阅读说明：本技术 用于磁共振成像系统的梯度系统 (Gradient system for magnetic resonance imaging system ) 是由 斯特凡·波佩斯库 于 2020-11-25 设计创作，主要内容包括：本发明描述了用于磁共振成像系统的梯度系统(6),梯度系统(6)包括：使用公共基本磁场(B0)的至少两个检查区域(M1,M2,M3,M4,M5,M6)和在至少两个检查区域(M1,M2,M3,M4,M5,M6)中的多个梯度线圈(25x,25y,25z)；以及梯度控制单元(22),其被设计成使得其以时间同步的方式控制流过不同检查区域(M1,M2,M3,M4,M5,M6)中的相似梯度轴的至少两个梯度线圈(25x,25y,25z)的电流。本发明还描述了控制这样的梯度系统的方法、用于磁共振成像系统的控制装置以及这样的磁共振成像系统。(The invention describes a gradient system (6) for a magnetic resonance imaging system, the gradient system (6) comprising: at least two examination zones (M1, M2, M3, M4, M5, M6) using a common basic magnetic field (B0) and a plurality of gradient coils (25x, 25y, 25z) in the at least two examination zones (M1, M2, M3, M4, M5, M6); and a gradient control unit (22) which is designed such that it controls the currents flowing through at least two gradient coils (25x, 25y, 25z) of similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6) in a time-synchronized manner. The invention also describes a method of controlling such a gradient system, a control device for a magnetic resonance imaging system and such a magnetic resonance imaging system.)

1. A gradient system (6) for a magnetic resonance imaging system, the gradient system (6) comprising: at least two examination zones (M1, M2, M3, M4, M5, M6) using a common basic magnetic field (B0) and a plurality of gradient coils (25x, 25y, 25z) in the at least two examination zones (M1, M2, M3, M4, M5, M6); and a gradient control unit (22), the gradient control unit (22) being designed such that the gradient control unit (22) controls the currents flowing through at least two gradient coils (25x, 25y, 25z) of similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6) in a time-synchronized manner.

2. Gradient system according to claim 1, comprising a set of gradient coils (25x, 25y, 25z) for similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6), wherein the gradient coils (25x, 25y, 25z) in the set are electrically connected in series and/or in parallel and the set is powered by one single power unit (23), preferably wherein the gradient coils (25x, 25y, 25z) for similar gradient axes in all examination zones (M1, M2, M3, M4, M5, M6) are powered by one single power unit (23).

3. Gradient system according to one of the preceding claims, wherein the gradient coils (25x, 25y, 25z) comprise a central plane and are arranged such that the central planes of two adjacent gradient coils (25x, 25y, 25z) are at an angle of more than 10 ° with respect to each other such that the coils (25x, 25y, 25z) are V-shaped.

4. Gradient system according to one of the preceding claims, wherein the gradient coils (25x, 25y, 25z) are arranged such that the resulting gradient fields (Gx, Gy, Gz) have a ring shape or a ring shape with straight channels, wherein the gradient coils (25x, 25y, 25z) are preferably arranged in a star shape, preferably rotationally symmetrically, around at least one central axis (a).

5. Gradient system according to one of the preceding claims, wherein the gradient coils (25x, 25y, 25z) are arranged such that the examination zone (M1, M2, M3, M4, M5, M6) comprises only gradient coils (25x, 25y, 25z) on one single side,

preferably, wherein the number of gradient coils (25x, 25y, 25z) for a group of examination zones (M1, M2, M3, M4, M5, M6) corresponds to the number of examination zones (M1, M2, M3, M4, M5, M6) in the group,

preferably, wherein the separating element between adjacent examination zones (M1, M2, M3, M4, M5, M6) comprises only one gradient coil (25x, 25y, 25z) for each gradient axis of the two examination zones (M1, M2, M3, M4, M5, M6).

6. Gradient system according to one of the preceding claims, wherein the gradient coils (25x, 25y, 25z) are biplane gradient coils (25x, 25y, 25z), preferably wherein a plurality of gradient coils (25x, 25y, 25z) are formed to cover one side of the entire field of view of the examination region (M1, M2, M3, M4, M5, M6), preferably mechanically and/or permanently attachable to the MRI scanner (2).

7. Gradient system according to one of the preceding claims, comprising magnetic field shim coils and/or active shield coils, wherein these coils are preferably arranged similar to the gradient coils (25x, 25y, 25 z).

8. Gradient system according to one of the preceding claims, wherein gradient coils (25Z) for gradients of the Z-axis are connected to the gradient control unit (22) such that adjacent gradient coils (25Z) apply mirrored magnetic fields.

9. Method, in particular for controlling a gradient system (6) according to the preceding claim, for applying gradient fields with gradient coils (25x, 25y, 25z) to at least two examination zones (M1, M2, M3, M4, M5, M6), the method comprising the steps of:

-applying currents through at least two gradient coils (25x, 25y, 25z) of similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6) in a time-synchronized manner.

10. A control device (13) for a magnetic resonance imaging system (1) is designed to apply currents flowing through at least two gradient coils (25x, 25y, 25z) of similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6) in a time-synchronized manner.

11. A magnetic resonance imaging system (1) comprising at least two examination zones (M1, M2, M3, M4, M5, M6), and a gradient system (6) according to one of the claims 1 to 8 and/or a control device according to claim 10.

12. The magnetic resonance imaging system of claim 11, comprising a power unit (23) for applying electrical currents to gradient coils (25x, 25y, 25z), wherein the magnetic resonance imaging system (1) comprises a set of gradient coils (25x, 25y, 25z) for similar gradient axes in different examination zones (M1, M2, M3, M4, M5, M6), wherein the gradient coils (25x, 25y, 25z) in the set are electrically connected in a series and/or parallel connection and the set is powered by the power unit (23).

13. The magnetic resonance imaging system of claim 11 or 12, comprising basic field magnets (44) between adjacent examination zones, wherein a single gradient coil (25x, 25y, 25z) for the gradient axis is connected to a set of basic field magnets (44).

14. The magnetic resonance imaging system of one of claims 11 to 13, wherein the gradient system (6) comprises a plurality of gradient coils (25x, 25y, 25z), the plurality of gradient coils (25x, 25y, 25z) covering one side of the entire field of view of an examination region (M1, M2, M3, M4, M5, M6), preferably the plurality of gradient coils (25x, 25y, 25z) are mechanically and/or permanently attached to an MRI scanner (2) of the magnetic resonance imaging system (1).

15. The magnetic resonance imaging system of one of claims 11 to 14, wherein a central plane of the gradient coils (25x, 25y, 25z) of the gradient system (6) on at least one side of the examination region (M1, M2, M3, M4, M5, M6) is parallel to a basic field magnet of an MRI scanner (2) of the magnetic resonance imaging system (1).

Technical Field

Gradient systems for magnetic resonance imaging systems ("MRI systems"), in particular for MRI systems having two or more examination zones, and such MRI systems are described. The invention also describes a method of controlling such a gradient system, a control device for an MRI system and such a magnetic resonance imaging system.

