may be quite strong. The usage of a high number of cycles
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narrows the pulse bandwidth to some extent. Nevertheless, low frequencies will be present and reflected from the ‘‘natural scatterers’’. In the in vivo study, the usage of a low frequency band (0.5–0.7 MHz) helped to efficiently attenuate the low frequency content of the transmission pulse that was reflected from the ‘‘natural scatterers.’’ CONCLUSIONS The behavior of the LFB energy measured from segments within the US scan plane during RFA was studied. Thermocouple measurements were used as a reference. In the in vitro studies, analysis of the LFB energies from segments, which were at the same distance from the RF needle as the thermocouples, showed a 3 dB increase at a typical temperature of 53 6 5 C. The measured LFB and TFB energies were higher at shorter distances from the RF needle, where the temperatures were higher. At temperatures above 60 C, the LFB and TFB energy measurements changed abruptly between adjacent time samples, producing a significantly higher variance. This indicates a highly dynamical process that probably resulted from tissue boiling. The LFB energy measurements were used by an algorithm, which performs a moving average calculation of these measurements to produce an image that highlights an area which corresponds to the treated area. A comparison between manual measurements and measurements obtained from the algorithm produced a linear correlation coefficient of r 5 0.743 and a slope of 0.94. This indicates a strong relation between the formation of microbubbles and the area being treated. Thus, LFB energy can provide useful information regarding the treatment progress and may help the physicians in controlling and monitoring RFA treatments. REFERENCES Amini AN, Ebbini ES, Georgiou TT. Noninvasive estimation of tissue temperature via high-resolution spectral analysis techniques. IEEE Trans Biomed Eng 2005;52:221–228. Arthur RM, Straube WL, Trobaugh JW, Moros EG. Noninvasive estimation of hyperthermia temperatures with ultrasound. Int J Hyperthermia 2005;21:589–600. Biagi E, Breschi L, Masotti L. Transient subharmonic and ultraharmonic acoustic emission during dissolution of free gas bubbles. IEEE Trans Ultrason Ferroelectr Freq control 2005;52:1048–1054. Buscarini E, Savoia A, Brambilla G, Menozzi F, Reduzzi L, Strobel D, H€ansler J, Buscarini L, Gaiti L, Zambelli A. Radio-frequency thermal ablation of liver tumors. Eur Radiol 2005;15:884–894. Chiou S, Liu J, Needleman L. Current status of sonographically guided radio-frequency ablation techniques. J Ultrasound Med 2007;26: 487–499. Daniels MJ, Varghese T, Madsen EL, Zagzebski JA. Noninvasive ultrasound-based temperature imaging for monitoring radiofrequency heating–Phantom results. Phys Med Biol 2007;52: 4827–4843.
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Fahey BJ, Nightingale KR, Stutz DL, Trahey GE. Acoustic radiation force impulse imaging of thermally and chemically induced lesions in soft tissues: Preliminary ex-vivo results. Ultrasound Med Biol 2004;30:321–328. Gaitini D, Zivari M, Abadi S, Goldberg NS, Adam D. Evaluating tissue changes with ultrasound during radio-frequency ablation. Ultrasound Med Biol 2005;34:586–597. Gazelle SG, Goldberg NS, Solbiati L, Livraghi T. Tumor ablation with radio-frequency energy. Radiology 2000;217:633–646. Gertner MR, Wilson BC, Sherar MD. Ultrasound properties of liver tissue during heating. Ultrasound Med Biol 1997;23:1395–1403. Kolokythas O, Gauthier T, Fernandez AT, Xie H, Timm BA, Cuevas C, Dighe MK, Mitsumori LM, Bruce MF, Herzka DA, Goswami GK, Andrews TR, Oas KM, Dubinsky TJ, Warren BH. Ultrasound based elastography: A novel approach to assess radio-frequency ablation of liver masses performed with expandable ablation probes. J Ultrasound Med 2008;27:935–946. Leighton TG. The acoustic bubble. San Diego: Academic Press; 1994. Liu X, Gong X, Yin C, Li J, Zhang D. Noninvasive estimation of temperature elevations in biological tissue using acoustic nonlinearity parameter imaging. Ultrasound Med Biol 2008;34:414–424. Lobik L, Leveillee RJ, Hoey MF. Geometry and temperature distribution during radio-frequency tissue ablation: An experimental ex vivo model. J Endourol 2005;19:242–247. Maass-Moreno R, Damianou CA. Noninvasive temperature estimation in tissue via ultrasound echo-shifts. Part 1. Analytical model. J Acoust Soc Am 1996;100:2514–2521. Maass-Moreno R, Damianou CA, Sanghvi NT. Noninvasive temperature estimation in tissue via ultrasound echo-shifts. Part 2. In vitro study. J Acoust Soc Am 1996;100:2522–2530. Mast TD, Salgaonkar VA, Karunakaran C, Besse JA, Datta S, Holland CK. Acoustic