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Attenuation and De-focusing During High-Intensity Focused Ultrasound Therapy Through Peri-nephric Fat
Ultrasound in Med. & Biol., Vol. 39, No. 10, pp. 1785–1793, 2013 Copyright Ó 2013 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/$ - see front matter
http://dx.doi.org/10.1016/j.ultrasmedbio.2013.04.010
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Original Contribution ATTENUATION AND DE-FOCUSING DURING HIGH-INTENSITY FOCUSED ULTRASOUND THERAPY THROUGH PERI-NEPHRIC FAT ROBERT RITCHIE,*yz JAMIE COLLIN,yz CONSTANTIN COUSSIOS,yz and TOM LESLIE*y * Nuffield Department of Surgical Sciences, Oxford University Hospitals, Oxford, UK; y Oxford Clinical HIFU Unit, Churchill Hospital, Oxford, UK; and z Department of Biomedical Engineering, University of Oxford, Oxford, UK (Received 20 March 2012; revised 1 April 2013; in final form 11 April 2013)
Abstract—High-intensity focused ultrasound (HIFU) is an attractive therapy for kidney cancer, but its efficacy can be limited by heat deposition in the pre-focal tissues, notably in fat around the kidney (peri-nephric fat), the acoustic properties of which have not been well characterized. Measurements of attenuation were made using a modified insertion-loss technique on fresh, unfixed peri-nephric fat obtained from patients undergoing kidney surgery for cancer. The de-focusing effect of changing the position of the fat layers was also investigated using fresh subcutaneous fat from euthanized pigs. The mean attenuation of human peri-nephric fat was found to be 11.9 ± 0.9 Np/m (n 5 10) at 0.8 MHz, the frequency typically used for HIFU ablation of kidney tumors, with a frequency dependence of f1.2. A typical 2- to 4-cm thickness of peri-nephric fat would result in a de-rated intensity of 3%–62% at 0.8 MHz compared with a hypothetical patient with no peri-nephric fat. Through the use of freshly excised porcine subcutaneous fat, the presence of fat 100 mm in front of the focus was found to have a defocusing effect of approximately 1 mm in both transverse directions, which corresponds to a full HIFU beam width off-target. Peri-nephric fat may significantly affect both the intensity and accuracy of HIFU fields used for the ablation of kidney cancer. (E-mail: [email protected] or [email protected]) Ó 2013 World Federation for Ultrasound in Medicine & Biology. Key Words: High-intensity focused ultrasound, Peri-nephric fat, Attenuation, De-focusing, Thermal ablation, Renal cancer, Kidney.
10-y cancer-specific survival rates greater than 90% (Van Poppel et al. 2011). However, kidney cancer surgery can be risky and is associated with complications in more than 20% of cases (Porpiglia et al. 2008). This rate rises even further in patients aged over 80 (Liguori et al. 2007). A number of energy ablative therapies have been used as minimally invasive treatments for kidney cancer, including cryotherapy (Gill et al. 2000), radiofrequency ablation (Yohannes et al. 2001), microwave ablation (Rehman et al. 2004) and high-intensity focused ultrasound (Illing et al. 2006). These modalities are associated with shorter hospital stays, lower transfusion rates, less pain, lower complication rates and quicker recovery. Currently oncologic outcomes fail to match those after surgery (Kunkle and Uzzo 2008). High-intensity focused ultrasound (HIFU) is an attractive kidney cancer therapy as it is entirely noninvasive, requiring no skin or tumor puncture. It is therefore not associated with a risk of bleeding or tumor spillage. HIFU causes minimal side effects and is generally acceptable to both patient and clinician. However, recent studies have reported that oncologic outcomes
INTRODUCTION Kidney cancer is the fifth and seventh most common cancer in the United Kingdom and United States, respectively (Cancer Research UK 2010). Its incidence is rising according to UK cancer statistics (Cancer Research UK 2010; Centers for Disease Control and Prevention 2012). This is due to increased detection resulting from the greater use of abdominal imaging such as ultrasound (US), X-ray computed tomography (CT) and magnetic resonance imaging (MRI) and to a real increase in the number of new cases. Increasingly, tumors are small (,4 cm, clinical stage T1a) when diagnosed. The vast majority of small tumors are locally confined at diagnosis and therefore curable (Bosniak et al. 1995). Typically, kidney cancer is treated surgically, with excision of the tumor from the kidney or complete kidney removal. Cure rates with surgery are high, with 5- and
Address correspondence to: Robert Ritchie, HIFU Unit, Churchill Hospital, Oxford OX3 7 LJ, UK. E-mail: [email protected] or [email protected] 1785
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are variable (Ritchie et al. 2010), especially compared with the much greater success rates of HIFU ablation in the liver (Illing et al. 2005). Delivery of HIFU energy to either the liver or the kidney is limited by transmission losses through the ribs, as well as attenuation through subcutaneous fat and the abdominal wall. However, a key difference between the two organs is that the kidney is surrounded by a dense layer of fat (peri-nephric fat), which serves to protect it from injury. As seen in Figure 1, it is typically 2–4 cm thick, although it can be significantly thicker (Morris et al. 2010; Ritchie 2012). Evidence from both extracorporeal and laparoscopic HIFU device trials in the kidney suggest that significant energy is deposited in this fat layer, limiting energy delivery and, thus, temperature rises at the focus (Ritchie et al. 2010, 2011). Additionally, biological tissue contains numerous inhomogeneities. Local areas of scarring, calcifications, blood vessels and lymph glands may result in variations in both the speed of sound and thickness of tissue. These local changes can result in phase aberrations, which may distort both the shape and the location of the focal region of a HIFU transducer (Fink et al. 2003). Both attenuation and aberration caused by peri-nephric fat are thought to be major contributors to the sub-optimal oncologic outcomes, and the objective of the present work is to verify this hypothesis through characterization of clinically relevant fat samples. The acoustic properties of both human and animal fat tissue have been investigated previously. Estimates of attenuation at a therapeutically relevant frequency of 1 MHz vary from 7 6 2 Np/m for fresh human fat to 9 Np/m for formalin-fixed human fat (Chivers and Parry 1978; Duck 1990; Goss et al. 1978). However, in many of these studies, the origin of the fat tissue is not recorded or the tissue was melted or placed in formalin before its attenuation was measured. No measurements of the acoustic properties of fresh, human, peri-nephric fat have previously been published, nor has aberration caused
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by fat layers been investigated in the context of HIFU therapy. The aim of this study was twofold: (i) to measure the attenuation coefficient and speed of sound through perinephric fat retrieved from patients undergoing kidney cancer surgery, and (ii) to quantify the extent of the spatial shift in focal position caused by the presence of a typical thickness of fat tissue at various positions in a HIFU field. De-focusing studies were conducted on freshly excised thoracic subcutaneous porcine fat because of the limited quantities of human tissue available for analysis, but only after verifying that this represented a model that closely mimics the relevant properties of human peri-nephric fat. It should also be noted that the aberration study was performed using a HIFU transducer in current clinical use (Ritchie et al. 2010; Ter Haar and Coussios 2007). METHODS Attenuation measurements Human peri-nephric fat. Fresh human peri-nephric fat was obtained from patients who voluntarily agreed to participate in a clinical trial. The trial was approved by the Oxford Research Ethics Committee (REC No. 10/H0604/34). Patients who were scheduled to undergo renal cancer surgery (partial or radical nephrectomy) were eligible for inclusion. Written informed consent was obtained from all patients. Participants were asked to donate, for use in experiments, a sample of their peri-nephric fat tissue, which was discarded thereafter. The samples obtained were typically 10–20 mm thick and up to 80 mm in length. Ten patients were recruited into the study. All had a pre-operative diagnosis of renal cancer on crosssectional imaging and/or histology. The surgical procedure was performed in a standard fashion and was not altered for the purpose of this trial. After excision, the fat was immediately placed into a saline-filled container
Fig. 1. Magnetic resonance images for two patients who underwent renal high-intensity focused ultrasound therapy. There is a significantly thicker layer of peri-nephric fat (arrowheads) in the right image compared with that on the left. The thickness of this layer has been observed to be typically between 2 and 4 cm (Ritchie 2012).