Background

For over forty years, the principles of magnetic resonance imaging ("MRI") have been used for imaging and other measurements. Despite the long and important time of this measurement method, only two magnet designs are currently used for clinically used MRI systems or MRI scanners: c-magnet form and solenoid. The operation of this type of MRI scanner remains problematic for clinical workflow.

The most serious problems occur with respect to the wide range of stray magnetic fields around these scanners. To deal with this problem and avoid accidents and damage, hospital authorities must demarcate tightly controlled areas within and near the MRI exam room by restricting the access of people and equipment. Damage may occur if a metal or magnetic component is attracted by the strong magnet of the MRI scanner and accelerates in the direction of the scanner volume.

Another problem is that MRI scanners using solenoid-magnet designs "wrap" the patient in a narrow patient passageway, which can lead to claustrophobia, among other things. This claustrophobia may be so intense in some patients that MRI scans cannot be performed. Furthermore, access to the patient by medical personnel is severely restricted due to the narrow examination channel, which is disadvantageous for interventional or therapeutic procedures, in particular for real-time MRI imaging.

Typically, MRI scanners use self-shielded solenoid-type superconducting magnets to reduce the strength of the leakage magnetic field produced by the coils of the basic field magnet. Actively shielded basic field magnets are much more expensive than unshielded basic field magnets. In addition, the shielding coil reduces the efficiency of the basic magnetic field that can be used for measurements in the examination channel. The active (active) bucked magnets have a larger diameter (about 220cm) than the unshielded magnets (about 145 cm).

An alternative design of MR scanner uses C-shaped magnets. This may be a permanent magnet or an electromagnet consisting of two helmholtz coils. The C-shaped magnet has two pole pieces which produce a perpendicular basic magnetic field in their space. A similar structure is a mechanically stronger door magnet (portal magnet), and may also be implemented with superconducting helmholtz coils in some embodiments. C-shaped and portal magnets have the advantage of open access to the patient and additionally reduce claustrophobia. However, such a structure requires a very strong mechanical structure to counteract the large magnetic attraction between the two opposing basic field magnets. To reduce the propagation of stray magnetic fields, these magnet architectures typically use iron yokes to guide and close magnetic field lines outside the imaging volume. The yoke is one of the most cost effective shields. A disadvantage of such a yoke is the large size, weight and volume of the MR scanner.

A method to solve these problems has been introduced a short time ago. The method is based on an MRI system with a toroidal magnetic field. Unlike prior art MR magnets that use solenoids or helmholtz-pair magnet coils, toroidal coils tend to confine the magnetic field inside the toroid, with only small and not as far stray magnetic fields as possible. This system not only overcomes the problems of stray magnetic fields and light weight construction, but also provides the opportunity to achieve two or more examination regions in one single MRI system. An example of such an MRI system is a basic field magnet arrangement with three, four, six or eight (e.g. identical) basic field magnet segments arranged in a star shape around a central axis with rotational symmetry (e.g. 60 ° for six magnets and six examination zones). The basic magnetic field has a main direction which extends in the form of a toroidal magnetic field.

Another approach to solve these problems has been introduced shortly ago. This alternative approach is based on a conventional MRI scanner surrounded by a "satellite scanner" which uses the stray fundamental magnetic field of the conventional MRI scanner. The additional scanners do not of course use as strong a basic magnetic field as inside a conventional MRI scanner, however, they provide the opportunity to realize a plurality of inexpensive examination zones in which the basic examination can be performed. For example, in case of a disaster or a remote area, many patients can be examined simultaneously using one single (mobile) MRT system.

There are local gradient systems with coil pairs arranged parallel to the left and right side of the patient. However, although such known gradient systems can also be used for these new MRI systems, there is currently no gradient system that works in an optimal way with these MRI systems. In particular, conventional gradient systems generate stray fields outside of conventional MRI scanners that may affect other examination regions of the above-described new MRI systems.

Disclosure of Invention

It is an object of the invention to improve the known MRI system in order to facilitate an improved gradient system which is suitable for MRI systems having two or more examination zones, in particular for generating a spatially non-constant magnetic field.

This object is achieved by a gradient system, a method, a control unit and a magnetic resonance imaging system according to the invention.

A gradient system for a magnetic resonance imaging system according to the invention comprises at least two examination zones and a plurality of gradient coils using a common basic magnetic field. Gradient coils are arranged in at least two examination zones. The system further comprises a gradient control unit which is designed such that it controls the currents flowing through at least two gradient coils of similar gradient axes in different examination zones in a time-synchronized manner.

Gradient systems typically include gradient coils disposed in an examination region (e.g., at a basic field magnet bounding the examination region). Although known gradient coils may be used, particular shapes and/or particular arrangements may provide additional advantages, as described further below. Although a coil may comprise many different conductor loops, all conductor loops on one side of the examination area, which produce a gradient on the same axis, are hereinafter considered to be one single coil.

The gradient control unit may be a component which provides control signals for the currents of the gradient coils (for example as a controller of a power amplifier) as well as a unit which itself provides the currents (for example in case the gradient control unit comprises a respective power amplifier). The gradient control unit is designed such that it controls the currents flowing through the at least two gradient coils in a time-synchronized manner, i.e. simultaneously. Preferably, the gradient control unit is designed to coordinate all gradient activities, preferably to coordinate independent or synchronized operation of the different gradient axes, in particular even minimization and/or correction of cross interference terms between the gradient coils.

With regard to the gradient axes in the different examination zones, it should be noted that the gradient axes follow the local coordinate system in the respective examination zone. Typically, the z-gradient axis follows the basic magnetic field. In a ring-shaped arrangement, the z-axis (all together) will extend in a circle or polygon, since the orientation of the basic magnetic field is different in each examination zone and it has the shape of a circle or polygon. Similarly, the X-axis, which is usually directed perpendicular to the basic magnetic field, parallel to the annular plane, will be different in each examination zone, always directed outside the annular basic magnetic field plane. The y-axis is generally directed perpendicular to the basic magnetic field and its annular plane. Preferably, for each examination zone there is associated with it a local coordinate system XYZ, wherein the local Z-axis extends parallel to the static basic magnetic field B0 and points in the same direction as the static basic magnetic field B0, the Y-axis is parallel to the vertical rotational symmetry axis of the MRI scanner, and the X-axis corresponds to the radial direction pointing from the magnet out through the vertical mid-plane of the imaging compartment from the center of symmetry.

The coils for similar gradient axes are two Z-gradient coils in the two examination zones, or two X-gradient coils, or two Y-gradient coils. The axes do not necessarily point in similar directions, they must be the same axes with respect to the local coordinate system of the gradient coil in the respective examination zone.

