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In vitro and in vivo ablation of porcine renal tissues using high-intensity focused ultrasound
Ultrasound in Med. & Biol., Vol. 29, No. 9, pp. 1321–1330, 2003 Copyright © 2003 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/03/$–see front matter
doi:10.1016/S0301-5629(03)00981-5
● Original Contribution IN VITRO AND IN VIVO ABLATION OF PORCINE RENAL TISSUES USING HIGH-INTENSITY FOCUSED ULTRASOUND CHRISTAKIS DAMIANOU Frederick Institute of Technology, Limassol, Cyprus (Received 13 November 2002; revised 25 April 2003; in final form 1 May 2003)
Abstract—The aim of this paper is to present issues regarding the thermal ablation of porcine renal tissues in vitro and in vivo using high-intensity focused ultrasound (HIFU). Production of lesions in the cortex in vitro is consistent, whereas lesions in the medulla are created whenever there are no air spaces in the medulla. Typically, the lesion length at 2000 W/cm and 5-s pulse duration is around 20 mm and the corresponding width around 3 mm. Lesioning of a large volume was achieved by moving the transducer in a grid formation. Lesioning through a fat layer is possible provided that there are no air spaces between the fat and kidney interface. It was found that, above 3200 W/cm with 5-s pulse duration at 4 MHz, cavitation activity occurred in most of the lesions created. (E-mail: [email protected]) © 2003 World Federation for Ultrasound in Medicine & Biology. Key Words: Ultrasound, Kidney, Attenuation, Simulation, Lesion, Cortex, Cavitation.
imaging. Ultrasonic imaging is the simplest and most inexpensive method; however, it has poor contrast between soft tissues. On the other hand, magnetic resonance imaging (MRI) offers superior contrast, but it is more expensive. Several trials have been conducted in the area of ultrasonic imaging (for example, Seip and Ebbini 1995; Maass-Moreno et al. 1996) and in the area of MRI (for example, Cline et al. 1992; Hynynen et al. 1993a, 1993b), thus enhancing the potential of HIFU. The main goal of HIFU is to maintain a temperature between 50 to 100°C for a few s (typically ⬍ 10 s), in order to cause tissue necrosis. Typically, focal peak intensity between 1000 to 10,000 W/cm2 is used with pulse duration between 1 to 10 s and a frequency of 1 to 5 MHz. This paper reports additional experience in the ablation of kidney tissue. For applications such as the treatment of benign prostatic hyperplasia (BPH), it is sufficient to destroy as much tissue as possible for the purpose of tissue debulking. For applications in oncology, destruction of all viable tumor cells is required and, therefore, protocols in this area must be very accurate and reliable. In this work, methods suggested by Malcolm and ter Haar (1996) were used to produce complete, reliable and consistent ablation of renal tissues (creation of a contiguous array of touching lesions of thermal origin, avoiding boiling and cavitation, and use of cooling to avoid merging of lesions in front of the focus).
INTRODUCTION High-intensity focused ultrasound (HIFU) is a noninvasive procedure for heating tumours without affecting the healthy tissue surrounding the tumour. Therefore, application of HIFU is being investigated as an alternative to standard surgical techniques. Although the idea of using HIFU was proposed in the middle of this century by Lynn et al. (1942), its maximum potential for clinical use has been established only recently, due to the developments of sophisticated systems (for example, Chapelon et al. 1992b; Hynynen et al. 1993a and Birhle et al. 1994). HIFU has been proven over nearly 60 years to be an effective and efficient method for ablating soft tissue, but its commercial success is still under trial. HIFU was explored in almost every tissue that is accessible by ultrasound. The following represent some examples of some applications explored: eye (Lizzi et al. 1984), prostate (Sanghvi et al. 1991; Chapelon et al. 1999), liver (ter Haar et al. 1989), brain (Fry et al. 1954; Lele 1962; Vydkotseva et al. 1994) and kidney (Linke et al. 1973; Chapelon et al. 1992a; Hynynen et al. 1995). Recently, the technology of HIFU systems has improved because the ultrasonic therapy can be guided by
Address correspondence to: Dr. C. Damianou, Frederick Institute of Technology (FIT), 18, Mariou Agathangelou, 3080, Limassol, Cyprus. E-mail: [email protected] 1321
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Experiments were performed using a generic HIFU system that includes a signal generator, a radiofrequency (RF) amplifier, a 3-D robotic system, a transducer and a personal computer (PC) that controls the entire system. The transducer of 4-cm diameter was spherically focused, operating at 4 MHz. Pulse duration of 5 s was used in all the experiments to minimise effects of blood perfusion. Results of HIFU ablation of porcine kidney in vitro and in vivo are presented. Single lesions in the cortex and medulla with or without a fat layer were created. The ultimate goal was to create large lesions, which was accomplished by moving the transducer in patterned schemes. The experimental results are compared with the results of a simulation model that includes the effect of rapid increase of attenuation during the transition of a tissue from healthy to necrotic. Thus, the kidney attenuations of porcine cortex, medulla, muscle and fat were measured as a function of thermal dose. The attenuation was measured using a system that includes two low-intensity transducers, a signal generator, a data-acquisition card and a PC. Attenuation includes absorption, scattering and reflection but, when minimising scattering and reflection, attenuation will reflect mostly losses due to absorption. Scattering and reflection was minimised by using a homogeneous sample with a minimum number of interfaces or air ducts, and by removing acoustically troublesome gas bubbles using degassing techniques. Using these precautions in the study by Damianou et al. (1997), it was found that the value of absorption was very close to the value of attenuation, indicating that the contribution of reflection and scattering was minimised. Absorption is an acoustical property of tissue with wide variation from tissue to tissue and represents the rate at which energy in tissue is converted to heat. For short pulses (⬍ 5 s), it is one of the primary factors contributing to the temperature elevation in tissue because the effect of blood flow is minimised.
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Fig. 1. Block diagram of the high-intensity focused ultrasound system.
the 3-D robotic system was controlled using the parallel port. A user-friendly program written in Visual Basic has been developed to control the system. Figure 2 shows the main window of the software. The software serves six main tasks through various windows or menus: 1. imaging (in this window US or MRI images can be displayed), 2. transducer movement (the user may move the robotic arm in a specific direction or customise the automatic movement of the robotic arm in any formation by specifying the pattern, the step and the number of steps), 3. messaging (starting time, treatment time left, etc.), 4. telematics (a physician in a remote site can monitor the treatment protocol and can suggest ideas during the treatment, 5. patient data (age, weight, etc.), and 6. online help (the help menus will serve two purposes: explain the various concepts and provide instructions for the various functions of the software).
MATERIALS AND METHODS Ultrasonic system Figure 1 shows the block diagram of the system, with photographs of the actual instruments. The system consists of a signal generator (HP 33120A Hewlett Packard, now Agilent technologies, Englewood, CO), a RF amplifier (LA 100-CE, Kalmus, Bothell, WA), a 3-D positioning system (MD-2, Arrick Robotics, Hurst, TX) and a 10-cm spherically shaped bowl transducer made from piezoelectric ceramic PZT4 (Etalon, Lebanon, IN). The transducer operates at 4 MHz, has a focal length of 10 cm and diameter of 4 cm. The transducer was rigidly mounted on the 3-D positioning system. The signal generator was controlled using the serial port of the PC and
Fig. 2. Main window of the software that controls the HIFU system.
