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Research Papers

# Enhanced Heat Transfer and Thermal Dose Using Magnetic Nanoparticles During HIFU Thermal Ablation—An In-Vitro StudyOPEN ACCESS

[+] Author and Article Information
Seyed Ahmad Reza Dibaji, Marwan F. Al-Rjoub

Department of Mechanical
and Materials Engineering,
College of Engineering and Applied Science,
University of Cincinnati,
2600 Clifton Avenue,
Cincinnati, OH 45221

Matthew R. Myers

Division of Solid and Fluid Mechanics,
Center for Devices and Radiological Health,
U. S. Food and Drug Administration,
10903 New Hampshire Avenue,
Silver Spring, MD 20993

Rupak K. Banerjee

Department of Mechanical
and Materials Engineering,
College of Engineering and Applied Science,
University of Cincinnati,
598 Rhodes Hall,
P.O. Box 210072,
Cincinnati, OH 45221
e-mail: rupak.banerjee@uc.edu

1Corresponding author.

Manuscript received January 20, 2014; final manuscript received March 19, 2014; published online April 15, 2014. Assoc. Editor: Sumanta Acharya. This material is declared a work of the US Government and is not subject to copyright protection in the United States. Approved for public release; distribution is unlimited.

J. Nanotechnol. Eng. Med 4(4), 040902 (Apr 15, 2014) (8 pages) Paper No: NANO-14-1005; doi: 10.1115/1.4027340 History: Received January 20, 2014; Revised March 19, 2014

## Abstract

Avoiding collateral damage to healthy tissues during the high intensity focused ultrasound (HIFU) ablation of malignant tumors is one of the major challenges for effective thermal therapy. Such collateral damage can originate out of the need for using higher acoustic powers to treat deep seated or highly vascularized tumors. The objective of this study is to assess the utility of using magnetic nanoparticles (mNPs) during HIFU procedures to locally enhance heating at low powers, thereby reducing the likelihood of collateral thermal damage and undesired destruction due to cavitation. Tissue phantoms with 0% (control), 1% and 3% mNPs concentrations by volume were fabricated. Each tissue phantom was embedded with four thermocouples (TCs) and sonicated using transducer acoustic powers of 5.15 W, 9.17 W, and 14.26 W. The temperature profiles during the heating and cooling periods were recorded for each embedded TC. The measured transient temperature profiles were used for thermal-dose calculations. The increase in the concentration of mNPs in the tissue phantoms, from 0% to 3%, resulted in the rise in the peak temperatures for all the TCs for each acoustic power. The thermal dose also increased with the rise in the concentration of mNPs in the tissue phantoms. For the highest applied acoustic power (14.26 W), the peak temperature at TC 1 (T1) in tissue phantoms with 1% and 3% mNPs concentrations increased (with respect to tissue phantom with 0% (control) mNPs concentration) by 1.59× and 2.09×, respectively. For an acoustic power of 14.26 W, the time required to achieve cellular necrosis as defined by a 240 equivalent min thermal dose was approximately 75 s in the absence of mNPs, 14 s for the 1% concentration, and 8 s for the 3% concentration. Magnetic nanoparticles have the potential to significantly reduce the time for HIFU thermal-ablation procedures. They can also decrease the likelihood of collateral damage by the propagating beam in HIFU procedures by reducing the intensity required to achieve cellular necrosis.

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## Introduction

HIFU is a noninvasive medical procedure with a significant potential for a variety of therapeutic applications. Some of the medical applications of HIFU include tumors ablation, drug delivery, hemostasis, and gene activation [1]. Unlike other hyperthermia techniques, HIFU ablation is completely noninvasive with minimum patient recovery time [2]. During HIFU ablation procedures, the mechanical energy of acoustic waves converts to thermal energy as the ultrasound propagates through the tissue. The localized temperature rise causes rapid cell death, or necrosis, in the targeted region [2]. In order to ensure safety and efficacy of HIFU devices, preclinical evaluations of the thermal and acoustic field generated by the HIFU transducers are necessary. Preclinical testing of HIFU systems has been performed using computational modeling [3-7], ex-vivo [8-11] or in-vivo [12-15] animal tissues, and tissue phantoms [16-19]. The advantage of using phantoms with tissue mimicking material (TMM) is that such test sections can be reused repeatedly without affecting acoustical and thermal properties, which can be made similar to that of human tissue [20].

Despite the advancements of HIFU in recent years, there are still several factors limiting its application in thermal ablation. For example, the efficacy of HIFU in ablating the larger and deeper tumors can be low, as the ultrasound intensity attenuates exponentially with the increase of depth in tissue [21,22]. In addition, the blood flow through a large vessel can significantly reduce the temperature elevation, when the HIFU focus is close to the vessel wall [23-25]. The HIFU energy deposition in larger or highly vascularized tumors can be enhanced by increasing either the ultrasound power or sonication time [22,26]. However, increasing the acoustic power or exposure time can lead to unnecessary heating of surrounding normal tissues, causing collateral damage to healthy tissues such as nerve injury and skin burns [26]. Therefore, it is advantageous to seek methods for better efficacy of HIFU thermal therapy using low ultrasound power and short exposure time.

Ultrasound contrast agents can be used to improve the therapeutic efficacy of HIFU procedures. In the studies by Luo et al. [26] and Tung et al. [27], microbubbles were used as ultrasound contrast agents to enhance energy deposition in the HIFU focal region. The presence of microbubbles near the HIFU focus also resulted in larger lesions [28]. Similarly, Tran et al. [29] reported that the requisite ultrasound intensity and exposure time can be reduced by the administration of contrast microbubbles.

