WO2009001084A1 - Système d'aimant pour utilisation dans l'imagerie par résonance magnétique - Google Patents

Système d'aimant pour utilisation dans l'imagerie par résonance magnétique Download PDF

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Publication number
WO2009001084A1
WO2009001084A1 PCT/GB2008/002182 GB2008002182W WO2009001084A1 WO 2009001084 A1 WO2009001084 A1 WO 2009001084A1 GB 2008002182 W GB2008002182 W GB 2008002182W WO 2009001084 A1 WO2009001084 A1 WO 2009001084A1
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Prior art keywords
coils
magnet
coil
shielding
compensation
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PCT/GB2008/002182
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English (en)
Inventor
Ian Leitch Mcdougall
Peter Hanley
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Oxford Instruments Plc
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Publication of WO2009001084A1 publication Critical patent/WO2009001084A1/fr

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/381—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets
    • G01R33/3815—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets with superconducting coils, e.g. power supply therefor
    • 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
    • G01R33/38—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
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G01R33/3875—Compensation of inhomogeneities using correction coil assemblies, e.g. active shimming

Definitions

  • the present invention relates to a magnet system for use in magnetic resonance imaging (MRI).
  • MRI magnetic resonance imaging
  • Magnets for MRI systems have evolved with a typically tubular examination space, derived from the use of "solenoid” magnet geometry. This is because the solenoid arrangement provides excellent uniformity of the magnetic field. For practical applications the "solenoid” arrangement has been effected using quasi-solenoids in which sets of circular coils are distributed along the axis of the solenoid.
  • values of the field derivatives describing rotation (uniformity) of the central magnetic field may be calculated and plotted against (for convenience) a given coil radius and current, at points along a uniform cylindrical bore tube. Such plots show that the value of field derivatives exhibit harmonic profiles with bore length, and that there are zero value positions, where the field derivative profile changes from a positive to negative value, or vice versa.
  • the design process seeks to balance a number of coils with respect to their strengths of field derivatives each side of such zero points, the result being that chosen field derivatives sum to zero for the coil set. Similar harmonic plots of field derivatives may be obtained against coil radius at a particular bore length. Again, sets of coils may be chosen to provide that the sum of a particular order field derivative is zero, and, consequently, this order does not contribute to field non-uniformity at the origin.
  • Standard tubular MRI magnets are therefore formed from a series of coaxial magnetic field coils, all of similar diameter.
  • the coils considered as pairs across the mid-plane, contribute magnetic field to the working total field for MRI at the origin, or centre of the magnet, along with characteristic non-uniformity in the central magnetic field.
  • non-uniformity due to one coil pair may be cancelled out by the non-uniformity of another pair, when all coils carry current in the same sense.
  • the non-uniformities can be described by nth orders of derivatives of the magnetic field, considered as an expansion of increasingly high orders.
  • a magnet system for use in magnetic resonance imaging, the magnet system being adapted when in use to generate a resultant magnetic field, within a working region, of sufficient homogeneity to enable magnetic resonance imaging to be performed, the system comprising:- a primary magnet comprising a pair of coils positioned upon a common axis passing through the centre of the working region, the coils being spaced symmetrically about an origin upon the common axis and having dimensions such that each coil is at or adjacent a nominal position at which the second order derivative of a corresponding magnetic field in the working region is substantially zero for a coil pair located symmetrically about the origin at such a position; a shielding magnet comprising a pair of coils positioned upon the common axis, the coils being spaced symmetrically about the origin and having dimensions such that each coil is at or adjacent a nominal position at which the second order derivative of a corresponding magnetic field in the working region is substantially zero for a coil pair located symmetrically about the origin at such a position, the shielding
  • the coils are at or adjacent in the sense that each coil has a corner of its cross-section within a distance of the nominal position line in radial and axial co-ordinates, wherein the said distance is not greater than the maximum of the radial or axial extent of the coil section, whichever is the larger. More preferably, the coils are at or adjacent the nominal position in the sense that the coils have at least part of the coil cross-section intersecting the nominal position.
  • this highest field derivative may be sixth order, or more typically, eighth, or even tenth order.
  • the first group comprise the primary magnet and the shielding magnet, whereas the second group comprise the compensation magnet.
  • the said "nominal positions" which cause the cancellation of nth order magnetic field gradients are nominal positions derivable from magnetic theory.
  • the coils of the magnets according to the invention are positioned at or adjacent these positions, the differences between the actual positions and the nominal positions being due in particular to the finite size of the coils and the fact that, in order to produce a desired geometry, whilst minimizing the use of coil material, some non-zero magnitude in the particular field derivative may be advantageous, practically.
  • the respective nominal position is typically defined by a frusto-conical surface of rotation about the common axis, wherein the said derivative of the respective magnetic field is zero for a nominal set of ideal coils defining a circle, each part of which intersects with said surface.
  • Such ideal coils have zero radial and axial thickness.
  • the frusto-conical surface is preferably defined by a cone angle, the cone angle being the angle at the cone vertex between the common axis and a line lying parallel to said surface and passing through the vertex, and wherein the cone angle for each of the pairs of coils of the compensation magnet is in excess of that of the primary magnet.
  • each coil is defined by a circle centred upon the common axis and wherein each part of said circle is equally at or adjacent the nominal position.