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Fig. 4. Block diagram of the attenuation measurement system.
Fig. 3. Acoustical field measurement system.
Acoustical field The size of the focal region produced by this transducer was obtained by mapping the acoustic pressure field with a needle hydrophone (Specialty Engineering Associates, San Jose, CA) having an active element of 1 mm in diameter. The transducer under test was driven by a pulser/receiver (Panametrics 5050R, Waltham, MA). The hydrophone was connected to the receiver input of the pulser/receiver. The output of the pulser/receiver was connected through an A/D card (CS1250, A/D 12 bit, 50-MHz, Gage, Lachine, Quebec, Canada) to the PC for signal processing. The transducer was moved automatically by the same robotic system as described in Fig. 1. The block diagram of the system that measures the acoustical field of ultrasonic transducers using the actual photographs of the instruments is shown in Fig. 3. Raum and O’Brien (1997) describe in more details the above principles used. The full width at half-maximum intensity (D) of the beam was estimated from the measured acoustical field. The spatial average in situ intensity was estimated by using the equation: I SAL ⫽ 0.87P/D 2,
Attenuation system The system include the signal generator, the A/D acquisition card and two identical plane circular transducers (Etalon) operating at 4 MHz. The diameter of each transducer was 10 mm. Figure 4 shows an illustrated block diagram of the system using the actual photographs of the instruments. One transducer was connected to the signal generator and functioned as the transmitter (T1). The transmitter was submerged in a small container that was filled with degassed water. The tissue under measurement was placed inside the small container. In the other side of the container, a receiving transducer, T2, was placed. The output of the receiver was connected to the A/D acquisition card. The temperature of the water bath was controlled by circulating thermally regulated water through a heating coil. A 50-m diameter T-type copper-constantan thermocouple (Physitemp, Clifton, NJ) was inserted in the tissue to measure the tissue temperature and then to estimate the thermal dose referenced at 43°C. The temperature was measured using an HP 7500 series B system and an HP 1326B multimeter. The thermal dose was calculated using the technique suggested by Sapareto and Dewey (1984). The technique uses numerical integration to calculate the time that would give an equivalent thermal dose at a reference temperature for different temperature profiles and is given by:
(1)
where P is the acoustic power of the transducer (Malcolm and ter Haar 1996). The total power delivered by the transducer was measured before the beginning of each experiment with an ultrasound power meter (model UPM-DT-100N, Ohmic Instruments, Easton, MD). The estimation of the spatial average intensity is strongly affected by the measurement of the full-width at half-maximum intensity (w). The maximum error for the intensity measurement was estimated to be 5%.
冘R
t⫽final
t 43 ⫽
共43⫺Tt兲⌬t
(2)
t⫽0
where t43 is the equivalent time at 43°C (thermal dose), and Tt is the average temperature during the interval ⌬t. The default value of R equal to 0.25 was chosen for temperatures smaller than 43°C and a value equal to 0.5 for temperatures higher than 43°C (Sapareto and Dewey 1984). Varying the time that the sample was immersed in the temperature-controlled water bath varied the thermal dose. The bath temperature was set at 46°C and a 100-
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min dose at 43°C was delivered in 11.1 min. Thus, the dose rate was 9 min of the dose at 43°C/min. The value of attenuation (which gives a very good indication of absorption) was measured using the transmission and reception method of Kossoff et al. (1973). Initially, the signal through the water VW was measured. Then, the signal VT was measured with the presence of a tissue sample. The following equation gives the value of attenuation ␣ in dB/cm:
␣ ⫽ 20 log 共V w/V T兲/d,
(3)
where d is the sample thickness measured in cm. The attenuation of kidney cortex, medulla, muscle and fat was measured from excised pig tissues. In vitro experiments The tissue under ablation was placed in the water tank of the system shown in Fig. 1. Boiling the water to 100°C degassed the water. The tissue was placed on top of an absorbing material (rubber pad) to shield adjacent tissue from stray radiation from the bottom of the plastic water tank. The transducer was placed on the 3-D robotic arm and was immersed in the water tank; thus, providing good acoustical coupling between tissue and transducer. Any bubbles that may have collected under the face of the transducer were removed to eliminate any reflections. The sample size was typically 60 mm ⫻ 40 mm with thickness of around 30 mm. The sample was gently massaged to remove any trapped air bubbles and the experiment was initiated 10 to 20 min after placing the sample in the tank. Although it is difficult to assess if this procedure was sufficient to remove very small air bubbles, in several samples where this procedure was not followed the presence of air bubbles was easily detected from the resulting lesion appearance. With a single kidney, it would have been impossible to test the system for producing deep-seated lesions. To emulate the case of ablating a deep-seated target (⬎ 4 cm), as in the case of humans, two kidneys were used with one placed on top of the other. The power used during the experiments was kept to a level that did not cause boiling of tissue water, which would lead to the formation of vapour bubbles and prevent the production of pure thermal lesions, whose size can be controlled. In vivo experiments The kidney was determined using X rays and using palpation performed by an experienced veterinarian. The coupling arrangement during animal experiments (Fig. 5) was slightly different from the method used in the in vitro experiments. A rectangular container 8 cm ⫻ 8 cm
Fig. 5. Experimental set-up for the in vivo experiments.
⫻ 10 cm was made out of plastic. A special holder was designed to hold the container in the laboratory table. The top and the bottom sides of the container were left open so that a low attenuation bag filled with degassed water could be inserted. The transducer was placed inside the water-filled bag (on the top side). The lower side was placed on the animal that lay on a surgical table. The kidney of adult pigs was selected as the target for ablation, because its size is similar to that of the human. The pig was shaved before the experiment. Between the shaved animal and the lower side of the bag, US gel was placed to ensure that no localised heating at that interface was produced due to reflections. The initial anesthetic was administered using an IM injection of a cocktail of xylazine (3 mg/kg), ketamine (15 mg/kg) and atropine (0.05 mg/kg). Then, an IV drip to the dorsal auricular vein was followed using ketamine (1 mg/mL), xylazine (1 mg/mL), guaifenesin (50 mg/ mL) and 5 % dexrose with a rate of 2.2 mL/kg/h. At euthaniasia, the animal was administered an overdose of anaesthetic. The animal protocols were approved by the national agency granting permission for animal experiments (Ministry of Agriculture). At the end of the experiment, the ablated tissue was serially sectioned, to acquire photographs showing the necrosis. Simulation model The lesion size was predicted using a model that uses the thermal dose concept described by Damianou and Hynynen (1993, 1994). The kidney attenuation used in the simulation model depended on the amount of the accumulated thermal dose. It is known that the thermal dose threshold of necrosis referenced at 43°C for kidney is 50 min (Borelli et al. 1990). For thermal doses below 50 min at 43°C, the attenuation at body temperature was used. For thermal doses higher than 50 min at 43°C, the attenuation of necrotic tissue was used (measured using the attenuation method described in this paper). Attenuation a affects the power deposition density q, which is given by the following equation:
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Fig. 6. Flow chart of the simulation model that includes variable attenuation.
q ⫽ 2aIe ⫺2ax,
(4)
where I is the intensity and x is the depth in tissue. For short pulses, the power deposition density is the main factor elevating the tissue temperature. Figure 6 shows the flow chart of the simulation model that includes the effect of variable attenuation during heating. In the flow chart, z denotes axial distance, r radial distance and t time.