An additional global coordinate system may be connected to the entire MRI scanner. The global coordinate system preferably includes a vertical Y-axis (identical to the rotational symmetry axis), a radial R-coordinate pointing radially outward from the symmetry axis, and an angular theta coordinate. For a star arrangement of basic field magnets of an MRT scanner, the spatial relationship between the global coordinate system and the local coordinate system is as follows: the local Y coordinate is always equal to the global Y coordinate. The local X-axis corresponds to a radial spoke (spoke) through the vertical midplane of the imaging compartment. All local Z-axes combine to construct the sides of a polygon that passes through the horizontal mid-plane of the magnet and is centered on the axis of symmetry.

Of course, the gradient system should also comprise further components for optimal operation, which the state of the art gradient systems also comprise. These components are, for example, dedicated gradient power amplifiers, shim coils or holding structures for each axis GPAx, GPAy and GPAz.

Such a gradient system has the advantage that it generates a special, synchronous gradient field. This is very advantageous for MRI scanners with basic field magnets in an inclined arrangement (e.g. a star arrangement). It is also advantageous for the "satellite scanner" described above.

The method according to the invention, in particular the method of controlling a gradient system according to the invention for applying a gradient field to at least two examination zones with gradient coils, comprises the steps of:

-applying currents through at least two gradient coils of similar gradient axes in different examination zones in a time-synchronized manner.

Thus, there are at least two gradient fields applied simultaneously in two different examination zones for similar axes (X-axis or Y-axis or Z-axis). Of course, the method may be applied to more than one axis, such that two axes or all three axes are driven synchronously, wherein a synchronous operation is necessary for any similar axis.

The control device for a magnetic resonance imaging system according to the invention is designed to apply currents through at least two gradient coils of similar gradient axes in different examination zones in a time-synchronized manner. The control means preferably comprise a system according to the invention. The control means may comprise additional units or means for controlling components of the magnetic resonance imaging system, such as a sequence control unit for measurement sequence control, a memory, radio frequency transmission means for generating, amplifying and transmitting RF pulses, a gradient system interface, radio frequency reception means for acquiring magnetic resonance signals and/or a reconstruction unit for reconstructing magnetic resonance image data.

A magnetic resonance imaging system comprises at least two examination zones and a gradient system according to the invention and/or a control device according to the invention. A preferred MRI scanner of such a magnetic resonance imaging system comprises an inclined arrangement, e.g. a star arrangement, of basic field magnets. Particularly preferred is an MRI scanner having a ring MRI scanner architecture. In a star arrangement of basic field magnets with a toroidal magnetic field, the front side of the gradient system should be directed outside the toroidal magnetic field.

Particularly advantageous embodiments and features of the invention are given by the dependent claims, as disclosed in the following description. Features from different claim categories may be combined as appropriate to give other embodiments not described herein.

A preferred gradient system comprises a set of gradient coils for similar gradient axes in different examination zones (in the examination zone all gradient coils for the X-axis, all gradient coils for the Y-axis or all gradient coils for the Z-axis), wherein the gradient coils of the set are electrically connected in series and/or parallel connection and the set is powered by a (in particular one single) power unit (e.g. a power amplifier of a gradient control unit). Preferably, the gradient coils for similar gradient axes (all X-coils, all Y-coils or all Z-coils) in all examination areas are powered by this principle (one single power unit per axis). This has the advantage that an easy synchronization operation can be achieved. The series connection has the further advantage that: the current in all coils connected in series is the same.

To clarify the position in the 3D coordinate system, a reference to the center plane of the gradient coil is present. The central plane of the coil is the plane (or at least the mean median plane) of the loops of the coil. Considering a planar gradient coil, the central plane is the plane of the coil, considering a helmholtz coil, the central plane is the plane of the coil windings, wherein the gradient magnetic field vector is perpendicular to the central plane. The gradient coils (i.e. the X-gradient coil, the Y-gradient coil and/or the Z-gradient coil) are preferably arranged at opposite walls of the examination zone. The examination region of the gradient system is the region between the gradient coils, since in normal use in an MRT scanner the examination region for an MRI examination will be between the gradient coils.

According to a preferred gradient system, the gradient coils comprise a central plane and are arranged such that the central planes of two adjacent gradient coils are at an angle of more than 10 ° with respect to each other, such that the gradient system is V-shaped. This arrangement is advantageous for a star MRI scanner arrangement. For better understanding, it is defined herein that the "front side" of the gradient system is the mouth of the V-shape, while the "back side" is the opposite side, with the pairs of gradient coils closest to each other. The vector of the X-axis points to the front side of the gradient system.

According to a preferred gradient system, the gradient coils are arranged such that the resulting gradient field has a ring shape or a ring shape with straight channels (hereinafter referred to as "ring arrangement"), wherein the gradient coils are preferably arranged in a star shape, in particular rotationally symmetric, around at least one central axis. The particular arrangement of the coils tends to confine the magnetic field to the interior of the toroid, with only small stray fields that cannot reach that far. Therefore, an expensive magnetic shield is not required.

Typically, the gradient coils are arranged in pairs on opposite sides of the examination region, so that on one side of the examination region there is a coil for the gradient axis and on the other side there is one coil for the axis. This means that in each examination zone there are two gradient coils for each gradient axis. Typically, the gradient system creates a gradient in the X, Y and Z directions to achieve a gradient vector with X, Y and Z contributions. Therefore, preferably, there are a pair of X-gradient coils, a pair of Y-gradient coils and a pair of Z-gradient coils. It is clear that two gradient coils designed to generate the same gradient (and arranged on different sides of the examination zone) form a pair.

Unlike the prior art, it is not necessary to provide a pair of gradient coils for each gradient axis in each imaging compartment. In particular with the described annular arrangement, one gradient coil of the gradient axis can be used for two adjacent examination regions, so that no pairs of gradient coils are required. For example, in an MRI scanner with six examination zones and a ring-shaped basic magnetic field, it is sufficient to provide a minimum of six gradient coils for each axis, which means that each imaging compartment (i.e. examination zone) accommodates a single gradient coil for the X-axis, a single gradient coil for the Y-axis and a single gradient coil for the Z-axis.

According to a preferred gradient system, the gradient coils are arranged such that the examination region comprises only a single gradient coil for the gradient axis (on one single side). This means that in one examination zone there is preferably only one gradient coil for the X-axis, one gradient coil for the Y-axis and one gradient coil for the Z-axis. The coils are arranged in particular on a "wall" (basic field magnet) between two adjacent examination zones. Preferably, the number of gradient coils of a set of examination regions corresponds to the number of examination regions in the set. This means that there are no pairs of gradient coils in the set. Preferably, the separating element between adjacent examination zones comprises only one gradient coil for each gradient axis of the two examination zones.

Preferably, the system of planar gradient coils implemented as planar solenoids cooperate to generate gradient fields over all imaging volumes. This cooperation is not only based on synchronous operation, but also on the arrangement of the coils.

Thus, it is not necessary to have two gradient coil housing units in each examination region. In some particular embodiments and in order to save space, it is foreseen that only a single housing may be sufficient, that is, only one side (e.g. the left part) or only the other side (e.g. the right part) may be used within the imaging compartment. Of course, specific combinations are possible without limiting the scope of the invention. For example, one side (e.g., the left portion) may accommodate two gradient coils (e.g., an X-gradient coil and a Y-gradient coil) for two different axes, while the other side (e.g., the right portion) accommodates only one gradient coil (e.g., a Z-gradient coil) for one axis.

