Classifications

• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
  • G01R33/385—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using gradient magnetic field coils

Definitions

• the following relates to the magnetic resonance imaging (MRI) arts. It has particular applicability to a magnet resonance imaging system having an open geometry and will be described with particular reference thereto.
• MRI magnetic resonance imaging
• MRI designers have been directing efforts towards open and patient-friendly geometries of the MRI system and, in particular, the main magnet.
• One of the challenges is that a more open geometry can come with a tradeoff in magnet performance and/or magnet cost.
• One type of MRI magnet uses a cylindrical shape, as illustrated in Figure 1.
• the magnetic field has a homogeneous volume, or examination region, 18, necessary for MRI imaging, which is often located in the center of the system along the z-axis of the cylinder.
• the z-axis of the cylindrical system is along the same direction as the main magnetic field.
• a patient lies in the cylinder at a position along the z-axis which places a region of interest of the patient within the homogeneous volume.
• a disadvantage is that many times the cylindrical system encloses the patient, leading to a patient unfriendly situation. Shortening the magnet length can improve the openness but it can also come with a penalty in cost price and/or magnetic field homogeneity.
• FIG. 2 An open magnet geometry is shown in Figure 2. Such a magnet includes two symmetric halves, an upper half and a lower half, with the patient positioned between the halves. In this geometry, assuming the patient is lying down, the magnetic field is in the vertical direction, hence the term "vertical field magnet” for this type of magnet.
• Another open geometry though relatively inefficient, is that of a so-called “projection magnet”. It includes a single-sided magnet, with a homogenous volume on the z-axis, in front of the magnet, projected outside the magnet as shown in Figure 3. If such a magnet were hidden behind a wall, the homogenous volume would be floating in the room without giving the patient the impression of being enclosed by a magnet. An extremely friendly "walk-up magnet” is the result. However, such projection magnets have not achieved much success due to problems with inefficiency.
• Figure 4 shows a projection magnet with a single supplementary coil 22. It leads to an asymmetric magnet design as shown in Figure 4. The efficiency and homogeneity are significantly improved over the pure projection magnet shown in Figure 3.
• the disadvantages of a projection magnet with a supplementary coil can be explained as follows.
• the supplementary coil 22 at the small pole (right side) have a current density comparable to that of the large pole (left side), or even higher. So, if the large pole includes a superconductive magnet, then the small pole also includes a superconductive magnet. Therefore, as shown in Figure 4 the main coils 20 are disposed within a first cryostat 30 and the supplementary coils 22 are disposed within a second cryostat 32.
• the cryostat at the small pole adds cost, reduces the openness, and reduces the gap between the poles, thereby reducing the space available to accommodate a patient.
• Other considerations when designing an MRI system relate to gradient coils. Shielded gradient coils are known for both cylindrical magnets as well as for open magnets. In Figures 5 and 6 shielded gradient coils for cylindrical and open systems, respectively, are shown. In both of the cases a magnetic field gradient within the imaging volume is created by a primary set of windings. In the case of a cylindrical system as shown in Figure 5, the primary set of windings are represented by a first, or inner, cylinder 50.
• the primary set of windings are represented by upper and lower inner surfaces 60, 61.
• the shielding is achieved by using a secondary set of windings, separated from the primary set by a certain distance, which generate a magnetic field that substantially cancels the magnetic fringe field generated by the first set of windings.
• the secondary set of windings are represented by a second, or outer, cylinder 52.
• the secondary windings are represented by upper and lower outer surfaces 62, 63.
• plane A contains the primary windings
• plane B contains the secondary windings
• rim C contains the cross-overs.
• This type of design is also used on cylindrical gradient coils in order to combine short length with good efficiency.
• the aforementioned shielded gradient coils have disadvantages.
• both sides of the gradient coil are actively shielded.
• Active shielding of a gradient coil typically requires extra space needed in comparison to an unshielded gradient coil.
• Active shielding of both sides is not necessarily needed for certain special open MRI systems.
• configurations where active shielding is not necessary for both sides for example because of the use of a non-superconducting side, the use of shielded gradient coils will lead to inefficient designs.
• the energy stored in that type of configuration is higher than needed, and space required for the gradient coil is more than necessary, and the openness is reduced.
• a gradient coil assembly for use in magnetic resonance imaging.
• the gradient coil assembly includes a first gradient coil portion and a second gradient coil portion disposed from the first gradient coil portion along a z-axis.
