US20110011102A1 - Cryogenically cooled superconductor rf head coil array and head-only magnetic resonance imaging (mri) system using same - Google Patents

Cryogenically cooled superconductor rf head coil array and head-only magnetic resonance imaging (mri) system using same Download PDF

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US20110011102A1
US20110011102A1 US12/764,044 US76404410A US2011011102A1 US 20110011102 A1 US20110011102 A1 US 20110011102A1 US 76404410 A US76404410 A US 76404410A US 2011011102 A1 US2011011102 A1 US 2011011102A1
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coils
head
radiofrequency
coil
module according
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Erzhen Gao
Qiyuan Ma
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TIME MEDICAL HOLDINGS Co Ltd
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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
    • 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
    • G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
    • G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
    • G01R33/34015—Temperature-controlled RF coils
    • G01R33/3403—Means for cooling of the RF coils, e.g. a refrigerator or a cooling vessel specially adapted for housing an RF coil
    • 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
    • G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
    • G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
    • G01R33/341—Constructional details, e.g. resonators, specially adapted to MR comprising surface coils
    • G01R33/3415—Constructional details, e.g. resonators, specially adapted to MR comprising surface coils comprising arrays of sub-coils, i.e. phased-array coils with flexible receiver channels
    • 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
    • G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
    • G01R33/36—Electrical details, e.g. matching or coupling of the coil to the receiver
    • G01R33/3628—Tuning/matching of the transmit/receive coil
    • G01R33/3635—Multi-frequency operation
    • 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
    • G01R33/3806—Open magnet assemblies for improved access to the sample, e.g. C-type or U-type magnets
    • 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
    • G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
    • G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
    • G01R33/34015—Temperature-controlled RF coils
    • G01R33/34023—Superconducting RF coils

Definitions

  • the present invention relates generally to magnetic resonance imaging and spectroscopy, and, more particularly, to magnetic resonance imaging and spectroscopy apparatus employing superconductor components, and to methods for manufacturing such apparatus.
  • Magnetic Resonance Imaging (MRI) technology is commonly used today in larger medical institutions worldwide, and has led to significant and unique benefits in the practice of medicine. While MRI has been developed as a well-established diagnostic tool for imaging structure and anatomy, it has also been developed for imaging functional activities and other biophysical and biochemical characteristics or processes (e.g., blood flow, metabolites/metabolism, diffusion), some of these magnetic resonance (MR) imaging techniques being known as functional MRI, spectroscopic MRI or Magnetic Resonance Spectroscopic Imaging (MRSI), diffusion weighted imaging (DWI), and diffusion tensor imaging (DTI). These magnetic resonance imaging techniques have broad clinical and research applications in addition to their medical diagnostic value for identifying and assessing pathology and determining the state of health of the tissue examined.
  • MR magnetic resonance
  • a patient's body (or a sample object) is placed within the examination region and is supported by a patient support in an MRI scanner where a substantially constant and uniform primary (main) magnetic field is provided by a primary (main) magnet.
  • the magnetic field aligns the nuclear magnetization of precessing atoms such as hydrogen (protons) in the body.
  • a gradient coil assembly within the magnet creates a small variation of the magnetic field in a given location, thus providing resonance frequency encoding in the imaging region.
  • a radio frequency (RF) coil is selectively driven under computer control according to a pulse sequence to generate in the patient a temporary oscillating transverse magnetization signal that is detected by the RF coil and that, by computer processing, may be mapped to spatially localized regions of the patient, thus providing an image of the region-of-interest under examination.
  • RF radio frequency
  • the static main magnetic field is typically produced by a solenoid magnet apparatus, and a patient platform is disposed in the cylindrical space bounded by the solenoid windings (i.e. the main magnet bore).
  • the windings of the main field are typically implemented as a low temperature superconductor (LTS) material, and are super-cooled with liquid helium in order to reduce resistance, and, therefore, to minimize the amount of heat generated and the amount of power necessary to create and maintain the main field.
  • LTS low temperature superconductor
  • the majority of existing LTS superconducting MRI magnets are made of a niobium-titanium (NbTi) and/or Nb 3 Sn material which is cooled with a cryostat to a temperature of 4.2 K.
  • the magnetic field gradient coils generally are configured to selectively provide linear magnetic field gradients along each of three principal Cartesian axes in space (one of these axes being the direction of the main magnetic field), so that the magnitude of the magnetic field varies with location inside the examination region, and characteristics of the magnetic resonance signals from different locations within the region of interest, such as the frequency and phase of the signals, are encoded according to position within the region (thus providing for spatial localization).
