WO2015189786A1 - Système d'imagerie à résonance magnétique transportable - Google Patents

Système d'imagerie à résonance magnétique transportable Download PDF

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Publication number
WO2015189786A1
WO2015189786A1 PCT/IB2015/054395 IB2015054395W WO2015189786A1 WO 2015189786 A1 WO2015189786 A1 WO 2015189786A1 IB 2015054395 W IB2015054395 W IB 2015054395W WO 2015189786 A1 WO2015189786 A1 WO 2015189786A1
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Prior art keywords
magnet
coils
superconducting magnet
superconducting
vacuum vessel
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PCT/IB2015/054395
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English (en)
Inventor
Robert Slade
Benjamin PARKINSON
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Victoria Link Ltd
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Priority to CN201580038720.XA priority Critical patent/CN106716166A/zh
Priority to BR112016028985A priority patent/BR112016028985A2/pt
Publication of WO2015189786A1 publication Critical patent/WO2015189786A1/fr

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/381—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets
    • G01R33/3815—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets with superconducting coils, e.g. power supply therefor
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01F—MAGNETS; INDUCTANCES; TRANSFORMERS; SELECTION OF MATERIALS FOR THEIR MAGNETIC PROPERTIES
    • H01F6/00—Superconducting magnets; Superconducting coils
    • H01F6/04—Cooling
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01F—MAGNETS; INDUCTANCES; TRANSFORMERS; SELECTION OF MATERIALS FOR THEIR MAGNETIC PROPERTIES
    • H01F6/00—Superconducting magnets; Superconducting coils
    • H01F6/06—Coils, e.g. winding, insulating, terminating or casing arrangements therefor
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/3802—Manufacture or installation of magnet assemblies; Additional hardware for transportation or installation of the magnet assembly or for providing mechanical support to components of the magnet assembly
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/3804—Additional hardware for cooling or heating of the magnet assembly, for housing a cooled or heated part of the magnet assembly or for temperature control of the magnet assembly
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/385—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using gradient magnetic field coils
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G01R33/3875—Compensation of inhomogeneities using correction coil assemblies, e.g. active shimming
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/42—Screening
    • G01R33/421—Screening of main or gradient magnetic field
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/42—Screening
    • G01R33/422—Screening of the radio frequency field

Definitions

  • This disclosure relates to a transportable magnetic resonance imaging (MRI) system.
  • MRI magnetic resonance imaging
  • Magnetic resonance imaging is a medical imaging technique for investigating the anatomy and function of the body of healthy and diseased patients.
  • MRI involves placing a patient in a strong magnetic field, exciting the patient with a radio-frequency signal to induce nuclear magnetic resonance, receiving a radio- frequency signal indicating the nuclear magnetic resonance, and producing an image of the patient from the received radio-frequency signal.
  • spatial information is encoded into the received radio-frequency signal by using pulsed magnetic field gradients.
  • MRI does not expose the patient to ionizing radiation, and provides good imaging of soft tissue.
  • MRI systems may use either permanent magnets or superconducting magnets, and may be designed for either whole body scanning or for scanning of just a body part such as a head, arm, or leg.
  • Permanent magnet MRI systems are commonly configured as "C" magnets, also called an open system in which a patient's body part to be imaged is inserted between the poles of the permanent magnet.
  • C C magnets
  • the practically achievable magnetic field strength of a whole body permanent magnet is limited to between 0.2 and 0.4 Tesla (“low field”) by the intrinsic properties of permanent magnet materials and practical weight limits.
  • a superconducting magnet is most commonly configured so that the MRI system is closed around the patient's body part to be imaged.
  • the superconducting magnet is in the form of a solenoid having a bore into which a patient's body part is inserted.
  • the superconducting magnet typically is a low temperature superconducting (LTS) solenoid, bath-cooled using liquid helium.
  • Operating field is typically 1.5 T to 3 T (“high field”), and magnet weight is in the range of 5-10 metric tons for whole body scanning.
  • LTS low temperature superconducting
  • a liquid helium bath-cooled magnet is complex, and system cost is typically two to three times that of a permanent magnet, but the higher field of the superconducting magnet provides higher signal-to- noise which can be used to provide increased image resolution or shorter scan times.
  • Transportable MRI systems are of interest for rapid deployment in disaster relief and military applications, as well as for general use in remote regions of developing countries.
  • the weight of a suitable permanent magnet is a clear disadvantage in comparison to a superconducting magnet, and rules out the use of a permanent magnet for a transportable MRI system.
  • Various aspects of transportable MRI systems are disclosed, for example, in Getz et al. U.S. Patent 5,727,353 issued Mar. 17, 1998, and Hobbs et al. U.S. Patent 7,733,089.
  • U.S. Patent 5,727,353 is directed to a portable self-contained diagnostic suite, which may include an MRI system, in a portable enclosure such as an International Standards Organization (ISO) approved shipping container.
  • ISO International Standards Organization
  • the present disclosure describes a self-contained transportable magnetic resonance imaging (MRI) system.
  • the MRI system includes a standard shipping container, and a vacuum vessel mounted within the shipping container.
  • the MRI system further includes a superconducting magnet suitable for magnetic resonance imaging (MRI), the superconducting magnet being contained within the vacuum vessel.
  • the superconducting magnet includes a stack of pre-formed coils of high temperature superconducting wire.
  • the vacuum vessel has an outer diameter wall of ferromagnetic material for magnetically shielding the superconducting magnet.
  • the MRI system further includes a cryogenic refrigerator mounted within the shipping container, the cryogenic refrigerator having at least one cold element in the vacuum vessel.
  • the MRI system further includes a cooling manifold in the vacuum vessel. The cooling manifold thermally couples the at least one cold element to the coils for cooling the superconducting magnet.
  • the MRI system also includes ferromagnetic material added to internal faces of the shipping container for reducing a fringe magnetic field footprint outside of the shipping container.
  • the present disclosure describes a magnetic resonance imaging (MRI) system including an imaging magnet having a bore and defining an imaging volume within the bore, and a patient table slidably mounted with respect to the imaging magnet for sliding a patient through the bore.
  • the imaging magnet defines an imaging volume in the bore, and a nuclear magnetic resonance (MR) probe is disposed within the imaging volume and offset from the patient when the patient is slid through the bore.
  • the NMR probe is comprised of a material capable of generating an NMR signal.
  • the present disclosure describes a magnetic resonance imaging (MRI) system including a vacuum vessel, a superconducting magnet suitable for magnetic resonance imaging (MRI) and contained within the vacuum vessel, a cryogenic refrigerator attached to the vacuum vessel, and a cooling manifold in the vacuum vessel.
  • the superconducting magnet includes a stack of pre-formed coils of high temperature superconducting wire.
  • the cryogenic refrigerator has at least one cold element in the vacuum vessel.
  • the cooling manifold thermally couples the at least one cold element to the coils for cooling the superconducting magnet.
  • the cooling manifold includes a thermal bus oriented along an axis of the superconducting magnet, and cooling plates sandwiched between neighboring coils of the superconducting magnet. Each of the cooling plates is thermally anchored to the thermal bus and contains at least one complete cut to disrupt formation of circumferential eddy currents in said each of the cooling plates, and the axial bus is connected to the at least one cold element of the cryogenic refrigerator.
  • the present disclosure describes a method of using a transportable magnetic resonance imaging (MRI) system.
  • the system has a superconducting magnet and a cryogenic refrigerator for cooling the magnet, and a controlled current source for supplying current to the magnet.
  • MRI magnetic resonance imaging
  • the method includes shipping the system to the target location over a distance requiring several days of travel when the system is de-energized, and at the target location: supplying power to the cryogenic refrigerator and cooling the magnet with the cryogenic refrigerator so that the magnet becomes superconductive; energizing the magnet using the controlled current source; checking the magnet's homogeneity using a standardized imaging phantom and diagnostic routines; and shimming the magnet's homogeneity when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity; and carrying out magnetic resonance imaging (MRI) scans using the magnet.
  • MRI magnetic resonance imaging
  • the present disclosure describes a method of using a transportable magnetic resonance imaging (MRI) system.
  • the system has a superconducting magnet, a cryogenic refrigerator for cooling the magnet, an on-board generator for powering the cryogenic refrigerator during transport of the system, and a controlled current source for supplying current to the magnet.
  • MRI magnetic resonance imaging
  • the method includes shipping the system to a target location with the onboard generator running to power the cryogenic refrigerator to keep the magnet cold but de-energized; and then, at the target location: energizing the magnet using the controlled current source; checking the magnet's homogeneity using a standardized imaging phantom and diagnostic routines and shimming the magnet's homogeneity when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity; and carrying out magnetic resonance imaging (MRI) scans using the magnet.
