Classifications

• • 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
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0033—Features or image-related aspects of imaging apparatus, e.g. for MRI, optical tomography or impedance tomography apparatus; Arrangements of imaging apparatus in a room
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/05—Detecting, measuring or recording for diagnosis by means of electric currents or magnetic fields; Measuring using microwaves or radio waves
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/05—Detecting, measuring or recording for diagnosis by means of electric currents or magnetic fields; Measuring using microwaves or radio waves
  • A61B5/055—Detecting, measuring or recording for diagnosis by means of electric currents or magnetic fields; Measuring using microwaves or radio waves involving electronic [EMR] or nuclear [NMR] magnetic resonance, e.g. magnetic resonance imaging
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/70—Means for positioning the patient in relation to the detecting, measuring or recording means
  • A61B5/704—Tables
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/0005—Geometrical arrangement of magnetic sensor elements; Apparatus combining different magnetic sensor types
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/02—Measuring direction or magnitude of magnetic fields or magnetic flux
  • G01R33/035—Measuring direction or magnitude of magnetic fields or magnetic flux using superconductive devices
• • 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
  • G01R33/30—Sample handling arrangements, e.g. sample cells, spinning mechanisms
  • G01R33/307—Sample handling arrangements, e.g. sample cells, spinning mechanisms specially adapted for moving the sample relative to the MR system, e.g. spinning mechanisms, flow cells or means for positioning the sample inside a spectrometer
• • 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
  • G01R33/421—Screening of main or gradient magnetic field
  • G01R33/4215—Screening of main or gradient magnetic field of the gradient magnetic field, e.g. using passive or active shielding of the gradient magnetic field
• • 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
  • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
  • G01R33/48—NMR imaging systems
  • G01R33/54—Signal processing systems, e.g. using pulse sequences ; Generation or control of pulse sequences; Operator console
  • G01R33/56—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution
  • G01R33/563—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution of moving material, e.g. flow contrast angiography
  • G01R33/56375—Intentional motion of the sample during MR, e.g. moving table imaging
  • G01R33/56383—Intentional motion of the sample during MR, e.g. moving table imaging involving motion of the sample as a whole, e.g. multistation MR or MR with continuous table motion

Definitions

• the present disclosure relates to magnets for use in magnetic resonance imaging and Magnetic Resonance Imaging (MRI) systems.
• an invention of the disclosure relates to disk-type magnetic resonance imaging systems.
• Magnetic resonance imaging was introduced in the 1980s, and has developed into a major medical imaging modality.
• MRI Magnetic spherical imaging volume
• Deviations from homogeneity of the static magnetic field over the DSV are typically required to be less than 20 parts per million peak-to-peak (or 2 parts per million rms).
• MRI equipment has undergone a number of refinements since the introduction of the first closed cylindrical systems.
• improvements have occurred in quality/resolution of images through improved signal to noise ratios and introduction of high and ultra-high field magnets.
• Improved resolution of images has led to MRI being a modality of choice for an increasing number of specialists for both structural anatomical and functional human MRI imaging.
• the basic components of a typical magnetic resonance system for producing diagnostic images for human studies include a main magnet (usually a superconducting magnet which produces the substantially homogeneous static magnetic field (the Bo field) in the DSV), one or more sets of shim coils, a set of gradient coils, and one or more RF coils.
• a main magnet usually a superconducting magnet which produces the substantially homogeneous static magnetic field (the Bo field) in the DSV
• the Bo field substantially homogeneous static magnetic field
• RF coils RF coils.
• the typical aperture of a conventional MRI machine after the addition of ancillary components is a cylindrical space having a diameter of about 0.6-0.8 meters, i.e., just large enough to accept the subject's shoulders, and a length of about 2.0 meters or more.
• ancillary components grades and radiofrequency coils
• many people suffer from claustrophobia when placed in such a space.
• the large distance between the portion of the subject’s body which is being imaged and the end of the magnet system means that physicians cannot easily assist or personally monitor a subject during an MRI procedure. Therefore, there is a need for a short or compact magnet system in clinical applications.
• the challenge in designing such a high-field system is maintaining both the field homogeneity and size of the DSV using the currently available, cost- effective, superconducting technology.