RESULTS Figure 7 shows the attenuation of porcine kidney (cortex and medulla), muscle and fat as a function of thermal dose. Note that the attenuation increases rapidly during necrosis and, eventually, stabilises after total necrosis has been produced. To eliminate the effect of blood perfusion, the pulse duration of 5 s was used (Billard et al. 1990) as the standard pulse duration to be used throughout the in vitro
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Fig. 7. Attenuation vs. thermal dose for (a) kidney cortex, (b) kidney medulla, (c) muscle, and (d) fat.
and in vivo experiments. Pulse duration of 1 s would be desired, because treatment time is decreased; however, at this pulse duration, the width of lesions is small. Table 1 shows the simulated length and width using the pure thermal model and using the model that accounts for the variable attenuation. The transducer parameters used were: frequency ⫽ 4 MHz, radius of curvature ⫽ 10 cm and transducer diameter ⫽ 4 cm. The focal in situ spatial-average intensity of 2000 W/cm was applied for 5 s at a focal depth of 15 mm. The zero perfusion represents the case of in vitro kidney, whereas the case of in vivo is modelled by using the perfusion of 70 kg/m3-s (Lehmann 1982). Figure 8 shows the length of a lesion created in excised porcine kidney using the in situ spatial-average intensity of 2000 W/cm for 5 s. The lesion produced was placed entirely in the cortex of the kidney. The length was measured along the transducer central axis, whereas the width was measured perpendicularly to the transducer central axis. The lesion length is about 20 mm and the width is about 3 mm. In a certain row of lesions, only
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Fig. 8. Thermal lesion placed in the cortex of pig kidney in vitro.
one dimension of the lesion (length or width) was measured because the sample was cut either along the length or the width. Figure 9 shows the lesion width with 2000 W/cm applied for 5 s. The lesion measured was between 2.8 to 3.4 mm. The lesions appeared to be repeatable for a given acoustic exposure. Some differences that we observed are attributed to the variation of focal depth arising from the kidney curvature. In most of the in vitro experiments, the propagation inside the medulla was difficult due to the air spaces in the medulla. Normally, in an in vivo kidney, there is blood inside the medulla, making the medulla a possible target for ultrasound. Figure 10 shows lesions that propagated inside the medulla region. Thus, in the case that there are no bubbles inside the medulla, ablation is feasible.
Table 1. Simulated length and width for the pure thermal model and the model of variable attenuation Pure thermal model, perfusion (kg/m3 -s)
Variable attenuation model, perfusion (kg/m3 -s)
Quantity
0 (in vitro)
70 (in vivo)
0 (in vitro)
70 (in vivo)
Length (mm) Width (mm)
17.2 2.6
16.7 2.4
20.6 3.3
19.7 3.1
The transducer parameters used are frequency ⫽ 4 MHz, radius of curvature ⫽ 10 cm and transducer diameter ⫽ 4 cm. The focal intensity of 2000 W/cm2 was applied for 5 s at a focal depth of 15 mm.
Fig. 9. Lesions created in pig kidney demonstrating the excellent repeatability.
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Fig. 10. Lesions propagating in the medulla of pig kidney in vitro. Fig. 12. Lesions in peritoneal fat.
Creation of lesions through a fat layer is possible if there are no bubbles within the fat-kidney interface. Figure 11 shows lesion length through a fat layer with a 2000 W/cm ablation for 5 s. Note that, compared with a similar ablation in kidney tissue, the lesion length is much smaller with the presence of a fat layer. This is attributed to the fact that, sometimes, bubbles exist in the kidney-fat interface. Figure 12 shows a 3 ⫻ 3 matrix of lesions in fat, displaying the lesion width. Lesions appeared to be light brown inside the white fat tissue. Figure 13 shows a lesion created by moving the transducer in a patterned movement (square grid of 8 ⫻ 8 with 3-mm step) in both directions. The intended target was the round shape in red colour. The power used was 2000 W/cm for 5 s. A delay of 10 s was used between the pulses, to eliminate the near-field heating (Damianou and Hynynen 1993). Based on the width obtained in Fig. 8, a 3-mm step will produce overlapping lesions. Necrosis with a 3 to 5 mm margin around the target is desired. For
Fig. 11. Lesions created in pig kidney in vitro through fat layer showing poor penetration.
the in vitro case, a larger step can be used, because ultrasound penetration is excellent and there is no blood flow. Based on the simulation study for the in vivo case, a 3 mm step must be used. Figure 14 shows the corresponding length of the large lesion. The lesion was extended up to the medulla. Figure 15 shows a large lesion created through fat using transducer movement in a grid pattern. Presumably, there were no air spaces between fat and kidney and, therefore, the intended target was completely covered with necrosis. Figure 16 shows a lesion placed 4 cm deep in kidney tissue. The thick kidney tissue was achieved by using two kidneys, with one kidney placed on top of the other. The importance in this result is that the lesion propagates through the medulla region of the top kidney. This animal model, which was used extensively, proved to be a very good verification tool for thick kidney heating.
Fig. 13. Large lesion in pig kidney in vitro by moving the transducer in a grid formation (8 ⫻ 8) (top view).
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Fig. 14. Large lesion in pig kidney in vitro by moving the transducer in a grid pattern (8 ⫻ 8) (side view).
The presence of cavitation was verified by using a wide-band receiver placed perpendicular to the insonation beam. The threshold of cavitation was found by detecting the in situ spatial average intensity, which produces a signal in the receiver indicative of cavitation. Thus, single lesions were created for this purpose by varying the in situ spatial average intensity from 2000 to 3500 W/cm2. For the given transducer and for a pulse duration of 5 s, it was observed that cavitation took place for in situ spatial-average intensity levels greater than 3200 W/cm2 (error was ⫾ 100 W/cm2). Figure 17 shows lesion length using 5-s pulse and 4000 W/cm2 in situ spatial-average intensity in kidney tissue in vitro. Note that, at this acoustic level, cavitational effect occurred, which was verified by the cavitational receiver. Tissue examination revealed gross tissue disruption, verifying the event of acoustic cavitation. In the in vivo study, a total of six animals experiments have been conducted. Both kidneys of each animal have been ablated; thus, ablation was performed in 12
Fig. 15. Large lesion in pig kidney in vitro through fat layer using a grid pattern (side view).
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Fig. 16. Deep-seated lesions in pig kidney in vitro using two kidneys placed on top of each other.
kidneys. Successful necrosis was achieved in all 12 kidneys. All the animals in this series tolerated the procedure well and remained in good condition throughout the period of the experiment. Figure 18 shows the length of an 8 ⫻ 8 matrix of lesions created using 2000 W/cm2 for 5 s in pig kidney in vivo. The transducer was moved with a step of 3 mm. The lesions were usually firm and white and, usually, no adjacent hemorrhagic zone was created. Only in one kidney was haemorrhage observed. Despite the thick fat layer, necrosis of the renal cortex was reliable. The lesion length in the in vivo case appeared to be smaller (about 18 mm compared to 20 mm in the in vitro case). Single lesions produced in the in vivo kidney revealed a lesion width of about 3 mm for 2000 W/cm2 at 5 s. Thus, the step of 3 mm produced overlapped lesions covering a large target volume. DISCUSSION The trend of attenuation with dose shows a similar trend to the results published by Damianou et al. (1997).
