Continuous thermometry during a hyperthermic procedure may help to correct for local differences in heat conduction and energy absorption, and thus allow optimization of the thermal therapy. Noninvasive, three‐dimensional mapping of temperature changes is feasible with magnetic resonance (MR) and may be based on the relaxation time T₁, the diffusion coefficient (D), or proton resonance frequency (PRF) of tissue water. The use of temperature‐sensitive contrast agents and proton spectroscopic imaging can provide absolute temperature measurements. The principles and performance of these methods are reviewed in this paper. The excellent linearity and near‐independence with respect to tissue type, together with good temperature sensitivity, make PRF‐based temperature MRI the preferred choice for many applications at mid to high field strength (≥ 1 T). The PRF methods employ radiofrequency spoiled gradient‐echo imaging methods. A standard deviation of less than 1°C, for a temporal resolution below 1 second and a spatial resolution of about 2 mm, is feasible for a single slice for immobile tissues. Corrections should be made for temperature‐induced susceptibility effects in the PRF method. If spin‐echo methods are preferred, for example when field homogeneity is poor due to small ferromagnetic parts in the needle, the D‐ and T₁‐based methods may give better results. The sensitivity of the D method is higher that that of the T₁ methods provided that motion artifacts are avoided and the trace of D is evaluated. Fat suppression is necessary for most tissues when T₁, D, or PRF methods are employed. The latter three methods require excellent registration to correct for displacements between scans. J. Magn. Reson. Imaging 2000;12:525–533.

Cell‐specific imaging has been proposed to increase the potential of magnetic resonance imaging (MRI) for tissue analysis. The hypothezis of the present work was that following intravenous injection of ultra‐small particle iron oxide, a contrast agent that accumulates in mononuclear phagocyte cells, macrophages with iron burden would be detectable by MRI within the central nervous system at sites of inflammatory cellular activity. In experimental autoimmune encephalomyelitis in Lewis rats (in which intense macrophage activity results from both hematogenous macrophages and activated microglia), lesions have been seen by MRI as low signal intensities related to magnetic susceptibility effects induced by iron particles. Electron microscopy has revealed the presence of such particles within the cytoplasm of cells that had the morphological aspect of macrophages. Macrophage activity imaging might increase MRI capability with regard to the in vivo pathophysiological aspects of central nervous system (CNS) diseases and might help in therapeutic trials in the numerous CNS diseases in which macrophages are involved. Magn Reson Med 41:329–333, 1999.

Continuous, real-time, 3D temperature mapping during a hyperthermic procedure may provide (i) enhanced safety by visualizing temperature maps in and around the treated region, (ii) improved efficiency by adapting local energy deposition with feedback coupling algorithms and (iii) therapy end-points based on the accumulated thermal dose. Non-invasive mapping of temperature changes can be achieved with MRI and may be based on temperature dependent MRI parameters. The excellent linearity of the temperature dependency of the proton resonance frequency (PRF) and its near-independence with respect to tissue type make the PRF-based methods the preferred choice for many applications, in particular at mid- to-high field strength (> or =0.5 T). The PRF methods employ RF-spoiled gradient echo imaging methods and incorporate fat suppression techniques for most organs. A standard deviation of less than 1 degrees C, for a temporal resolution below 1 s and a spatial resolution of approximately 2 mm is feasible for immobile tissues. Special attention is paid to methods for reducing artifacts in MR temperature mapping caused by intra-scan and inter-scan motion and motion and temperature-induced susceptibility effects in mobile tissues. Real-time image processing and visualization techniques, together with accelerated MRI acquisition techniques, are described because of their potential for therapy guidance.

The objective of hyperthermia treatment is to deliver a similar therapeutic thermal dose throughout the target volume within a minimum amount of time. We describe a noninvasive approach to this goal based on magnetic resonance imaging (MRI)‐guided focused ultrasound (FUS) with a spherical transducer that can be moved along two directions inside the bed of a clinical MR imager and that has an adjustable focal length in the third dimension. Absorption of FUS gives rise to a highly localized thermal buildup, which then spreads by heat diffusion and blood perfusion. A uniform temperature within a large target volume can be obtained using a double spiral trajectory of the transducer focal point together with constant and maximum FUS power. Differences between the real and target temperatures during the first spiral are evaluated in real time with temperature MRI and corrected for during the second spiral trajectory employing FUS focal point velocity modulation. Once a uniform temperature distribution is reached within the entire volume, FUS heating is applied only at the region's boundaries to maintain the raised temperature levels. Heat conduction, together with the design and timing of the trajectories, therefore ensures a similar thermal dose for the entire target region. Good agreement is obtained between theory and experimental results in vitro on gel phantoms, ex vivo on meat samples, and in vivo on rabbit thigh muscle. Edema in muscle was visible 1 hour after hyperthermia as a spatially uniform rise of the signal intensity in T₂‐weighted images. J. Magn. Reson. Imaging 2000;12:571–583.