• the magnet performance is largely related to the bore size in both axial and radial directions. Short or compact magnets are very difficult to design and build. This is mainly because the dense coil structure produced by conventional designs will lead to unacceptable peak field values and stress for the superconducting coil bundles. Normally, an engineering compromise in DSV size has to be made relative to the size of the magnet.
• the size of the magnet is a primary factor in determining the cost of an MRI machine, as well as the costs involved in the installation of such a machine. Another important consideration is the volume of helium needed to maintain the system at cryogenic temperatures. Due to their large size, such whole body magnets are expensive for use in producing images of small sizes of objects, such as, heads, extremities and neonates, etc.
• MRI machines In order to be used safely, MRI machines often need to be shielded so that the magnetic fields surrounding the machine at the location of the operator are below regulatory agency-specified exposure levels. By means of shielding, the operator can be safely sited much closer to the magnet than in an unshielded system. Longer magnets require more shielding and larger shielded rooms for such safe usage, thus leading to higher costs.
• the invention resides in a magnetic resonance imaging system for producing magnetic resonance images through moving human subjects.
• the system comprises a magnet that produces a disk-type homogeneous magnetic field region; a gradient coil; and a radio frequency (RF) coil.
• RF radio frequency
• the invention resides in method for moving a patient through a magnetic resonance imaging system to conduct imaging during movement and/or at multiple positions.
• the invention resides in a magnet suitable for use in a Magnetic Resonance Imaging (MRI) system, the magnet having a magnet body having a bore extending therethrough along an axis of the body, the magnet comprising: a primary coil structure having at least four primary coils positioned along an axis, including a first end coil adjacent a first end of the bore of the magnet and a second end coil adjacent a second end of the magnet,
• MRI Magnetic Resonance Imaging
• first end coil and the second end coil are spaced apart by no more than 1000mm
• an imaging region produced by the primary coils is of a disk- type.
• the disk-type imaging region has an axial diameter (Dz) and a radial diameter (Dr), wherein the axial diameter is less than the radial diameter. More preferably, a ratio of the axial diameter to the radial diameter of the disk-type imaging region is equal to or less than 0.75.
• a diameter of the imaging region along an x-axis is between 100mm and 500mm.
• a diameter of the imaging region along a y- axis is between 100mm and 500mm.
• a diameter of the imaging region along a z-axis is between 20mm and 350mm.
• the imaging region has dimensions of 250mm(x-)x250mm(y- )x40mm(z-).
• the imaging region has dimensions of 320mm(x-)x320mm(y-)x100mm(z-).
• the imaging region has dimensions of 450mm(x-)x450mm(y- )x100mm(z-).
• the imaging region has dimensions of 300mm(x-)x300mm(y-)x100mm(z-).
• the first end coil and the second end coil are spaced apart by between 300mm and 1000mm.
• the primary coil structure has between four primary coils and eight primary coils.
• the primary coil structure has four primary coils, five primary coils, six primary coils or seven primary coils.
• each of the primary coils has a same current polarity.
• one of the at least four primary coils adjacent the second end coil has an opposite current polarity to the second end coil.
• the primary coil structure is symmetric relative to an axial centre of the imaging region.
• the primary coil structure is asymmetric relative to an axial centre of the imaging region.
• the magnet body and bore are cylindrical, conical, frustoconical or stepped.
• the magnet body and bore comprise at least one cylindrical portion.
• a cylindrical portion adjoins a frustoconical portion.
• a first cylindrical portion having a diameter adjoins a second cylindrical portion having a diameter, wherein the diameter of the first cylindrical portion is greater than the diameter of the second cylindrical portion.
• a plurality of frustoconical portions and/or cylindrical portions define a stepped-diameter bore.
• an inner diameter of the bore is between 200mm and 1 100mm.
• the bore is frustoconical having a largest diameter of 820mm and a smaller diameter of 282mm.
• the primary coils are cylindrical, conical, frustoconical or stepped.
• the magnet comprises a stepped diameter bore having at least one primary coil located about a first step of the stepped diameter bore and at least one primary coil located about a second step of the stepped diameter bore.
• the magnet is capable of producing a magnetic field of at least 1.0 Tesla and more preferably a magnetic field of at least 3.0 Tesla.
• the magnetic field is substantially homogenous over a predetermined imaging region.
