WO2011135312A2 - Détecteur irm - Google Patents

Détecteur irm Download PDF

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Publication number
WO2011135312A2
WO2011135312A2 PCT/GB2011/000667 GB2011000667W WO2011135312A2 WO 2011135312 A2 WO2011135312 A2 WO 2011135312A2 GB 2011000667 W GB2011000667 W GB 2011000667W WO 2011135312 A2 WO2011135312 A2 WO 2011135312A2
Authority
WO
WIPO (PCT)
Prior art keywords
coil
pin diode
solenoid
sub
mri detector
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Ceased
Application number
PCT/GB2011/000667
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English (en)
Other versions
WO2011135312A3 (fr
Inventor
Nandita Maria Desouza
David John Gilderdale
Maria Angelica Schmidt
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Institute of Cancer Research Royal Cancer Hospital
Royal Marsden NHS Foundation Trust
Original Assignee
Institute of Cancer Research Royal Cancer Hospital
Royal Marsden NHS Foundation Trust
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Institute of Cancer Research Royal Cancer Hospital, Royal Marsden NHS Foundation Trust filed Critical Institute of Cancer Research Royal Cancer Hospital
Publication of WO2011135312A2 publication Critical patent/WO2011135312A2/fr
Publication of WO2011135312A3 publication Critical patent/WO2011135312A3/fr
Anticipated expiration legal-status Critical
Ceased legal-status Critical Current

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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
    • G01R33/32—Excitation or detection systems, e.g. using radio frequency signals
    • G01R33/34—Constructional details, e.g. resonators, specially adapted to MR
    • G01R33/34046—Volume type coils, e.g. bird-cage coils; Quadrature bird-cage coils; Circularly polarised coils
    • G01R33/34053—Solenoid coils; Toroidal 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
    • G01R33/32—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
    • G01R33/3642—Mutual coupling or decoupling of multiple coils, e.g. decoupling of a receive coil from a transmission coil, or intentional coupling of RF coils, e.g. for RF magnetic field amplification
    • G01R33/3657—Decoupling of multiple RF coils wherein the multiple RF coils do not have the same function in MR, e.g. decoupling of a transmission coil from a receive 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
    • G01R33/34084—Constructional details, e.g. resonators, specially adapted to MR implantable coils or coils being geometrically adaptable to the sample, e.g. flexible coils or coils comprising mutually movable parts

