Patent Document

BACKGROUND OF THE INVENTION 
   This invention relates to a Magnetic Resonance Imaging (MRI) system. More particularly, this invention relates to superconducting coils used in MRI systems for correcting central magnetic field temporal shift, and for shielding external magnetic disturbances from large electromagnetic fields. 
   A highly uniform magnetic field is useful for nuclear resonance image (MRI) and nuclear magnetic resonance (NMR) systems as medical devices or chemical/biological devices. Most popular systems currently available worldwide use a superconducting magnet system which creates a highly uniform field in a pre-determined space (imaging volume). A superconducting magnet system usually uses multiple superconducting coils (main coil system) to achieve a desired highly uniform magnetic field in the imaging volume. More advanced superconducting MRI and NMR magnet systems also use an active shielding technique which adds a second set of multiple coils (shielding coil system) which creates a reverse direction magnetic field to reduce the fringe magnetic field and to achieve a significant reduction of the external magnetic field in the surrounding space of the magnet system. Depending on the design, the main coils system and shielding coils system can use a single circuit running the same electrical current, or two individual circuits running either the same current or two different currents. From the laws of physics, one knows that for a single superconductive closed loop, the total magnetic flux inside of the loop does not change. However, for a multi-coil system, especially for an actively shielded MRI magnet system with main coils and shielding coils connected in series, the situation is a little different. 
   Due to the environment disturbances, such as a train and/or other moving vehicles, rotating machinery, elevators, etc, in the surrounding area, the magnetic field of the system will have a corresponding temporal magnetic flux change. Practically, all magnet systems are subject to such temporal field instability ranging from ppm (parts per million) to ppb (parts per billion). But for actively shielded magnet systems, this change is more severe. For good image quality, the temporal field variation of a typical MRI should normally be less than 0.05 to 0.1 ppm/hour. The stability of the magnet center field is, however, highly affected by the environment disturbances, especially for those actively shielded magnets. The magnitude of the field fluctuations depends on both the size of the object and the distance away from the magnet system. For example, a typical elevator 20 feet away from the magnet can cause a field fluctuation of about 0.01 Gauss or 1.0E-6 Tesla, a subway can also cause a 0.1 Gauss field fluctuation. 
   Clearly, these environment disturbances included changes in both center magnetic field and its homogeneity will cause detectable deviation of the nuclear imaging quality (imaging distortion) for MRI and NMR devices. 
   In order to minimize such effects caused by environment changes and other disturbances, the electrical currents changing in both main coils and shielding coils should be controlled or limited to some prescribed acceptable level such that the environment disturbance is compensated and the center magnetic field remains constant and uniform. One structure and method has been described in U.S. Pat. No. 4,926,289 for such purpose by using a single filament or a few filaments of superconducting wire for the purpose of having low critical current. However, it would be desirable to provide methods and apparatus which are not constrained to filament(s) with low critical current. 
   In one aspect, a method of operating an imaging system having a main coil and a shield coil electromagnetically coupled to the main coil is provided. The method includes monitoring for an external environmental fluctuation of electromagnetism, and controlling current flow through the main and shield coils based upon the monitoring using a quench heater. 
   In another aspect, an imaging system includes a main coil, a shield coil positioned to shield an electromagnetic field generated by the main coil, and at least one environmental fluctuation circuit operationally coupled to at least one of the shield coil, the main coil the circuit including at least one detection coil, and a quench heater positioned proximate the detection coil. 
   In a further aspect, a method of operating an imaging system comprising a main coil, a shield coil positioned to shield an electromagnetic field generated by the main coil, and at least one environmental fluctuation circuit operationally coupled to at least one of the main coil and the shield coil, the circuit comprising at least one detection coil, and a quench heater positioned proximate the detection coil is provided. The method includes energizing the quench heater such that the detection coil is in a non-superconductive state, supplying current to the main coil and the shield coil until a predetermined current is reached while the detection coil is in the non-superconductive state, activating a persistence switch to a superconductive state, and de-energizing the quench heater when the persistence switch is in the superconductive state. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
       FIG. 1  is a block schematic diagram of a Magnetic Resonance Imaging (MRI) system. 