• the imaging region has an external surface defined by a computed variation of a longitudinal magnetic field relative to the longitudinal magnetic field at an imaging centre of less than 20 parts per million peak-to- peak.
• the invention resides in a magnetic resonance imaging system having a magnet suitable for use in a Magnetic Resonance Imaging (MRI) system, the magnet having a bore extending along an axis of the magnet, the magnet comprising:
• a primary coil structure having at least four primary coils positioned along an axis, including a first end coil adjacent a first end of the bore of the magnet and a second end coil adjacent a second end of the magnet,
• first end coil and the second end coil are spaced apart by no more than 1000mm
• the magnet or the MRI system further comprises a shield coil structure.
• the shield coil structure is located around the primary coil structure (i.e. has a larger diameter than the primary coil structure).
• the shield coil structure comprises at least one shield coil having a greater diameter than the primary coils.
• the shield coil structure is located radially outwardly of the primary coil structure.
• the shield coil structure has at least two shield coils.
• each of the shield coils carry current in a direction opposite to a direction of current in the first and second end coils of the primary coil structure.
• each of the shield coils are superconducting or ferromagnetic.
• the shield coils tailor the magnetic fields within the imaging region.
• the magnet or the MRI system further comprises a gradient coil structure comprising a primary coil layer and a shield coil layer.
• a length of the primary coil layer of the gradient coil structure is less than a length of the shield coil layer of the gradient coil structure. More preferably, the length of the primary coil layer of the gradient coil structure is significantly less than the length of the shield coil layer of the gradient coil structure.
• the shield coil layer can be closer to the primary coil layer to reduce the gradient coil thickness while still maintaining good shielding performance.
• the gradient coil structure is located within a gradient body located within the magnet.
• the gradient body is located between the bore and the magnet body.
• the magnet or the MRI system further comprises one or more Radio Frequency (RF) coils located between the gradient coil structure and the bore.
• the RF coils are frustoconical and/or cylindrical confirming to a shape of the bore.
• the RF coils are located on an inner surface of the gradient body surrounding the bore.
• the system further comprises one or more shim pockets.
• the shim pockets are frustoconical and/or cylindrical.
• a shim portion is located in each shim pocket.
• the shim portion comprises ferrous or ferromagnetic material.
• each primary coil has an associated shim pocket and shim portion having a shape conforming to the shape of the magnet body and/or the bore.
• the shim portions passively shim the imaging region to achieve a preferred field (Bo) homogeneity level.
• the shim device is located between the primary coil structure and the shield coil structure. In some embodiments, the shim device is located outside of the shield coil structure. Preferably, the shim device is located between the magnet and gradient coils.
• a length of the bore is between 250mm and 1000mm.
• the length of the bore is 300mm, 570mm, 600m, 800mm or 900mm.
• a size of a five Gauss line is between 1.5m and 6m in a radial direction and between 2.5m and 9m in an axial direction.
• the five Gauss line has dimensions of:
• the magnetic resonance imaging system comprises a movable platform or portion adapted to support a patient.
• the movable platform or portion is adapted to move through the bore of the magnetic resonance imaging system.
• the invention resides in a method of magnetic resonance imaging scanning, the method comprising the steps of:
• the magnet resonance imaging system having a magnet as described above.
• a magnet suitable for use in a Magnetic Resonance Imaging (MRI) system having a magnet body with a bore extended therethrough along an axis of the body, the magnet comprising:
• a primary coil structure having at least four primary coils positioned coaxially along the axis, including a first end coil adjacent a first end of the bore and a second end coil adjacent a second end of the bore, the primary coil structure arranged to generate a substantially homogenous disk shaped imaging region within the bore.
• Figure 1 illustrates a schematic cross-section view of a magnetic resonance imaging system according to a first embodiment of the present invention
• Figure 1 A illustrates a magnet coil configuration and DSV dimensions of the magnetic resonance imaging system of Figure 1 ;
• Figure 1 B illustrates the stray field outside the magnet and the 5 Gauss line of the magnetic resonance imaging system of Figure 1 ;
• Figure 2 illustrates a schematic cross-section view of a magnet for use in a magnetic resonance imaging system according to a second embodiment of the present invention
• Figure 3 illustrates a schematic cross-section view of a magnet for use in a magnetic resonance imaging system according to a third embodiment of the present invention
• Figure 3A illustrates a magnet coil configuration and DSV dimensions of the magnetic resonance imaging system of Figure 3;
• Figure 3B illustrates the stray field outside the magnet and the 5 Gauss line of the magnetic resonance imaging system of Figure 3;
• Figure 4 illustrates a schematic cross-section view of a magnet for use in a magnetic resonance imaging system according to a fourth embodiment of the present invention
• Figure 4A illustrates a magnet coil configuration and DSV dimensions of the magnetic resonance imaging system of Figure 4.