  • the gradient fields are created by current passing through coiled saddle or solenoid windings, which are affixed to cylinders concentric with and fitted within a larger cylinder containing the windings of the main magnetic field.
  • the coils used to create the gradient fields typically are common room temperature copper windings.
  • the gradient strength and field linearity are of fundamental importance both to the accuracy of the details of the image produced and to the information on tissue chemistry (e.g., in MRSI).
  • imaging (acquisition) speed is desired to minimize imaging blurring caused by temporal variations in the imaged region during image acquisition, such as variations due to patient movement, natural anatomical and/or functional movements (e.g., heart beat, respiration, blood flow), and/or natural biochemical variations (e.g., caused by metabolism during MRSI).
  • the pulse sequence for acquiring data encodes spectral information in addition to spatial information
  • minimizing the time required for acquiring sufficient spectral and spatial information to provide desired spectral resolution and spatial localization is particularly important for improving the clinical practicality and utility of spectroscopic MRI.
  • SNR signal-to-noise ratio
  • Increasing SNR by increasing the signal before the preamplifier of the MRI system is important in terms of increasing the quality of the image.
  • One way to improve SNR is to increase the magnetic field strength of the magnet as the SNR is proportional to the magnitude of the magnetic field. In clinical applications, however, MRI has a ceiling on the field strength of the magnet (the US FDA's current ceiling is 3 T (Tesla)).
  • Other ways of improving the SNR involve, where possible, reducing sample noise by reducing the field-of-view (where possible), decreasing the distance between the sample and the RF coils, and/or reducing RF coil noise.
  • Various embodiments of the present invention provide a cryogenically cooled superconducting RF head-coil array which may be used in whole-body MRI scanners and/or in dedicated, head-only MRI systems (also referred to herein as “head-dedicated MRI systems,” “head-only MRI systems,” or the like).
  • Some embodiments of the invention provide a head-dedicated MRI system and, more particularly, various embodiments provide a superconducting main magnet for a head-dedicated MRI system which, in some embodiments, further comprises a cryogenically-cooled superconducting RF head-coil array according to embodiments of the present invention.
  • the port may be coupled to a cryocooler that is thermally coupled to the at least one thermal sink member.
  • each radiofrequency coil is in direct thermal contact with a respective one of the thermal sink members that are each in direct thermal contact with another of the thermal sink members that is in thermal contact with the cryocooler.
  • the radiofrequency coils may comprise at least eight radiofrequency coils that are azimuthally displaced about a common longitudinal axis at a substantially common displacement along the longitudinal axis, and are configured for imaging a region surrounded by the radiofrequency coils.
  • Each of the radiofrequency coils may be configured to receive and not transmit radiofrequency signals.
  • the vacuum thermal isolation housing and radiofrequency coils may be dimensioned and configured for head imaging and not whole body imaging.
  • the radiofrequency coil array module is dimensioned and configured for use in a head-only magnetic resonance imaging system that comprises a main electromagnet system comprising: a first and second set of high temperature superconductor coils which are configured to be coaxial relative to a common longitudinal axis; wherein the first coil set includes at least two coils having an inner radius and disposed in a first region of a length along the common axis to cover a head and neck of a human body, and the second coil set includes at least one coil having an inner radius and disposed in a second region of a length along the common axis to cover a portion of a human torso; and wherein the first and second coils are configured to provide a uniform magnetic field in the first region to provide for imaging a region of interest of the individual's head when positioned within the first region.
  • FIGS. 1A and 1B schematically depict orthogonal views of an illustrative cryogenically cooled superconducting RF head coil array, in accordance with some embodiments of the present invention
  • FIG. 2 schematically illustrates wall(s) of the vacuum chamber depicted in FIG. 1A being implemented as a double-walled glass Dewar, in accordance with some embodiments of the present invention
  • FIG. 3 schematically depicts an illustrative cross-sectional view along the longitudinal axis of a superconductor RF head coil array corresponding to embodiments depicted in FIGS. 1A and 1B with the vacuum chamber comprising a Dewar 1 according to various embodiments represented by FIG. 2 , in accordance with some embodiments of the present invention;
  • FIGS. 4A and 4B depict an illustrative alternative implementation of a superconductor RF head coil array (module), in accordance with some embodiments of the present invention
  • FIG. 5 schematically depicts a cross section of an illustrative MRI system, in accordance with some embodiments of the present invention
  • FIG. 6 schematically depicts an illustrative RF head coil array that includes thermal radiation screening, in accordance with some embodiments of the present invention
  • FIG. 7 schematically depicts a cross-sectional view of a superconducting main magnet of a head-only MRI system, in accordance with some embodiments of the present invention
  • FIG. 8 depicts with reference to the z-r plane a coil configuration of a superconducting main magnet system, in accordance with some embodiments of the present invention.