In addition to ultrasound contrast agents, mNPs can also be used to enhance the HIFU thermal ablation of cancer tumors. In the study by Quanyi et al. [30], mNPs were incorporated into an interface layer in an egg white phantom radiated by HIFU. It was reported that the lesion volume was enlarged when the acoustic focal region was close to the interface. More recently, Ho et al. [31] examined the effect of magnetic (Fe3O4) nanoparticle agglomerates on the destruction of tumor spheroids using HIFU ablation. Hela multicellular spheroids were insonated in the presence and absence of mNPs agglomerates. They found that the magnetic nanoparticle agglomerates can increase the degree of HIFU induced inertial cavitation, possibly at higher temperatures, and consequently enhance the rate of destruction of tumor spheroids. However, no examination was performed to investigate the effect of mNPs on the HIFU induced temperature rise in the absence of inertial cavitation. In another study, supermagnetic poly(lactic-co-glycolic) acid (PLGA) microcapsules (Fe3O4/PLGA) were administered into a rabbit breast cancer model to evaluate the in-vivo HIFU synergistic ablation efficiency caused by the introduction of such microcapsules [32]. It was found that the volume of coagulative necrosis was substantially larger after the injection of Fe3O4/PLGA microcapsules compared to the group without this agent. The results showed that Fe3O4/PLGA microcapsules could improve the ablation efficiency of HIFU at a lower power and shorter exposure duration. However, the impact of using different amounts of Fe3O4 in the synthesis process was not studied.

Gold nanoparticles-coated, perfluorohexane-encapsulated, and PEGylated mesoporous silica nanocapsule-based enhancement agents (MSNC@Au-PFH-PEG, abb. as MAPP), were also used as an intensified ultrasound-guided HIFU enhancement agent [33]. It was found that the ultrasound-guided HIFU therapy ex-vivo and in-vivo with MAPP can be highly efficient on rabbit VX2 xenograft tumor ablation due to both thermal energy accumulation and PFH bubble cavitation.

In the present study, the effect of mNPs (Fe3O4) on the HIFU induced temperature rise and thermal dose was assessed. Tissue phantoms with 0% (control), 1% and 3% mNPs concentrations by volume were fabricated. Each tissue phantom was embedded with four TCs and sonicated using transducer acoustic powers of 5.15 W, 9.17 W, and 14.26 W. The temperature profiles during the heating and cooling periods were recorded at each embedded TC. The measured transient temperature profiles were used for thermal dose calculation. The temperature profiles as well as the thermal doses at each TC were compared for the three different concentrations for each acoustic power to quantify the effect of mNPs.

## Methods

###### Fabrication of Tissue Phantoms With 0% (Control), 1% and 3% mNPs Concentrations.

Three cylindrical fixtures with the length of 5 cm and the inner diameter of 3 cm (volume = 35.34 cm3) were developed. Each fixture was embedded with an array of four thin-wire (Chromega-Constantine) TCs with the diameter of 0.003 in, labeled T1-T4, arranged in two layers (Fig. 1). The TC standard limit of error (above 0 °C) was greater of 1.7 °C or 0.5% of the measured temperature based on the TC accuracy chart provided by the company. Each layer had two TCs with wires that were parallel to each other, and the TC wires in one layer were oriented 90 deg to those in the other layer. The two TCs in each layer were separated by a distance of 4 mm, and each layer was 3 mm in axial extent away from the adjacent layer.

A 40 mL gelrite-based TMM was prepared according to the protocol of King et al. [19]. Properties of the TMM are presented in King et al. [19]. To prepare a tissue phantom with 0% mNPs concentration by volume, the 40 mL liquid TMM was poured into the first fixture (volume = 35.34 cm3) until the fixture was filled. To construct the tissue phantom with 1% mNPs concentration, the following mixture was first created. 10.26 mL of water-based ferrofluid (EMG705 series, Ferrotec (USA) Corporation, Nashua, NH) with a magnetic (Fe3O4) biocompatible particles concentration of 3.9% by volume and a particle size of 10 nm was diluted with 29.74 mL of water (10.26 mL + 29.74 mL = 40 mL) and then mixed with the TMM raw materials following the King et al. [19] protocol. The resulting mixture was poured into the second fixture (volume = 35.34 cm3) until the fixture was filled. In order to prepare a mNPs infused tissue phantom with 3% concentration by volume, a 30.77 mL of the provided EMG705-ferrofluid was diluted with 9.23 mL of water (30.77 mL + 9.23 mL = 40 mL) and then mixed with the TMM raw materials following the King et al. protocol [19]. The resulting 40 mL liquid mixture was then poured into the third fixture (volume = 35.34 cm3) until the fixture was filled. Thus, three tissue phantoms with the mNPs concentrations (by volume) of 0%, 1%, and 3% were developed and each tissue phantom was embedded with four TCs. It should be noted that the three tissue phantoms were kept at room temperature for about 12 h so that the poured liquid in the fixtures solidified completely.

###### Micro-Computed Tomography (Micro-CT) Imaging.

High resolution micro-CT (Inveon CT, Siemens, Germany) was used to scan the three tissue phantoms with 0% (control), 1% and 3% mNPs concentrations (Fig. 2), to differentiate between the tissue phantoms using imaging modalities. The micro-CT images (100 μm/voxel) were brighter for tissue phantoms with higher mNPs concentration (Fig. 2). The Hounsfield Unit (HU) values were -191, -34, and 285 for tissue phantoms with 0%, 1%, and 3% mNPs concentration, respectively.

###### Sonication Procedure.

Figure 3(a) shows the experimental setup used for performing HIFU sonications. The HIFU source was a H102 transducer (Sonic Concepts Inc., Bothell, WA) with a focal length of 6.26 cm, outer diameter of 6.4 cm, and inner diameter of 2.2 cm. The operating frequency was 1.025 MHz. The transducer was driven in continuous-wave mode by a signal generator (33220A, Agilent Technologies). The driving signal was amplified through the use of a 150-Watt amplifier (150A100B, Amplifier Research). The transducer and the tissue phantoms resided in a tank of degassed water. A holding rod was mounted to the transducer housing and was parallel to the transducer axis of symmetry (Fig. 3(a)). The holding rod was attached to an x-y-z positioning system that was capable of adjusting any of the coordinates in discrete 0.025 mm increments.