Definitions

  • the present invention relates to a magnetic resonance imaging (MRI) detector for detecting radio frequency (RF) signals.
  • MRI magnetic resonance imaging
  • RF radio frequency
  • MR magnetic resonance
  • SNR signal-to-noise ratio
  • N.M de Souza et al . , .American Journal of Roentgenology, 163 (1994), 607-612 describes, in particular, in vivo imaging of the uterine cervix using a ring-design solenoid receiver coil which is placed intravaginally to envelope the cervix.
  • Such intravaginal coils have been successfully used at B 0 magnetic field strengths of up to 1.5 T (64 MHz).
  • Figure 1 shows schematically an MRI detector having a single turn solenoid coil 1 which was used successfully at Bo field strengths of 0.5 T (21 MHz) and 1.5 T (64 MHz) .
  • Tuning capacitor C t ' at the output port of the coil determines the resonance frequency of the coil, giving a correspondingly high dynamic impedance across the coil.
  • the capacitors C m i and C m 2 are then used to provide the required 50 ohm (resistive) impedance to match the co-axial transmission cable 2.
  • the coil 1 operates in receive-only mode and must therefore present a high series impedance during the excitation phase of the MR sequence, so as to avoid distortion of the Bi field and heating of local tissue.
  • the blocking impedance is generated by resonating the capacitor C t ' with a short length 3 of coaxial cable, a condition achieved when the PIN diode 4 is forward biased (so that the PIN diode behaves as a low value resistance) .
  • the PIN diode switch 4 is usually reversed biased (so that the PIN diode behaves as a low capacitance parallel plate capacitor) , which removes the blocking impedance.
  • the solenoid coil 1 of the MRI detector of Figure 1 typically has a short cylindrical structure.
  • a 37 mm diameter, 15 mm deep, copper foil cylinder has been found suitable for cervical imaging.
  • the single turn, short cylindrical structure helps to reduce inductance and hence to reduce conservative electric fields.
  • a receive-only coil must also cause minimal disturbance to the uniformity of the Bi excitation field.
  • a typical priority therefore, is to minimise circulating currents around the receive coil loop during Bi excitation. Nevertheless, as the short cylindrical structure presents a significant cross- sectional area perpendicular to the Bi flux, additional field distortion results, even when the coil loop is open-circuited.
  • a short cylinder solenoid coil is employed as an external surface coil
  • the field distortion is usually remote from the volume of interest (VOI) and can be ignored.
  • VOI volume of interest
  • Bi flux may be partially shielded within the VOI, and unacceptable field distortions may result.
  • flip angle accuracy is increasingly important for quantitative studies and advanced imaging pulse sequences, but this accuracy can be degraded by Bi field distortions.
  • a first aspect of the present invention provides an MRI detector for detecting RF signals having a solenoid coil which completes a single solenoid turn, the coil being split into a plurality of coaxial sub-coils, each of which has an output port, completes the single solenoid turn and has at least one tuning capacitor positioned therein for tuning the resonance frequency of the sub-coil, the sub-coils being spaced along the axis of the coil and being connected in parallel with each other at their output ports.
  • the cross-sectional area of the coil can be decreased, which reduces eddy currents in the coil. In this way, distortion of the Bi excitation field can be significantly reduced.
  • the RI detector may have any one or, to the extent that they are compatible, any combination of the following optional features .
  • the sub-coils are only connected with each other at their output ports.
  • the coil may have two, three, four, five or six sub-coils.
  • each sub-coil has a plurality of tuning capacitors positioned therein which divide the sub-coil into a plurality of sub-coil sections connected in series by the tuning
  • each sub-coil By providing each sub-coil with a plurality of tuning capacitors, for a given coil resonance frequency it is possible to increase the value of the individual tuning capacitors. This allows higher resonance frequencies to be achieved by improving the tuning stability of the coil with body loading, and also by reducing dissipated tissue and radiation losses.
  • the tuning capacitors are circumferentially spaced around each sub-coil to provide sub- coil sections of about equal length.
  • a second aspect of the present invention provides an MRI detector for detecting RF signals having:
  • solenoid coil which completes one or more solenoid turns (and preferably a single solenoid turn)
  • the tuning capacitors dividing the coil into a plurality of coil sections connected in series by the tuning capacitors.
  • the tuning capacitors are circumferentially spaced around the coil to provide coil sections of about equal length.
  • the MRI detector of the first or second aspect may have any one or, to the extent that they are compatible, any
  • each sub-coil may have at least two, preferably at least four, and more preferably at least six or eight of the tuning capacitors.
  • the MRI detector further has:
  • the PIN diode can be at an output port of the coil, or (when the coil is split into a plurality of coaxial sub-coils) the PIN diode can be at the output ports of the sub-coils .
  • the PIN diode behaves as a low value capacitor which does not pass much RF signal due to its high impedance.
  • the PIN diode has, in contrast, a low resistance to RF signals .
  • a third aspect of the present invention provides an MRI detector for detecting RF signals having:
  • solenoid coil which completes one or more solenoid turns (and preferably a single solenoid turn)
  • one or more tuning capacitors positioned in the coil for tuning the resonance frequency of the coil
  • a PIN diode positioned in the coil (e.g. at an output port of the coil) and arranged in series with the tuning
  • circuitry for controlling the bias of the PIN diode, whereby when the PIN diode is forward biased the coil is able to detect RF signals emitted by spins from a subject in response to an excitation Bi magnetic field, and when the PIN diode is zero or reverse biased the coil is substantially prevented from distorting the excitation Bi magnetic field.
  • the MRI detector of the first or second aspect which has a PIN diode positioned in the coil, or the MRI detector of the third aspect, may have any one or, to the extent that they are compatible, any combination of the following optional