  • the coils are typically circular in geometry, they have optimum (well known) mechanical properties. Thus high values of MRI field can be obtained when the coils will operate at high internal stress values.
  • a hitherto unobtainable combination of access to the subject being examined is achieved, with very high image quality, because of the high signal to noise provided by the high magnetic field strength obtainable.
  • a high image quality and access to the "patient" subject is particularly valuable to the research community who wish to use MRI as a gauge of experiments in molecular medicine. For example it may be used to observe brain activity during investigation of the effect that drug compounds may have on cognitive ability.
  • the net current within at least one pair of the coils of the compensation magnet is caused when in use to flow about the common axis in a counter-running manner with respect to the net current in the primary magnet.
  • the coil arrangement of the present invention provides for only a small amount of counter-running current which prevents the magnets unnecessarily working against one another. The efficiency of the system is therefore improved.
  • the compensation magnet comprises a number of pairs of coils, such as three pairs, each pair having a different radius.
  • the pair of coils having the smallest radius is preferably counter-running.
  • a pair of coils for one or more of the primary, shielding or compensation magnets typically has coil windings at the respective nominal position. This is particularly the case for the compensation magnet where it is preferred that each coil has windings at the nominal position.
  • the windings may be distributed substantially uniformly within the coil cross-section and it is the centre of the said cross-section (at the average position of the windings) that may be used to define the position of the coil.
  • the shielding magnet has coils having the largest radius.
  • Each of the shielding and primary magnets typically has coils having a radius in excess of those of the compensation magnet.
  • the large radius coils of the shielding and primary magnets is operative to produce a relatively small high order (such as 8 th order) derivative of the magnetic field in the working region.
  • the primary magnet typically provides the largest contribution to the resultant magnetic field, such as in excess of 75% of the resultant magnetic field strength.
  • the primary, shielding and compensation magnets are operative together to produce a sufficiently homogeneous magnetic field in the working region for performing magnetic resonance imaging. It should be noted however that the arrangement is such that, when taken alone, each of the primary, shielding and compensation magnets do not produce a sufficiently homogeneous magnetic field in the working region for imaging.
  • the shielding magnet is provided so as to limit the extent of the resultant magnetic field at locations distal from the magnet system.
  • the shielding magnet is adapted when in use to have a net current flow in its coils which is counter- running with respect to the net current flowing within the coils of the primary magnet.
  • the coil pair of the shielding magnet is placed adjacent the nominal second order derivative position so as to cause a non-zero magnitude second order derivative having a first polarity and the coil pair of the primary magnet is placed adjacent the nominal second order derivative position so as to cause a non-zero magnitude second order derivative having a polarity opposite to the first polarity, such that the second order magnetic field derivatives of the primary and shielding magnets cancel.
  • the combined effect of the primary and shielding magnets is to cause a non-zero value of the second order magnetic field derivative.
  • the coil pair of the shielding magnet is placed adjacent the nominal second order derivative position so as to cause a non-zero magnitude second order derivative which is not fully cancelled by the primary magnet.
  • the shielding magnet may comprise two sets of coils spaced apart along the common axis. These may each be upon one side of the nominal position. Each of the sets of coils of the shielding magnet may be positioned closer to the origin at the centre of the working region than the nominal second order derivative position.
  • each of the magnets could be powered separately, typically they are each powered together by electrically connecting them in series.
  • a suitable controller is preferably provided to control the electrical current within at least one and preferably each of the primary, shielding and compensation magnets.
  • the compensation magnet In order to produce a magnetic field within the working region of high uniformity it is necessary to accurately position the respective coils of the magnets.
  • the additional coils may be used in one of two ways. They may be adapted to allow the positions of the magnetic centres of the said at least one pair of coils to be modified by controlling the current within the additional coils.
  • the additional coils may alternatively, although preferably additionally, be adapted to allow the strength of the magnetic field of the compensation magnet to be modified by controlling the current within the additional coils.
  • the additional coils may therefore modify the strength and/or position of the effective centre of the pairs of coils.
  • At least three additional coils may be provided for each coil in question, these being preferably distributed symmetrically about each coil.
  • Four additional coils are preferred, these additional coils being positioned upon either side of each of the said at least one pair both radially and axially with respect to the common axis.
  • the powering of the coils independently of the coils to which they relate can be performed using the controller (or the controller and second controller) and suitable switching apparatus.
  • the second controller is adapted in use to provide electrical current to the said coils such that the corresponding strength of the magnetic field generated by the said additional coils alone is up to about 10% of the magnetic field strength of the compensation magnet.
  • the system further comprises a set of gradient coils, each gradient coil being positioned at or adjacent the nominal second order derivative position.
  • each magnet coil within the system is positioned substantially within a geometrical envelope defined by two symmetrical conical surfaces representing the second order position that extend away from the centre of the working region.
  • Such an envelope may be further bounded by a cylindrical surface at the largest radius of the magnet coils.
  • the coils are enclosed within a common housing.
  • the geometry of the apparatus may be such that the primary and shielding magnets are arranged in a first part of the housing which has a substantially annular geometry.
  • the compensation coils may be arranged in a second part of the housing having substantially annular geometry of smaller axial and radial dimensions than the first part.
  • the second part of the housing may be provided with walls which taper in axial dimension towards the common axis.