The following procedure was performed for sonicating each of the phantoms. First, the rod holding the transducer, the tissue phantom axis of symmetry and the z-axis of the positioning system were placed parallel to each other (Fig. 3(a)). The HIFU beam axis was thus parallel to the axis of symmetry of tissue phantom, while the TC wires were perpendicular to the beam axis. T1 and T2 TC wires were parallel to the y-axis while T3 and T4 wires were parallel to the x-axis of the positioning system (Fig. 3(b)). In order to find the T1 junction, the beam was moved inside the tissue phantom, using the 3D positioning system, until the maximum temperature rise for T1 was observed during a brief sonication period (10 s). After positioning the beam on the T1 junction (Fig. 3(b)), the transducer was activated in a continuous-wave mode at a time of 5 s for a period of 30 s. The temperature on the full array, i.e., T1, T2, T3, and T4, was recorded using an OMB-DAQ-56 (Omega Engg. Inc., Stamford, CT) data acquisition system over the 5 s presonication period, 30 s heating period, and 40 s cooling period (total time = 75 s). The temporal resolution of the temperature measurements was 0.5 s. A heating period of 30 s was chosen because it yielded a temperature just short of boiling when the highest power and highest mNPs concentration were used. Three transducer acoustic powers of 5.15 W, 9.17 W, and 14.26 W were used. These powers correspond to values at which the transducer was previously calibrated using a radiation-force balance. Three trials (n = 3) were performed for each power level. After each trial, the tissue phantoms were allowed to cool down to the ambient water temperature.

###### Thermal Dose.

The thermal dose at each TC (T1, T2, T3, and T4) was calculated using the corresponding transient temperature profile, according to the method developed by Sapareto and Dewey [34]. The thermal dose parameter is expressed asDisplay Formula

(1)$t43(x,y,z)=∫t=0t=tfinalR43-T(t)dt$

where t43 is the thermal dose at the reference temperature of 43 °C, tfinal is the treatment (sonication) time, T(t) is the temperature (in  °C) as a function of time obtained experimentally, and

$R={0.5 if T(t)≥43 °C0.25 otherwise}$

A trapezoidal scheme was used to perform the integration shown in Eq. (1) with dt = 0.5 s.

## Results

The HIFU induced temperature rise was measured using embedded TCs in tissue phantoms with three different mNPs concentrations (0% (control), 1% and 3%). The temperatures of the degassed water and the presonication tissue phantoms were 23.6 °C. The measured temperatures at each TC (T1, T2, T3, and T4) were averaged over three trials for each acoustic power level (5.15 W, 9.17 W, and 14.26 W). Although the temporal resolution of measurements was 0.5 s, the error bars were shown at every 2 s for all the temperature plots to provide better clarity. Results are presented as mean ± SD.

Figures 4(a)4(c) show the HIFU induced transient temperature profiles at T1 in tissue phantoms with various mNPs concentrations (0%, 1%, and 3%) using acoustic powers of 5.15 W, 9.17 W, and 14.26 W, respectively. Since these sonications were performed directly atop T1, the temperature readings are subject to TC artifact (Morris et al. [35]). However, it was assumed that the artifact was comparable for the three nanoparticle concentrations. The peak temperature at T1 for the acoustic power of 5.15 W (Fig. 4(a)), increased from 30.87 ± 0.03 °C for 0% mNPs concentration to 40.71 ± 0.05 °C for 1% mNPs concentration, and then to 52.42 ± 0.19 °C for 3% mNPs concentration. Using the acoustic power of 9.17 W (Fig. 4(b)), the measured peak temperature at T1 increased from 38.08 ± 0.07 °C for 0% mNPs concentration to 52.87 ± 0.04 °C for 1% mNPs concentration, and then to 85.32 ± 0.28 °C for 3% mNPs concentration. Similarly, the T1 measured peak temperature for the power of 14.26 W (Fig. 4(c)) increased from 44.67 ± 0.18 °C to 70.97 ± 0.14 °C and then to 93.16 ± 0.13 °C for mNPs concentrations of 0%, 1%, and 3%, respectively.

The temperatures as a function of time for T2 in tissue phantoms with different mNPs concentrations are plotted in Figs. 5(a)5(c) for the power levels of 5.15 W, 9.17 W, and 14.26 W, respectively. T2 was the farthest TC from the beam axis and the measured temperature at T2 was not affected by the artifact as for T1 TC. The peak temperature at T2 (Fig. 5) shifted to times greater than 35 s (i.e., beyond the end of sonication) as heat diffused from the most intense part of the beam near T1. The peak temperature at T2 for the power of 5.15 W (Fig. 5(a)) increased from 24.85 ± 0.04 °C for 0% mNPs concentration to 26.15 ± 0.04 °C for 1% mNPs concentration, and then to 28.52 ± 0.25 °C for 3% mNPs concentration. For the acoustic power of 9.17 W (Fig. 5(b)), the peak temperature at T2 increased from 25.74 ± 0.11 °C to 28.07 ± 0.03 °C, and then to 32.93 ± 0.001 °C for mNPs concentrations of 0%, 1%, and 3%, respectively. For the power level of 14.26 W (Fig. 5(c)), the measured peak temperatures at T2 were 26.94 ± 0.04 °C, 30.68 ± 0.07 °C, and 37.23 ± 0.04 °C for 0%, 1%, and 3% mNPs concentrations, respectively.

The transient temperature profiles for T3 in tissue phantoms with three different mNPs concentrations are plotted in Figs. 6(a)6(c) using acoustic powers of 5.15 W, 9.17 W, and 14.26 W, respectively. Similar to the temperature profiles at T1 and T2, the measured peak temperature at T3 increased with the increase in mNPs concentration for each applied power level. The rise in the measured peak temperatures at T4 with the increase in mNPs concentration are shown in Figs. 7(a)7(c) for the power levels of 5.15 W, 9.17 W, and 14.26 W, respectively.

The increases in peak temperature at T1 in tissue phantoms with 1% and 3% mNPs concentrations with respect to tissue phantom with 0% (control) mNPs concentration are summarized in Table 1, for all three powers. For the acoustic power of 5.15 W, the peak temperature at T1 increased (with respect to 0% mNPs concentration case) by 1.32× and 1.70× for mNPs concentrations of 1% and 3%, respectively. For the acoustic power of 9.17 W, the peak temperature at T1 increased (with respect to 0% mNPs concentration case) by 1.39× for 1% mNPs concentration and 2.24× for 3% mNPs concentration. Similarly, the peak temperature increase at T1 for the power level of 14.26 W was 1.59× and 2.09× for mNPs concentrations of 1% and 3%, respectively.