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at 37 C and transported to the laboratory on the hospital site for experimentation. All measurements were made within 1 h of excision, and the temperature was maintained at 37 C throughout the transport and experimentation period, ensuring that the sample had reached thermal equilibrium by the time the measurement was performed. The excised tissue was transferred from the container to a water bath (47 3 26 3 26 cm; 27-L capacity) containing isotonic phosphate-buffered saline (PBS) at 37 C. The PBS solution was made using 192 g PBS powder (GIBCO DPBS powder, Life Technologies, Paisley, UK) dissolved in 20 L of de-gassed, de-ionized water obtained fresh from a clinical therapy device purifying system (Model JC Tumor System, HAIFU, Chongqing, China). An immersion heater (GD 100, Grant Instruments, Shepreth, UK) was used to maintain a constant temperature (37 C) of the water bath. After the tissue was placed in the water bath, it was continuously held under the surface of the water to prevent any contact of the tissue with air and hence avoid bubble formation. Apparatus and procedure for attenuation measurements. As sound passes through an attenuating medium, the amplitude of the signal decays exponentially. Measurements of this decay are complicated by two factors—diffraction of the ultrasound beam and transmission losses on passing between media with different acoustic impedances—and these should be eliminated as far as possible. The measurement system used is illustrated in Figure 2. It consisted of two unfocused 2.25-MHz immersion transducers of the same type (V306, Olympus NDT, Rotherham, UK). Each transducer had a 12.5-mm aperture and –6-dB bandwidth of approximately 100%; The transducers were inserted into a custom-built holder comprising two blocks mounted on horizontal sliders. This allowed both transducers to be moved along their axes while maintaining their transverse alignment. A pulser-receiver (DPR 300, JSR Ultrasonics, Pittsford, NY, USA) was used to drive the transmitting transducer with a single short pulse, and the output was received by an 8-bit digital oscilloscope (Waverunner 44 Xi, LeCroy, Chestnut Ridge, NY, USA). The oscilloscope was controlled by a personal computer (PC) running LabView (Version 11, National Instruments, Newbury, UK). A sampling rate of 100 MS/s was used, and 32 averages were performed in the time domain before storage for post-processing. Needle hydrophone measurements (HNA, Onda, Sunnyvale, CA, USA) revealed the free field pressure to be approximately 160 kPa at a distance of 2 cm from the transmitting transducer. The excised peri-nephric fat was relatively malleable and therefore could be manipulated into different positions to allow propagation through a range of path lengths
Fig. 2. Top: Experimental apparatus for measurement of perinephric fat attenuation; source (TxS) and receiver TxR) transducers (V306, Olympus NDT, Rotherham, UK) are mounted on movable blocks to contact with and fix fat sample at varying path lengths, ensuring planar surfaces with constant transducer loading. Bottom: Schematic of the acquisition system.
to be observed within each sample. The sample was positioned between the transducers, ensuring that the surface of each was in contact with the fat across the full width of their faces and that the sample was significantly wider than the diameter of the transducers; consistent loading of the transducers without propagation through different layers could thus be achieved. After successful acquisition, the mounts were kept in place and the sample was removed from between the transducers using surgical forceps, ensuring that the positions of the transducers were not altered in the process. Propagation was then through the PBS alone, and the received signal was recorded. A typical acquired signal is provided in Figure 3, indicating the cross-correlation time delay determination and the usable bandwidth of the system from 0.5 to 4 MHz. The fat sample was then manipulated into an alternative shape while underwater and re-positioned between the transducers such that the path length was different. PBS-filled voids appeared within the sample as a distortion of the waveform. This was due to the difference in speed of sound between the fat and the water. Manipulation of the sample was repeated to remove these from the acoustic path. This was easily performed because of the size and consistency of the fat sample. The above process was repeated to acquire 10 measurements of propagation
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Fig. 3. Example traces acquired during the attenuation measurements. Top: Time traces from a reference measurement through phosphate-buffered saline (PBS) and an example measurement with propagation through peri-nephric fat. The vertical lines are assigned time delays by cross-correlation between the reference and sample trace. Bottom: The digital Fourier transform of the time trace through fat, illustrating the bandwidth of the system and justifying the choice of frequency range from 0.5 to 4 MHz.
loading of the transducer. The path lengths over which attenuation measurements were taken ranged from 18 to 60 mm. After excluding measurements more than three standard deviations from the mean, the speed of sound in peri-nephric fat at 37 C was found to be 1461.3 6 3.6 m/s, with a comparatively large standard deviation of 27 m/s. This measured value of the speed of sound in peri-nephric fat was then used to determine the effect of the separation of the transducers on the sensitivity of the system. By cooling de-ionized water to 13.5 6 0.1 C, the speed of sound was matched to that in the fat samples to 1460.6 6 0.4 m/s (Marczak 1997). Given the small error in the speed of sound in the diffraction measurement, the variation in the speed of sound in the peri-nephric fat dominates the error in the diffraction profile. The frequency-dependent diffraction profile of the attenuation setup was measured by displacing the mounting blocks along the axis of the transducers. This profile was used to correct the amplitude-propagation length relationship measured in each fat sample so as to yield values for attenuation. This measurement also allows us to adjust for any variation in alignment with the displacement of the transducers along their axes.
through varying path lengths in each sample. This was performed on different days for samples from 10 patients. After completion of data acquisition for each sample, the tissue was immediately placed into clinical waste for incineration to comply with the trial protocol and ethical approval. No human tissue was retained after completion of the attenuation measurements. In a separate measurement, the speed of sound in PBS at 37 C was measured by displacing a hydrophone (HNA, Onda) along the axis of one of the unfocused transducers by known intervals using a micrometer stage (Newport Corporation, Irvine, CA, USA) and finding the slope of the delays measured using the same acquisition system described earlier, giving a value of 1538.9 6 0.3 m/s, where the error is quoted as the standard error of the regression found from the residuals of a linear fit to the data. It should be noted here that the error in this measurement is only at one temperature. During the experiments, the temperature was observed to vary between 36 C and 37 C. Over this range, the speed of sound in pure water is known to vary by approximately 2 m/s (Marczak 1997), and the error in the speed of sound in the PBS reference medium is expected to reflect this scale of fluctuation. With use of this reference speed of sound, the thickness of the fat and the speed of sound in the sample were determined for each measurement, and the amplitudepropagation length relationship was found without any propagation through different layers or with different
Post-processing of attenuation data Post-processing was performed using MATLAB (Version R2009 b, Mathworks, Natick, MA, USA). Each set of data from one sample of fat was considered separately. The maximum in the cross-correlation between a reference measurement and each subsequent measurement was used to find the arrival time. A comparison between the paired measurements in PBS and the fat samples was used to determine both the speed of sound and the thickness of the sample. Discrete Fourier transforms (DFTs) of the timedomain waveforms were then taken. The amplitude of the signal was de-rated by the frequency-dependent diffraction loss measured in water for the relevant path length. An attenuation function of the form ln[V(f)] 5 ln[V0(f)] – a(f)x was then fitted to the measurements in each sample of fat, where V is the amplitude of the DFT at frequency f, a is the frequency-dependent attenuation coefficient, and x is the thickness of the sample. Averaged values for attenuation were then obtained by taking the mean of the attenuation values at each frequency across all samples. Finally, an average attenuation function of the form a 5 Afb was obtained by linear regression. All attenuation results are presented as mean values, with error bars representing the standard error of this quantity, including all errors propagated from contributing measurements. This was chosen as the most appropriate statistical analysis because it reflects the extent to which the sample mean accurately reflects the