  • FIG. 9 depicts a normalized current distribution for the main magnet coil arrangement corresponding to the illustrative embodiment of FIGS. 7 and 8 , in accordance with some embodiments of the present invention.
  • FIG. 10 is an illustrative coil pattern (depicted in the z-r plane, with units normalized to meters) of a 3 T head magnetic resonance imaging scanner, in accordance with various embodiments of the present invention.
  • FIG. 11 is a plot showing the magnetic field distribution for the illustrative embodiment depicted in FIG. 10 , in accordance with some embodiments of the present invention.
  • FIG. 12 shows the fringe fields of one Gauss (1 G), three Gauss (3 G) and five Gauss (5 G) lines for the field distribution of FIG. 11 , in accordance with an illustrative embodiment of the present invention.
  • cryogenically cooled superconducting RF head-coil array which may be used in whole-body MRI scanners and/or in dedicated, head-only MRI systems (also referred to herein as “head-dedicated MRI systems,” “head-only MRI systems,” or the like) and (ii) various embodiments of a head-dedicated MRI system and, more particularly, various embodiments of a superconducting main magnet for a head-dedicated MRI system which, in some embodiments, further comprises a cryogenically-cooled superconducting RF head-coil array according to embodiments of the present invention.
  • a cryogenically-cooled superconducting RF head-coil array coil may be implemented in myriad magnetic resonance imaging and spectroscopy systems, such as systems employing conventional copper gradient coils, systems employing superconducting gradient coils (e.g., such as disclosed in U.S. patent application Ser. No. 12/416,606, filed April 1, 2009, and in Provisional Application No. 61/170,135, filed Apr. 17, 2009, each of which is hereby incorporated by reference in its entirety), whole body systems, dedicated head-only systems, systems with a vertically or horizontally oriented main magnetic field, open or closed systems, etc.
  • myriad magnetic resonance imaging and spectroscopy systems such as systems employing conventional copper gradient coils, systems employing superconducting gradient coils (e.g., such as disclosed in U.S. patent application Ser. No. 12/416,606, filed April 1, 2009, and in Provisional Application No. 61/170,135, filed Apr. 17, 2009, each of which is hereby incorporated by reference in its entirety), whole body
  • a head-dedicated MRI system employing a superconducting main magnet may be implemented in myriad magnetic resonance imaging and spectroscopy systems, such as systems employing conventional copper gradient coils, systems employing superconducting gradient coils (e.g., such as disclosed in U.S. patent application Ser. No. 12/416,606, filed Apr. 1, 2009, and in Provisional Application No. 61/170,135, filed Apr.
  • MRI magnetic resonance
  • DTI diffusion tensor imaging
  • FIGS. 1A and 1B schematically depict orthogonal views of an illustrative cryogenically cooled superconducting RF head coil array 10 , in accordance with some embodiments of the present invention.
  • orthogonal x, y, z coordinates are depicted as a reference frame.
  • FIG. 1A is a cross-sectional view in the x-y plane indicated by reference IA-IA′ in FIG.
  • FIG. 1B illustrates a configuration of eight superconducting RF coils 3 a - 3 h (also referred to herein collectively as superconductor RF coils 3 or RF coil array 3 ) each disposed in thermal contact with a respective one of eight thermal conductors 5 a - 5 h (e.g., non-metallic high thermal conductivity materials, such as high thermal conductivity ceramic, such as sapphire or alumina), with the RF coils 3 a - 3 h and thermal conductors 5 a - 5 h being disposed within a sealed vacuum chamber having vacuum chamber wall(s) 2 .
  • thermal conductors 5 a - 3 h e.g., non-metallic high thermal conductivity materials, such as high thermal conductivity ceramic, such as sapphire or alumina
  • FIG. 1B is a side view along the longitudinal axis (i.e., z axis) viewed from the direction indicated by reference IB in FIG. 1A , and illustrates components comprising the cooling system of superconducting RF head coil array 10 , the cooling system including thermal conductor 15 (e.g., non-metallic high thermal conductivity materials, such as high thermal conductivity ceramic, such as sapphire or alumina) in thermal contact with each of thermal conductors 5 a - 5 h, cold head 9 in thermal contact with thermal conductor (sink) 15 , and cryocooler 7 configured for maintaining the cold head 9 at a desired cryogenic temperature.