• Figure 4B illustrates the stray field outside the magnet and the 5 Gauss line of the magnetic resonance imaging system of Figure 4;
• Figure 5 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to a fifth embodiment of the present invention
• Figure 6 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to a sixth embodiment of the present invention
• Figure 6A illustrates a magnet coil configuration and DSV dimensions of the magnetic resonance imaging system of Figure 6;
• Figure 6B illustrates the stray field outside the magnet and the 5 Gauss line of the magnetic resonance imaging system of Figure 6;
• Figures 7 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to a seventh embodiment of the present invention
• Figure 8 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to an eighth embodiment of the present invention.
• Figure 9 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to a ninth embodiment of the present invention.
• Figure 10 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to a tenth embodiment of the present invention.
• Figure 1 1 illustrates a schematic cross-section of a magnet for use in a magnetic resonance imaging system according to an eleventh embodiment of the present invention.
• the system 01 includes a magnet 101 being a superconducting magnet, a gradient body 102, passive shim pockets 103, a RF coil 104 and a bore 105 extending axially through the magnet 101 parallel to longitudinal axis 106.
• the bore 105 of the magnet 101 has a cold bore length L1 (preferably between 250mm and 1000mm) and an inner bore diameter D1 [0068]
• the magnet 101 is symmetric with respect to the centre of the Diameter of Spherical Volume (DSV)/imaging region and comprises a primary coil structure 1 10 having four primary coils: a first end coil 1 12a, a second end coil 1 12b, and two middle coils 1 13a, 1 13b located between the first end coil 1 12a and second end coil 1 12b.
• Each of the coils 1 12a, 1 12b, 1 13a, 1 13b have the same (positive) polarity.
• the respective end coils 112a, 1 12b, located at opposite ends of the bore 105 are spaced apart by distance C1 , which is preferably between 300mm and 1000mm. It will be appreciated that the distance between the end coils 1 12a, 112b is measured from a centre of one coil to a centre of the other coil.
• the system 01 also includes two shield coils 1 14a and 114b with opposite (negative) polarity to the polarity of the primary coils 1 12a, 1 12b, 1 13a, 1 13b, which produces a disk-type DSV 1 11 having an axial diameter Dzi less than the radial diameter Dn, where the ratio of the axial diameter Dzi to the radial diameter Dn of the DSV 21 1 is equal to or less than 0.75.
• a gradient body 102 is located within the magnet 101. Inside the gradient body 102, there is located a primary gradient coil layer 121 and a shield gradient coil layer 122, generating three orthogonal gradient fields in three orthogonal z, x and y axes.
• the electric current directions in the shield gradient coils of the shield gradient coil layer 122 are opposite to those in the relevant primary gradient coils of the primary gradient coil layer 121 .
• the length (Lp) of the primary gradient coil layer 121 is significantly shorter than the length (Ls) of the shield gradient coil layer 122, allowing closer radial distance between the primary and shield layers while still maintaining good shielding performance.
• the RF coil 104 is also short and therefore efficient due to a shorter DSV.
• Figures 1 A and 1 B respectively illustrate the magnet configuration and DSV, and the 5 Gauss lines of the 1.5 Tesla superconducting magnet 101 a.
• Superconducting magnet 101 a has the same primary coil structure 1 10 as superconducting magnet 101. As shown, the size of the DSV 1 1 1 a is 250mm (X-) x 250mm (Y-) x 40mm(Z-), providing a dimensional ratio of Dzi/Dn of 40mm/250mm which is equal to 0.16, and peak to peak homogeneity of 10ppm. [0074]
• the bore 105a of the magnet 101 a has a cold bore length of 570mm and an inner bore diameter of 570mm.
• the size of the 5 Gauss line is 3 meters in the radial direction and 5 meters in the axial direction, which demonstrates well controlled stray fields.