  • MRI magnetic resonance imaging
  • FIG. 1 is perspective view of a self-contained transportable magnetic resonance imaging (MRI) system including a standard shipping container, with a side wall of the shipping container removed to expose components within the shipping container;
  • MRI magnetic resonance imaging
  • FIG. 2 is a top view of the MRI system in FIG. 1 in schematic form to show a fringe magnetic field footprint outside of the shipping container;
  • FIG. 3 is a side view of the MRI system in FIG. 1 in schematic form to show various components in an examination room and in a control room inside the shipping container;
  • FIG. 4 is transverse cross-section of a vacuum vessel of the MRI system and components associated with a superconducting magnet in the vacuum vessel;
  • FIG. 5 is a longitudinal cross-section of the vacuum vessel and components associated with the superconducting magnet in the vacuum vessel;
  • FIG. 6 is an end view of a cryogenic refrigerator, cooling manifold, and superconducting magnet assembly;
  • FIG. 7 is a side view of the cryogenic refrigerator, cooling manifold, and superconducting magnet assembly of FIG. 5, showing the cooling manifold and superconducting magnet assembly in cross-section;
  • FIG. 8 is a bottom perspective view of a portion of the cooling manifold
  • FIG. 9 is an electrical schematic showing a circuit for energizing the superconducting magnet
  • FIG. 10 is a schematic diagram of electrical current and a magnetic field component generated by a gradient coil producing a magnetic field along the Z- direction and a gradient of the magnetic field along the Y-direction;
  • FIG. 11 is a is a schematic diagram of electrical current and a magnetic field component generated by a gradient coil producing a magnetic field along the Z- direction and a gradient of the magnetic field along the Z-direction;
  • FIG. 12 is a perspective view of an embodiment of the gradient coil represented schematically in FIG. 10;
  • FIG. 13 is a perspective view of an embodiment of the gradient coil represented schematically in FIG. 11 ;
  • FIG. 14 shows a Y gradient coil assembled into a cooling manifold
  • FIG. 15 shows an assembly of shimming coils
  • FIG. 16 shows an assembly of movable ferromagnetic shimming elements
  • FIG. 17 shows a flowchart of a method of transporting the transportable MRI system over a distance of several days travel to a target location and using the transportable MRI system at the target location;
  • FIG. 18 shows a flowchart of a method of transporting the transportable MRI system to a target location while a generator is running, and using the transportable MRI system at the target location
  • FIG. 19 is an end view of a cryogenic refrigerator, cooling manifold, and superconducting magnet assembly using layer wound winding blocks;
  • FIG. 20 is a cross-section side view of the cryogenic refrigerator, cooling manifold, and superconducting magnet assembly along line 20-20 of FIG. 19; and [00032] FIG. 21 is a side view of a cylindrical tube of the cooling manifold in FIG. 20.
  • a self-contained transportable magnetic resonance imaging (MRI) system 100 including a standard shipping container 101.
  • the standard shipping container is a standardized steel box that is transportable within the global containerized intermodal freight transport system, so that the standard shipping container is easily moved from ship, to rail, and to truck.
  • the standard shipping container is a twenty-foot ISO approved shipping container having external dimensions of 19' 10 1/2" (6.058 m) in length, 8' 6" (2.591 m) in height, and 8 ⁇ " (2.438 m) in width.
  • the interior of the shipping container 101 is subdivided by a partition wall 102 into an examination room 103 and a control room 104.
  • the separate rooms are accessible from outside of the shipping container through doors 105, 106 at opposite ends of the shipping container 105, 106.
  • the doors 105, 106 are in the two end walls of the shipping container 101.
  • FIG. 1 shows that the examination room 103 contains a torroidal vacuum vessel 107 containing a superconducting MRI imaging magnet.
  • the vacuum vessel 107 is supported on a pedestal 108 mounted to the bottom wall of the shipping container, which serves as the floor of the examination room 103 and the floor of the control room 104.
  • the pedestal 108 mounts the vacuum vessel and superconducting magnet to the shipping container 101.
  • the vacuum vessel would be hung from the top wall of the shipping container, which serves as the ceiling of the examination room 103 and the control room 104.
  • the vacuum vessel would be both supported by the pedestal 108 and hung from the top wall, so that it would be in the middle of a pillar between the top wall and the bottom wall. Such a pillar would provide better control over movement of the magnet relative to the walls of the shipping container 101 and ferromagnetic shielding 120, 157 mounted on the faces of the shipping container.
  • a patient table 109 is slidably mounted to the vacuum vessel for receiving a human patient 110 and sliding the whole body of the patient into an MRI imaging volume that is outside of the vacuum vessel 107 and within a bore of the superconducting magnet.
  • the imaging volume is defined by the region of magnetic field that has been designed to provide homogeneity suitable for imaging, which is typically a sphere of 40-45cm diameter with maximum deviation in field strength of +/-20 parts per million (ppm) peak-to-peak.
  • the internal diameter of the solenoid coils required to achieve this imaging volume is typically 84-93 cm, and the imaging volume within the bore is at least forty percent of the bore diameter.
  • the gradient coil assembly 133 and radio frequency body coil 142 are arranged inside the bore so that the tunnel that receives the patient is typically 60-70 cm in diameter.
  • the patient table 109 would be mounted on rails which would be attached to the pedestal and would extend from the pedestal to the partition wall 102 and would also be attached to the partition wall, so that the partition wall would help support the patient and avoid placing the patient's weight on the RF body coil 142.
  • FIG. 2 is a top view of the MRI system 100 in schematic form to show a fringe magnetic field footprint 121 extending outside of the shipping container 101.
  • the magnetic field at the periphery of the footprint 121 is five Gauss (0.5 Tesla), which is the internationally accepted safety limit for pedestrians outside the MRI examination room 103, and is predominantly defined to protect members of the public fitted with older heart pacemakers, which are sensitive to weak magnetic fields.
  • the MRI system 100 includes a set of barrier panels 125, 126, 127, etc. that are transportable within the shipping container 101 and are configured for being set up outside of the shipping container for constructing a barrier as shown in FIG. 2.
  • FIG. 3 is a side view of the MRI system 100 in schematic form to show various components in the examination room 103 and in the control room 104.
  • the torroidal vacuum vessel 107 contains the superconducting magnet 131 and includes a primary ferromagnetic shield 132, which may also function as the outer annular wall and annular end plates of the vacuum vessel.
  • the inner bore of the vacuum vessel may be formed by an impermeable tube made of relatively low conductivity material, such as glass fibre or stainless steel, which is sealed to the end plates by means of glue or o-rings to form a vacuum-tight seal.
  • a gradient coil and shim assembly 133 includes at least three nested coaxial tubular gradient coils disposed outside of the vacuum vessel and within the bore of the superconducting magnet 131.
  • the gradient coil and shim assembly 133 may also include shimming coils or ferromagnetic shimming elements.
  • the gradient coil and shim assembly 133 is in the ambient environment of room temperature and atmospheric pressure.
  • a Gifford-McMahon (GM) cryogenic refrigerator 134 is mounted to the pedestal 108 and has a cold element 135 within the vacuum vessel 107 in order to cool the superconducting magnet 131.
  • GM Gifford-McMahon
  • a pulse tube refrigerator (PTR) is used instead of a GM refrigerator.
  • the PTR would have a cold element in the vacuum vessel, but the PTR would also be mounted above the vacuum vessel.
  • the PTR would be better suited in the alternative embodiments in which the vacuum vessel 107 is suspended from the upper wall of the shipping container 101 or is in a pillar between the upper wall and the lower wall of the shipping container.
  • a GM refrigerator may be mounted above or below the vacuum vessel 107, so it is also suitable for the alternative embodiments in which the vacuum vessel 107 is suspended from the upper wall of the shipping container 101 or is in a pillar between the upper wall and the lower wall of the shipping container.
  • the vacuum vessel 107, superconducting magnet 131, and a thermal connection between the superconducting magnet and the cryogenic refrigerator 134 are configured to be relatively inexpensive, compact, light weight, easy to transport, and quick to set up.
  • FIG. 3 also shows various standard components used for MRI. These components include gradient power amplifiers 141 for driving pulsed current waveforms through the gradient coils in the gradient coil and shim assembly 133, a radiofrequency (RF) transmit body coil 142 in the bore of the superconducting magnet 131, an RF power amplifier 143 to excite resonance in nuclear spins of a body part in an MRI imaging volume 144, at least one body-part specific RF receive coil 145 to acquire the nuclear magnetic resonance ( MR) signal from the body part being imaged, a transmit-receive switch 146, a RF pre-amplifier 147, a spectrometer 148 to execute pulse sequences and digitize data, a control and system monitor computer 149, a magnet power supply 150 including a controlled current source, a compressor 151 for supplying compressed helium gas to the cryogenic refrigerator 134, and active shim power supplies 152 for powering shimming coils in the gradient coil and shimming assembly 133.
  • RF radio
  • FIG. 3 also shows components that are often found in shipping containers for refrigerated shipping of perishable items. These components include an uninterruptible power supply 152 for backup power and power during transport, such as a diesel generator set, and a climate control unit 154.
  • the uninterruptable power supply is used in the MRI system 100 to provide power to the cryogenic refrigerator and optionally to the complete MRI system when an external power source is not available.
  • the climate control unit 154 maintains the examination room 103 and the control room 104 at a comfortable room temperature.
  • the walls of the examination room 103 are configured as a Faraday cage 155 so that the superconducting magnet 131 and low-signal level RF components such as the RF receiver 145 are protected from pick-up of extraneous RF interference, such as interference produced by electrical components in the control room 104 or RF sources outside of the shipping container 101.
  • Parts of this Faraday cage may be formed by the steel walls of the shipping container 101 and/or secondary ferromagnetic shielding 156, 157.
  • a filter box 158 allows control and monitoring cables, RF cables, magnet and active shim electrical power and helium gas lines to and from the compressor 151 to pass through the Faraday cage without compromising its electromagnetic integrity.
  • the area around the filter box 154 and around the door 105 to the examination room 103 may include localized radio frequency shielding, such as copper screen and copper flashing, added to the shipping container to close radio frequency (RF) signal leaks.
  • RF radio frequency
  • the majority of electrically noisy or high power electronic sub- systems are housed in the control room 104 at one end of the shipping container, outside the Faraday cage.
  • the walls of the shipping container are filled with thermal insulation 159 so that the climate control system 154 can maintain a uniform temperature during operation and transportation of the system. A uniform temperature during operation is desired because the ferromagnetic properties of steel vary markedly with temperature, which will affect the homogeneity of the magnetic field in the imaging volume.