  • the circuitry for controlling the bias of the PIN diode may comprise an inductor connected in parallel with the PIN diode, the inductor conveniently providing a current path for biasing the PIN diode.
  • the zero or reverse biased PIN diode has a residual capacitance.
  • the inductor can therefore also resonate with the residual capacitance of the PIN diode when the PIN diode is zero or reverse biased to produce a blocking impedance which substantially prevents the coil from distorting the excitation Bi magnetic field.
  • the MRI detector of the first, second or third aspect may have any one or, to the extent that they are compatible, any combination of the following optional features.
  • the MRI detector may further have:
  • a transmission cable extending from an output port of the coil, or (when the coil is split into a plurality of coaxial sub-coils) from the output ports of the sub-coils, for transmitting RF signals detected by the coil, and
  • a further capacitor positioned in the coil at the output port or ports, wherein the further capacitor matches the coil output impedance to the impedance of the transmission cable.
  • the value of the further capacitor does not, typically, have a significant effect on the resonance frequency of the coil. However, the further capacitor does help to avoid the
  • the transmission cable is a co-axial cable.
  • the solenoid coil may have an internal diameter of less than 40 mm.
  • the solenoid coil may have a depth (i.e. length in the axial direction of the coil) of less than 20 mm.
  • the solenoid coil is tuned to give a resonance frequency for the coil of at least 100 MHz, and more
  • the solenoid coil is adapted for internal in vivo use.
  • the coil can be contained in a suitable housing which, typically, has to: protect the coil
  • the housing can be formed, for example, of an acetyl homopolymer.
  • the coil may be embedded in epoxy resin for water resistance and to seal the electronics.
  • the solenoid coil is adapted for vaginal insertion.
  • the coil may be located, for example, at the end of an approximately 30 cm long elongate member which can provide a guide for the transmission cable and can
  • the solenoid coil can be adapted to be positioned around the cervix, i.e. to envelope the cervix.
  • the coil has an internal diameter of about 37 mm and a depth of about 12 mm.
  • the solenoid coil is a receive-only coil.
  • a further aspect of the invention provides an apparatus for generating a magnetic resonance image of a subject, the apparatus comprising:
  • a transmit coil for generating an excitation ⁇ magnetic field within the subject, the Bi magnetic field being
  • the MRI detector of any one of the first, second and third aspects for detecting RF signals emitted by spins within the subject .
  • the magnet produces a B 0 magnetic field having a field strength of at least 3 T.
  • a further aspect of the invention provides the use of the MRI detector of any one of the first, second and third aspects for internal in vivo imaging of a human or animal body.
  • the use can be in vivo imaging of a cervix.
  • a further aspect of the invention provides a method of imaging a subject comprising:
  • the Bo magnetic field has a field strength of at least 3 T.
  • Figure 1 shows schematically an MRI detector having a single turn solenoid coil
  • Figure 2 shows results for a hybrid method of moments (HMOM) electromagnetic (EM) simulation of the vertical (y-axis) component of the Bi receive field of a single turn, copper, solenoid coil of the type shown Figure 1;
  • Figure 3 shows results for an HMOM EM simulation of the horizontal (x-axis) component of a ⁇ excitation field produced by an excitation loop;
  • HMOM hybrid method of moments
  • EM electromagnetic
  • Figure 4 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when the solenoid coil of Figure 2 is positioned therein;
  • Figure 5 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when a detector coil according to the present invention is positioned therein;
  • Figure 6 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when a further detector coil according to the present invention is positioned therein;
  • Figure 7 shows schematically test objects used to load a body coil and cylindrical copper structure
  • Figure 8 shows images of the Bi excitation field within the dashed outline in Figure 7 when (a) no cylindrical copper structure was in place, (b) a cylindrical copper structure with a single turn was in place, (c) a cylindrical copper structure with two parallel turns was in place, and (d) a cylindrical copper structure with four parallel turns was in place; and
  • Figure 9 shows schematically an MRI detector having a single turn solenoid coil with distributed tuning capacitors.
  • Figure 2 shows results for an H OM EM simulation of the vertical (y-axis) component of the Bi receive field of a single turn, copper, solenoid detector coil 11 of the type shown Figure 1, the gap at the front of the coil corresponding to the position of the coil output port at C t ' .
  • Figure 3 shows results for an HMOM EM simulation of the horizontal (x-axis) component of a Bi excitation field produced by an excitation loop 12. For modelling convenience, a relatively small excitation loop was modelled, which resulted in a less uniform excitation field than would normally be the case.
  • the Bi excitation field is typically circularly polarised in the x-y plane, and the orientation of the detector coil 11 ensures that the cross-sectional area presented by the coil perpendicular to the ⁇ excitation field varies throughout each cycle of the Bi excitation field, being at a maximum with Bi in the x direction and at a minimum with ⁇ in the y direction.
  • Faraday's Law dictates that any reduction in the cross- sectional area of the copper presented perpendicular to Bi commensurately will reduce eddy currents.
  • the coil was split into two coaxial sub-coils.
  • Figure 5 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when such a detector coil 21 is positioned therein.
  • the coaxial sub-coils are spaced from each other along the y- axis, and both complete the single solenoid turn of the detector coil 11 of Figures 2 and 4.
  • the overall cross-sectional area in the x direction is reduced. Relative to the detector coil 11, this leads to a decrease in the distortion of the Bi excitation field, and also to an increase in the field strength throughout the volume encircled by the detector coil.