  • thermal conductor 15 e.g., non-metallic high thermal conductivity materials, such as high thermal conductivity ceramic, such as sapphire or alumina
  • FIG. 1B does not show (i) the vacuum chamber comprising vacuum chamber wall(s) 2 , (ii) coils 3 b and 3 d, and (iii) thermal conductors 5 b and 5 d (as will be further understood from the ensuing description (e.g., in connection with FIG. 3 ), FIG. 1B also does not show a vacuum chamber portion into which cryocooler 7 is mounted).
  • coils 3 a - 3 h are in vacuum and cooled by the thermal conductors 5 a - 5 b, which conduct heat away from the coils to the thermal conductor/sink 15 , which is thermally coupled with a cryogenic cooler 7 .
  • a cryogenic cooler 7 As will be understood by those skilled in the art, in some embodiments (e.g., low main magnetic field implementations, such as less than 3 T, or less than 1.5 T, etc.) small amounts of metal, such as copper, may be used for thermal conductor/sink 15 and/or possibly thermal conductors 5 a - 5 h.
  • thermal conductors 5 a - 5 h may be integrally formed with thermal conductor/sink 15 , whereas in some embodiments, one or more of thermal conductors 5 a - 5 h are distinct members that are mechanically joined (e.g., using epoxy, etc.) to thermal conductor/sink 15 to provide a good thermal conduction therebetween.
  • the coils 3 a - 3 h may be cooled to a temperature in the range of about 4K to 100K, and more particularly, to a temperature below the critical temperature of the superconducting material (e.g., in some embodiments, below the critical temperature of a high temperature superconductor (HTS) material used for the RF coils 3 a - 3 h ).
  • a temperature below the critical temperature of the superconducting material e.g., in some embodiments, below the critical temperature of a high temperature superconductor (HTS) material used for the RF coils 3 a - 3 h .
  • HTS high temperature superconductor
  • each of RF coil elements 3 a - 3 h is implemented as a high temperature superconductor (HTS), such as YBCO and/or BSCCO, etc. (e.g., using an HTS thin film or HTS tape), though a low temperature superconductor (LTS) may be used in various embodiments.
  • HTS high temperature superconductor
  • LTS low temperature superconductor
  • each of RF coil elements 3 a - 3 h is an HTS thin film spiral coil and/or an HTS thin film spiral-interdigitated coil on a substrate such as sapphire or lanthanum aluminate.
  • superconducting RF head coil array 10 is implemented as an HTS thin film RF head coil array.
  • vacuum chamber comprising wall(s) 2 may comprise a double-walled Dewar 1 made of glass and/or other non-conductive, mechanically strong material(s), such as G10, RF4, plastic, and/or ceramic. More specifically, FIG. 2 schematically illustrates wall(s) 2 of the vacuum chamber depicted in FIG. 1A being implemented as a double-walled glass Dewar 1 , in accordance with some embodiments of the present invention. It will be understood that the dimensions and shape of a cryogenically cooled superconducting RF head-coil array module may be modified according to various implementations of the present invention. In accordance with some implementations, FIG.
  • cylinder 60 has an inner diameter, outer diameter, and axial length of 230 mm, 236 mm, and 254 mm, respectively;
  • cylinder 62 has an inner diameter, outer diameter, and axial length of 246 mm, 252 mm, and 254 mm, respectively;
  • cylinder 64 has an inner diameter, outer diameter, and axial length of 280 mm, 286 mm, and 312 mm, respectively;
  • cylinder 66 has an inner diameter, outer diameter, and axial length of 296 mm, 302 mm, and 330 mm, respectively;
  • inner bottom plate (circular/cylindrical) 74 has a diameter of 236 mm and a thickness of 12.7 mm;
  • outer bottom plate (circular/cylindrical) 74 has a diameter of 236 mm and a thickness of 12.7 mm;
  • a plug 70 seals off a standard vacuum port in ring 68 through which the intra-dewar cavity is evacuated.