Fig. 17. Large lesion in pig kidney in vivo by moving the transducer in grid pattern showing cavitational activity.
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Table 2. Summary of simplistic protocol for treating renal tissues Pulse duration Delay between US firings Intensity for thermal lesions Spatial step Treatment time for a 20 mm ⫻ 20 mm ⫻ 20 mm target
5s 10 s ⬍ 3200 W/cm2 3 mm 8 ⫻ 8 (10 ⫹ 5) ⫽ 960 s ⫽ 16 min
Transducer parameters are frequency ⫽ 4 MHz, diameter ⫽ 4 cm and radius of curvuture ⫽ 10 cm. Focal depth ⫽ 15 mm.
Fig. 18. Large lesion in pig kidney in vivo (through fat layer) by moving the transducer in grid formation.
The normal kidney tissue included both medulla and renal cortex, for which we found no significant difference in terms of ultrasonic attenuation. The research on muscle and fat is justified because these two tissues usually surround the kidney. The trend of attenuation with dose in all tissues was similar. There was a rapid increase of attenuation with thermal dose and, when necrosis occurred, attenuation stabilised. At that point, the attenuation was roughly double the attenuation at room temperature. Creation of thermal lesions in the cortex is very consistent and repeatable. For the given transducer (f ⫽ 4 MHz, r ⫽ 10 cm and d ⫽ 4 cm) and the given exposure (in situ spatial-average intensity ⫽ 2000 W/cm2, pulse duration 5 s and 15 mm focal depth), the lesion length was around 20 mm and the lesion width was about 3 mm. The lesion length and width of Fig. 8, which was produced in a tissue in vitro (i.e., perfusion rate is 0) is best estimated using the variable attenuation model. The experimental length was 20 mm and the width was 3 mm. The simulated length was 20.6 mm and the width 3.3 mm, using the variable attenuation model. Note that, at perfusion of 70 kg/m3-s (in vivo case), the simulated lesion length and width is decreased compared with the in vitro case. The lesion length and width was generally bigger in the variable attenuation model, because the increased absorption raised the power density in the tissue and, thus, the lesion size increased. This is true for this type of transducer and focal depth. From the equation of absorbed power density, it is clear that the effect of attenuation depends on the transducer frequency and focal depth. Lesions were created all the way to the medulla, provided that there were no air spaces in the medulla. Air spaces are created because of the absence of blood that normally flows in this region. However, ablation of the cortex tissue is of primary concern because the renal
carcinoma (most important renal cancer) grows inside the cortex and then extends to the peritoneal fat. Lesions can be created through fat layers, provided that there are no air spaces between the fat and kidney interface. The lesion size when the beam goes through a fat layer is decreased, because of the high attenuation of fat, which acts likes an acoustic barrier. Deep-seated lesions in the kidney can be created by using two kidneys on top of each other. With a 4-MHz transducer, 5-s pulses and in situ spatial-average intensity levels below about 3200 W/cm2, the lesions created are based solely on thermal effects. Data from Hynynen (1991) suggest that the threshold spatial peak intensity at 4 MHz is about 6000 W/cm2 (extrapolated because data up to 2 MHz is provided) or 3333 W/cm2 (in situ spatial-average intensity). Thus, with the in situ spatial-average intensity of 4000 W/cm2 used in some experiments, cavitation effects were not avoided. In the cavitation range, lesions have irregular boundaries, as opposed to the case of operating below the cavitation range, where the lesions have smooth boundaries, exhibit less mechanical damage to the tissues and always occur at the centre of the focal region. Lesions were smaller in the in vivo experiments compared with lesions created in dead tissue for the same exposure. The reduced size of the in vivo exposures is attributed to the blood flow and possible reduction of the energy due to attenuation at various interfaces, including the skin. Table 2 summarises the treatment protocol for a shallow target in kidney (about 15 mm deep) for a transducer operating at 4 MHz (a frequency with high cavitational threshold). The protocol is simplistic in the sense that integer numbers are used. The determination of these numbers gives a good indication of how to derive a treatment protocol in humans. In the case of humans, the main difference from this animal model is certainly the location of the tumours, which are deeper in the body. Thus, to achieve satisfactory heating in the human case, probably a more powerful transducer needs to be used. The treatment time for a 20 mm ⫻ 20 mm ⫻ 20 mm target is around 16 min. This treatment time can
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be reduced by using, for example, spatial average intensity of 3000 W/cm2 (spacing between transducer movement is around 3.8 mm) and, therefore, treatment time is 6 ⫻ 6 ⫻ 15 ⫽ 735 s ⫽ 9 min (a reduction of 44%). This is just one example of reducing the treatment time. Acknowledgements—This work was supported by the Research Promotion Foundation (RPF) of Cyprus (contract 25/99).
REFERENCES Bihrle R, Foster RS, Sanghvi NT, Donohue JP, Hood PJ. High-intensity focused ultrasound for the treatment of benign prostatic hyperplasia: Early United States clinical experience. J Urol 1994;151(5): 1271–1275. Billard BE, Hynynen K, Roemer RB. Effects of physical parameters on high temperature ultrasound hyperthermia. Ultrasound Med Biol 1990;16:409–420. Borrelli M, Thompson L, Cain C, Dewey W. Time temperature analysis of cell killing of BHK cells heated at temperatures in the range of 43.5°C to 57 °C. J Radiat Oncol Biol Phys 1990;19:389–399. Chapelon JY, Margonari J, Theillere Y, et al. A. Effects of high-energy focused ultrasound on kidney tissue in the rat and the dog. J Eur Urol 1992a;22(2):147–152. Chapelon JY, Margonari J, Vernier F, et al. A. In vivo effects of high-intensity ultrasound on prostatic adenocarcinoma Dunning R3327. Cancer Res 1992b;52(22):6353–6357. Chapelon JY, Ribault M, Vernier F, Souchon R, et al. Treatment of localised prostate cancer with transrectal high intensity focused ultrasound. Eur J Ultrasound 1999;9(1):31–38. Cline HE, Schenck JF, Hynynen K, et al. MR-guided focused ultrasound surgery. J Comput Assist Tomogr 1992;16:956–965. Damianou C, Hynynen K. Focal spacing and near-field heating during pulsed high temperature ultrasound therapy. Ultrasound Med Biol 1993;19(9):777–787. Damianou C, Hynynen K. The effect of various physical parameters on the size and shape of necrosed tissue volume during ultrasound surgery. J Acoust Soc Am 1994;95(3):1641–1649. Damianou C, Sanghvi N, Fry F, Maass R. Dependence of ultrasonic attenuation and absorption in dog soft tissues on temperature and thermal dose. J Acoust Soc Am 1997;102(2):628–634. Fry W, Mosberg W, Barnard J, Fry F. Production of focal destructive lesions in the central nervous system with ultrasound. J Neurosurg 1954;11:471–478. Hynynen K. The threshold for thermally significant caviation in dog’s thigh muscle in vivo. Ultrasound Med Biol 1991;17(2):157–169.
