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Assessment of Gold Nano-particle mediated Enhanced Hyperthermia using MR Guided High-Intensity Focused Ultrasound Ablation Procedure Surendra Balaji Devarakonda, Matthew R Myers, Matthew Lanier, Charles Dumoulin, and Rupak Kumar Banerjee Nano Lett., Just Accepted Manuscript • DOI: 10.1021/acs.nanolett.7b00272 • Publication Date (Web): 13 Mar 2017 Downloaded from http://pubs.acs.org on March 13, 2017
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Nano Letters
Assessment of Gold Nano-particle mediated Enhanced Hyperthermia using MR Guided High-Intensity Focused Ultrasound Ablation Procedure
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Surendra B. Devarakonda1, Matthew R. Myers2, Mathew Lanier3, Charles Dumoulin3, and Rupak K. Banerjee1*
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Department of Mechanical, Materials Engineering1
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College of Engineering and Applied Science,
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University of Cincinnati,
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Cincinnati, OH 45221
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Tel: 513-556-2124; Fax: 513-556-3390; email:
[email protected] 14
Division of Solid and Fluid Mechanics2
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Center for Devices and Radiological Health,
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U.S. Food and Drug Administration,
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Silver Spring, MD 20993
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Department of Radiology3
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Cincinnati Children's Hospital Medical Center,
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Cincinnati, Ohio, USA
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Keywords: High Intensity Focused Ultrasound; HIFU; Hyperthermia; Gold Nano-particles; Cancer therapy, MRI
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Abstract High-intensity focused ultrasound (HIFU) has gained increasing popularity as a non-invasive
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therapeutic procedure to treat solid tumors. However, collateral damage due to the use of high
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acoustic powers during HIFU procedures remains a challenge. The objective of this
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study is to assess the utility of using gold nano-particles (gNPs) during HIFU procedures to
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locally enhance heating at low powers, thereby reducing the likelihood of collateral damage.
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Phantoms containing tissue-mimicking material (TMM) and physiologically relevant
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concentrations (0%, 0.0625%, and 0.125%) of gNPs were fabricated. Sonications at acoustic
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powers of 10 W, 15 W, and 20 W were performed for a duration of 16 sec using an MR-HIFU
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system. Temperature rises and lesion volumes were calculated and compared for phantoms with
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and without gNPs. For an acoustic power of 10 W, the maximum temperature rise increased by
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32% and 43% for gNPs concentrations of 0.0625% and 0.125%, respectively, when compared to
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the 0% gNPs concentration. For the power of 15 W, a lesion volume of 0 mm3, 44.5 ± 7 mm3,
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and 63.4 ± 32 mm3 was calculated for the gNPs concentration of 0%, 0.0625%, and 0.125%,
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respectively. For a power of 20 W, it was found that the lesion volume doubled and tripled for
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concentrations of 0.0625% and 0.125% gNPs, respectively, when compared to the concentration
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of 0% gNPs. We conclude that gNPs have the potential to locally enhance the heating and reduce
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damage to healthy tissue during tumor ablation using HIFU.
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Keywords: High Intensity Focused Ultrasound; HIFU; Hyperthermia; Gold Nano-particles;
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Cancer therapy; MRI
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In high-intensity focused ultrasound (HIFU) ablation procedures, ultrasound energy from an
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external transducer is focused onto target tissue within the body. HIFU therapy has increasingly
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gained clinical interest as a novel tool for ablation procedures. Using focused ultrasound energy,
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temperatures greater than 60°C can be induced in tissues essentially instantaneously, achieving
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cellular necrosis and vascular cauterization1-3. A large tumor area can be treated by placing
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multiple focal zones sequentially and in close proximity until the entire area is ablated3.
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Characterization and testing of HIFU procedures has been performed using computational
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methods4-8, in-vitro tissue phantoms, and animal models9-14.
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Furusawa et al.15, performed MR-guided HIFU on patients with breast cancer and were able
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to successfully ablate tumors using acoustic powers of greater than 60 W. However, they
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reported third degree skin burns in 3% of the patients. Li et al.16, performed HIFU ablations on
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17 human patients with recurrent and metastatic abdominal tumors to determine the
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complications arising during and after the HIFU treatment. They used focal peak intensities of
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2000 to 5000 Wcm-2and the treatment time was between 3120 sec to 8950 sec. Skin burns were
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found in all 17 patients and mild enteroparalysis was found in 15 patients. These adverse events
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are possibly due to the usage of higher acoustic power. Although HIFU is experiencing higher
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clinical application, issues related to high acoustic intensity, such as skin burns and damage to
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neighboring healthy tissue17,18 remain.
