By: Dan Sperling, MD

During the past two decades, the evolution of minimal-to-noninvasive tumor ablation by high heat (hyperthermia) has been explored and improved in various organs: liver, kidney, brain, pancreas, breast and prostate are the most common organ sites in which malignant and benign tumors have been treated. Early efforts had uneven success due to the problem of mapping and tracking temperature changes. Heat distribution in healthy and tumor tissue can be affected by factors such as the size and shape of the lesion to be treated, tissue density, water or fat content. In addition, physiological changes as ablation progresses (e.g. increased blood flow, increased scatter space) may modify how the ablation energy is being transferred throughout the tissue. Uneven ablation can render the treatment ineffective and pose risks to nearby health tissue. However, the ability to accurately monitor and map temperatures may be a key element in assuring the safety and durable success of treatment. The term for such monitoring is thermometry. A noninvasive thermometry device would minimize intrusion into the patient’s body; a technology that would offer “treatment control and heating field localization” was considered greatly desirable by the hyperthermia community of the mid-1990s.[i]

Given the relative economy, convenience and portability of ultrasound devices, many researchers at that time believed it offered the greatest promise of noninvasive temperature monitoring during heat ablation. When ultrasound is applied in medicine it is called therapeutic ultrasound. Generally speaking, therapeutic ultrasound equipment is portable and relatively inexpensive compared with other image-based thermometry such as CT and MRI.

 Therapeutic ultrasound involves passing sound waves, or sonic energy, through tissue. Changing the wavelength and the velocity with which ultrasound passes into tissue creates a range of effects from viewing differences in tissue (imaging) to actually accomplishing thermal ablation (high intensity focused ultrasound). In the 1990s ultrasound imaging was used to guide the placement of radiofrequency probes (RF, or thermal ablation by running electric current through tissue). Thermocouples in the RF probes could record temperature in the immediate vicinity but were not reliable for measuring zonal temperatures or dangerous heat in nearby healthy tissue. Experiments with using ultrasound to measure differences in tissue as temperatures rose were based in physics: sound waves lose some of their energy as they strike and enter tissue (attenuation); they scatter in all directions, including backward toward the source (backscatter, tracked as a hyperechoic or echogenic effect). The waves that bounce back are recorded, and differences become visible on the monitor. Therefore, early studies were done in the laboratory with different materials such as a variety of organ tissues, muscle tissues, and phantoms (laboratory substances such as gel that are useful for calibrating energy wave levels).[ii]^,[iii] It was essential to establish two types of calibration:

1. Wavelength levels sufficient to compensate for attenuation in targeted tissues, and
2. The effect of heating tissues and how those changes appear in ultrasound images.

Subsequent studies during live ablation treatments involved correlating real-time temperatures (thermocouples attached to the ablation device tips) with changes on the ultrasound images. Initially, there were problems adapting traditional diagnostic gray scale ultrasound (called B-mode) to accurately depict the size and position of the thermal lesion; in animal studies, the images did not correspond well to pathology specimens removed after ablation. However, adapting the ultrasound during RF ablation allowed the capture and accumulation of time-shift estimates obtained at extremely brief and rapid intervals during the heating period; the improvements were based on tracking how the speed of sound changes as tissue expands due to heating, and quantifying the echo shifts in the backscatter signal over time. The cumulative time-shift estimates provided an estimate of the 3-dimensional temperature distribution at any given time. According to Varghese et al (2002), “The maximum temperatures in these thermal maps correspond closely with the invasive measurements of the temperature obtained from the thermosensors.”[iv] Thus, investigators could predict tissue temperature during ablation of the region of interest.[v]

During the past decade, ultrasound thermometry appears to have been overshadowed by the accuracy of magnetic resonance thermometry (MRI thermometry). Furthermore, much of the ablation technology materials (cryoablation or freezing, high intensity focused ultrasound, radiofrequency ablation, laser ablation) has been adapted for use in the MRI suite and within the magnet’s bore, or tunnel in which the patient is positioned. Still, given the portability and inexpensiveness of ultrasound compared with MRI, it is worth continued research and development to create a way to monitor ablation temperatures that could be used in practice setting.

[i] Seip R, Ebbini ES. Noninvasive estimation of tissue temperature response to heating fields using diagnostic ultrasound. IEEE Trans Biomed Eng. 1995 Aug;42(8):828-39.

[ii] Damianou CA, Sanghvi NT, Fry FJ, Maass-Moreno R. Dependence of ultrasonic attenuation and absorption in dog soft tissues on temperature and thermal dose. J Acoust Soc Am. 1997 Jul;102(1):628-34.

[iii] Arthur RM, Straube WL, Trobaugh JW, Moros EG. Non-invasive estimation of hyperthermia temperatures with ultrasound. Int J Hyperthermia. 2005 Sep;21(6):589-600.

[iv] Varghese T, Zagzebski JA, Chen Q, Techavipoo U, Frank G, Johnson C, Wright A, Lee FT Jr. Ultrasound monitoring of temperature change during radiofrequency ablation: preliminary in-vivo results. Ultrasound Med Biol. 2002 Mar;28(3):321-9.

[v] Arthur RM et al. Ibid.