Volume 29, Number 9, 2003 Hynynen K, Damianou CA, Colucci V, et al. MR monitoring of focused ultrasonic surgery of renal cortex: Experimental and simulation studies. J Magn Reson Imaging 1995;5(3):259–266. Hynynen K, Damianou C, Darkazanli A, Unger E, Schenck JF. The feasibility of using MRI to monitor and guide noninvasive ultrasound surgery (letter). Ultrasound Med Biol 1993a;19(1):91– 92. Hynynen K, Darkazanli A, Unger E, Schenck JF. MRI-guided noninvasive ultrasound surgery. Med Phys 1993b;20(1):107–115. Kossoff G, Kelly-Fry E, Jellins J. Average velocity of ultrasound in the human female breast. J Acoust Soc Am 1973;53(6):1730–1736. Lehmann J. Therapeutic heat and cold. Baltimore, MD: Williams and Wilkins, 1982, 38. Lele PP. A simple method for production of trackless focal lesions with focused ultrasound. J Physiol 1962;160:494–512. Linke C, Carstensen E, Frizzell L, Elbadiwi A, Fridd C. Localized tissue destruction by high-intensity focused ultrasound. Arch Surg 1973;107:887–891. Lizzi F, Coleman D, Driller J, et al. Ultrasonic hyperthermia for ophthalmic therapy. IEEE Trans Son Ultrason 1984;SU-31(5):473– 481. Lynn JG, Zwemer RL, Chick AJ, Miller AE. A new method for the generation and use of focused ultrasound in experimental biology. J Gen Phyciology 1942;26:179–193. Maass-Moreno R, Damianou CA, Sanghvi NT. Noninvasive temperature estimation in tissue via ultrasound echo-shifts: Part II. In vitro study. J Acoust Soc Am 1996;100(4 Pt 1):2522–2530. Malcolm AL, ter Haar GR. Ablation of tissue volumes using high intensity focused ultrasound. Ultrasound Med Biol 1996;22(5): 659–669. Raum K, O’Brien D. Pulse-echo field distribution measurement technique for high-frequency ultrasound sources. IEEE Trans Ultrason Ferroelec Freq Control 1997;44(4):810–815. Sanghvi N, Fry F, Foster R, et al. System design and considerations for high intensity focused ultrasound device for the treatment of tissue in vivo. Med Biol Eng Comp 1991;29:748. Sapareto S, Dewey W. Thermal dose determination in cancer therapy. Int J Radiat Oncol Biol Phys 1984;10:787–800. Seip R, Ebbini E. Non-invasive estimation of tissue temperature response to heating using diagnostic ultrasound. IEEE Trans Biomed Eng 1995;42(8):828–839. ter Haar G, Sinnett D, Rivens I. High intensity focused ultrasound—A surgical technique for the teatment of discrete liver tumors. Phys Med Biol 1989;34(11):1743–1750. Vykhodtseva NI, Hynynen K, Damianou C. Pulse duration and peak intensity during focused ultrasound surgery: Theoretical and experimental effects in rabbit brain in vivo. Ultrasound Med Biol 1994; 20(9):987–1000.
doi:10.1016/S0301-5629(03)00981-5
● Original Contribution IN VITRO AND IN VIVO ABLATION OF PORCINE RENAL TISSUES USING HIGH-INTENSITY FOCUSED ULTRASOUND CHRISTAKIS DAMIANOU Frederick Institute of Technology, Limassol, Cyprus (Received 13 November 2002; revised 25 April 2003; in final form 1 May 2003)
Abstract—The aim of this paper is to present issues regarding the thermal ablation of porcine renal tissues in vitro and in vivo using high-intensity focused ultrasound (HIFU). Production of lesions in the cortex in vitro is consistent, whereas lesions in the medulla are created whenever there are no air spaces in the medulla. Typically, the lesion length at 2000 W/cm and 5-s pulse duration is around 20 mm and the corresponding width around 3 mm. Lesioning of a large volume was achieved by moving the transducer in a grid formation. Lesioning through a fat layer is possible provided that there are no air spaces between the fat and kidney interface. It was found that, above 3200 W/cm with 5-s pulse duration at 4 MHz, cavitation activity occurred in most of the lesions created. (E-mail: [email protected]) © 2003 World Federation for Ultrasound in Medicine & Biology. Key Words: Ultrasound, Kidney, Attenuation, Simulation, Lesion, Cortex, Cavitation.
imaging. Ultrasonic imaging is the simplest and most inexpensive method; however, it has poor contrast between soft tissues. On the other hand, magnetic resonance imaging (MRI) offers superior contrast, but it is more expensive. Several trials have been conducted in the area of ultrasonic imaging (for example, Seip and Ebbini 1995; Maass-Moreno et al. 1996) and in the area of MRI (for example, Cline et al. 1992; Hynynen et al. 1993a, 1993b), thus enhancing the potential of HIFU. The main goal of HIFU is to maintain a temperature between 50 to 100°C for a few s (typically ⬍ 10 s), in order to cause tissue necrosis. Typically, focal peak intensity between 1000 to 10,000 W/cm2 is used with pulse duration between 1 to 10 s and a frequency of 1 to 5 MHz. This paper reports additional experience in the ablation of kidney tissue. For applications such as the treatment of benign prostatic hyperplasia (BPH), it is sufficient to destroy as much tissue as possible for the purpose of tissue debulking. For applications in oncology, destruction of all viable tumor cells is required and, therefore, protocols in this area must be very accurate and reliable. In this work, methods suggested by Malcolm and ter Haar (1996) were used to produce complete, reliable and consistent ablation of renal tissues (creation of a contiguous array of touching lesions of thermal origin, avoiding boiling and cavitation, and use of cooling to avoid merging of lesions in front of the focus).
INTRODUCTION High-intensity focused ultrasound (HIFU) is a noninvasive procedure for heating tumours without affecting the healthy tissue surrounding the tumour. Therefore, application of HIFU is being investigated as an alternative to standard surgical techniques. Although the idea of using HIFU was proposed in the middle of this century by Lynn et al. (1942), its maximum potential for clinical use has been established only recently, due to the developments of sophisticated systems (for example, Chapelon et al. 1992b; Hynynen et al. 1993a and Birhle et al. 1994). HIFU has been proven over nearly 60 years to be an effective and efficient method for ablating soft tissue, but its commercial success is still under trial. HIFU was explored in almost every tissue that is accessible by ultrasound. The following represent some examples of some applications explored: eye (Lizzi et al. 1984), prostate (Sanghvi et al. 1991; Chapelon et al. 1999), liver (ter Haar et al. 1989), brain (Fry et al. 1954; Lele 1962; Vydkotseva et al. 1994) and kidney (Linke et al. 1973; Chapelon et al. 1992a; Hynynen et al. 1995). Recently, the technology of HIFU systems has improved because the ultrasonic therapy can be guided by
Address correspondence to: Dr. C. Damianou, Frederick Institute of Technology (FIT), 18, Mariou Agathangelou, 3080, Limassol, Cyprus. E-mail: [email protected] 1321
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Experiments were performed using a generic HIFU system that includes a signal generator, a radiofrequency (RF) amplifier, a 3-D robotic system, a transducer and a personal computer (PC) that controls the entire system. The transducer of 4-cm diameter was spherically focused, operating at 4 MHz. Pulse duration of 5 s was used in all the experiments to minimise effects of blood perfusion. Results of HIFU ablation of porcine kidney in vitro and in vivo are presented. Single lesions in the cortex and medulla with or without a fat layer were created. The ultimate goal was to create large lesions, which was accomplished by moving the transducer in patterned schemes. The experimental results are compared with the results of a simulation model that includes the effect of rapid increase of attenuation during the transition of a tissue from healthy to necrotic. Thus, the kidney attenuations of porcine cortex, medulla, muscle and fat were measured as a function of thermal dose. The attenuation was measured using a system that includes two low-intensity transducers, a signal generator, a data-acquisition card and a PC. Attenuation includes absorption, scattering and reflection but, when minimising scattering and reflection, attenuation will reflect mostly losses due to absorption. Scattering and reflection was minimised by using a homogeneous sample with a minimum number of interfaces or air ducts, and by removing acoustically troublesome gas bubbles using degassing techniques. Using these precautions in the study by Damianou et al. (1997), it was found that the value of absorption was very close to the value of attenuation, indicating that the contribution of reflection and scattering was minimised. Absorption is an acoustical property of tissue with wide variation from tissue to tissue and represents the rate at which energy in tissue is converted to heat. For short pulses (⬍ 5 s), it is one of the primary factors contributing to the temperature elevation in tissue because the effect of blood flow is minimised.
