WO2007004109A1

WO2007004109A1 - Asymmetric magnet for mri

Info

Publication number: WO2007004109A1
Application number: PCT/IB2006/052084
Authority: WO
Inventors: Gert B. J. Mulder; Robert-Paul Buitendijk
Original assignee: Koninklijke Philips Electronics NV; US Philips Corp
Current assignee: Koninklijke Philips NV; US Philips Corp
Priority date: 2005-06-30
Filing date: 2006-06-26
Publication date: 2007-01-11
Anticipated expiration: 2007-12-30

Abstract

An MRI apparatus is provided. The appartaus includes a main magnet system having a first main magnet portion (120) and a second main magnet portion (130) disposed away from the first main magnet portion along a z-axis. The first main magnet portion is superconductive and the second main magnet portion is non-superconductive.

Description

ASYMMETRIC MAGNET FOR MRI

DESCRIPTION

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 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. As shown in Figure 1, the z-axis of the cylindrical system is along the same direction as the main magnetic field. As a result, 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. 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. In the case of an open system, as shown in Figure 6, 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. In the case of the cylindrical system shown in Figure 5, the secondary set of windings are represented by a second, or outer, cylinder 51. In the case of the open system shown in Figure 6, the secondary windings are represented by upper and lower outer surfaces 62, 63.

An important design consideration for gradient coils is the stored energy in the coil. A small amount of stored energy allows fast on and off switching of the gradient resulting in a very efficient gradient coil suitable for fast MR imaging. In general shielding the gradient coil will lead to higher stored energy and therefore a decreased efficiency. Although active shielding of the gradient coil is less efficient it is generally necessary when there is a cryostat adjacent to the gradient coil in order to reduce eddy currents in the cryostat to an acceptable level. In another design it is possible for the windings to cross over between the inner and outer cylinders, in the case of a cylindrical system, or surfaces, in the case of an open system. These cross-overs can increases the efficiency of the gradient coils. Depending on the efficiency, without cross-overs, it may or may not be necessary to allow the windings to cross over.

In Figure 7, plane A contains the primary windings, plane B contains the secondary windings and rim C contains the cross-overs. Depending on the diameter of the gradient coils this increases the efficiency. 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. In open systems, both sides of the gradient coil are actively shielded. Active shielding of a gradient coil 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. In 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 will be higher than needed, and space required for the gradient coil is more than necessary, and the openness is reduced.

In accordance with one aspect of an embodiment of the invention, an MRI apparatus is provided. The apparatus includes a main magnet system having a first main magnet portion and a second main magnet portion disposed away from the first main magnet portion along a z-axis. The first main magnet portion is superconductive and the second main magnet portion is non-superconductive.

In accordance with another aspect of an embodiment of the present invention, a magnetic resonance main magnet having a main field direction B₀, is provided. The magnet includes a large pole portion having a first diameter; a small pole portion disposed away from the large pole along a z-axis in the main field direction and having a second diameter, the large and small pole portions defining an examination region. The ratio between the first and second diameters is less than 0.7.

One advantage of an embodiment of the present invention is that the openness of a magnetic resonance imaging apparatus is facilitated.

Another advantage is that a reduction in cost is facilitated.

Another advantage is that an increase in magnet performance is facilitated. Another advantage is that increased space for a patient is facilitated.

Numerous additional advantages and benefits will become apparent to those of ordinary skill in the art upon reading the following detailed description of the preferred embodiments.

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. With reference to Figure 8, an example of an embodiment of the present invention is shown. More specifically, 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. In the embodiment shown, 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. However, the small pole is not disposed within a cryostat. Further, 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.

Accordingly, the embodiment shown in Figure 8 includes a superconducting large pole in combination with a non- superconducting small pole. In other words, the two poles are not only different in size, but also different in applied technology. In designing such a main magnet system, multiple coils on the small pole are allowed. The use of current on the small pole, however, is discouraged as much as possible. This is done by means of a cost function of the design solver. For example, the dissipation in the entire magnet system is minimized while all coils of the small pole are given a much higher resistivity (e.g. factor 100) than the coils in the large pole. This approach results in a design where most of the field is generated by the big pole, using a relatively large number of ampere-turns, while the small pole becomes a set of coils carrying relatively few ampere-turns, which can be realized with resistive coils or magnetic rings. In such asymmetric geometries 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. The strategy for designing such an asymmetric magnetic resonance imaging system is described in more detail as follows. 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. On the surface of the intended homogeneous volume 200 a large number of points, e.g. 50, are selected in which the magnetic field is pre-scribed with a certain tolerance, e.g. B_z = 1000000 ± 10 μT. 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 μT.

Next the necessary currents in the mesh are determined using a linear solver, which solves the currents while satisfying the target field constraints and simultaneously minimizing a cost function. 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 α times higher resistivity than those in the superconducting regions e.g. the large pole. The larger the factor α, the more current is pushed from the normal conducting side towards the superconducting side. In the examples α 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 α 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. During this process, 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. Finally, if needed, 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.

The above design strategy leads to designs as illustrated in Figures 8 and 10-12. These magnets all have the following features in common: - 1.0 T central field;

- field homogeneity of ± 10 ppm on a sphere of 20 cm diameter;

- expected net patient gap 35 cm;

- diameter of the small pole ranging from 0.45 to 0.60 m; - approximate outer dimensions of big pole arel .8 m diameter and 0.7 m width;

- 5 cm deep recess in the big pole for, for example, a gradient coil; and

- actively shielded, 0.5 mT contour smaller than RxZ=6x7 m.