According to a preferred gradient system, the gradient coils are biplane gradient coils. This brings the advantage that the gradient system does not require much space. Preferably, the central planes of the gradient coils on one side of the examination zone are parallel to each other, in particular in case the gradient coils are biplane gradient coils.

It is further preferred that the gradient system comprises two or three pairs of gradient coils (e.g. X-, Y-and Z-gradient coils), wherein all pairs of gradient coils are arranged at the same angle to each other (i.e. the angle between the central planes is equal).

Further preferably, a plurality of gradient coils are formed to cover one side of the entire field of view (FoV) of the examination region or at least a part of the FoV, preferably the gradient coils are mechanically and/or permanently attachable to the MRI scanner. Preferably, the gradient coil system has two building blocks which are located symmetrically (mirror-symmetrical or rotationally symmetrical) to the right (e.g. block 1) and to the left (e.g. block 2) of the examination region. The right block 1 and the left block 2 are preferably mirror image pairs or rotation pairs. Each block preferably integrates at least one gradient coil for one axis. Typically, three gradient coils for all three axes are integrated per block. This means that the block 1 comprises a stack of planar gradient coils, for example the right half of an X-gradient coil, a Y-gradient coil and a Z-gradient coil. In particular in the case of a V-shaped basic magnet arrangement, the examination zone (or each examination zone) has a V-shaped gradient coil system which is attached to the basic field magnet and comprises the two blocks. Such a V-shaped architecture of the gradient coil better utilizes the available installation space within the examination region, which is shaped like a triangular or trapezoidal prism. For a local gradient system, each block (1, 2) consists of a stack of planar gradient coils, for example each block integrates half of an X-gradient coil pair, a Y-gradient coil pair and a Z-gradient coil pair.

It should be noted that the multiple gradient coils may also be formed as local gradient coils, preferably as local gradient coils for head imaging. Preferably, the gradient coils are integrated into and/or mechanically attached to a headrest of the patient seat.

Preferred gradient systems comprise magnetic field shim coils and/or active shield coils, wherein the coils are preferably arranged similar to the gradient coils (e.g. the central planes of the coils are parallel to the central plane of the gradient coils).

With respect to the above examples, the gradient system blocks (1, 2) may also integrate magnetic field shim coils and/or active shield coils designed to attenuate stray gradient fields outside the imaging volume that would otherwise penetrate the adjacent examination region and/or imaging volume.

According to a preferred gradient system, the gradient coils for the gradient of the Z-axis are connected to a gradient control unit such that adjacent coils apply mirror magnetic fields.

It should be noted that in the case of a coil and an examination area "adjacent" means with respect to the basic magnetic field. Thus, the predetermined direction of the basic magnetic field defines the order of adjacent elements.

According to a preferred gradient system, the gradient coil is formed by a plurality of conductor loops, preferably wherein,

the loops of the gradient coil for the X-gradient comprise two sets of counter-rotating (constraining) loops adjacent in the X-direction, preferably wherein the radius of one set of loops increases, wherein the outer conductor in the X-direction remains substantially at the side of the gradient coil,

the loops of the gradient coil for the Y-gradient comprise two sets of counter-rotating loops adjacent in the Y-direction, preferably wherein the radius of one set of loops increases, wherein the outer conductor in the Y-direction remains substantially at the side of the gradient coil,

the loops of the gradient coil for the Z-gradient comprise an incremental set of loops, preferably wherein the center of the loops remains substantially at the center of the gradient coil.

According to a preferred gradient system, the distance of the field-dependent conductors of a set of loops steadily decreases at least in the direction of the aperture (aperture) of the V-shape of the gradient system, with respect to the gradient coils for X-gradients and/or the gradient coils for Z-gradients. Alternatively or additionally, with regard to the gradient coil for the Y-gradient, the radius of the set of loops also increases in the X-direction to the aperture of the V-shape, wherein the outer conductor remains substantially at the aperture.

A preferred magnetic resonance imaging system comprises a power unit for applying electrical currents to the gradient coils, wherein the magnetic resonance imaging system is designed such that a set of gradient coils of the magnetic resonance imaging system are directed to similar gradient axes in different examination zones, wherein the gradient coils of the set are electrically connected in series and/or in parallel and the set is powered by the power unit.

A preferred gradient system comprises basic field magnets between adjacent examination zones, wherein a single gradient coil for the gradient axis is connected to a set of basic field magnets.

The preferred X-gradient field has an annular shape over the examination region. The preferred X-gradient field has a dipole distribution in which the positive component (added to and increased by the basic magnetic field B0) is at the positive X-value of the local X-coordinate, and the negative component (subtracted from and decreased by the basic magnetic field B0) is at the negative X-value of the local X-coordinate. For example, at a negative X-coordinate, that is to say at a spatial position lying between the origin of the local coordinate system, for example the isocenter (isocenter) of the examination zone, and the central axis of the MRI scanner, the X-gradient field is negative and cancels out (attenuates) the static basic magnetic field B0. At positive X-coordinates, that is to say at spatial positions located between the origin of the examination zone and the entrance, the X-gradient field is positive and it is added to the (enhanced) static basic magnetic field B0. The global distribution of the X-gradient field preferably follows the same rule in all examination regions and is preferably identical there (viewed from the respective local coordinate system). Preferably, the gradient field lines are closed following a polygonal (e.g. hexagonal) contour over the imaging compartment. At inner spatial locations (x < 0), the polygon and closed constant field lines extend in the opposite direction to B0, while at outer spatial locations (x > 0), the gradient field lines extend with the B0 lines. For an ideal gradient distribution, the resultant magnetic field at any spatial location in the local coordinate system is given by the following expression: b (X, y, z) ═ B0+ Gx × X, where Gx is the selected strength of the X gradient controlled by the strength of the current flowing in the gradient coil.

All of the inventive features disclosed above for the X-gradient system are also applicable to the Y-gradient field, the only difference being that the Y-gradient field is implemented along the vertical Y-axis rather than along the horizontal X-axis. Thus, the Y-gradient field is also preferably closed in a ring or polygon (e.g. hexagon) over and throughout all imaging compartments. The Y-gradient field intensifies the static basic magnetic field at those spatial regions whose local Y-coordinate is positive (Y > 0) and weakens the static basic magnetic field in those spatial regions whose local Y-coordinate is negative (Y < 0). It should be noted that the direction of reinforcement/weakening may also be the opposite direction.

In contrast to X-gradient systems, Y-gradient coils are preferably equivalent to these X-gradient coils, in which the gradient currents and the associated conductor patterns are rotated by 90 ° in the plane in order to achieve the target field distribution of the Y-gradient. For an ideal gradient distribution, the resultant magnetic field at any spatial location in the local coordinate system is given by the following expression: b (x, Y, z) ═ B0+ Gy × Y, where Gy is the selected strength of the Y gradient controlled by the strength of the current flowing in the gradient coil.