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Fig. 1. Block diagram of the high-intensity focused ultrasound system.
the 3-D robotic system was controlled using the parallel port. A user-friendly program written in Visual Basic has been developed to control the system. Figure 2 shows the main window of the software. The software serves six main tasks through various windows or menus: 1. imaging (in this window US or MRI images can be displayed), 2. transducer movement (the user may move the robotic arm in a specific direction or customise the automatic movement of the robotic arm in any formation by specifying the pattern, the step and the number of steps), 3. messaging (starting time, treatment time left, etc.), 4. telematics (a physician in a remote site can monitor the treatment protocol and can suggest ideas during the treatment, 5. patient data (age, weight, etc.), and 6. online help (the help menus will serve two purposes: explain the various concepts and provide instructions for the various functions of the software).
MATERIALS AND METHODS Ultrasonic system Figure 1 shows the block diagram of the system, with photographs of the actual instruments. The system consists of a signal generator (HP 33120A Hewlett Packard, now Agilent technologies, Englewood, CO), a RF amplifier (LA 100-CE, Kalmus, Bothell, WA), a 3-D positioning system (MD-2, Arrick Robotics, Hurst, TX) and a 10-cm spherically shaped bowl transducer made from piezoelectric ceramic PZT4 (Etalon, Lebanon, IN). The transducer operates at 4 MHz, has a focal length of 10 cm and diameter of 4 cm. The transducer was rigidly mounted on the 3-D positioning system. The signal generator was controlled using the serial port of the PC and
Fig. 2. Main window of the software that controls the HIFU system.
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Fig. 4. Block diagram of the attenuation measurement system.
Fig. 3. Acoustical field measurement system.
Acoustical field The size of the focal region produced by this transducer was obtained by mapping the acoustic pressure field with a needle hydrophone (Specialty Engineering Associates, San Jose, CA) having an active element of 1 mm in diameter. The transducer under test was driven by a pulser/receiver (Panametrics 5050R, Waltham, MA). The hydrophone was connected to the receiver input of the pulser/receiver. The output of the pulser/receiver was connected through an A/D card (CS1250, A/D 12 bit, 50-MHz, Gage, Lachine, Quebec, Canada) to the PC for signal processing. The transducer was moved automatically by the same robotic system as described in Fig. 1. The block diagram of the system that measures the acoustical field of ultrasonic transducers using the actual photographs of the instruments is shown in Fig. 3. Raum and O’Brien (1997) describe in more details the above principles used. The full width at half-maximum intensity (D) of the beam was estimated from the measured acoustical field. The spatial average in situ intensity was estimated by using the equation: I SAL ⫽ 0.87P/D 2,
Attenuation system The system include the signal generator, the A/D acquisition card and two identical plane circular transducers (Etalon) operating at 4 MHz. The diameter of each transducer was 10 mm. Figure 4 shows an illustrated block diagram of the system using the actual photographs of the instruments. One transducer was connected to the signal generator and functioned as the transmitter (T1). The transmitter was submerged in a small container that was filled with degassed water. The tissue under measurement was placed inside the small container. In the other side of the container, a receiving transducer, T2, was placed. The output of the receiver was connected to the A/D acquisition card. The temperature of the water bath was controlled by circulating thermally regulated water through a heating coil. A 50-m diameter T-type copper-constantan thermocouple (Physitemp, Clifton, NJ) was inserted in the tissue to measure the tissue temperature and then to estimate the thermal dose referenced at 43°C. The temperature was measured using an HP 7500 series B system and an HP 1326B multimeter. The thermal dose was calculated using the technique suggested by Sapareto and Dewey (1984). The technique uses numerical integration to calculate the time that would give an equivalent thermal dose at a reference temperature for different temperature profiles and is given by:
(1)
where P is the acoustic power of the transducer (Malcolm and ter Haar 1996). The total power delivered by the transducer was measured before the beginning of each experiment with an ultrasound power meter (model UPM-DT-100N, Ohmic Instruments, Easton, MD). The estimation of the spatial average intensity is strongly affected by the measurement of the full-width at half-maximum intensity (w). The maximum error for the intensity measurement was estimated to be 5%.
冘R
t⫽final
t 43 ⫽
共43⫺Tt兲⌬t
(2)
t⫽0
where t43 is the equivalent time at 43°C (thermal dose), and Tt is the average temperature during the interval ⌬t. The default value of R equal to 0.25 was chosen for temperatures smaller than 43°C and a value equal to 0.5 for temperatures higher than 43°C (Sapareto and Dewey 1984). Varying the time that the sample was immersed in the temperature-controlled water bath varied the thermal dose. The bath temperature was set at 46°C and a 100-
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min dose at 43°C was delivered in 11.1 min. Thus, the dose rate was 9 min of the dose at 43°C/min. The value of attenuation (which gives a very good indication of absorption) was measured using the transmission and reception method of Kossoff et al. (1973). Initially, the signal through the water VW was measured. Then, the signal VT was measured with the presence of a tissue sample. The following equation gives the value of attenuation ␣ in dB/cm:
␣ ⫽ 20 log 共V w/V T兲/d,
(3)
where d is the sample thickness measured in cm. The attenuation of kidney cortex, medulla, muscle and fat was measured from excised pig tissues. In vitro experiments The tissue under ablation was placed in the water tank of the system shown in Fig. 1. Boiling the water to 100°C degassed the water. The tissue was placed on top of an absorbing material (rubber pad) to shield adjacent tissue from stray radiation from the bottom of the plastic water tank. The transducer was placed on the 3-D robotic arm and was immersed in the water tank; thus, providing good acoustical coupling between tissue and transducer. Any bubbles that may have collected under the face of the transducer were removed to eliminate any reflections. The sample size was typically 60 mm ⫻ 40 mm with thickness of around 30 mm. The sample was gently massaged to remove any trapped air bubbles and the experiment was initiated 10 to 20 min after placing the sample in the tank. Although it is difficult to assess if this procedure was sufficient to remove very small air bubbles, in several samples where this procedure was not followed the presence of air bubbles was easily detected from the resulting lesion appearance. With a single kidney, it would have been impossible to test the system for producing deep-seated lesions. To emulate the case of ablating a deep-seated target (⬎ 4 cm), as in the case of humans, two kidneys were used with one placed on top of the other. The power used during the experiments was kept to a level that did not cause boiling of tissue water, which would lead to the formation of vapour bubbles and prevent the production of pure thermal lesions, whose size can be controlled. In vivo experiments The kidney was determined using X rays and using palpation performed by an experienced veterinarian. The coupling arrangement during animal experiments (Fig. 5) was slightly different from the method used in the in vitro experiments. A rectangular container 8 cm ⫻ 8 cm
Fig. 5. Experimental set-up for the in vivo experiments.
