A Cylindrical Phased-Array Ultrasound Transducer for Breast Tumor Thermal Therapy Cheng-Shiao Ho1, Kuen-Cheng Ju1, Yung-Yaw Chen2, Win-Li Lin1 1 2

Institute of Biomedical Engineering, National Taiwan University, Taipei, Taiwan Department of Electrical Engineering, National Taiwan University, Taipei, Taiwan

Abstract— The purpose of this study is to investigate the feasibility of a cylindrical phased-array ultrasound transducer for producing a conformal breast tumor thermal therapy. In this study, a 1-MHz cylindrical phased-array transducer consisting of 200 elements with a radius of 10 cm and a width of 2 cm surrounds the breast. The phased-array transducer has the ability to dynamically focus the acoustic energy on the planning target volume (PTV). Different PTVs are evaluated and the simulations show that the volume percentage of the PTV with thermal dose higher than 240min is over 90% and the overheating volume outside the PTV is less than 30%. This method provides good conformal heating for breast tumor thermal therapy. Keywords: cylindrical ultrasound phased-array,breast tumor, scan,thermal therapy

I.

INTRODUCTION

Ultrasound thermal therapy is a truly non-invasive treatment method for tissue ablation deep in the body. The main goal is to destroy cells by raising the temperature and thermal dose of the target volume to a level that denature cell proteins and cause cellular death. Theoretically, tissue temperature rises above a threshold 56℃ for 1 second leads to irreversible cellular death. Results of early clinical studies have shown the potential of the ultrasound thermal therapy for solid tumors （both malignant and benign）, such as liver [12], prostate [3], breast [4-5], kidney [6], uterine fibroids [7] bone, and bladder. The advantages of ultrasound thermal therapy include the ability to focus the ultrasound power on the target region precisely. Ultrasound heating is a nonradiating modality, and therefore the treatment of ultrasound thermal therapy can be repeated if required. The first clinical studies of ultrasound thermal therapy for treatment of breast tumor using single element transducer [810] or phased-array transducer [11] have been accomplished. The main limitation of the single element transducer is that the size of the focal region is too small, approximately 1-2㎜ in diameter. Therefore, the treatment time is significantly increased when the treatment target volume is large and an undesired heating may appear. The increasing treatment time is mainly because that not only the ultrasound focal region is small but also some cooling time is required between the heating lesions. Even the ultrasound thermal therapy is developed rapidly in the past decade, the total treatment time is still long and there are some overheating problems when the

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target volume is close to the ribs or skin. The relevant anatomy in the treatment for breast tumor is commonly to take account of the ribs. The acoustic absorption in ribs at frequencies of 0.5–5 MHz generally used for the ultrasound thermal therapy is one or two order(s) of magnitude greater than that in the soft tissue [12]. Not only the absorption but also the impedance in ribs is larger than that in soft tissue. Therefore, ultrasound wave reflects significantly when propagating through the interface of soft tissue and ribs. These ultrasound physical phenomena may cause the limitation of the treatment such as the patient pain by generating localized high temperature. In this paper our goal is to eliminate these problems and make treatments more efficient. To achieve this goal, a cylindrical phased-array ultrasound transducer with an appropriate heating strategy was proposed. The emitted wave from the transducer approximately paralleled to the ribs. Therefore, not only the power deposition but also the temperature rise of ribs can be effectively eliminated. The preliminary simulations were performed to investigate the feasibility of this ultrasound transducer for producing a conformal heating while preventing overheating or overdosing the ribs and normal tissue for breast tumor thermal therapy. II.