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true population mean. Furthermore, in treatment planning of HIFU, mean values rather than patient-specific values tend to be used. Aberration measurements Ultrasound beam aberration by layers of porcine subcutaneous fat. Aberration experiments required a sufficiently large surface area of fat to cover the majority of the HIFU beam cross section at various positions along the HIFU axis. It was clinically not possible to obtain sufficient quantities of fresh human fat to conduct these measurements. Therefore, we used porcine fat, which has a similar macro- and microscopic structure (Hausman and Kauffman 1986). To maintain maximum clinical applicability, porcine fat was obtained from freshly (within 1 h) slaughtered pigs used for research. Sections of skin and subcutaneous fat down to muscle, approximately 30–40 cm 3 30–40 cm, were excised from the back and belly. To ensure maintenance of osmotic equilibrium across adipose cells during transport, the tissue was then immediately covered with 0.9% saline, wrapped in saline-soaked swabs, transferred to self-sealing tissue bags and transported to the laboratory. Despite their size (60–80 kg), pigs are generally lean animals; the sections of excised fat were approximately 3–5 mm thick. Identical experiments were conducted on tissue from two separate pigs on separate days. To justify the use of porcine fat as a surrogate for human fat, both the speed of sound and attenuation coefficient of porcine fat were measured in two additional samples, one taken from each animal, using the technique described above. These measurements were compared with literature values for porcine and human fat. Apparatus and procedure for aberration measurements. Two sections of tissue were cut from the fat from each animal, each approximately 15 3 15 cm to ensure that entire cone of the HIFU beam from the transducer passed through the sample. Each sample was
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then positioned in a custom-built holder, which allowed the tissue to be held normal to the transducer axis without the holder interfering with the ultrasound propagation as seen in Figure 4a. They are held normal so as to introduce only the phase aberration caused by variations in thickness of the fat layer and variations in speed of sound in the fat layer. It is expected that presenting the layer at an angle would introduce further phase aberration because of the differences in path length from the transducer to the focus across the surface of the transducer. Each holder was then attached to a positioning system in a large water tank, which allowed each holder to be moved relative to the other along the axis of the transducer. On clinical HIFU devices, a significant proportion of the propagation path length is through freshly filtered and de-gassed water. The water within the tank was therefore freshly filtered and de-gassed, as well as heated using an immersion heater to 37 C; the tissue was given sufficient time in the tank to reach thermal equilibrium. It should be noted that the approximate speed of sound of water at 37 C of 1520 m/s is comparable to if somewhat lower than, the speed of sound in human muscle that may form a gap in a patient of 1510–1580 m/s (Goss et al. 1978). Because the speed of sound in the water is closer to the speed of sound in the fat samples, the phase aberration effects observed will be an underestimate. To maintain clinical applicability, the transducer from a clinical HIFU device (20099–44, Model JC-200 Tumor System, HAIFU) for the extra-corporeal treatment of abdominal tumors was used. This transducer has a driving frequency of 0.95 MHz, a focal length of 145 mm and an aperture of 200 mm, resulting in a focal spot with a transverse –3-dB beamwidth of 1.2 mm and an axial –3-dB length of 9.5 mm. The transducer was positioned on the floor of the tank, with the propagating surface facing vertically upward toward the surface of the water. A 0.5- mm-diameter needle hydrophone (HPM05/3, Precision Acoustics, Dorchester, UK) was positioned at
Fig. 4. (a) Diagram of experimental setup with HIFU transducer facing upward within the water tank; the needle hydrophone is attached to a positioning system to enable exact location at the transducer focus. (b) Schematic of the acquisition system.
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the focus of the transducer by scanning to find the beam maximum using a positioning system controlled by computer software (UMS 3, Precision Acoustics) on a desktop PC (see Fig. 4b). Twenty-five cycle bursts of ultrasound at 10 Vpeak-to-peak were transmitted by the transducer via a signal generator (33250 A, Agilent Technologies, Santa Clara, CA, USA). The hydrophone signal received 100 MS/s on an 8-bit digital oscilloscope (Waverunner 44 Xi) with a synchronized trigger from the signal generator and acquired on the desktop PC. A low-output pressure amplitude of 40 kPa (corresponding to a spatial peak time-averaged intensity ISPTA of 0.05 W/cm2) was used to avoid the introduction of any nonlinear effects. The 25-cycle burst length corresponds to a spatial pulse length of approximately 40 mm, thus ensuring that standing waves are avoided from reflectors more than 20 mm behind the hydrophone tip; this means that reflections from the hydrophone mount do not interfere with the measurement. The pulse length of 25 cycles also ensures that the transducer output reaches steady state. During an axial scan, data were acquired at 51 points separated by 0.5 mm centered over the focus, that is, 12.5 mm pre- and post-focal. During a planar scan, data were acquired at 31 3 31 points separated by 0.2 mm centered over the focus, that is, a plane extending 3 mm either side of the focus. Axial and planar scans were undertaken as follows: (i) reference scan, with no fat and propagation through pure water alone; (ii) fat sample A placed 15 mm in front of the focus (cf. peri-nephric fat); (iii) fat sample A placed 100 mm in front of the focus (cf. subcutaneous fat); (iv) fat sample A 100 mm in front of the focus and fat sample B placed 15 mm in front of the focus. The raw data were transferred to a PC running MATLAB software (Version 2011b). With the use of in-house MATLAB scripts, transverse planar and axial plots were plotted for all fat positions. For ease of comparison, all scans were normalized to the maximum of the reference scan, and difference images produced by scans (ii), (iii) and (iv) were digitally subtracted from the reference scan (i).
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includes propagated errors caused by variations in the reference speed of sound with temperature. The mean attenuation coefficient across the range 0.5–4 MHz is illustrated in Figure 5, and standard errors once again include propagation of the estimated error in speed of sound caused by temperature variations. The mean error in the fitting of the attenuation coefficient to the data within each sample, being the standard error of the regression for each fit, when averaged over all 10 samples, was 6.4% at the resonance frequency of 2.2 MHz, which increased to 16.1% at 0.5 MHz and 14.9% at 4 MHz because of the lower signal levels at frequencies away from the resonance of the system. Also plotted is a power law least-squares fit to the attenuation data. This yields an attenuation function of 15.6 Np/m/MHz1.2 that is within the standard error of the mean attenuation over this range. At clinically relevant frequencies for abdominal HIFU of 0.8 and 1 MHz (Ter Haar and Coussios 2007), this represents attenuation coefficients of 11.9 6 0.9 Np/m and 15.6 6 1.0 Np/m (mean 6 standard error of the mean). Porcine fat aberration The mean speed of sound in porcine fat was found to be 1492 6 15 m/s over 20 measurements. The measured attenuation coefficient at 1 MHz, the frequency at which the aberration experiments were conducted, was found to be 9.6 6 1.0 Np/m. This compares well with reported values for porcine fat and human fat (Duck 1990; Goss et al. 1978), and justifies the use of porcine thoracic subcutaneous fat as a model for determining aberration effects.
RESULTS Peri-nephric fat attenuation After exclusion of results more than three standard deviations from the mean, the measured speed of sound in human peri-nephric fat was 1461.3 6 3.6 m/s, averaged across 84 measurements in 10 different tissue samples. The mean standard deviation within each sample was 27 m/s and includes any error in maintaining the transducer positions when withdrawing the sample, as well as any intrinsic variation in speed of sound. It should be noted that the reported standard error of 3.6 m/s also
Fig. 5. Graph of the mean attenuation coefficient of human peri-nephric fat across a range of frequencies plotted with upper and lower limits of the standard error of that quantity (n 5 10), including errors resulting from time delay uncertainty and uncertainty in the attenuation determined for each sample.
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Fig. 6. Example aberration introduced by insertion of porcine fat layers into the beam path of a clinical high-intensity focused ultrasound transducer (20099–44, 0.95 Mz, focal length 145 mm, HAIFU, Chongqing, China) for various scenarios. Top: Transverse beam profile at the nominal focal length. Bottom: Normalized difference images. (i) Beam with only de-ionized water (ii) A single fat layer 12.5 mm in front of the focus. (iii) A single fat layer 100 mm in front of the focus. (iv) Fat layers at both locations.