With regard to the Z-gradient, with reference to the examination region, the Z-gradient preferably weakens the static basic magnetic field B0 at those spatial regions where the local Z-coordinate is positive (Z > 0) and strengthens the static basic magnetic field B0 in those spatial regions where the local Z-coordinate is negative (Z < 0). For an ideal gradient distribution, the resultant magnetic field at any spatial location in the local coordinate system is given by the following expression: in any second examination zone, B (x, y, Z) ═ B0-gzxz, and in any other second examination zone, B (x, y, Z) ═ B0+ gzxz, where Gz is the selected strength of the Z gradient field controlled by the strength of the currents flowing in the gradient coils. Similar to the X-and Y-gradients, the Z-gradient field is also closed in a ring or polygon (here a hexagon) over and throughout all examination regions.

It should be noted that there are significant differences that avoid strong variations in the magnetic field strength at the boundary between the two examination zones. By a new and further inventive method, preferably applied only to Z-gradient systems, the direction of flow of the coil currents is preferably alternated for each coil, so that the (Z) -gradient fields are mirrored in adjacent examination zones.

Preferably, the coil current flows in a counter-clockwise direction in the coil of any second examination region and in a clockwise direction (or in the opposite direction) in any other second examination region. This facilitates the generation of a spatial distribution of the Z-gradient fields within the imaging compartment and, in addition, allows the use of only six Z-gradient coils for all imaging compartments. As a result of this solution, the slope of the resulting Z-gradient field alternates per imaging compartment.

It should be noted that at the boundary between two adjacent imaging regions, the magnetic field strength does not change abruptly, but passes continuously from one region to the next. For example, in the right half of the first examination zone, when the local Z-coordinate (Z < 0) approaches a negative peak, the static basic magnetic field increases according to the expression B (Z) -B0-gzxz, and it reaches a maximum amplitude at the boundary to the first examination zone. In the left half of the second examination zone, in which the local Z coordinate Z > 0, the static basic magnetic field decreases according to the expression B (Z) ═ B0+ Gz × Z, and reaches zero amplitude in the middle of the second examination zone when Z ═ 0.

Referring now to the second examination zone, the Z-gradient intensifies the static basic magnetic field at those spatial regions whose local Z-coordinate is positive (Z > 0) and weakens the static basic magnetic field B0 in those spatial regions whose local Z-coordinate is negative (Z < 0). For an ideal gradient distribution, the resultant magnetic field at any spatial location in the local coordinate system is given by the following expression: b (x, y, Z) ═ B0+ Gz × Z, where Gz is the selected strength of the Z gradient field controlled by the strength of the currents flowing in the gradient coils.

This inversion of the Z-gradient field does not significantly complicate the scanning sequence running synchronously over all imaging compartments, since it can be corrected by a coordinate transformation from a logical gradient coordinate system to a physical gradient coordinate system, so that the Z-direction is inverted only for any second examination zone.

According to a preferred gradient system, the gradient coil is formed by a plurality of conductor loops. It is clear that preferably only one long conductor is wound into a plurality of loops, however, open loops (open loops) connecting to each other may also occur. In the following, the loops of the coil are designated as "set of loops", wherein any reference to the movement action is meant to be understood as a change of the subsequent loop. The following design is preferred (alternatively or additionally).

The loops of the gradient coil for the X-gradient comprise two sets of counter-rotating loops adjacent in the X-direction, preferably wherein the radius of one set of loops increases while the outer conductor in the X-direction (front and back) remains substantially at the side of the gradient coil. This means that the shape of such a coil is reminiscent of the shape of a butterfly.

The loops of the gradient coil for the Y-gradient comprise two sets of counter-rotating loops adjacent in the Y-direction, preferably wherein the radius of one set of loops increases while the outer conductor in the Y-direction remains substantially at the sides of the gradient coil (sides perpendicular to the front and back sides). This means that the coil can be seen as a coil for an X-gradient rotated only 90 °.

The loops of the gradient coil for the Z-gradient comprise an incremental set of loops, preferably wherein the center of the loops remains substantially at the center of the gradient coil. This means that the coils can be coaxial but become larger at least in the X-direction.

The specific design of the more preferred coil is described below. This special design results in a linear gradient field which increases in the direction of the V-shaped mouth (to the front side) of the gradient system to compensate for the radial effect.

With regard to the gradient coil for the X-gradient, the distance of the field-dependent conductors of a set of loops steadily decreases at least in the direction of the aperture of the V-shape of the gradient system (in the direction to the front side). The field dependent conductors are those parts of the loop of the magnetic field that determine the gradient.

With regard to the gradient coil for the Z-gradient, the distance of the field-dependent conductors of a set of loops steadily decreases at least in the direction of the aperture of the V-shape of the gradient system (in the direction to the front side).

With regard to the gradient coil for the Y-gradient, the radius of the set of loops increases in the X-direction as well as in the Y-direction, wherein the outer conductor at the V-shaped aperture remains substantially at the sides of the aperture (i.e., at the front side) as well as at the sides perpendicular to the front side.

Various hardware or software tools may be used to further fine-tune these lead patterns in order to satisfy some additional constraints (e.g., gradient linearity) to reduce the magnitude of stray fields, mechanical vibrations, and the level of acoustic noise or peripheral nerve stimulation.

The wire pattern distribution eliminates inherent non-linearity of the gradient field along the radial or local X-axis. This non-linearity is only produced by V-shaped gradient coils when the conductor pattern density along the X-axis is approximately constant.

Special loop-shaping (loop-shaping) as described above may eliminate the non-linear components of the X-and Z-gradient coils along the X-axis by modifying the wire spacing along the X-axis from a constant wire spacing to a more quadratic wire spacing, where the wire density increases approximately quadratically with radial distance from the axis of symmetry (e.g., the axis of symmetry of the annular basic magnet). For a Y-gradient coil, the exemplary solution adds an additional wire distribution with constant spacing along the X-axis. This is similar to the conductor pattern of the magnet coils used for generating the static basic magnetic field B0.

According to a preferred magnetic resonance imaging system, the gradient system comprises a plurality of gradient coils covering one side (or at least a part of one side) of the entire field of view of the examination zone, preferably mechanically and/or permanently attached to an MRI scanner of the magnetic resonance imaging system.

According to a preferred magnetic resonance imaging system, the central plane of the gradient coils of the gradient system on at least one side of the examination zone is arranged parallel to the basic field magnetic coils of the MRI scanner of the magnetic resonance imaging system. Thus, the MRI scanner has a tilted basic field magnet, and the gradient system follows the arrangement of the basic field magnet.

The preferred opening angle of the V-shape of the gradient system is 120 ° (for example for a star MRI scanner with 3 examination zones), 90 ° (for example for a star MRI scanner with 4 examination zones), 60 ° (for example for a star MRI scanner with 6 examination zones) or 45 ° (for example for a star MRI scanner with 8 examination zones).

