A method of operating a magnetic resonance imaging (MRI) apparatus includes exciting a body coil of the MRI apparatus to emit a radio-frequency signal, determining a center frequency of a resonance curve of the body coil, and calculating a magnet target frequency based on the determined center frequency. A magnet is ramped to the magnet target frequency.

An ambulance-compatible magnetic resonance imaging (MRI) system for on-site emergency diagnosis includes a mid-field super-conducting head-only magnet including a bore and an active shield arranged relative to the magnet, a passive shield arranged relative to the magnet, the passive shield including a first flange arranged adjacent to a first side of the magnet bore, a second flange arranged adjacent to a second side of the magnet bore, wherein the first flange and the second flange are electrically connected to each other, and wherein the passive shield is operative to capture flux extending out from the magnet bore and return the flux to the magnet. An asymmetric head gradient assembly for generating magnetic gradient field in the mid-field super-conducting magnet is also provided, the magnetic gradient field being between 100-150 mT/m or having a slew rate between 400-800 T/m/s. The MRI system includes a receiver coil and a controller operatively coupled to the receive coil, the controller configured to produce an image based on data obtained from the receive coil. The MRI system is mountable in an ambulance vehicle.

An ambulance-compatible magnetic resonance imaging (MRI) system for on-site emergency diagnosis includes a mid-field super-conducting head-only magnet including a bore and an active shield arranged relative to the magnet, a passive shield arranged relative to the magnet, the passive shield including a first flange arranged adjacent to a first side of the magnet bore, a second flange arranged adjacent to a second side of the magnet bore, wherein the first flange and the second flange are electrically connected to each other, and wherein the passive shield is operative to capture flux extending out from the magnet bore and return the flux to the magnet. An asymmetric head gradient assembly for generating magnetic gradient field in the mid-field super-conducting magnet is also provided, the magnetic gradient field being between 100-150 mT/m or having a slew rate between 400-800 T/m/s. The MRI system includes a receiver coil and a controller operatively coupled to the receive coil, the controller configured to produce an image based on data obtained from the receive coil. The MRI system is mountable in an ambulance vehicle.

According to one embodiment, a magnetic resonance imaging system includes a first imaging apparatus, a first cooling system, a second imaging apparatus, a second cooling system and a cooling control device. The first imaging apparatus includes a first magnet configured to generate a static magnetic field. The first cooling system is configured to cool the first magnet. The second imaging apparatus includes a second magnet configured to generate a static magnetic field. The second cooling system is configured to cool the second magnet. The cooling control device is configured to switch a cooling target of each of the first cooling system and the second cooling system.

According to one embodiment, a magnetic resonance imaging system includes a first imaging apparatus, a first cooling system, a second imaging apparatus, a second cooling system and a cooling control device. The first imaging apparatus includes a first magnet configured to generate a static magnetic field. The first cooling system is configured to cool the first magnet. The second imaging apparatus includes a second magnet configured to generate a static magnetic field. The second cooling system is configured to cool the second magnet. The cooling control device is configured to switch a cooling target of each of the first cooling system and the second cooling system.

An NMR measuring arrangement (20) includes a cryostat (1), a superconducting magnet coil system (2) and an NMR probe (3). The cryostat has an evacuated vacuum container (5) and forms a bore (10). A wall (12) of the bore delimits the vacuum container. The cryostat forms only one evacuated gap (13) in a space (18) between the magnet coil system and the wall of the bore. At least a segment of the wall of the bore is thermally coupled to a heat sink of the cryostat. As a result, the NMR measurement arrangement provides more precise NMR measurements (in particular with a higher spectral resolution and/or a higher signal-to-noise ratio) on measurement samples.

An NMR measuring arrangement (20) includes a cryostat (1), a superconducting magnet coil system (2) and an NMR probe (3). The cryostat has an evacuated vacuum container (5) and forms a bore (10). A wall (12) of the bore delimits the vacuum container. The cryostat forms only one evacuated gap (13) in a space (18) between the magnet coil system and the wall of the bore. At least a segment of the wall of the bore is thermally coupled to a heat sink of the cryostat. As a result, the NMR measurement arrangement provides more precise NMR measurements (in particular with a higher spectral resolution and/or a higher signal-to-noise ratio) on measurement samples.

A superconducting magnet (10) includes a cryogenic container (22, 32) containing a superconducting magnet winding (20). A sealed electrical feedthrough (36) passes through the cryogenic container. A contactor (40) inside the cryogenic container has an actuator (42) and feedthrough-side and magnet-side electrical terminals (46, 47). A high temperature superconductor (HTS) lead (60) also disposed in the cryogenic container has a first end (62) electrically connected with the magnet-side electrical terminal of the contactor and a second end (64) electrically connected to the superconducting magnet winding. A first stage thermal station (52) thermally connected with the first end of the HTS lead has a temperature (T1) lower than the critical temperature (TC,HTS) of the HTS lead. A second stage thermal station (54) thermally connected with the second end of the HTS lead has a temperature (T2) lower than a critical temperature (TC) of the superconducting magnet winding (20).

A superconducting magnet (10) includes a cryogenic container (22, 32) containing a superconducting magnet winding (20). A sealed electrical feedthrough (36) passes through the cryogenic container. A contactor (40) inside the cryogenic container has an actuator (42) and feedthrough-side and magnet-side electrical terminals (46, 47). A high temperature superconductor (HTS) lead (60) also disposed in the cryogenic container has a first end (62) electrically connected with the magnet-side electrical terminal of the contactor and a second end (64) electrically connected to the superconducting magnet winding. A first stage thermal station (52) thermally connected with the first end of the HTS lead has a temperature (T1) lower than the critical temperature (TC,HTS) of the HTS lead. A second stage thermal station (54) thermally connected with the second end of the HTS lead has a temperature (T2) lower than a critical temperature (TC) of the superconducting magnet winding (20).

The present application provides a method of designing a high shielding gradient coil for a planar superconducting magnetic resonance imaging (MRI) system and a gradient coil thereof, the method determines a shielding area according to an outer profile of a metal conductor around the position of the gradient coil in the planar superconducting MRI system, and performs partitioned shielding of a stray field. The constraint values of stray fields at different partitioned zones of the shielding area are adjusted according to the shielding requirements. The primary coils of both the transverse gradient coil and the longitudinal gradient coil optimized by the design method of the high shielding gradient coil contain a reverse coil, which generates a magnetic field that offsets leakage magnetic field of other coils, thus achieving the purpose of reducing the stray field of the gradient coil.