Patent Publication Number: US-2013241558-A1

Title: Magnetic Resonance Tomograph with Cooling Device for Gradient Coils

Description:
This application claims the benefit of DE 10 2012 203 974.0, filed on Mar. 14, 2012, which is hereby incorporated by reference. 
     BACKGROUND 
     The present embodiments relate to a magnetic resonance tomography device. 
     Magnetic resonance devices (MRTs) for the examination of objects or patients using magnetic resonance tomography are known, for example, from DE10314215B4. 
     A magnetic resonance tomography device MRT has, for example, three-axle gradient coils (e.g., GC or Gradient Coil) that are employed to generate magnetic fields in the direction of three Cartesian spatial axes, for example. In order to generate the desired field strengths, currents of several hundred amperes may be used. The gradient coil conductors may be placed layer by layer on cylindrical surfaces. The gradient coil conductors are exposed to high alternating forces (e.g., Lorentz forces) on account of the arrangement in the base field of the MRT magnets. In order to achieve a mechanical fixing of the conductors and a good thermal coupling with the cooling device, the conductors are, for example, embedded in a resin matrix (e.g., epoxy). The high electrical currents generate thermal losses up to 25 kW. 
     In order to be able to discharge dissipative power as effectively as possible, cooling hoses are embedded in the resin between the individual coil layers (e.g., several hundred meters of cooling hose per coil and several parallel cooling circuits). The thermal losses formed in the coil windings may be discharged to the heat sink (e.g., a cooling medium such as water) with as minimal a thermal resistance as possible. At the same time, electrical insulation may be established between the copper coils and, if necessary, an electrically conductive cooling medium. 
     Considerable care is therefore taken in terms of optimizing the space requirement for the individual layers. If a large conductor cross-section is selected for the coil conductor in order to generate less thermal losses, this results in an increased radial space requirement for the overall coil. The larger the radius of a coil layer is selected, the more current is expended to generate the desired magnetic field. The current requirement may be somewhat proportional to the fifth power of the radius (I˜R 5 ). The radii may be kept as small as possible, and the layer structure may be provided in as compact a manner as possible. The conductor cross-sections are, for example, selected to be as large in order to achieve an operating temperature of approximately 85° C. during nominal output operation. 
     SUMMARY AND DESCRIPTION 
     The present embodiments may obviate one or more of the drawbacks or limitations in the related art. For example, the cooling of gradient coils in a magnetic resonance tomography (MRT) device may be optimized. 
     Without necessarily changing the thickness of the cooling layers, the flow rate or the cooling medium, embodiments optimize the cooling of the innermost gradient coil layer compared with known conventional structures. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  shows a magnetic resonance tomography (MRT) system; 
         FIG. 2  shows simplified partial coil layers of a gradient system of an MRT; 
         FIG. 3  shows a systematic partial section through a known gradient coil cooling of three Cartesian coil layers, having an integrated cooling device and a typical temperature curve in a warm state; 
         FIG. 4  shows one embodiment of a gradient coil cooling as a systematic, simplified partial section; and 
         FIG. 5  shows one embodiment of a gradient coil cooling as in  FIG. 4 , as a systematic, simplified partial section extended by a temperature profile. 
     
    
    
     DETAILED DESCRIPTION 
       FIG. 1  shows a magnetic resonance imaging device MRT  10  disposed in a shielded room or Faraday cage F. The magnetic resonance imaging device MRT  10  includes a whole body coil  102  with a tubular room  103 , for example, into which a patient couch  104  with a body  105  (e.g., of an examination object such as a patient; with or without a local coil arrangement  106 ) may be moved in the direction of arrow z in order to generate recordings of the patient  105  using an imaging method. The local coil arrangement  106  is arranged on the patient. In a local region (e.g., field of view (FOV)) of the MRT, recordings of a subarea of the body  105  may be generated in the FOV with the local coil arrangement  106 . Signals from the local coil arrangement  106  may be evaluated (e.g., converted into images, stored or displayed) by an evaluation device (e.g., including elements  168 ,  115 ,  117 ,  119 ,  120 ,  121 ) of the MRT  101  that may be connected to the local coil arrangement  106  via coaxial cables or radio (e.g., element  167 ), for example. 