  • double-walled Dewar 1 may be constructed, in a variety of ways, as a continuous, hermetically sealed glass housing enclosing an interior chamber (or cavity) 4 in which at least a low vacuum condition and, in accordance with some embodiments, preferably at least a high vacuum condition (e.g., about 10 ⁇ 6 Torr or lower pressure) is maintained.
  • a low vacuum condition e.g., about 10 ⁇ 6 Torr or lower pressure
  • double-walled Dewar 1 may be manufactured as follows: (i) forming two generally cylindrical (e.g., but hexagonal in cross-section transverse to the longitudinal/cylindrical access) double-walled structures each having a generally U-shaped wall cross-section, the first corresponding to continuous glass wall portion 1 a (comprising cylinders 60 and 66 , ring 68 and plate 74 ) and the second corresponding to continuous wall portion 1 b (comprising cylinders 62 and 64 , ring 66 , and plate 76 ), (ii) fitting the generally cylindrical continuous glass wall portion 1 b into the annular space of generally cylindrical continuous glass wall portion 1 a, possibly using glass spacers therebetween (e.g., identified in FIG.
  • double-walled Dewar 1 may be implemented in accordance with, or similar to, the hermetically sealed double-walled structures (and vacuum thermal isolation housing) described in U.S. application Ser. No. 12/212,122, filed Sep. 17, 2008, and in U.S. application Ser. No. 12/212,147, filed Sep. 17, 2008, each of which is herein incorporated by reference in its entirety.
  • FIG. 3 schematically depicts an illustrative cross-sectional view along the longitudinal axis of a superconductor (e.g., HTS) RF head coil array corresponding to embodiments depicted in FIGS. 1A and 1B with the vacuum chamber comprising a Dewar 1 according to various embodiments represented by FIG. 2 .
  • Dewar 1 is sealably joined to a double-walled stainless steel chamber 8 that includes a flange to which cryocooler 7 is sealably mounted.
  • double-walled stainless steel chamber 8 is hermetically sealed, enclosing an interior chamber (or cavity) 12 in which at least a low vacuum condition and, in accordance with some embodiments, preferably at least a high vacuum condition (e.g., about 10 ⁇ 6 Torr or lower pressure) is maintained.
  • a low vacuum condition e.g., about 10 ⁇ 6 Torr or lower pressure
  • the joint between the hermetically sealed double-walled Dewar 1 (e.g., glass) and the stainless steel chamber may be formed by epoxy bonding, welding, or other hermetically sealed flange connection, providing a sufficient seal to maintain at least a low vacuum condition (e.g., about 10 ⁇ 2 to about 10 ⁇ 5 Torr) in the interior chamber portion 6 that houses the superconducting RF coils 3 and thermal conductors 5 (i.e., 5 a - 5 h ) and 15 .
  • a low vacuum condition e.g., about 10 ⁇ 2 to about 10 ⁇ 5 Torr
  • cryocooler 7 and the flange of stainless steel chamber 8 may be provided by an O-ring or other sealing mechanism (e.g., metal gasket/knife-edge connection) to, similarly, maintain the at least low vacuum condition in the interior chamber portion 6 that houses the RF coils 3 and thermal conductors 5 and 15 .
  • chamber 8 may be made of materials other than stainless steel, e.g., aluminum or other metallic or other non-metallic material, such as glass, ceramic, plastics, or combination of these materials, and such other materials may be appropriately joined to Dewar 1 and cryocooler 7 .
  • cryocooler 7 may be implemented as any of various single stage or multi-stage cryocoolers, such as, for example, a Gifford McMahon (GM) cryocooler, a pulse tube (PT) cooler, a Joule-Thomson (JT) cooler, a Stirling cooler, or other cryocooler.
  • GM Gifford McMahon
  • PT pulse tube
  • JT Joule-Thomson
  • the superconductor RF head coil array 10 may be configured for cooling such that coils 3 are cooled by a cryogen, such as liquid helium and liquid nitrogen.
  • a cryogenically cooled superconductor RF coil array (e.g., array 10 ) in accordance with various embodiments of the present invention includes at least one electrical feedthrough (e.g., through chamber 8 ) to provide for coupling electrical signals into and/or out of the array (e.g., for the RF coils, for controlling and/or monitoring any sensors (e.g., pressure and/or temperature, etc.) that may be provided in the module).
  • receiver and/or, if applicable, transmitter circuitry e.g., amplifiers and/or filters and/or appropriate matching and/or decoupling circuitry
  • transmitter circuitry e.g., amplifiers and/or filters and/or appropriate matching and/or decoupling circuitry
  • each of the RF coils may be provided within the vacuum chamber; for example, it may be disposed on and in thermal contact with thermal conductors 5 a - 5 h, wherein such cooling may provide for improving noise properties and/or for using superconducting components for at least a portion of such circuitry.