In contrast to heating, cooling within the phantom was essentially unaffected by the presence of the nanoparticles. While a larger decrease in temperature occurs in a given period of cooling time for the 1% and 3% mNPs cases, when normalized by the higher temperature rises occurring during the heating period, the cooling rate is essentially independent of mNPs concentration. Table 2 provides the time required for the temperature to drop to 50% of its peak value, for T1. No significant trends can be identified, either with acoustic power or nanoparticle concentration.

The thermal dose at each TC was computed using Eq. (1) and averaged over three trials. For all three powers (5.15 W, 9.17 W, and 14.26 W) and all three mNPs concentrations (0%, 1%, and 3%), the computed thermal doses at T1 are presented in logarithmic scale in Fig. 8(a). The calculated thermal dose at T1 for the power of 5.15 W increased from 6.12 × 10−09 ± 0.20 × 10−09 min for 0% mNPs concentration to 2.59 × 10−03 ± 0.22 × 10−03 min for 1% mNPs concentration, and then to 5.01 × 10+01 ± 0.66 × 10+01 min for 3% mNPs concentration. The calculated thermal dose at T1 for the power of 9.17 W also increased from 8.17 × 10−05 ± 0.39 × 10−05 min for 0% mNPs concentration to 6.42 × 10+01 ± 0.20 × 10+01 min for 1% mNPs concentration, and then to 2.43 × 10+11 ± 0.48 × 10+11 min for 3% mNPs concentration. The thermal doses at T1 for the power of 14.26 W were 2.57 × 10−01 ± 0.33 × 10−01 min, 1.22 × 10+07 ± 0.06 × 10+07 min and 4.07 × 10+13 ± 0.74 × 10+13 min for the mNPs concentration of 0%, 1%, and 3%, respectively.

Figure 8(b) (logarithmic scale) shows the computed thermal doses at T2 for the three different mNPs concentrations using acoustic powers of 5.15 W, 9.17 W, and 14.26 W. The calculated thermal dose at T2 for the power of 5.15 W increased from 9.57 × 10−12 ± 0.41 × 10−12 min to 4.28 × 10−11 ± 0.21 × 10−11 min, and then to 8.47 × 10−10 ± 2.68 × 10−10 min for the mNPs concentrations of 0%, 1%, and 3%, respectively. The thermal dose elevated from 2.81 × 10−11 ± 0.44 × 10−11 min for 0% mNPs concentration to 5.09 × 10−10 ± 0.14 × 10−10 min for 1% mNPs concentration, and then to 2.70 × 10−7 ± 0.02 × 10−7 min for 3% mNPs concentration for the power of 9.17 W. Similarly, the thermal dose value at T2 for the power of 14.26 W increased from 1.27 × 10−10 ± 0.08 × 10−10 min to 1.55 × 10−08 ± 0.16 × 10−08 min, and then to 8.47 × 10−05 ± 0.30 × 10−05 min for the mNPs concentrations of 0%, 1%, and 3%, respectively.

Using acoustic powers of 5.15 W, 9.17 W, and 14.26 W, the computed thermal doses at T3 and T4 in tissue phantoms with three different mNPs concentrations are plotted in logarithmic scale in Figs. 8(c) and 8(d), respectively. Similar to T1 and T2, the thermal doses for both T3 and T4 increased with the increase in mNPs concentration for each applied power.

In addition to thermal doses, the time required to achieve cellular necrosis—defined as a thermal dose of 240 equivalent min—was also computed. In cases (low power or low mNPs concentration) where the temperature rise was insufficient to produce an equivalent dose of 240 min, the heating curve was extrapolated to longer times using a curve-fitting procedure. A logarithmic fitting function was used; the regression coefficient R2 was greater than 0.98 in each case. The time to achieve cellular necrosis was computed only for T1; the other locations were remote enough from the beam that necrosis was not achieved for the powers considered. In principle, the threshold for cell necrosis established for the particular organ and species should be used in thermal dose calculations. As these values are often not available, a value of 240 equivalent min is usually used. In the HIFU ablation studies by Righetti et al. [36] involving excised livers, a thermal dose of 243 min (∼240 min) was used as the threshold for cell necrosis.

The time to achieve cellular necrosis is shown in Table 3. At the lowest power (5.15 W), a 100-fold decrease in the necrosis time was observed between the 0% concentration and the 3% concentration. At the highest power (14.26 W), the reduction in time-to-necrosis achieved by using a 3% mNPs concentration (relative to no nanoparticles) was less than at lower powers, but still about a factor of 10.

## Discussion

The increases in temperature rise and thermal dose (Figs. 48) associated with small concentrations of nanoparticles are manifestations of a substantial increase in acoustic attenuation due to the presence of mNPs. Dąbek et al. [37] report an increase of more than a factor of 2 in the acoustic attenuation of a liquid nanoparticle suspension, when the concentration of Fe3O4 is increased to only 1%. Dąbek et al. [37] cite two contributions to the attenuation that increase with increasing nanoparticle concentration: a viscous component, involving the difference in density between the nanoparticles and the surrounding medium, and a thermal attenuation, involving the difference in thermal properties between the nanoparticles and their surroundings. Given the small particle size relative to the acoustic wavelength, enhanced scattering due to the presence of the nanoparticles can be neglected, though this may not be possible if the nanoparticles form agglomerations within the phantom.

The increase in heat production associated with the enhanced attenuation can be observed in the initial slope of the temperature traces in Figs. 47. For short times during early phase of sonication, thermal diffusion can be neglected, and the change in temperature is due solely to the absorption of acoustic energy. Mathematically, the temperature rise is linearly proportional to the heat production for small values of time. Table 4 provides the ratio of the initial slope of the temperature trace for a given nanoparticle concentration, normalized by the slope for the case of 0% mNPs. An initial time of 2.5 s was chosen to compute the initial slopes. The 2.5 s time was chosen because it was long enough that the temperature rise due to ultrasound attenuation elevated above the TC noise level on the remote TCs, yet short enough that appreciable diffusion had not yet occurred. For T1, heat production in the presence of nanoparticles is between 1.7 and 6 as intense as the heat production without nanoparticles. For the 3% concentration, the heat-production ratio (with nanoparticles compared to without) greater than 1 for all TCs. At the 1% concentration, ratios less than 1 were observed on T2 and T4; this was likely due to the fact that the beam was positioned slightly closer to T2 and T4, and this positioning difference overwhelmed the effect of the increased attenuation in the presence of the mNPs. Given this uncertainty due to positioning differences, exact numerical values should not be inferred from Table 4, however the trend of increase in heat production with increasing nanoparticle concentration seems clear.

