Coil assembly for a magnetic resonance examination system having unshielded gradient coils

Abstract

The present invention relates to the field of magnet coil assemblies for magnetic resonance examination systems, which aim to minimize voltages induced in each winding of a superconducting magnet 2 surrounding a gradient system. This is achieved by providing a magnet coil assembly 1 having an unshielded gradient coil 3 having two separate sections 4, 5, including an imaging section 4 around an examination zone of the main magnet and at least an auxiliary section 5 offset from the imaging section 4 in a radial or axial direction of the main magnet 2 and adjacent to the imaging section, the imaging section 4 and the auxiliary section 5 having coil windings 7, 8 disposed on the surfaces of the imaging section 4 and the auxiliary section 5, and along a boundary 10 between the imaging section 4 and the auxiliary section 5.

Description

[Technical Field]

[0001] The present invention relates to the field of magnet coil assemblies for magnetic resonance examination systems, and more particularly to magnet coil assemblies having unshielded gradient coils. [Background technology]

[0002] Unshielded gradient coils are attractive because they offer better gradient efficiency and, at the same time, more patient space within a given magnet bore. Unshielded z-gradients are problematic because they inductively couple with the magnet, potentially inducing extremely large voltages. It would be highly advantageous to design the magnet / gradient coil so that the gradient coil's external magnetic field penetrates the windings of the superconducting magnet. If the gradient coil does not require an active shielding layer, the magnet warm bore diameter can be sized slightly larger than the gradient coil's outer diameter. This can be used to reduce the magnet size for a given patient bore size or to increase the patient bore diameter for a given magnet inner diameter. In all cases, the efficiency of the gradient coil is significantly increased compared to standard actively shielded gradient coils. Summary of the Invention [Problem to be solved by the invention]

[0003] A prerequisite for systems with unshielded gradient coils tightly coupled to the magnet is that all cryostat shells immediately outside the gradient coils must be transparent to the changing magnetic field. This implies a non-conductive magnet bore tube (a very thin metal layer for vacuum sealing is acceptable as long as the electrical time constant is much shorter than the typical switching time of the gradients). The radiation shield must be constructed of strips or wires to remove radiated heat without creating conductive patches that would block the gradient field, and can also be combined with convection cooling loops or heat pipes. Conductive loops in the interior regions of the cooling body must also be avoided.

[0004] Exposing magnet windings to the gradient stray fields presents a number of problems. First, the alternating magnetic field experienced by the conductors results in eddy current losses, which must be kept low enough to avoid overloading the cooling system. This heating problem exists in principle equally for all three gradient axes. A second problem, which concerns only the z-gradient axis, is that the gradient alternating magnetic field induces large voltages in magnet coils with many windings. Transverse coils do not induce a net magnetic flux through the magnet windings and therefore do not result in significantly large induced voltages (if the system is constructed symmetrically).

[0005] The ISMRM abstract "Concurrent use of 4 gradient axes enables eddy current compensation of an unshielded gradient insert coil" by TA van der Velden et al. in Proc. Intl. Soc. Mag. Reson. Med. 25 (2017)4328 discloses using a whole-body gradient coil to drive pre-emphasis currents to compensate for eddy currents induced by an unshielded gradient insert coil.

[0006] The present invention aims to generate a homogeneous static magnetic field in an examination zone with an unshielded z-gradient coil. [Means for solving the problem]

[0007] According to the present invention, this object is addressed by the subject matter of the independent claims. Preferred embodiments of the invention are set forth in the dependent claims.

[0008] Therefore, according to the present invention, there is provided a magnet coil assembly for a magnetic resonance examination system, comprising a main magnet having superconducting windings for generating a uniform static main magnetic field within an examination zone, and unshielded gradient coils for generating gradient magnetic field waveforms, the unshielded gradient coil having two separate sections: an imaging section around the examination zone of the main magnet, and at least an auxiliary section offset from the imaging section in a radial or axial direction of the main magnet and adjacent to the imaging section, the imaging section and the auxiliary section having coil windings, the coil windings being arranged on surfaces of the imaging section and the auxiliary section, and each being arranged at least partially along the boundary between the imaging section and the auxiliary section.