An advantage of the gradient system according to the invention is that a (annular) parallel MRI scanner with global gradient coils operating synchronously in all examination regions can be realized to generate global gradient fields, in particular with an annular distribution. A particular advantage of this solution is that it requires only a minimum of three Gradient Power Amplifiers (GPA) for all examination regions, while further reducing the cost and installation space of the gradient coils. Furthermore, with respect to the annular arrangement, because all gradient coils are operated together and co-operate and all act on the same annularly distributed global gradient field, there will be no stray gradient fields leaking from one imaging region into the other.

Description of the drawings

Other objects and features of the present invention will become apparent from the following detailed description considered in conjunction with the accompanying drawings. It is to be understood, however, that the drawings are designed solely for the purposes of illustration and not as a definition of the limits of the invention.

FIG. 1 shows a simplified MRI system according to an embodiment of the present invention.

Fig. 2 shows an exemplary embodiment of a magnetic resonance tomography system with a star-shaped basic field magnet arrangement with circular closed basic magnetic field lines.

Fig. 3 shows a detailed schematic representation of individual basic field magnet segments of a star-shaped basic field magnet arrangement with polygonal closed basic field lines.

Figure 4 shows a configuration of a gradient coil driven by a single gradient power amplifier.

Fig. 5 shows an exemplary X-gradient field.

Fig. 6 shows an exemplary y-gradient field.

Fig. 7 shows an exemplary z-gradient field.

Fig. 8 graphically illustrates an exemplary z-gradient field.

In the drawings, like numerals refer to like elements throughout. The objects in the drawings are not necessarily to scale.

Detailed Description

Fig. 1 shows a schematic representation of a magnetic resonance imaging system 1 ("MRI system"). The MRI system 1 comprises an actual magnetic resonance scanner (data acquisition unit) 2 with an examination space 3 or patient tunnel, wherein a patient or test person is located on a drive bed 8, in which an actual examination object O is located.

The magnetic resonance scanner 2 is typically equipped with a basic field magnet system 4, a gradient system 6, and an RF transmit antenna system 5 and an RF receive antenna system 7. In the exemplary embodiment shown, the RF transmit antenna system 5 is a whole-body coil (whole-body coil) which is permanently mounted in the magnetic resonance scanner 2, in contrast to the RF receive antenna system 7 which is formed as a local coil (here represented by only a single local coil) to be arranged on the patient or test object. In principle, however, the whole-body coil can also be used as an RF receiving antenna system, and the local coils can be switched to different operating modes, respectively.

The basic field magnet system is typically designed such that it generates a basic magnetic field which advances in the z direction in the longitudinal direction of the patient (i.e. along the longitudinal axis of the magnetic resonance scanner 2). The gradient system 6 typically comprises individually controllable gradient coils in order to be able to switch (activate) the gradients in the X-direction, the y-direction or the z-direction independently of one another.

The MRI system 1 shown here is a whole-body system with a patient access into which the patient can enter completely. In principle, however, the invention may also be used for other MRI systems (e.g. MRI systems with a laterally open C-shaped housing), as well as for smaller magnetic resonance scanners in which only one body part can be placed.

The MRI system 1 also has a central control device 13 for controlling the MRI system 1. The central control device 13 comprises a sequence control unit 14 for measurement sequence control. With the sequence control unit 14, a series of radio frequency pulses (RF pulses) and gradient pulses can be controlled according to the selected pulse sequence.

For outputting the individual RF pulses of the pulse sequence, the central control device 13 has a radio frequency transmission device 15, which generates and amplifies the RF pulses and feeds them into the RF transmission antenna system 5 via a suitable interface (not shown in detail). For controlling the gradient coils of the gradient system 6, the control device 13 has a gradient system interface 16. The sequence control unit 14 communicates with the radio frequency transmission means 15 and the gradient system interface 16 in a suitable manner to transmit the pulse sequence.

Furthermore, the control device 13 has a radio frequency receiving device 17 (which likewise communicates in a suitable manner with the sequence control unit 14) in order to acquire magnetic resonance signals (i.e. raw data) for the individual measurements, which are received in a coordinated manner from the RF receiving antenna system 7 over the range of the pulse sequence.

A reconstruction unit 18 receives the acquired raw data and reconstructs magnetic resonance image data therefrom for measurement. Such reconstruction is typically performed based on parameters that may be specified in the respective measurement or control protocol. The image data may then be stored in the memory 19, for example.

The central control device 13 can be operated via a terminal 10 having an input unit and a display unit 9, and therefore the entire MRI system 1 can also be operated by an operator via the terminal 10. The MR images can also be displayed on the display unit 9 and measurements can be planned and started by means of the input unit (possibly in combination with the display unit 9) and, in particular, an appropriate control protocol can be selected (and possibly modified) with an appropriate series of pulse sequences PS as described above.

The MRI system 1 according to the invention, and in particular the control device 13, may have a number of additional components, not shown in detail but normally present in such systems, such as a network interface, in order to connect the entire system with a network and to be able to exchange raw data and/or image data, or parameter maps, respectively, but also additional data (e.g. patient-related data or control protocols).

The way in which appropriate raw data are acquired by irradiation of RF pulses and generation of gradient fields and MR images are reconstructed from the raw data is known to the person skilled in the art and therefore does not need to be explained in detail here.

Fig. 2 shows an exemplary embodiment of a magnetic resonance tomography system 1 with a star-shaped basic field magnet arrangement 40. The MRI system comprises a ring MRI scanner 2 which allows scanning up to six patients O simultaneously.

Here, a magnetic resonance scanner 2 is shown, the function of which can be controlled by a control device 13. The control device 13 can in principle be constructed in a similar manner and have the same components as the control device 13 in the conventional MR system according to fig. 1. Likewise, the control device 13 may also have a suitable terminal or the like (not shown here).

The basic field magnet arrangement 40 of the magnetic resonance scanner 2 in this figure comprises: six (here identical) basic field magnet segments 44, which are arranged in a star shape about the central axis a with a rotational symmetry of 60 ° in this embodiment. The basic magnetic field B0 indicated by an arrow has a basic field main direction R0, which extends in the form of a circular or ring-shaped magnetic field.

This magnet solution provides six examination zones (imaging volumes) M1, M2, M3, M4, M5, M6, wherein the homogeneity of the magnetic field B0 is sufficiently high for performing MR imaging sequences known from the prior art of MRI. Thus, the scanner is assumed to acquire raw data and reconstruct an image acquisition sequence and image reconstruction methods known in the art (e.g., fig. 1).

For each examination region M1, M2, M3, M4, M5, M6, there is associated therewith a local coordinate system XYZ, as exemplarily depicted for the two examination regions M3, M5. The local Z-axis extends parallel to the static basic magnetic field B0 and points in the same direction as the static basic magnetic field B0. The Y-axis is parallel to the vertical rotational symmetry axis (central axis a) of the MRI scanner 2, while the X-axis corresponds to the radial direction pointing outwards from the magnet through the vertical mid-plane of the imaging compartment from the symmetry center.