Besides uncertainties associated with positioning the beam atop a given thermocouple, slight differences in thermocouple location between fixtures is an additional source of uncertainty in the comparison between the experiments with and without nanoparticles. Care was taken to construct all phantoms in the same manner, and to pour the liquid gel into the fixtures without directly contacting thermocouples. Still, some thermocouple displacement during the fabrication process was possible. One measure indicating that thermocouples were located in approximately the same location in the different phantoms is the time required for the peak temperature to occur at remote thermocouple locations. At the locations remote from the beam, the peak temperature appeared at approximately the same time for the three fixtures. These times were later than the end of sonication time, since the peak temperature resulted from diffusion of heat generated near the beam axis. In the future, experiments will be performed using the same fixture with and without nanoparticles. After the control experiments without nanoparticles, nanoparticles will be injected into the region of interest with a syringe.

Further work is required to expand the single-location thermal-dose computations featured in Fig. 8 into actual lesion volumes. This could be done with an expanded array, in conjunction with techniques for interpolating values within the array [38]. The results of Fig. 8 do indicate that mNPs infusion can decrease the acoustic power required to achieve cellular necrosis, thereby reducing the risk of collateral damage. This advantage is particularly valuable for deep seated or highly vascularized tumors. An additional advantage of nanoparticle infusion can be seen in the necrosis times in Table 3: the presence of mNPs has the potential to substantially reduce the time for HIFU thermal ablation procedures. This is particularly valuable for large-tumor ablation [21].

## Conclusion

The HIFU induced temperature rise was measured using embedded TCs in tissue phantoms with different concentrations of mNPs for three distinct acoustic powers. The transient temperature profiles were then used to calculate the thermal doses for each power level. The mNPs increased the attenuation of tissue phantoms so that higher peak temperatures were achieved for the TMMs with greater concentrations of mNPs for all different powers. Consequently, higher thermal doses were obtained for greater concentrations of mNPs for all the applied powers. Therefore, in the presence of mNPs, lower acoustic powers can be used to achieve higher thermal doses. Thus, the required power to obtain the adequate thermal dose that is capable of causing cell necrosis in tumors can be reduced substantially with the use of mNPs. This could reduce the likelihood of damage to the healthy tissues caused by the application of higher acoustic powers for thermal therapy of deep seated or highly vascularized tumors. It can also reduce the time required for HIFU ablation procedures.

## Acknowledgements

Financial support from the National Science Foundation (Grant No. 1137166) is gratefully acknowledged. The initial assessments with Micro-CT on tissue phantoms were conducted in collaboration with Dr. Lisa Lemen and Mrs. Kathleen Lasance in Vontz Core Imaging Laboratory at the University of Cincinnati. We appreciate Mr. Dushyanth Giridhar for helping us in developing tissue phantoms and performing HIFU sonications.