• FIG. 2 there is shown a superconducting magnet 02 suitable for use in a magnetic resonance imaging system having a primary coil structure 210 according to an embodiment of the present invention.
• the magnet 02 includes a bore 201 extending axially through the magnet 02 parallel to longitudinal axis 202.
• the bore 201 of the magnet 02 has a cold bore length L2 (preferably between 250mm and 1000mm) and an inner bore diameter D2.
• the primary coil structure 210 of the magnet 02 includes five primary coils: a first end coil 212a, a second end coil 212b, and three intermediate coils 213a, 213b and 214 arranged symmetrically with respect to the centre of DSV 21 1 , with intermediate coil 214 spanning over the centre of the DSV 21 1.
• the first end coil 212a and the second end coil 212b are spaced apart by distance C2, which is preferably between 300mm and 1000mm.
• All of the five coils 212a, 212b, 213a, 213b and 214 of the primary coil structure 210 have the same polarity, producing the disk-type DSV 21 1 having an axial diameter Dz2 less than the radial diameter Dr2, where the ratio of the axial diameter Dz2 to the radial diameter Dr2 of the DSV 21 1 is equal to or less than 0.75.
• FIG. 3 there is shown a superconducting magnet 03 suitable for use in a magnetic resonance imaging system, the magnet 03 having a primary coil structure 310 according to an embodiment of the present invention.
• the magnet 03 includes a bore 301 extending axially through the magnet 03 parallel to longitudinal axis 302.
• the magnet 03 has a cold bore length L3 (preferably between 250mm and 1000mm) and an inner bore diameter D3.
• the primary coil structure 310 includes six primary coils in the magnet 03 distributed symmetrically over the centre of DSV, four of which 312a, 312b, 313a, 313b have the same polarity. The remaining two coils 314a and 314b have an opposite polarity to the other four which produces disk-type DSV 311 having an axial diameter Dz3 less than the radial diameter Dr3, where the ratio of the axial diameter Dz3 to the radial diameter Dr3 of the DSV 21 1 is equal to or less than 0.75.
• the respective end coils 312a, 312b, located at opposite ends of the bore 301 are spaced apart by distance C3, which is preferably between 300mm and 1000mm. It will be appreciated that the distance between the end coils 312a, 312b is measured from a centre of one coil to a centre of the other coil.
• Figures 3A and 3B illustrate the magnet configuration and DSV, and the 5 Gauss lines of the 3.0 Tesla superconducting magnet 03a, which is substantially the same as superconducting magnet 03 having the same primary coil structure 310.
• the magnet 03 (and thus magnet 03a) employs six primary coils 312a, 312b, 313a, 313b, 314a, 314b in which four primary coils 312a, 312b, 313a, 313b have positive polarity and two primary coils 314a, 314b have negative polarity.
• magnet 03a includes two shield coils 315a, 315b also having a negative polarity.
• the bore 305a of the magnet 03a has a cold bore length of 800mm and an inner bore diameter of 760mm.
• this configuration provides a 5 Gauss line that is 4.6 meters in the radial direction and 7.9 meters in the axial direction, which again demonstrates that the stray fields are well controlled.
• Figure 4 shows the primary coil structure 410 of the fourth preferred embodiment of the present invention in form of a superconducting magnet 04.
• the magnet 04 has a cold bore length L4 (preferably between 250mm and 1000mm) and an inner bore diameter D4.
• the magnet 04 includes a primary coil structure 410 having seven primary coils: a first end coil 412a, a second end coil 412b, and five intermediate coils 413a, 413b, 414, 415a and 415b.
• the primary coil structure 410 is arranged symmetrically with respect to the centre of DSV 41 1 , with coil 414 spanning over the centre of the DSV 41 1.
• the first end coil 412a and the second end coil 412b are located at opposing ends of the bore 401 and are spaced apart by distance C4, which is preferably between 300mm and 1000mm.
• Coils 412a, 412b, 413a, 413b, 414 have the same polarity, while coils 415a and 415b have the opposite polarity, which, in combination, produces the disk-type DSV 411 having an axial diameter Dz4 less than the radial diameter Dr4, where the ratio of the axial diameter Dz4 to the radial diameter Dr4 of the DSV 41 1 is equal to or less than 0.75.