The presence of fat layers in all positions had a significant effect on the location of the focus, because of phase aberrations, as well as an expected amplitude attenuation effect. An example set of aberration images are plotted in Figure 6 as both absolute and difference transverse beam profiles. In the transverse plane, a focal shift of 0.5–1 mm was commonly seen, as indicated in Table 1, comparable to the beamwidth of 1.5 mm, where the measurements were taken with a precision of 60.2 mm. The fat layer close to the transducer, approximately 100 mm in front of the focus, resulted in a greater shift in the location of the focus than the layer only 15 mm in front of the focus. In the axial plane, a maximum focal deviation of 1.5 mm was seen, with a measurement precision of 60.5 mm, which, relative to the 9-mm axial length Table 1. Beam distortions introduced by the insertion of fat layers into the acoustic path of a clinical highintensity focused ultrasound transducer*
Sample Name Set 1 No fat Fat near hydrophone Fat near transducer Two fat layers Set 2 No fat Fat near hydrophone Fat near transducer Two fat layers
Axial offset Transverse (mm) offset (mm)
Ratio of maximum intensity to free field intensity
0
0
1
0 0 0.5 0
0.6 0.89 1.08 0
0.66 0.57 0.37 1.00
20.5 1.5 1.5
0 0.45 0.45
0.71 0.68 0.46
* 20099-44, Model JC-200 Tumor System, HAIFU, Chongqing, China. Axial measurements are accurate to 60.5 mm and transverse to within 60.2 mm.
of the transducer focus, is a considerably smaller deviation than that in the transverse plane. DISCUSSION The measured attenuation of 15.6 Np/m/MHz1.2 in fresh human peri-nephric fat in this study is significant. It is greater than the reported attenuation of human fat from other locations (9.3 6 2 Np/m, 1 MHz) (Chivers and Hill 1975) and muscle perpendicular to the fibers (6.9 6 0.7 Np/m/MHz, 1–7 MHz) (Shore et al. 1986) through which ultrasound must pass to reach the kidney during HIFU. A typical thickness of human peri-nephric fat, on the basis of cross-sectional imaging, is 2–4 cm (Ritchie 2012). This can be significantly greater in the obese patient. On the basis of these typical values, the HIFU driving signal intensity (at 0.8 MHz) would be derated to 62%–39% compared with a similar propagation distance through water. The fat thickness may account for less than one-third of the propagation distance, but its attenuation may lead to significant energy deposition. The consequences of this attenuation are significant. Higher transducer outputs are required to achieve adequate heating at the desired focus position. The higher output has the potential to cause greater pre-focal heating risking injury to surrounding vital structures. In addition, abdominal wall swelling resulting from pre-focal heating may increase propagation distances and result in deterioration of the B-mode diagnostic image. This makes HIFU treatment more risky and, ultimately, necessitates a further increase in HIFU output to maintain energy deposition at the focus. The vicious cycle goes against the principle of using the lowest ‘‘dose’’ necessary to
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achieve the desired outcome so as to minimize side effects. A similar problem with image degradation has been encountered during kidney radiofrequency ablation, where the formation of pre-focal steam bubbles led to loss of tumor visualization and sub-optimal results in 5 of 31 kidney masses (Park et al. 2010). The spatial offset between the intended and actual focal positions caused by the presence of fat layers is small in absolute terms (circa 1.5 mm axial, 1 mm transverse). However, given the size of the focus of the typical HIFU transducer (10–12 mm axial, 2–3 mm transverse), it represents a significant displacement in relative terms. Correcting for this shift in focus is likely to necessitate the application of phase aberration correction procedures (Fink et al. 2003), which are possible only on arraybased HIFU devices. HIFU is ideally suited for kidney tumors 20–30 mm in maximum dimension; tumors up to 40 mm are possible but potentially time consuming. Given the size of the tumor treated, together with the close proximity of vital structures, the estimated focus displayed within the device software cannot be assumed to be the true focus. Particular attention must be given to the gray-scale hyper-echo feedback obtaining during treatment to ensure that the zone of ablation is within the target region. In light of the limited effectiveness of gray-scale hyper-echo monitoring, these results strengthen the argument for new, more accurate methods of monitoring ultrasound-guided HIFU, particularly in terms of identifying the location of the HIFU focus non-destructively (Jensen et al. 2012; Yu and Xu 2008) The fat layer close to the transducer resulted in a greater focal aberration effect than that close to the hydrophone. The clinical analogy in this circumstance would suggest that subcutaneous fat within the abdominal wall would result in a greater de-focusing effect than the peri-nephric fat. It is likely that this is due to the size of the incident HIFU beam close to the hydrophone: its wider cone necessarily passes through a greater area of inhomogeneous fat tissue, leading to a more significant phase aberration effect. It should be noted that the effect of the curvature of the fat was not taken into account in the present study and that this could lead to further aberration effects. Furthermore, the fields employed in the present study were of sufficiently low amplitude to be deemed linear, and increasing levels of non-linearity would undoubtedly result in aberration effects greater than those reported here. CONCLUSIONS The attenuation of peri-nephric fat is significant and greatly impedes focal heating in HIFU treatment of kidney cancer. The difference in speed of sound between fat and the surrounding tissue and variations in thickness
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also result in a de-focusing of the HIFU focus. Kidney HIFU may be challenging in patients with thick layers of peri-nephric fat, and pre-treatment imaging should be carefully evaluated in this respect. It should be possible to improve patient selection for kidney HIFU with a quantitative assessment of fat thickness, and this should be the focus of future research. Acknowledgments—The authors acknowledge the Oxford Biomedical Research Centre (Grant A90104), Northwick Park Institute for Medical Research, and Urology Cancer Research & Education (UCARE).
REFERENCES Bosniak M, Birnbaum B, Krinsky G, Waisman J. Small renal parenchymal neoplasms: Further observations on growth. Radiology 1995;197:589–597. Chivers RC, Hill CR. Ultrasonic attenuation in human tissue. Ultrasound Med Biol 1975;2:25–29. Chivers RC, Parry RJ. Ultrasonic velocity and attenuation in mammalian tissues. J Acoust Soc Am 1978;63:940–953. Duck FA. Physical properties of tissue: A comprehensive reference book. London: Academic Press; 1990. Fink M, Montaldo G, Tanter M. Time-reversal acoustics in biomedical engineering. Annu Rev Biomed Eng 2003;5:465–497. Goss SA, Johnston RL, Dunn F. Comprehensive compilation of empirical ultrasonic properties of mammalian tissues. J Acoust Soc Am 1978;64:423–457. Hausman G, Kauffman R. The histology of developing porcine adipose tissue. J Anim Sci 1986;63:642–658. Illing R, Kennedy J, Wu F, ter Haar G, Protheroe A, Friend P, Gleeson F, Cranston D, Phillips R, Middleton M. The safety and feasibility of extracorporeal high-intensity focused ultrasound (HIFU) for the treatment of liver and kidney tumours in a Western population. Br J Cancer 2005;93:890–895. Illing RO, Leslie TA, Kennedy JE, Wu F, Ter Haar GR, Phillips RR, Protheroe A, Cranston DW. High-intensity focused ultrasound for the treatment of small renal tumours: The Oxford experience. J Urol 2006;175:236–237. Jensen CR, Ritchie RW, Gyongy M, Collin JR, Leslie T, Coussios CC. Spatiotemporal monitoring of high-intensity focused ultrasound therapy with passive acoustic mapping. Radiology 2012;262:252–261. Kunkle DA, Uzzo RG. Cryoablation or radiofrequency ablation of the small renal mass: A meta-analysis. Cancer 2008;113:2671–2680. Liguori G, Trombetta C, Pomara G, Amodeo A, Bucci S, Garaffa G, Francesca F, Belgrano E. Major invasive surgery for urologic cancer in octogenarians with comorbid medical conditions. Eur Urol 2007; 51:1600–1604. discussion 1605. Marczak W. Water as a standard in the measurements of speed of sound in liquids. J Acoust Soc Am 1997;102:2776–2779. Morris K, Tuorto S, Gonen M, Schwartz L, DeMatteo R, D’Angelica M, Jarnagin WR, Fong Y. Simple measurement of intra-abdominal fat for abdominal surgery outcome prediction. Arch Surg 2010;145: 1069–1073. Park BK, Kim CK, Choi HY, Lee HM, Jeon SS, Seo SI, Han DH. Limitation for performing ultrasound-guided radiofrequency ablation of small renal masses. Eur J Radiol 2010;75:248–252. Porpiglia F, Volpe A, Billia M, Scarpa RM. Laparoscopic versus open partial nephrectomy: Analysis of the current literature. Eur Urol 2008;53:732–742. discussion 742–743. Rehman J, Landman J, Lee D, Venkatesh R, Bostwick D, Sundaram C, Clayman R. Needle-based ablation of renal parenchyma using microwave, cryoablation, impedance- and temperature-based monopolar and bipolar radiofrequency, and liquid and gel chemoablation: Laboratory studies and review of the literature. J Endourol 2004;18:83–104. Ritchie RW, Leslie T, Phillips R, Wu F, Illing R, Ter Haar G, Protheroe A, Cranston D. Extracorporeal high intensity focused
Attenuation/de-focusing during HIFU through peri-nephric fat d R. RITCHIE et al. ultrasound for renal tumours: A 3-year follow-up. BJU Int 2010;106: 1004–1009. Ritchie RW, Leslie TA, Turner GD, Roberts IS, D’Urso L, Collura D, Demarchi A, Muto G, Sullivan ME. Laparoscopic high-intensity focused ultrasound for renal tumours: A proof of concept study. BJU Int 2011;107:1290–1296. Ritchie RW. Improving clinical outcomes in High Intensity Focused Ultrasound [DPhil]. Bodleian Library, Oxford: University of Oxford; 2012. Available from: http://www.ora.ox.ac.uk/. Shore D, Woods MO, Miles CA. Attenuation of ultrasound in post rigor bovine skeletal muscle. Ultrasonics 1986;24:81–87. Ter Haar G, Coussios C. High intensity focused ultrasound: Physical principles and devices. Int J Hyperthermia 2007;23:89–104.