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Dasgupta, S., Das, P., Wansapura, J., Hariharan, P., Pratt, R., Witte, D., Myers, M. R., and Banerjee, R. K., 2011, “Reduction of Noise From MR Thermometry Measurements During HIFU Characterization Procedures,” ASME J. Nanotechnol. Eng. Med., 2(2), p. 024501.
Mesiwala, A. H., Farrell, L., Wenzel, H. J., Silbergeld, D. L., Crum, L. A., Winn, H. R., and Mourad, P. D., 2002, “High-Intensity Focused Ultrasound Selectively Disrupts the Blood-Brain Barrier in vivo,” Ultrasound Med. Biol., 28(3), pp. 389–400. [PubMed]
Solomon, S. B., Nicol, T. L., Chan, D. Y., Fjield, T., Fried, N., and Kavoussi, L. R., 2003, “Histologic Evolution of High-Intensity Focused Ultrasound in Rabbit Muscle,” Invest. Radiol., 38(5), pp. 293–301. [PubMed]
Köhler, M. O., Mougenot, C., Quesson, B., Enholm, J., Le Bail, B., Laurent, C., Moonen, C. T. W., and Ehnholm, G. J., 2009, “Volumetric HIFU Ablation Under 3D Guidance of Rapid MRI Thermometry,” Med. Phys., 36(8), pp. 3521–3535. [PubMed]
Quesson, B., Laurent, C., Maclair, G., de Senneville, B. D., Mougenot, C., Ries, M., Carteret, T., Rullier, A., and Moonen, C. T. W., 2011, “Real-Time Volumetric MRI Thermometry of Focused Ultrasound Ablation in vivo: A Feasibility Study in Pig Liver and Kidney,” NMR Biomed., 24(2), pp. 145–153. [PubMed]
Canney, M. S., Bailey, M. R., Crum, L. A., Khokhlova, V. A., and Sapozhnikov, O. A., 2008, “Acoustic Characterization of High Intensity Focused Ultrasound Fields: A Combined Measurement and Modeling Approach,” J. Acoust. Soc. Am., 124(4), pp. 2406–2420. [PubMed]
Chen, D., Fan, T., Zhang, D., and Wu, J., 2009, “A Feasibility Study of Temperature Rise Measurement in a Tissue Phantom as an Alternative Way for Characterization of the Therapeutic High Intensity Focused Ultrasonic Field,” Ultrasonics, 49(8), pp. 733–742. [PubMed]
Farny, C. H., Holt, R. G., and Roy, R. A., 2009, “Temporal and Spatial Detection of HIFU-Induced Inertial and Hot-Vapor Cavitation With a Diagnostic Ultrasound System,” Ultrasound Med. Biol., 35(4), pp. 603–615. [PubMed]
King, R. L., Yunbo, L., Maruvada, S., Herman, B. A., Wear, K. A., and Harris, G. R., 2011, “Development and Characterization of a Tissue-Mimicking Material for High-Intensity Focused Ultrasound,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 58(7), pp. 1397–1405. [PubMed]
Maruvada, S., Liu, Y., Pritchard, W., Herman, B., and Harris, G., 2012, “Comparative Study of Temperature Measurements in Ex Vivo Swine Muscle and a Tissue-Mimicking Material during High Intensity Focused Ultrasound Exposures,” Phys. Med. Biol., 57(1), pp. 1–19. [PubMed]
Kennedy, J. E., 2005, “High-Intensity Focused Ultrasound in the Treatment of Solid Tumours,” Nat. Rev. Cancer, 5(4), pp. 321–327. [PubMed]
Dasgupta, S., Wansapura, J., Hariharan, P., Pratt, R., Witte, D., Myers, M. R., and Banerjee, R. K., 2010, “HIFU Lesion Volume as a Function of Sonication Time, as Determined by MRI, Histology, and Computations,” ASME J. Biomech. Eng., 132(8), p. 081005.
Huang, J., Holt, R. G., Cleveland, R. O., and Roy, R. A., 2004, “Experimental Validation of a Tractable Numerical Model for Focused Ultrasound Heating in Flow-Through Tissue Phantoms,” J. Acoust. Soc. Am., 116(4), pp. 2451–2458. [PubMed]
ter Haar, G., 2008, “Harnessing the Interaction of Ultrasound With Tissue for Therapeutic Benefit: High-Intensity Focused Ultrasound,” Ultrasound Obstet. Gynecol., 32(5), pp. 601–604. [PubMed]
Dasgupta, S., Banerjee, R. K., Hariharan, P., and Myers, M. R., 2011, “Beam Localization in HIFU Temperature Measurements Using Thermocouples, With Application to Cooling by Large Blood Vessels,” Ultrasonics, 51(2), pp. 171–180. [PubMed]
Luo, W., Zhou, X., Tian, X., Ren, X., Zheng, M., Gu, K., and He, G., 2006, “Enhancement of Ultrasound Contrast Agent in High-Intensity Focused Ultrasound Ablation,” Adv. Ther., 23(6), pp. 861–868. [PubMed]
Tung, Y.-S., Liu, H.-L., Wu, C.-C., Ju, K.-C., Chen, W.-S., and Lin, W.-L., 2006, “Contrast-Agent-Enhanced Ultrasound Thermal Ablation,” Ultrasound Med. Biol., 32(7), pp. 1103–1110. [PubMed]
Yu, T., Wang, G., Hu, K., Ma, P., Bai, J., and Wang, Z., 2004, “A Microbubble Agent Improves the Therapeutic Efficiency of High Intensity Focused Ultrasound: A Rabbit Kidney Study,” Urol. Res., 32(1), pp. 14–19. [PubMed]
Tran, B. C., Jongbum, S., Hall, T. L., Fowlkes, J. B., and Cain, C. A., 2003, “Microbubble-Enhanced Cavitation for Noninvasive Ultrasound Surgery,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 50(10), pp. 1296–1304. [PubMed]
Quanyi, L., Liyuan, F., Yan, Q., Faqi, L., and Zhibiao, W., 2008, “Role of Acoustic Interface Layer during High Intensity Focused Ultrasound Therapeutics,” J. Med. Coll. PLA, 23(4), pp. 223–227.
Ho, V. H. B., Smith, M. J., and Slater, N. K. H., 2011, “Effect of Magnetite Nanoparticle Agglomerates on the Destruction of Tumor Spheroids Using High Intensity Focused Ultrasound,” Ultrasound Med. Biol, 37(1), pp. 169–175. [PubMed]
Sun, Y., Zheng, Y., Ran, H., Zhou, Y., Shen, H., Chen, Y., Chen, H., Krupka, T. M., Li, A., Li, P., Wang, Z., and Wang, Z., 2012, “Superparamagnetic PLGA-Iron Oxide Microcapsules for Dual-Modality US/MR Imaging and High Intensity Focused US Breast Cancer Ablation,” Biomaterials, 33(24), pp. 5854–5864. [PubMed]
Wang, X., Chen, H., Zheng, Y., Ma, M., Chen, Y., Zhang, K., Zeng, D., and Shi, J., 2013, “Au-Nanoparticle Coated Mesoporous Silica Nanocapsule-Based Multifunctional Platform for Ultrasound Mediated Imaging, Cytoclasis and Tumor Ablation,” Biomaterials, 34(8), pp. 2057–2068. [PubMed]
Sapareto, S. A., and Dewey, W. C., 1984, “Thermal Dose Determination in Cancer Therapy,” Int. J. Radiat. Oncol., Biol., Phys., 10(6), pp. 787–800.
Morris, H., Rivens, I., Shaw, A., and Haar, G. T., 2008, “Investigation of the Viscous Heating Artefact Arising from the Use of Thermocouples in a Focused Ultrasound Field,” Phys. Med. Biol., 53(17), pp. 4759–4776. [PubMed]
Righetti, R., Kallel, F., Stafford, R. J., Price, R. E., Krouskop, T. A., Hazle, J. D., and Ophir, J., 1999, “Elastographic Characterization of HIFU-Induced Lesions in Canine Livers,” Ultrasound Med. Biol., 25(7), pp. 1099–1113. [PubMed]
Dąbek, L., Hornowski, T., Józefczak, A., and Skumiel, A., 2013, “Ultrasonic Properties of Magnetic Nanoparticles with an Additional Biocompatible Dextrane Layer,” Arch. Acoust., 38(1), pp. 93–98.
Hariharan, P., Dibaji, S. A. R., Banerjee, R. K., Nagaraja, S., and Myers, M. R., 2014, “Evaluation of Targeting Accuracy in Focused-Ultrasound Procedures, Using Remote Thermocouple Arrays,” J. Acoust. Soc. Am. (submitted).
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Ter Haar, G., 2001, “Acoustic Surgery,” Phys. Today, 54(12), pp. 29–34.
Wu, J., and Du, G., 1990, “Temperature Elevation Generated by a Focused Gaussian Beam of Ultrasound,” Ultrasound Med. Biol., 16(5), pp. 489–498. [PubMed]
Curra, F. P., Mourad, P. D., Khokhlova, V. A., Cleveland, R. O., and Crum, L. A., 2000, “Numerical Simulations of Heating Patterns and Tissue Temperature Response Due to High-Intensity Focused Ultrasound,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 47(4), pp. 1077–1089. [PubMed]
Soneson, J. E., 2009, “A User-Friendly Software Package for HIFU Simulation,” AIP Conf. Proc., 1113(1), pp. 165–169.