  • various components of the MRI system 100 are configured to be relatively inexpensive, compact, light weight, easy to transport, and quick to set up, so that the system is especially suited for providing routine medical care in rural regions of developing countries.
  • the MRI system 100 provides a relocatable or mobile whole body clinical MR imaging system at significantly higher field strength and much lower weight than a permanent magnet MRI, and with substantially lower installed cost than a high field superconducting magnet MRI.
  • the superconducting magnet 131 uses high temperature superconductor and is coupled to the cryogenic refrigerator 134 in such a way that the superconducting magnet 131 and the vacuum vessel 107 are significantly smaller than normal for conventional whole body MRI, and the complete MRI system fits inside a standard 20 foot shipping container 101.
  • no liquid cryogens are required, which is a major advantage over fixed and mobile superconducting MRI, particularly in a transportable system.
  • the cryogen-free high-temperature superconducting magnet 131 overcomes various problems with the low-temperature superconductor (LTS) magnets used in the vast majority of commercial superconducting MRI systems.
  • LTS low-temperature superconductor
  • These commercial superconducting MRI systems typically use solenoid magnets wound from NbTi LTS conductor, bath-cooled in liquid helium (LHe) boiling at atmospheric pressure at 4.2 Kelvin.
  • LHe liquid helium
  • higher field LTS solenoids > 1 T
  • they have a set of secondary counter-running shielding coils placed at a substantially larger radius than the primary field coils to constrain the magnet's return flux and minimize the dimensions of the 5 Gauss fringe field footprint.
  • a consequence of adding the active shield coils is that whole-body solenoids typically have large outside diameters, exceeding 2.2 m. Both the primary and shielding coils must be kept below ⁇ 6 Kelvin over their entire circumference for stable, quench-free superconducting operation. This is achieved by partially submersing them in LHe.
  • the large size of the active shielding coils and the low critical temperature of LTS conductor mean that immersion in, or intimate contact with liquid cryogen over at least part of the coils circumference, is necessary to maintain all of the coils below the critical temperature of the superconductor.
  • An active shielded LTS solenoid magnet is therefore conventionally contained in a large annular helium-tight vessel, itself placed inside an annular vacuum vessel. Due to the presence of liquid cryogen, the cryostat must be certified to meet the regulatory pressure vessel safety standards which prevail in each territory in which the magnet will be sold.
  • the radial space inside the helium vessel between the primary and shield coils is used as a reservoir for several thousand liters of boiling LHe.
  • a suitable cryogenic refrigerator such as a Gifford-McMahon (GM) or pulse tube refrigerator (PTR).
  • GM Gifford-McMahon
  • PTR pulse tube refrigerator
  • the helium vessel is wrapped in multi-layer insulation and surrounded by at least one highly conductive thermal shield, which is cooled by the first stage of the refrigerator.
  • a large stored volume of LHe is necessary for practical operation of a conventional LTS magnet. Its latent heat of evaporation provides a distributed source of cooling power which allows the magnet to remain at 4.2 Kelvin without operation of the cryo-refrigerator. The magnet will not warm up until all the stored LHe has boiled, so a large reservoir results in a long "hold time”. This allows manufacturers to cool the magnet to 4.2 Kelvin in the factory, energize it temporarily for testing, and then ship it to the customer site cold and full of LHe. This approach is often chosen because cooling down the magnet cold mass from room temperature to 4.2 Kelvin consumes several thousand liters of LHe, at significant cost, and it is financially undesirable to do this more times than necessary.
  • the large size of an active shielded magnet results in a large reservoir and a hold-time of about 30 days.
  • transport of the magnet from the factory to the customer site, installation, helium re- fill and power-up of the refrigerator must all be accomplished within the cryogen hold-time. If the magnet warms up due to transport delays, additional LHe must be used to re-cool it, at significant expense and delay, and with some risk that movement of coils during the consequent thermal cycle will affect quench behavior or magnetic field homogeneity on subsequent ramps to field.
  • the magnet is equipped with an emergency run-down switch, which allows the user to force a quench.
  • the quench duct must be fitted with a one-way valve to prevent ingress of air into the cryostat, which could freeze and block the quench duct.
  • An LTS magnet's thermal shield is typically manufactured from aluminum alloy to facilitate heat flow to the refrigerator and minimize temperature variation across its surface area.
  • the thermal shield bore tube presents a high conductivity surface to the stray magnetic fields emanating from the gradient coils used in an MRI system, which leads to generation of large transient eddy currents when current pulses are applied to the gradient coils during imaging; these eddy currents in turn generate unwanted transient magnetic fields in the imaging volume, which create image artifacts.
  • A/S gradients also known as self- shielded gradients, were introduced in the 1980s to minimize induction of eddy currents in the thermal shield.
  • A/S gradients are significantly more complex and costly than conventional unshielded gradient coils. They are also much less efficient in terms of gradient strength per amp (mT/m/A) and hence significantly increase the cost of the gradient power amplifiers used to drive pulsed current through the coils. It is therefore desirable to avoid the use of actively shielded gradient coils in a low cost MRI system.
  • a further complication with installation of LTS magnets is introduced by the sensitivity of an actively shielded magnet to sources of magnetic fields and ferrous objects in its immediate environment.
  • the homogeneity of the magnetic field over the imaging volume which is critical for imaging, is influenced by the presence of ferromagnetic components in the building structure, such as structural steel beams often used to reinforce floors. Combined with the heavy weight of existing whole body MRI magnets this makes installation above ground floor level considerably more costly.
  • An individual site survey is carried out for each magnet installation to assess the suitability of the site. When the magnet has been installed, topped up with LHe and ramped back to field, it is then necessary to check and adjust the homogeneity of the magnetic field, since this will have been affected by its new environment.
  • the process of "shimming" the magnet involves mapping the magnetic field variation over the imaging volume, then either adjusting the currents in superconducting shim coils ("active shimming"), or calculating a distribution of small iron pieces (shims) which must be loaded into the magnet bore to compensate for field inhomogeneity (“passive shimming”).
  • Active shimming adjusting the currents in superconducting shim coils
  • shims small iron pieces which must be loaded into the magnet bore to compensate for field inhomogeneity
  • Passive shimming Loading these ferromagnetic shims cannot be carried out at field, for safety reasons because the non-uniform magnetic field has the potential to accelerate the ferromagnetic parts to high velocity, so the magnet must be ramp cycled for each shim iteration. Typically several shim iterations are necessary, with a small risk that the magnet will quench each time it is ramped to field.
  • the magnet must be topped up again with L
  • An actively shielded superconducting magnet is also electromagnetically "transparent" to time varying external magnetic fields.
  • iron shielded magnets are inherently more immune to the influence of external fields, because the superconducting coils tend to exclude external flux; the counter-running coils in an active shielded magnet unfortunately negate this effect.
  • EIS coil external interference shield coil
  • An additional problem created by active shielding of high field magnets is the generation of intense static field gradients around the open bore of the magnet, where the magnetic field direction turns through 180 degrees as magnetic flux passes from flowing inside the primary magnet coils to flow between the primary and shield coils.
  • These static field gradients create a significant safety issue because the force on a ferromagnetic object depends on field gradient.
  • the attractive force experienced when a ferromagnetic object such as a tool or oxygen cylinder is introduced into the fringe field of an actively shielded magnet varies rapidly with distance, particularly around the open bore, giving little time for an inattentive operator to react before the object is pulled from their grasp, with potentially serious safety consequences.
  • the same field gradients can also induce significant currents in conductive objects which move near the open bore of the magnet, such as the tissues of the patient or MRI technicians. Therefore is desirable to minimize the generation of intense static field gradients around the magnet.
  • the magnet can be conduction-cooled and have a minimized cold mass which speeds up cool-down time and dispenses with the very significant technical and logistical complexities introduced by use of liquid cryogens.
  • the magnet can be very compact yet have whole-body access for human patients and a -45 cm diameter imaging volume similar to a high field LTS solenoid, yet having significantly smaller outside diameter and lower weight than existing LTS magnets. This is achieved by eliminating active shield coils, dropping field strength to 0.5 - 1 T, and using minimal iron shielding to control the stray magnetic field.
  • the combination of wide bore, short length and significantly smaller outside diameter provides a very appealing aesthetic appearance and a reduction in patient anxiety.
  • the magnet can have a compact 5 gauss fringe field, achieved by use of primary and secondary iron shields, one mounted on the magnet and the second mounted on some faces shipping container, such as the side walls of the examination room.
  • This approach also avoids generation of the strong magnetic field gradients associated with active shielded magnets, improving safety, and also improving the immunity of the system to perturbation by external sources of magnetic field and/or ferrous items in the local environment. This in turn allows the magnet to be passive shimmed in the factory.
  • the magnet can use high temperature superconducting wire (HTS) which allows the magnet to operate at higher temperature (typically 15-30 K) using conduction cooling from a cryogenic refrigerator, which in turn allows the elimination of liquid cryogen and the complex design features and logistics involved with use of liquid cryogen. It also enables the cryostat construction to be substantially simplified and more robust, reducing cost, and significantly increasing tolerance to shock and vibration.
  • HTS magnet is therefore much better suited to use in a mobile or transportable application than an LTS or permanent magnet.
  • the system can provide flexible operation modes for transport, including the ability to transport the magnet cold and de-energized, for rapid deployment at the new location (in a matter of hours), or allowing the magnet to warm up during longer journeys, with a 2-3 day cool-down time once reconnected to electrical power.
  • Cold- transport is achieved by using a shipping container incorporating a diesel generator to run the magnet refrigerator.