• Figures 4A and 4B illustrate the magnet configuration and DSV, and the 5 Gauss lines of the 1.5 Tesla superconducting magnet 04a, which is substantially the same as superconducting magnet 04 having the same primary coil structure 410.
• the magnet 04a employs seven primary coils 412a, 412b, 413a, 413b, 414, 415a, 415b (as in magnet 04) and two shield coils 416a, 416b, in which five primary coils 412a, 412b, 413a, 413b, 414 have positive polarity and two primary coils 415a, 415b as well as the two shield coils 416a, 416b have negative polarity.
• the magnet 04a has a cold bore length of 899mm and an inner bore diameter of 900mm.
• Figure 5 shows the primary coil structure 510 of superconducting magnet 05.
• the magnet 05 includes a frustoconical bore 501 extending axially through the magnet 05 about longitudinal axis 502.
• the magnet 05 includes a primary coil structure 510 having five primary coils: a first end coil 512, a second end coil 513, and three intermediate coils 514, 515 and 516 located between the first end coil 512 and the second end coil 513. All of the coils 512-516 are of the same polarity, having an asymmetric, frustoconical arrangement.
• the bore 501 of the magnet 05 has a cold bore length L5 (preferably between 250mm and 1000mm).
• the first end coil 512a and the second end coil 513 are spaced apart by distance C5, which is preferably between 300mm and 1000mm.
• the end coil 512 has the largest inner diameter D5a about cylindrical portion 503, while another end coil 513 has the smallest inner diameter D5b about cylindrical portion 505 which is significantly smaller than that of end coil 512.
• the coil 516 has a similar inner diameter as that of end coil 512, allowing access for human shoulders.
• coils 513, 514 are located about the angled portion/frustoconical portion 504 of the bore 502 having an intermediate inner diameter D5c which is between largest inner diameter D5a and smaller inner diameter D5b.
• the magnet 05 produces disk-type DSV 51 1 having an axial diameter DZ5 less than the radial diameter Drs, where the ratio of the axial diameter Dzs to the radial diameter Drs of the DSV 51 1 is equal to or less than 0.75.
• Figure 6 shows the primary coil structure of the sixth preferred embodiment of the present invention, in which the superconducting magnet 06 has a primary coil structure 610.
• the magnet 06 includes a narrowing bore 601 extending axially through the magnet 06 about longitudinal axis 602.
• the primary coil structure 610 of magnet 06 comprises six primary coils: a first end coil 612, a second end coil 613, and four intermediate coils 614, 615, 616 and 617 located therebetween.
• the primary coils 612-617 have an asymmetric, frustoconical arrangement. Five coils 612, 613, 614, 615, 616 have the same polarity, while one primary coil 617 adjacent the end coil 612 has an opposing, negative polarity. In some alternative embodiments, the primary coil 617 can be located adjacent end coil 613.
• the first end coil 612 and the second end coil 613 are located at opposing ends of the bore 601 and spaced apart by distance C6, which is preferably between 300mm and 1000mm. It will be appreciated that the distance between the end coils 612, 613 is measured from a centre of one coil to a centre of the other coil.
• the end coil 612 is located about the widest portion (first cylindrical portion 603) of the bore 601 having the largest inner diameter D6a, while the opposing end coil 613 is located about a narrower portion (second cylindrical portion 605) of the bore 601 having the smallest inner diameter D6b which is significantly smaller than that of the largest inner diameter D6a.
• coils 614, 615 are located about the angled portion/frustoconical portion 604 of the bore 601 having an intermediate diameter D6c which is between largest inner diameter D6a and smaller inner diameter D6b.
• the coil 616 has a similar inner diameter to that of end coil 612, allowing access for human shoulders.
• the magnet 06 produces disk-type DSV 61 1 having an axial diameter DZ6 less than the radial diameter Dr6, where the ratio of the axial diameter Dz6 to the radial diameter Dr6 of the DSV 61 1 is equal to or less than 0.75.
• the magnet 06 has a cold bore length of L6 (preferably between 250mm and 1000mm), a largest inner bore diameter D6a and a smaller inner bore diameter D6b.