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Dibaji, S. A. R., Banerjee, R. K., Soneson, J. E., and Myers, M. R., 2013, “Nonlinear Derating of High-Intensity Focused Ultrasound Beams Using Gaussian Modal Sums,” J. Acoust. Soc. Am., 134(5), pp. 3435–3445. [PubMed]
Bailey, M. R., Couret, L. N., Sapozhnikov, O. A., Khokhlova, V. A., ter Haar, G., Vaezy, S., Shi, X., Martin, R., and Crum, L. A., 2001, “Use of Overpressure to Assess the Role of Bubbles in Focused Ultrasound Lesion Shape in vitro,” Ultrasound Med. Biol., 27(5), pp. 695–708. [PubMed]
McLaughlan, J., Rivens, I., Leighton, T., and ter Haar, G., 2010, “A Study of Bubble Activity Generated in Ex Vivo Tissue by High Intensity Focused Ultrasound,” Ultrasound Med. Biol., 36(8), pp. 1327–1344. [PubMed]
Kyriakou, Z., Corral-Baques, M. I., Amat, A., and Coussios, C.-C., 2011, “HIFU-Induced Cavitation and Heating in Ex Vivo Porcine Subcutaneous Fat,” Ultrasound Med. Biol., 37(4), pp. 568–579. [PubMed]
Dasgupta, S., Das, P., Wansapura, J., Hariharan, P., Pratt, R., Witte, D., Myers, M. R., and Banerjee, R. K., 2011, “Reduction of Noise From MR Thermometry Measurements During HIFU Characterization Procedures,” ASME J. Nanotechnol. Eng. Med., 2(2), p. 024501.
Mesiwala, A. H., Farrell, L., Wenzel, H. J., Silbergeld, D. L., Crum, L. A., Winn, H. R., and Mourad, P. D., 2002, “High-Intensity Focused Ultrasound Selectively Disrupts the Blood-Brain Barrier in vivo,” Ultrasound Med. Biol., 28(3), pp. 389–400. [PubMed]
Solomon, S. B., Nicol, T. L., Chan, D. Y., Fjield, T., Fried, N., and Kavoussi, L. R., 2003, “Histologic Evolution of High-Intensity Focused Ultrasound in Rabbit Muscle,” Invest. Radiol., 38(5), pp. 293–301. [PubMed]
Köhler, M. O., Mougenot, C., Quesson, B., Enholm, J., Le Bail, B., Laurent, C., Moonen, C. T. W., and Ehnholm, G. J., 2009, “Volumetric HIFU Ablation Under 3D Guidance of Rapid MRI Thermometry,” Med. Phys., 36(8), pp. 3521–3535. [PubMed]
Quesson, B., Laurent, C., Maclair, G., de Senneville, B. D., Mougenot, C., Ries, M., Carteret, T., Rullier, A., and Moonen, C. T. W., 2011, “Real-Time Volumetric MRI Thermometry of Focused Ultrasound Ablation in vivo: A Feasibility Study in Pig Liver and Kidney,” NMR Biomed., 24(2), pp. 145–153. [PubMed]
Canney, M. S., Bailey, M. R., Crum, L. A., Khokhlova, V. A., and Sapozhnikov, O. A., 2008, “Acoustic Characterization of High Intensity Focused Ultrasound Fields: A Combined Measurement and Modeling Approach,” J. Acoust. Soc. Am., 124(4), pp. 2406–2420. [PubMed]
Chen, D., Fan, T., Zhang, D., and Wu, J., 2009, “A Feasibility Study of Temperature Rise Measurement in a Tissue Phantom as an Alternative Way for Characterization of the Therapeutic High Intensity Focused Ultrasonic Field,” Ultrasonics, 49(8), pp. 733–742. [PubMed]
Farny, C. H., Holt, R. G., and Roy, R. A., 2009, “Temporal and Spatial Detection of HIFU-Induced Inertial and Hot-Vapor Cavitation With a Diagnostic Ultrasound System,” Ultrasound Med. Biol., 35(4), pp. 603–615. [PubMed]
King, R. L., Yunbo, L., Maruvada, S., Herman, B. A., Wear, K. A., and Harris, G. R., 2011, “Development and Characterization of a Tissue-Mimicking Material for High-Intensity Focused Ultrasound,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 58(7), pp. 1397–1405. [PubMed]
Maruvada, S., Liu, Y., Pritchard, W., Herman, B., and Harris, G., 2012, “Comparative Study of Temperature Measurements in Ex Vivo Swine Muscle and a Tissue-Mimicking Material during High Intensity Focused Ultrasound Exposures,” Phys. Med. Biol., 57(1), pp. 1–19. [PubMed]
Kennedy, J. E., 2005, “High-Intensity Focused Ultrasound in the Treatment of Solid Tumours,” Nat. Rev. Cancer, 5(4), pp. 321–327. [PubMed]
Dasgupta, S., Wansapura, J., Hariharan, P., Pratt, R., Witte, D., Myers, M. R., and Banerjee, R. K., 2010, “HIFU Lesion Volume as a Function of Sonication Time, as Determined by MRI, Histology, and Computations,” ASME J. Biomech. Eng., 132(8), p. 081005.
Huang, J., Holt, R. G., Cleveland, R. O., and Roy, R. A., 2004, “Experimental Validation of a Tractable Numerical Model for Focused Ultrasound Heating in Flow-Through Tissue Phantoms,” J. Acoust. Soc. Am., 116(4), pp. 2451–2458. [PubMed]
ter Haar, G., 2008, “Harnessing the Interaction of Ultrasound With Tissue for Therapeutic Benefit: High-Intensity Focused Ultrasound,” Ultrasound Obstet. Gynecol., 32(5), pp. 601–604. [PubMed]
Dasgupta, S., Banerjee, R. K., Hariharan, P., and Myers, M. R., 2011, “Beam Localization in HIFU Temperature Measurements Using Thermocouples, With Application to Cooling by Large Blood Vessels,” Ultrasonics, 51(2), pp. 171–180. [PubMed]
Luo, W., Zhou, X., Tian, X., Ren, X., Zheng, M., Gu, K., and He, G., 2006, “Enhancement of Ultrasound Contrast Agent in High-Intensity Focused Ultrasound Ablation,” Adv. Ther., 23(6), pp. 861–868. [PubMed]
Tung, Y.-S., Liu, H.-L., Wu, C.-C., Ju, K.-C., Chen, W.-S., and Lin, W.-L., 2006, “Contrast-Agent-Enhanced Ultrasound Thermal Ablation,” Ultrasound Med. Biol., 32(7), pp. 1103–1110. [PubMed]
Yu, T., Wang, G., Hu, K., Ma, P., Bai, J., and Wang, Z., 2004, “A Microbubble Agent Improves the Therapeutic Efficiency of High Intensity Focused Ultrasound: A Rabbit Kidney Study,” Urol. Res., 32(1), pp. 14–19. [PubMed]
Tran, B. C., Jongbum, S., Hall, T. L., Fowlkes, J. B., and Cain, C. A., 2003, “Microbubble-Enhanced Cavitation for Noninvasive Ultrasound Surgery,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 50(10), pp. 1296–1304. [PubMed]
Quanyi, L., Liyuan, F., Yan, Q., Faqi, L., and Zhibiao, W., 2008, “Role of Acoustic Interface Layer during High Intensity Focused Ultrasound Therapeutics,” J. Med. Coll. PLA, 23(4), pp. 223–227.
Ho, V. H. B., Smith, M. J., and Slater, N. K. H., 2011, “Effect of Magnetite Nanoparticle Agglomerates on the Destruction of Tumor Spheroids Using High Intensity Focused Ultrasound,” Ultrasound Med. Biol, 37(1), pp. 169–175. [PubMed]
Sun, Y., Zheng, Y., Ran, H., Zhou, Y., Shen, H., Chen, Y., Chen, H., Krupka, T. M., Li, A., Li, P., Wang, Z., and Wang, Z., 2012, “Superparamagnetic PLGA-Iron Oxide Microcapsules for Dual-Modality US/MR Imaging and High Intensity Focused US Breast Cancer Ablation,” Biomaterials, 33(24), pp. 5854–5864. [PubMed]
Wang, X., Chen, H., Zheng, Y., Ma, M., Chen, Y., Zhang, K., Zeng, D., and Shi, J., 2013, “Au-Nanoparticle Coated Mesoporous Silica Nanocapsule-Based Multifunctional Platform for Ultrasound Mediated Imaging, Cytoclasis and Tumor Ablation,” Biomaterials, 34(8), pp. 2057–2068. [PubMed]
Sapareto, S. A., and Dewey, W. C., 1984, “Thermal Dose Determination in Cancer Therapy,” Int. J. Radiat. Oncol., Biol., Phys., 10(6), pp. 787–800.
Morris, H., Rivens, I., Shaw, A., and Haar, G. T., 2008, “Investigation of the Viscous Heating Artefact Arising from the Use of Thermocouples in a Focused Ultrasound Field,” Phys. Med. Biol., 53(17), pp. 4759–4776. [PubMed]
Righetti, R., Kallel, F., Stafford, R. J., Price, R. E., Krouskop, T. A., Hazle, J. D., and Ophir, J., 1999, “Elastographic Characterization of HIFU-Induced Lesions in Canine Livers,” Ultrasound Med. Biol., 25(7), pp. 1099–1113. [PubMed]
Dąbek, L., Hornowski, T., Józefczak, A., and Skumiel, A., 2013, “Ultrasonic Properties of Magnetic Nanoparticles with an Additional Biocompatible Dextrane Layer,” Arch. Acoust., 38(1), pp. 93–98.
Hariharan, P., Dibaji, S. A. R., Banerjee, R. K., Nagaraja, S., and Myers, M. R., 2014, “Evaluation of Targeting Accuracy in Focused-Ultrasound Procedures, Using Remote Thermocouple Arrays,” J. Acoust. Soc. Am. (submitted).