⫻ 10 cm was made out of plastic. A special holder was designed to hold the container in the laboratory table. The top and the bottom sides of the container were left open so that a low attenuation bag filled with degassed water could be inserted. The transducer was placed inside the water-filled bag (on the top side). The lower side was placed on the animal that lay on a surgical table. The kidney of adult pigs was selected as the target for ablation, because its size is similar to that of the human. The pig was shaved before the experiment. Between the shaved animal and the lower side of the bag, US gel was placed to ensure that no localised heating at that interface was produced due to reflections. The initial anesthetic was administered using an IM injection of a cocktail of xylazine (3 mg/kg), ketamine (15 mg/kg) and atropine (0.05 mg/kg). Then, an IV drip to the dorsal auricular vein was followed using ketamine (1 mg/mL), xylazine (1 mg/mL), guaifenesin (50 mg/ mL) and 5 % dexrose with a rate of 2.2 mL/kg/h. At euthaniasia, the animal was administered an overdose of anaesthetic. The animal protocols were approved by the national agency granting permission for animal experiments (Ministry of Agriculture). At the end of the experiment, the ablated tissue was serially sectioned, to acquire photographs showing the necrosis. Simulation model The lesion size was predicted using a model that uses the thermal dose concept described by Damianou and Hynynen (1993, 1994). The kidney attenuation used in the simulation model depended on the amount of the accumulated thermal dose. It is known that the thermal dose threshold of necrosis referenced at 43°C for kidney is 50 min (Borelli et al. 1990). For thermal doses below 50 min at 43°C, the attenuation at body temperature was used. For thermal doses higher than 50 min at 43°C, the attenuation of necrotic tissue was used (measured using the attenuation method described in this paper). Attenuation a affects the power deposition density q, which is given by the following equation:
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Fig. 6. Flow chart of the simulation model that includes variable attenuation.
q ⫽ 2aIe ⫺2ax,
(4)
where I is the intensity and x is the depth in tissue. For short pulses, the power deposition density is the main factor elevating the tissue temperature. Figure 6 shows the flow chart of the simulation model that includes the effect of variable attenuation during heating. In the flow chart, z denotes axial distance, r radial distance and t time.
RESULTS Figure 7 shows the attenuation of porcine kidney (cortex and medulla), muscle and fat as a function of thermal dose. Note that the attenuation increases rapidly during necrosis and, eventually, stabilises after total necrosis has been produced. To eliminate the effect of blood perfusion, the pulse duration of 5 s was used (Billard et al. 1990) as the standard pulse duration to be used throughout the in vitro
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Fig. 7. Attenuation vs. thermal dose for (a) kidney cortex, (b) kidney medulla, (c) muscle, and (d) fat.
and in vivo experiments. Pulse duration of 1 s would be desired, because treatment time is decreased; however, at this pulse duration, the width of lesions is small. Table 1 shows the simulated length and width using the pure thermal model and using the model that accounts for the variable attenuation. The transducer parameters used were: frequency ⫽ 4 MHz, radius of curvature ⫽ 10 cm and transducer diameter ⫽ 4 cm. The focal in situ spatial-average intensity of 2000 W/cm was applied for 5 s at a focal depth of 15 mm. The zero perfusion represents the case of in vitro kidney, whereas the case of in vivo is modelled by using the perfusion of 70 kg/m3-s (Lehmann 1982). Figure 8 shows the length of a lesion created in excised porcine kidney using the in situ spatial-average intensity of 2000 W/cm for 5 s. The lesion produced was placed entirely in the cortex of the kidney. The length was measured along the transducer central axis, whereas the width was measured perpendicularly to the transducer central axis. The lesion length is about 20 mm and the width is about 3 mm. In a certain row of lesions, only
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Fig. 8. Thermal lesion placed in the cortex of pig kidney in vitro.
one dimension of the lesion (length or width) was measured because the sample was cut either along the length or the width. Figure 9 shows the lesion width with 2000 W/cm applied for 5 s. The lesion measured was between 2.8 to 3.4 mm. The lesions appeared to be repeatable for a given acoustic exposure. Some differences that we observed are attributed to the variation of focal depth arising from the kidney curvature. In most of the in vitro experiments, the propagation inside the medulla was difficult due to the air spaces in the medulla. Normally, in an in vivo kidney, there is blood inside the medulla, making the medulla a possible target for ultrasound. Figure 10 shows lesions that propagated inside the medulla region. Thus, in the case that there are no bubbles inside the medulla, ablation is feasible.
Table 1. Simulated length and width for the pure thermal model and the model of variable attenuation Pure thermal model, perfusion (kg/m3 -s)
Variable attenuation model, perfusion (kg/m3 -s)
Quantity
0 (in vitro)
70 (in vivo)
0 (in vitro)
70 (in vivo)
Length (mm) Width (mm)
17.2 2.6
16.7 2.4
20.6 3.3
19.7 3.1
The transducer parameters used are frequency ⫽ 4 MHz, radius of curvature ⫽ 10 cm and transducer diameter ⫽ 4 cm. The focal intensity of 2000 W/cm2 was applied for 5 s at a focal depth of 15 mm.
Fig. 9. Lesions created in pig kidney demonstrating the excellent repeatability.
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Fig. 10. Lesions propagating in the medulla of pig kidney in vitro. Fig. 12. Lesions in peritoneal fat.
Creation of lesions through a fat layer is possible if there are no bubbles within the fat-kidney interface. Figure 11 shows lesion length through a fat layer with a 2000 W/cm ablation for 5 s. Note that, compared with a similar ablation in kidney tissue, the lesion length is much smaller with the presence of a fat layer. This is attributed to the fact that, sometimes, bubbles exist in the kidney-fat interface. Figure 12 shows a 3 ⫻ 3 matrix of lesions in fat, displaying the lesion width. Lesions appeared to be light brown inside the white fat tissue. Figure 13 shows a lesion created by moving the transducer in a patterned movement (square grid of 8 ⫻ 8 with 3-mm step) in both directions. The intended target was the round shape in red colour. The power used was 2000 W/cm for 5 s. A delay of 10 s was used between the pulses, to eliminate the near-field heating (Damianou and Hynynen 1993). Based on the width obtained in Fig. 8, a 3-mm step will produce overlapping lesions. Necrosis with a 3 to 5 mm margin around the target is desired. For
Fig. 11. Lesions created in pig kidney in vitro through fat layer showing poor penetration.
the in vitro case, a larger step can be used, because ultrasound penetration is excellent and there is no blood flow. Based on the simulation study for the in vivo case, a 3 mm step must be used. Figure 14 shows the corresponding length of the large lesion. The lesion was extended up to the medulla. Figure 15 shows a large lesion created through fat using transducer movement in a grid pattern. Presumably, there were no air spaces between fat and kidney and, therefore, the intended target was completely covered with necrosis. Figure 16 shows a lesion placed 4 cm deep in kidney tissue. The thick kidney tissue was achieved by using two kidneys, with one kidney placed on top of the other. The importance in this result is that the lesion propagates through the medulla region of the top kidney. This animal model, which was used extensively, proved to be a very good verification tool for thick kidney heating.