       FIG. 2  is a schematic diagram of a conventional circuitry of a superconducting current limiter for superconducting MRI, NMR magnet system. 
       FIG. 3  is a schematic diagram of a circuitry of the superconducting MRI, NMR magnet system illustrated in  FIG. 1  with a detection system. 
       FIG. 4  is a schematic diagram of another circuitry of the superconducting MRI, NMR magnet system shown in  FIG. 1  with a detection system. 
       FIG. 5  is a schematic diagram of a circuitry of the superconducting detection system shown in FIG.  3 . 
       FIG. 6  is a schematic diagram of a circuitry of the superconducting detection system shown in FIG.  4 . 
       FIG. 7  is a schematic diagram of the detection systems with a mechanical sensor system shown in  FIGS. 5 and 6 . 
       FIG. 8  is a schematic diagram of the detection systems with an electronic sensors system shown in FIGS.  5  and  6 . 
   

   DETAILED DESCRIPTION OF THE INVENTION 
   Herein described are methods and apparatus which utilize a current limiter for active shielding of a superconducting magnet system used in MRI and NMR magnetic field generators. More specifically, in one embodiment, a detection system is provided for an active shielding of superconducting magnet systems which use a single electrical current as explained in greater detail below. In another embodiment, a detection system is provided for an active shielding of a multiple electrical circuits superconducting magnet system as also explained in greater detail below. The herein described methods and apparatus use a combination of a detection mechanism and a controlled triggering level to limit the electrical current induced by environment disturbances. 
   As used herein, an element or step recited in the singular and proceeded with the word “a” or “an” should be understood as not excluding plural said elements or steps, unless such exclusion is explicitly recited. Furthermore, references to “one embodiment” of the present invention are not intended to be interpreted as excluding the existence of additional embodiments that also incorporate the recited features. Additionally, as is known in the art, a reference to a main coil contemplates a plurality of coils, and therefore the terms main coil and main coils are used interchangeably herein. For the same reason, the terms shield coil and shield coils are also interchangeable herein. 
     FIG. 1  is a block diagram of an embodiment of a magnetic resonance imaging (MRI) system  10  in which the herein described systems and methods are implemented. MRI  10  includes an operator console  12  which includes a keyboard and control panel  14  and a display  16 . Operator console  12  communicates through a link  18  with a separate computer system  20  thereby enabling an operator to control the production and display of images on screen  16 . Computer system  20  includes a plurality of modules  22  which communicate with each other through a backplane. In the exemplary embodiment, modules  22  include an image processor module  24 , a CPU module  26  and a memory module  28 , also referred to herein as a frame buffer for storing image data arrays. Computer system  20  is linked to a disk storage  30  and a tape drive  32  to facilitate storing image data and programs. Computer system  20  communicates with a separate system control  34  through a high speed serial link  36 . 
   System control  34  includes a plurality of modules  38  electrically coupled using a backplane (not shown). In the exemplary embodiment, modules  38  include a CPU module  40  and a pulse generator module  42  that is electrically coupled to operator console  12  using a serial link  44 . Link  44  facilitates transmitting and receiving commands between operator console  12  and system command  34  thereby allowing the operator to input a scan sequence that MRI system  10  is to perform. Pulse generator module  42  operates the system components to carry out the desired scan sequence, and generates data which indicative of the timing, strength and shape of the RF pulses which are to be produced, and the timing of and length of a data acquisition window. Pulse generator module  42  is electrically coupled to a gradient amplifier system  46  and provides gradient amplifier system  46  with a signal indicative of the timing and shape of the gradient pulses to be produced during the scan. Pulse generator module  42  is also configured to receive patient data from a physiological acquisition controller  48 . In the exemplary embodiment, physiological acquisition controller  48  is configured to receive inputs from a plurality of sensors indicative of a patient&#39;s physiological condition such as, but not limited to, ECG signals from electrodes attached to the patient. Pulse generator module  42  is electrically coupled to a scan room interface circuit  50  which is configured to receive signals from various sensors indicative of the patient condition and the magnet system. Scan room interface circuit  50  is also configured to transmit command signals such as, but not limited to, a command signal to move the patient to a desired position with a patient positioning system  52 . 