[0009] The voltages induced in sections of a superconducting magnet by single-layer gradient coils can be reduced to acceptable levels by several methods. Active axial shielding can effectively decouple the end sections of the magnet. Axially offset sections are formed in the longitudinal end regions of the gradient coil, with gradient field windings running in opposite directions to actively shield the magnetic field from the central winding at the end sections of the magnet. These axially offset sections can be arranged in an axial region extending from the axial ends over one-quarter of the magnet's length. Coupling with the central magnet section can be avoided by adapting the windings of these sections layer by layer repeatedly across the magnet's central plane. The gradient coil winding pattern is modeled by a continuous stream function on the surface where conductors are allowed to be placed. Given appropriate target values ​​for the magnetic field within the imaging volume, the external magnetic field, and optionally other constraints such as maximum current density or forces in the main magnet's magnetic field, a quadratic programming optimization algorithm generates the optimal stream function distribution. The stream function contours are then taken as 3D windings. This modeling approach for gradient coil windings is known per se from WO 2018 / 033548. The mutual inductance between the gradient windings and the magnet coils is reduced by reducing the net magnetic flux of the gradient windings in the magnet coil windings. Another way to maximize gradient efficiency while essentially isolating the gradient coils from the external magnet is to divide the gradient space into two compartments and place the windings around them, thereby simultaneously achieving a good imaging field of view and net magnetic flux compensation.

[0010] Thus, the transformer formed by the main magnet windings and the unshielded z-gradient coil are weakly coupled or completely separated. The present invention avoids high induced voltages in the superconducting windings of the main magnet. The unshielded z-gradient coil does not require a surrounding shielding layer. This allows for a wider bore size, e.g., 80 cm. Note that the saddle coil configuration of the transverse gradient coil inherently has low magnetic flux coupling with the main magnet windings, as long as the main magnet windings are asymmetric relative to the magnet. The key element for separating the single-layer gradient coil from the complete magnet is dividing the space inside the coil into two sections and placing the windings on the boundary between these two sections.

[0011] In a technically advantageous embodiment of the magnet coil assembly, the coil windings of the auxiliary section carry the same current as the coil windings of the imaging section.

[0012] In another technically advantageous embodiment of the magnet coil assembly, the coil windings of the auxiliary section and the coil windings of the imaging section are connected in series, whereby the series connection of the windings makes it easy to achieve that the same current flows through the windings.

[0013] In a technically advantageous embodiment of the magnet coil assembly, the imaging compartment surrounds the examination space of the magnetic resonance examination system and the auxiliary compartment is the remaining part of the cylindrical coil volume below the patient table of the magnetic resonance examination system.

[0014] In another technically advantageous embodiment of the unshielded gradient coil, the unshielded gradient coil is a longitudinal z-gradient coil or an asymmetric superior-inferior gradient coil.

[0015] In an advantageous embodiment of the invention, the boundary of the imaging compartment adjacent to the auxiliary compartment is located in the range of 100-200 mm below the isocenter of the main magnet.

[0016] The present invention also provides a magnetic resonance examination system, which comprises a magnet coil assembly as described above.

[0017] These and other aspects of the invention will be apparent from and elucidated with reference to the embodiments described hereinafter, but such embodiments do not necessarily represent the full scope of the invention, and reference should therefore be made to the claims and this specification for interpreting the scope of the invention. [Brief explanation of the drawings]