emissions during 3.1 MHz ultrasound bulk ablation in vitro. Ultrasound Med Biol 2008;34:1434–1448. Mclaughlan J, Rivens I, Haar G. Cavitation detection in ex vivo bovine liver tissue exposed to high-intensity focused ultrasound (HIFU). IEEE Int Symp Biomed Imaging 2007;1124–1127. Miller NR, Bamber JC, Meaney PM. Fundamnetal limitations of noninvasive temperature imaging by means of ultrasound echo strain estimation. Ultrasound Med Biol 2002;28:1319–1333. Rhim H, Dodd GD. Radio-frequency thermal ablation of liver tumors. J Clin Ultrasound 1999;27:221–229. Sanghvi NT, Fry FJ, Bihrle R, Foster RS, Phillips MH, Syrus J, Zaitsev A, Hennige C. Microbubbles during tissue treatment using high intensity focused ultrasound. IEEE Ultrason Symp 1995;1571–1574. Seip R, Ebbini ES. Noninvasive estimation of tissue temperature response to heating fields using diagnostic ultrasound. IEEE Trans Biomed Eng 1995;42:828–839. Simon C, VanBaren P, Ebbini ES. Two-dimensional temperature estimation using diagnostic ultrasound. IEEE Trans Ultrason Ferroelectr Freq Control 1998;45:1088–1099. Souchon R, Bouchoux G, Maciejko E, Lafon C, Cathignol D, Bertrand M, Chapelon J. Monitoring the formation of thermal lesions with heat induced echo-strain imaging: A feasibility study. Ultrasound Med Biol 2005;31:251–259. Tateishi R, Shiina S, Teratani T, Obi S, Sato S, Koike Y, Fujishima T, Yoshida H, Kawabe T, Omata M. Percutaneous radio-frequency ablation of hepatocellular carcinoma–An analysis of 1000 cases. Cancer 2005;103:1201–1209. Techavipoo U, Vaghese T, Chen Q, Stiles TA, Zagzebski JA, Frank GR. Temperature dependence of ultrasonic propagation speed and attenuation in excised canine liver tissue measured using transmitter and reflected pulses. J Acoust Soc Am 2004; 115:2859–2865. Varghese T, Zagzebski JA, Chen Q, Techavipoo U, Frank G, Johnson C, Wright A, Lee FT. Ultrasound monitoring of temperature change during radio-frequency ablation: Preliminary in vivo results. Ultrasound Med Biol 2001;28:321–329.
doi:10.1016/j.ultrasmedbio.2010.11.008
d
Original Contribution MONITORING RADIO-FREQUENCY THERMAL ABLATION WITH ULTRASOUND BY LOW FREQUENCY ACOUSTIC EMISSIONS – IN VITRO AND IN VIVO STUDY ITAI WINKLER and DAN ADAM Biomedical Engineering, Technion City, Haifa, Israel (Received 11 June 2010; revised 5 November 2010; in final form 12 November 2010)
Abstract—The object of this study was to evaluate the monitoring of thermal ablation therapy by measuring the nonlinear response to ultrasound insonation at the region being treated. Previous reports have shown that during tissue heating, microbubbles are formed. Under the application of ultrasound, these microbubbles may be driven into nonlinear motion that produces acoustic emissions at sub-harmonic frequencies and a general increase of emissions at low frequencies. These low frequency emissions may be used to monitor ablation surgery. In this study, a modified commercial ultrasound system was used for transmitting ultrasound pulses and for recording raw RF-lines from a scan plane in porcine (in vitro) and rabbit (in vivo) livers during radio-frequency ablation (RFA). The transmission pulse was 15 cycles in length at 4 MHz (in vitro) and 3.6 MHz (in vivo). Thermocouples were used for monitoring temperatures during the RFA treatment. In the in vitro experiments, recorded RF signals (A-lines) were segmented, and the total energy was measured at two different frequency bands: at a low frequency band (LFB) of 1–2.5 MHz and at the transmission frequency band (TFB) of 3.5–4.5 MHz. The mean energy at the LFB and at the TFB increased substantially in areas adjacent to the RF needle. These energies also changed abruptly at higher temperatures, thus, producing great variance in the received energy. Mean energies in areas distant from RF needle showed little change and variation during treatment. It was also shown that a 3 dB increase of energy at the low frequency band was typically obtained in regions in which temperature was above 53.3 ± 5 C. Thus, this may help in evaluating regions undergoing hyperthermia. In the in vivo experiments, an imaging algorithm based on measuring the LFB energy was used. The algorithm performs a moving average of the LFB energies measured at segments within the scan plane.Results show that a colored region is formed on the image and that it is similar in size to a measurement of the lesion from gross pathology, with a correlation coefficient of 0.743. (E-mail: [email protected]) Ó 2011 Published by Elsevier Inc. on behalf of World Federation for Ultrasound in Medicine & Biology. Key Words: Radio-frequency ablation (RFA), Treatment monitoring, Acoustic emissions, Ultrasound, Signal processing.