Fig. 3 shows a detailed schematic representation of the individual basic field magnet segments 44 of the star-shaped basic field magnet arrangement 40. Here, the six coil arrangements can be regarded as basic field magnet segments 44 of the basic field magnet arrangement 40.

It should be noted that in fig. 3, the lines of the basic magnetic field B0 do not form a circle, but rather a hexagonal outline in each examination zone M1, M2, M3, M4, M5, M6. The field lines of the basic magnetic field B0 form parallel lines.

Such a "six-pack" ring MRI scanner allows scanning up to six patients simultaneously in six imaging volumes (examination zones), wherein the homogeneity of the magnetic field B0 is sufficiently high to perform MR imaging sequences known in the art. The scanner is assumed to acquire raw data and reconstruct MR images by the acquisition and image reconstruction method as explained in the procedure of fig. 1.

As already indicated in fig. 2, for each imaging volume there is associated with it a local coordinate system XYZ, as exemplarily depicted here for only two imaging volumes. Again, the local Z-axis extends parallel to the static basic magnetic field B0 and points in the same direction as the static basic magnetic field B0. The Y-axis is parallel to the axis of rotational symmetry of the six-set magnet system, while the X-axis corresponds to the radial direction pointing outward from the center of symmetry by the magnets.

The arrows and contour lines depict the overall distribution of the static basic magnetic field B0 in a cross-sectional plane through the middle of the magnet. The local magnetic field vector depicted by the small arrows shows the distribution of the local magnetic field magnitude (arrow magnitude) and its direction (arrow orientation).

A significant advantage of such a symmetrical arrangement is the structural stability when the basic magnetic field B0 is switched on. The magnetic forces between the individual basic field magnet segments 44 cancel one another in the direction of the main magnetic field direction R0. Each basic field magnet segment 44 is attracted by its two adjacent basic field magnet segments each with the same force. The resulting force acts inwardly towards the column 43 and can be compensated well there by suitable structural reinforcements.

Such a magnetic resonance imaging system 1 with a basic field magnet arrangement 40 according to fig. 2 and 3 allows measurements to be taken at six different examination zones M1, M2, M3, M4, M5, M6 (see fig. 2), wherein, in the example shown, measurements of an object O (a patient as shown or an inanimate object) are taken at the examination zone M4, wherein the patient stands upright here on a vertical wall of the basic field magnet arrangement 40. In principle, measurements can be carried out simultaneously at all six examination zones M1, M2, M3, M4, M5, M6. The central column 43 holds the basic field magnet segments 44 in place and may also comprise technical components, such as electrical connections or even power supplies.

In each case, the measuring apparatus 12 (only symbolically shown) or components (for example HF coils) required for this purpose at the examination zones M1, M2, M3, M4, M5, M6, respectively, can be located at the examination zones M1, M2, M3, M4, M5, M6 transmit coils of the RF transmit system, the RF receive coils of the RF receive system and/or the common RF transmit/receive coils. Again, this may include gradient and/or shim coils. All these components may be controlled in coordination by a common control device 13.

Of course, the magnetic resonance scanner 2 may also have more than six examination zones M1, M2, M3, M4, M5, M6, the height of the magnetic resonance scanner 2 may be lower, or the magnetic resonance scanner 2 may be designed for examining small regions of the body, for example for head examinations or examinations of limbs, female breasts, prostate, liver, kidney or other organs. The star-shaped basic field magnet arrangement 40 may also be positioned lying flat.

In fig. 2, an example of a local gradient system 6 is shown in one examination region M4 surrounding the head of a patient. The V-shape of the gradient system 6 follows the angle between the two basic field magnet segments 44, i.e. 60 °.

Another example of a gradient system 6 in an adjacent examination region M3 is also shown in fig. 2. The gradient system 6 comprises: large gradient coils, which cover the entire field of view of the examination region M3 and are mechanically and permanently attached to the basic field magnet segment 44 of the MRI scanner 2. Each examination region M1, M2, M3, M4, M5, M6 may have such a V-shaped gradient system 6 attached thereto and consisting of: a right block and a mirrored pair of left blocks. The V-shaped configuration of the gradient system 6 better utilizes the magnetic field space available in the examination zone, which is shaped like a triangular or trapezoidal prism. For the local gradient system 6, each block comprises a stack of planar gradient coils 25x, 25y, 25Z (see the figures below), for example each block integrates X, Y and half of a Z-gradient coil pair. Furthermore, the blocks may also integrate magnetic field shim coils and/or active shielding coils such that these coils are designed to attenuate stray gradient fields outside the imaging volume that would otherwise penetrate the adjacent examination region and/or imaging volume. The gradient system 6 may further comprise: dedicated gradient power amplifiers for each axis GPAx, GPAy and GPAz and for each examination region, which are preferably all controlled by a common central unit.

The principles of the present invention are illustrated in several examples below. It should be noted that in these examples for gradient systems, the same reference numerals are used as for the gradient systems mentioned before. The reason is that the intention and purpose of the gradient system 6 is always the same, i.e. to generate a suitable gradient field for the measurement. The gradient system 6 according to the invention may absolutely use a design as described above, wherein, however, the gradient coil arrangement should be installed in most (all) examination regions M1, M2, M3, M4, M5, M6. It would be advantageous to use the same or at least similar gradient coil arrangements in the examination regions M1, M2, M3, M4, M5, M6, since this would result in a periodic gradient field. The gradient system 6 of the state of the art differs from the gradient system 6 according to the invention in the functional arrangement (e.g. internal wiring) of the coils 25a, 25b, 25c such that the coils in the different examination zones M1, M2, M3, M4, M5, M6 can be driven in a synchronized manner. This is explained below.

Fig. 4 shows a configuration of a Z-gradient coil 25Z driven by a single gradient power unit 23 (e.g., a power amplifier). In this example, the gradient coil 25z is a z-axis gradient coil 25z for a gradient system 6 (see, e.g., fig. 2) of the six-bank parallel MRI scanner 2, wherein the gradient control unit 22 of the gradient system 6 provides gradient signals to one single gradient power amplifier ("GPA") of the z-axis gradient coil 25 z. Since in this example the MRI scanner 2 comprises only one gradient coil 25z per examination region M1, M2, M3, M4, M5, M6, the GPA drives a total of six gradient coils per axis, each gradient coil being individually accommodated within one examination region M1, M2, M3, M4, M5, M6. In fig. 2, in which gradient coil pairs are shown, the number of gradient coils 25z driven by one single GPA will be 12.

Fig. 5 shows an exemplary X-gradient field Gx. The figure shows the target spatial distribution of the global gradient field Gx for the X-axis over the examination zone M4, which, however, as shown in the top view at the bottom right of the figure, should indeed extend over all six examination zones M1, M2, M3, M4, M5, M6. The gradient field has a bipolar distribution, with a positive component (added to and increased by the basic magnetic field B0) at a positive X value of the local X coordinate and a negative component (subtracted from and decreased by the basic magnetic field B0) at a negative X value of the local X coordinate. This distribution is exemplarily depicted in the examination region M4 by an arrow, the thickness of which indicates the field strength of the X-gradient field Gx. The thicker arrows indicate stronger local gradient strength.