[0018] [Figure 1] FIG. 1 shows an unshielded Z-gradient coil with closely fitted magnet coils and its magnetic field map. [Figure 2] Diagram showing the induced voltage per magnet turn as a function of axial position. [Figure 3] FIG. 1 is a schematic diagram of a single-layer Z-gradient coil with axial shielding. [Figure 4] Diagram showing the induced voltage per turn of an axially shielded gradient coil. [Figure 5] Illustrates the Z gradient field in the inner layer of the winding. [Figure 6] FIG. 10 illustrates a magnet design with a long central solenoid section. [Figure 7] FIG. 2 is a schematic cross-sectional view of a two-section gradient coil according to an embodiment of the present invention. [Figure 8a] Figure 1 shows the magnetic field map of a two-compartment inductively separated Z-gradient design. [Figure 8b] FIG. 2 shows individual coil sections according to one embodiment of the present invention. [Figure 8c] FIG. 2 shows individual coil sections according to one embodiment of the present invention. [Figure 8d] FIG. 2 shows individual coil sections according to one embodiment of the present invention. [Figure 9] FIG. 10 shows the induced voltage per turn of a magnet at 10 mT / m and 1 kHz according to an embodiment of the present invention. [Figure 10] FIG. 10 is a diagram illustrating the magnitude of the Z-gradient magnetic field in the inner layer of the magnet winding according to one embodiment of the present invention. [Figure 11]FIG. 10 shows the magnetic field map of an asymmetric Y-coil around the upper gradient section. [Figure 12a] 10 shows the magnetic field map of an asymmetrical up-down gradient with a decoupling coil in the lower section. [Figure 12b] FIG. 2 shows individual coil sections according to one embodiment of the present invention. [Figure 12c] FIG. 2 shows individual coil sections according to one embodiment of the present invention. [Figure 12d] FIG. 2 shows individual coil sections according to one embodiment of the present invention. DETAILED DESCRIPTION OF THE INVENTION

[0019] Below, a typical example of an unshielded z-gradient coil 3, as shown in Figure 1, is analyzed to obtain an estimate of the induced voltage problem to be solved. The exemplary coil 3 is assumed to have a diameter of 800 mm and to be combined with a magnet winding 6 having an inner diameter of 840 mm. In this example, the gradient coil windings 6 can be distributed over a coil length of 1500 mm, and the magnet windings can also be arranged over this entire length. If there are no external magnetic field constraints and a good imaging field of view is required within an ellipsoid with a diameter of 500 mm and a length of 450 mm, the optimal winding distribution would be as shown in Figure 1. Here, and in all the following gradient coil examples, the windings are shown as closed contours. It should be clear to those skilled in the art that these windings are all connected in series and that the same current flows through all windings. The arrows indicate the magnetic field generated by the gradient coil 3. It is clear from the diagram that this field results in a significant net magnetic flux through each magnet winding. For this coil, the magnetic flux through a single turn of the magnet was calculated as a function of the turn's axial position, with a gradient amplitude of 10 mT / m. Multiplying this flux by 2π.f yields the peak voltage induced in that turn by a sinusoidal gradient of frequency f. Figure 2 shows the simulation results. The simulation shows that for a 100-turn magnet section, the induced voltage exceeds 1 kV. Magnet coils with several hundred turns are at risk of electrical breakdown. An additional problem is that the frequencies at which magnet sections exhibit self-resonances are within the frequency range of the gradient system, and if such resonances are excited at very high amplitudes, this could very well lead to a quench.

[0020] The reduction of coupling with the z-gradient coil is achieved by axially or radially offsetting sections without adding a larger diameter shield layer, so that adjacent sections have opposite current directions, i.e., adjacent sections have windings running in opposite directions. This reduces coupling between the z-gradient coil and the magnet coils by actively shielding in the axial direction with the shield windings on the primary gradient diameter that form the axially offset gradient coil section, and by compensating for the mutual inductance between the gradient coils of the two radially offset sections and the complete magnet.

[0021] Mutual inductance can be further reduced by strict layer-by-layer windings due to voltage cancellation within the magnet windings. This winding pattern can be optimized to reduce mutual inductance between the gradient coils and the magnet coils by modeling it with a continuous stream function.

[0022] These approaches are described in more detail in the next section.

[0023] The coil 3 shown in FIG. 1 has a very low winding density 6 in regions far from the central plane. This is advantageous if the goal is to obtain the most efficient gradient coil. However, considering that the efficiency of Z-gradient coils is typically much greater than that of transverse gradient coils, it is acceptable to deviate from this optimal shape if it solves other problems. If the gradient coil length is constrained to one meter or less, it is still possible to generate a good imaging field, thereby reducing coupling with the magnet windings in the outermost sections. This reduction in gradient coil length is a good approach, as these are the sections with the most turns. However, one step can be taken further. In this case, the empty end regions of the gradient coil 3 can be used to place reverse-running gradient windings 6, forming axially offset sections that actively shield the magnetic field from the central winding at the magnet's end sections. FIG. 3 shows an example of such an axially shielded single-layer coil 3, providing the same imaging field as the unconstrained coil of FIG. 1.