ultrasound (HIFU), have been suggested and introduced into clinics. RFA is considered the most commonly used approach for liver tumor ablation. The RFA procedure includes the insertion of a radiofrequency (RF) electrode (RF needle) into the tumor, usually aided by ultrasound imaging guidance. After the placement of the RF needle at the target location, a strong alternating electric current is delivered to the tumor. The treatment goal is to cause immediate cell death by heat to the tumor cells and their adjacent marginal area to ensure treatment success but with minimal damage to nearby healthy tissue. The electrical current causes agitation of ions that causes frictional heat, leading to quick temperature elevation and tissue ablation around the RF needle and in adjacent areas (Daniels et al. 2007; Gazelle et al. 2000). Lesion size depends on the RF output power and duration and it
INTRODUCTION Treatment of primary and secondary malignant hepatic tumors, considered to be one of the most common malignant tumors worldwide, is limited. Chemotherapy and radiation therapy are in many cases ineffective against primary and secondary hepatic tumors, thus, resection surgery is a potential solution. However, most patients are not candidates for this surgery for various reasons (Buscarini et al. 2005; Chiou et al. 2007). In recent years, minimally invasive ablation approaches, such as radio-frequency ablation (RFA) and laser ablation, or noninvasive approaches such as high-intensity focused
Address correspondence to: Itai Winkler, Biomedical Engineering, 32000 Technion City, Haifa 32000, Israel. E-mail: Itai. [email protected] 755
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varies greatly between different individuals and different tissue types. Blood perfusion, tissue properties and anisotropy make it difficult to predict for each case the resultant lesion size and shape. Ultrasound imaging is the most common technique for guiding the insertion of the RF needle. (Buscarini et al. 2005; Chiou et al. 2007; Rhim and Dodd 1999; Tateishi et al. 2005). However, B-mode imaging does not provide sufficient visualization of the lesion formation and treatment progress (Buscarini et al. 2005; Chiou et al. 2007; Rhim and Dodd 1999; Varghese et al. 2001). Eventually, the lack of sufficient real-time monitoring ability leads to a high recurrence rate, greater than 33% (Varghese et al. 2001), and repeated surgeries. During the clinical ablative process, numerous microbubbles are formed because usually, the heating is fast and temperatures reach boiling temperatures. These microbubbles are strong reflectors and, thus, produce strong echoes that enable their visualization in regular B-mode imaging. This appearance of new reflectors in the treated area is currently used to assess the heated region. However, this visualization technique is poor for assessing the true damage caused to the tumor and the completion of the treatment (Gazelle et al. 2000). Many studies have attempted to use ultrasound (US) thermometry to image temperature changes. The most common approaches are based on measuring echo shifts (Maass-Moreno and Damianou 1996; Maass-Moreno et al. 1996; Simon et al. 1998; Souchon et al. 2005). Other approaches include measuring frequency shifts (Amini et al. 2005; Gaitini et al. 2005; Seip and Ebbini 1995) and changes in backscattered energy (Arthur et al. 2005). Echo shifts result from two main mechanisms: changes in the speed of sound (SOS) and tissue expansion. However, the SOS does not change monotonically with temperature (Techvipoo et al. 2004). Moreover, the natural variation between individuals creates the need for specific calibration curves for each treatment (Arthur et al. 2005). Another problem with echo shift and frequency shift methods is their sensitivity to movements and bubble formation. Both methods are based on tracking echo or frequency displacements by some correlation technique. Therefore, both techniques are very sensitive to tissue motions and bubble formation that lead to severe decorrelation between measurements. Therefore, the echo shift and frequency shift methods work well for small temperature elevations but fail for high temperature elevations typical in ablation surgeries (Gaitini et al. 2005; Miller et al. 2002). Other studies have directly addressed the problem of treatment monitoring, e.g., using elastography (Fahey et al. 2004; Kolokythas et al. 2008; Varghese et al. 2001) or the acoustic nonlinearity parameter (b/a) (Liu
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et al. 2008). Elastography also suffers from the decorrelation of signals due to bubble formation. Therefore, elastography measurements are usually made after a cool down period, when most microbubbles have been dissolved and the lesion becomes stiffer than its surrounding. During heating, even at small temperature elevations (at the early stages of the ablation treatment), microbubbles are formed as the temperature rises (Mclaughlan et al. 2007; Sanghvi et al. 1995). These microbubbles can be detected and distinguished from the surrounding tissue by their typical emissions, caused when the microbubbles are insonified by US (Biagi et al 2005; Leighton 1997; Mast et al. 2008). A free bubble in liquid media, insonified by US with sufficient intensity and duration (number of cycles), may emit different subharmonics (Biagi et al. 2005; Leighton 1997) as well as other low frequency emissions (Leighton 1997; Mast et al. 2008). In this study, an attempt was made to exploit these emissions at low frequencies for the assessment of the treatment progress. A modified commercial US system was used for transmitting and receiving US waves, so as to interrogate a scan plane during RFA. From the acquired RF signals, low frequency band (LFB) energies were evaluated at different regions within the scan plane. Thermocouple measurements were used as control measurements. MATERIALS AND METHODS Experiment set-up In vitro. Fresh whole porcine livers were purchased from a local slaughter house less than 48 h post mortem and were used on the same day of purchase. Each liver was brought to room temperature (19–23 C) and two to three samples, no smaller than 7 3 4 3 2 cm3, were cut from it. The samples were visually checked to be certain they contained a minimum amount of large blood vessels. The samples were held by a specially designed sample holder, which has several holes in the front to allow insertion of the RF needle and two needle thermocouples (Fig. 1). The sample holder has a hollow floor to allow contact between the tissue samples and a metal plate that serves as the dispersive electrode. The metal plate was connected in series to a power resistor, to achieve total impedance of around 120 U. The RF ablation system used with an internally cooled needle (cool-tip RF, Vallylab; Tyco Healthcare Group LP, Boulder, CO, USA) theoretically produces a lesion of a cylindrical shape (Lobik et al. 2005). Two needle thermocouples, with a diameter of 0.25 mm (Type T; Omega Engineering, Stamford, CT, USA) and placed inside 10 cm syringe needles, were inserted into
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Fig. 1. (a) Experimental set-up. The liver samples were fixed by the sample holder and were in direct contact with the metal plate under it, serving as the dispersive electrode. The radio-frequency (RF) needle with the two needle thermocouples were placed through designated holes in front of the sample holder and perpendicular to US beam direction. (b) Cross-section after a typical radio-frequency ablation (RFA) treatment.