  • the system can provide full body scanning while having a low total system weight of less than twenty metric tons (20,000 kg), including all electronics, ancillary systems, and a standard shipping container equipped with a diesel generator, which allows low-cost and straight-forward transport and rapid deployment in a variety of locations, ranging from military battle-field hospitals or disaster relief, to semi-permanent deployment in remote or rural areas.
  • This concept is particularly well suited to developing nations which do not have well developed healthcare or transportation infrastructure.
  • a significant reduction in system power consumption and cost is achieved by use of unshielded gradient coils combined with unconventional construction of the HTS magnet, which minimizes eddy current interactions between the magnet and gradient coils.
  • High temperature superconductor is a superconductor having a transition temperature above thirty degrees Kelvin (-243.2 °C). It is desirable to operate a HTS MRI magnet at a temperature substantially lower than the transition temperature in order to achieve a higher critical current and hence a high magnetic field with minimized length of superconducting wire. For this reason, it is desirable to cool the HTS magnet to about 20 degrees Kelvin before operating the HTS magnet, and to maintain the HTS magnet between ten degrees Kelvin and thirty degrees Kelvin during operation, and to keep the magnetic field to which the coils are exposed well below the critical field at the operating temperature.
  • Suitable HTS conductors for use in the HTS magnet are ceramic superconductors such as BSCCO (e.g., Bi-2223) and REBCO family of superconductors, and conductors using magnesium diboride (MgB 2 ).
  • REBCO superconductor contains barium, copper, oxygen, and one or more of the rare earth elements (represented by the acronym RE).
  • a typical REBCO superconductor has a Perovskite lattice structure (the so-called 123 structure) and the chemical formula REBa 2 Cu 3 0 7 , where RE is Yttrium, Europium, or Erbium.
  • MgB 2 has some distinct advantages, particularly for operation in low to mid-field magnets; namely, negligible dependence of critical current on magnetic field angle and substantially lower cost. MgB 2 is therefore the preferred conductor for the superconducting magnet, but REBCO as well as BSCCO and YBCO are technically suitable choices, which are simply less desirable for cost reasons.
  • the preferred wire format for the MgB 2 magnet is flat tape manufactured by a pre-reacted powder-in-tube method.
  • the magnet is formed from a set of winding blocks, and each winding block includes one coil or a number of contiguous coils.
  • each winding block includes at least one set of double pancake coils.
  • One way of forming the individual coils is by winding flat HTS tape onto a former to create a single turn-per-layer pancake coil, impregnating the coil with suitable resin or wax to provide structural integrity in a cryogenic environment, and removing the coil from the former. Pairs of pancake coils with opposing winding directions are then connected to form double pancakes.
  • FIGS. 4 and 5 show details of the HTS magnet 131 in the vacuum vessel 107
  • the HTS magnet 131 is part of a cold-mass assembly 170, and the HTS magnet 131 in this example includes four winding blocks 301, 302, 303, 304.
  • a typical magnet would contain at least five or six winding blocks to achieve the desired magnetic field homogeneity, but in the figures this has been simplified to four winding blocks, the outer pair 301, 304, containing two double pancakes coils each and the inner pair 302, 303, just one double-pancake coil each.
  • each winding block will contain at least one but more commonly multiple double pancake coils, stacked coaxially.
  • the cold-mass assembly 170 is then wrapped in multi-layer insulation (MLI) 177 reducing radiation heat load by reflecting thermal radiation, and loaded into the vacuum vessel 107.
  • the outer tube of the vacuum vessel 107 is the primary ferromagnetic shield 132, made of iron or mild steel, which also forms a primary flux return path.
  • the vacuum vessel 107 has an inner bore 178 made of low-conductivity non-magnetic material such as stainless steel, or optionally non- conductive but vacuum-tight composite material such as fiberglass.
  • the annular vacuum vessel 107 is closed with annular end-plates 179, 180 which may be ferromagnetic.
  • a cooling manifold 182 includes flexible copper braid 305 connected to a second stage of the cryo-refrigerator 134 to decouple the magnet 131 from vibration from the cryo-refrigerator.
  • a two-stage Gifford McMahon (GM) refrigerator or pulse tube refrigerator (PTR) is suitable.
  • a thermal shield 181 is optionally provided on the outer radial surface of the cold mass 170.
  • the thermal shield 181 is attached to a first stage cold element 136 of the cryogenic refrigerator 134 to balance the radiation heat load and maintain the thermal shield at a constant temperature.
  • the cryogenic refrigerator 134 keeps the first stage cold element 136 at a higher temperature and with a higher cooling power than the second stage cold element 135, and the thermal radiation shield is thermally anchored to the first stage cold element. This arrangement considerably simplifies magnet construction and reduces system cost compared to a conventional helium bath-cooled LTS magnet.
  • the cold mass assembly 170 is attached rigidly to the end plates 179, 180 using posts 191, 192, 193, 194 of non-conductive material such as G10 epoxy- fiberglass arranged in compression. Contraction of the cold mass during cooldown is compensated by use of pre-compressed Bellville washers 185, 196, 197, 198.
  • the Bellville washers 185, 196, 197, 198 are coaxial with the posts 191, 192, 193, 194 and are disposed between the cold mass assembly 170 and ends of the posts 191, 192, 193, 194 adjacent to the cold mass assembly.
  • the end plates 179, 180 and posts 191, 192 rigidly locate the superconducting magnet 131 within the vacuum vessel 107, and the Bellville washers 191, 192, 193, 194 take up the slack created by contraction of the cold mass when cooled, ensuring that the superconducting magnet 131 remains accurately located within the vacuum vessel 107 during cool down, energization of the superconducting magnet, and under exposure to external vibration.
  • This design results in an increase in conducted heat load compared to the conventional suspension system used on LTS magnets, but this is tolerable thanks to the higher operation temperature of the magnet and hence greater cooling power available from the refrigerator 134.
  • a particular benefit of this construction method is the substantially increased ability of the magnet 131 to withstand shock loads without damage, and a marked reduction in susceptibility of the temporal stability of the magnetic field to vibration from external sources and from the refrigerator 134.
  • FIGS. 6 and 7 show details of the cryogenic refrigerator 134 and the cold mass assembly 170 including the cooling manifold 182 and the winding blocks 301, 302, 303, 304.
  • a distributed heat conduction pathway called the cooling manifold 182 is used to transport heat from the winding blocks to the cold element 135 of the cryogenic refrigerator 134.
  • the cooling manifold 182 includes annular plates 201, 202, 203, 204, 205, 206, 207, 208, 209, 210 and an axial thermal bus 211 thermally coupled to the cold element 135 of the second stage of the cryogenic refrigerator 134.
  • the annular plates 202, 203, 204, 205, 206, 207, 208, and 209 are disposed on the faces of respective neighboring pairs of the pancake coils 171, 172, (that comprise winding block 301), 173 (in 302), 174, (in 303) and 175 and 176 (in 304).
  • the annular plates 201, 202, 203, 204, 205, 206, 207, 208, 209, 210 and thermal bus 211 are constructed from a material with good thermal conductivity, such as copper, aluminum or sapphire, or a combination of these materials.
  • the annular plates and the thermal bus 211 are constructed from oxygen-free copper plate, and the annular plates are thermally connected to the thermal bus.
  • the thermal connection may use indium shims or grease to reduce thermal resistance across joints between the annular plates and the thermal bus.
  • a good thermal connection can also be obtained by gold plating the annular plates and the thermal bus at the regions of the joints, and configuring the joints for applying a high pressure for pressing the gold-plated parts together.
  • the thermal bus is connected to the second stage cold head of the cryogenic refrigerator by oxygen-free copper braid 305, which permits a small amount of relative motion between the cold head and magnet assembly and therefore isolates the double pancake coils from vibration of the cryogenic refrigerator 134.
  • the magnetic field homogeneity in the imaging volume (144 in FIGS. 3 and 4) is controlled by leaving axial gaps between the winding blocks and/or selecting the inner and/or outer radii and number of turns on each double pancake to achieve a uniformity of typically 20 ppm peak-to-peak over the imaging volume 144. Large gaps are left between winding blocks to achieve the desired magnetic field homogeneity and these gaps are filled with spacers 212, 213, 214 to support the electromagnetic forces acting axially on the coils.
  • the annular disks, pancake coils, and spacers can be held together by axial tie rods or banding.
  • the spacers 212, 213, 214 can be made from a non-conductive structural material suitable for use in a cryogenic environment, such as engineering composites like G10 epoxy-fiberglass. Stainless steel is also suitable, having the advantage of better defined thermal contraction but the disadvantage of greater electrical conductivity.
  • the annular spacers may be continuous disks or segments of disks.
  • the latter may be preferred as it is advantageous to minimize the mass of material used in the magnet to speed up cool down time, and to break circumferential eddy current paths if a conductive material such as stainless steel is used. For the same reason it may be preferred to use material with a low density and/or low heat capacity, contingent on having sufficient mechanical strength, to minimize the cold mass.
  • a high volume of heat capacity material to provide a thermal reservoir to extend the ride-through time in the event of cooling power failure, because the HTS magnet 131 is provided with an uninterruptable power supply in the form of a diesel generator (153 in FIG. 3).
  • the cooling manifold's annular plates 204, 206 have complete slots 215, 216 to prevent circumferential eddy currents from being induced in the annular plates when the gradient coils are energized. This will be further described below with reference to FIG. 14. For this reason, all of the annular plates 201, 202, 203, 204, 205, 206, 207, 208, 209, 210 shown in FIG. 7 have similar complete slots diametrically opposite from the thermal bus 211.