The arrows point in the direction of the local gradient field. At the negative X-coordinate, that is to say at a spatial position located between the origin of the local coordinate system (e.g. the isocenter of the examination zones M1, M2, M3, M4, M5, M6) and the central axis a of the MRI scanner 2, the X-gradient field Gx is negative and cancels (attenuates) the static basic magnetic field B0. At positive X-coordinates, that is to say at spatial positions located between the origin and the entrance of the examination zones M1, M2, M3, M4, M5, M6, the X-gradient field Gx is positive and it is added to the (intensified) static basic magnetic field B0. The global distribution of the X-gradient field Gx preferably follows the same rule in all examination regions M1, M2, M3, M4, M5, M6, and is preferably identical there (from the respective local coordinate system). As shown in the bottom right top view, the gradient field lines are closed on the imaging compartment with a hexagonal outline. At inner spatial locations (x < 0), the polygon and closed constant field lines extend in the opposite direction to B0, while at outer spatial locations (x > 0), the gradient field lines extend with the B0 lines.

The figure also exemplarily depicts the main gradient current pattern and the gradient coils 25X for generating this spatial distribution of the X-gradient field only for the two examination regions M4, M5. In this configuration, the paired current loops generate a circularly distributed positive X gradient field Gx, where the field vector is depicted by the arrow pointing from right to left. It should be noted that the pairs of current loops belong to a single gradient coil 25 x.

The paired current loops generate a circularly distributed negative X-gradient field, wherein the field vector is again depicted by the arrow pointing from the left to the right. This wire and current pattern is repeated identically for all other X gradient coils 25X (see lower right picture). All the X-gradient coils 25X work together in coordination to generate an X-gradient field having a circular distribution as depicted in the figure.

It should be noted that fig. 5 depicts a simplified, more ideal-like spatial distribution of the gradient field. In effect, the gradient field is non-linear across the imaging compartment and away from the center/symmetry axis. Furthermore, the so-called accompanying terms arise under the intensification of the laws of physics. Whenever a linear gradient is activated, a concomitant magnetic field with a non-linear spatial correlation result occurs. This is a result of maxwell's equations, i.e., the magnetic field must have zero divergence and negligible curl within the imaging volume.

Fig. 6 shows the distribution of the Y gradient field Gy. All of the inventive features disclosed above for the X-gradient system and depicted by fig. 5 also apply to the Y-gradient field Gy, with the only difference that the Y-gradient field Gy is implemented along the vertical Y-axis instead of along the horizontal X-axis. Thus, the Y gradient field Gy is also preferably a ring or polygon (e.g., a hexagon) that is closed over and throughout all imaging compartments. The Y-gradient field Gy intensifies the static basic magnetic field B0 in those spatial regions whose local Y-coordinate is positive (Y > 0) and attenuates the static basic magnetic field B0 in those spatial regions whose local Y-coordinate is negative (Y < 0). It should be noted that the direction of reinforcement/weakening may also be the opposite direction.

Similar to the X-gradient system, there are preferably six Y-gradient coils 25Y, which are identical to the X-gradient coils 25X in fig. 5. The exemplarily depicted gradient currents and the associated conductor patterns are rotated by 90 ° in the plane in order to achieve a target field distribution of the Y gradient field Gy.

Fig. 7 shows an exemplary Z-gradient field. The figure depicts the target magnetic field distribution that is generated when the Z-gradient field Gz is active. The Z-gradient system preferably comprises six gradient coils 25Z with corresponding wire patterns and gradient currents. The Z-gradient field Gz is depicted only for the three examination zones M3, M4, M5. However, as shown in the lower right top view, they should be present in all examination regions M1, M2, M3, M4, M5, M6.

Referring to the front examination region M4, the Z-gradient attenuates the static basic magnetic field B0 at those spatial regions whose local Z-coordinate is positive (Z > 0) and enhances the static basic magnetic field B0 in those spatial regions whose local Z-coordinate is negative (Z < 0). Similar to the X and Y gradients, the Z gradient field Gz is also closed in a ring or polygon (here a hexagon) over and throughout all examination regions M1, M2, M3, M4, M5, M6.

It should be noted that there is a significant difference that strong variations in the magnetic field strength at the boundary between the two examination zones M1, M2, M3, M4, M5, M6 are avoided. By the new and further inventive method (applied only to Z-gradient systems), the flow direction of the coil current is preferably alternated for each coil, so that the Z-gradient fields are mirrored in the adjacent examination zones M1, M2, M3, M4, M5, M6.

The coil current flows in the counterclockwise direction in the coil 25z of any second examination region M1, M3, M5 and flows in the clockwise direction (or the opposite direction) in any other second examination region M2, M4, M6. This facilitates the generation of a spatial distribution of the Z-gradient fields within the imaging compartment and, in addition, allows the use of only six Z-gradient coils for all imaging compartments. The result of this solution is that the slope of the Z-gradient field alternates per imaging compartment, as further depicted in fig. 8.

This figure 8 clearly depicts the practical advantages of the gradient system of the present invention. It shows that at the boundary between two adjacent imaging regions the magnetic field strength does not change abruptly, but passes continuously from one region to the next. For example, in the right half of the first examination zone M1, when the local Z coordinate (Z < 0) approaches a negative peak, the static basic magnetic field increases and reaches a maximum amplitude at the boundary of the first examination zone M1 to the second examination zone M2. In the left half of the first examination zone M1, in which the local Z coordinate Z > 0, the static basic magnetic field decreases and, when Z is 0, zero amplitude is reached in the middle of the first examination zone M1.

Referring now to the second examination zone M2, the Z-gradient intensifies the static basic magnetic field B0 at those spatial regions whose local Z-coordinate is positive (Z > 0), and weakens the static basic magnetic field B0 in those spatial regions whose local Z-coordinate is negative (Z < 0).

This inversion of the Z-gradient field Gz does not significantly complicate the scanning sequence running synchronously over all imaging compartments, since it can be corrected by a coordinate transformation from a logical gradient coordinate system to a physical gradient coordinate system, thereby inverting the Z-direction only for any second examination zone.

When the ring-shaped gradient fields penetrate the magnet coils (these are the coils generating the static basic magnetic field B0), this is a limitation of the solution requiring the magnet coils to withstand the switched gradient fields without the associated difficulties. For example, an electromagnet for low-field MRI would not have the problems associated with this requirement. Superconducting magnets will require special protection measures to reduce the heat introduced into, for example, Cu wires supporting SC material. These magnets are for example referred to as "transparent" SC magnets.

Although the present invention has been disclosed in the form of preferred embodiments and variations thereon, it will be understood that numerous additional modifications and variations could be made thereto without departing from the scope of the invention. For the sake of clarity, it is to be understood that the use of "a" or "an" throughout this application does not exclude a plurality, and "comprising" does not exclude other steps or elements. Reference to "a unit" or "a device" does not preclude the use of more than one unit or device.