   The gradient waveforms produced by pulse generator module  42  are input to gradient amplifier system  46  that includes a G x  amplifier  54 , a G Y  amplifier  56 , and a G Z  amplifier  58 . Amplifiers  54 ,  56 , and  58  each excite a corresponding gradient coil in gradient coil assembly  60  to generate a plurality of magnetic field gradients used for position encoding acquired signals. In the exemplary embodiment, gradient coil assembly  60  includes a magnet assembly  62  that includes a polarizing magnet  64  and a whole-body RF coil  66 . 
   In use, a transceiver module  70  positioned in system control  34  generates a plurality of electrical pulses which are amplified by an RF amplifier  72  that is electrically coupled to RF coil  66  using a transmit/receive switch  74 . The resulting signals radiated by the excited nuclei in the patient are sensed by RF coil  66  and transmitted to a preamplifier  76  through transmit/receive switch  74 . The amplified NMR (nuclear magnetic resonance) signals are then demodulated, filtered, and digitized in a receiver section of transceiver  70 . Transmit/receive switch  74  is controlled by a signal from pulse generator module  42  to electrically connect RF amplifier  72  to coil  66  during the transmit mode and to connect preamplifier  76  during the receive mode. Transmit/receive switch  74  also enables a separate RF coil (for example, a surface coil) to be used in either the transmit or receive mode. 
   The NMR signals received by RF coil  66  are digitized by transceiver module  70  and transferred to a memory module  78  in system control  34 . When the scan is completed and an array of raw k-space data has been acquired in the memory module  78 , the raw k-space data is rearranged into separate k-space data arrays for each cardiac phase image to be reconstructed, and each of these arrays is input to an array processor  80  configured to Fourier transform the data into an array of image data. This image data is transmitted through serial link  36  to computer system  20  where it is stored in disk memory  30 . In response to commands received from operator console  12 , this image data may be archived on tape drive  32 , or it may be further processed by image processor  24  and transmitted to operator console  12  and presented on display  16 . 
     FIG. 2  illustrates a conventional circuitry of a superconducting MRI system  100  including a cryogenic temperature cryostat  102  in which a main coil  104 , a shielding coil  106 , a quench protection system  110 , and a superconducting persistent switch  112  are positioned. A power supply  108  is typically positioned outside cryostat  102 . During a magnet system energizing process, persistent switch  112  is in an off mode (i.e., a resistive state). Energy is supplied to main coil  104  and shielding coil  106  from power supply  108  until a desired magnetic field is produced, then persistent switch  112  is switched to an on mode (i.e., a superconductive state). Without electromagnetic disturbance, electrical current I a  of main coils  104 , and electrical current I b  of shielding coils  106  is the same in persistent mode. Upon an environment disturbance occurring, main coil electrical current I a  and shielding coil electrical current I b  can change slightly since the laws of physics necessitates only that a total magnetic flux of both main and shielding coils  104  and  106  together will attempt to remain constant. 