the tissue through the sample holder holes, 2.5 mm and 5 mm away from the RF needle (Fig. 1). The RF needle and the two thermocouples were oriented along a line parallel to the ultrasound transducer’s surface and perpendicular to the imaged plane. The temperature was recorded every 1 s during the whole experiment by software (Virtual Bench Logger; National Instruments, Austin, TX, USA). A Vivid III US imaging system with a S5 phased array probe (GE Healthcare, Giles, UK) was used for transmitting and receiving the US beams at the selected scan plane. The transmission frequency was set to be 4 MHz with bursts of 15 cycles in length. The transducer’s focus was set to a depth similar to the depth of the RF needle or at most 1 cm below it. Output intensity was set to produce a mechanical index of MI 5 0.6. The transducer was held fixed during the entire experiment. The raw US analog signals (RF lines) were acquired using custom software, designed for acquisition of the Vivid III analog signals before they are processed by the imaging system. Because the current produced by the RF generator disrupts the US data acquisition, a switch box was designed to turn off the ablation current for brief periods during data acquisition. The switch box was controlled by the custom data acquisition software. The current was turned off every 5 s for 0.2 s, during which the RF signals from the whole scan plane, as well as the thermocouple measurements, were acquired. The resultant duty cycle of the RFA treatment was 96%. Twelve acquisitions (frames), consisting of all the RF lines comprising a scan plane, were made as reference, before the ablation was started, and 24 were made after the ablation current was terminated. The RF needle is 15 cm long and has an exposed tip of 3 cm in length. The tips of the thermocouples were aligned with the center of the exposed part of the RF needle. Insertion of the RF needle and the thermocouples was
done under regular B-mode imaging, to confirm the correct placement of the instruments. The US imaging was also used to make sure that the scan plane and the chosen target for ablation consisted of a minimal amount of large scatterers. The metal plate located under the tissue sample was clearly visualized in the B-mode image and served as an indicator that the US beams were not blocked from reaching the region-of-interest (ROI). The RF generator output was manually set to 16 watts. The duration of each RF treatment was around 15 min, which was the time needed to achieve a desired temperature of 80–90 C at thermocouple 1 (closer to the RF needle). In vivo. The in vivo experiments were carried out using rabbits (NZW, 3.5–4.5 kg). The protocol for the experiments was approved by the Technion committee for the supervision of animal experiments. Rabbits were anesthetized with a mixture of ketamine (35 mg/kg), xylazine (2 mg/kg), acepromazine (0.3 mg/kg) and butorphanol (0.5 mg/kg) (SC). An additional dose of 3 mg/kg ketamine was injected (IV) every 10 min. Vital signs and anesthesia strength were checked frequently. The liver was exposed by an abdominal section. The liver was pulled outward slightly and two pads were inserted beneath the liver to expose a larger surface area to allow better placement of the instruments (US transducer, RF needle and thermocouples). The US transducer, RF needle and thermocouples were fixed during the entire procedure. As detailed in the in vitro section, a B-mode image was acquired before the treatment started to identify the locations of the RF needle and the thermocouples within the tissue. RF current was administered for 5–10 min with power output of 20–40 Watts. Fifteen rabbits were included in the study. Each rabbit was subjected to a single RFA treatment.
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After the procedure and a cool down period, the rabbits were euthanized (Thiopental, 650 mL/kg, IV). The ablated areas were incised for visual examination of the treatment outcome and for lesion measurements. The in vivo set-up differed from the in vitro phase for certain parameters. The transmission frequency was 3.6 MHz with bursts of 15 cycles in length and with an intensity of MI 5 0.6. Insertions of the RF needle and the thermocouples were made through a custom built holder with the same spacing between the RF needle and the thermocouples as was described in the in vitro section. Data processing and analysis In vitro. The acquired RF lines were sampled at 20 MHz (by the Vivid III) and were filtered with two band pass filters (BPF) with pass bands in the following range: low frequency band (LFB) 5 1–2.5 MHz, transmission frequency band (TFB) 5 3.5–4.5 MHz. The LFB was measured to study the behavior of the low frequency emissions from different areas within the imaging plane during RFA. The higher frequency band (TFB), which includes the transmission frequency, was measured to study the changes of the tissue echogenicity during RFA. The US transmit pulse had a bandwidth of 0.5 MHz. The filtered signals were uniformly segmented to 64 sample windows, which correspond to 2.5 mm in length, without overlaps. LFB and TFB energies in each window were calculated using the L2 norm:
Eðf1 ; f2 Þ5
ð f2 f1
jXðf Þj df z 2
N 1 X n50
jx½n hBPF j
2
(1)
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where Eðf1 ; f2 Þ denotes the energy in the band specified by f1 ; f2 , Xðf Þ is the Fourier transform of a time signal, x½n is the sampled signal and hBPF is the filter impulse response. Because the US probe used here was a phased array probe, energy calculations made from the acquired A-lines are given in a polar coordinate systemðr; qÞ, where r represents the distance from the center of the transducer center and q represents the angle of the beam in relation to the normal to the transducer’s surface. The energy measurements were first transformed to a Cartesian coordinate system using a weighted average of the two nearest energy measurements in the polar representation. Weights were assigned according to the distances as presented in eqn (2): Eðx; yÞ5W1 E1 ðr; qÞ1W2 E2 ðr; qÞ d2 W1 5 ; d1 1d2
W2 5
d1 d1 1d2
(2) (3)
where E1 ; E2 represent the energies of the neighboring locations in polar coordinates, and d1 ; d2 are the Euclidean distances between the vertex in the Cartesian grid to the closest polar neighbors. After obtaining the LFB and TFB energy maps for the entire scan plane, data was analyzed in three ways, as follows. (1) The average LFB energy was calculated in two windows of sizes 1.6 mm 3 2.5 mm (lateral x axial), as described in Figure 2. Each of the two window centers is positioned at the same distance from the RF needle as that of the thermocouples and at the
Fig. 2. Experiment set-up (cross-section view). The radio-frequency (RF) needle and the two thermocouples separated by 2.5 mm spacing on a line parallel to the transducers surface. Two regions (windows) are centered in the same distances from the RF needle as the thermocouples and on the same line parallel to the transducers surface.