  • superconducting RF head coil array 10 is implemented as a receive-only array, with an RF transmitter being implemented as a separate RF coil (not shown), which in various embodiments may be a conventional (e.g., non-superconducting, such as a conventional copper RF coil) RF transmitter coil or a superconducting RF transmitting coil.
  • a separate transmitter coil may be configured external to the vacuum chamber comprising wall(s) 2 (e.g., external to Dewar 1 ) or, in some embodiments, within the vacuum chamber comprising wall(s) 2 (e.g., within Dewar 1 ).
  • an RF transmission coil is implemented as one or more superconducting RF transmission coils (e.g., a high temperature superconductor (HTS) RF transmitter) that are separate from the RF receiver coils
  • one or more superconducting RF transmission coils may be disposed in thermal contact with one or more of thermal conductors 5 a - 5 h.
  • superconducting RF head coil array 10 may be implemented as a transmit and receive coil array (a transceiver array), with each of one or more of the superconducting RF coils 3 a - 3 h being used for both transmission and reception of RF signals.
  • one or more of the superconducting RF coil elements 3 a - 3 h may be implemented as a multiple resonance RF coil element (e.g., comprising two or more receiving coils having different resonant frequencies, such as for detecting sodium and hydrogen resonances at a given magnetic field (e.g., at 3 Tesla (T)).
  • a multiple resonance RF coil element e.g., comprising two or more receiving coils having different resonant frequencies, such as for detecting sodium and hydrogen resonances at a given magnetic field (e.g., at 3 Tesla (T)).
  • two or more different ones of superconducting RF coil elements 3 a - 3 h may be designed to have different resonant frequencies; for example, RF coil elements 3 a, 3 c, 3 e, and 3 g may be tuned to a first resonant frequency (e.g., that of hydrogen nuclei at 3 T) and RF coil elements 3 b, 3 d, 3 f, and 3 h may be tuned to a second resonant frequency (e.g., that of sodium nuclei at 3 T).
  • a superconducting RF head coil array in accordance with various embodiments of the present invention may be used for acquiring magnetic resonance signals from different types of nuclei in a simultaneous or time-multiplexed manner.
  • a cryogenically-cooled superconducting RF head-coil array coil may be implemented in a magnetic resonance imaging system that employs superconducting gradient coils such as those disclosed in U.S. patent application Ser. No. 12/416,606, filed Apr. 1, 2009, and in Provisional Application No. 61/170,135, filed Apr. 17, 2009, each of which is hereby incorporated by reference in its entirety.
  • one or more of the superconducting gradient coils may be disposed within the same vacuum chamber as the superconducting RF coils (e.g., the gradient coils may be in thermal contact with the surfaces of thermal conductors 5 a - 5 h that are opposite the surfaces in contact with coils 3 a - 3 h ).
  • FIGS. 4A and 4B there is shown an illustrative alternative implementation of a superconductor RF head coil array (module), in accordance with some embodiments of the present invention. More specifically, FIG. 4A schematically depicts a cross-sectional view in a plane containing the longitudinal axis, similar to the cross-sectional view depicted with respect to the embodiment of FIG. 3 (e.g., viewing an x-z plane cross-section, using a coordinate system oriented similarly to that for the embodiment of FIGS. 1A , 1 B, 2 and 3 ), while FIG. 4B generally depicts a plan or end-on view, viewed from the left-hand side of FIG.
  • FIG. 4A schematically depicts a cross-sectional view in a plane containing the longitudinal axis, similar to the cross-sectional view depicted with respect to the embodiment of FIG. 3 (e.g., viewing an x-z plane cross-section, using a coordinate system oriented similarly to that for the embodiment of FIGS.
  • FIGS. 4A and 4B are similar to that of FIGS. 1A , 1 B, 2 and 3 , for convenience and ease of reference, identical reference numerals have been used to identify corresponding or similar elements. As may also be understood, a difference between the embodiment depicted in FIGS. 1B , 2 and 3 and the embodiment depicted in FIG.
  • a thermal conductive ring 25 (cylindrical ring) is thermally coupled to each thermal conductor 5 a - h ( 5 a and 5 e shown in FIG. 4A ) and to cryocooler 7 , which is sealably mounted (e.g., via an O-ring sealed flange 19 ) to chamber 8 .