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Van Poppel H, Da Pozzo L, Albrecht W, Matveev V, Bono A, Borkowski A, Colombel M, Klotz L, Skinner E, Keane T, Marreaud S, Collette S, Sylvester R. A prospective, randomised EORTC intergroup phase 3 study comparing the oncologic outcome of elective nephron-sparing surgery and radical nephrectomy for low-stage renal cell carcinoma. Eur Urol 2011;59:543–552. Yohannes P, Pinto P, Rotariu P, Smith A, Lee B. Retroperitoneoscopic radiofrequency ablation of a solid renal mass. J Endourol 2001;15: 845–849. Yu T, Xu C. Hyperecho as the indicator of tissue necrosis during microbubble-assisted high intensity focused ultrasound: Sensitivity, specificity and predictive value. Ultrasound Med Biol 2008;34: 1343–1347.
http://dx.doi.org/10.1016/j.ultrasmedbio.2013.04.010
d
Original Contribution ATTENUATION AND DE-FOCUSING DURING HIGH-INTENSITY FOCUSED ULTRASOUND THERAPY THROUGH PERI-NEPHRIC FAT ROBERT RITCHIE,*yz JAMIE COLLIN,yz CONSTANTIN COUSSIOS,yz and TOM LESLIE*y * Nuffield Department of Surgical Sciences, Oxford University Hospitals, Oxford, UK; y Oxford Clinical HIFU Unit, Churchill Hospital, Oxford, UK; and z Department of Biomedical Engineering, University of Oxford, Oxford, UK (Received 20 March 2012; revised 1 April 2013; in final form 11 April 2013)
Abstract—High-intensity focused ultrasound (HIFU) is an attractive therapy for kidney cancer, but its efficacy can be limited by heat deposition in the pre-focal tissues, notably in fat around the kidney (peri-nephric fat), the acoustic properties of which have not been well characterized. Measurements of attenuation were made using a modified insertion-loss technique on fresh, unfixed peri-nephric fat obtained from patients undergoing kidney surgery for cancer. The de-focusing effect of changing the position of the fat layers was also investigated using fresh subcutaneous fat from euthanized pigs. The mean attenuation of human peri-nephric fat was found to be 11.9 ± 0.9 Np/m (n 5 10) at 0.8 MHz, the frequency typically used for HIFU ablation of kidney tumors, with a frequency dependence of f1.2. A typical 2- to 4-cm thickness of peri-nephric fat would result in a de-rated intensity of 3%–62% at 0.8 MHz compared with a hypothetical patient with no peri-nephric fat. Through the use of freshly excised porcine subcutaneous fat, the presence of fat 100 mm in front of the focus was found to have a defocusing effect of approximately 1 mm in both transverse directions, which corresponds to a full HIFU beam width off-target. Peri-nephric fat may significantly affect both the intensity and accuracy of HIFU fields used for the ablation of kidney cancer. (E-mail: [email protected] or [email protected]) Ó 2013 World Federation for Ultrasound in Medicine & Biology. Key Words: High-intensity focused ultrasound, Peri-nephric fat, Attenuation, De-focusing, Thermal ablation, Renal cancer, Kidney.
10-y cancer-specific survival rates greater than 90% (Van Poppel et al. 2011). However, kidney cancer surgery can be risky and is associated with complications in more than 20% of cases (Porpiglia et al. 2008). This rate rises even further in patients aged over 80 (Liguori et al. 2007). A number of energy ablative therapies have been used as minimally invasive treatments for kidney cancer, including cryotherapy (Gill et al. 2000), radiofrequency ablation (Yohannes et al. 2001), microwave ablation (Rehman et al. 2004) and high-intensity focused ultrasound (Illing et al. 2006). These modalities are associated with shorter hospital stays, lower transfusion rates, less pain, lower complication rates and quicker recovery. Currently oncologic outcomes fail to match those after surgery (Kunkle and Uzzo 2008). High-intensity focused ultrasound (HIFU) is an attractive kidney cancer therapy as it is entirely noninvasive, requiring no skin or tumor puncture. It is therefore not associated with a risk of bleeding or tumor spillage. HIFU causes minimal side effects and is generally acceptable to both patient and clinician. However, recent studies have reported that oncologic outcomes
INTRODUCTION Kidney cancer is the fifth and seventh most common cancer in the United Kingdom and United States, respectively (Cancer Research UK 2010). Its incidence is rising according to UK cancer statistics (Cancer Research UK 2010; Centers for Disease Control and Prevention 2012). This is due to increased detection resulting from the greater use of abdominal imaging such as ultrasound (US), X-ray computed tomography (CT) and magnetic resonance imaging (MRI) and to a real increase in the number of new cases. Increasingly, tumors are small (,4 cm, clinical stage T1a) when diagnosed. The vast majority of small tumors are locally confined at diagnosis and therefore curable (Bosniak et al. 1995). Typically, kidney cancer is treated surgically, with excision of the tumor from the kidney or complete kidney removal. Cure rates with surgery are high, with 5- and
Address correspondence to: Robert Ritchie, HIFU Unit, Churchill Hospital, Oxford OX3 7 LJ, UK. E-mail: [email protected] or [email protected] 1785
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are variable (Ritchie et al. 2010), especially compared with the much greater success rates of HIFU ablation in the liver (Illing et al. 2005). Delivery of HIFU energy to either the liver or the kidney is limited by transmission losses through the ribs, as well as attenuation through subcutaneous fat and the abdominal wall. However, a key difference between the two organs is that the kidney is surrounded by a dense layer of fat (peri-nephric fat), which serves to protect it from injury. As seen in Figure 1, it is typically 2–4 cm thick, although it can be significantly thicker (Morris et al. 2010; Ritchie 2012). Evidence from both extracorporeal and laparoscopic HIFU device trials in the kidney suggest that significant energy is deposited in this fat layer, limiting energy delivery and, thus, temperature rises at the focus (Ritchie et al. 2010, 2011). Additionally, biological tissue contains numerous inhomogeneities. Local areas of scarring, calcifications, blood vessels and lymph glands may result in variations in both the speed of sound and thickness of tissue. These local changes can result in phase aberrations, which may distort both the shape and the location of the focal region of a HIFU transducer (Fink et al. 2003). Both attenuation and aberration caused by peri-nephric fat are thought to be major contributors to the sub-optimal oncologic outcomes, and the objective of the present work is to verify this hypothesis through characterization of clinically relevant fat samples. The acoustic properties of both human and animal fat tissue have been investigated previously. Estimates of attenuation at a therapeutically relevant frequency of 1 MHz vary from 7 6 2 Np/m for fresh human fat to 9 Np/m for formalin-fixed human fat (Chivers and Parry 1978; Duck 1990; Goss et al. 1978). However, in many of these studies, the origin of the fat tissue is not recorded or the tissue was melted or placed in formalin before its attenuation was measured. No measurements of the acoustic properties of fresh, human, peri-nephric fat have previously been published, nor has aberration caused