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same depth. When the average LFB energy within a window increased by 3 dB, the temperature at that moment, as measured by the corresponding thermocouple, was registered. This enabled us to correlate between the LFB energy and the temperature, assuming that the temperature at the thermocouple sites is similar to the temperature at the symmetrical points on the other side of the RF needle. This assumption is likely valid because liver tissue is relatively homogenous, as evident by the circular symmetry of the resultant lesion. (2) The average LFB and TFB energies were calculated for seven adjacent windows, as described in Figure 3. All windows were on the same line formed by the RF needle and the thermocouples and at the same depth. All windows were non-overlapping (Fig. 3). The window size was 2 mm 3 2.5 mm (lateral 3 axial), thus enabling a reasonable spatial resolution. The axial size of the window was limited by the number of samples used for the energy calculation in each window. Average energies for these windows were calculated over 50 consecutive frames from the moment thermocouple 1 reached 60 C in each of the samples. (3) Two large windows within the scan plane of 5 mm 3 5 mm were selected. The first window was a large region within the ablated area. The second window was centered at a distant location from the RF needle, assumed not to be affected by the treatment (Fig. 4). The average LFB and TFB energies and energy variances were calculated for the two regions and for seven consecutive frames from three different start-
Fig. 3. Illustration of the location of the windows where the low frequency band (LFB) and transmission frequency band (TFB) energies were measured. In each of the windows numbered 1–7, the mean LFB and TFB energies were calculated.
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ing points. The starting points were chosen to be the moments thermocouple 1 reached temperatures of 50, 60 and 70 C. All data processing was performed using MATLAB (Mathworks, Natick, MA, USA). In vivo. The data measured during the in vivo experiments were processed similarly to that obtained during the in vitro experiments and they were also processed to provide images of the calculated parameters. The formation of image sequences that allow monitoring of the treatment first requires the computation of the LFB energy from all segments within the scan plane or in a predefined ROI. The energy band that was studied, 0.5–0.7 MHz, was somewhat different than during the in vitro studies. This band produced better results due to a better rejection of natural scatterers. To confirm that the energy in this frequency range was due to microbubbles, a single element transducer with a center frequency of 0.5 MHz (V318; Panametrics, Waltham, MA, USA) was used to passively record signals during the experiment. For each of the Vivid III recordings, the control system issued a trigger to a computer scope (Compuscope CS14100; Gage Applied Inc., Lachine, Canada) that recorded 540 K samples at a sampling rate of 25 MHz. The first 120 K samples were recorded before the Vivid III received the trigger and the rest were recorded after the trigger. Energy in the 0.5–0.7 MHz band was calculated from two different segments in the recorded signal. The first segment was before the Vivid III was activated and the second was after it was activated, i.e., during ultrasound transmission. The segments were 110 K samples in length. At each time step, an image was created by calculating the average of the last 10 frames. Because the ablated regions were expected to produce higher LFB energies than untreated regions and natural scatterers are severely attenuated due to the low frequency range used for the LFB energy calculations, the resultant image is expected to produce a contrast between treated and untreated areas. After calculating all the pixel values that comprise an image, spatial filtering was performed by a median filter (a square mask of 5 3 5 pixels, lateral 3 axial), followed by a spatial averaging filter (of 5 3 3 pixels, lateral 3 axial), to improve the image and reduce artifacts, and by gamma correction ðg51=2Þ. The median filter was used to remove high energy measurements from the images. These high energy measurements are caused by the wide energy distribution. Removal of these high energies is desired for creating a tighter energy range before assigning colors according to the energy levels. The averaging filter was implemented to smooth the image and produce a visually clearer result.
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Fig. 4. The two selected windows: each window was of a size of 5 3 5 mm. The ablation window corresponds to an area that is subjected to ablation. The reference window is at a distant location from the radio-frequency (RF) needle and is not subjected to ablation.