  • a generally cylindrically shaped RF head coil array module such as depicted in the foregoing described embodiments may be well suited for use, for example, in an MRI system that employs a cylindrical, solenoid main magnet structure that generates a substantially uniform, horizontal magnetic field.
  • an MRI system is schematically depicted in FIG. 5 in longitudinal cross section, and includes cylindrical main magnet 17 having a bore in which a superconductor RF head coil array (module) 10 corresponding to that of FIGS. 4A and 4B is disposed, and which also includes gradient coil(s) 13 .
  • cryogenically cooled superconducting RF head coil array 10 may be implemented with main magnet configurations other than a cylindrical, solenoid magnet that provides horizontal fields and/or, for example, may be implemented with open magnet configurations, such as vertical magnet or a double-donut magnet.
  • main magnet 17 may be the main magnet of a whole-body scanner or may be the main magnet of a dedicated (e.g., head-only) system (e.g., such as the main magnet described hereinbelow in connection with FIGS. 7-12 ).
  • FIG. 6 schematically depicts an illustrative RF head coil array that includes thermal radiation screening, in accordance with some embodiments of the present invention. More specifically, FIG. 6 depicts the upper half of the coil depicted in FIG. 4A , further showing thermal radiation screens 17 that are used as an option to further protect the low temperature of the RF coil 3 a and the non-metallic thermal conductor 5 a from heating by the radiation from the outer wall of the double-walled glass dewar and the environment outside the dewar.
  • Thermal radiation screen 17 may be made from one or more materials, such as foam, fabricate, cotton, or other non-metallic, good thermal insulation materials or combinations thereof.
  • a superconductor RF head coil array in accordance with the hereinabove embodiments may be implemented in connection with a whole-body MRI scanner, such RF head coil arrays may alternatively be used in dedicated, head-only MRI scanners.
  • a dedicated head-only scanner may implement a superconductor main magnet in accordance with embodiments represented by, and described in connection with, the following drawings.
  • MRI scanners employing a superconductor main magnet may employ various RF coil configurations (e.g., array, non-array type, superconducting, non-superconducting, etc.), though some embodiments may employ superconducting RF head coil arrays implemented in accordance with embodiments described hereinabove.
  • RF coil configurations e.g., array, non-array type, superconducting, non-superconducting, etc.
  • FIG. 7 schematically depicts a cross-sectional view of a superconducting main magnet of a head-only MRI system, the superconducting main magnet comprising double-walled housing 41 and solenoid/helical coils 42 , with a subject illustrated disposed therein with the subject's head arranged within the diameter-sensitive volume 43 of the main magnet.
  • double-walled housing 41 encloses a hermetically sealed region 47 that is under at least a low vacuum condition, but preferably is under high vacuum (e.g., 10 ⁇ 6 to 10 ⁇ 12 Torr), and also encloses an interior chamber region 45 in which superconducting coils 42 are disposed and which is under at least a low vacuum condition (e.g., 10 ⁇ 3 to 10 ⁇ 6 Torr).
  • a hermetically sealed region 47 that is under at least a low vacuum condition, but preferably is under high vacuum (e.g., 10 ⁇ 6 to 10 ⁇ 12 Torr)
  • an interior chamber region 45 in which superconducting coils 42 are disposed and which is under at least a low vacuum condition (e.g., 10 ⁇ 3 to 10 ⁇ 6 Torr).
  • the superconducting main magnet is an electromagnet system comprising a vacuum thermal isolation housing 41 (e.g., a dewar) that is integrated with a cryogenic system (not shown) to provide for cooling superconducting coils 42 via a heat pipe (not shown) and a heat sink assembly (not shown) in thermal contact with the superconducting coils.
  • Superconducting coils may be implemented as high temperature superconductor (HTS) coils and, in some embodiments, may comprise at least one of the following superconductor materials: YBaCuO, BiSrCaCuO, TIBiCaCuO, and MgB 2 .
  • the temperature in the interior chamber region in which the coils are disposed may be in the range of about 77K-80K.
  • the coils are configured as (i) a first coil set that is disposed in a first region to cover or surround or otherwise be disposed adjacent to an individual's head, and (ii) a second coil set that is coaxial with the first coil set and is disposed in a second region to cover or surround or otherwise be disposed adjacent to the individuals shoulders or upper torso, wherein the inner radius of the first set of coils is less than the inner radius of the second set of coils, and the coils are configured to provide a uniform magnetic field in the region of the individual's head.