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To reduce the required acoustic intensity and consequently the likelihood of adverse events
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such as skin burns, absorption-enhancing agents can be used. One promising agent is
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nanoparticles, which can be injected to the tumor site to achieve targeted and intense heating of
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the tumor cells 17. Sun et al.18, 19, reported the effects of magnetic microcapsules (PLGA-coated
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Fe3O4) during HIFU therapy, and found enhancement of hyperthermia due to the microcapsules. 3 ACS Paragon Plus Environment
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However, the authors themselves pointed out that the microcapsules size was very high (500 –
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800 nm). The power used was extremely high (180 W - 250 W) which can lead to skin burns
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and collateral damage to neighboring tissue17. Also, the MR thermometry images reported in this
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study have limited clarity regarding the temperature increase in the tumor.
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Using 1% and 3% concentrations of magnetic nanoparticles (mNPs) embedded in tissue
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mimicking material (TMM), Dibaji et al.20, measured the HIFU-induced temperature rise using
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embedded thermocouples (TCs). They determined that the peak temperature rise increased by 1.6
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and 2 times when mNPs concentration of 1% and 3% were used, respectively, for an acoustic
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power of 14.2 W. Temperature changes in HIFU can also be measured using temperature
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sensitive MR. However, the magnetism and magnetic susceptibility of mNPs have been
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reported to interfere with Magnetic Resonance Imaging (MRI). This results in decreased signal
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to noise ratio, geometric distortions and low quality images21. It may be noted that the mNPs
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concentration of 1% and 3% for the pilot study was relatively high from the perspective of
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physiological relevance.
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To help assess the efficacy of nanoparticle introduction pre-clinically, it is useful to make
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thermal measurements in tissue phantoms. HIFU temperature measurements can be performed
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using invasive modalities such as TCs, or by non-invasive modalities such as magnetic resonance
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(MR) thermometry. Direct measurement of temperature using TCs can lead to errors such as
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viscous-heating22, 23 artifacts and positioning errors24 caused by misalignment between the HIFU
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beam and the TCs. Dasgupta et al.23, have conducted HIFU experiments in phantoms using
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powers of 5 W, 10.3 W, 17.3 W, and 24.8 W and measured the temperature rise using TCs.
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They have reported that the thermal artifact can be up to 2 times the local tissue temperature rise
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when direct sonications were performed on TCs. Alternatively, with the use of MR thermometry, 4 ACS Paragon Plus Environment
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thermocouple artifacts can be avoided25-27. By optimizing the MR thermometry parameters such
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as T1 and T2 relaxation times, proton resonance frequency (PRF), and the temperature-dependent
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diffusion coefficient, temperature measurement can be achieved during HIFU sonication with
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reasonable accuracy28-32.
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In the past, researchers from our lab have performed MR thermometry on in-vivo porcine
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livers using acoustic powers of 10 W, 30 W, and 40 W20. This study reported that MR
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thermometry can be used during HIFU sonication for acquiring the maximum temperature rise of
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57 °C with localized cavitation. On a similar note, MR-guided HIFU procedures33 were
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performed in-vitro in turkey breast samples and in-vivo in transcranial rat brains with acoustic
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powers of 40 W. In this study, tissue displacements due to acoustic force were detected.
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In the present study, the effect of gold nanoparticles (gNPs) on the HIFU-induced
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temperature rise and lesion volume have been assessed using MR thermometry. To the authors’
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knowledge, quantification of temperature rise and lesion volume using MR thermometry with
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nano-particles, for example, gNPs has not been previously reported. The TMM phantoms with
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physiologically relevant concentrations of 0% (control), 0.0625%, and 0.125% gNPs by volume
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were fabricated. Low acoustic powers - 10 W, 15 W, and 20 W - were used for the HIFU
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sonications. The temperature rises and the lesion volumes were calculated from the MR
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thermometry and were compared for different concentrations of gNPs.
