Calibration of Ultrasound Backscatter Temperature Imaging for High-Intensity Focused Ultrasound Treatment Planning

Ultrasound in Med. & Biol., Vol. 39, No. 9, pp. 1596–1612, 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.2013.04.001

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Original Contribution CALIBRATION OF ULTRASOUND BACKSCATTER TEMPERATURE IMAGING FOR HIGH-INTENSITY FOCUSED ULTRASOUND TREATMENT PLANNING JOHN CIVALE,* IAN RIVENS,* GAIL TER HAAR,* HUGH MORRIS,y CONSTANTIN COUSSIOS,z PETER FRIEND,x and JEFFREY BAMBER* * Joint Department of Physics, Institute of Cancer Research and Royal Marsden NHS Trust, Sutton, Surrey, UK; y Biomedical Engineering Department, Ohio State University, Columbus, Ohio, USA; z Institute of Biomedical Engineering, Department of Engineering Science, Headington, Oxford, UK; and x Nuffield Department of Surgical Sciences, Oxford Transplant Centre, Churchill Hospital, Headington, Oxford, UK (Received 6 November 2012; revised 6 March 2013; in final form 1 April 2013)

Abstract—High-intensity focused ultrasound (HIFU) is rapidly gaining acceptance as a non-invasive method for soft tissue tumor ablation, but improvements in the methods of treatment delivery, planning and monitoring are still required. Backscatter temperature imaging (BTI) uses ultrasound to visualize heating-induced echo strain and may be used to indicate the position of the HIFU focal region using low-power ‘‘sub-lesioning’’ exposure. The technique may also provide a quantitative tool for assessing the efficacy of treatment delivery if apparent strain measurements can be related to the underlying temperature rise. To obtain temperature estimates from strain measurements, the relationship between these variables has to be either measured or otherwise assumed from previous calibrations in similar tissues. This article describes experimental measurements aimed at deriving the relationship between temperature rise and apparent strain in the laboratory environment using both ex vivo bovine liver tissue samples and normothermically perfused porcine livers. A BTI algorithm was applied to radiofrequency ultrasound echo data acquired from a clinical ultrasound scanner (Z.One, Zonare Medical Systems, Mountain View, CA, USA) where the imaging probe was aligned with the focal region of a HIFU transducer. Temperature measurements were obtained using needle thermocouples implanted in the liver tissue. A series of ‘‘non-ablative’’ HIFU exposures giving peak temperatures below 10 C were made in three separate ex vivo bovine livers, yielding an average strain/temperature coefficient of 0.126 ± 0.088 percentage strain per degree Celsius. In the perfused porcine livers at a starting temperature of 38 C (normal body temperature) the strain/temperature coefficients were found to be 0.040 ± 0.029 percentage strain per degree Celsius. The uncertainty in these results is directly linked to the precision of the strain measurement, as well as the naturally occurring variance between different tissue samples, indicating that BTI may lack the accuracy required to be implemented successfully in practice as a quantitative treatment planning technique at a sub-lesioning exposure level. This is because, to be of use in treatment planning, temperature-rise estimates may require an accuracy greater (,10%) than that offered by BTI measurement. BTI may, however, still play a role in ensuring the correct positioning of the focal region and as a treatment monitoring modality capable of detecting an increased rate of heating in tissue after HIFU ablation. (E-mail: [email protected]) Ó 2013 World Federation for Ultrasound in Medicine & Biology. Key Words: High-intensity focused ultrasound, Ultrasound thermometry, Diagnostic ultrasound, Liver, Perfusion.

et al. 2005; Leslie et al. 2012; Ritchie et al. 2012; ter Haar et al. 1998; Vallancien et al. 1992; Visioli et al. 1999; Wu 2006; Wu et al. 2004). An important requirement for maximizing the efficacy of HIFU is a non-invasive method for obtaining information from the target tissue that allows control over the location and extent of the region of HIFUinduced tissue necrosis. Approaches that have been taken to provide such information include attempts either to detect when and where tissue necrosis occurs or to predict the formation of necrosis. The former may be achieved using imaging techniques that are sensitive to tissue

INTRODUCTION High-intensity focused ultrasound (HIFU) has emerged in recent years as a technique for non-invasive treatment of soft tissue tumors. Its feasibility has been reported in clinical trials for the treatment of benign and malignant tumors of organs including the kidney, liver, bone, breast, bladder, uterus and pancreas (Hindley et al. 2004; Illing

Address correspondence to: John Civale, Division of Radiotherapy and Imaging, Institute of Cancer Research, Cotswold Road, Sutton, Surrey, SM2 5 NG, UK. E-mail: [email protected] 1596

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changes during ablation (Liu et al. 2010; Nandlall et al. 2011; Rabkin et al. 2006; Righetti et al. 1999). The latter may be accomplished by using a model for necrosis in combination with either a model of the factors that influence the temperatures achieved with HIFU or real-time monitoring of the tissue temperature during treatment. Although magnetic resonance imaging has been found to have considerable promise in monitoring temperature during HIFU (Furusawa et al. 2006, Hindley et al. 2004), it presents problems in terms of expense and portability and the need for the HIFU delivery and positioning systems to be magnetic resonance compatible. Ultrasound image-based thermometry (Seip et al. 1996), which we refer to as backscatter temperature imaging (BTI), could overcome these disadvantages of MR thermometry and provide improved temporal and spatial information, but is restricted to detecting spatially localized changes in temperature and does not provide an absolute measurement of temperature. Miller et al. (2002, 2004, 2005) confined their interest in BTI to its use for detecting thermal test exposures to position the HIFU focal region in the target volume. Here we extend the work of Miller et al. toward a more quantitative use of BTI. Specifically, our interest is in understanding whether BTI of one or more thermal test exposures could be used to estimate the HIFU parameters required to ablate a region of tissue of desired size and shape in a controlled manner (i.e., in effect to provide a means of compensating for variations in the ultrasonic absorption coefficient and the thermal dissipation characteristics at the target site, as well as ultrasonic attenuation in the intervening tissue). Before this question can be fully investigated, however, it is necessary to consider whether the tissue properties on which BTI depends can be adequately calibrated. Backscatter temperature imaging relies on the temperature dependence of the speed of sound in tissue, which, for non-fatty tissues, increases with temperature between 37 C and 50 C (Bamber and Hill 1979). Thus, the pulse-echo time is shorter for a reflected ultrasound pulse propagating through a heated region of non-fatty tissue than for one propagating through the same tissue at a lower temperature. For an ultrasound system that assumes a fixed value for the speed of sound, the apparent position of any reflecting structure is therefore temperature dependent. The apparent displacements of radiofrequency (RF) ultrasound echoes in frames obtained pre- and post-heating can be determined using crosscorrelation techniques similar to those used by Ramamurthy and Trahey (1991) for measuring blood velocity. Images of apparent axial strain (the gradient of apparent displacement in the direction of imaging axis) can be produced from the axial displacement profiles using a noise-smoothing gradient estimator, such as