  • the present system preferably further comprises a table upon which a subject is positioned for imaging when the system is in use, the table defining a table plane, and a system mounting, adapted to allow the magnets to be rotated with respect to the table plane, such that the common axis is rotated within a plane substantially normal to the table plane.
  • the magnets may be arranged to be rotated in a plane substantially parallel to the table plane.
  • the coils of the magnets described above may be formed from high conductivity resistive wire. Preferably however, they are formed from superconducting wire, this being cooled when in use, using a coolant.
  • the axial space between the respective coils of particularly the shielding and primary magnets can be used to accommodate a reservoir for coolant, such as a cryogen in the form of helium or nitrogen.
  • the reservoir may be divided into two separate reservoirs, the reservoirs being positioned upon either side of a mid-plane of the system, the said plane passing through the centre of the working region and having a plane normal parallel to the common axis. This allows the system to operate in a wide range of orientations since the use of two dedicated reservoirs, one for each half of the magnet system, ensures a supply of coolant to the magnets during use.
  • we also provide a method of constructing such a magnet system comprising the steps of:- i) mounting one coil of each pair of coils of the primary, shielding and compensation magnets to a first support; ii) mounting a first coolant reservoir to the first support, the first coolant reservoir being adapted to supply the said one coil of each of the primary, shielding and compensation magnets with coolant during use; iii) mounting the other coil of each pair of coils of the primary, shielding and compensation magnets to a second support; iv) mounting a second coolant reservoir to the second support, the second coolant reservoir being adapted to supply the said other coil of each of the primary, shielding and compensation magnets with coolant during use; and, v) mounting the first and second supports together.
  • first support, coils and first reservoir may comprise a first half of the magnet system and the second support, coils and second reservoir may comprise a second half of the magnet system.
  • the first and second halves are preferably arranged as substantially mirror images of one another.
  • Figure 1 is a schematic section through a first example system
  • Figure 2 is a table showing the field derivatives and required compensation magnet for a first coil pair
  • Figure 3 is a table showing the field derivatives and required compensation magnet for a second coil pair
  • Figure 4 is a table showing the field derivatives and required compensation magnet for a third coil pair
  • Figure 5 is a table showing the field derivatives and required compensation magnet for a fourth coil pair
  • Figure 6 is a table showing the field derivatives and required compensation magnet for a fifth coil pair
  • Figure 7 is a table showing the field derivatives and required compensation magnet for a sixth coil pair arranged as a shielding magnet;
  • Figure 8 is a table showing the field derivatives and required compensation magnet for a seventh coil pair arranged as a shielding magnet;
  • Figure 9 is a table showing details of an example full magnet system;
  • Figure 10 is a table showing details of other example full magnet systems
  • Figure 11 shows a coil with a set of additional coils
  • Figure 12 shows the steps of a method of constructing the magnet systems
  • Figure 13 shows a section through a magnet system
  • Figure 14 shows relative movement between the system and a subject
  • Figure 15 shows a first schematic perspective view of a magnet system
  • Figure 16 shows a second schematic perspective view of the magnet system
  • Figure 17 is a table showing calculated magnetic field values
  • Figure 18 is a table comparing exact and calculated field values
  • Figure 19 shows the effect of tolerances in coil fabrication
  • Figure 20 shows the compensation of the tolerances using additional coils.
  • the magnets comprise three magnets.
  • the first is a primary magnet (for generating the majority of the magnetic field) and the second is a shielding magnet to reduce the stray field.
  • This combination produces a very axially short magnet system where all the coil winding volumes lie outside a cone shaped region whose apex is the magnetic centre of the magnet system.
  • the angles between the lines joining the geometric centres of the coil sections (the winding cross sections) and the magnetic centre of the system, and the magnetic axis is greater than about 63 degrees.
  • all the winding volumes typically lie outside a cone shaped region whose apex is the magnetic centre, whose axis is the magnetic axis and whose angle is greater than 45 degrees.
  • Each of the magnet systems described below produce a working region of sufficient uniformity to be suitable for MRI.
  • Some of the coils of the magnet system are energised in opposition to others so that the combination produces the required magnetic field homogeneity and at the same time reduces the net magnetic moment of the system so as to effectively screen the system when viewed from a distance substantially greater than the radius of the largest coil (shielding).
  • the coils described may be resistive and continuously energised although they may be superconductive.
  • Superconductive coils may be fitted with persistent mode switches, allowing them to be used in persistent mode (without being in constant connection to a power supply).
  • Conventional superconductors or high temperature superconductors may be used to construct some or all of the magnet coils.
  • the systems described produce a short MRI magnet, with the length of the "patient" examination bore less than V ⁇ the diameter of the examination bore (typically VA the diameter).
  • VA the diameter
  • the magnetic fields from each of the magnets need to be combined to produce a designed central field (at the origin of the short tubular bore) that is sufficiently uniform for MRI.
  • the short nature of the tube allows a patient with their head in the central field to have line-of-sight to the examination room, external to the magnet system. This is a great advantage in the reduction of physiological stresses which otherwise may distort results from MRI observations.
  • Figure 1 is schematic and representative of all examples described. It shows the system in section when viewed from the side.
  • the system coordinates are defined by a common axis (Z) upon which the geometrical centre of each coil of the magnets are positioned.
  • a mid-plane XY is also shown, this having a plane normal coaxial with the axis Z.