It is to be understood that various alterations to the above parameters are contemplated. The figures show a variety of clinical applications possible with this type of asymmetric magnet. In most of the cases, the large pole can be hidden behind a wall, in the ceiling, or in the floor, in order to improve the perception of openness.

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/mm²) is much higher than the current density in the normal conducting coils (9-24 A A/mm²). Design 1:

# z[m] r[m] w [m] t[m] J[A/mm2]

1 -0.250000 0.006856 0.019833 0.035401 -114.0390

2 -0.250000 0.045603 0.025667 0.033443 114.0390

3 -0.250000 0.080445 0.027483 0.032442 -114.0390

4 -0.250000 0.132073 0.030221 0.041746 114.0390

5 -0.250000 0.200522 0.037466 0.046324 -114.0390

6 -0.250000 0.285693 0.057945 0.057945 114.0390

7 -0.170000 0.342044 0.021285 0.041325 114.0390

8 -0.170000 0.405800 0.069891 0.088924 -114.0390

9 -0.170000 0.628364 0.103344 0.206687 114.0390

10 0.300000 0.009956 0.072915 0.074729 9.1207

11 0.300000 0.106653 0.065498 0.056657 -9.1207

12 0.300000 0.178114 0.086429 0.151880 9.1207

13 -0.700000 0.674464 0.089098 0.178197 -114.0390

Design 2:

# z[m] r[m] w [m] t[m] J[A/mm2]

1 -0.325000 0.002825 0.044798 0.071180 -119.9570

2 -0.325000 0.084148 0.041736 0.057811 119.9570

3 -0.325000 0.164410 0.043147 0.061437 -119.9570

4 -0.325000 0.266472 0.058382 0.073209 119.9570

5 -0.245000 0.346340 0.028114 0.028114 119.9570

6 -0.245000 0.401760 0.074918 0.086880 -119.9570

7 -0.245000 0.630751 0.102570 0.205141 119.9570

8 0.225000 0.012281 0.020570 0.027684 23.9929

9 0.225000 0.054820 0.014331 0.026502 -23.9929

10 0.225000 0.099509 0.027073 0.034952 23.9929

11 0.225000 0.161785 0.028467 0.035194 -23.9929

12 0.225000 0.221281 0.046841 0.095310 23.9929

13 -0.775000 0.668719 0.089610 0.179220 -119.9570

Design 3:

# z[m] r[m] w [m] t[m] J[A/mm2]

1 -0.400000 0.009633 0.042581 0.085982 117.5530

2 -0.400000 0.118450 0.039519 0.070578 -117.5530

3 -0.400000 0.243919 0.058403 0.081064 117.5530

4 -0.320000 0.389294 0.066842 0.084562 -117.5530

5 -0.320000 0.629736 0.102262 0.204525 117.5530

6 0.150000 0.009038 0.011183 0.011183 23.5120

7 0.150000 0.020493 0.011026 0.011026 -23.5120

8 0.150000 0.037331 0.013008 0.013008 23.5120

9 0.150000 0.057466 0.013652 0.012056 -23.5120

10 0.150000 0.072347 0.015006 0.026391 23.5120

11 0.150000 0.102543 0.017798 0.017986 -23.5120

12 0.150000 0.131306 0.030358 0.030698 23.5120

13 0.150000 0.180185 0.033878 0.031838 -23.5120

14 0.150000 0.227283 0.048844 0.090898 23.5120

15 -0.850000 0.663642 0.090061 0.180122 -117.5530

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.

The invention has been described with reference to the preferred embodiments. Obviously, modifications and alterations will occur to others upon reading and understanding the preceding detailed description. It is intended that the invention be construed as including all such modifications and alterations insofar as they come within the scope of the appended claims or the equivalents thereof.

Claims

1. An MRI apparatus comprising: a main magnet system having a first main magnet portion (120) and a second main magnet portion (130) disposed away from the first main magnet portion along a z-axis, wherein the first main magnet portion is superconductive and the second main magnet portion is non-superconductive.

2. The MRI apparatus as set forth in claim 1 wherein the second main magnet portion has a substantially smaller diameter than the first main magnet portion.

3. The MRI apparatus as set forth in claim 2 where in a ratio of the diameters of of the first and second main magnet portions is less than 0.7.

4. The MRI apparatus as set forth in claim 1 wherein the second main magnet portion includes conductive coils.

5. The MRI apparatus as set forth in claim 1 wherein the second main magnet portion includes ferro-magnetic rings.

6. The MRI apparatus as set forth in claim 1 wherein the first main magnet portion includes coil elements disposed within a cryostat.

7. The MRI apparatus as set forth in claim 6 wherein the second main magnet portion are not disposed within a cryostat.

8. The MRI apparatus as set forth in claim 1 wherein the main magnet portion includes at least two areas of coil windings (252, 253) which define a recess (260) within the first main magnet portion.

9. A magnetic resonance main magnet having a main field direction B₀, the magnet comprising: a large pole portion having a first diameter; a small pole portion disposed away from the large pole along a z-axis in the main field direction and having a second diameter, the large and small pole portions defining an examination region (18) wherein the ratio between the first and second diameters is less than 0.7.

10. The magnetic resonance main magnet as set forth in claim 9 wherein the large pole portion includes superconductive coil elements and the small pole portion is non- superconductive.

11. The magnetic resonance main magnet as set forth in claim 10 wherein the small pole portion includes a conductive magnet.

12. The magnetic resonance main magnet as set forth in claim 10 wherein the small pole portion includes a ferro-magnetic magnet.

13. The magnetic resonance main magnet as set forth in claim 10 wherein the magnetic elements of the small pole portion includes a conductive magnet.