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a moving local least-squares linear regression, in a manner similar to that used in imaging tissue elasticity (Kallel et al. 1999; Ophir et al. 1991). For the case of temperature imaging, assuming there is no mechanical source of tissue strain, strain is related to localized changes in the speed of sound and to thermal expansion, and is not due to tissue elasticity. We therefore refer to it as apparent strain, or echo strain, throughout this article. The apparent strain value can be related to the local temperature change if the relationship between speed of sound and temperature is known, a relationship that we refer to as a BTI calibration, and provided thermal expansion effects can be considered negligible, as described below. Preliminary experiments (Seip et al. 1996) indicated that strain estimated in this way changed with temperature, both in a rubber phantom and in excised liver tissue, for rises of a few degrees (,10 C) above room temperature. Maass-Moreno and Damianou (1996) studied the problem using an analytical model and, using published thermal expansion coefficients for muscle and water, suggested that tissue strain caused by thermal expansion (0.01%/ C) is approximately one order of magnitude less than the apparent strain caused by speed of sound changes (0.1%/ C), in the temperature range 20240 C, and hence could be neglected. Similar conclusions were drawn by Sun and Ying (1999). MaassMoreno et al. (1996) reported experimental pulse-echo time-of-flight measurements in HIFU-heated turkey breast muscle tissue that confirmed that for temperature rises of up to 10 C above room temperature, the change in the time of arrival of echoes was well described by a linear fit to the peak temperature rise, with thermal expansion effects dominated by the speed-of-sound temperature dependency. Souchon et al. (2005), using simulations and experimental measurements, reported that thermal expansion could not be ignored in the case of highly elevated (lesioning) temperature rises induced in ex vivo liver samples. In their study, echo-strain measurements exhibited some evidence of tissue expansion which, the authors hypothesized, was due to the effects of coagulative necrosis. Miller et al. (2004) also observed no evidence, in excised liver tissue, of detectable displacements caused by thermal expansion when low-temperature rises were induced. Miller et al. (2002, 2004) also identified fundamental limitations of this technique such that (i) the temperature dependence of the speed of sound varies substantially between tissues with different fat contents; and (ii) given the dependence of the sound speed on temperature and a starting temperature of 37 C (body temperature), a maximum apparent strain value of only 0.5% would be expected for a temperature rise of 50 C (assuming the speed-of-sound maximum in human liver reaches

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a maximum at this temperature [Bamber and Hill 1979]). They found, however, that the technique should be feasible for a baseline temperature of 37 C, as apparent strains lower than 0.2% were measurable using BTI in excised bovine liver. More recently, Liu et al. (2010) implemented a high-frame-rate 2-D real-time temperature-rise imaging ultrasound system and reported temperature-rise estimates with a precision better than 1 C. The same publication also described the ability of BTI to detect changes in the ultrasonic absorption coefficient caused by tissue necrosis, observed as a higher BTI-measured temperature rise for a test exposure after ablation, compared with a similar exposure before ablation. Although Liu et al. intended to use this as a HIFU treatment monitoring technique, that is, to detect necrosis, the demonstration that BTI is sensitive to changes in local absorption coefficient is an important finding for our proposed application of BTI, because absorption is one of the tissue properties required for HIFU treatment planning. However, as with all of the aforementioned studies, temperature-rise estimates were based on an assumed knowledge of the relationship between sound speed and temperature. It therefore appears that there is sufficient evidence to suggest that the ability of BTI to measure the temperature rises generated by sub-lesioning HIFU exposures could be used to improve HIFU treatment planning by providing a relative measure of the combined effects of variables affecting HIFU heat deposition, such as ultrasound attenuation in the propagation path, local ultrasound absorption and thermal diffusivity, all of which influence lesion formation. A major difficulty, however, is that in practice, a calibration of the relationship between sound speed and temperature cannot be obtained for the tissue region to be measured. This raises the question of whether a calibration derived from sample tissues would provide sufficient accuracy when applied to determine the properties of an unknown test tissue, an important aspect of BTI that has not previously been addressed. One aim of this article is therefore to add knowledge to the field of BTI research by determining a calibration of apparent measured strain as a function of temperature rise for low-power (sub-lesioning) HIFU exposures in liver, and examining its variation within, and between, tissue samples. The method described by Miller et al. (2004) has been used to compute apparent strain from RF echo data acquired from excised liver using a clinical ultrasound imaging scanner. The strain has then been compared with thermocouple temperature measurements for a range of HIFU-induced temperature rises. In addition, to gain early insight into the use of excised tissue for deriving such calibrations and into the relationship between ex vivo and in vivo use of BTI, the experiments were conducted for two different tissue configurations.

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The first was ex vivo liver tissue with an imaging geometry that was optimum for BTI; the second was a living perfused whole excised liver with an imaging geometry suitable for in vivo application, which simulated the in vivo use of BTI. The perfused liver measurements therefore provide valuable new data on the use of BTI in a clinically relevant functioning living organ. EXPERIMENTAL STRATEGY, EQUIPMENT AND METHODS Experimental strategy For the majority of ultrasound-guided HIFU applications, the most practical configuration is a co-axial alignment of the therapy and imaging beams. This allows the use of a single treatment head, whether it be a large therapy transducer with an imaging probe built into the center or a combined therapy-imaging array. This maximizes the use of the available acoustic window on the skin for focusing and manipulating the therapy beam. It also means that imaging can be performed along the therapy ultrasound beam axis, enabling the clinician to visualize potential obstacles or structures in the HIFU beam that may need to be spared. Unfortunately, this co-axial alignment also maximizes thermal refraction and aberration of the diagnostic beam and, importantly, strain noise that arises from strong lateral gradients in axial displacement (Bamber et al. 1997; Miller et al. 2004, 2005). The thermal gradients leading to the thermal refraction effect are directly related to the shape and size of the hot spot generated by the HIFU and its orientation to the imaging beam. For extracorporeal HIFU transducers such as those used in these experiments, the length of the focal region is typically one order of magnitude greater than the width. The hot spots generated by these focal regions will start to develop thermal profiles of dimensions similar to those of the focal region. The greatest thermal gradients are therefore obtained in the lateral direction, and again, these would be approximately one order of magnitude greater than in the axial direction. The use of a co-axial alignment of therapy and imaging beams therefore maximizes the effect of these thermal gradients. As a consequence, the noise in strain estimates using this alignment configuration might be expected to increase by approximately one order of magnitude, although other sources of error will also contribute. A comparison of the standard deviation in peak strain estimates as measured by Miller et al. (2005), using different therapy-imaging beam orientations, confirmed an increase of approximately this magnitude using a co-axial alignment compared with a longitudinal transverse alignment. Experiments were therefore performed in two stages. Initially, calibration measurements were

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made on ex vivo non-perfused bovine liver tissue at room temperature. The purpose of these measurements was to show the maximum potential of the BTI technique, that is, with as many as possible of the known sources of error in strain measurements removed and the corresponding errors minimized. This was achieved by (i) using a transverse alignment in which the imaging plane is aligned with the HIFU trans-axial plane and centered on the HIFU focal plane (this was easier to align than arranging for the lateral direction of the imaging plane to be parallel with the HIFU beam axis); (ii) maximizing sensitivity by starting at a base temperature of about 20 C, where the temperature gradient of sound speed (1.7 m/s/ C or 0.1%/ C) is greater than that at normal body temperature (1.0 m/s/ C or 0.6%/ C); and (iii) removing errors resulting from motion by having the tissue constrained by a holder. Then calibration measurements were made in living excised perfused whole porcine livers to mimic the more realistic clinical tissue conditions of normal body temperature, intact cellular structure and tissue composition and normal intracellular, extracellular and vascular fluid contents. In addition, the clinically preferred co-axial alignment of the therapy and imaging transducers was used, and the liver was free to move. In these experiments, the accuracy and precision of BTI strain measurements were expected to be reduced by maintaining the tissue at body temperature, employing the co-axial configuration and allowing the organ to move. This part of the study therefore also allowed us to investigate some of the practical difficulties that are likely to be encountered when employing BTI in the clinic. HIFU transducers Two HIFU transducers were used for the calibrations. The first device was a prototype piezo-composite segmented extracorporeal HIFU transducer (Imasonic, Besanc¸on, France) comprising a spherical bowl divided into 10 parallel linear segments (strips), as described by Civale et al. (2006). Although this probe had the potential to be used as a phased array, in these experiments all segments were driven in phase, at their fundamental resonant frequency of 1.7 MHz. This transducer had a focal length of 15 cm and an outer diameter of 11 cm. It also had a 5-cm-diameter central hole that allowed insertion of an ultrasound imaging probe to permit co-axial imaging for HIFU treatment guidance and monitoring; for the remainder of article, we therefore refer to this device as the spherical annulus HIFU transducer. The full width at half-maximum dimensions of the pressure amplitude distribution at the focal region were measured under free field conditions with a membrane hydrophone having a sensitive area 0.4 mm in diameter (UC1604, Precision Acoustics, Dorchester, UK) to be 1.5 mm and