  • FIG. 6 shows results for an HMOM EM simulation of the horizontal (x-axis) component of the Bi excitation field produced by the excitation loop of Figure 3 when such a detector coil 31 is positioned therein.
  • the detector coil 31 now has four coaxial sub-coils spaced from each other along the y-axis, and again the depth of the detector coil 31 in the y direction is not increased.
  • cylindrical copper structures were built on 37mm diameter x 12mm deep acetyl formers, the structures consisting of, respectively, one, two and four parallel turns.
  • Figure 8 shows images of the Bi excitation field within the dashed outline of Figure 7 when (a) no cylindrical copper structure was in place, (b) the "unsplit” cylindrical copper structure with a single turn was in place, (c) the “split” cylindrical copper structure with two parallel turns was in place, and (d) the "split” cylindrical copper structure with four parallel turns was in place.
  • the single-turn 12mm deep, open-circuit, receive-only coil
  • Figure 9 shows schematically an MRI detector according to the present invention having a single turn solenoid coil 51 in which, instead of a single tuning
  • N tuning capacitors C t are circumferentially
  • each of these coil sections has an inductance which is approximately 1/(N+1) the inductance of the equivalent
  • each tuning capacitor has a value that is N times the capacitance of the single tuning capacitor of the equivalent undivided coil.
  • the values of the tuning capacitors can be significantly increased, such that the tuning stability of the coil 51 with body loading is substantially improved.
  • the coil 51 can be tuned to provide a resonance at 3 T (128 MHz), well above the 1.5 T (64 MHz) achieved with the coil 1 of Figure 1. Dissipated tissue and radiation losses can also be reduced.
  • each sub-coil can be divided into sub-coil sections by respective tuning capacitors
  • a high series impedance during the excitation phase of the MR sequence is achieved by forward biasing the PIN diode 4 so that the capacitor C t ' at the output port of the coil resonates with the short length 3 of coaxial cable to generate a blocking impedance.
  • the performance of this circuit is dependent, however, on the physically
  • the capacitor C m is positioned at the output port of the coil 51.
  • the value of C m can conveniently be set to match the, typically 50 ohm, impedance of co-axial transmission cable 53 which extends from the output port.
  • the approach to producing a blocking impedance used in respect of the detector of Figure 1 is adopted for the detector of Figure 9, it becomes
  • a PIN diode 54 is positioned in the coil 51, in series with the tuning capacitors C t and impedance-matching capacitor C m , rather than being positioned outside the coil as in the detector of Figure 1. Further, an inductor L is connected in parallel across the PIN diode 54 and capacitor C m to provide a current path for the diode d.c. drive.
  • the PIN diode 54 is zero or reversed biased.
  • One option is to choose the inductor L so that it is close to self-resonance at the resonance frequency of the coil 51 (i.e. about 128 MHz at 3 T) .
  • the zero or reversed biased PIN diode has a very low residual capacitance, typically of about 2.2 pF, which gives an impedance of about 566 ohms at 128 MHz. Therefore, preferably the inductor L is chosen to have a similar impedance (e.g. about 566 ohms at 128 MHz) to the zero or reversed biased PIN diode so that the inductor resonates with this residual capacitance. In this way, impedances of about 15 kohms in series with the coil 51 between positions A and Ai can be achieved, which provides effective detuning of the coil.
  • a preamplifier input at the other end of transmission cable 53 is typically protected by shunting its input with a protecting PIN diode during the excitation phase. This causes the impedance "seen" at the coil end of the transmission line to depend on the line length.
  • the line length By adjusting the line length to be a half-wavelength (approximately 0.775 metres at 128 MHz), a low resistance is presented across C m , which avoids dependency of the detuning circuit on the value of Cm.
  • the PIN diode 54 is forward biased to provide a low value resistance-
  • the presence of the inductor L (providing an impedance of e.g. about 566 ohms) connected across the 50 ohm output port has little influence on circuit performance.
  • the inductor L can be chosen so as to resonate with the capacitance of the PIN diode 54 when not forward-biassed. This maximises the impedance of the blocking circuit when in the excitation phase.
  • the r.f. current circulating within L is increased, and significant magnetic field can exist around L. This field can distort the local Bi excitation field, leading to image artefacts.
  • This problem can be reduced or overcome by positioning L away from the imaging field of view within, e.g. along the coil "handle" which also carries the co-axial transmission cable 53.
  • a conductive shield may be employed around L.
  • the Bi excitation field produced by the transmit coil has both electrostatic and magnetic components. Although not directly involved in the MR imaging process, the electrostatic
  • a conductor running parallel to the E-field can act as short- circuit which may carry a potentially large "common mode" current. This can be the case when a cable connecting a receive-only coil passes between a patient and scanner input electronics outside the magnet bore. At frequencies at or above 100 MHz, such cables may have a physical length close to a half wavelength and a resonance can occur, amplifying the induced r.f. current and increasing the hazard.
  • An insertable MR probe of the type shown schematically in Figure 9 (adapted to be positioned around the cervix), may be more subject to the effects of these induced r.f. currents (compared to a standard surface coil) due to stronger capacitive coupling with the body.
  • Baluns As well as reducing the induced current in general, a number of these circuits, often referred to as Baluns, may be
  • baluns are generally employed, one close to the coil, within the handle, so as to provide a common mode impedance between the cable and the patient, and a second balun along the cable approximately one quarter wavelength from the first. See D.M. Peterson et al. Common Mode Signal Rejection Methods for MRI: Reduction of Cable Shield Currents for High Static Magnetic Field Systems,
  • Magnetic Resonance Part B Magnetic Resonance Engineering