RESULTS In vitro Energy elevations of 3 dB in the LFB were measured from the graphs of energy vs. time for the two windows and the temperatures associated with these 3 dB increases were found from the thermocouple recordings to be 53.3 6 5 C (mean 6 std, n 5 36). At higher temperatures, rapid changes in the LFB energy could be observed, which may be explained by rapid formation and implosion or dissolution of microbubbles. Figure 5 presents averaged results of the LFB and TFB energies, recorded from the moment that thermocouple 1 reached 60 C, as a function of the distance from the RF needle, calculated over 18 experiments. All energies were normalized by the energy measured before the beginning of the heating within a rectangular
window whose center was the RF needle, averaged over the first 10 recorded frames. Mean energies are presented in arbitrary units. The bars represent the standard error of the measurements. It can be seen in Figure 5 that both the LFB energy and TFB energy decrease as the distance from the RF needle increases. Figure 6 depicts representative recordings of the LFB and TFB energies measured at different single ‘‘pixels’’, i.e., energy calculated from a single RF segment, taken from different locations within the windows illustrated in Figure 4 during RFA treatment. Energies are displayed in dB relative to the mean noise level, which was measured in a large region of the image before the heating started. The energy measured in the ablated areas tended to vary abruptly from frame to frame, thus, producing
Fig. 5. Mean energy measured during 50 frames from the moment thermocouple 1 reached 60 C. (a) Mean low frequency band (TFB) energy. (b) Mean transmission frequency band (LFB) energy. (average 6 standard error).
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Fig. 6. Representative recordings of energy measured from four segments during a typical experiment. Each pixel represents a small area in the scan plane. (a) Low frequency band (LFB) energies measured from two nonablated segments. (b) LFB energies measured from two ablated segments. (c) Transmission frequency band (TFB) energies measured from two nonablated segments. (d) TFB energies measured from two ablated segments. (e) Temperature measured by the two thermocouples during the experiment.
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a high variance of energy at both frequency bands compared to the base level variance. In the non-heated areas, the LFB and TFB energies changed slowly and were usually lower than the measured energies in the heated areas. In the non-heated regions, rapid changes in the energy were recorded from time to time, but in general, the energy graphs were smoother. In Figure 6, some jumps can be seen to occur sporadically at the time interval 400–800 s. The cause of these occasional rapid changes is not clear but we speculate that it may result from some type of treatment induced movements. The LFB energy measured in the non-heated areas was typically smoother than the TFB energy recorded from those regions. These points are depicted in Figure 7. Figure 7 presents energies (average 6 standard error) calculated from two large regions (windows) within the scan plane at different temperatures. Results are for 18 measurements; the bars represent the standard error of the measurements. The average normalized LFB and TFB energies measured in the window within the heated region increased as the temperature increased. Significant increases of LFB energy can be seen in ablated areas when heating from 50 to 70 C (p , 0.05). A significant increase of the TFB energy was also observed at these temperatures (p , 0.04). Both energies showed similar behavior. The variance of the LFB and TFB energies, measured at segments within the heated region and the reference one, also both increased with temperature. A significant increase in the mean LFB and TFB variances was found between the base level of 22 to 70 C (p , 0.02 for the LFB and p , 0.01 for the TFB).
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This was probably due to large bubbles being trapped in the tissue. The diameter of the region that changed its color was measured and compared with the manual pathology measurements made after a cool down period, performed after the animal was euthanized. The manual measurement of the lesion was taken along a line parallel to the liver lobe surface (the surface that is distal from the US transducer) and through the point of the RF needle location. This liver lobe surface was in contact with the surface of the rabbit’s stomach during the experiment. Because the rabbit stomach surface was visible from the B-mode image taken before the initiation of ablation, the orientation of that surface in the algorithm images could be found. Thus, the image measurement was performed along a line passing through the RF needle and with the same direction as the stomach surface. The boundaries for the lesion measurements were found by including all pixel values in the range of 5% to 100% along the line of measurement and were taken from the frame that presented the largest measurement along this line. The comparison between manual and image measurements is presented in Figure 10. As can be seen in Figure 10, a correlation coefficient of 0.743 was found between the lesion diameter, as obtained by the manual measurement, and the region in which the color changed, as measured from the images of the algorithm. The error calculated from these 14 experiments is –0.3 6 1.4 [mm] (average 6 std) or 3% 6 13.7%. Two measurements were removed as outliers (see Discussion section). DISCUSSION
In vivo Figure 8 presents the energy recorded within the 0.5–0.7 MHz band as a function of time (frame), which increased with time from the moment the ablation procedure was initiated and decreased once the ablation was terminated. This indicates that the emissions in the low frequency band were caused by the treatment. Some residual energy, caused by the RF generator, was also present, but this value was constant as can be seen from the graphs. An example of the algorithm output, together with the temperature measurements and a B-mode image acquired before the initiation of the ablation, are presented in Figure 9. As can be seen, change in color that correspond to an increase of the LFB energies occur in the region associated with the probe location. The region in which color changed increased in size as long as the ablation current was on and it started to shrink once the ablation current was stopped. Even after a cool down period, some areas still produced increased LFB energy.
In vitro In vitro tissue samples had large bubbles and air cavities trapped in the blood vessels. These bubbles appear as strong echoes in the B-mode image and produce artifacts in the experiments. When the air cavity is large enough it can cause shadowing and block the US waves. To decrease the probability of having trapped air, the liver samples were taken from the distal part of each liver lobe, where smaller blood vessels are present. Tissue samples were also gently massaged to remove the trapped air from inside the large vessels. Despite these efforts, six experiments were excluded due to strong echoes in areas near the RF needle. The liver samples used in the in vitro experiments were not degassed. From our experience, degassing could only be partially performed and may damage the sample. The degassing process may also change the ability to form microbubbles within the tissue due to the decreased gas concentration in the liquids. In previous papers, degassing was applied only for short durations (Mclaughlan et al.