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Three cylindrical fixtures with a length of 3 cm and inner diameter of 2.5 cm (volume ~ 15
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cm3) were developed (Fig. 1A). The protocol reported by King et al.34 was used to fabricate the
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tissue phantom with 0% gNPs concentration. To prepare the tissue phantoms with 0.0625% and
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0.125% gNPs, the method explained in Dibaji et al.20 was used. The size of the gNPs embedded 5 ACS Paragon Plus Environment
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in the TMM was 15 nm (US Research Nanomaterials, Inc., Houston, TX). Thorough mixing
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using a magnetic stir bar was performed for 2 hours to achieve uniform distribution of the gNPs
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inside the TMM during the fabrication of the phantoms21. All the fabricated phantoms were kept
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for 12 hours at room temperature to ensure complete solidification of TMM.
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A clinical MR-HIFU system (SonalleveTM V2, Philips Medical Systems, Vantaa, Finland),
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integrated into a 1.5 Tesla (T) whole-body scanner (Philips IngeniaTM, Healthcare, Best, The
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Netherlands) was used for scanning the phantom (Fig. 1B). The MR-HIFU system contains a
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256-element phased array HIFU transducer which can be used to focus the ultrasound energy
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into small volumes within the phantom. Radius of curvature and the aperture of the transducer
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array are 120 mm and 130 mm, respectively35. The operating frequency was 1.2 MHz. Focal
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spot of the HIFU beam was approximately 2 mm in the radial direction.
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For targeting and sonicating inside the TMM, phantoms were placed above the HIFU
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transducer on the MR-HIFU tabletop in a vertical position. Scans were performed to ensure
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proper acoustic coupling and for detecting air bubbles. Any air bubbles detected near the surface
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of the TMM were removed. The phantoms were placed inside a 75 mm x 75 mm single loop
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tunnel and matched imaging coil to achieve an optimized resolution. A resolution of 0.7 mm x
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0.7 mm in the radial direction and a slice thickness of 6 mm were achieved.
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The electrical powers used by the MR-HIFU system to generate the desired acoustic powers
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were derived from radiation force balance measurements. In order to optimize the conversion of
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electrical field to acoustic waves, an automated electrical matching routine was performed by the
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system after placing the phantom inside the imaging coil.
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After aligning the phantoms inside the imaging coil, MR images using T1-weighted and T2weighted sequences were obtained. These images were obtained in both sagittal and coronal 6 ACS Paragon Plus Environment
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planes. Using the T1-weighted images, it was possible to avoid any obstacles in the path of the
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HIFU beam such as the plastic fixture of the phantoms. The imaging technique was two
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dimensional (2D) “fast field echo” (FFE) and the flip angle was 20°. The repetition time (TR)
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and the echo time (TE) were 53 msec and 23 msec, respectively. Additionally, for the T2-
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weighted sequence, the TR, TE, and flip angle were 1300 msec, 130 msec, and 90°, respectively.
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The HIFU beam was oriented inside the tissue phantom in such a way that the outer plastic
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fixture was not in the path of incoming or outgoing beam. Three acoustic powers of 10 W, 15 W,
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and 20 W were used for the HIFU sonication. A period of 16 sec was chosen for the HIFU
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sonication of the phantoms.
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During the sonication and cooling periods of 16 sec and 15 sec, respectively, MR
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thermometry was used to acquire the temperature data. The initial temperature of the phantoms
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and the surrounding medium was fixed at 37 °C. Using a fast field echo (FFE) and echo planar
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technique (EPI), the temperature maps at the focus in both radial and axial direction were
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captured. The temperature changes inside the phantom were calculated using the proton
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resonance frequency-shift (PRFS) method. The optimized temporal resolution of the
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temperature measurements was 6 sec. The maximum temperature values discussed in the
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manuscript are the volumetric average of all the temperature values inside a pixel. Temperature
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measurements were recorded from three trials (n=3) for each acoustic power (m=3).
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For cases where temperature values were desired outside the 6-second intervals, the
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temperature values obtained from the MR thermometry were curve fit to a function representing
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a solution to the heat equation. The functional form of the fitting function is exponential-
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integral36. The exponential-integral solution to the heat equation was derived assuming a heat
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source with a Gaussian profile in radial direction37. The heat source is assumed to be constant in 7 ACS Paragon Plus Environment
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the axial direction over the distances of interest. The solution is applicable to cases where the
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radial dimension of the heat source is much smaller than the axial dimension, as is the case with
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the HIFU beam. The temperature distribution as a function of time at a distance r from the beam
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axis is:
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167
, =
−
!