• Figures 6A and 6B illustrate the magnet configuration and DSV, and the 5 Gauss lines of the 1.5 Tesla superconducting magnet 06a, which is substantially the same as superconducting magnet 06 having the same primary coil structure 610. Similar to magnet 06, magnet 06a employs six primary coils 612, 613, 614, 615, 616, 617. Magnet 06a also includes one shield coil 618. Five primary coils 612, 613, 614, 615, 616 have a positive polarity and one primary coil 617 as well as the shield coil 618 have an opposing, negative polarity.
• the DSV 61 1 a produced by magnet 06a has dimensions of 300mm (X- ) x 300mm (Y-) x 100mm(Z-) producing a D ⁇ 6/Dr6 ratio of 0.333 (100mm/300mm) and peak to peak homogeneity of 10ppm.
• the magnet 06a has a cold bore length of 600 mm, a largest inner bore diameter of 820mm and a smallest inner bore diameter of 282mm.
• Figure 7 shows another magnet 07 having a primary coil structure 710.
• the magnet 07 includes a stepped bore 701 extending axially through the magnet 07 about longitudinal axis 702.
• the magnet 07 has a cold bore length of L7 (preferably between 250mm and 1000mm).
• the primary coil structure 710 has five primary coils 712, 713, 714, 715 and 716 of the same polarity, having an asymmetric, three-stepped arrangement.
• the end coil 712 is located about the widest portion (first step 703) of the bore 701 having the largest inner diameter D7a, while opposing end coil 713 is located about the narrowest portion (third step 705) of the bore 701 having the smallest inner diameter D7b which is significantly smaller than that largest inner diameter D7a.
• the first end coil 712 and the second end coil 713 are spaced apart by distance C7, which is preferably between 300mm and 1000mm.
• the coil 716 has a similar inner diameter as that of end coil 712, allowing access for human shoulders.
• the two medium sized coils 714 and 715, adjacent the end coil 713, located about the second step (second step 704) of the bore 701 have the same or similar inner diameter D7c that is significantly smaller than the largest inner diameter D7a and significantly larger than the smallest inner diameter D7b.
• the magnet 07 produces disk-type DSV 71 1 having an axial diameter DZ7 less than the radial diameter Drz, where the ratio of the axial diameter Dzz to the radial diameter Drz of the DSV 71 1 is equal to or less than 0.75.
• Figure 8 illustrates a primary coil structure 810 of magnet 08.
• the magnet 08 includes a stepped bore 801 extending axially through the magnet 08 about longitudinal axis 802.
• the magnet 08 has a cold bore length of L8 (preferably between 250mm and 1000mm).
• the primary coil structure 810 comprises six primary coils 812, 813, 814, 815, 816 and 817, having an asymmetric, three-stepped, configuration.
• Five coils 812, 813, 814, 815, 816 have the same polarity, while one primary coil 817 adjacent the end coil 812 has an opposing polarity.
• primary coil 817 can be located adjacent opposing end coil 813.
• the end coil 812 is located about the widest portion (first step 803) of the bore 801 having the largest inner diameter D8a, while opposing end coil 813 is located about the narrowest portion (third step 805) of the bore 801 having the smallest inner diameter D8b which is significantly smaller than the largest inner diameter D8a.
• the coil 816 has a similar inner diameter to that of end coil 812, allowing access for human shoulders.
• the first end coil 812 and the second end coil 81 are spaced apart by distance C8, which is preferably between 300mm and 1000mm.
• the two medium sized coils 814 and 815, adjacent the end coil 813, located about the second step 804 of the bore 801 have the same or similar inner diameter D8c that is significantly smaller than the largest inner diameter D8a and significantly larger than the smallest inner diameter D8b.
• the magnet 08 produces disk-type DSV 81 1 having an axial diameter DZ8 less than the radial diameter Drs, where the ratio of the axial diameter Dzs to the radial diameter Drs of the DSV 81 1 is equal to or less than 0.75.
• Figure 9 shows a magnet 09 according to another embodiment of the present invention.
• the magnet 09 includes a stepped bore 901 extending axially through the magnet 09 about longitudinal axis 902. [0142] In the illustrated embodiment, the magnet 09 has a cold bore length of L9 (preferably between 250mm and 1000mm).
• the magnet 09 comprises a primary coil structure 910 having five primary coils 912, 913, 914, 915 and 916 of the same polarity, having an asymmetric, two-step arrangement. It will be appreciated that the magnet could have more primary coils.
• the inner diameter of coils 912, 916 at the first step 903 is significantly larger than that of coils 913, 914, 915 at the second step 904.