  • the position of intersection between the plane XY and the axis Z defines an origin of the system which is also, by design, the centre of a working region of the magnet system 1.
  • the coils which comprise the respective magnets are of circular geometry having a centre upon the common axis and lying in a plane normal thereto (parallel with the mid-plane XY). Each coil is described by a radius "a" and a half-length "b".
  • the radius a defines the distance of the effective centre of the multiple windings of the coil from the common axis Z
  • the half-length b describes the distance along the common axis Z from the origin at which the geometrical centre of the respective coil is located.
  • magnet coils discussed herein are symmetrical about the magnetic centre, and thus in order to produce a B2 of zero for example, an ideal coil is placed at the respective position on each side of the origin.
  • the ideal coil zero lines are shown in Figure 1 as B2, B4, B6, B8.
  • coils of the primary, shielding and compensation magnets are positioned at or adjacent the B2 and B8 lines.
  • the primary magnet coils are illustrated at 2. Note that only one quadrant of the coil set is shown in detail. Thus the coil 2 encircles the axis Z and the second coil 2 (of the pair) is positioned at a similar mirrored position upon the other side of the origin. In this example the cone of rotational positions about the Z axis defining the line B2 passes through the middle of the primary magnet coil 2.
  • the coils of the shielding magnet are shown at 3,4. It will be noted that this comprises a set of two coils, with each having a similar radius a (in excess of the coils 2) but positioned in a split formation axially (different b).
  • the coils 3 intersect the B2 line, whereas the coils 4 are positioned adjacent the line B2 (a short distance therefrom). Since the coils 3,4 form the shielding magnet, the current within them is counter-running with respect to that within the primary coils 2.
  • each of the coils of the compensation magnet is intersected by the B8 line.
  • the B8 line intersects the coil windings of the compensation coils, although it should be noted that, due to their finite size, the B8 line may not pass through the geometrical centre of each coil section.
  • the coils 6 and 7 each carry an electrical current having a flow sense similar to that of the primary magnet 2.
  • the coil 5, is counter-running with respect to the coils 2,6,7.
  • the coils are enclose in a support structure or housing 10, this having an approximate "T" shape in section (above the Z axis).
  • the upper part or cross piece of the "T” accommodates at least the coils at or adjacent the B2 line, whereas the stem of the "T” accommodates the B8 coils.
  • the "T” has a distal part from the Z axis which has a greater axial dimension than the proximal part nearer the axis.
  • FIG. 1 A subject under examination (which is typically an animal, or more often a human patient) is illustrated in Figure 1.
  • the head of a human subject 12 is shown positioned at the origin of the system 1.
  • the subject has line-of-sight out of the bore of the magnet system, in particular due to the axially short arrangement of the housing 10.
  • the axially narrow part of the housing enhances the open nature of the system.
  • a person located as illustrated may therefore read a book held slightly above them (adjacent the axially short part of the housing).
  • other measurements and interventions may be performed upon the subject located in this position because the short arrangement of the magnet now makes this possible.
  • the design can be arranged in this manner since the positioning of the coils on the B8 line allows very good access to the patient.
  • the primary and shielding magnets are anchored about the B2 line on a larger diameter, and have a lower a/b ratio (which can also be thought of as a lower conic angle). This is less convenient for access, so the diameter is made large.
  • the large diameter magnets comprise the coils that provide the main proportion of the MRI field intensity (more than 75% from the primary magnet), as well as the external shielding field coils (for the shielding magnet). It is known that the magnetic moments of the primary and shield magnets are similar for optimum shielding (Bm x Area m ⁇ Bs x Area s) and this approach is adopted here also.
  • B2 is that the intermediate even orders of n, B4 and B6, are well within the range of values that the compensation magnet will compensate.
  • the values of the field derivative B2 can be chosen to cover a wide range.
  • the magnet anchored on B2 is physically much shorter in axial length than a standard MRI magnet, but it can produce values of B2 typical of a standard magnet. This can make the compensation of B4 and B6 errors by the smaller magnet, with larger conic angle, particularly efficient in terms of ampere turns. This is because the smaller diameter magnet does not have to "waste" its own ampere turns compensating for its own value of field derivative B2, as it compensates B4 and B6 of the larger diameter magnet.
  • the latter requirement usually means locating some coils remote from the mid-plane so as to provide possibilities for cancellation of particular field derivatives between sets of coils, without these coils producing their own high order field derivatives.
  • Our basic approach to reconcile these requirements is to use both cancellation of field derivatives between, at least two coils, and to avoid values of certain field derivatives by placing some coils on or adjacent "zero locations". In particular, higher order field derivatives are also reduced to acceptable values by the use of large diameter coils.
  • the primary and active shield magnets on the smaller conic angle provide the main field for MRI and the external shielding.
  • Each has only the second order field derivative controlled to be either zero, or a finite target value. Residual errors in field uniformity, described by 4 th , 6 th and 8 th order field derivatives are minimised as far as is practical by using large diameter coils, but achieving zero value for these orders only occurs when the fields of the first and second magnet are combined.
  • the compensation magnet is chosen such that the radial and axial dimensions of are located on the zero of the highest order of field derivative required to be cancelled (eighth order for example).
  • the primary magnet may have a non-zero negative B2 value for a positive ampere turns ("Nl) value, whereas the shielding magnet has a a positive B2 for negative Nl.