• the first gradient coil portion being shielded and the second gradient coil portion being unshielded.
• a gradient coil assembly for use in magnetic resonance imaging.
• the gradient coil includes a shielded gradient coil portion disposed in a first plane; and an unshielded gradient coil portion disposed in a second plane, the second plane being substantially parallel to the first plane and the shielded gradient portion acting in combination with the unshielded gradient coil portion to generate a first gradient magnetic field.
• the invention may take form in various components and arrangements of components, and in various process operations and arrangements of process operations.
• the drawings are only for the purpose of illustrating preferred embodiments and are not to be construed as limiting the invention.
• Figure 1 shows an illustration of a cylindrical magnetic resonance imaging system.
• Figure 2 shows an illustration of an open magnetic resonance imaging system.
• Figure 3 shows an illustration of a projection magnet.
• Figure 4 is an illustration of an asymmetric open magnet.
• Figure 5 shows an illustration of cylindrical shielded gradient coil.
• Figure 6 is an illustration of a shielded gradient coil for an open MR system.
• Figure 7 is an illustration of one side of a shielded gradient coil with cross-overs for an open MR system.
• Figure 8 is an illustration of an asymmetric open system having superconducting coils at the large pole and non-superconducting coils at the small pole.
• Figure 9 shows an illustration of a design strategy for designing asymmetric main magnets for an MR system.
• Figure 10 shows an asymmetric open system using normal coils at the small pole in which the homogeneous volume is centered and having a patient in a sitting position whose head is being scanned.
• Figure 11 shows an asymmetric open system, using normal coils at the small pole, in which the homogeneous volume is near the normal pole (right side).
• Figure 12 shows an asymmetric open system, using ferro-magnetic rings (iron) at the small pole.
• Figure 13 shows a cross section of a coil element.
• Figure 14 shows an open gradient coil which is shielded on only one side.
• Figure 15 shows an exploded view of a gradient coil.
• Figure 16 shows windings of a gradient coil which is shielded on one side.
• Figure 17 shows a view of the windings in the gradient coil of Figure 16 showing both sets of primary windings and the secondary windings of the shielded part.
• Figure 18 shows an overview of the fields generated by the gradient coils.
• Figure 19 shows windings of the Baseline and Correction coils.
• Figure 8 shows an asymmetric main magnet system of a magnetic resonance imaging apparatus 100.
• the asymmetric main magnet system includes a first main magnet portion 120 and a second main magnet portion 130 disposed away from the first main magnet portion along the z-axis.
• the first main magnet portion, or large pole includes a number of coils 1-9, 13 disposed within a cryostat 125.
• the second main magnet portion, or small pole includes a number of coils 10-12.
• the small pole is not disposed within a cryostat.
• the total diameter of the small pole is substantially smaller than that of the large pole. In one embodiment, the ratio of diameters is less than 0.7.
• the embodiment shown in Figure 8 includes a superconducting large pole in combination with a non- superconducting small pole.
• the two poles are not only different in size, but also different in applied technology.
• the field contributions generated by the small and the big pole are manipulated. For example, during the design process, for a given target field (homogeneity and fringe field) current flow in the big pole can be encouraged, and currents on the small pole can be discouraged. By doing this, a situation where most of the field is generated by the large pole (necessitating superconductivity), while the small pole provides significantly less field (achievable with non-superconducting technology) can be achieved.
• FIG. 9 First the areas at the large and small pole where current is in principle allowed to flow are set. This is illustrated in Figure 9. As shown in Figure 9, there are first 251, second 252, and third 253 areas in the large pole where current is allowed to flow. With respect to the small pole shown in this embodiment, there is one area, or fourth area 254, in which current is allowed to flow. By doing this, the dimensions of the large and small poles respectively are defined. As can also be seen in Figure 9, whether the large pole should define a recess 260 or other geometrical configuration can be established. For example, a recess can be established to accommodate a gradient coil, radio-frequency coil, or for any other reason.
• the current-carrying areas are then meshed in small elements.
• the field targets are also defined.
• B z 1000000 ⁇ 10 ⁇ T.
• To obtain active shielding, several fringe field points 210 in the external region around the magnet are selected and the fringe field is prescribed with certain tolerance, e.g. B 0 ⁇ 500 ⁇ T.