(1a)
where Ei(x) is the exponential integral: ( $ %&
" = − #)
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'
*+
(1b)
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where r is the radial coordinate, r0 is the beam radius width of the Gaussian intensity distribution,
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α is the absorption coefficient (dB/m), ,- is the intensity on the beam axis (W/mm3), κ is the
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thermal diffusivity (m2/s), ρ0 is the density (kg/m3), and cp is the specific heat (J/kg.K) of TMM.
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The properties of the TMM are presented in Table 1.
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In order to calculate the lesion volume, the MR-HIFU algorithm calculates the thermal dose
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using the method developed by Sapareto and Dewey 38. The thermal dose parameter is given by
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3
. ", /, 0 = # 3- 45678 1 .2 *
(2)
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where . is the thermal dose at the reference temperature of 43 ºC, 9:; M 0.25 IℎKL+K
(3)
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Thermal dose contours calculated by the treatment planning algorithm for the temperature maps
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were used in determining the region that has been subjected to a thermal dose of at least 240
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equivalent minutes. This volume is defined as the lesion volume. Line contours in the middle 8 ACS Paragon Plus Environment
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zone of the temperature maps were measured using the MR software across the focal diameter
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(d) and along the beam axial path (l) at the focus. Symmetry was assumed in the radial direction
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as the properties of the TMM did not vary in the radial direction. These dimensions were then
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used in equation 2, which calculates the lesion volume by assuming an ellipsoid shape for the
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lesions inside the phantom:
N = P* Q O
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(4)
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Here d (mm) is the focal diameter of the lesion in radial direction, l (mm) is the length of the
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lesion in axial direction. Lesions were measured from three trials (n=3) for each acoustic power
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(m=3). Uncertainties in lesion volume represent the standard deviation of the volumes computed
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from the three measurements.
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In addition to temperature rise and lesion volume, absorbed ultrasound energy due to the
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presence of gNPs was also calculated. This was computed from the initial temperature values
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calculated from the MR thermometry, as heat conduction can be neglected for initial duration.
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Thus, temperature rise is related to the absorbed energy Q by:
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U2
R- ST U = V = 2 W ,
(5)
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The ratio of the absorbed ultrasound energy with and without gNPs was computed for the
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different powers considered. Statistical analysis has been performed on the maximum
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temperature rise and lesion volume data using a t-test to determine if the enhancement of heating
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by gNPs is statistically significant.
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At each power level (10 W, 15 W, 20 W), the measured temperatures were averaged over
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three trials (n=3). Results are presented as mean ± SD. The temperature maps (Fig. 2) show the 9 ACS Paragon Plus Environment
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variation of temperature for the concentrations of 0%, 0.0625%, and 0.125% gNPs and for the
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powers of 10 W, 15 W, and 20 W at the plane of maximum pixel temperature. A trending
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increase in temperature (increase in number of red pixels) can be seen in moving from left to
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right (increasing gNP) concentration, particularly at the two lower powers. Figure 3 shows the
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temperature traces derived using the exponential-integral function (1) fitted using the MR
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thermometry values. The temperature rise at 7 seconds (the second MR data value) for the
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0.0625% gNP concentration was about 1.1 times that for no gNP’s, and for 0.125% gNP
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concentration the ratio was about 1.2. The acoustic power was 10 W.
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increase in temperature rise at the end of 7 seconds of sonication was nearly 1.2 times and 1.6
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times for the concentration of 0.0625% and 0.125% gNPs, respectively, when compared to 0%
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gNPs. Finally, for a power of 20 W, the increase in temperature rise for the initial 7 seconds of
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sonication was nearly 1.2 times and 1.5 times for the concentration of 0.0625% and 0.125%
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gNPs, respectively, when compared to 0% gNPs.
For a power of 15 W, the
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Figure 4A shows the comparison of maximum temperature rise, which occurs at the end of
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the 13-second sonication time. For a power of 10 W, the maximum temperature rises of 16±1
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°C, 21±3 °C, and 23±3 °C were observed for the concentrations of 0%, 0.0625%, and 0.125%
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gNPs, respectively. The changes relative to 0% gNP’s were 32% and 43% for the two phantoms
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with gNP’s. For a power of 15 W, the maximum temperature rises of 25±0.7 °C, 29±4 °C, and
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36±5 °C were observed for the concentrations of 0%, 0.0625%, and 0.125% gNPs. The
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increases relative to baseline (0% gNP) were 16% and 43% for the 0.0625% and 0.125% gNP
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values. For a power of 20 W, the maximum temperature rises of 34±3 °C (0% gNP), 43±0.4 °C
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(0.0625% gNP), and 46±0.7 °C (0.125% gNP). The increase relative to baseline for the 0.0625%
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concentration was 24% and for the 0.125% concentration was 34%. It is also evident that the
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enhancement of heating by gNPs is statistically significant (p < 0.05).