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Fig. 7. (a) and (b) Low frequency band (LFB) and transmission frequency band (TFB) energy ratio at 50, 60, and 70 C. Ratio is calculated in relations to mean energy at the same region prior to RFA treatment. (c) and (d) Averaged energy variance of in the selected region. All values in these graphs are calculated from 18 experiments (average 6 standard error).
2007; Souchon et al. 2005). Other researchers simply suspended the tissue in a partially degassed buffer (Mast et al. 2008). The purpose of the first set of experiments was to find a typical temperature where a 3 dB increase in the measured LFB energy was observed. It was shown that such an increase usually occurs when temperatures of 53.3 6 5 C are reached. This result indicates that at these temperatures microbubbles were already formed, which increased the back scattered energy. This comparison is made under the assumption that the temperatures on both sides of the RF needle are symmetrically distributed. It would have been more accurate to measure the energy right at the thermocouples, however, echoes and reverber-
ations from the thermocouples and the RF needle disrupt the energy measurements. Moreover, the insertion of the thermocouples into the tissue may produce cavities filled with air, further affecting energy measurements from this area. Therefore, all energy measurements were made at regions that did not include the thermocouples but were at symmetrical distances on the other side of the RF needle and with a safety margin of 2 mm from it, assuming that liver tissues are relatively homogenous and that there is no asymmetric perfusion involved. The lesions that were created by the ablation were measured after each experiment and were found to be symmetrical on both sides of the RF needle, thus, supporting the assumption of a symmetrical heating profile.
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Fig. 8. (a) In vivo experimental set-up. The imaging probe can be seen on top with the Panametrics V318 probe that is located on the side. (b) (top) Energy calculated from the 0.5–0.7MHz band vs. frame as was recorded by the passive transducer during transmission (green) and before transmission (blue). (bottom) A subtraction of the two.
The second set of measurements was made to help establish a relationship between the distance from the RF needle and the average LFB energy. A general increase of the mean LFB energy was observed as the distance from the RF needle decreased. To combine results from different experiments, normalization had to be performed; the energy values were normalized by the energy reflected from the RF needle itself. This was found to work well when there was no obstruction of the acoustic path between the US transducer and the RF needle, but when a large air cavity blocked part of the transmitted waves, the normalization produced erroneous values. To exclude calculations based on improper normalization, it was assumed that in the last window (No. 7) the LFB energies should have comparable values because the region inside window 7 was not exposed to the ablation. Three measurements were excluded based on this criterion. Calculations of the mean energy ratios and their variances were based on measurements performed over a relatively large region. The variance of the LFB and TFB energies, as described in the graphs of Figure 7, shows a substantial increase at temperatures over 60 C. A similar result was obtained by Gertner et al. (1997), who showed that a large variability of the backscatter coefficient occurs at temperatures above 60 C. The variance of the TFB energy in the reference window, which was not subjected to the heating, increased as well. The reason for this increase is probably due to motion that was induced by changes in the SOS and thermal expan-
sion. This result exemplifies the advantage of using the LFB energy over the TFB energy because the latter is more sensitive to natural scatterers. When treating a large region with high temperatures, the large amount of bubbles that are formed may create an acoustic barrier that prevents the ability to measure the LFB energy from areas beyond this barrier. Such acoustic barriers were clearly visualized in the in vitro experiments, where large lesions were produced (1–2 cm in diameter). When the ablation process ends and the tissue starts to cool down, the microbubbles created during the ablation dissolve quickly, leading to improved acoustic penetration. Large bubbles dissolve much slower and lead to the creation of the hyperechoic regions typically seen after ablation treatments. Echoes from these large bubbles appear as strong and sparse speckles in B-mode imaging, but they cannot really help in estimating lesion size and shape (Rhim and Dodd 1999). The mechanical index (MI) used during the experiments was 0.6. Thus, US induced cavitation is assumed to be only a minor contributor to the formation of microbubbles in the tissue. The thermal index (TI) used during US exposures was high (TI 5 3.0) due to the long transmission of 15 cycles. However, the duty cycle (4%) of the US transmission that was used throughout the experiments was low, thus, preventing significant tissue heating due to the US transmission. Even though all experiments were performed under the same parameters for the RFA instrument (output
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Fig. 9. Measurements performed during the treatment. (a)–(i) Display of the processed measurements performed during one of the ablation treatments. Images were improved using a median filter followed by an averaging filter and gamma correction (g 5 1/2). The dark red dot marks the location of the radio-frequency (RF) needle. (j) B-mode image of the tissue before the experiment. The RF needle is visualized as the bright dot near the middle of the image. (k) Section of the ablated liver after the procedure. (l) Temperature recorded during this experiment.
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Fig. 10. A comparison between manual and algorithmic lesion diameters from 13 measurements.
power, duration), the thermocouple recordings showed that different liver samples were heated at different rates and reached different maximal temperatures. Lesion diameters measured after a cool down period also varied significantly. This is probably due to the natural variability of the tissue samples, and may help explain some of the high variance that was observed in the in vitro experiments. However, the circular symmetry was observed in all of the measurements regardless of the produced lesion size. The turn-off periods, required for the acquisition of the US data, resulted in a duty cycle of 96%, which did not allow time for a significant cooling of the tissue. This is demonstrated in the two examples of temperature vs. time graphs depicted in Figures 6 and 9. It is assumed, therefore, that because the temperature remains elevated, additional microbubbles are formed and the existing ones mostly do not dissolve. Moreover, the US transmission that was applied during that period may have resulted in an increase in size of the microbubbles, due to the rectified diffusion mechanism. In this study spectrums were not included due to the usage of 64 samples segments, which leads to frequency resolution of 312.5 KHz (at 20 MHz sampling frequency). At this resolution it is difficult to interpret the spectrums and to obtain any conclusion regarding the source of the low frequency emissions.