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2.0 cm in the radial and axial directions respectively. The second transducer was a single-element spherically focused piezoceramic device, developed in-house. It had a focal length of 15 cm and a diameter of 8.6 cm. The device was driven at 1.693 MHz, the third harmonic of its natural resonant frequency. The full widths at halfmaximum of its pressure amplitude distribution at the focal region were 2.1 mm and 2.3 cm in the radial and axial directions, respectively. Because it lacked a central hole for co-axial imaging, we refer to this device as the spherical bowl HIFU transducer. Both HIFU transducers were used for the non-perfused (transversely aligned) tissue experiments. Only the spherical annulus HIFU transducer was used for the perfused (co-axially aligned) liver work. The radial and axial pressure profiles of both transducers in the focal plane are shown in Figure 1. The HIFU transducers were driven using a signal generator (HP 33120A, Agilent Technologies, Wokingham, UK), the output of which was amplified using a broadband power amplifier (A300, ENI, Rochester, NY, USA). Voltages at the transducer in the range 10 to 30 V were used to produce sub-lesioning exposures (Ispta , 500 W/cm2) with peak temperature rises between 2 C and 10 C (see Experimental Procedure). The HIFU transducer position was controlled using a computercontrolled 3-D motorized gantry. Ultrasound imaging A Zonare ultrasound scanner (Z.One, Zonare Medical Systems, Mountain View, CA, USA) was used for all ultrasound backscatter temperature imaging. For the non-perfused tissue calibration experiments, because the transverse orientation was employed, it was possible to mount the imaging probe close to the HIFU focus. A linear (L10–5, Zonare) imaging probe with a centre frequency of 5.9 MHz (4.2–6.5 MHz, 12-dB bandwidth [Civale 2008]) was therefore used as it provided good spatial resolution. In the perfused liver calibrations, a curvilinear (C5–2, Zonare) imaging probe with a center frequency of 2.8 MHz (2.1–3.3 MHz, 12-dB bandwidth) was used as this provided the necessary imaging depth (.18 cm) for imaging the HIFU focal region when inserted into the imaging aperture of the spherical annulus HIFU transducer. In this arrangement, the center of the front face of the probe was positioned approximately 16.5 cm from the HIFU focal region. The lateral and elevational resolutions at this imaging depth were estimated, by measurement of the transmitted acoustic pulse with a 0.4-mm membrane hydrophone (UC1604, Precision Acoustics, Dorchester, UK), to be 7 and 9 mm, respectively. When the curvilinear probe was used, the sweep angle was minimized to reduce the total amount of image

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Fig. 1. Hydrophone pressure scans of the acoustic pressure at the focus of the high-intensity focused ultrasound transducers driven in burst mode at low power. Plots show the pressure profiles in the trans-axial (left) and axial (right) planes for the spherical bowl (a, c) and annulus (b, d) transducers. RMS 5 root mean squared.

data acquired because echoes from regions of tissue well away from the HIFU focal region were of no interest. Frame rates of 5 and 10 Hz were used for the nonperfused and perfused liver calibrations, respectively. Digital in-phase and quadrature (IQ) (2.0-MHz sampling rate) sampled beam-formed echo image data were saved and processed offline to reconstruct RF images before BTI processing described below. Interference, which prevented echo displacement measurement, was observed during the HIFU exposure when imaging with the C5–2 probe. Thus, for the perfused liver experiments, strain images could be obtained only after the HIFU exposure, that is, during the cooling phase. When the L10–5 probe was used, however, such interference was not apparent. BTI was therefore possible during both the heating and cooling times for the non-perfused tissue calibrations. Non-perfused tissue calibrations Figure 2 is a diagram of the apparatus used for the non-perfused tissue calibration experiments. For each experiment, bovine liver tissue was collected from a local abattoir and stored in a cold room overnight. Two blocks of tissue (approximately 8 3 15 3 15 cm) were cut on the day of the experiment. Each sample was degassed by placing it in degassed tap water at reduced pressure (85 kPa vacuum pressure) for at least 30 min. The degassed samples were immediately transferred under degassed tap water into the experimental tank illustrated in Figure 2 and placed on top of each other in a holder, resulting in one large block of liver of approximate

dimensions 16 3 15 3 15 cm. The HIFU beam entered the tissue through the open side of the Perspex holder; an acoustic absorber was placed on the opposite side to prevent reflections. The focal depth was set to be approximately 2 cm. Imaging was performed from above the sample, through an acoustically transparent window. The linear array imaging probe was placed against the top acoustic window to prevent multiple reflections. The attenuation provided by 16 cm of liver was useful in reducing reflections from the bottom of the tank. The imaging probe was aligned in the transaxial plane at the HIFU focus so that the tip of a needle thermocouple inserted into the tissue, as described below, was visible in the ultrasound image, thus

Fig. 2. Diagram of the arrangement of the spherical bowl highintensity focused ultrasound (HIFU) transducer and the L10–5 imaging probe for ex vivo non-perfused backscatter temperature imaging calibrations. The image plane (seen here end-on) is aligned with the HIFU trans-axial plane.

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Fig. 3. Diagram of the perfused liver system. Blood flow is maintained using a centrifugal pump. The heat exchanger maintains normal body temperature (38 C). Blood parameters such as oxygenation and glucose level are maintained close to their physiologic values. The IMED (Alaris IMED Gemini PC-4, Carefusion, San Diego, CA, USA) provides a constant infusion of nutrients (total parenteral nutrition) and prostacyclin. An anticoagulant (heparin) and antibiotics were regularly administered as bolus injections.

ensuring that thermocouple measurements were in the BTI plane. Perfused liver system Measurements were made in isolated perfused porcine livers. Five to eight units of blood was obtained from a first donor animal, whereas the liver from a second pig was surgically excised and placed on a blood perfusion circuit as described by Butler et al. (2002) (Fig. 3). The blood from two donors was required to provide sufficient volume to both fill and run the perfusion system; both the liver-donor and blood-donor pigs were in the range 35–50 kg. The perfusion system was designed to maintain normal liver blood flow and metabolism (Butler et al. 2002). This was achieved by maintaining normal blood flow rates, blood oxygenation and body temperature (38 C) and by infusing nutrients (total parenteral nutrition), prostacyclin (for improved vasodilation) and taurocholic acid (for bile salt replacement) into the blood. Hourly measurements of blood glucose level, gas levels and flow rates were made, and levels of the above controlled variables were adjusted manually as necessary to maintain the correct values as follows. Partial oxygen and carbon dioxide pressures in blood were maintained between 10 and 25 kPa and between 4.0 and 6.0 kPa,

respectively. A blood pH level of 7.35 to 7.45 was maintained by controlling the air flow to vary the removal rate. Arterial pressure was maintained between 65 and 110 mm Hg, and hepatic portal vein pressure was maintained below 11 mm Hg. Blood glucose levels were between 8 and 15 mmol/L. Heparin was injected every 4 h. Blood flow rates were maintained at approximately 0.5 L/min for the hepatic artery and 1.5 L/min for the portal vein (values similar to normal physiologic values in vivo). The perfusion system could be run for at least 24 h. The BTI calibration data presented here were obtained within the first 12 h of the perfusion. The arrangement for imaging and HIFU exposures of the perfused liver is depicted in Figure 4. The liver was placed in a metal bowl with a central drain through which the ascites was collected and re-circulated into the blood perfusion system. Acoustic coupling for the 15-cm-focal-length HIFU transducer and imaging probe was achieved using 10–15 cm of degassed water in a polyethylene bag (40-mm thickness). To minimize pressure on the liver, this bag was not placed in contact with the liver, but was immersed in a 5-cm layer of degassed water that was located above the liver and separated from it by a sterile sheath of 0.1-mm-thick Saran wrap, with the liver resting on a number of sterile perfusate-filled latex gloves. Finally, a layer of acoustically absorbing