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  • Physics & Mathematics (AREA)
  • Condensed Matter Physics & Semiconductors (AREA)
  • General Physics & Mathematics (AREA)
  • Magnetic Resonance Imaging Apparatus (AREA)

Abstract

L'invention concerne un détecteur IRM, destiné à détecter des signaux RF, comportant une bobine solénoïde comprenant une ou plusieurs spires solénoïdales. Le détecteur comprend en outre un ou plusieurs condensateurs d'accord placés dans la bobine et destinés à ajuster la fréquence de résonance de la bobine. Le détecteur possède en outre une diode PIN placée dans la bobine et montée en série avec les condensateurs d'accord. Le détecteur comprend également un circuit de commande de la polarisation de la diode PIN. Lorsque la diode PIN est polarisée dans le sens direct, la bobine est en mesure de détecter des signaux RF émis par les spins d'un sujet en réponse à un champ magnétique d'excitation Bi. Lorsque la diode PIN n'est pas polarisée ou est polarisée en sens inverse, la bobine ne peut fondamentalement pas provoquer de distorsion du champ magnétique d'excitation.
PCT/GB2011/000667 2010-04-29 2011-04-28 Détecteur irm Ceased WO2011135312A2 (fr)

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
GBGB1007198.3A GB201007198D0 (en) 2010-04-29 2010-04-29 MRI detector
GB1007198.3 2010-04-29

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Publication Number Publication Date
WO2011135312A2 true WO2011135312A2 (fr) 2011-11-03
WO2011135312A3 WO2011135312A3 (fr) 2012-01-05

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WO (1) WO2011135312A2 (fr)

Cited By (1)

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* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
WO2024012921A1 (fr) * 2022-07-14 2024-01-18 Koninklijke Philips N.V. Bobine réceptrice de radiofréquence à fusible électronique

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