  • the sum of the coils B2 can be made to be zero.
  • the compensation magnet has only to compensate B4 and B6 while presenting no B8, which requires fewer coils in the compensation magnet; and,
  • the shielding magnet coils (inside the line) have a positive B2 for negative Nl.
  • the sum of the coils B2 values is that normally provided by a tubular magnet of approximately twice the length, when using two coils with NI in the same sense.
  • the conical magnet "looks" longer at its origin with respect to field uniformity described by B2 than it actually is. This reduces the ampere turn correction "load" required of the compensation magnet that compensates high order field derivatives.
  • the coils of the shielding magnet are placed in the vicinity of the conic angle projected from the origin of the bore tube that coincides with the lowest even order of field derivative (second order), so that adjustments of each coil in radial and axial space, along with adjustments in relative ampere turns provide controlled target values of the second order field derivative.
  • second order field derivative can be summed to zero by choice of current ratio and radial/axial space coordinates. It is also possible to create a sum value of second order field derivatives which is typical of a longer tubular magnet. This makes subsequent field uniformity at the origin easier to achieve when the compensation magnet is used to generate the full resultant field.
  • the compensation magnet needs fewer ampere turns than otherwise needed, although the precision of the coil build is more demanding.
  • the shield magnet is preferably divided into two coils, both with negative current (with respect to the primary magnet coils). The purpose of this is to achieve optimum values for the radial and axial shielding of magnetic field, while also achieving, where required, the option of a maximum value of B2.
  • the compensation magnet is preferable to arrange the compensation magnet to have at least one coil carrying current in a direction opposed to the current producing the main field.
  • the compensation magnet has the coils arranged on the radius (from the origin of the bore tube) that coincides with a zero value for the field derivative of the highest order of field error that the magnet is intended to have near to or equal to zero.
  • this is eighth order (B8), although tenth order is feasible also, particularly using the current and geometry control technique described below.
  • the primary and shielding magnets be chosen at such a coil diameter that the 8th order field derivative tends to zero, and in any case the 8th order is chosen smaller than would produce unacceptable MRI image distortion.
  • the compensation magnet is placed on the conic angle projected from the origin of the bore tube that coincides with the first zero value of the 8th order field derivative.
  • the compensation magnet coil set can be organised to compensate the intermediate orders of field derivative produced by the larger diameter coil set, without contributing unwanted high order field errors.
  • the primary magnet has a positive ampere turns values with the shielding magnet providing a negative ampere turns value.
  • the primary coils are positioned slightly closer to the origin.
  • the shielding magnet coils are also closer to the origin. This produces an overall shorter magnet axially.
  • the resultant B2 of the primary and shielding magnets together is cancelled by non-zero (positive and negative values of the ampere turns of the primary and shielding magnets). Again there is some residual B4,B6.
  • each (primary and shielding) magnet is on the B2 line the diameter is chosen such that B8 and above are close to zero.
  • the primary and shielding coils have a radius of the order 80cm.
  • the compensation magnet comprises, in this case, two positive NI coils and a small counter-running trimming coil on the smallest diameter.
  • the B4 and B6 produced by the compensation magnet are equal and opposite to the combined primary and shielding magnet, B4 and B6.
  • the compensation magnet coils are set in a radius range from 28cm to 50cm.
  • the combined magnet system balances B2 to B6 to zero, and controls B8 to an acceptably low level.
  • the compensation magnet may also be provided with additional coils (see below) to produce a net resultant B6 or B8 to "sharpen the imaging field of view".
  • Figures 2 to 10 Some tables of field profiles are shown in Figures 2 to 10. These are expressed in terms of field derivatives to show how a system of coils can be organised to give a uniform central field, and how grouping coils in particular regions of radial/axial space give emphasis to control of either low or high orders of field derivatives.
  • the tables each comprise the error term orders (calculated field derivatives)
  • BO to either B6, or in some tables to B10 or B12, for finite coil cross-sections (radial build and length), where the Bn (with “n” being the derivative order) are tabulated against axial position of the coil.
  • Each table represents one coil of one mean radius.
  • Figures 2 to 6 illustrate tables for coils configured as a primary magnet
  • BO to Bn are for half of the magnet, these being the sets of coils along the axis of the magnet to one side of the mid-plane. All Bo values are given in Gauss, with dimensions in centimetres.
  • Bn values given are those taken on a 15cm radius position from the origin, that is, at the edge of the working region 15.
  • the coil finite dimensions are typically those of a whole body, 2 Tesla, magnet, for main field and a shield field.
  • Cross-section inner radius a1 ; outer radius a2; inner half-length b1 ; outer half-length b2), ampere turns (Nl), turns per coil (“T coil-1”), turns per square centimetre ( 1 T cm-2”) and current (I) are appropriate examples for this type of magnet.
  • Each table also shows the necessary strength in terms of ampere turns (Nl) required by a three-finite-coil compensation magnet.
  • the dimensions of the compensation coils are given on the left hand side of each table (these being "L” for large radius, "M” for middle radius and “S” for small radius). Note that the same compensation coil dimensions are used for all of the included tables.
  • B8 is zero in the compensation arrangement, and small for the set of main coil radii in the tables by virtue of the absolute magnitude of main coil radius, a mean.