Fig. 13. Large lesion in pig kidney in vitro by moving the transducer in a grid formation (8 ⫻ 8) (top view).
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Fig. 14. Large lesion in pig kidney in vitro by moving the transducer in a grid pattern (8 ⫻ 8) (side view).
The presence of cavitation was verified by using a wide-band receiver placed perpendicular to the insonation beam. The threshold of cavitation was found by detecting the in situ spatial average intensity, which produces a signal in the receiver indicative of cavitation. Thus, single lesions were created for this purpose by varying the in situ spatial average intensity from 2000 to 3500 W/cm2. For the given transducer and for a pulse duration of 5 s, it was observed that cavitation took place for in situ spatial-average intensity levels greater than 3200 W/cm2 (error was ⫾ 100 W/cm2). Figure 17 shows lesion length using 5-s pulse and 4000 W/cm2 in situ spatial-average intensity in kidney tissue in vitro. Note that, at this acoustic level, cavitational effect occurred, which was verified by the cavitational receiver. Tissue examination revealed gross tissue disruption, verifying the event of acoustic cavitation. In the in vivo study, a total of six animals experiments have been conducted. Both kidneys of each animal have been ablated; thus, ablation was performed in 12
Fig. 15. Large lesion in pig kidney in vitro through fat layer using a grid pattern (side view).
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Fig. 16. Deep-seated lesions in pig kidney in vitro using two kidneys placed on top of each other.
kidneys. Successful necrosis was achieved in all 12 kidneys. All the animals in this series tolerated the procedure well and remained in good condition throughout the period of the experiment. Figure 18 shows the length of an 8 ⫻ 8 matrix of lesions created using 2000 W/cm2 for 5 s in pig kidney in vivo. The transducer was moved with a step of 3 mm. The lesions were usually firm and white and, usually, no adjacent hemorrhagic zone was created. Only in one kidney was haemorrhage observed. Despite the thick fat layer, necrosis of the renal cortex was reliable. The lesion length in the in vivo case appeared to be smaller (about 18 mm compared to 20 mm in the in vitro case). Single lesions produced in the in vivo kidney revealed a lesion width of about 3 mm for 2000 W/cm2 at 5 s. Thus, the step of 3 mm produced overlapped lesions covering a large target volume. DISCUSSION The trend of attenuation with dose shows a similar trend to the results published by Damianou et al. (1997).
Fig. 17. Large lesion in pig kidney in vivo by moving the transducer in grid pattern showing cavitational activity.
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Table 2. Summary of simplistic protocol for treating renal tissues Pulse duration Delay between US firings Intensity for thermal lesions Spatial step Treatment time for a 20 mm ⫻ 20 mm ⫻ 20 mm target
5s 10 s ⬍ 3200 W/cm2 3 mm 8 ⫻ 8 (10 ⫹ 5) ⫽ 960 s ⫽ 16 min
Transducer parameters are frequency ⫽ 4 MHz, diameter ⫽ 4 cm and radius of curvuture ⫽ 10 cm. Focal depth ⫽ 15 mm.
Fig. 18. Large lesion in pig kidney in vivo (through fat layer) by moving the transducer in grid formation.
The normal kidney tissue included both medulla and renal cortex, for which we found no significant difference in terms of ultrasonic attenuation. The research on muscle and fat is justified because these two tissues usually surround the kidney. The trend of attenuation with dose in all tissues was similar. There was a rapid increase of attenuation with thermal dose and, when necrosis occurred, attenuation stabilised. At that point, the attenuation was roughly double the attenuation at room temperature. Creation of thermal lesions in the cortex is very consistent and repeatable. For the given transducer (f ⫽ 4 MHz, r ⫽ 10 cm and d ⫽ 4 cm) and the given exposure (in situ spatial-average intensity ⫽ 2000 W/cm2, pulse duration 5 s and 15 mm focal depth), the lesion length was around 20 mm and the lesion width was about 3 mm. The lesion length and width of Fig. 8, which was produced in a tissue in vitro (i.e., perfusion rate is 0) is best estimated using the variable attenuation model. The experimental length was 20 mm and the width was 3 mm. The simulated length was 20.6 mm and the width 3.3 mm, using the variable attenuation model. Note that, at perfusion of 70 kg/m3-s (in vivo case), the simulated lesion length and width is decreased compared with the in vitro case. The lesion length and width was generally bigger in the variable attenuation model, because the increased absorption raised the power density in the tissue and, thus, the lesion size increased. This is true for this type of transducer and focal depth. From the equation of absorbed power density, it is clear that the effect of attenuation depends on the transducer frequency and focal depth. Lesions were created all the way to the medulla, provided that there were no air spaces in the medulla. Air spaces are created because of the absence of blood that normally flows in this region. However, ablation of the cortex tissue is of primary concern because the renal
carcinoma (most important renal cancer) grows inside the cortex and then extends to the peritoneal fat. Lesions can be created through fat layers, provided that there are no air spaces between the fat and kidney interface. The lesion size when the beam goes through a fat layer is decreased, because of the high attenuation of fat, which acts likes an acoustic barrier. Deep-seated lesions in the kidney can be created by using two kidneys on top of each other. With a 4-MHz transducer, 5-s pulses and in situ spatial-average intensity levels below about 3200 W/cm2, the lesions created are based solely on thermal effects. Data from Hynynen (1991) suggest that the threshold spatial peak intensity at 4 MHz is about 6000 W/cm2 (extrapolated because data up to 2 MHz is provided) or 3333 W/cm2 (in situ spatial-average intensity). Thus, with the in situ spatial-average intensity of 4000 W/cm2 used in some experiments, cavitation effects were not avoided. In the cavitation range, lesions have irregular boundaries, as opposed to the case of operating below the cavitation range, where the lesions have smooth boundaries, exhibit less mechanical damage to the tissues and always occur at the centre of the focal region. Lesions were smaller in the in vivo experiments compared with lesions created in dead tissue for the same exposure. The reduced size of the in vivo exposures is attributed to the blood flow and possible reduction of the energy due to attenuation at various interfaces, including the skin. Table 2 summarises the treatment protocol for a shallow target in kidney (about 15 mm deep) for a transducer operating at 4 MHz (a frequency with high cavitational threshold). The protocol is simplistic in the sense that integer numbers are used. The determination of these numbers gives a good indication of how to derive a treatment protocol in humans. In the case of humans, the main difference from this animal model is certainly the location of the tumours, which are deeper in the body. Thus, to achieve satisfactory heating in the human case, probably a more powerful transducer needs to be used. The treatment time for a 20 mm ⫻ 20 mm ⫻ 20 mm target is around 16 min. This treatment time can
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be reduced by using, for example, spatial average intensity of 3000 W/cm2 (spacing between transducer movement is around 3.8 mm) and, therefore, treatment time is 6 ⫻ 6 ⫻ 15 ⫽ 735 s ⫽ 9 min (a reduction of 44%). This is just one example of reducing the treatment time. Acknowledgements—This work was supported by the Research Promotion Foundation (RPF) of Cyprus (contract 25/99).
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