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Fig. 4. Arrangement of the perfused liver and ultrasound transducers. The 3-D gantry was used to control positioning of the therapy transducer, which also incorporated the co-axially aligned ultrasound imaging probe (C5–2, Zonare Medical Systems, Mountain View, CA, USA).

and anechoic rubber was placed behind the liver to prevent reflections from the metal bowl. Experimental procedure All calibration experiments consisted of measuring the apparent echo strain as a function of HIFU-induced temperature rise measured with a needle thermocouple placed at the same position as the strain measurement. Preliminary studies (Civale 2008) indicated that the presence of a needle thermocouple during ultrasound echo data collection had significant adverse effects on the quality of strain images. Therefore, for the studies reported here, but with exceptions as described below, the following two-stage procedure was used which relied on the reversibility, and hence repeatability, of change in sound speed with the temperature rises induced. A needle thermocouple (‘K’ type, Model KMTSS-010(G)-(6), Omega Engineering, Manchester, UK) of diameter 0.25 mm was inserted under ultrasound guidance so that its tip was at the elevational center of the image plane at a depth between 1 and 2 cm beneath the surface of the liver, away from any major blood vessels. The imaging probe was then clamped, and the HIFU focus was localized on the thermocouple (within 0.1 mm radially and 1 mm axially) by measuring the spatial peak temperature rise obtained from multiple low intensity (100 W/cm2 Isp) exposures. During and after each exposure within a first set of HIFU exposures, temperature was recorded on a PC at 30 Hz using a digital multimeter (2000 DMM, Keithley Instruments, Cleveland, OH,

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USA). Each exposure within the set was at a different HIFU intensity, but all exposures were of the same exposure duration. The time between exposures varied, with a maximum of several minutes, so as to allow the temperature to return to the pre-exposure temperature before beginning the next exposure. The thermocouple was then removed, and the same sequence of HIFU exposures and cooling times repeated while IQ echo data were acquired continuously for strain imaging. Strain was calculated from the echo displacements measured relative to the echo positions immediately before each exposure. Three non-perfused tissue calibration experiments were performed, one on each of three bovine liver tissue samples. Each tissue sample came from a different animal. Each experiment began with the liver at 20 C, and an increment in intensity between exposures was chosen with the aim of increasing the peak temperature rise by 0.5 C more than that achieved during the preceding exposure. All experiments employed the L10–5 probe imaging the trans-axial cross section at the HIFU focus. The first experiment was performed using the spherical annulus HIFU transducer. Temperature and strain data were obtained once for each 5-s exposure and during the first 5 s of the post-exposure cooling period. The second and third experiments employed the spherical bowl HIFU transducer, with exposure times of 3 and 10 s and post-exposure data collection periods of 7 and 10 s, respectively. For the first two experiments, an additional set of exposures were performed with the HIFU focus positioned a short lateral distance (1 cm) from the thermocouple position, at a constant depth in tissue. Strain was imaged at this new focal position. This additional set of strain data was used to study whether spatial variations in measured strain, which might indicate differences in local tissue properties (e.g., absorption, thermal conductivity), could be observed. Three perfused liver experiments were carried out, each with a different liver, maintained at 38 C. During each experiment, a set of 10-s HIFU exposures, giving peak temperature rises between 0 and 15 C, were performed at increments of approximately 2.5 C using the spherical annulus HIFU transducer. The spatial peak intensity did not exceed 400 W/cm2. These relatively long exposures at low power were chosen in an attempt to minimize bulk tissue displacement due to the effect of acoustic radiation force created by the HIFU beam acting on the tissue. Thermocouple data were obtained during the HIFU exposure and for the first 10 s of cooling; strain images were calculated for the cooling period. For the first two experiments, the needle thermocouple had to remain during ultrasound echo data acquisition because it was required for another study, outside the scope of this article. It was therefore not possible to adopt the

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two-stage procedure described above. Instead, the second set of exposures took place with the HIFU focus repositioned 1 cm from the thermocouple laterally, whilst maintaining the same focal depth and strain measured at this position. For the third experiment it was possible to remove the thermocouple after acquisition of the thermocouple temperature data; therefore, it was possible to acquire strain data at the same position with the thermocouple withdrawn and, subsequently, with both the HIFU focus and the point of strain measurement repositioned by 1 cm as described above. Comparison of these two sets of measurements in the third liver allowed preliminary assessment of the validity of the method used for the first two perfused livers. Thermocouple measurements Thermocouple measurements of the temperature rise caused by the absorption of ultrasound are known to exhibit a viscous heating artifact (Fry and Fry 1954) that arises from the relative motion between the thermocouple and the tissue that surrounds it when an ultrasound field is applied. This motion leads to increased heating at the thermocouple, which would not occur if the device was not present. The result is that temperature-versustime curves measured with thermocouples show disproportionately large initial heating. As the exposure continues, this component occupies a progressively smaller proportion of the measured temperature rise, which tends toward the value that would be measured in the absence of viscous heating. To make a fair comparison between strain and thermocouple temperature measurements, it is therefore necessary to estimate and remove the viscous heating artifact from the thermocouple measurements or, if this is not possible, to compare temperature measurements with strain at time points at which viscous heating can be assumed to contribute only a small percentage (,10%) of the overall temperature rise. The second-order differential method of Morris et al. (2006) was used to determine the viscous heating component of the temperature rise. Briefly, this method provides the time point during the heating phase at which the temperature rise caused by viscous heating appears to stabilize. The rate of change in temperature at this time point is estimated by means of a linear regression applied to temperature data selected over a short time window (0.45 s) centered at the selected time point. This rate of change in temperature is then used to extrapolate back to the beginning of the exposure (t 5 0) with a second linear regression to estimate the viscous heating artifact. The true spatial peak temperature rise was therefore estimated by subtracting the viscous heating component from the measured value. This was done for all temperature measurements during the heating phase, except those during the initial 0.5 s, when the viscous

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heating component was not considered to have reached stability; hence these data points were not used in the comparison with strain measurements described below. The second-order differential method was also used to determine the time during cooling at which the component resulting from the presence of the thermocouple could be considered to be negligible (equivalent to approximately less than 10% of the total temperature change). Figure 5 provides an example of thermocouple measurement data from the perfused liver calibrations. In this figure are uncorrected temperature-versus-time data and the values of the viscous heating corrections at the ends of the HIFU exposures. Figure 5 also indicates the region (rectangle) that was used to select data for comparison with strain measurements during the cooling phase (corresponding to times between 14 and 20 s), which ignored the first 4 s of cooling during which the artifact resulting from the presence of the thermocouple was considered unacceptable. Backscatter temperature imaging Echo image acquisition was started shortly before the start of HIFU sonication, the last pre-HIFU frame being chosen as a reference frame. Axial echo strain images were produced as described by Miller et al. (2002). Briefly, the axial displacements of echoes between the reference frame and frames during or after HIFU were estimated using a one-dimensional crosscorrelation search of the RF A-lines in each image (Ophir et al. 1991). RF echo data, required by the algorithm used for echo displacement tracking, were obtained

Fig. 5. Set of thermocouple temperature measurements obtained in perfused liver tissue. The series represents temperature measurements from six successive high-intensity focused ultrasound exposures at 1.7 MHz with corresponding approximate spatial peak intensities of 50, 100, 150, 200, 250 and 300 W/cm2, yielding increments of 2.5 C in peak temperature rise between exposures. The arrows and dots illustrate the effect of applying the viscous heating correction to the peak temperature rises. The dashed box represents the period during the cooling phase when the temperature and strain measurements were compared.