METHODOLOGY

A. System description A schematic diagram of the proposed system is illustrated in Fig. 1. A 1-MHz cylindrical phased-array transducer consisted of 200 elements with a radius of 10 cm and a width of 2 cm. The breast was taken as a 6 cm- radius hemisphere and was surrounded by the transducer. The geometrical focus of the cylindrical phased-array was mechanically located at the center of the target volume. Water was filled between the transducer and breast. Foci were electrically scanned to cover the entire target volume. The main advantage of this arrangement is that the emitted wave from the transducer approximately parallels to the ribs. Therefore, not only the power deposition but also the temperature rise of ribs can be minimized. B. Ultrasonic pressure field calculations Continuous wave sonication was used in the simulation. The ultrasonic pressure field was calculated using the Rayleigh-Sommerfeld to integral the contribution of each

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A. Temperature and thermal dose calculations The temperature response in tissue induced by thermal conduction, blood perfusion and ultrasonic energy absorption is modeled by using the Pennes bioheat transfer equation（BHTE ） [14]

TableⅠ Acoustic and thermal properties of the tissue used in simulation Tissue parameters Tissue density , ρ（kg m-3） Speed of sound , c（m s -1） Ultrasound absorption coefficient, α （Np m-1 @ 1 MHz） Thermal conductivity, k （W m-1 ℃ -1） Specific heat capacity (tissue/blood), Blood perfusion,

wb

ct / cb （J kg

（kg m-3 s-1）

-1

-1

℃ ）

Value 1000 1500 8.86 0.5 3770

ρ ct

1.7

∂T = k ∇2T − wb cb (T − Ta ) + q ∂t

(2)

where T and Ta are the temperature of tissue and arterial blood (37℃), respectively, k is the thermal conductivity, ct and cb are the specific heat capacity of tissue and blood, respectively, and wb is the blood perfusion rate. The parameters used are listed in Table I. The power deposition q obtained from Eq. (1). The BHTE was solved by using a numerical finite difference scheme. The thermal dose (TD) of the tissue was calculated by using Sapareto and Dewey’s thermal dose function [15]. tf

TD= ∫ R (T − 43) dt

(3)

t0

where R=2 for T ≥ 43℃, and R=4 for T＜43℃. Tissue was considered to be necrotic when TD exceeded 240 minutes. [16].

Fig 1: Schematic diagram of cylindrical phased-array ultrasound transducer system: (a) top view (b) side view.

point source on the surface of the transducer. The acoustic power deposition q is given as [13]

q=

αp 2 ρc

(1)

where α is the ultrasound absorption coefficient of tissue, p is the ultrasonic pressure, ρ is the tissue density, and c is the speed of sound in tissue. Values for α, ρ, c used in simulation are listed in Table I. The driving signals for the transducer elements that produce a specific focused pattern are calculated by a pseudo inverse method.

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B. The heating strategy To produce a conformal heating, the predetermined focal spots were uniformly distributed within the planning target volume (PTV) in X-Y plane. The thickness of the heating pattern is dependent on the size of the focal spot in Z-direction. The size of -6dB intensity of the focal spot in Z-direction (~0.8cm) is a good approximation of the thickness of the resulting heating pattern. If the thickness of PTV is greater than 0.8 cm, the PTV was divided into several stacks and then heated separately to cover the whole target region. Phased-array transducer has the ability to produce multiple focus power patterns (MFPP). To reduce the overall treatment time, a multiple-focus scan strategy was employed [17]. Each MFPP consists of four foci with 0.2 cm apart. At the center of each MFPP, a temperature control point (TCP) was placed. Fig. 2 shows the arrangement of the MFPPs and TCPs. During the heating process, MFPPs were sequentially and repeatedly scanned with a 0.1s sonication time at each position. No cooling phase was used between sonications. The maximum permissible intensity of foci is based on the restriction of 5 W cm-2 output powers on the transducer surface. When the temperature at a TCP reached the target temperature (Tg), the corresponding MFPP would not be sonicated at the following scanning. The scanning heating process stopped when all TCPs had ever reached the Tg. The coverage volume index (CI) and the external volume index (EI) was used to evaluate the conformability of the resulting heated patterns. CI = (PTVTD≧240)/PTV*100%

(4)

EI = (OTVTD≧240)/PTV*100%

(5)

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where PTVTD≧240 is the volume inside the PTV with TD≧240 minutes, and OTVTD≧240 is the volume outside the PTV with TD≧240 minutes.

