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Bones are one of the most common sites of cancer metastasis, which usually causes pain and impairs quality of life. Radiation therapy (RT) combined with opioids is the standard treatment for painful bone metastases. This treatment achieves effective pain control in 60−74% of patients1. However, limited treatment choices are available for recurrent or residual metastatic bone pain after RT. Reirradiation, surgical intervention, percutaneous cryoablation, or radiofrequency ablation and increased doses of systemic opioids and analgesics are options with limited indications and usually with side effects. Moreover, these secondary treatments have yielded unsatisfactory results: more than 40% of patients continue to experience moderate to severe bone pain after reirradiation2.
High-intensity focused ultrasound systems integrate ultrasounds from multiple angles into one spot, transferring acoustic energy at ablative temperatures of more than 65 °C3. This noninvasive technique has been used for thermal ablation at various sites and for various types of lesions4,5. Generally, focused ultrasound systems generate acoustic energy at frequencies of 200 kHz-4 MHz6,7, producing an intensity in the focal point on the order of 100-10,000 W/cm2. At these energy levels, the focused ultrasound beams trigger a rise in cell temperature over the treated volume of tissue. The temperature rise varies according to the tissue absorption coefficient, predicted using Arrhenius analysis or the Sapareto-Dewey isoeffect thermal dose relationship. To achieve better control and a more rapid temperature increase, focal volumes of 0.2−5 mm3 are suggested for each sonication. Therefore, the ablation of larger areas requires tiling of multiple sonications to cover a large volume and to create homogeneous thermal damage. In addition to causing damage as a result of thermal effects, focused ultrasound also creates microbubbles because of physical factors such as rectified diffusion in the treated area. When the size of microbubbles reaches a cutoff, they eventually implode, causing microshock waves and affecting surrounding tissues. This parallel nonthermal effect also contributes to tissue injury and tumor necrosis.
Unlike other image guidance techniques, such as ultrasound imaging, magnetic resonance (MR) imaging provides a three-dimensional image of anatomy with clear resolution images of soft tissue and quantitative temperature monitoring. The mapping software of quantitative MR thermometry can calculate the thermal change in degrees Celsius and then superimpose the respective locations onto the anatomic MR images8. By detecting the proton resonance frequency shift in water hydrogen, which corresponds to approximately 0.01 ppm per degree Celsius, the temperature-sensitive MR sequence can control energy deposition, with an accuracy of 1 °C for measurement of thermal changes, a spatial resolution of 1 mm, and a temporal resolution within 3 s9,10. With this extended software, the MR device could provide diagnostic images and also detect thermal changes within seconds, mapping these onto the anatomical images during the whole treatment course. Despite the development of such an innovative technique, few articles describe qualitative security during each treatment course. Here we aim to share our protocol and experiences with MRgFUS.