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as the real part of the product of the IQ complex beamformed echo data, after up-sampling to 40 MHz using cubic interpolation, and the complex conjugate of the phasor that the scanner had used to derive the IQ data. This was calculated using a modified version of a MATLAB script provided by the scanner’s manufacturers. Correlation window sizes in the reference frame were chosen as the shortest that would cover more than one echo pulse; that is, 71 samples (1.35 mm) for data acquired using the L10–5 imaging probe and 141 samples (2.7 mm) for the C5–2 probe. Overlaps between successive reference windows, 85% and 70%, provided displacement images with pixel separations of 0.2 and 0.8 mm for the L10–5 and C5–2 probes, respectively. An axial strain image was computed from each displacement image using a least-squares strain estimator (Kallel et al. 1999), with window sizes of 21 and 29 displacement values (4.3 and 24 mm) for the L10–5 and C5–2 probes, chosen as the largest (for reduction of strain noise) that would retain adequate spatial resolution for visualizing the width of the focal region of the spherical bowl HIFU transducer and the length of the focal region of the spherical annulus HIFU transducer, respectively. Strain values for comparison with thermocouple measurements of temperature rise were then extracted from these images as the spatial peak strain at the location of the HIFU focus. To improve the precision of this measurement the mean value within a 7 3 5 strain pixel window centered on the focal peak was used. Analysis of strain as a function of temperature rise As described by Miller et al. (2002), induced apparent strain can be considered to vary linearly with changes in sound speed, provided the relative change in sound speed is small (,1%). It follows, therefore, that strain should provide a good quantitative measure of temperature rise if the relationship between speed of sound and temperature can also be assumed to be linear, with known slope. As described earlier however, the speed of sound-versus-temperature relationship in liver tissue, even when reversible, is known to be non-linear. The strain-versus-temperature rise relationship will therefore also be non-linear. In non-fatty liver tissue, the sound speed gradient (Dc/DT) is greatest at lower temperatures (room temperature) and decreases with increasing temperature until it reaches zero (typically between 50 C and 60 C) and becomes negative thereafter. Because the BTI strain measurements of interest here were limited to a narrow temperature range suitable for low-power HIFU test exposures (and, therefore, to a narrow, approximately linear, portion of the c–T curve), second-order polynomial fits were considered sufficient to describe the general trend of the measured strain (ε) versus temperature rise. These fits were described in

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terms of the coefficients a and b, where Tr is the temperature rise: εðTr Þ 5 aTr2 1bTr

(1)

The polynomial coefficients were derived for strain and temperature measurements made during the heating phase for the non-perfused tissue calibrations (i.e., when the echo data were not affected by HIFU interference) and during the artifact-free period of the cooling phase described above for both non-perfused and perfused tissue calibrations. Data were analyzed in this way separately for each liver and by pooling data into two groups: (i) the non-perfused livers and (ii) the perfused livers. RESULTS Ex vivo (non-perfused) bovine liver In Figure 6, the measured spatial peak strain is plotted as a function of temperature rise for all exposures,

Fig. 6. Measured strain as a function of temperature rise measured with a thermocouple placed at the high-intensity focused ultrasound focal peak in the ex vivo non-perfused bovine liver tissue samples: for (a) liver 1, 5-s exposures; (b) liver 2, 3-s exposures; (c) liver 3, 10-s exposures. Symbols represent whether strain data were acquired at the thermocouple position (black) or away from the thermocouple position (gray); see text for details. Second-order polynomial lines of best fit are included (dashed lines).

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Table 1. Second-order polynomial fit coefficients (a and b) for the non-perfused bovine liver tissue calibrations illustrated in Figure 5 (transverse alignment)* s

Polynomial coefficient Liver

HIFU transducer

Exposure time (s)

Site

a

b

s (%)

s ( C)

1 1 2 2 3 Mean SD All data

Annulus Annulus Bowl Bowl Bowl

5 5 3 3 10

T A T A T

0.0028 20.0021 20.0039 20.0165 20.0017 20.0043 60.007 20.0128

0.1259 0.114 0.0865 0.1933 0.0564 0.1152 60.051 0.1259

0.0181 0.0159 0.0173 0.0313 0.0282 0.0222 60.007 0.0880

0.1438 0.1395 0.2000 0.1619 0.5000 0.2290 60.153 0.6990

SD 5 standard deviation; T 5 thermocouple site; A 5 approximately 1 cm away from thermocouple site (but at same depth in tissue). * Root-mean-squared deviations (s) in measured strain, about the lines of best fit, provide a measure of uncertainty in the strain measurement, which is scaled as s(%)/b to provide a prediction of the uncertainty that would be experienced if strain were to be used as a measure of temperature rise (assuming a linear relationship). Results are provided for individual experiments (with means and standard deviations) and for all data in a pooled analysis.

in each of the three bovine liver samples. The graphs include the second-order polynomial lines of best fit to the data. The polynomial fit coefficients are reported in Table 1. Second-order coefficients (a) of the polynomial fits are typically small compared with the first-order coefficient (b), indicating that for small temperature rises (,5 C), the strain has a quasi-linear relationship with temperature rise, as can be seen in Figure 6. The average value of the parameter b, the linear term, is 0.115 (range: 0.056%–0.193% strain/ C temperature rise). Within each calibration data set, the root-mean-squared (RMS) deviation of the measured strain values about the line of best fit

provides a measure of uncertainty, and hence sensitivity, in the strain measurements. On average, this is 0.022% strain (range: 0.016–0.031), with a corresponding temperature sensitivity (assuming a linear relationship between strain and temperature rise) of 0.23 C. For a pooled analysis of all calibration data, the b value is 0.1259% strain/ C. The RMS deviation for all strain data is 0.088% strain, which provides a measure of uncertainty that would be associated with a new measurement using the calibration from the pooled data for temperature estimation in any of the individual tissues. It reflects differences in the underlying speed of sound-versus-temperature

Fig. 7. Example regions of apparent displacement (a), correlation coefficient (b) and apparent strain (c) images and their corresponding pre-heating B-mode image (d) in a perfused liver 1 s after the end of a high-intensity focused ultrasound exposure that resulted in a peak temperature rise of 9 C. The location of the measured peak strain is indicated by the cross in the strain image.

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Fig. 8. Spatial peak strain as a function of temperature rise, measured during the cooling phase (the period 14–20 seconds after the start of the exposure) for liver 1 (a), liver 2 (b), liver 3 at the thermocouple position (c) and liver 3 at a position 1 cm away from the thermocouple position (d). Data series represent the six different exposures (with peak temperature rises of 2.5, 5, 7.5, 10, 12.5 and 15 C). The polynomial fits for each experiment (dashed line) and for the pooled data from all four experiments (dotted line) are included.

relationship across the different livers due to heterogeneity as well as the error in strain measurement. When divided by the slope b, this gives a temperature equivalent uncertainty for the pooled calibration of approximately 0.7 C. The typical RMS strain uncertainty for single livers stated above (0.0222%) is smaller than that for the pooled analysis (0.088%), as indicated in Table 1. Furthermore, close inspection of the measurements within Figure 6 reveals that their relatively small deviation about the best-fit predicted values does not depend on the time of measurement during any exposure and that there is no substantial contribution from variation in measurements between exposures. Results for liver 2 (Fig. 6b and Table 1) indicate that strain values and the temperature coefficient of strain b can more than double from one spatial location to another, even when the two locations are separated by a distance of only 1 cm. It was, however, not possible

to verify the temperature rise at the second location, and therefore, it is not possible to separate conclusively any perceived differences in the strain-versus-temperature rise curve from any local changes in the HIFU heating rate at this new position. Perfused bovine liver Examples of the displacement, correlation and strain images obtained from the perfused livers are provided in Figure 7. Negative apparent displacements (toward the imaging transducer) were observed within, and distal to, the heated region (Fig. 7a), consistent with an increase in the speed of sound in liver after HIFU heating. In the strain image, the location of the HIFU focus is depicted by the bright region (Fig. 7c). The expected position of the peak temperature rise, determined by visualizing the localized thermocouple on the ultrasound B-mode image that was registered with the strain image, was found always to be within 3 mm of the peak in the strain image.

Ultrasound backscatter temperature imaging calibration for HIFU d J. CIVALE et al.