     FIG. 3  illustrates a circuitry of MRI system  10  including a two coil detection system  118 . MRI system  10  includes a cryogenic temperature cryostat  120  in which a main coil  122 , a shielding coil  124 , a quench protection system  128 , and a superconducting persistent switch  134  are positioned. A power supply  126  is typically positioned outside cryostat  120 . Detection system  118  includes an environmental fluctuation circuit  130 . In an exemplary embodiment, main coil  122  and shield coil  124  are wired in series receiving the same current, and environmental fluctuation circuit  130  includes two environmental fluctuation circuits  132 , one for main coil  122 , and one for shield coil  124 . During a magnet system energizing process, persistent switch  134  is in an off mode (i.e., a resistive state). Energy is supplied to main coil  122  and shielding coil  124  from power supply  126  until a desired magnetic field is produced, then persistent switch  134  is switched to an on mode (i.e., a superconductive state). During the just described magnet ramping, a pair of quench heaters (not shown in  FIG. 3 ) are turned on, thus the sections of CC′D′D and DD″E′E are resistive and prevent electrical current to flow therethrough, and all electrical current flows through main coil  122  and shielding coil  124 . After the magnet (coils  122  and  124 ) reaches a desired field level, and are shimmed and parked using conventional methods, the quench heaters of environmental fluctuation circuits  132  are turned off, and sections CC′D′D and DD″E′E return to a superconductive state. When an outside disturbance is present, both electrical currents in main coil  122  and shield coil  124  may start to change. Since coils  122  and  124  and environmental fluctuation circuits  132  are in the same circuit, any induced current flows through either CC′D′D, or DD′E′E circuit, or both circuits. Thus with the aid of a detection and controlling scheme identical or similar to that illustrated in  FIG. 5 , currents I c  and I d  are detected, limited, and/or controlled as explained below in greater detail. 
     FIG. 4  illustrates a one coil detection system  150  in which MRI system  10  includes a cryogenic temperature cryostat  152  in which a main coil  154 , a shielding coil  156 , a quench protection system  158 , and a superconducting persistent switch  160  are positioned. A power supply  161  is typically positioned outside cryostat  152 . System  150  also includes an environmental fluctuation circuit  162 . In an exemplary embodiment, main coil  154  and shield coil  156  are wired in series receiving the same current, and environmental fluctuation circuit  162  is wired in parallel to one of main coil  154  and shielding coil  156 . As illustrated in  FIG. 4 , environmental fluctuation circuit  162  is wired in parallel to main coil  154 . When electrical current I a  and I b  are not equal due to outside electromagnetic disturbances, the differential current of main coils I a  and shielding coils I b  flows through superconducting circuit CC′D′D, thus with aid of a detection and controlling scheme identical or similar to that illustrated in  FIG. 6 , a differential current I c  is detected, limited, and/or controlled. Although  FIG. 4  illustrates that superconducting wire is connected to main coil  154  at points C and D in  FIG. 4 , the superconducting wire alternatively can be connected to shield coil  156  similarly, or be connected to the points within the coil. For example, in  FIG. 4  points C and D are located at a plurality of edges of coil  154 , points C and D may be located within coil  154  and coil  156  respectively (i.e., points C and/or D are located in a coiled section of coil(s)  154  and/or  156 ). The exact position of points C and D for example depends entirely on a particular magnet design and the requirements for environment disturbance compensation. FIG.  5  through  FIG. 8  explain in additional detail how to detect these induced currents and how to control/eliminate these currents. 
     FIG. 5  is a detailed illustration of a detection circuit  170  having two parts, one part is connected to points C, D, and E of  FIG. 1 , with two pieces of superconducting wire  176  and  178  wound on a single mandrel in bifilar fashion, the other part is a plurality of quench heaters  174  with a controlling switch  180  and a resistive quench heater power supply  172 . A sensor  182  is positioned to sense electromagnetic fields. When the current either in CC′D′D circuit (I c ) or DD″E′E (I d ) or both starts to flow, and with the aid of detection sensor  182  (either mechanical or electronic as detailed below) and control switch K, quench heaters  174  are energized to heat the superconducting wires CC′D′D and DD″E′E and cause the superconducting wire to quench when current I c  and/or I d  reaches above a predetermined level (e.g., 2 amperes), and thus reduce the electrical currents I c  and I d  to zero, which forces electrical currents in main coil  122  I a  and shield coil  124  I b  to be the same. After sensor  182  detects zero current in I c  and/or in I d , control switch  180  switches off the current in the quench heaters  174 . Thus the electrical currents of main coil  122  and shield coil  124  are the same again. A similar construction is also shown in  FIG. 6  for one coil detection circuit  150  (shown in FIG.  4 ). 