  • small (sub-millimetre) b changes of the inner radius compensation coil can cancel small B8 arising from the main coil, without significant changes in the NI values of the compensation coils required to compensate B2, B4, B6 of the main coil.
  • the compensation magnet produces its own Bo value when the NI set provides the required B2, B4, B6 compensation.
  • the combined main and compensation magnet have B2, B4, B6, B8 cancelled to zero, and the combined Bo is the signed sum of the respective value in the Bo column and that of the "Comp Bo" column. Whilst the procedure may be used to compensate a primary magnet, similarly it is used to compensate a shielding magnet (whilst taking the polarity reversal into account).
  • the compensation three coil set for all tables has a minimum radius "a" mean of 41 .48cm. This sets the internal bore as about 40cm radius.
  • other compensation magnet dimensions could be calculated to use in look-up tables.
  • the mean half-length b towards the top of the table where compensation NI are smaller and Bo from the compensation magnet is positive, for efficiency, but not the shortest possible length. For example, we take "b mean” 54cm.
  • the only counter-running coils are those of the shielding magnet and the innermost coil of the compensation magnet.
  • the homogeneity of the resultant field is shown to be about 5 parts per million at a working region radius of 15cm, and about 100 parts per million at 20cm radius.
  • the compensation magnet coils L and M both have positive NI values, and because these coils are closest to the main coil, peak fields should be low, meaning high field systems are economically practical. Note further that, by interpolation of the tables it would be possible to exactly cancel the NI of the smallest compensation coil, thereby achieving a system magnet with increased bore radius.
  • Figure 10 shows some details of other example systems derived using a similar process.
  • FIG. 11 shows the cross-section of a coil 20 which is representative of a generic coil of the compensation magnet (although similar principles can be applied to the primary and shielding magnets).
  • the coil 20 has a rectangular cross-section with dimensions defined by maximum and minimum radial (a2 and a1 respectively) and axial dimensions (b2 and b1 respectively).
  • the Z direction is also indicated in Figure 11. Additional coils 21 ,22,23,24 are positioned upon each face of the coil 20.
  • the coils 21 ,22,23,24 can be used in two possible modes. The first is to provide a net change to the effective ampere turns of the coil 20 in total, without changing the centre of the coils section as a whole. It will be appreciated that all main coils of the magnets are preferably connected in series and therefore adding currents to specific coils is not possible in such an arrangement. It is also undesirable to have separate main coil power supplies.
  • the arrangement in Figure 11 solves this problem by allowing a current flow within the additional coils 21 ,22,23,24, and this may be in the same sense as that flowing in coil 20 so as to increase the strength of the magnetic field from coil assembly as a whole.
  • the net current may also be reduced by providing a counter-running current in each of coils 21 ,22,23,24.
  • the second mode is the powering one or more individual coils of the additional coils 21 ,22,23,24.
  • This may also include providing normal running current in one or more coils and none or counter-running current in others.
  • the effect of this is to change the effective geometrical centre of the coil cross-section which provides the ability to produce very small effective movements of the coil without any physical movement being required.
  • the additional coils are distributed symmetrically with respect to the cross section of each coil. Specifically, additional windings controlled by switches are placed on the inner and outer radial faces of each coil cross section in equal amount, and on the axial inner most and axial outer most longitudinal faces of each coil in equal amount. The radial and longitudinal additional coil faces may be separately adjusted.
  • an additional (second) coil controller 25 is illustrated to control the current flowing in each of the additional coils. This may form part of a general controller for the other coils of the magnets, or it may be a separate device. Note that this includes suitable switching apparatus, particularly to effect the use of superconductors in persistent mode (if required).
  • Each coil may therefore have its net ampere turns increased by switches across the auxiliary turns of the additional coils, independently from the other coils of the set.
  • each of the coils of the compensation magnet is preferably organised such that with no current through its own additional coils, the coil in question provides a lower value of the field derivatives (that it contributes) than is required for full compensation.
  • the compensation magnet "under-compensates” the intermediate field derivatives of the combined primary and shielding magnets.
  • the network of switches is therefore preferably used to increase the effective ampere turns in each coil of the compensation magnet. It should be noted that the changes in effective ampere turns do not produce pure changes in field derivatives for compensation of the field derivatives of the primary and shielding magnets.
  • the ampere turn changes for each coil are made to adjust the sum of the field derivatives for the compensation magnet to "meet and compensate” as a whole the field derivatives of the primary and shielding magnets in combination.
  • the compensation of the field derivatives of the shielding and primary coils is improved as the relative strength of intermediate order field derivative for each coil of the compensation magnet may be changed.
  • Additional field derivatives may be included in the correction list by moving the electric centres. This allows for change to be made in the volume of the working region at the origin of the tubular bore of the system. For example, if the compensation magnet coil set is placed to coincide with the first zero value of the 8th order field derivative, the volume of the uniform zone suitable for MRI is at a maximum (all other lower even orders being also zero in sum between the magnets). In some experimental situations it may be desirable to reduce the field of view derived by reducing the volume of uniformity.
  • Figure 6 shows the major steps in the construction of a generic magnet system of the types described above, each of the coils being superconducting.
  • the principle is to construct the magnet in two halves, these being two halves upon either side of the mid-plane XY.
  • one set of primary coils 2 of the primary magnet are positioned horizontally upon a main support structure such that the their common axis points upwards.