Speckle de-correlation can be seen distal to the focus. This worsened with increasing HIFU power/peak temperature, consistent with the observations of Miller et al. (2005), who also explain its origin. Other areas of de-correlation may be due to poor echo signal-to-noise ratio and image clutter from unwanted reflections or reverberations. The measured spatial peak strain is plotted as a function of temperature rise in Figure 8, for the cooling phase for all six exposures in all four calibration data sets. The second-order polynomial fits applied to the data from each exposure reveal a quasi-linear relationship between strain and temperature rise, over the range of temperature rises experienced during the cooling period studied. The strain-temperature polynomial coefficients are summarized in Table 2. The average value of the linear coefficient b is 0.040% strain/ C (range: 0.025–0.062), which is lower than for the ex vivo tissues. The average value of the single liver RMS deviation, or strain sensitivity, is 0.013% strain (range: 0.009%–0.02%), with a corresponding average temperature sensitivity of 0.33 C. These values are similar to those for the ex vivo tissues. The polynomial fit to all data from all experiments, represented by the faint dotted line in Figure 8, provides, in the linear approximation, a strain temperature coefficient b of 0.040% strain/ C, which, as for the mean of the individual b values, is lower than the value for the ex vivo calibrations. The RMS deviation of all strain data about this best fit relationship is 0.029% strain. When divided by the slope b, this gives a final temperature equivalent uncertainty of approximately 0.7 C, a value very similar to that obtained from the ex vivo calibrations.

Table 2. Second-order polynomial fit coefficients (a and b) for the perfused porcine liver calibrations illustrated in Figure 8 (co-axial alignment)* s

Polynomial coefficient Liver

Site

a

b

s (%)

s ( C)

1 2 3 3 Mean SD All data

A A A T

20.0022 20.0046 20.000077 20.0021 20.0022 60.002 20.0022

0.0355 0.0623 0.0247 0.0381 0.0402 60.016 0.0402

0.0089 0.0197 0.0121 0.0101 0.0127 60.005 0.029

0.2507 0.3162 0.4899 0.2651 0.3304 60.110 0.7214

SD 5 standard deviation; T 5 thermocouple site; A 5 approximately 1 cm away from thermocouple site (but at same depth in tissue). * Root-mean-squared deviations (s) in measured strain, about the lines of best fit, provide a measure of uncertainty in the strain measurement, which is scaled as s(%)/b to provide a prediction of the uncertainty that would be experienced if strain were to be used as a measure of temperature rise (assuming a linear relationship). Results are provided for individual experiments (with means and standard deviations) and for all data in a pooled analysis.

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As for the ex vivo experiments, the typical RMS strain uncertainty for single livers is less than that for the pooled analysis. For the perfused livers, however, this occurs despite the fact that close inspection of the strain data from the cooling curves for individual livers (Fig. 8) indicates that the variation in strain within single livers arises in part from the variations between the successive exposures. This is a major difference from the ex vivo calibration measurements (Fig. 6), and possible explanations are provided in the Discussion. Spatial variation of results within a liver is observed in Figure 8 and Table 2; in the third liver, the two sets of strain measurements, one obtained at the thermocouple location and the other 1 cm away, were very different. At the location away from the thermocouple, the bestfit polynomial was more linear over the entire temperature-rise range studied and had a lower slope over half the range (below 3 C), than that at the thermocouple position. For all temperature rises, the measured strain away from the thermocouple position was substantially lower. DISCUSSION The results described in this article confirm the ability to visualize apparent strain caused by temperature-dependent changes in sound speed in nonperfused ex vivo bovine liver and extend previous studies in demonstrating that this is also possible in a perfused living porcine liver model maintained at normal body temperature. To obtain strain images with low noise, leastsquares strain estimator (LSSE) window sizes were used that were similar in dimensions to the 6-dB HIFU focal peak in the direction of the imaging axis. One of the consequences of this is that the measured spatial peak strain is underestimated by a factor that is dependent on the relative dimensions of the HIFU focal region and the LSSE window (Civale 2008; Miller et al. 2004). The peak strain values reported in this article are therefore representative only of those to be expected for the HIFU fields and LSSE windows employed. The use of a different LSSE window, or different HIFU focal dimensions, would lead to different peak strain values when applied to the same raw RF data. As shown by Miller et al. (2004), however, the factor by which the peak strain is reduced by the smoothing effect of the LSSE window is independent of the induced temperature rise. It was thus considered acceptable here to ignore the effect of the LSSE window when analyzing the relationship between measured strain and temperature rise. This underestimate was thought to be less than 50% and 35%, respectively, for the trans-axial (non-perfused studies) and co-axial (perfused studies)

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alignments and their respective strain processing parameters (Civale 2008). The measurements of temperature-induced apparent echo strain in three samples of ex vivo non-perfused bovine liver and three living perfused whole livers indicated that the exact relationship between strain and temperature rise is highly tissue sample dependent, even within each of these two groups. This finding may be consistent with the variation found between liver samples in the relationship between speed of sound and temperature that has been observed by others, such as Bamber and Hill (1979) (0.1% strain/ C at 20 C), Miller et al. (2004) (0.07% strain/ C at 20 C), Varghese et al. (2002a) (0.13% strain/ C at 20 C), Techavipoo et al. (2004) (0.08% strain/ C at 20 C) and Ghoshal et al. (2011) (0.06% strain/ C at 37 C). Varghese and Daniels (2004) attempted an echo strain calibration for the purposes of monitoring heating during RF ablation procedures. They reported linear trends of relative changes in echo shifts with increasing temperature over the range 37–100 C. This linearity was, however, mostly attributed to tissue expansion effects after prolonged heating (20 min) of the tissue to high temperatures. This is in contrast to our calibrations, which used short, lower-magnitude temperature changes. The data presented here also indicate spatial variation in measured strain, which can be as large within a liver sample as between livers. Further research is needed to determine the cause of this variation. Possible causes include changes in localized HIFU heating rate (temperature) and heterogeneity in the relationship between tissue sound speed and temperature. Strain measurement in the perfused porcine livers posed a number of difficulties. The clinically desirable co-axial alignment of therapy and imaging beams is not ideal for BTI because, as observed previously (Bamber et al. 1997; Miller et al. 2005), this configuration maximizes the speckle de-correlation at, and distal to, the focus, which could have led to noisier echo strain estimates in this region compared to those obtained with transverse alignment. In addition, the co-axial configuration placed the HIFU focus at a depth (15 cm) for which the imaging system has to operate at low ultrasonic frequencies and has very poor resolution in comparison to the high-frequency linear probe that was positioned close to the HIFU focal region in the ex vivo nonperfused tissues. Nevertheless, the spatial peak strain typically occurred in a region that provided a sufficiently high echo correlation to enable quantitative analysis to be carried out, despite the use of this co-axial alignment (Fig. 7). Indeed, the peak strain measurements for perfused livers had a within-sample precision similar to that of the measurements for the ex vivo non-perfused tissue (Tables 1 and 2).