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The amount of absorbed ultrasound energy, determined using an initial time of 1 sec for
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slope calculation, is shown in Fig. 4B. For the power of 10 W, the relative energy absorption
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was similar for all the concentrations. It is possible that the temperature rise for the time period
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of 1 sec is very small in relation to 6 sec, particularly for the lower power of 10 W. Therefore,
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the small temperature rise may have remained undetected by the MRI sensor. Consequently, the
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change in the slope of temperature rise between no gNPs (0%) and with gNPs (00625% and
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0.125%) for the power of 10 W is not apparent. For the power of 15 W, approximately 2 times
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more energy was absorbed with the 0.0625% gNPs concentration than with no gNPs, and 2.7
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times more energy relative to control for the 0.125% concentration. For the 20 W power level,
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the relative energy absorption was 4 for the 0.0625% gNPs concentration and 5.7 for the 0.125%
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gNPs concentration.
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Figure 4C presents the lesion volumes obtained for the 0%, 0.0625%, and 0.125% gNPs
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concentrations for powers of 10 W, 15 W, and 20 W. The white line contours in the middle zone
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of the temperature maps (Fig 2) indicate the predicted lesion area at that cross-section. The
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lesion area has been exposed to at least a thermal dose of 240 equivalent minutes41. To calculate
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the predicted lesion volumes, the white lines from the temperature maps in both radial and axial
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directions have been measured.
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For the concentration of 0% gNPs, at powers of 10 W (Fig 2A1) and 15 W (Fig 2B1),
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insufficient heat was generated to produce a predicted lesion. For the power of 15 W, a predicted
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lesion volume of 44.5 ± 7 mm3 and 63.4 ± 32 mm3 was calculated for the gNPs concentration of
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0.0625% and 0.125%, respectively. For 20 W, the lesion volume increased roughly two and three 11 ACS Paragon Plus Environment
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times when gNPs concentration of 0.0625% and 0.125% were used compared to 0% gNPs
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concentration.
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For the same power, the lesion volume increased nearly 3 fold with the addition of gNP’s
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(Fig. 4C). Thus, it can be observed from Fig. 4A and 4C that the enhancement of heating by
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gNPs (0.0625% and 0.125%) is statistically significant (p < 0.05) when compared to 0% gNPs at
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13 sec time point. Through the use of lower acoustic powers (e.g. 15 W) and shorter sonication
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times afforded by optimized concentrations of gNP’s, undesired damage can potentially be
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reduced in a clinical setting. Currently, researchers have used acoustic powers in excess of 60 W
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- 250 W to achieve cell necrosis during HIFU procedures15.
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The enhancement of ablation efficiency does not appear to be linear with gNP concentration.
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For example, for a power of 15 W, the relative absorbed energy (Fig. 4B) is 2.0 for a gNP
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concentration of 0.0625%, but doubling this concentration increases the absorbed energy to only
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2.7. A similar trend holds at 20 W. Likewise, a 0.0625% gNP concentration produces a lesion
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volume of 44 mm3, but doubling the gNP concentration increases the lesion volume by a factor
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of 1.4 (power = 15 W). At 20 W, doubling the gNP concentration also leads to a factor of 1.4
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increase in lesion volume. Based upon the current data of our study, enhancement in ablation
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efficiency scales roughly as the square root of the gNP concentration.
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Temperature rise in the presence of gNP’s is a linear function of power, as it is for no gNP’s
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(Fig. 3A). In that sense, the effect of incorporating gNP’s in tissue volume can be modeled as a
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local increase in acoustic absorption.
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The assumption of an ellipsoid shape for a lesion at higher powers may lead to some error in
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calculation of lesion volume. This is because the ellipsoid shape is known to get distorted due to
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cavitation at higher energy deposition, although no cavitation was detected in the present
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experiments.