  • the coils 3,4 of the shielding magnet are likewise positioned co-axially upon the main support.
  • a compensation support is added, to which coils 5,6,7 of the compensation magnet are attached.
  • the main and compensation supports are mounted together such that all coils have a common axis orientated upwards.
  • the first part of the cooling system is added, this being the helium system which provides helium coolant for each of the coils so that they may operate in superconducting mode.
  • the helium system is shown schematically as a bold outline in Figure 12, and is illustrated at 30. This is achieved by placing the magnet coils into a bottom half of a horizontal "helium can"; followed by placing a top half of the helium can on top of the bottom half. The two halves of the helium can are then closed by welding.
  • a liquid nitrogen cooling system 31 is added, this having an annular geometry analogous to that of the coils.
  • the liquid nitrogen system comprises a liquid nitrogen coolant reservoir which is positioned at a similar axial location, although radially outwardly of the compensation magnet coils (with respect to the common axis).
  • a “can” for liquid nitrogen for absorbing environmental heat at intermediate cryogenic temperatures
  • appropriate nitrogen temperature (and other lower intermediate temperature) shields are assembled to surround the helium can, to further improve thermal insulation.
  • a second instance of the system produced by step 203 is generated, this forming the second of the two halves of the system.
  • the two halves are then brought together such that they represent mirror images of one another across the mid-plane XY (requiring the inversion of the second).
  • the halves are appropriately located by a fiduciary plate on the mid-plane of the magnet system.
  • the two halves are encapsulated within a cryostat 32.
  • the two halves are mounted horizontally in the bottom half of the vacuum cryostat structure, and appropriately fixed by a support insulation structure.
  • the top half of the vacuum cryostat is subsequently mounted horizontally.
  • the two halves of the vacuum cryostat are joined by welding at the mid-plane. Thereafter the system is ready for additional support apparatus to be added and it will be understood that detailed electrical and cryogenic liquid handling conduits are installed as appropriate during steps 200 to 205.
  • Figure 13 shows a further diagram of the magnet system, similar to that of Figure 1 , although in this case with all four quadrants illustrated. It will be noted that the system is indeed significantly shorter than a more conventional tubular magnet illustrated at 11.
  • the location of gradient coils 60 is also illustrated.
  • the gradient coils are situated on the face of the magnet where the conic angle is largest.
  • the gradient coils arcs contributing the main gradient field components are located on a similar diameter to standard systems, but do not intrude into the bore of the short magnet. Thus similar gradient field strength to a standard system may be achieved without adverse impact on the view to the exam room afforded the patient.
  • the gradient coils arcs are provided with additional turns under separate control (using suitable apparatus) so as to provide odd order field derivatives, both in an axial direction and a transverse direction.
  • Normally only first order gradient fields are provided in scanners to reference the source of NMR signals in the imaging volume.
  • there are desirable uses that require the field of view to be limited to a small central volume to speed up the overall imaging process, and focus on a particular volume of interest. This is facilitated by adjustment of the values of high order field derivatives using the compensate magnet coil set on the larger conic angle (small diameter coils). This produces a focussing effect on the image field of view by reducing the volume of the uniform zone suitable for MRI.
  • a radio frequency (rf) receive coil may be able to receive from these alias regions. It is an intention of the present invention to apply imaging pulse gradients that are non- linear in the hollow sphere surrounding the imaging field of view such that alias regions are overcome and unwanted NMR resonances avoided.
  • non-linear pulse gradients may be used in this invention either alone, or in conjunction with rf "preparation pulses" which can augment the avoidance of alias effects when using a focused field of imaging view inside a large diameter rf coil.
  • the outer hollow spherical volume where the nmr resonance frequency generally has a value different to that of the central uniform zone (or "sweet spot"), may be prepared by a saturation rf pulse, before application of the spatial reference gradient pulses. Then, when the reference gradients are applied for localisation of the nmr signal, there is no magnetisation located outside the sweet spot which can be brought back into the rf band width detected for constructing image data.
  • Figure 14 shows an example use of the system in a medical procedure.
  • a human subject 50 is positioned upon a horizontal table and undergoes a head scan whilst the head of the subject is located within the working region of the magnet system 1.
  • the common axis of the magnet system is not horizontal, but angled with respect to the horizontal, such that the upper part of the system moves away from the feet of the subject. This, coupled with a recess in the floor where the medical practitioner or technician is standing (or alternatively by raising the scanner and subject), allows good access to the head of the subject.
  • FIGS. 15 and 16 are schematic perspective views of example systems which illustrate the extremely short axial nature of the magnets, allowing the much improved access to the subject whilst providing line-of-sight outside the magnet system for the subject.
  • the invention provides benefits in that the working region produced by the magnet system is similar in form to that produced by a conventional tubular magnet.
  • further discussion is provided relating to the mathematical theory underlying the invention, together with a consideration of the accuracy of the calculations used in designing the disclosed short tubular magnets.
  • the object is then to achieve field uniformity by arranging for as many of the derivatives as possible to be zero.
  • the design calculations assume uniform current density within a cylindrical region of conductor characterised by inner and outer radii, a x , a 2 , and the positions of the ends b ⁇ , b 2 .
  • the derivatives have been calculated using a general-purpose program which allows for non-coaxial elements of arbitrary orientation.