[0024] Constraining the coupling with the magnet's end coils increases the stored energy at 10 mT / m from 3.7 J to 6.7 J, which is significantly lower than the stored energy of actively shielded gradient coils of the same inner diameter. Figure 4 shows the voltage induced in a single turn of the tightly packed magnet for this axially shielded z-gradient. The graph shows that all magnet windings axially beyond 550 mm from the midplane are very effectively isolated from the gradient field, and no significant induced voltages are expected. It is also unlikely that special measures to avoid self-resonance in these end coils will be necessary when using this axially shielded approach.

[0025] The return flux of gradient coil 3 is forced out of the coil cylinder over a small area, resulting in an increased flux density. This results in an increase in the amplitude of the gradient field towards the magnet's central windings, as shown in Figure 5. At the end coils of magnet 2, the gradient field amplitude is significantly reduced due to the active axial shielding. These magnet end coils have the largest amount of windings and the smallest temperature margin.

[0026] In the next section we will explain how improvements can be made in the central coil section.

[0027] Assuming that the induced voltage in the windings in the region of peak gradient coupling is on the order of 10 V per turn and that the ground potential at the windings should be limited to 1–2 kV, some measure is required if the central section of the coil exceeds 100–200 turns. This is typically the case in medium- and high-field MRI magnets. The solution to this problem is to cross the mid-plane for every 100–200 turns in the central section of the magnet and continue the coil circuit with a corresponding 100–200 turns in the other half of the magnet. By repeating this process, the net voltage stored in such layers spanning multiple coil sections is zero. Magnet designs with an odd number of coil sections (one of which is mid-plane) are advantageous here because these central sections are automatically inductively isolated. Figure 6 shows an example of such a design in a 0.6 T field.

[0028] The end coils of this magnet have a fairly small aspect ratio to keep all windings in the low-field region of the axially shielded z-gradient coil. The central solenoid is 800 mm long. Even if wound with a pitch on the order of 1 mm, the accumulated voltage per half-layer would still be too large. This must be avoided by increasing the winding pitch. In the design shown in Figure 6, the long solenoid has four layers, each with only 290 turns (i.e., the gradient voltage accumulation reverses every 145 turns). This number of turns implies a turn pitch of 2.75 mm. This can best be manufactured by winding on a template, with one such template required for each layer. This large separation between individual turns in both the axial and radial directions allows the gradient field to penetrate the magnet windings with acceptable eddy current losses, minimizing parasitic capacitance between the wires and resulting in a high self-resonant frequency. The small field-shaped section at the top of the solenoid has a sufficiently small number of turns that it does not require special winding procedures.

[0029] By combining the methods described in the previous two sections, a system combining a small bore magnet with a tightly fitting single-layer cylindrical Z-gradient coil can be realized without excessive voltage on the magnet. However, there is another way to achieve the same goal, and in this approach, the magnets can be essentially separated and wound to best suit them.

[0030] The key element in separating the single-layer gradient coil 3 from the complete magnet 2 is to divide the space inside the coil 3 into two sections 4, 5 and place the windings on the boundary 10 between these two, as shown in FIG. 7. One section 4, in one embodiment, is the space above the patient table (with extra margin to allow for the thickness of the tabletop and the RF coil below the table), and the other section 5 is the entire bottom area below this separation. FIG. 7 shows the cross-sectional shape of the coil 3. In the embodiment shown in FIG. 7, the lower boundary of the imaging section is located 180 mm below the isocenter.

[0031] In one embodiment of the present invention, the imaging compartment 4 surrounds the examination space of the magnetic resonance examination system, and the auxiliary compartment 5 is the remainder of the cylindrical coil volume below the patient table of the magnetic resonance examination system.