• the first end coil 912 and the second end coil 913 are spaced apart by distance C9, which is preferably between 300mm and 1000mm.
• the magnet 09 produces disk-type DSV 91 1 having an axial diameter DZ9 less than the radial diameter Drg, where the ratio of the axial diameter Dzg to the radial diameter Drg of the DSV 91 1 is equal to or less than 0.75.
• Figure 10 shows a magnet 010 which includes a primary coil structure 1010.
• the magnet 010 includes a stepped bore 1001 extending axially through the magnet 010 about longitudinal axis 1002.
• the magnet 010 has a cold bore length of L10 (preferably between 250mm and 1000mm).
• the magnet 010 comprises seven primary coils 1012, 1013, 1014, 1015, 1016, 1017 and 1018, having an asymmetric, two-step arrangement.
• Five coils 1012, 1013, 1014, 1015, 1016 of the primary coil structure 1010 have the same polarity, while the two primary coils 1017, 1018, adjacent the two end coils 1012 and 1013 respectively, have opposite polarities to the five coils 1012, 1013, 1014, 1015, 1016.
• the first end coil 1012 and the second end coil 1013 are spaced apart by distance C10, which is preferably between 300mm and 1000mm.
• the inner diameter of coils 1012, 1016, 1017 at the first step 1003 is significantly larger than that of coils 1013, 1014, 1015, 1018 at the second step 1004.
• the magnet 010 produces disk-type DSV 1011 having an axial diameter Dzio less than the radial diameter Dno, where the ratio of the axial diameter Dzio to the radial diameter Dno of the DSV 1011 is equal to or less than 0.75.
• Figure 11 shows a magnet 011 which includes a primary coil structure 1110.
• the magnet 011 includes a two-stepped bore 1101 extending axially through the magnet 011 about longitudinal axis 1102.
• the magnet 011 has a cold bore length of L11 (preferably between 250mm and 1000mm).
• the magnet 011 comprises seven primary coils 1112, 1113, 1114, 1115, 1116, 1117 and 1118, having an asymmetric, two-step arrangement.
• the first end coil 1112 and the opposing second end coil 1113 are spaced apart by distance C11, which is preferably between 300mm and 1000mm.
• the inner diameter of coils 1112, 1114, 1115, 1116, 1117 and 1118 about the first step 1103 is significantly larger than that of coil 1113 about the second step 1104.
• the magnet 011 produces disk-type DSV 1111 having an axial diameter Dzu less than the radial diameter Dm, where the ratio of the axial diameter Dzu to the radial diameter Dm of the DSV 1111 is equal to or less than 0.75.
• an inner diameter of the bore is preferably between 200mm and 1 100mm.
• the bore is frustoconical having a largest inner diameter of 820mm and a smallest inner diameter of 282mm.
• the DSV in an MRI magnet is designed to encompass the whole organ under study (for example, the whole brain).
• Typical DSV sizes are 45-50cm in diameter having spherical or slightly ellipsoidal volumes such that the patient remains positioned at the centre of the DSV while the organ is scanned/imaged. Thus the patient is not moved during the examination of that organ.
• the size of the DSV in the axial direction has a strong influence on the length of the magnet. It is a feature of this invention that the DSV has a much smaller axial extent than its radial extent. This allows the magnet to be shorter than conventional systems and, in use, whole organ imaging is achieved by moving the patient in an automatically controlled fashion through the magnet system.
• the patient can be moved through the imaging region/DSV, thereby allowing the length of the imaging device to be significantly shorter as the organ of the patient that is to be imaged does not need to be centred within the DSV.
• the magnet is capable of producing a magnetic field of at least 1.0 Tesla, and preferably at least 3.0 Tesla, which is substantially homogeneous over a predetermined imaging region or volume (also called the‘homogeneous region’ or‘DSV’).
• the imaging region has an external surface defined by a computed variation of the longitudinal magnetic field relative to the longitudinal magnetic field at the imaging centre of less than 20 parts per million peak-to-peak.
• the terms ‘comprises’, ‘comprising’, ‘includes’, ‘including’, or similar terms are intended to mean a non-exclusive inclusion, such that a method, system or apparatus that comprises a list of elements does not include those elements solely, but may well include other elements not listed.

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