  • the method is to calculate In + 1 values of B over a length 2r 0 and to fit a Taylor series of nth order to these values.
  • the accuracy of the results can be influenced by the choice of n and r 0 and also by the precision of the calculation of B. For example. ro must not extend into the conductor, where there would be a discontinuity in the derivatives, nor should it be so small that the variation of the n th derivative is lost in the error in B.
  • the field values are calculated by integrating the contributions from elementary hoops over the conductor cross-sectional area.
  • the engineering tolerances are concerned with the positions of the coils, the dimensions of the coils and spatial variations of the winding densities. Having evaluated the possible errors arising from these illustrate how a strategy for correcting them is devised.
  • ⁇ , -(n + l)B n+ ⁇ ⁇ b
  • Figure 19 shows a misplaced thick coil (solid line) superimposed on the correct size and position (broken line). The regions contributing to the errors are confined to the hatched areas.
  • the lower order gradients are relatively easy to correct using "shim" coils.
  • First, second and third order correction coils can be made sufficiently powerful to correct those gradients in short magnets.
  • the size of the useful working volume is determined by the higher order gradients, such as the fourth, fifth, sixth and seventh, and these are much more difficult to correct. In the very short magnets under consideration these can be significant due to the practical working tolerances.
  • the four coils could be at
  • correction coils are close to, or even coincident with, the regions responsible for errors. Correction is achieved by choosing the currents in the correction coils, /, I 2 / 3 / 4 so as to cancel the errors from the misplacement of the main coil. There are four degrees of freedom so that a linear combination can always be found to cancel at least four orders of gradients.
  • the correct strategy is to correct the most significant high orders, for example 4th, 5th, 6th and 7th order gradients. Because the lower orders, although strong, are relatively less sensitive to the position of conductors, the discrepancy between the positions of the correction coils and the regions responsible for the errors is less significant than the higher orders. Consequently, fairly good correction is achieved for the lower orders, with imperfections being manageable by conventional shim coils.
  • the magnetic field vector, H can be defined in terms of a scalar potential V :
  • V ⁇ A lm p ⁇ Y,'"( ⁇ , ⁇ )
  • an arbitrary magnetic potential, or its magnetic field can be expressed as a series of spherical harmonics.
  • the first few terms of a series are generally all that is required. These are:

Abstract

L'invention concerne un système d'aimant pour utilisation dans l'imagerie par résonance magnétique. Pendant l'utilisation, le système génère un champ magnétique résultant, à l'intérieur d'une zone de travail, d'une homogénéité suffisante pour permettre à l'imagerie par résonance magnétique d'être exécutée. Le système comprend un aimant principal comprenant une paire de bobines positionnées sur un axe commun passant à travers le centre de la zone de travail, les bobines étant espacées symétriquement de part et d'autre d'une origine sur l'axe commun et ayant des dimensions telles que chaque bobine se trouve à une position nominale ou adjacente à cette position nominale dans laquelle la dérivée de second ordre d'un champ magnétique correspondant dans la zone de travail est sensiblement nulle pour une paire de bobines situées symétriquement de part et d'autre de l'origine à une telle position. Un aimant de protection comprend une paire de bobines positionnées sur l'axe commun, les bobines étant espacées symétriquement de part et d'autre de l'origine et ayant des dimensions telles que chaque bobine se trouve à une position nominale ou adjacente à cette position nominale dans laquelle la dérivée de second ordre d'un champ magnétique correspondant dans la zone de travail est sensiblement nulle pour une paire de bobines situées symétriquement de part et d'autre de l'origine à une telle position, l'aimant de protection servant à réduire le champ magnétique résultant à des emplacements distaux du système d'aimant. Un aimant de compensation comprend au moins une paire de bobines positionnées sur l'axe commun, les bobines étant espacées symétriquement de part et d'autre de l'origine et ayant des dimensions telles que chaque bobine se trouve à une position nominale ou adjacente à cette position nominale dans laquelle au moins une dérivée d'ordre pair d'un champ magnétique correspondant ayant une amplitude supérieure ou égale à 6, dans la zone de travail, est sensiblement nulle pour une paire de bobines situées symétriquement de part et d'autre de l'origine à une telle position.
PCT/GB2008/002182 2007-06-26 2008-06-25 Système d'aimant pour utilisation dans l'imagerie par résonance magnétique WO2009001084A1 (fr)

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WO2010123939A3 (fr) * 2009-04-20 2010-12-09 Time Medical Holdings Company Limited Réseau de bobine crâne rf supraconducteur refroidi de façon cryogénique et système d'imagerie par résonance magnétique (irm) pour tête uniquement l'utilisant
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CN102597794A (zh) * 2009-04-20 2012-07-18 美时医疗控股有限公司 低温冷却超导体rf头部线圈阵列和具有超导的头部专用mri系统
RU2570219C2 (ru) * 2009-04-20 2015-12-10 Тайм Медикал Холдингз Компани Лимитед Комплект сверхпроводящих рч-катушек с криогенным охлаждением для головы и система магнитно-резонансной томографии (мрт) только для головы, использующая такой комплект рч-катушек
CN102597794B (zh) * 2009-04-20 2016-08-10 美时医疗控股有限公司 低温冷却超导体rf头部线圈阵列和具有超导的头部专用mri系统

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