  • the coils are connected in series by low resistance solder joints. While it is potentially possible to connect MgB 2 wire with superconducting joints and achieve a magnet with ability to operate in persistent current mode, this technology is currently poorly developed, particularly for use in a strong background field. Naturally a magnet with resistive joints cannot be operated in persistent current mode (as done with conventional LTS magnets), but this is addressed by operating the magnet in driven mode using an external current source. Various features and advantages of this approach are described below. Care is taken to minimize the resistance of the joints to reduce additional heat load to the cold mass.
  • the superconducting magnet 131 is provided with a primary ferromagnetic shield (132 in FIGS. 3, 4 and 5) and also by a secondary ferromagnetic shield including the shielding plates 156, 157 shown in FIG. 3 to provide a return path for magnetic flux and reduce the fringe field footprint of the magnet.
  • the secondary ferromagnetic shield also includes similar shielding plates on the walls of the examination room (103 in FIG. 1) coplanar with the side walls of the shipping container (101 in FIG. 1). For example, these shielding plates are constructed very simply from single plates of mild steel typically 5-20 mm thick. No shielding is necessary on the two faces of the cuboid shield which are aligned with the magnet axis.
  • the axial length of the shielding plates is determined by finite element modeling to optimize the constraint of the magnet's stray field while minimizing the mass of steel used.
  • the shielding plates may coincide with the end wall of the shipping container that is a wall of the examination room 103, and the partition wall 102 between the examination room 103 and the control room 104, or may be shorter.
  • the shielding plates (156, 157 in FIG. 3) on the floor and ceiling are also optional if it is deemed acceptable to allow the magnet's fringe field to extend outside the container on these faces, as would be the case in installations where there is no pedestrian access to regions above or below the container.
  • the combined effect of the primary and secondary iron shields is sufficient to pull the 5 gauss contour line (121 in FIG. 2) of the magnet in to within 50 cm of the external walls of the shipping container using a total mass of steel less than 15,000 kg.
  • the elimination of large diameter active shield coils substantially reduces the size, cost and complexity of the vacuum vessel.
  • the largest temperature difference across the diameter of the small diameter main magnet coils occurs at the point furthest from the second-stage of the refrigerator. With appropriate choice of materials and dimensions in the cooling manifold this temperature difference can be small enough such that all parts of the coils can remain substantially below the critical temperature of the high temperature superconductor, for example, at no more than 75% of the critical current of the superconductor as defined by the local temperature and magnetic field to which the superconducting wire is exposed at any position within the magnet structure. If conduction cooling of large diameter shielding coils were attempted from a single cold head then excessive temperature rise would result at the further point from the cold head, which would limit the coil's current capability.
  • the use of two iron shields does require the magnet and primary shield to be accurately located with respect to the secondary shield to avoid generation of large field inhomogeneities in the imaging volume.
  • Small misalignments introduced during manufacture can be compensated during passive shimming of the magnet. Any small variations in the relative positions of the primary and secondary shields which occur during operation, such as those potentially caused by mechanical distortions during system relocation or thermal effects due to ambient weather conditions (e.g. : sunlight shining on one side of the shipping container), can be minimized by incorporating insulation between the secondary shield plates and the container wall. Any residual variations may be compensated by adjusting currents in the active shim coils fitted to the magnet.
  • Active shim coils are additional coils which generate magnetic fields of known geometry, typically in the form of low order and degree spherical harmonics.
  • the gradient coils themselves can be used as active shim coils for X, Y and Z spherical harmonics.
  • the necessary field corrections can be generated by applying appropriate DC currents to these coils, up to a few ppm of the centre field with practical coil currents in copper. If it is necessary to compensate for larger variations, the magnet can be fitted with HTS shim coils located in the vacuum space and cooled using the cryogenic refrigerator in the same way as the magnet coils.
  • Another feature of the transportable MRI system 100 is the use of an external power supply (or current source).
  • an external power supply or current source.
  • Conventional LTS magnets are operated in persistent mode. The magnet is ramped to field using a power supply located outside the cryostat and connected to the superconducting circuit via current leads which penetrate the cryostat. A superconducting switch is then closed and the power supply de-energized. The power supply is often removed and sometimes the current leads may be physically disconnected if it is desired to absolutely minimize conducted heat leak to the 4.2 K cold mass.
  • a quench protection circuit commonly comprising resistors and diodes, is connected across the LTS magnet to dissipate its stored energy in the event of a quench.
  • the HTS magnet 131 is permanently connected to a controlled current source 221 of the magnet power supply unit (150 in FIG. 3) via permanently connected current leads 223, 224.
  • the controlled current source 221 is stabilized by feedback of an NMR frequency of the superconducting magnet.
  • feedback of an NMR frequency stabilizes a controlled current source (275, 276 in FIG. 15) providing current to shim coils (273, 274 in FIG. 15) further described below with respect to FIG. 15. In either case, this enables the magnetic field strength in the bore of the magnet to be kept to a constant value.
  • a dump resistor 222 is connected across the magnet 131 to safely dissipate its stored energy if the external power supply is switched off or becomes disconnected.
  • the dump resistor 222 also presents a lower impedance to high frequency noise from the power supply than the highly reactive HTS magnet 131, which helps ensure good short term stability of the magnet current.
  • the controlled current source 221 is used both to energize the HTS magnet 131 and also maintain current during operation, offsetting losses in the slightly resistive joints between coils.
  • the controlled current source 221 has the ability to stabilize the magnet's field strength rapidly after ramping. Ramping induces screening currents in the individual turns of superconducting coils by self-field induction. These parasitic screening currents generate a magnetic field which opposes and hence reduces the magnet's as-designed field strength and also may slightly change the field homogeneity across the imaging volume. However, the screening current field decays with time constant of hours or days as flux pinning centers creep and screening currents within each superconductor strand redistributes.
  • the NMR feedback signal may be supplied from the object being imaged, or from an additional an NMR probe (161 in FIG. 3) placed within the imaging volume 144. A suitable place to locate this probe would be under or inside the patient table 109.
  • the NMR probe 161 comprises a small container of liquid, gel or any material which can generate an NMR signal, with a tuned RF pickup coil 162 used to detect the NMR frequency.
  • the NMR nucleus chosen to provide the feedback signal could be hydrogen, or another active NMR nucleus, such as deuterium or sodium, or any nucleus with non-integer spin.
  • the sample is isotopically enriched D 2 0 (heavy water).
  • Deuterium is chosen because its NMR frequency is substantially lower than that of hydrogen and thus avoids any complications with interference between the NMR signal used to lock the magnet current and the hydrogen signal from the patient.
  • the NMR signal from the probe 161 may be acquired using pulsed NMR or continuous wave NMR, as established in the literature, and converted to an analog voltage signal or a digital representation of the NMR probe sample's frequency, which is proportional to the magnetic field strength integrated across the probe's sample volume. This approach to locking the magnet field is routinely used in NMR spectroscopy magnets, where the chemical sample being analyzed is dissolved in a deuterated solvent, however the difference here is that the NMR probe 161 is located physically separately from the sample.
  • the NMR probe 161 naturally cannot occupy the same space as the patient (110 in FIG. 1) so the NMR probe 161 will be offset from the magnet's isocentre, for example, under the patient table where it is still inside the imaging region having sufficient homogeneity to provide an NMR signal with high frequency resolution.
  • a bath-cooled LTS magnet operated in persistent mode is able to "ride-though" mains power failure for days, or even weeks.
  • the magnet remains cold thanks to the cooling power of boiling liquid cryogen, and remains at field thanks to persistent current operation.
  • a conduction-cooled magnet operated in driven mode will immediately ramp down through the dump resistor when mains power is interrupted and the current source ceases to supply current.
  • the magnet will also begin to gradually warm up due to loss of cooling power from the refrigerator. This scenario is undesirable because the magnet must be cooled down again and ramped back to field before imaging can recommence.
  • an aspect of the present disclosure is the use of a diesel generator equipped shipping container, a standard option for the transport of refrigerated goods, to provide an uninterruptible power supply for the MRI system.
  • the generator enables the MRI superconducting magnet to remain cold during short-term power failure and also to be transported with the cryo-refrigerator running if necessary. This would be the preferred way to transport the system over shorter distances as it allows rapid deployment at the destination without waiting for the magnet to cool down.
  • An added benefit of this feature is the ability to operate the MRI system without connection to an external power source, which will be of great benefit in military or humanitarian relief scenarios.
  • Another feature of the transportable MRI system 100 is minimization of eddy currents in the magnet structure during imaging. Eddy currents induced in conductive magnet components during pulsing of the gradient coil can create transient distortions of the magnetic field homogeneity over the imaging volume, which in turn cause image artifacts.
  • the elimination of the inner thermal shield removes a major source of eddy currents and potentially allows the use of unshielded gradient coils, which are significantly simpler and mre power efficient than shielded gradient coils used in conventional LTS magnets. However, this exposes the magnet directly to the stray pulsed magnetic flux from the gradient coils, which can induce eddy currents in any conductive components.
  • the construction of the magnet previously described ensures that the majority of magnet components are made from materials with zero or poor electrical conductivity, such as stainless steel. However, a significant exception is the distributed copper cooling manifold.
  • FIGS. 10 and 11 The basic conductor patterns for generating the linear gradient magnetic fields required in conventional MR imaging are depicted in FIGS. 10 and 11. These figures use the conventional Cartesian axis nomenclature, with the Z axis oriented along the direction of magnetic field at the magnet centre, and hence along the axial bore of the superconducting solenoid magnet (131 in FIG. 3).