     In order to examine the body  105  (e.g., an examination object or a patient) using a magnetic resonance device MRT  101  using a magnetic resonance imaging, different magnetic fields attuned as precisely as possible to one another in terms of temporal and spatial characteristics are irradiated onto the body  105 . A strong magnet (e.g., a cryomagnet  107 ) in a measuring cabin with a tunnel-type opening  103  generates a strong static main magnetic field B 0  that amounts, for example, to 0.2 Tesla to 3 or more Tesla. The body  105  to be examined, mounted on a patient couch  104 , is moved into an approximately homogenous region of the main magnetic field B 0  in the FoV. Excitation of the nuclear spin of atomic nuclei of the body  105  takes place via high frequency magnetic excitation pulses B 1 ( x, y, z, t ) that are irradiated via a high frequency antenna (and/or, if necessary, a local coil arrangement) that is shown in  FIG. 1  in a very simplified manner as a body coil  108  (e.g. a multipart body coil  108   a ,  108   b ,  108   c ). High frequency excitation pulses are generated, for example, by a pulse generation unit  109  that is controlled by a pulse sequence control unit  110 . After amplification by a high frequency amplifier  111 , the high frequency excitation pulses are routed to the high frequency antenna  108 . The high frequency system shown in  FIG. 1  is only indicated schematically. In other embodiments, more than one pulse generation unit  109 , more than one high frequency amplifier  111  and a number of high frequency antennas  108 , a, b, c are used in a magnetic resonance device  101 . 
     The magnetic resonance device  101  has gradient coils  112   x ,  112   y ,  112   z , with which magnetic gradient fields B G (x, y, z, t) are irradiated during a measurement for selective slice excitation and local encoding of the measuring signal. The gradient coils  112   x ,  112   y ,  112   z  are controlled by a gradient coil control unit  114  that, similarly to the pulse generation unit  109 , is likewise connected to the pulse sequence control unit  110 . 
     Signals emitted by the excited nuclear spin (e.g., the atomic nuclei in the examination object) are received by the body coil  108  and/or at least one local coil arrangement  106 , amplified by an associated high frequency preamplifier  116  and further processed and digitalized by a receive unit  117 . The recorded measurement data is digitalized and stored as complex numerical values in a k-space matrix. An associated MR image may be reconstructed from the k-space matrix populated with values using a multidimensional Fourier transformation. 
     For a coil, which may be operated both in transmit and also in receive mode (e.g., the body coil  108  or a local coil  106 ), the correct signal forwarding is controlled by an upstream transmit/receive switch  118 . An image processing unit  119  generates an image from the measurement data. The generated image is shown to a user via a console terminal  120  and/or is stored in a storage unit  121 . A central computing unit  122  controls the individual system components. 
     Images with a high signal/noise ratio (SNR) may be recorded in MR tomography with local coil arrangements (e.g., coils, local coils). The local coil arrangements are antenna systems that are applied in the immediate vicinity on (anterior) and/or below (posterior), on, or in the body  105 . With an MR measurement, the excited nuclei induce a voltage into the individual antennas of the local coil. The induced voltage is amplified with a low noise preamplifier (e.g., LNA, preamp) and forwarded to the receive electronics. In order to improve the signal/noise ratio even with highly resolved images, high field systems are used (e.g., 1.5T-12T or more). If more individual antennas may be connected to an MR receive system than there are receivers present, a switching matrix (e.g., RCCS) is integrated between the receive antennas and the receiver. The switching matrix routes the currently active receive channels (e.g., the receive channels that currently lie in the field of view of the magnet) to the existing receiver. More coil elements than there are receivers present may thus be connected, since with a whole body coverage, only the coils that are disposed in the field of view and/or in the homogeneity volume of the magnet are to be read out. 
     An antenna system may be referred to as a local coil arrangement  106 , for example, which may include an antenna element or, as an array coil, a number of antenna elements (e.g., coil elements). These individual antenna elements are embodied, for example, as loop antennas (loops), butterfly, flexible coils or saddle coils. A local coil arrangement includes, for example, coil elements, a pre-amplifier, further electronics (e.g., a balun), a housing, supports and may include a cable with a plug, by which the local coil arrangement is connected to the MRT system. A receiver  168  attached on the system side filters and digitalizes a signal received by a local coil  106  (e.g., by radio) and transfers the data to a digital signal processing device. The digital signal processing device may derive an image or a spectrum from the data obtained by measurement and provides the user with the image and the spectrum for a subsequent diagnosis and/or storage purposes, for example. 
       FIG. 2  shows schematic gradient coil layers a, b, c (e.g., a and b for the generation of magnetic fields in the x and y direction) of a gradient system of an MRT  101 . 
     Coils in coil layers a, b, c are embodied by their arrangement for the generation of a gradient magnetic field (BG (x, y, z, t)) in one of three directions x, y, z (e.g., the coil  112   z  for the generation of a gradient magnetic field in the direction z such that the coil  112   z  has windings arranged in an approximately circular manner about the axis Ax, z; the coil  112   y  for the generation of a gradient magnetic field in direction y; and the coil  112   x  for the generation of a gradient magnetic field in direction x). 
       FIG. 3  shows a known gradient coil cooling having transversal coil layers a, b, c for the generation of magnetic fields in x-, y-, and z-directions. 