In vivo A linear correlation coefficient of 0.743 was obtained between the manual measurements of the diameter of the lesion and measurements of the size of the region that underwent color change. This suggests that there is a correlation between the ablated areas and areas in which bubble formation occurred. Thus, detec-
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tion of tissue regions in which bubbles are being generated can be used to estimate lesion size. Because two different measurements types of the same parameter were compared, the optimal fit was forced to intercept the origin to appreciate the deviation of the slope from unity. It is clear, however, that the optimal fit may not necessarily intercept the origin. The cross-section produced by the sectioning of the liver, which was carried out after the experiment, probably did not always coincide with the scanned plane because it was performed manually. As the lesions showed great variation in shape, it is possible that even a millimeter shift between the planes could introduce large errors. We believe that the two outliers obtained in these measurements are due to lack of agreement between the scan plane and the sectioned plane. The algorithm performs temporal averaging of the LFB energy measurements, thus there is an inherent sensitivity to movements. However, if a higher US frame rate is used, the effect of movements may become negligible. To produce subharmonic responses and other low frequency emissions, which are produced by nonlinear bubble oscillations, several US cycles should be used (Biagi et al. 2005; Leighton 1994). Here, the transmission wave used consisted of 15 cycles to sufficiently excite the microbubbles. Increasing the number of cycles would increase the low frequency content of the backscattered signal. However, this long transmission wave leads to a degraded axial resolution. To efficiently produce subharmonic emissions and other low frequency emissions, the transmission frequency should be related to the bubble sizes in the media. Previous studies showed the ability of liver tissues to produce low frequency emissions during heating, when insonified with frequencies at the 3–4 MHz range (Mast et al. 2008, Mclaughlan et al. 2007). This frequency range imposes a limitation on the selection of the transducer, so that both LFB and TFB should be included in the transducer frequency band. It would be better to select the LFB as far from the TFB as possible to obtain a better rejection of natural scatterers. The bandwidth of the ultrasound probe limits this choice, as preserving a sufficiently high signal to noise ratio (SNR) is also required. Therefore, there is a tradeoff between the need to select the LFB as low as possible and maintaining a reasonable SNR for the measurements. In general, naturally occurring scatterers are not expected to produce any LFB emissions. However, due to the fact that the pulse has finite duration it contains low frequencies in it. If these low frequencies are close to the transmission frequency and are included in the transducer bandwidth, they are barely attenuated and may be quite strong. The usage of a high number of cycles
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narrows the pulse bandwidth to some extent. Nevertheless, low frequencies will be present and reflected from the ‘‘natural scatterers’’. In the in vivo study, the usage of a low frequency band (0.5–0.7 MHz) helped to efficiently attenuate the low frequency content of the transmission pulse that was reflected from the ‘‘natural scatterers.’’ CONCLUSIONS The behavior of the LFB energy measured from segments within the US scan plane during RFA was studied. Thermocouple measurements were used as a reference. In the in vitro studies, analysis of the LFB energies from segments, which were at the same distance from the RF needle as the thermocouples, showed a 3 dB increase at a typical temperature of 53 6 5 C. The measured LFB and TFB energies were higher at shorter distances from the RF needle, where the temperatures were higher. At temperatures above 60 C, the LFB and TFB energy measurements changed abruptly between adjacent time samples, producing a significantly higher variance. This indicates a highly dynamical process that probably resulted from tissue boiling. The LFB energy measurements were used by an algorithm, which performs a moving average calculation of these measurements to produce an image that highlights an area which corresponds to the treated area. A comparison between manual measurements and measurements obtained from the algorithm produced a linear correlation coefficient of r 5 0.743 and a slope of 0.94. This indicates a strong relation between the formation of microbubbles and the area being treated. Thus, LFB energy can provide useful information regarding the treatment progress and may help the physicians in controlling and monitoring RFA treatments. REFERENCES Amini AN, Ebbini ES, Georgiou TT. Noninvasive estimation of tissue temperature via high-resolution spectral analysis techniques. IEEE Trans Biomed Eng 2005;52:221–228. Arthur RM, Straube WL, Trobaugh JW, Moros EG. Noninvasive estimation of hyperthermia temperatures with ultrasound. Int J Hyperthermia 2005;21:589–600. Biagi E, Breschi L, Masotti L. Transient subharmonic and ultraharmonic acoustic emission during dissolution of free gas bubbles. IEEE Trans Ultrason Ferroelectr Freq control 2005;52:1048–1054. Buscarini E, Savoia A, Brambilla G, Menozzi F, Reduzzi L, Strobel D, H€ansler J, Buscarini L, Gaiti L, Zambelli A. Radio-frequency thermal ablation of liver tumors. Eur Radiol 2005;15:884–894. Chiou S, Liu J, Needleman L. Current status of sonographically guided radio-frequency ablation techniques. J Ultrasound Med 2007;26: 487–499. Daniels MJ, Varghese T, Madsen EL, Zagzebski JA. Noninvasive ultrasound-based temperature imaging for monitoring radiofrequency heating–Phantom results. Phys Med Biol 2007;52: 4827–4843.
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