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by fat layers been investigated in the context of HIFU therapy. The aim of this study was twofold: (i) to measure the attenuation coefficient and speed of sound through perinephric fat retrieved from patients undergoing kidney cancer surgery, and (ii) to quantify the extent of the spatial shift in focal position caused by the presence of a typical thickness of fat tissue at various positions in a HIFU field. De-focusing studies were conducted on freshly excised thoracic subcutaneous porcine fat because of the limited quantities of human tissue available for analysis, but only after verifying that this represented a model that closely mimics the relevant properties of human peri-nephric fat. It should also be noted that the aberration study was performed using a HIFU transducer in current clinical use (Ritchie et al. 2010; Ter Haar and Coussios 2007). METHODS Attenuation measurements Human peri-nephric fat. Fresh human peri-nephric fat was obtained from patients who voluntarily agreed to participate in a clinical trial. The trial was approved by the Oxford Research Ethics Committee (REC No. 10/H0604/34). Patients who were scheduled to undergo renal cancer surgery (partial or radical nephrectomy) were eligible for inclusion. Written informed consent was obtained from all patients. Participants were asked to donate, for use in experiments, a sample of their peri-nephric fat tissue, which was discarded thereafter. The samples obtained were typically 10–20 mm thick and up to 80 mm in length. Ten patients were recruited into the study. All had a pre-operative diagnosis of renal cancer on crosssectional imaging and/or histology. The surgical procedure was performed in a standard fashion and was not altered for the purpose of this trial. After excision, the fat was immediately placed into a saline-filled container
Fig. 1. Magnetic resonance images for two patients who underwent renal high-intensity focused ultrasound therapy. There is a significantly thicker layer of peri-nephric fat (arrowheads) in the right image compared with that on the left. The thickness of this layer has been observed to be typically between 2 and 4 cm (Ritchie 2012).
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at 37 C and transported to the laboratory on the hospital site for experimentation. All measurements were made within 1 h of excision, and the temperature was maintained at 37 C throughout the transport and experimentation period, ensuring that the sample had reached thermal equilibrium by the time the measurement was performed. The excised tissue was transferred from the container to a water bath (47 3 26 3 26 cm; 27-L capacity) containing isotonic phosphate-buffered saline (PBS) at 37 C. The PBS solution was made using 192 g PBS powder (GIBCO DPBS powder, Life Technologies, Paisley, UK) dissolved in 20 L of de-gassed, de-ionized water obtained fresh from a clinical therapy device purifying system (Model JC Tumor System, HAIFU, Chongqing, China). An immersion heater (GD 100, Grant Instruments, Shepreth, UK) was used to maintain a constant temperature (37 C) of the water bath. After the tissue was placed in the water bath, it was continuously held under the surface of the water to prevent any contact of the tissue with air and hence avoid bubble formation. Apparatus and procedure for attenuation measurements. As sound passes through an attenuating medium, the amplitude of the signal decays exponentially. Measurements of this decay are complicated by two factors—diffraction of the ultrasound beam and transmission losses on passing between media with different acoustic impedances—and these should be eliminated as far as possible. The measurement system used is illustrated in Figure 2. It consisted of two unfocused 2.25-MHz immersion transducers of the same type (V306, Olympus NDT, Rotherham, UK). Each transducer had a 12.5-mm aperture and –6-dB bandwidth of approximately 100%; The transducers were inserted into a custom-built holder comprising two blocks mounted on horizontal sliders. This allowed both transducers to be moved along their axes while maintaining their transverse alignment. A pulser-receiver (DPR 300, JSR Ultrasonics, Pittsford, NY, USA) was used to drive the transmitting transducer with a single short pulse, and the output was received by an 8-bit digital oscilloscope (Waverunner 44 Xi, LeCroy, Chestnut Ridge, NY, USA). The oscilloscope was controlled by a personal computer (PC) running LabView (Version 11, National Instruments, Newbury, UK). A sampling rate of 100 MS/s was used, and 32 averages were performed in the time domain before storage for post-processing. Needle hydrophone measurements (HNA, Onda, Sunnyvale, CA, USA) revealed the free field pressure to be approximately 160 kPa at a distance of 2 cm from the transmitting transducer. The excised peri-nephric fat was relatively malleable and therefore could be manipulated into different positions to allow propagation through a range of path lengths
Fig. 2. Top: Experimental apparatus for measurement of perinephric fat attenuation; source (TxS) and receiver TxR) transducers (V306, Olympus NDT, Rotherham, UK) are mounted on movable blocks to contact with and fix fat sample at varying path lengths, ensuring planar surfaces with constant transducer loading. Bottom: Schematic of the acquisition system.
to be observed within each sample. The sample was positioned between the transducers, ensuring that the surface of each was in contact with the fat across the full width of their faces and that the sample was significantly wider than the diameter of the transducers; consistent loading of the transducers without propagation through different layers could thus be achieved. After successful acquisition, the mounts were kept in place and the sample was removed from between the transducers using surgical forceps, ensuring that the positions of the transducers were not altered in the process. Propagation was then through the PBS alone, and the received signal was recorded. A typical acquired signal is provided in Figure 3, indicating the cross-correlation time delay determination and the usable bandwidth of the system from 0.5 to 4 MHz. The fat sample was then manipulated into an alternative shape while underwater and re-positioned between the transducers such that the path length was different. PBS-filled voids appeared within the sample as a distortion of the waveform. This was due to the difference in speed of sound between the fat and the water. Manipulation of the sample was repeated to remove these from the acoustic path. This was easily performed because of the size and consistency of the fat sample. The above process was repeated to acquire 10 measurements of propagation
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Fig. 3. Example traces acquired during the attenuation measurements. Top: Time traces from a reference measurement through phosphate-buffered saline (PBS) and an example measurement with propagation through peri-nephric fat. The vertical lines are assigned time delays by cross-correlation between the reference and sample trace. Bottom: The digital Fourier transform of the time trace through fat, illustrating the bandwidth of the system and justifying the choice of frequency range from 0.5 to 4 MHz.
loading of the transducer. The path lengths over which attenuation measurements were taken ranged from 18 to 60 mm. After excluding measurements more than three standard deviations from the mean, the speed of sound in peri-nephric fat at 37 C was found to be 1461.3 6 3.6 m/s, with a comparatively large standard deviation of 27 m/s. This measured value of the speed of sound in peri-nephric fat was then used to determine the effect of the separation of the transducers on the sensitivity of the system. By cooling de-ionized water to 13.5 6 0.1 C, the speed of sound was matched to that in the fat samples to 1460.6 6 0.4 m/s (Marczak 1997). Given the small error in the speed of sound in the diffraction measurement, the variation in the speed of sound in the peri-nephric fat dominates the error in the diffraction profile. The frequency-dependent diffraction profile of the attenuation setup was measured by displacing the mounting blocks along the axis of the transducers. This profile was used to correct the amplitude-propagation length relationship measured in each fat sample so as to yield values for attenuation. This measurement also allows us to adjust for any variation in alignment with the displacement of the transducers along their axes.