[0032] In the upper or imaging section, the field of view must be of the same quality as for an equivalent cylindrical gradient coil. The purpose of the windings around the lower section is to balance the magnetic flux through the magnet windings at every axial position of the windings. Figure 8(a) shows the z-gradient coil 3 with an optimized winding pattern. To obtain the optimal winding pattern for the two sets of coils, the total rms magnetic flux coupled to a set of test windings placed on a cylinder of the same diameter as the superconducting magnet coil was added to the numerical optimization cost function. Other factors in this cost function are the deviation of the magnetic field within the imaging volume, magnetic stored energy and dissipation, local surface current density, net magnetic force in the field of the main magnet, and other requirements that may be added by the gradient coil designer.

[0033] FIG. 8( a) shows a magnet coil assembly 1 including a main magnet 2 having superconducting windings 9 for generating a uniform static main magnetic field in an examination region, and an unshielded gradient coil 3 for generating magnetic field gradient waveforms. The unshielded gradient coil 3 has two separate sections 4, 5: an imaging section 4 around an examination zone of the main magnet, and at least an auxiliary section 5 offset from the imaging section 4 in the radial or axial direction of the main magnet 2 and adjacent to the imaging space 4. The imaging section 4 and the auxiliary section 5 have coil windings 7, 8 disposed on the surfaces of the imaging section 4 and the auxiliary section 5, respectively, at least partially along the boundary 10 between the sections 4, 5. In one embodiment, the coil winding 8 of the auxiliary section 5 carries the same current as the coil winding 7 of the imaging section 4. This can be achieved, for example, by connecting the coil winding 8 of the auxiliary section 5 and the coil winding 7 of the imaging section 4 in series. However, the currents in the coil windings 7, 8 may be different.

[0034] In Figures 8b)-8c), only a portion of the coil is shown to provide a clearer overall view. In Figure 8b), the imaging section 4 and the auxiliary section 5 are shown in cross section. In particular, the coil windings 7 and 8 can be seen on the surfaces of the sections 4 and 5. In particular, the coil windings 7 and 8 run along the boundary 10 between the sections 4 and 5. In Figure 8c), only the auxiliary section 5 with the coil winding 8 is shown in cross section. In Figure 8d), only the imaging section 4 with the coil winding 7 is shown in cross section.

[0035] Figure 9 shows that the compensating winding in the lower section balances the mutual inductance between the gradient coils and the magnet surprisingly well. Even with a 10x vertical scale increase compared to Figure 4, there is very little visible.

[0036] The absence of net magnetic flux through the magnet windings does not mean that this gradient coil does not generate a magnetic field at the location of the magnet windings. Figure 10 shows a comparison of this field magnitude with an equivalent unconstrained and axially compensated gradient coil. The stored energy at 10 mT / m of the two-section coil of Figure 8 is 5.8 J, and therefore it is more efficient than the axially shielded coil described above.

[0037] In principle, a two-section z-gradient could be combined with a cylindrical transverse gradient coil wrapped around the outer surface of the z-gradient. Alternatively, the presence of the lower boundary of the imaging section could be utilized to significantly improve the efficiency of these transverse coils while simultaneously reducing the magnetic field generated in the magnet windings. A cylindrical X- or Y-gradient coil approximately 800 mm in diameter would be compatible with the z-coil shown in Figure 8 and would have a stored energy of 6.8 Joules at 10 mT / m. An equivalent Y-coil that follows the contours of the upper section of the z-gradient coil would have a stored energy of only 4.1 J, meaning that 40% less amplifiers would be required to drive it. Here, the Y-direction refers to the transverse direction parallel to the surface of the patient table, and the X-direction refers to the transverse direction perpendicular to the surface of the patient table.

[0038] For symmetry reasons, the imaging section 4 of the unshielded gradient coil shown in Figure 11 does not inductively couple to any of the magnet windings. An asymmetric upper and lower gradient coil on this same surface would be more efficient, but this coil would also couple strongly to the magnet. However, similar to how the Z gradient can be decoupled, an additional set of balance windings around the lower gradient coil section can fully compensate for this net coupling. Even with these additional windings, the stored energy of the decoupled upper and lower gradients is only 3.9 J, a significant reduction from the 6.9 J of the purely cylindrical gradient.