  • FIG. 10 shows that transverse gradients X and Y (i.e. : ⁇ ⁇ / ⁇ and d z /dy, depicted as vectors 245) may be generated by pairs of Golay saddle coils 246, 247 with appropriate orientation.
  • FIG. 11 shows that a simple Z gradient (i.e.
  • ⁇ /dz with the change in magnetic field depicted along the z axis by vectors 250) may be generated by a Maxwell pair of opposed conductor loops 251, 252.
  • the conductor paths for a typical Y coil 251 are shown in FIG. 12.
  • the conductor paths for a typical X coil are the same, except that they are rotated 90 degrees about the Z axis.
  • the conductor paths for a typical Z coil 252 are shown in FIG. 13.
  • a feature of the HTS magnet 131 is that all components with high electrical conductivity, such as the copper cooling manifold, are segmented such that there are no pathways which support eddy currents with the same symmetry as the X, Y or Z gradient coils.
  • This principle can be extended to cover gradient coils which generate other field shapes if necessary, such as required for PATLOC imaging.
  • composite materials made from a mesh of conductive metal fibers within a non-conducting matrix, such as resin it is possible to use machined copper components with strategically placed breaks which disrupt the eddy current symmetry.
  • FIG. 14 depict only the highly conductive cooling manifold components of a magnet cold mass 262 including cooling plates 263, 264, 265, and 266, and an axial thermal bus 267, which are used to thermally connect all elements of the magnet cold mass.
  • Pancake coils of a superconducting magnet and structural components spacing and supporting the cooling plates have been omitted.
  • FIG. 14 the conductor patterns for a Y gradient coil 268 are shown located inside the cooling manifold.
  • an X gradient coil and also a Z gradient coil would be nested with the Y gradient coil in a coaxial fashion, and these coils could be nested in each other in any order.
  • the cooling plates 263, 264, 265, and 266 all have at least one complete cut 267 which prevents eddy currents from flowing around their periphery from being induced. It also breaks coupling to the Z gradient coil, preventing formation of a mirror eddy current when pulsing the Z coil.
  • the preferred structure of the cooling manifold 262 in FIG. 14 can be described as a "spine” (the axial thermal bus 267) and "ribs" (the cooling plates 263, 264, 265, 266 between coils, which have been omitted). While the figures show an embodiment with one spine and one set of ribs encompassing the full circumference of the coils (broken only by the small cut, at the position of the "sternum", it is possible to have more than one spine, each with its own set of ribs which subtend a smaller angle. For example, it may be advantageous to have two sets, with their spines diametrically opposed on opposite sides of the magnet coils and ribs subtending slightly less than 180 degrees. Each spine could be connected to its own refrigerator cold element. This arrangement would reduce the temperature variation around the coils and allow the magnet to be run at a higher percentage of critical current, thereby generating a stronger magnetic field.
  • LTS low temperature superconducting
  • the superconducting wire filaments typically of NbTi
  • NbTi are embedded in a copper matrix to provide stability from flux jumping and provide an alternative path for the magnet current in the event that a local quench occurs. If a quench does occur in an LTS magnet is propagates quickly over the whole winding (typically in a few seconds), and the magnet current is temporarily carried by the copper matrix and flows through the quench protection circuit resistors, which both dissipate the magnet's stored energy as heat, boiling and ejecting the stored LHe. A correctly designed magnet will safely dissipate its stored energy in this way and survive a quench without damage. However, it needs to be re-cooled and re-energized before it can be used again.
  • HTS conductors behave quite differently. Firstly they are stable against flux jumping without requiring a large quantity of copper stabilizer in the wire structure. However, if a localized quench occurs, it propagates very slowly (minutes). This creates a problem with protection, because the magnet's inductance will generate a very large voltage across the portion of wire which is no longer superconducting, until an arc occurs, which will destroy the coil. Simply providing an alternative path for the current though a copper co-winding is not a satisfactory solution, because the majority of the magnet's HTS wire persists in superconducting mode, so current only flows through the copper around the local quench.
  • the resistance of the copper causes local heating and it quickly melts, creating a high resistance, leading to an arc and destruction of the coil.
  • Others have proposed systems of heaters which are triggered when a quench is detected to rapidly warm up the magnet and force the current to commute more globally to the copper co-wind.
  • This approach is complex and adds significant cost.
  • the matrix used for HTS wire is generally of low conductivity (e.g. : stainless steel for REBCO and nickel for MgB 2 , although BSCCO uses a silver matrix). It is desirable to minimize the bulk conductivity of the pancake coils in the magnet to prevent formation of eddy currents when pulsing the gradient coils.
  • the coils are made from REBCO or MgB 2 tape without copper stabilizer. This keeps their bulk conductivity low, as desired. It also increases engineering current density and simplifies coil winding. However, in the event of a local quench, the coil will burn out. This risk is mitigated by running the magnet at no more than 75% of critical current and by continuously monitoring the coil temperatures and voltages. If either starts to rise, the magnet current source goes open circuit and the magnet discharges through the dump resistor (222 in FIG. 9). A spontaneous quench in an LTS magnet is a rare, but not impossible event, so LTS magnets must be designed to survive this. However, a spontaneous quench in a correctly designed HTS magnet has a much lower probability, so this approach to quench protection is acceptable in an HTS magnet.
  • FIG. 15 shows an example of an active shimming coil assembly 271 used for fine correction of the homogeneity in the imaging volume.
  • the active shim has a set of coils 273, 274 for producing a relatively pure Z 2 spherical harmonic, and a set of saddle coils 277, 278 for producing spherical harmonics of the sine and cosine type.
  • the shimming coils 273, 274, 277, 278 are driven by respective controlled current sources 275, 276, 277, 278.
  • the shimming coil assembly 271 is tubular so that it can be nested in a coaxial fashion with the X, Y, and Z gradient coils in the gradient coil and shimming assembly (133 in FIG. 3).
  • the shimming coil assembly 271 is mounted in the vacuum vessel (107 in FIG. 3), so that the cold mass assembly (170 in FIG. 3) is nested in the assembly 271 of shimming coils.
  • the shimming coil assembly includes a tubular electrically non- conductive support 272, and the shimming coils 273, 274 mounted on or integral with the electrically non-conductive support.
  • the electrically non-conductive support 273 is a fiberglass tube made by winding resin impregnated fiberglass cloth around a cylindrical mandrel, and then curing the resin.
  • the shimming coils 272, 273 are wound circumferentially around the electrically non-conductive support 273.
  • these coils function in a way similar to the Z gradient coil (252 in FIG. 13).
  • the shimming coils 277, 278 are not wound circumferentially around the electrical non-conductive support 272.
  • these coils function in a way similar to an X or Y gradient coil (251 in FIG. 12).
  • the shimming coil assembly 271 is mounted outside of the vacuum vessel, between the magnet bore and gradient coil assembly, or inside the gradient coil assembly, then the shimming coils are made of copper wire or copper foil strip.
  • the shim coils can alternatively be manufactured as flexible printed circuits. If the shimming coil assembly 271 is mounted inside the vacuum vessel, for example inside the inner diameter of the magnet coils, then the shimming coils may be cooled and made of HTS wire or tape allowing them to carry a significantly higher current.
  • a shimming element assembly 291 includes a cylindrical former 282 provided with a regular array of pockets 283, 284 arranged in axial rows and circumferential rings.
  • the cylindrical former 282 is an electrically non-conductive support, for example, a fiberglass tube.
  • Each pocket (typically of rectangular shape) may be filled with thin steel plates 285, 286 (called “shim plates”) to provide a depth of iron. The iron becomes axially magnetized when the magnet is energized and the each shim's small dipole field adds to the magnet field, and locally changes the magnetic field strength.
  • the size, shape, and location of the ferromagnetic shimming elements 285, 286 on the cylindrical former 282 can be determined empirically for each particular transportable MRI system.
  • the superconducting magnet would be perfectly symmetrical about the axis of the magnet bore, and the center of the superconducting magnet (131 in FIG. 3) would be located at the center of the shipping container (101 in FIG.
  • the superconducting magnet is mounted to the shipping container to locate the superconducting magnet with respect to the ferromagnetic sheets of the secondary shielding and the walls of the shipping container with an accuracy of better than 10 mm in all dimensions.
  • the system is assembled and tested without the shimming element assembly 281 and when the shimming coils 273, 274 are not energized.
  • the depth of shimming elements in each pocket required to correct the measured inhomogeneity in the imaging volume are determined using conventional electromagnetic modeling and optimization methods. The modeling techniques assume that the ferromagnetic shimming elements 285, 286 become uniformly magnetized due to saturation caused by the magnetic field of the superconducting magnet. Once the distribution of the shimming elements 285, 286 has been loaded into the array of pockets, the shimming element assembly 281 is inserted in the magnet bore. Fine adjustment of the magnetic field inhomogeneity is achieved by driving appropriate direct current through the active shim coils (273, 274 in FIG. 15) using an electrical shim power supply unit (152 in FIG. 3). The currents are adjusted to obtain the best homogeneity of the magnetic field in the imaging volume.
  • Pulsing of gradient coils during imaging can induce transient eddy currents in the magnet structure which distort the homogeneity in the imaging volume such that the resultant field is not the desired superposition of the uniform static background field from the magnet and the linear gradient field from the gradient coil being pulsed.
  • pulsing a Y gradient coil can introduce eddy currents which generate a transient X or Z or Z 2 field inhomogeneity, or indeed fields with arbitrary shape, depending on the distribution of the eddy currents.
  • FIG. 17 shows a method of transporting the transportable MRI system over a distance of several days travel to a target location and using the transportable MRI system at the target location.