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Co-axial alignment is also more likely to produce a peak strain measurement bias than trans-axial alignment, because the center of the imaging beam corresponding to the image line for which the peak strain occurs may not coincide exactly with the spatial peak temperature rise, and even if it does, the line is formed with an imaging beam that covers regions of temperature rise that are lower than the peak. Although this is true for both configurations, the lower frequencies (and hence wider beams) and greater depths (and, hence, greater line spacing for a sector scan) imposed by co-axial alignment will make the phenomenon more severe than for trans-axial alignment. Furthermore, in the image elevation direction, the phenomenon will be particularly severe for co-axial alignment and almost non-existent for trans-axial alignment (because near its focus, the HIFU beam’s cross section varies slowly with axial position). As far as we are aware, this source of error has not previously been described. Further study is required to estimate its severity and properties and, if considered necessary, investigate correction methods. It is a source of variation in measured echo strain between experiments, and because the spatial distribution of temperature changes in width and shape (kurtosis) during heating or cooling, it may cause the shape of the curves of echo strain versus temperature rise to be different when the spatial peak temperature rise is varied by using different time points during heating or cooling, compared with when it is varied by changing HIFU intensity. It is difficult to be certain whether the measurements reported here display evidence of this error; perhaps it could explain why (e.g., in Fig. 8c) the Dε/DTr slopes for individual exposures using co-axial alignment often appear to be visually lower than the slope for the pooled data for each experiment. There was no evidence of this phenomenon within the individual exposure data for the non-perfused liver experiments, which may be expected because they were all carried out using trans-axial alignment and narrow imaging beams. The phenomenon may also have contributed to reducing the b-values of the polynomial fits for all co-axially aligned experiments (Table 2), relative to those for the trans-axially aligned experiments (Table 1), although this trend is most likely due to the higher starting temperature of the former experiments, as described below. The starting temperature of 38 C, used to simulate in vivo conditions, limits the temperature-rise range over which sub-lesioning exposures can be made and also limits the maximum strain that can be generated because it lies close to the peak of the speed of soundversus-temperature curve, typically at about 50 C (Bamber and Hill 1979). This is the most likely reason why experiments with perfused livers consistently produced lower gradients, for the temperature-rise dependence of echo strain, than those with non-perfused livers.

Ultrasound backscatter temperature imaging calibration for HIFU d J. CIVALE et al.

The echo strain-versus-temperature rise curves for perfused livers were continuous (i.e., exhibited little RMS residual variation) for data acquired within an exposure but discontinuous (i.e., exhibited considerable RMS residual variation) between exposures (Fig. 8). This apparent behavior was not observed for experiments with non-perfused livers. In addition to the possible explanation given above for the use of co-axial alignment for the perfused liver studies, these discontinuities may be explained by strains that were observed to occur in the absence of heating over periods of 10 s or longer. Highly preliminary evidence for this came from tracking tissue displacement using frames acquired when the tissue was not heated by HIFU exposures. Strain images under these conditions exhibited spatial in-homogeneity, with peak values estimated to be in the region of 0.02% for a time interval (10 s) equivalent to the interval between BTI reference and measurement frames, which is comparable to the magnitude of the differences between strain curves measured for separate exposures in the perfused liver tissue. The experiments were not designed to acquire data for an analysis of the temporal dependence of strain in the absence of heating, which is why the data are not provided here. These strains are believed to have been real tissue strain, because they were not temperature related and were not confined to a single location. They could have been due to blood vessel pulsation or movement of the unconstrained liver. During HIFU exposures, this situation is not likely to be improved as the radiation force associated with the HIFU beam could provide an additional source of tissue motion. Although further work is needed if the true sources of bias observed in the strain measurements in the perfused liver model are to be fully understood, these measurements highlight a need to overcome the effect of real tissue motion, particularly for organs such as the liver that lie close to the heart and lungs, if BTI is to be implemented successfully for guidance and control of clinical HIFU therapy. Speckle tracking methods for measuring pre-existing tissue motion and strain in the liver, for correction of elastograms, have been described (Kolen et al. 2004) and correction methods implemented (Kolen and Bamber 2003). Similar techniques have been implemented to provide motion correction for HIFU (Pernot et al. 2004) and radiotherapy (Harris et al. 2010). There is potential to extend these techniques to BTI, and early work in this direction has been carried out (Liu and Ebbini 2009). Predicted RMS uncertainties in the temperature-rise estimates that would be experienced if echo strain were to be used as a measure of temperature rise, assuming a linear relationship between strain and temperature rise and without knowing the correct calibration parameters, were found to be in the region of 60.7 C and were similar

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for the perfused and non-perfused liver studies (i.e., values for pooled data in Tables 1 and 2). This similarity was brought about, however, by the fact that the bvalue was lower for perfused liver than non-perfused liver studies; the RMS uncertainties in echo strain alone were in fact greatest for the pooled non-perfused liver experiments, which is surprising given the greater number of sources of error in strain estimation known to exist in the perfused liver studies, and it probably represents some evidence toward heterogeneity in the tissue Dc/ DT properties. Indeed, although insufficient data exist to conduct a full analysis of variance, it may be seen in Tables 1 and 2 that the RMS deviations about the bestfit polynomial lines are greater for pooled data than for individual experiments; that is, the between-experiment variance is larger than the within-experiment variance. Finally, the observed difference in the relative strain RMS uncertainties between the non-perfused and perfused studies may be due to differences in the relative amount of heterogeneity between the population of livers from which the tissues were sourced; for example, bovine and porcine livers were used exclusively for the non-perfused and perfused studies respectively. Tissue autolysis may also be a factor for the non-perfused calibrations, as these livers were stored refrigerated overnight before the experiment. Returning to the stated aim of the work reported here, if BTI is to have value as a pre-ablation means of estimating the HIFU parameters required to carry out a controlled ablation, taking the 60.7 C uncertainty in temperature-rise estimates mentioned above as an a example, this quantity is large for a technique designed to measure low (5 C) temperature rises for HIFU treatment planning. In addition, however, it is instructive to consider a reasonable worst-case use of an incorrect calibration. It would be reasonable, for example, to consider the consequences of using the polynomial fit to the pooled data from the perfused liver study (Table 2) as a calibration for a future ablation study. As a worst-case example of applying this calibration, taken from the measurements reported in Figure 8, it can be seen that a HIFU exposure designed to produce a spatial peak temperature rise of about 8 C would in fact, for liver 2 (Fig. 8b), have produced a temperature rise in the region of 4 C. In other words, if 8 C were to be the HIFU-induced increase in temperature at the focal peak required to produce the desired ablation, a BTI test exposure would substantially underestimate the HIFU source power required to achieve this in liver 2. In practice, the consequences of using this calibration are likely to be more severe than this, as the calibration does not account for the fact that ultrasound of the intensity required to cause tissue ablation will propagate with greater non-linearity than needed for the test

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exposure. These observations appear to suggest that BTI of non-ablative test exposures will not be able to provide an adequate method for pre-ablation control of HIFU therapy exposure parameters to adjust for small (normal) variations in liver attenuation, absorption coefficient and thermal dissipation. Nevertheless, it remains plausible that the BTI test exposure approach could have a use in supplementing conventional imaging, in detecting large departures from normal levels of the relevant tissue properties, such as unusually high intervening attenuation from gas or bone, or unusually low absorption at the HIFU focus such as within a large blood vessel or a necrotic region of a tumor, or unusually rapid local heat dissipation caused by blood flow. None of these situations were studied in the experiments reported here and, therefore, would represent topics for further study. The mean within-experiment uncertainty (Table 2) is about 60.3 C. This can be thought of as a measure of the sensitivity of BTI, that is, the smallest detectable temperature rise between exposures within a given tissue region. Such a value for sensitivity is consistent with the ability of BTI, even with co-axial alignment, to visualize the location of the HIFU focus using non-ablative test exposures (Miller et al. 2005), confirming that tissue targeting is a viable application for BTI. Yet another application for BTI may be in HIFU treatment monitoring, to detect when irreversible tissue damage has occurred and to confirm the region of tissue that has been thermally ablated. This would be as an alternative to the direct imaging and measurement of tissue properties such as attenuation coefficient (Baker 2003; Bush et al. 1993; Lemor et al. 2002; Zderic et al. 2004) and elasticity (Bercoff et al. 2004; Doyley et al. 1999; Kallel et al. 1999; Varghese et al. 2002b). The temperature coefficient of sound speed undergoes no obvious sudden transition in the 40–50 C temperature range, where biomolecules tend to denature, whereas heat denaturation causes a substantial increase in ultrasonic attenuation (Bamber and Hill 1979). The apparent strain-versus-temperature rise curve for the healthy tissue, which may be unknown, may therefore be applicable to the same tissue when ablated. If so, an assumed BTI calibration would no longer be needed, and for HIFU test exposures of equal power, BTI would become a method for detecting the increase in ultrasonic absorption coefficient, and possibly also decreased blood perfusion, caused by ablation (as long it is ensured that blood perfusion effects do not modulate or bias the pre-ablation BTI measurement in any way). This hypothesis would appear to be confirmed by the results obtained by Liu et al. (2010), who observed an increase in peak measured strain a few minutes after the end of an ablative exposure in ex vivo myocardial tissue. These