• a suitable cost function is for example the total dissipation in the magnet system, where sections in the normal (non-superconducting) conducting area, e.g. the small pole, are given an ⁇ times higher resistivity than those in the superconducting regions e.g. the large pole.
• ⁇ has been chosen to be 250, but other values are applicable.
• the result is a homogeneous magnet design having currents in many small mesh elements. Due to the design approach, the current density in the normal conducting side is substantially smaller than in the superconducting side. If the current density in the normal conducting region is still too high the dimensions of this side can either be increased, or the value of ⁇ can be increased and the solver can be run again.
• the next step is to group clusters of current elements with equal current direction into discrete rectangular coils that can be practically wound.
• dimensions, positions, and current densities of the new coils are chosen such that the target fields are preserved. In the examples a total of 13 to 15 coils is reached. Tiny coils can often be eliminated to reduce the number of coils (i.e. less complex magnet) without giving up too much of the target field.
• normal conducting coils can replaced by ring-shaped elements of ferro-magnetic or permanently magnetic materials. This requires shift in positions and dimensions in order to preserve the target field.
• Figures 8, 10, and 11 show how the homogeneous volume can be chosen near one of the magnet sides or simply in the middle, the optimum choice depending on the intended clinical application. These three designs, which employ coils in the main magnet portions, are shown in more detail in Table 1. As shown in Figure 13 the values for r correspond to the radial distance from the z-axis to the individual coil element, the values for w correspond to the width of the coil element, and the values for t correspond to the thickness of the coil elements. As can be seen, the current density in the superconducting coils (100- 120 A/mm 2 ) is much higher than the current density in the normal conducting coils (9-24 A A/mm 2 ). Design 1:
• Table 1 Values relating to the designs of Figures 8 (Design 1), 10 (Design 2), and 11 (Design 3).
• Figure 12 shows a main magnet where the coil elements 9, 10 are ferro-magnetic (iron) rings rather than the normal conducting coils discussed above.
• Figure 14 shows an embodiment of a gradient coil assembly 500 which leaves one side unshielded.
• the gradient coil assembly includes a first, or baseline, gradient coil portion 510 and a second, or correction, gradient coil portion 520.
• the first and second gradient coil portions are disposed apart from each other along the z-axis.
• the first gradient coil generates most of the gradient field and acts in combination with the second gradient coil. Also as shown in Figure 14, the first gradient coil is shielded.
• the gradient coil assembly is build up out of two coaxial planes (planes A) containing primary sets of winding that generate the main gradient field.
• a secondary (shield) plane (plane B) is added coaxial to the two primary planes, separated by a distance from one of the primary planes.
• the embodiment shown in Figure 15 also includes a connecting rim C, on the shielded side. This rim connects the primary plane with the secondary plane, and contains the windings that cross over between planes A and B.
• the outer plane represents the conducting outer surface of a magnet cryostat. During the design a limitation is set on the amplitude of eddy currents that are induced in this extra plane, or on the fields that are generated by these eddy currents.
• Figure 17 shows the discretised windings in the primary planes and the shielding plane for the example.
• the homogeneous area created by the gradient coil can be placed at an arbitrary position, on the axis, in the gap between the two primary planes. It is not necessary to have the same position as in the example.
• a 10mT/m gradient field is created with a 360A operating current.
• a gradient coil assembly with a switchable homogeneous area is provided.
• This gradient coil is build up out of two gradient coils described above.
• the baseline coil carries a larger part of the current to generate the gradient field (>60%) and the correction coil generates a smaller part of the current to generate the gradient field ( ⁇ 40%).
• Baseline and “Correction” are used to denote the current patterns of the Baseline and Correction coil respectively.
• BC and CC are obtained by adding and subtracting the current patterns as follows:
• BC and CC can be designed in such a way that they generate the desired magnetic field. BC would then make the average field of A and B: %(A+B).
• the Correction coil Because the Correction coil generates far less magnetic field than the Baseline coil it has a very low amount of stored energy, 0.074J vs. 0.843 J. Therefore by using the
• Figure 18 shows the magnetic fields generated by the Baseline coil, Correction coil and both coils A and B.
• the windings of the primary planes of the Baseline and Correction coil are displayed in Figure 19. To generate 10mT/m one has to operate the gradient coils at 170A.
• the Correction coil contains far less windings and is less complex.

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  • 2006-06-29 EP EP06765955A patent/EP1924867A2/de not_active Withdrawn
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