     In accordance with the gradient coil cooling shown in  FIG. 3 , coil layers are arranged far radially inwards in order to retain an efficient structure. A radially outer lying cooling layer discharges heat produced in the coil layers through current into gradient coils. A further coil layer is provided on the cooling layer. The Helmholtz-type wound c-coil, which intrinsically also involves the highest efficiency of the field generation, may be selected. The advantage of the described arrangement may be a high layer efficiency of the two transversal coil layers. The relatively high thermal resistance of the layer remote from the cooling relative to the heat sink is disadvantageous with respect to a possible nominal current load. 
       FIG. 4  shows a schematic and simplified view of one embodiment of a gradient coil system GS (e.g., of a magnetic resonance tomography device  101 ) having three coil layers a, b, c. The three coil layers a, b, c are provided with gradient coils  112   x ,  12   y ,  112   z  therein for the generation, in each case, of a temporally changeable gradient magnetic field B G (x, y, z, t) in one of three directions (e.g., arranged orthogonal to one another; in the x-direction, y-direction, z-direction). 
     The coil layers a, b, c and cooling layers KL 1 , KL 2  may be arranged, for example, so as to surround a cylindrical axis Ax of the MRT bore  103  (e.g., MRT opening; including a radius Ra). 
     A first cooling layer KL 1 , which includes, for example, one or a number of cooling hoses KS 1  as a cooling element (e), is arranged between a first (a) and a second (b) of the coil layers a, b, c. A cooling medium (e.g., water Wa 1 ) passes through the cooling hoses KS 1 . 
     A second cooling layer KL 2 , which includes, for example, one or a number of cooling hoses KS 2  as a cooling element(s), is arranged between a second (b) and a third (c) of the coil layers a, b, c. A cooling medium (e.g., water Wa 2 ) passes through the cooling hoses KS 2 . 
     For the sake of clarity, only one winding of a cooling hose KS 1  (similarly KS 2 ) is shown in  FIG. 4  in each instance in a cross-section. A number of windings (e.g., in the z-direction) may be adjacent to one another, or several cooling hoses KS 1  (e.g., in the z-direction) may adjacent to one another in a cooling layer KL 1 , KL 2 . 
     Cooling hoses KS may be integrated in a manner known, for example, and/or may be connected to a circulating pump and/or a cooling unit. 
     At least one cooling layer KL 1 , KL 2  is arranged in the immediate vicinity of each of the three coil layers a, b, c (e.g., rests directly thereupon or is separated by a thin electrically insulating layer and/or supporting arrangement). 
     Radial conductor cross-sections of conductors  112   x ,  112   y ,  112   z  in the coil layers a, b, c may be smaller than the radial conductor cross-sections would be without two cooling layers on account of the two cooling layers KL 1 , KL 2 . 
     One advantage may be that a structure of a gradient coil arrangement is layered, which, compared with the prior art, may be energized more significantly with a similar permissible operating temperature (e.g., may be applied with current) and may thus enable higher nominal gradient fields. A cooling layer KL 1 , KL 2  may be arranged in the immediate vicinity of each coil layer a, b, c. The thermal transition resistance between a cooling layer KL 1  of the cooling layers and the coils (e.g., 12×) remote from the cooling according to  FIG. 2  may be reduced as a result. 
     In order not to increase the overall installation space and/or be able to disadvantageously shift conductor radii outwards, the conductor cross-sections (radial) may be reduced in this embodiment in order to obtain space (e.g., compared with the known prior art with only one cooling layer) for the additional cooling layers. If the cooling layers are embodied to be very thin, the reduction in the conductor height with the accompanying increased loss of power may be secondary compared with the gain in the heat reducing performance (or the cooling output). 
     Possible advantages may be a more effective cooling of the coil windings and/or reduced thermal resistance of the coil axis remote from the cooler relative to the cooling medium. As a result, operation of the gradient coil with higher current strengths (e.g., higher nominal gradient strengths with the same permissible maximum temperature) may be provided. Temperature peaks may be avoided in the region of tightly wound conductors in the coil planes. As a result, more even temperature distribution and less thermomechanical voltages in the coil structure may be provided. Optimization may be provided in terms of assembly of high-performance coils in a small installation space. 
       FIG. 5  shows a view as in  FIG. 4 , extended by a simplified, schematic temperature profile TP. On account of the coolant, the temperature in the cooling layers is at its lowest and is higher than in the coolant in the coils  112   x ,  112   y ,  112   z  of the coil layers a, b, c, and is just as high as in the two outer coil layers a, b in the innermost coil layer c. 
     While the present invention has been described above by reference to various embodiments, it should be understood that many changes and modifications can be made to the described embodiments. It is therefore intended that the foregoing description be regarded as illustrative rather than limiting, and that it be understood that all equivalents and/or combinations of embodiments are intended to be included in this description.