  • various embodiments may vary the number of coils per set, the coil radii, number of turns, longitudinal position and length, and electric current magnitude and direction in each coil to provide a desired magnetic field distribution.
  • the longitudinal position and extension, the number of turns, and electric current direction of each coil are designed to provide 1-10 ppm uniform magnetic field within the first region for head imaging.
  • the first set of coils may include at least two coils having an inner radius in a range of about 25-35 cm and disposed in a first region of a length along the common axis in a range of 40-60 cm to cover a head and neck of a human body
  • the second set of coils may include at least one coil having an inner radius in a range of about 30-40 cm and disposed in a second region of a length along the common axis in a range of 15-25 cm to cover a portion of a human torso.
  • the length of the first and second regions may, for example, range from about 20-70 cm and 10-40 cm, respectively, and the inner radii of the first and second set of coils may range from about 10-40 com and 20-50 cm, respectively.
  • Some embodiments may employ a length of the first and second regions in a range from about 10-20 cm and 20-30 cm respectively.
  • some embodiments may employ an inner radius of the first and second coils of about 10-20 cm and 20-30 cm, respectively.
  • FIG. 8 depicts with reference to the z-r plane, with dimensions in meters (m), the longitudinal extent L 2 of a first set of coils (e.g., corresponding to the four leftmost coil sets depicted in FIG. 7 ) having an inner radius of 0.28 meters, the longitudinal extent L 1 of a second coil set (e.g., corresponding to the rightmost coil set in FIG. 7 ) having an inner radius of 0.38 meters, and DSV 43 having a radius that is about 0.1 meters and offset by about 0.05 meters from the transition from the first to second set of coils (from L 2 to L 1 ) along the z-axis, in accordance with an illustrative example according to some embodiments of the present invention.
  • FIG. 9 depicts a normalized current distribution for the main magnet coil arrangement corresponding to the illustrative embodiment of FIGS. 7 and 8 . As shown, in accordance with some embodiments, at least one coil is wound to carry current in the reverse direction relative to other coils.
  • FIG. 10 is an illustrative coil pattern (depicted in the z-r plane, with units normalized to meters) of a 3 T head magnetic resonance imaging scanner, in accordance with various embodiments of the present invention. More specifically, active shield coil 51 is disposed at the outer side, main magnet coils 52 comprise eight coil sets, and a diameter-sensitive volume (DSV) 53 of homogeneous fields is about 200 mm in diameter (i.e., a radius of about 0.1 meter). The shield coil 51 may have a radius, for example, in the range of about 60-70 cm, though other radii are possible depending on the particular implementation.
  • the following table provides dimensions and current direction for coils arranged according to the embodiment of FIG.
  • the first set of coils comprise coil numbers 1 through 6
  • the second set of coils comprise coil numbers 7 and 8
  • the shielding coil is identified as coil 9
  • R 1 is the inner radius
  • R 2 is the outer radius
  • Z 1 is the first longitudinal position
  • Z 2 is the second longitudinal position
  • the current direction J is identified as positive (+) or negative ( ⁇ ):
  • FIG. 11 is a plot showing the magnetic field distribution for the illustrative embodiment depicted in FIG. 10 , with illustrative dimensions and current directions as per the foregoing table. As shown, a 3 T homogeneous field provides a 200 mm DSV.
  • FIG. 12 shows the fringe fields of one Gauss, three Gauss and five Gauss lines for the field distribution of FIG. 11 , in accordance with an illustrative embodiment of the present invention.
  • FIG. 10 illustrates a non-limiting example of an embodiment according to the present invention.
  • the outer layer is an active shield coil 51
  • the depicted inner layer comprises main magnet coils 52 having eight coil sets providing an asymmetric structure, with the coils on the right hand side (towards increasing z) having a bigger diameter for accommodating a patient's shoulders.
  • total length of the magnet is 0.86 m
  • the DSV 53 is 200 mm in diameter. According to these parameters, FIG.
  • the bore surrounding a DSV 43 of homogeneous fields is preferably not much larger in diameter than what is necessary to fit a patient's head, while the main magnet bore also includes a portion designed with a diameter having an appropriate size to accommodate the shoulder as shown in FIG. 7 .
  • a head-only main magnet in accordance with some embodiments of the present invention has a smaller DSV, so the size of superconducting magnet can be reduced, and a smaller Dewar and magnet system can be achieved, and the costs can be thus also be reduced.

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