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The temperatures obtained from the MR thermometry are an average of all the temperatures
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with in a pixel. To estimate the effect of averaging, the temperature distribution was assumed to
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possess the same functional form as the intensity distribution. Such an approximation would be
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most accurate at the beginning of sonication, prior to significant heat conduction. For the
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calculations, the intensity profile provided by Wu and Du39 (based on a Gaussian-beam
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assumption) was used. Averaging over the 6 mm pixel thickness in the axial direction results in
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a temperature error of approximately 4.2% at the focus. Away from the focus the error is less, as
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a change in temperature in the +z direction (relative to the center of the pixel) is partially
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compensated by a change of the opposite sign in the –z direction. In the radial direction, the 0.7
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mm resolution results in an averaging error of approximately 1%.
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The HIFU absorption in media embedded with gNPs depend on the thermal processes within
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the viscous and phonon layers at the interface of gNPs as well as on the intrinsic absorption
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properties of the media40-43. Propagating HIFU waves in a medium interact with the thermal
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phonons and consequently, a part of the wave is absorbed. Due to the wave absorption, there is
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an increase in the momentum of thermal phonons inside the medium leading to temperature rise.
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It is expected that the attenuation due to phonon layer is the dominating mechanism for the gNPs
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size of 15 nm leading to temperature rise, but further analysis is required. Such analysis may
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need an independent theoretical-experimental characterization study that we have embarked on.
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Gold nanoparticles can be made to accumulate at a desired location in the tumor tissue by
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injecting the nanoparticles directly into the tumor tissue intravenously. The gNPs are expected
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to accumulate in the tumor tissue due to the enhanced permeability and retention effect (EPR)44-
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performed using nanoparticles injected into the tumor site intravenously. Also, localized
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injection of nanoparticles could be performed in superficial tumors. MR thermometry can then
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be performed similar to the work presented in this study to evaluate the influence of gNPs.
. To demonstrate this technique, future in-vivo experiments, preferably in animals, can be
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Corresponding Author
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Rupak K. Banerjee
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Department of Mechanical, Materials Engineering1
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College of Engineering and Applied Science,
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University of Cincinnati,
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Cincinnati, OH 45221
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Tel: 513-556-2124; Fax: 513-556-3390; email:
[email protected] 307
Author contributions
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Surendra B. Devarakonda performed the experiments and wrote the manuscript. Rupak Banerjee
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supervised the project. Mathew R. Myers, Mathew Lanier, Charles Dumoulin and Rupak K.
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Banerjee reviewed and revised the manuscript.
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Competing financial interests
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The authors declare no competing financial interests.
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Funding Sources
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This work was supported by NSF grant 1403356.
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Acknowledgements
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This work was supported by NSF grant 1403356.
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Table 1: Properties of the TMM. Property
Value
Density, R-
1040 kg/m3
Absorption Coefficient, α Specific Heat, Cp
45 dB/m 4064 J/kg.K 1.4 x 10-7 m2/s
Thermal Diffusivity, κ 318 319
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A
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B
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Figure 1: A) MR image of the phantom filled with tissue-mimicking material (TMM) (diameter
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of 25 mm) and B) MR-HIFU setup - vertical cross-sectional view of the TMM within MR coil
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Figure 2: Representative temperature maps of TMM phantoms at maximum temperature with
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0%, 0.0625%, and 0.125% gNPs concentration for the acoustic powers of 10 W, 15 W, and 20
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W. The line contours in the center shows calculated lesion volume.
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Figure 23: Comparison of temporal variation of temperature rise (°C) using MR thermometry for the concentrations of 0%, 0.0625%, and 0.125% gNPs and powers of A) 10 W, B) 15 W, and C) 20 W. The sonication period was 16 sec. Measurements for each time point was conducted in triplicate (n=3). 19 ACS Paragon Plus Environment
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Figure 4: A) Comparison of maximum temperature rise (°C) for 0%, 0.0625%, and 0.125% gNPs concentration for powers of 10 W, 15 W, and 20 W, B) The ratio of the initial slope of the temperature trace, normalized by the slope for the case of 0% gNPs at each selected (10 W, 15 W, and 20 W) power, and C) Comparison of lesion volume (mm3) for 0%, 0.0625%, and 0.125% gNPs concentration for the acoustic powers of 15 W and 20 W. 20 ACS Paragon Plus Environment
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Table of contents Graphic
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References:
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