  • a first box 301 the transportable MRI system is shipped to the target location over the distance requiring several days of travel when the system is de-energized.
  • box 302 at the target location, power is supplied to the cryogenic refrigerator and the magnet is cooled with the cryogenic refrigerator so that the magnet becomes superconductive.
  • the magnet is energized using the controlled current source.
  • the magnet's homogeneity is checked using a standardized imaging phantom and diagnostic routines, and the magnet's homogeneity is shimmed using the electrical active shims when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity.
  • MRI magnetic resonance imaging
  • FIG. 18 shows a method of transporting the transportable MRI system to a target location while the generator is running, and using the transportable MRI system at the target location.
  • a certain duration of time that is acceptable for continuous use of the generator which may depend of the size of the fuel tank, the refill frequency, and possibly a need for maintenance.
  • the time that is acceptable for continuous use of the generator is five days.
  • the system is shipped to the target location over a distance requiring no more than five days travel, with the on-board generator running to power the cryogenic refrigerator to keep the magnet cold but de-energized.
  • the magnet is energized using the controlled current source.
  • the magnet's homogeneity is checked using a standardized imaging phantom and diagnostic routines, and the magnet's homogeneity is shimmed using the electrical shims when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity.
  • magnetic resonance imaging (MRI) scans are carried out using the magnet.
  • FIGS. 19 and 20 show an alternative construction for a cooling manifold 382 and a superconducting magnet assembly of layer wound winding blocks 371, 372, 373, 374.
  • This alternative construction can be substituted for the construction shown in FIGS. 6 and 7 using pancake coils.
  • Pancake coils are best suited to superconducting wire having a high aspect ratio cross section (i.e., tape).
  • Layer wound winding blocks are best suited to superconducting wire having a low aspect ratio, such as low aspect ratio MgB2 conductor.
  • Each layer of a layer wound winding block is a helical coil of the superconducting wire, and there can be multiple layers and many turns per layer, in order to produce a winding block having a much lower diameter to length ratio than pancake coils.
  • the layer wound winding blocks are easier to construct than the pancake coils by winding the superconducting wire over a cylindrical mandrel or bobbin.
  • the lower diameter to length ratio of the layer wound winding blocks presents less surface area at their ends for thermal coupling of the winding blocks to annular cooling plates 401, 402, 403, 405, 406, 407, 408, of the cooling manifold 382.
  • the cooling manifold 382 comprises a cylindrical tube 400, and the winding blocks 371, 372, 373, 375 are fitted inside the cylindrical tube for heat conduction from an outer diameter of the winding blocks to an inner diameter of the cylindrical tube.
  • the cylindrical tube 400 is seated on and thermally coupled to an axial thermal bus 411.
  • the thermal bus 411 is thermally coupled by a flexible copper braid 505 to the cold element 335 of the second stage of a cryogenic refrigerator 334.
  • the cylindrical tube 400 and the annular plates 401, 402, 403, 404, 405, 406, 407, 408 and thermal bus 411 are constructed from a material with good thermal conductivity, such as copper, aluminum, or sapphire, or a combination of these materials.
  • the annular plates 401, 402, 403, 405, 406, 407, 408, have a snug fit inside the outer cylindrical tube 400.
  • the annular plates are soldered to the internal diameter of the cylindrical tube, for good heat transfer between the plates and the tube.
  • one of the end plates 407 or 408 is first soldered to an end of the tube 400, and then one of the end winding blocks 371 or 374 is inserted into the tube, and then a next one of the plates 402 or 407 is inserted into the tube 400 and soldered to the internal diameter of the tube.
  • one of the annular spacers 412 or 414 is inserted into the tube 400, and then a next of the plates 403 or 406 is inserted into the tube 400 and soldered to the internal diameter of the tube.
  • the central annular spacer 413 is inserted into the tube, and a next one of the plates 406 or 403 is inserted into the tube and soldered to the internal diameter of the tube 400.
  • the outer diameter and end faces of the winding blocks 371, 372, 373, 374 can be coated with thermally conductive grease to assist heat transfer from the winding blocks to the tube 400 and to the plates.
  • the tube 400 and the flexible copper braid 505 are soldered to the axial thermal bus 411 to assist heat transfer to the cold element 335 of the second stage of the cryogenic refrigerator 334.
  • the cooling manifold 382 is geometrically configured and constructed not only to effectively carry heat from the winding blocks 371, 372, 373, 374 to the cold element 225, but also to reduce eddy currents from the gradient coils during imaging.
  • the cylindrical tube 400 and each of the annular plates 401, 402, 403, 404, 405, 406 contains at least one complete cut (421, 422 in FIG. 19) to disrupt formation of circumferential eddy currents in the cylindrical tube and in each of the cooling plates.
  • the complete cuts (421, 422 in FIG. 19) are diametrically opposite from the thermal bus 411.
  • the cylindrical tube 400 and each of the annular plates 401, 402, 403, 404, 405, 406 may have additional cuts to further disrupt formation of eddy currents.
  • the number and arrangement of the additional cuts may be chosen to significantly disrupt the formation of eddy currents while retaining some of the structural rigidity of the cylindrical tube 400 and the annular plates 401, 402, 403, 404, 405, 406.
  • the annular plate 401 has a circular array of radial slits 423, 424 extending from its internal diameter to nearly its outer diameter.
  • the radial cut 422 and the radial slits 423, 424 are spaced at equal angular intervals around the annular plate 401.
  • the other annular plates may have a similar array of radial slits, and each of the other annular plates may have the same geometrical configuration and orientation as the annular plate 401.
  • the cylindrical tube 400 may also have a circular array of axial slits 425, 426 extending from one or both of its ends.
  • the cylindrical tube 400 has a circular array of axial slits 425, 426 extending from a first end of the tube 400 nearly to a second end of the tube, and a circular array of axial slits 427, 428 extending from the second end of the tube nearly to the first end of the tube.
  • the slits are arranged so that the cut (421 in FIG. 19) and the slits are spaced at equal angular intervals round the circumference of the tube 400, alternate slits extend from alternate ends of the tube 400, and contiguous material of the tube 400 forms a serpentine path around the tube.
  • the superconducting magnet overcomes the complex logistics associated with shipping and installing a conventional liquid helium cooled superconducting magnet. Routine diagnostic MRI is made affordable in developing economies, or in locations where installation of conventional fixed or mobile MRIs is uneconomic or impractical.
  • a self-contained transportable magnetic resonance imaging (MRI) system comprising: a standard shipping container; a vacuum vessel mounted within the shipping container; a superconducting magnet suitable for magnetic resonance imaging (MRI), the superconducting magnet being contained within the vacuum vessel, the superconducting magnet including a stack of pre-formed coils of high temperature superconducting wire, and the vacuum vessel having an outer diameter wall of ferromagnetic material for magnetically shielding the superconducting magnet; a cryogenic refrigerator mounted within the shipping container, the cryogenic refrigerator having at least one cold element in the vacuum vessel; a cooling manifold in the vacuum vessel, the cooling manifold thermally coupling the at least one cold element to the coils for cooling the superconducting magnet; and ferromagnetic material added to faces of the shipping container for reducing a fringe magnetic field footprint outside of the shipping container.
  • MRI magnetic resonance imaging
  • cryogenic refrigerator and the cooling manifold are configured to maintain the temperature of the coils between ten degrees Kelvin and thirty degrees Kelvin when the superconducting magnet is energized.
  • the superconducting magnet has a bore
  • the bore has a diameter
  • the coils are positioned coaxially around the bore
  • the coils are configured in winding blocks thermally coupled to the cooling manifold
  • the winding blocks are dimensioned to provide a magnetic field with homogeneity of less than twenty parts per million (ppm) peak-to-peak over an imaging volume within the bore of at least forty percent of the bore diameter.
  • a seventh example there is disclosed a system according to the preceding fifth example, wherein the winding blocks are made from stacks of double pancake coils.
  • cooling manifold includes annular plates of high thermal conductivity material disposed on faces of the double pancake coils, and a thermal bus thermally connecting the annular plates to the at least one cold element of the cryogenic refrigerator.
  • cooling manifold comprises a cylindrical tube of high thermal conductivity material, and a thermal bus thermally connecting the cylindrical tube to the at least one cold element of the cryogenic refrigerator, and the winding blocks are fitted inside the cylindrical tube for heat conduction from an outer diameter of the winding blocks to an inner diameter of the cylindrical tube.
  • a system according to the preceding twelfth example further including a thermal radiation shield disposed outside an outer diameter of the coils and the multi-layer insulation blanket, and wherein the cryogenic refrigerator has a second cold element and is configured to maintain the second cold element at a higher temperature and with a higher cooling power than said at least one cold element, and the thermal radiation shield is thermally anchored to the second cold element.
  • a fourteenth example there is disclosed a system according to any of the preceding first to thirteenth examples, further including a combination of robust axial and radial supports rigidly locating the superconducting magnet within the vacuum vessel, wherein at least the axial supports are arranged in compression for ensuring that the superconducting magnet remains accurately located within the vacuum vessel during cool down, energization of the superconducting magnet, and under exposure to external vibration.
  • a fifteenth example there is disclosed a system according to any of the preceding first to fourteenth examples, wherein the superconducting magnet has a central flux density between 0.5 Tesla and 1 Tesla.
  • the ferromagnetic material added to the faces of the shipping container includes sheets of mild steel, and the ferromagnetic material is added only to some faces of the container.
  • a seventeenth example there is disclosed a system according to any of the preceding first to sixteenth examples, further including localized radio frequency shielding added to the shipping container to close radio frequency (RF) signal leaks and configure the portion of the shipping container containing the superconducting magnet to be a Faraday cage.