through varying path lengths in each sample. This was performed on different days for samples from 10 patients. After completion of data acquisition for each sample, the tissue was immediately placed into clinical waste for incineration to comply with the trial protocol and ethical approval. No human tissue was retained after completion of the attenuation measurements. In a separate measurement, the speed of sound in PBS at 37 C was measured by displacing a hydrophone (HNA, Onda) along the axis of one of the unfocused transducers by known intervals using a micrometer stage (Newport Corporation, Irvine, CA, USA) and finding the slope of the delays measured using the same acquisition system described earlier, giving a value of 1538.9 6 0.3 m/s, where the error is quoted as the standard error of the regression found from the residuals of a linear fit to the data. It should be noted here that the error in this measurement is only at one temperature. During the experiments, the temperature was observed to vary between 36 C and 37 C. Over this range, the speed of sound in pure water is known to vary by approximately 2 m/s (Marczak 1997), and the error in the speed of sound in the PBS reference medium is expected to reflect this scale of fluctuation. With use of this reference speed of sound, the thickness of the fat and the speed of sound in the sample were determined for each measurement, and the amplitudepropagation length relationship was found without any propagation through different layers or with different
Post-processing of attenuation data Post-processing was performed using MATLAB (Version R2009 b, Mathworks, Natick, MA, USA). Each set of data from one sample of fat was considered separately. The maximum in the cross-correlation between a reference measurement and each subsequent measurement was used to find the arrival time. A comparison between the paired measurements in PBS and the fat samples was used to determine both the speed of sound and the thickness of the sample. Discrete Fourier transforms (DFTs) of the timedomain waveforms were then taken. The amplitude of the signal was de-rated by the frequency-dependent diffraction loss measured in water for the relevant path length. An attenuation function of the form ln[V(f)] 5 ln[V0(f)] – a(f)x was then fitted to the measurements in each sample of fat, where V is the amplitude of the DFT at frequency f, a is the frequency-dependent attenuation coefficient, and x is the thickness of the sample. Averaged values for attenuation were then obtained by taking the mean of the attenuation values at each frequency across all samples. Finally, an average attenuation function of the form a 5 Afb was obtained by linear regression. All attenuation results are presented as mean values, with error bars representing the standard error of this quantity, including all errors propagated from contributing measurements. This was chosen as the most appropriate statistical analysis because it reflects the extent to which the sample mean accurately reflects the
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true population mean. Furthermore, in treatment planning of HIFU, mean values rather than patient-specific values tend to be used. Aberration measurements Ultrasound beam aberration by layers of porcine subcutaneous fat. Aberration experiments required a sufficiently large surface area of fat to cover the majority of the HIFU beam cross section at various positions along the HIFU axis. It was clinically not possible to obtain sufficient quantities of fresh human fat to conduct these measurements. Therefore, we used porcine fat, which has a similar macro- and microscopic structure (Hausman and Kauffman 1986). To maintain maximum clinical applicability, porcine fat was obtained from freshly (within 1 h) slaughtered pigs used for research. Sections of skin and subcutaneous fat down to muscle, approximately 30–40 cm 3 30–40 cm, were excised from the back and belly. To ensure maintenance of osmotic equilibrium across adipose cells during transport, the tissue was then immediately covered with 0.9% saline, wrapped in saline-soaked swabs, transferred to self-sealing tissue bags and transported to the laboratory. Despite their size (60–80 kg), pigs are generally lean animals; the sections of excised fat were approximately 3–5 mm thick. Identical experiments were conducted on tissue from two separate pigs on separate days. To justify the use of porcine fat as a surrogate for human fat, both the speed of sound and attenuation coefficient of porcine fat were measured in two additional samples, one taken from each animal, using the technique described above. These measurements were compared with literature values for porcine and human fat. Apparatus and procedure for aberration measurements. Two sections of tissue were cut from the fat from each animal, each approximately 15 3 15 cm to ensure that entire cone of the HIFU beam from the transducer passed through the sample. Each sample was
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then positioned in a custom-built holder, which allowed the tissue to be held normal to the transducer axis without the holder interfering with the ultrasound propagation as seen in Figure 4a. They are held normal so as to introduce only the phase aberration caused by variations in thickness of the fat layer and variations in speed of sound in the fat layer. It is expected that presenting the layer at an angle would introduce further phase aberration because of the differences in path length from the transducer to the focus across the surface of the transducer. Each holder was then attached to a positioning system in a large water tank, which allowed each holder to be moved relative to the other along the axis of the transducer. On clinical HIFU devices, a significant proportion of the propagation path length is through freshly filtered and de-gassed water. The water within the tank was therefore freshly filtered and de-gassed, as well as heated using an immersion heater to 37 C; the tissue was given sufficient time in the tank to reach thermal equilibrium. It should be noted that the approximate speed of sound of water at 37 C of 1520 m/s is comparable to if somewhat lower than, the speed of sound in human muscle that may form a gap in a patient of 1510–1580 m/s (Goss et al. 1978). Because the speed of sound in the water is closer to the speed of sound in the fat samples, the phase aberration effects observed will be an underestimate. To maintain clinical applicability, the transducer from a clinical HIFU device (20099–44, Model JC-200 Tumor System, HAIFU) for the extra-corporeal treatment of abdominal tumors was used. This transducer has a driving frequency of 0.95 MHz, a focal length of 145 mm and an aperture of 200 mm, resulting in a focal spot with a transverse –3-dB beamwidth of 1.2 mm and an axial –3-dB length of 9.5 mm. The transducer was positioned on the floor of the tank, with the propagating surface facing vertically upward toward the surface of the water. A 0.5- mm-diameter needle hydrophone (HPM05/3, Precision Acoustics, Dorchester, UK) was positioned at
Fig. 4. (a) Diagram of experimental setup with HIFU transducer facing upward within the water tank; the needle hydrophone is attached to a positioning system to enable exact location at the transducer focus. (b) Schematic of the acquisition system.
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the focus of the transducer by scanning to find the beam maximum using a positioning system controlled by computer software (UMS 3, Precision Acoustics) on a desktop PC (see Fig. 4b). Twenty-five cycle bursts of ultrasound at 10 Vpeak-to-peak were transmitted by the transducer via a signal generator (33250 A, Agilent Technologies, Santa Clara, CA, USA). The hydrophone signal received 100 MS/s on an 8-bit digital oscilloscope (Waverunner 44 Xi) with a synchronized trigger from the signal generator and acquired on the desktop PC. A low-output pressure amplitude of 40 kPa (corresponding to a spatial peak time-averaged intensity ISPTA of 0.05 W/cm2) was used to avoid the introduction of any nonlinear effects. The 25-cycle burst length corresponds to a spatial pulse length of approximately 40 mm, thus ensuring that standing waves are avoided from reflectors more than 20 mm behind the hydrophone tip; this means that reflections from the hydrophone mount do not interfere with the measurement. The pulse length of 25 cycles also ensures that the transducer output reaches steady state. During an axial scan, data were acquired at 51 points separated by 0.5 mm centered over the focus, that is, 12.5 mm pre- and post-focal. During a planar scan, data were acquired at 31 3 31 points separated by 0.2 mm centered over the focus, that is, a plane extending 3 mm either side of the focus. Axial and planar scans were undertaken as follows: (i) reference scan, with no fat and propagation through pure water alone; (ii) fat sample A placed 15 mm in front of the focus (cf. peri-nephric fat); (iii) fat sample A placed 100 mm in front of the focus (cf. subcutaneous fat); (iv) fat sample A 100 mm in front of the focus and fat sample B placed 15 mm in front of the focus. The raw data were transferred to a PC running MATLAB software (Version 2011b). With the use of in-house MATLAB scripts, transverse planar and axial plots were plotted for all fat positions. For ease of comparison, all scans were normalized to the maximum of the reference scan, and difference images produced by scans (ii), (iii) and (iv) were digitally subtracted from the reference scan (i).
