Figures & data
Table 1. Overview of relevant parameters for temperature error estimations: static magnetic field B0, chemical shift δ, permittivity ε, electric conductivity σel, permeability μ and magnetic volume susceptibility χ. The time dependent changes in B0 are based on a magnetic field drift of the magnet over time. Electromagnetic property (ɛ, σ, μ) changes with temperature influence phase velocity of the RF pulse, B1 and RF coil loading introducing additional echo time dependent phase changes. Magnetic susceptibility might also change due to perfusion and diffusion changes.
Figure 1. (a–c) Clinical example of MR-guided (B0 = 1.5 T) RF hyperthermia treatment of recurrent rectal carcinoma. (a) T1-weighted anatomical reference. (b) Uncorrected temperature distribution together with the contours delineating anatomical fat regions used for phase drift correction. (c) Corrected temperature distribution together with the contours delineating the tumour. For PRF thermometry, spoiled gradient echo images were acquired after an RF heating duration of 57 min, and temperature was calculated with regard to a reference phase map prior to the thermal therapy [Citation49]. (d–f) Influence of perfusion changes on PRF thermometry in comparison to fibre-optic temperature sensor readings in RF hyperthermia treatment of soft tissue sarcoma of the lower extremities. (d) T1-weighted anatomical reference and location of the fibre-optic temperature reading track. (e) PRF thermometry showing the temperature hotspot. For PRF thermometry drift corrected phase difference of spoiled gradient echo images towards a reference prior to the thermal treatment was evaluated [Citation106]. (f) Comparison of PRF thermometry and fibre-optic thermometry. Sincere under- and overestimation in PRF thermometry in the presence of perfusion decrease and increase can be seen. In the muscle reactive perfusion increases during RF hyperthermia (on the left side away from the tip). In tumours no or low reactive perfusion is expected, showing very good accuracy (<1 °C) between fibre-optic measurements and PRF thermometry. At the tumour margin perfusion declines (at the tip of the catheter), probably caused by the adjacent hyperperfused muscle (steal effect).
![Figure 1. (a–c) Clinical example of MR-guided (B0 = 1.5 T) RF hyperthermia treatment of recurrent rectal carcinoma. (a) T1-weighted anatomical reference. (b) Uncorrected temperature distribution together with the contours delineating anatomical fat regions used for phase drift correction. (c) Corrected temperature distribution together with the contours delineating the tumour. For PRF thermometry, spoiled gradient echo images were acquired after an RF heating duration of 57 min, and temperature was calculated with regard to a reference phase map prior to the thermal therapy [Citation49]. (d–f) Influence of perfusion changes on PRF thermometry in comparison to fibre-optic temperature sensor readings in RF hyperthermia treatment of soft tissue sarcoma of the lower extremities. (d) T1-weighted anatomical reference and location of the fibre-optic temperature reading track. (e) PRF thermometry showing the temperature hotspot. For PRF thermometry drift corrected phase difference of spoiled gradient echo images towards a reference prior to the thermal treatment was evaluated [Citation106]. (f) Comparison of PRF thermometry and fibre-optic thermometry. Sincere under- and overestimation in PRF thermometry in the presence of perfusion decrease and increase can be seen. In the muscle reactive perfusion increases during RF hyperthermia (on the left side away from the tip). In tumours no or low reactive perfusion is expected, showing very good accuracy (<1 °C) between fibre-optic measurements and PRF thermometry. At the tumour margin perfusion declines (at the tip of the catheter), probably caused by the adjacent hyperperfused muscle (steal effect).](/cms/asset/1082292d-9091-496b-bce4-8f9956dde6d1/ihyt_a_1108462_f0001_c.jpg)
Figure 2. (a) Echo time-dependent magnitude and phase images at 7.0 T of a coronary stent placed in an agar phantom. Signal voids and phase wraps evolve with prolonged echo times. (b) MR photograph of a coronary stent (left) used in RF heating experiments. PRF thermometry through the plane of the stent (middle) and in the plane slightly above the stent (right) using a double echo method (TE1 and TE2), where imaging artefacts are reduced. (c) Comparison of MR thermometry and fibre-optic thermometry in regions of the stent and surface regions of a homogeneous phantom [Citation87].
![Figure 2. (a) Echo time-dependent magnitude and phase images at 7.0 T of a coronary stent placed in an agar phantom. Signal voids and phase wraps evolve with prolonged echo times. (b) MR photograph of a coronary stent (left) used in RF heating experiments. PRF thermometry through the plane of the stent (middle) and in the plane slightly above the stent (right) using a double echo method (TE1 and TE2), where imaging artefacts are reduced. (c) Comparison of MR thermometry and fibre-optic thermometry in regions of the stent and surface regions of a homogeneous phantom [Citation87].](/cms/asset/1977bb56-2a3e-4bc6-bf38-f9816354c9d1/ihyt_a_1108462_f0002_c.jpg)
Figure 3. Schematic of a hybrid MR-guided RF hyperthermia system [Citation1] and an integrated system presented recently [Citation31]. While the hybrid system requires two separate RF amplifiers and control units plus additional hardware filters that need to be incorporated into the standard MR hardware, the integrated system utilises the proton excitation frequency of the MR system as well as its pulsed power amplifier and control unit to perform RF-induced heating. Only transmit receive switches are added to the integrated system, which can be done externally as an integral part of the applicator and does not involve changes in the standard MR architecture itself [Citation31].
![Figure 3. Schematic of a hybrid MR-guided RF hyperthermia system [Citation1] and an integrated system presented recently [Citation31]. While the hybrid system requires two separate RF amplifiers and control units plus additional hardware filters that need to be incorporated into the standard MR hardware, the integrated system utilises the proton excitation frequency of the MR system as well as its pulsed power amplifier and control unit to perform RF-induced heating. Only transmit receive switches are added to the integrated system, which can be done externally as an integral part of the applicator and does not involve changes in the standard MR architecture itself [Citation31].](/cms/asset/7fb4fbb4-92ee-4ef3-a808-4709fcef10d5/ihyt_a_1108462_f0003_c.jpg)
Figure 4. SNR comparison of a 16-channel bow tie dipole array [Citation121] and a 32-channel loop coil array [Citation125] for cardiac imaging at 7.0 T. A 2D CiNE FLASH of a short axis view (SAX) (1.1×1.1×2.5) mm2 and its corresponding SNR map [Citation123,Citation124] is displayed for a GRAPPA reduction factor R = 2. The 16-channel bow tie dipole array shows higher mean SNR values as compared to the 32-channel loop coil indicating its feasibility for MRI at UHF (B0 = 7.0 T) [Citation121]. The excellent image quality obtained from using dipole antennas for MRI demonstrates the rationale of these elements to be used for an integrated system.
![Figure 4. SNR comparison of a 16-channel bow tie dipole array [Citation121] and a 32-channel loop coil array [Citation125] for cardiac imaging at 7.0 T. A 2D CiNE FLASH of a short axis view (SAX) (1.1×1.1×2.5) mm2 and its corresponding SNR map [Citation123,Citation124] is displayed for a GRAPPA reduction factor R = 2. The 16-channel bow tie dipole array shows higher mean SNR values as compared to the 32-channel loop coil indicating its feasibility for MRI at UHF (B0 = 7.0 T) [Citation121]. The excellent image quality obtained from using dipole antennas for MRI demonstrates the rationale of these elements to be used for an integrated system.](/cms/asset/08942fb0-19b3-4204-869f-0ed2fd272d87/ihyt_a_1108462_f0004_c.jpg)