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authors also identified an increase in tissue absorption as the principal cause of the increase in measured strain they observed. CONCLUSIONS A reliable calibration of echo strain against temperature rise is of fundamental importance if BTI is to play a quantitative role in HIFU treatment planning. The primary purpose of the study described in this article was to determine whether apparent spatial peak echo strain measured in tissues can provide a reliable estimate of peak temperature rise generated by low-power HIFU exposures. Verification of whether a suitable calibration can be derived is needed, and hence our results can be considered as a first step in attempting to answer this question. BTI calibration measurements were performed in a number of tissue samples to determine whether the measured strain is a sufficiently reliable measure of HIFU focal peak temperature rise for BTI to have potential as a quantitative HIFU treatment planning method. An important and novel feature of our work is that we report measurements performed in living perfused porcine livers, an experimental model designed to simulate some in vivo conditions. The main finding from our calibrations was that the relationship between echo strain and temperature rise is highly tissue sample dependent. The total uncertainty (RMS deviation) of a BTI measurement of temperature rise in liver, when the only calibration information available is knowledge of tissue type (non-fatty liver) and the starting temperature, was estimated to be approximately 60.7 C, for both the ex vivo and perfused livers. If this level of estimated uncertainty is truly indicative of the overall uncertainty, it is large for a technique designed to register low (5 C) temperature rises for HIFU treatment planning. In fact, the worst example in this study, of using the calibration derived from our pooled data, gave rise to an overestimate of nearly 100% in one tissue sample. Quantitative BTI in its present form, therefore, seems to be fundamentally limited by tissue heterogeneity in the calibration between strain and temperature rise. One major source of variation in calibration was the starting temperature, that is, room temperature for the ex vivo non-perfused tissues and body temperature for perfused organs. Few other differences in results between the ex vivo and simulated in vivo tissues were noted, despite the fact that the latter deliberately employed less favorable experimental conditions. In particular, although a co-axial alignment of the therapy and imaging beams used in the simulated in vivo experiments produced strain images that were noisier in the distal part of the HIFU focal region than those for transverse alignment (also noted

Ultrasound backscatter temperature imaging calibration for HIFU d J. CIVALE et al.

previously by Miller et al. 2005), spatial peak strain was unaffected, and the strain precision and sensitivity were similar for the two alignment methods. This is a potentially helpful finding, given the clinical usefulness of the co-axial alignment method. Our findings therefore suggest that BTI may be confined to the detection of situations where extremely unusual values have been encountered for tissue properties (acoustic attenuation in the path and acoustic absorption or thermal dissipation at the focus) or equipment has malfunctioned. Alternatively, they indicate the need for further work to determine whether much larger temperature rises can be employed in the test exposures. This will require techniques to avoid irreversible effects on the tissue, perhaps by the use of very short but highintensity exposures, and methods to exceed the peak in the curve of sound speed versus temperature (Miller et al. 2002), perhaps by using multiple exposures of different intensity to ‘‘unwrap’’ the temperature ambiguity or by using the strain image to spatially unwrap it. Any future clinical application of quantitative BTI for HIFU treatment planning will also require detection of, and compensation for, pre-existing mechanical strain, such as may arise from cardiovascular sources and respiration. It would also benefit from the use of an echo imaging system designed specifically for this purpose, that is, one that provides adequate lateral and elevational resolution at the HIFU focal depth in a co-axial configuration or one that uses a non-co-axial approach. Even if BTI cannot be used reliably for our intended application of delivering quantitative evaluation of temperature rise from low-power HIFU exposures, continued development of BTI is worthwhile if only because of its useful role in visualizing the HIFU focal position before lesion formation (Miller et al. 2005) or its potential for detecting the change in acoustic absorption that occurs at the point of thermal coagulation (Liu et al. 2010). Acknowledgments—This work was supported by the Institute of Cancer Research and EPSRC Grants GR/64042, EP/F02617 X/1 and EP/ F029217/1. The authors thank Zonare Medical Systems for support with the Z.One ultrasound imaging system, David Napolitano (Zonare Medical Systems) and Nigel Bush (Institute of Cancer Research) for providing the software for conversion of IQ image data to radiofrequency.

REFERENCES Baker LAS. Ultrasonic reflex transmission imaging with commercial linear arrays for monitoring minimally invasive ablation therapies [PhD thesis]. London: University of London; 2003. Bamber JC, Hill CR. Ultrasonic attenuation and propagation speed in mammalian tissues as a function of temperature. Ultrasound Med Biol 1979;5:149–157. Bamber JC, Meaney PM, Doyley MM, Clarke RL, ter Haar GR. Noninvasive temperature imaging using ultrasound echo strain: Prelim-

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Volume 39, Number 9, 2013 Souchon R, Bouchoux G, Macieko E, Lafon C, Cathignol D, Bertrand M, Chapelon JY. Monitoring the formation of thermal lesions with heat-induced echo-strain imaging: A feasibility study. Ultrasound Med Biol 2005;31:251–259. Sun Z, Ying H. A multi-gate time-of-flight technique for estimation of temperature distribution in heated tissue: Theory and computer simulation. Ultrasonics 1999;37:107–122. Techavipoo U, Varghese T, Chen Q, Stiles TA, Zagzebski JA, Frank GR. Temperature dependence of ultrasonic propagation speed and attenuation in excised canine liver tissue measured using transmitted and reflected pulses. J Acoust Soc Am 2004;115:2859–2865. Ter Haar GR, Rivens IH, Moskovic E, Huddart R, Visioli AG. Phase one clinical trial of the use of focused ultrasound surgery for the treatment of soft-tissue tumours. SPIE 1998;3249:270–276. Vallancien G, Harouni M, Veillon B, Mombet A, Prapotnich D, Brisset JM, Bougaran J. Focused extracorporeal pyrotherapy: Feasibility study in man. J Endourol 1992;6:173–180. Varghese T, Daniels MJ. Real-time calibration of temperature estimates during radiofrequency ablation. Ultrasonic Imaging 2004;26: 185–200. Varghese T, Zagzebski JA, Chen Q, Techavipoo U, Frank G, Johnson C, Wright A, Lee FT Jr. Ultrasound monitoring of temperature change during radiofrequency ablation: Preliminary in-vivo result. Ultrasound Med Biol 2002a;28:321–329. Varghese T, Zagzebski JA, Lee FT. Elastographic imaging of thermal lesions in the liver in vivo following radiofrequency ablation: Preliminary results. Ultrasound Med Biol 2002b;28:1467–1473. Visioli AG, Rivens IH, ter Haar GR, Horwich A, Huddart RA, Moskovic E, Padhani A, Glees J. Preliminary results of a phase I dose escalation clinical trial using focused ultrasound in the treatment of localised tumours. Eur J Ultrasound 1999;9:11–18. Wu F. Extracorporeal high intensity focused ultrasound in the treatment of patients with solid malignancy. Minimally Invasive Ther Allied Technol 2006;15:26–35. Wu F, Wang ZB, Chen WZ, Wang W, Gui Y, Zhang M, Zheng G, Zhou Y, Xu G, Li M, Zhang C, Ye H, Feng R. Extracorporeal high intensity focused ultrasound ablation in the treatment of 1038 patients with solid carcinomas in China: An overview. Ultrasonics Sonochem 2004;11:149–154. Zderic V, Keshavarzi A, Andrew MA, Vaezy S, Martin RW. Attenuation of porcine tissues in vivo after high intensity ultrasound treatment. Ultrasound Med Biol 2004;30:61–66.