     FIG. 6  illustrates a single coil detection system  190  including a quench heater power supply  192  coupled to a quench heater  194  and a sensor  196  via a switch  198 . When the current in CC′D′D circuit (I c ) starts to flow, and with the aid of detection sensor  196  (either mechanical or electronic as detailed below) and control switch K, quench heater  194  is energized to heat the superconducting wires CC′D′D and cause the superconducting wire to quench when current I c  reaches above a predetermined level (e.g., 2 Amperes), and thus reduce the electrical currents I c  to zero, which forces electrical currents in main coil  154  I a  and shield coil  156  I b  to be the same. After sensor  196  detects a zero current I c  switch  198  switches off the current in quench heater  194 . Thus the electrical currents of main coil  154  and shielding coil  156  are the same again. 
     FIG. 7  is a schematic of a mechanical sensor  200  for detection systems  118  and  150  (e.g., sensors  182  and  196 ), employed in some embodiments. A power source  201  is coupled to a quench heater  202  via wires  208  to a piston assembly  209 . Mechanical sensor  200  includes a solenoid  204  which can be either a bifilar winding (as shown in  FIG. 4 ) or a simple winding (as shown in FIG.  6 ). A plurality of mechanical springs  206  regulate a null level and a trigger level to control a metal piston on/off condition. Mounted within piston assembly  209  is a plurality of pistons  210 . When no net magnetic field disturbances except original magnetic field created by the main and shielding coils present in solenoid  204 , mechanical springs  206  are at a pre-set null level, and metal pistons  210  do not contact a stator, and hence, no current goes through the resistive quench heater(s)  202 . When electrical current reaches a pre-set level (e.g., 2 amps) in solenoid  204  by the environment disturbances, the electromagnetic force on pistons  210  pulls one of the pistons  210  toward the stator, and the quench heater circuit engages, causing the superconducting wires (CC′D′D and/or D′D″E′E) to quench. When the current drops to zero after quench, piston  210  returns to its null position, and the quench circuit is disengaged. In one embodiment, pistons  210  are positioned opposing each other such that current flow in either direction CC′D′D or DD′C′C causes one of pistons  210  to move tow&amp;d a center of assembly  209  to complete the circuit between power supply  201  and heater  202 . In an alternative embodiment, only a single piston  210  is used. 
     FIG. 8  is a schematic of an electronic sensor circuit  220  that is used in detection systems  118  and  150  (e.g., sensors  182  and  196 ), in some embodiments. Circuit  220  includes a quench heater  222  coupled to a power source  224  via a switch  226 . An electronic sensor  228  is positioned within a solenoid  230 . Detection sensor  228  is, in one embodiment, a Hall effect element. In an alternative embodiment, sensor  228  is other means of semiconductor elements or a pickup coil. With the presence of electrical current in solenoid  230 , a net magnetic field fluctuation is detected by sensor  228 . Sensor  228  outputs a related voltage (or a related current) signal to control switch  226  in an on state and an off state. If sensor  228  detects the current in solenoid  230  reaching a predetermined level, the corresponding output signal triggers switch  226  to close, and thus, current flows through quench heater  222 , which starts to heat the superconducting wire to cause the superconducting wire to quench. When sensor  228  detects a zero current in solenoid  230 , switch  226  is opened to de-energize heater  222  allowing any superconductive wires proximate heater  222  to return to a superconductive state. The predetermined level can be set electronically. 
   If the main coils and shielding coils operate on different currents, the above described detection methods and systems are employable with only a slight modification. For example, with both coils operational electrical currents I m , I s  known, and with their respective preset current changing limits known, a ratio of the currents p=(I m /I s ) is determined. Then the number of turns of CC′D′D superconducting wire to the number of turns of DD′E′E superconducting wire can be selected such that (CC′D′D turn number)/(DD′E′E turn number) is equal to p and wound in bifilar fashion, and then the above described methods and apparatus are used to detect environmental disturbances as described above. 
   While the invention has been described in terms of various specific embodiments, those skilled in the art will recognize that the invention can be practiced with modification within the spirit and scope of the claims.

Technology Category: g