Fig. 4 shows the corresponding CI and EI values of Fig. 3. Fig. 4(a) shows the variation of CI and EI with time. The heating phase ended at 32.3 sec, and the CI increased sharply within this period. After 120 sec, the CI reached to 92.3%, and EI converged to 18.7%. The target temperature, Tg, is also a parameter that may affect the heating results. Fig. 4(b) displays the values of CI and EI as a function of Tg and shows that CI approaches to 100% for Tg equal to 54℃ while EI increases to 40%. This indicates that when entire PTV is heated to have a thermal dose higher than TD240, the resulting cost is to increase the damaged volume in normal tissue.

Fig 2: Scheme of the multiple-focus scan and the locations of temperature control points.

II. RESULTS AND DISCUSSION Figure 3 shows the simulation results for a 1×1×0.8 cm3 PTV located at (0,0,-1) for a Tg of 52oC. Figure 3(a) shows the 3-D, isosurface of the thermal dose 240min (TD240), and figures 3(b) to 3(d) shows the thermal dose contours on X=0, Y=0 and Z=-1 planes, respectively. It can be seen that the TD240 almost completely covers the entire PTV region except the corner regions, which is that due to the strong conduction effect. At the thermal lesion boundary, TD100 and TD240 contours are very close and it indicates that the thermal dose decreases sharply and the thermal effects on normal tissue can be reduced.

Fig 3: Heating results for a 1×1×0.8 cm3 PTV at (0,0,-1) with a Tg of 52℃. (a) TD240 isosurface contours. TD100 and TD240 contours on (b) X=0 plane, (c) Y=0 plane, (d) Z=-1 plane. The blue solid lines indicate the PTV.

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Fig 4: Coverage volume index (CI) and external volume index (EI) for the PTV in Fig. 3. (a) CI and EI as a function of time for Tg = 52℃, (b) CI and EI for various Tg.

For testing the flexibility of this cylindrical phased-array system, Fig. 5 shows the heating results for different types of PTVs with a Tg of 52℃. Figures 5(a)-(d) display the resulting 3D isosurface of TD100 and TD240 for PTV of 1×1×0.8 cm3 located at (4,0,-1) and 2×2×0.8 cm3 located at (0,0,-1) and (4,0,-1), respectively. Quantitative indices are listed in Table II. A rather high CI value greater than 90% is obtained with an EI value less than 30% and a maximum temperature less than 57 ℃ . The heating time increased with a smaller maximum permissible intensity, Imax. When the PTV is larger, it takes a longer time to cool the tissue to below 43℃.

Fig 5: Isosurfaces of TD=240 minutes for three PTV types: (a) 1×1×0.8 cm3 at (4,0,-1), (b) 2×2×0.8 cm3 at (0,0,-1), (c) 2×2×0.8 cm3 at (4,0,-1). The blue solid lines indicate the PTV.

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TableⅡ Simulation results for different planning target volumes [7]

[8]

[9]

III.

CONCLUSION

In this study, a cylindrical phased-array ultrasound transducer for breast tumor thermal therapy was proposed. Four types of PTV were used to evaluate the feasibility of the proposed system and the heating strategy. The main advantage of the proposed system is that the ultrasound beam path was approximately parallel to the ribs; hence it is capable of preventing the undesired temperature rise in the ribs during the heating process. Simulation results show that the coverage volume index (CI) are over 90％ in all cases and the external volume index (EI) ranges from 18.7％ to 29.7％. Comparing to the previous studies, this system provides an acceptable heating results in ultrasound thermal therapy. The preliminary simulation results presented in this paper indicate that the proposed cylindrical phased-array ultrasound transducer have the potential for breast tumor thermal therapy.

[10]

[11]

[12]

[13] [14] [15]

Acknowledgment [16]

The authors would like to thank the National Science Council of Taiwan for partially supporting this research under contract no. NSC 93-2213-E-002-069.

[17]

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