## Figures

Fig. 1

Schematic of tissue phantom with four embedded TCs

Fig. 2

Micro-CT image of tissue phantom with mNPs concentration of (a) 0%, (b) 1%, and (c) 3%. The initial assessments with Micro-CT on tissue phantoms were conducted in collaboration with Dr. Lisa Lemen and Mrs. Kathleen Lasance in Vontz Core Imaging Laboratory at the University of Cincinnati.

Fig. 3

(a) Schematic of the experimental setup showing the HIFU transducer aligned with the tissue phantom in degassed water medium. (b) Schematic of the HIFU beam positioning on T1 junction.

Fig. 4

Temperature profiles at T1 in tissue phantoms with 0%, 1%, and 3% mNPs concentrations using acoustic power of (a) 5.15 W, (b) 9.17 W, and (c) 14.26 W

Fig. 5

Temperature profiles at T2 in tissue phantoms with 0%, 1%, and 3% mNPs concentrations using acoustic power of (a) 5.15 W, (b) 9.17 W, and (c) 14.26 W

Fig. 6

Temperature profiles at T3 in tissue phantoms with 0%, 1%, and 3% mNPs concentrations using acoustic power of (a) 5.15 W, (b) 9.17 W, and (c) 14.26 W

Fig. 7

Temperature profiles at T4 in tissue phantoms with 0%, 1%, and 3% mNPs concentrations using acoustic power of (a) 5.15 W, (b) 9.17 W, and (c) 14.26 W

Fig. 8

Thermal doses at (a) T1, (b) T2, (c) T3, and (d) T4 in tissue phantoms with 0%, 1%, and 3% mNPs concentrations using acoustic powers of 5.15 W, 9.17 W, and 14.26 W

## Tables

Table 1 The increase in peak temperature at T1 in tissue phantoms with 1% and 3% mNPs concentrations with respect to tissue phantom with 0% (control) mNPs concentration. ×; times.
Table 2 The time required for the temperature at T1 to drop to 50% of its peak value
Table 3 The estimated sonication time required to obtain the thermal dose of 240 equivalent min at T1. The sonication time for complete cell necrosis for “*” can be achieved below the heating period of 30 s selected for this study.
Table 4 The ratio of the initial slope of the temperature trace, normalized by the slope for the case of 0% mNPs at each selected power. The “+” is indicative of increase in initial slope of the temperature trace signifying enhanced acoustic attenuation with elevated mNPs concentration.

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