[0039] With reference to Figure 12, Figure 12a) shows a magnetic field map of an asymmetric upper-lower gradient with a decoupling coil in the lower section, and Figures 12b-d) show separate coil sections according to an embodiment of the present invention. Figure 12A) shows, similar to Figure 8A), a magnet coil assembly 1 having a main magnet 2 with superconducting windings 9 for generating a uniform static main magnetic field within an examination zone, and unshielded gradient coils 3 for generating gradient magnetic field waveforms. The unshielded gradient coil 3 has two separate sections 4, 5: an imaging section 4 around the examination zone of the main magnet, and at least an auxiliary section 5 offset from the imaging section 4 in the radial or axial direction of the main magnet 2 and adjacent to the imaging section. The imaging section 4 and the auxiliary section 5 have coil windings 7, 8 disposed on the surfaces of the imaging section 4 and the auxiliary section 5, respectively, and at least partially along the boundary 10 between the sections 4, 5.

[0040] 12(b)-(c), only a portion of the coil is shown to facilitate a comprehensive understanding of the coil. In FIG. 12(b), the imaging section 4 and the auxiliary section 5 are shown in cross section. In particular, the coil windings 7 and 8 can be recognized on the surfaces of the sections 4 and 5. In particular, the coil windings 7 and 8 run along the boundary 10 between the sections 4 and 5. In FIG. 12(c), only the auxiliary section 5 with the coil winding 8 is shown in cross section. In FIG. 12(d), only the imaging section 4 with the coil winding 7 is shown in cross section.

[0041] While the invention has been illustrated and described in detail in the drawings and the foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive. The invention is not limited to the disclosed embodiments. Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. In the claims, the word "comprising" does not exclude other elements or steps, and the indefinite article "a" or "an" does not exclude a plurality. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. Any reference signs in the claims are not to be construed as limiting the scope. Moreover, for the sake of clarity, not all components in the drawings have been labeled with reference signs. [Explanation of symbols]

[0042] 1 Gradient Coil Assembly 2 Main magnet 3 Unshielded Gradient Coils 4 Imaging section 5. Auxiliary Sections 6 Gradient Coil Windings 7. Coil winding for the first section 8. Coil winding of the second section 9 Main magnet winding 10 Boundary between the first and second compartments

Claims

1. 1. A magnet coil assembly for a magnetic resonance examination system, comprising: a main magnet having superconducting windings that generate a uniform static main magnetic field in the examination region; an unshielded gradient coil for generating gradient magnetic field waveforms; and the unshielded gradient coil has two separate sections: an imaging section located around an examination region of the main magnet, and at least one auxiliary section offset from the imaging section in a radial or axial direction of the main magnet and adjacent to the imaging section; a magnet coil assembly, wherein the imaging section and the auxiliary section have coil windings disposed on surfaces of the imaging section and the auxiliary section, each at least partially disposed along a boundary between the imaging section and the auxiliary section.

2. The magnet coil assembly of claim 1 , wherein the coil windings of the auxiliary section carry the same current as the coil windings of the imaging section.

3. The magnet coil assembly of claim 2 , wherein the coil windings of the auxiliary section and the coil windings of the imaging section are connected in series.

4. 4. The magnet coil assembly according to claim 1, wherein the imaging section surrounds an examination space of the magnetic resonance examination system, and the auxiliary section is the remaining part of a cylindrical coil volume of the magnetic resonance examination system below a patient table of the magnetic resonance examination system.

5. 5. A magnet coil assembly according to claim 1, wherein the unshielded gradient coil is a longitudinal z-gradient coil or an asymmetric superior-inferior gradient coil.

6. 6. A magnet coil assembly according to claim 1, wherein the boundary between the imaging section and the auxiliary section is located in a range of 150 to 200 mm below the isocenter of the main magnet.

7. A magnetic resonance examination system comprising a magnet coil assembly according to any one of claims 1 to 6.

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