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includes propagated errors caused by variations in the reference speed of sound with temperature. The mean attenuation coefficient across the range 0.5–4 MHz is illustrated in Figure 5, and standard errors once again include propagation of the estimated error in speed of sound caused by temperature variations. The mean error in the fitting of the attenuation coefficient to the data within each sample, being the standard error of the regression for each fit, when averaged over all 10 samples, was 6.4% at the resonance frequency of 2.2 MHz, which increased to 16.1% at 0.5 MHz and 14.9% at 4 MHz because of the lower signal levels at frequencies away from the resonance of the system. Also plotted is a power law least-squares fit to the attenuation data. This yields an attenuation function of 15.6 Np/m/MHz1.2 that is within the standard error of the mean attenuation over this range. At clinically relevant frequencies for abdominal HIFU of 0.8 and 1 MHz (Ter Haar and Coussios 2007), this represents attenuation coefficients of 11.9 6 0.9 Np/m and 15.6 6 1.0 Np/m (mean 6 standard error of the mean). Porcine fat aberration The mean speed of sound in porcine fat was found to be 1492 6 15 m/s over 20 measurements. The measured attenuation coefficient at 1 MHz, the frequency at which the aberration experiments were conducted, was found to be 9.6 6 1.0 Np/m. This compares well with reported values for porcine fat and human fat (Duck 1990; Goss et al. 1978), and justifies the use of porcine thoracic subcutaneous fat as a model for determining aberration effects.
RESULTS Peri-nephric fat attenuation After exclusion of results more than three standard deviations from the mean, the measured speed of sound in human peri-nephric fat was 1461.3 6 3.6 m/s, averaged across 84 measurements in 10 different tissue samples. The mean standard deviation within each sample was 27 m/s and includes any error in maintaining the transducer positions when withdrawing the sample, as well as any intrinsic variation in speed of sound. It should be noted that the reported standard error of 3.6 m/s also
Fig. 5. Graph of the mean attenuation coefficient of human peri-nephric fat across a range of frequencies plotted with upper and lower limits of the standard error of that quantity (n 5 10), including errors resulting from time delay uncertainty and uncertainty in the attenuation determined for each sample.
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Fig. 6. Example aberration introduced by insertion of porcine fat layers into the beam path of a clinical high-intensity focused ultrasound transducer (20099–44, 0.95 Mz, focal length 145 mm, HAIFU, Chongqing, China) for various scenarios. Top: Transverse beam profile at the nominal focal length. Bottom: Normalized difference images. (i) Beam with only de-ionized water (ii) A single fat layer 12.5 mm in front of the focus. (iii) A single fat layer 100 mm in front of the focus. (iv) Fat layers at both locations.
The presence of fat layers in all positions had a significant effect on the location of the focus, because of phase aberrations, as well as an expected amplitude attenuation effect. An example set of aberration images are plotted in Figure 6 as both absolute and difference transverse beam profiles. In the transverse plane, a focal shift of 0.5–1 mm was commonly seen, as indicated in Table 1, comparable to the beamwidth of 1.5 mm, where the measurements were taken with a precision of 60.2 mm. The fat layer close to the transducer, approximately 100 mm in front of the focus, resulted in a greater shift in the location of the focus than the layer only 15 mm in front of the focus. In the axial plane, a maximum focal deviation of 1.5 mm was seen, with a measurement precision of 60.5 mm, which, relative to the 9-mm axial length Table 1. Beam distortions introduced by the insertion of fat layers into the acoustic path of a clinical highintensity focused ultrasound transducer*
Sample Name Set 1 No fat Fat near hydrophone Fat near transducer Two fat layers Set 2 No fat Fat near hydrophone Fat near transducer Two fat layers
Axial offset Transverse (mm) offset (mm)
Ratio of maximum intensity to free field intensity
0
0
1
0 0 0.5 0
0.6 0.89 1.08 0
0.66 0.57 0.37 1.00
20.5 1.5 1.5
0 0.45 0.45
0.71 0.68 0.46
* 20099-44, Model JC-200 Tumor System, HAIFU, Chongqing, China. Axial measurements are accurate to 60.5 mm and transverse to within 60.2 mm.
of the transducer focus, is a considerably smaller deviation than that in the transverse plane. DISCUSSION The measured attenuation of 15.6 Np/m/MHz1.2 in fresh human peri-nephric fat in this study is significant. It is greater than the reported attenuation of human fat from other locations (9.3 6 2 Np/m, 1 MHz) (Chivers and Hill 1975) and muscle perpendicular to the fibers (6.9 6 0.7 Np/m/MHz, 1–7 MHz) (Shore et al. 1986) through which ultrasound must pass to reach the kidney during HIFU. A typical thickness of human peri-nephric fat, on the basis of cross-sectional imaging, is 2–4 cm (Ritchie 2012). This can be significantly greater in the obese patient. On the basis of these typical values, the HIFU driving signal intensity (at 0.8 MHz) would be derated to 62%–39% compared with a similar propagation distance through water. The fat thickness may account for less than one-third of the propagation distance, but its attenuation may lead to significant energy deposition. The consequences of this attenuation are significant. Higher transducer outputs are required to achieve adequate heating at the desired focus position. The higher output has the potential to cause greater pre-focal heating risking injury to surrounding vital structures. In addition, abdominal wall swelling resulting from pre-focal heating may increase propagation distances and result in deterioration of the B-mode diagnostic image. This makes HIFU treatment more risky and, ultimately, necessitates a further increase in HIFU output to maintain energy deposition at the focus. The vicious cycle goes against the principle of using the lowest ‘‘dose’’ necessary to
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achieve the desired outcome so as to minimize side effects. A similar problem with image degradation has been encountered during kidney radiofrequency ablation, where the formation of pre-focal steam bubbles led to loss of tumor visualization and sub-optimal results in 5 of 31 kidney masses (Park et al. 2010). The spatial offset between the intended and actual focal positions caused by the presence of fat layers is small in absolute terms (circa 1.5 mm axial, 1 mm transverse). However, given the size of the focus of the typical HIFU transducer (10–12 mm axial, 2–3 mm transverse), it represents a significant displacement in relative terms. Correcting for this shift in focus is likely to necessitate the application of phase aberration correction procedures (Fink et al. 2003), which are possible only on arraybased HIFU devices. HIFU is ideally suited for kidney tumors 20–30 mm in maximum dimension; tumors up to 40 mm are possible but potentially time consuming. Given the size of the tumor treated, together with the close proximity of vital structures, the estimated focus displayed within the device software cannot be assumed to be the true focus. Particular attention must be given to the gray-scale hyper-echo feedback obtaining during treatment to ensure that the zone of ablation is within the target region. In light of the limited effectiveness of gray-scale hyper-echo monitoring, these results strengthen the argument for new, more accurate methods of monitoring ultrasound-guided HIFU, particularly in terms of identifying the location of the HIFU focus non-destructively (Jensen et al. 2012; Yu and Xu 2008) The fat layer close to the transducer resulted in a greater focal aberration effect than that close to the hydrophone. The clinical analogy in this circumstance would suggest that subcutaneous fat within the abdominal wall would result in a greater de-focusing effect than the peri-nephric fat. It is likely that this is due to the size of the incident HIFU beam close to the hydrophone: its wider cone necessarily passes through a greater area of inhomogeneous fat tissue, leading to a more significant phase aberration effect. It should be noted that the effect of the curvature of the fat was not taken into account in the present study and that this could lead to further aberration effects. Furthermore, the fields employed in the present study were of sufficiently low amplitude to be deemed linear, and increasing levels of non-linearity would undoubtedly result in aberration effects greater than those reported here. CONCLUSIONS The attenuation of peri-nephric fat is significant and greatly impedes focal heating in HIFU treatment of kidney cancer. The difference in speed of sound between fat and the surrounding tissue and variations in thickness
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also result in a de-focusing of the HIFU focus. Kidney HIFU may be challenging in patients with thick layers of peri-nephric fat, and pre-treatment imaging should be carefully evaluated in this respect. It should be possible to improve patient selection for kidney HIFU with a quantitative assessment of fat thickness, and this should be the focus of future research. Acknowledgments—The authors acknowledge the Oxford Biomedical Research Centre (Grant A90104), Northwick Park Institute for Medical Research, and Urology Cancer Research & Education (UCARE).
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