  • RF radio frequency
  • the ferromagnetic material added to the faces of the shipping container includes ferromagnetic sheets
  • the superconducting magnet is mounted to the shipping container to locate the superconducting magnet with respect to the ferromagnetic sheets and the walls of the shipping container with an accuracy of better than 10 mm in all dimensions.
  • the shimming elements include movable ferromagnetic shimming elements mounted on the surface of the ferromagnetic vacuum vessel outer shell.
  • shimming elements include electrical shim coils comprising at least one of superconducting coils in the vacuum vessel and copper coils within the bore and outside of the vacuum vessel.
  • a system according to any of the preceding first to twenty-first examples further including a current source outside of the vacuum vessel and electrically connected to at least one of the magnet coils and an electrical shim coil and arranged to stabilize a magnetic field in a bore of the superconducting magnet by feedback of a nuclear magnetic resonance (MR) frequency of the superconducting magnet.
  • MR nuclear magnetic resonance
  • a system according to any of the preceding first to twenty-second examples further including a backup uninterruptable power supply to maintain power to the cryogenic refrigerator when there is a loss of external power to the system.
  • cooling manifold includes a thermal bus oriented along an axis of the superconducting magnet, and cooling plates sandwiched between neighboring coils of the superconducting magnet, and each of the cooling plates is thermally anchored to the thermal bus and contains at least one complete cut to disrupt formation of circumferential eddy currents in said each of the cooling plates, and the axial bus is connected to the at least one cold element of the cryogenic refrigerator.
  • a twenty-fifth example there is disclosed a system according to any of the preceding first to twenty-fourth examples, wherein the high temperature superconducting wire is at least one of MgB 2 and REBCO, and the high temperature superconducting wire does not incorporate a copper co-winding for the purpose of carrying magnet current in the event of a quench.
  • a twenty-sixth example there is disclosed a system according to any of the preceding first to twenty-fifth examples, in which the gradient coils are at least partly unshielded such that return flux from the gradient coils penetrates the vacuum vessel, yet the return flux does not generate eddy currents in the magnet structure which at least partly mirror the conductor paths of the gradient coils.
  • a twenty- seventh example there is disclosed a system according to any of the preceding first to twenty-sixth examples, wherein the use of gradient pulse pre- emphasis and pulsed currents applied to selected shim coils are sufficient to counteract impact of eddy currents induced in at least one of the superconducting magnet and the vacuum vessel in routine clinical imaging sequences.
  • the superconducting magnet has a bore, and there is an imaging volume within the bore and outside of the vacuum vessel for imaging the whole body of the patient, and the imaging volume has a diameter of at least forty centimeters as defined by a peak to peak variation of +/- 20 ppm, and the MRI system has a total weight of less than twenty metric tons.
  • a system according to any of the preceding first to twenty-eighth examples, which further includes a set of barrier panels transportable within the shipping container and configured for being set up outside of the shipping container for constructing a barrier to exclude people from a magnetic field in excess of five gauss produced by the superconducting magnet and extending outside of the shipping container.
  • a system according to any of the preceding first to twenty-ninth examples, which further includes a patient table slidably mounted with respect to the superconducting magnet for sliding a patient through the bore, and the superconducting magnet defining an imaging volume in the bore, and an NMR probe disposed within the imaging volume and offset from the patient when the patient is slid through the bore, wherein the NMR probe is comprised of a material capable of generating an NMR signal.
  • a magnetic resonance imaging (MRI) system comprising: an imaging magnet having a bore and defining an imaging volume within the bore; a patient table slidably mounted with respect to the imaging magnet for sliding a patient through the bore, and the imaging magnet defining an imaging volume in the bore, and a nuclear magnetic resonance (NMR) probe disposed within the imaging volume and offset from the patient when the patient is slid through the bore, wherein the NMR probe is comprised of a material capable of generating an NMR signal.
  • MRI magnetic resonance imaging
  • a system according to any of the preceding examples thirty-second to thirty-fourth, further including an NMR signal receiving coil disposed about the NMR probe for receiving the NMR signal generated by the NMR probe.
  • a magnetic resonance imaging (MRI) system comprising: a vacuum vessel; a superconducting magnet suitable for magnetic resonance imaging (MRI), the superconducting magnet being contained within the vacuum vessel, the superconducting magnet including a stack of pre- formed coils of high temperature superconducting wire;
  • cryogenic refrigerator attached to the vacuum vessel and having at least one cold element in the vacuum vessel; and a cooling manifold in the vacuum vessel, the cooling manifold thermally coupling the at least one cold element to the coils for cooling the superconducting magnet
  • the cooling manifold includes a thermal bus oriented along an axis of the superconducting magnet, and cooling plates sandwiched between neighboring coils of the superconducting magnet, and each of the cooling plates is thermally anchored to the thermal bus and contains at least one complete cut to disrupt formation of circumferential eddy currents in said each of the cooling plates, and the axial bus is connected to the at least one cold element of the cryogenic refrigerator.
  • a method of using a transportable magnetic resonance imaging (MRI) system comprising: shipping the system to the target location over a distance requiring several days of travel when the system is de-energized, and at the target location: supplying power to the cryogenic refrigerator and cooling the magnet with the cryogenic refrigerator so that the magnet becomes superconductive; energizing the magnet using the controlled current source; checking the magnet's homogeneity using a standardized imaging phantom and diagnostic routines; and shimming the magnet's homogeneity when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity; and carrying out magnetic resonance imaging (MRI) scans using the magnet.
  • MRI magnetic resonance imaging
  • a method of using a transportable magnetic resonance imaging (MRI) system comprising: shipping the system to a target location with the on-board generator running to power the cryogenic refrigerator to keep the magnet cold but de-energized; and then, at the target location: energizing the magnet using the controlled current source; checking the magnet's homogeneity using a standardized imaging phantom and diagnostic routines and shimming the magnet's homogeneity when the checking of the magnet's homogeneity indicates a need for shimming the magnet's homogeneity; and carrying out magnetic resonance imaging (MRI) scans using the magnet.
  • MRI magnetic resonance imaging

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  • Engineering & Computer Science (AREA)
  • Power Engineering (AREA)
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Abstract

L'invention porte sur un système d'imagerie à résonance magnétique pour le corps tout entier, qui est peu coûteux, compact et facilement transportable, lequel système se loge entièrement dans un conteneur d'expédition standard. Dans ce système, un aimant supraconducteur à haute température est refroidi par conduction, n'utilise pas de liquide cryogénique et présente une force de champ comprise entre 0,5 tesla et 1 tesla. L'aimant supraconducteur évite la logistique complexe associée à l'expédition et à l'installation d'un aimant supraconducteur refroidi à l'hélium liquide classique. Une imagerie à résonance magnétique de diagnostic de routine est rendue abordable dans des économies en développement, ou dans des endroits où l'installation d'appareils d'imagerie à résonance magnétique fixes ou mobiles classiques n'est pas économique ni pratique.
PCT/IB2015/054395 2014-06-11 2015-06-10 Système d'imagerie à résonance magnétique transportable WO2015189786A1 (fr)

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CN113330525A (zh) * 2018-11-22 2021-08-31 托卡马克能量有限公司 部分绝缘超导磁体的快速泄放
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CN113631940A (zh) * 2019-03-22 2021-11-09 皇家飞利浦有限公司 用于控制持续电流开关的温度的系统
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WO2023094611A1 (fr) * 2021-11-25 2023-06-01 Tokamak Energy Ltd Exploitation en rampe d'aimant hts pour réduire les courants d'écrantage
WO2023129010A1 (fr) * 2021-12-30 2023-07-06 T.C. Ankara Universitesi Rektorlugu Aimant destiné à être utilisé dans des dispositifs d'imagerie par résonance magnétique
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CN113330525B (zh) * 2018-11-22 2023-09-29 托卡马克能量有限公司 部分绝缘超导磁体的快速泄放
CN113330525A (zh) * 2018-11-22 2021-08-31 托卡马克能量有限公司 部分绝缘超导磁体的快速泄放
CN111505549A (zh) * 2019-01-31 2020-08-07 中加健康工程研究院(合肥)有限公司 一种可移动的mri系统
CN113631940A (zh) * 2019-03-22 2021-11-09 皇家飞利浦有限公司 用于控制持续电流开关的温度的系统
CN113631940B (zh) * 2019-03-22 2024-04-05 皇家飞利浦有限公司 用于控制持续电流开关的温度的系统
CN113495238A (zh) * 2020-04-07 2021-10-12 中国航天科工飞航技术研究院（中国航天海鹰机电技术研究院） 有背景磁场的动态超导磁体热负载测试系统
US20220340274A1 (en) * 2021-04-26 2022-10-27 Joulea Llc Remote deployable transient sensory kit
EP4159641A1 (fr) * 2021-09-29 2023-04-05 David Carmine Raymond Rovere Dispositif de transport et de traitement d au moins un aimant
FR3127625A1 (fr) * 2021-09-29 2023-03-31 Carmine Raymond ROVERE David Dispositif de transport et de traitement d’au moins un aimant
WO2023094611A1 (fr) * 2021-11-25 2023-06-01 Tokamak Energy Ltd Exploitation en rampe d'aimant hts pour réduire les courants d'écrantage
WO2023129010A1 (fr) * 2021-12-30 2023-07-06 T.C. Ankara Universitesi Rektorlugu Aimant destiné à être utilisé dans des dispositifs d'imagerie par résonance magnétique
EP4300119A1 (fr) * 2022-06-29 2024-01-03 Siemens Healthcare Limited Structure de bus thermique pour dispositif d'imagerie par résonance magnétique

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