Method and apparatus of nuclear magnetic resonance imaging with optimized gradient magnetic field power source

Abstract: A method and an apparatus of nuclear magnetic resonance imaging, capable of minimizing a wasteful unproductive power capacity of the gradient magnetic field power source. A required amount of current is supplied to the gradient magnetic field coil (22) by a basic power source (41) for supplying the current up to a predetermined amount to the gradient magnetic field coil, and by a plurality of power source elements (42), each capable of supplying a prescribed amount of current, for supplementing the basic power source by using a necessary number of the power source elements such that a total amount of current supplied by the basic power source and the power source elements becomes equal to the required amount of current.

Description: The present invention relates to a nuclear magnetic resonance imaging apparatus for obtaining nuclear magnetic resonance images of an object to be examined. As is well known, in a nuclear magnetic resonance imaging apparatus, a nuclear magnetic resonance image is obtained by placing an object to be examined in a static magnetic field; applying a high frequency magnetic field (RF pulse) in a direction perpendicular to that of the static magnetic field in order to induce a nuclear magnetic resonance phenomenon in the object to be examined; superposing gradient magnetic fields G X , G Y , and G Z in X, Y, and Z directions, respectively, onto the static magnetic field, for the sake of tomographic imaging; collecting nuclear magnetic resonance signals due to the induced nuclear magnetic resonance phenomenon from the object to be examined; and image processing the collected nuclear magnetic resonance signals. In such a nuclear magnetic resonance imaging apparatus, the gradient magnetic fields G X , G Y , and G Z are produced by using X-, Y-, and Z-gradient magnetic field coils provided in correspondence with X, Y, and Z axes, respectively, each of which is equipped with an independent power source of the same power capacity. In terms of functions, the gradient magnetic fields can be considered as comprising three orthogonal fields of a slicing gradient magnetic field G S for determining a slicing plane of tomographic imaging, a phase encoding gradient magnetic field G E for providing coordinate information on the slicing plane, and a reading gradient magnetic field G R for tomographic extraction of the nuclear magnetic resonance signals. These gradient magnetic fields are obtained as a field given by superposition of three orthogonal gradient magnetic fields G X , G Y , and G Z in X, Y , and Z directions, respectively. For example, a total gradient magnetic field G ⊘ shown in Fig. 1 can be obtained from three orthogonal gradient magnetic fields G X , G Y , and G Z , and this total gradient magnetic field G ⊘ can be taken as being composed from three components corresponding to the slicing gradient magnetic field G S , the phase encoding gradient magnetic field G E , and the reading gradient magnetic field G R . Now, in a conventional nuclear magnetic resonance imaging apparatus, each of X-, Y-, and Z-gradient magnetic field coils is equipped with an independent power source of the same power capacity, so that, by assuming that a maximum power of each power source to be 1, a total power capacity of these power sources is equal to 3. However, in order to be able to take an image at an arbitrary cross section, a maximum total power required from these power sources is at most 1/√3 × 3 = √3 ≈ 1.73 occurring in a case of a total gradient magnetic field obliquely inclined by 45° from all of X, Y, and Z axes. Thus, in a conventional nuclear magnetic resonance imaging apparatus, over 40% of the total power capacity of the power sources for the gradient magnetic field coils has always been wasted as unproductive power capacity. This situation is particularly problematic when using a modern imaging technique such as an echo planer method in which the required power for the reading gradient magnetic field is much larger than the required powers for the slicing gradient magnetic field and the phase encoding gradient magnetic field, because in such a case, the conventional provision of providing three independent power sources of the same power capacities produces a waste of a very large amount of power. EP 0 250 718 discloses a variable current source comprising, for a rapid and low-loss setting of a voltage at an inductive load necessary for the generation, the upholding and the reduction of desired current impulses a so-called "step switch" and a variable voltage source. The step switch is composed of groups of voltage sources serially arranged, wherein the output voltages of all sources of each group have the same amount. A switch network composed of switch groups connected thereto is controlled by a control device in such a manner that the load voltage necessary for the generation/decrease, and for the upholding of a respective load current pulse is set by sequentially coupling a suitable combination of voltage sources. For fine tuning the exact amount of the load voltage, the adjustable voltage source is also controlled by said control device. This structure of the voltage supply may be used for feeding further current supply systems. GB-A-2 208 768 is directed to the eddy current correction of electromagnetic fields in a magnetic resonance imager, wherein an eddy current calibration circuit comprising a plurality of parallel individual calibration circuits modifies current pulse profiles to remove the effect of eddy currents such that the resultant gradient magnetic field has a selected profile. An automatic calibration circuit adjusts the decay constant or center frequency of a band pass filter and the gain of an amplifier. A sine wave generator applies a sine wave voltage to the gradient coils as a search coil monitors the resultant magnetic field and a current probe monitors the current. The magnetic field and current are phase sensitively detected and the relative amplitude and phase are compared as the frequency of the sine wave signal is swept. It is an object of the present invention to provide a method and an apparatus of nuclear magnetic resonance imaging, capable of minimizing a wasteful unproductive power capacity of the gradient magnetic field power source. According to one aspect of the present invention there is provided a nuclear magnetic resonance imaging apparatus, comprising the features mentioned in claim 1. According to another aspect of the present invention there is provid d a method of nuclear magnetic resonance imaging as defined in claim 4. Other features and advantages of the present invention will become apparent from the following description taken in conjunction with the accompanying drawings. Fig. 1 is a vector diagram for explaining gradient magnetic fields in a conventional nuclear magnetic resonance imaging apparatus. Fig. 2 is a schematic block diagram of one embodiment of a nuclear magnetic resonance imaging apparatus according to the present invention. Fig. 3 is a detailed block diagram of a main portion of the embodiment of a nuclear magnetic resonance imaging apparatus of Fig. 2. Fig. 4 is a block diagram for a basic power source unit of the embodiment of a nuclear magnetic resonance imaging apparatus of Fig. 2. Fig. 5 is a vector diagram for explaining gradient magnetic fields generated by the embodiment of a nuclear magnetic resonance imaging apparatus of Fig. 2. Fig. 6 is a schematic block diagram of another embodiment of a nuclear magnetic resonance imaging apparatus according to the present invention. Fig. 7 is a detailed block diagram of a main portion of the embodiment of a nuclear magnetic resonance imaging apparatus of Fig. 6. Referring now to Fig. 2, there is shown one embodiment of a nuclear magnetic resonance imaging apparatus according to the present invention. In this embodiment, the nuclear magnetic resonance imaging apparatus 1 comprises a magnet unit 2 containing a static magnetic field coil 21, gradient magnetic field coils 22, a nuclear magnetic resonance (NMR) signal detector coil 23, and an RF pulse transmitter coil 24, all of which are provided on a coil core 20; a gradient magnetic field power source 4 for supplying currents to the gradient magnetic field coils 22, which includes a basic power source unit 41 and a power source elements unit 42 to be described in detail below; an RF pulse transmitter 5 for applying RF pulses to the RF pulse transmitter coil 24; a nuclear magnetic resonance signal receiver 6 for receiving the nuclear magnetic resonance signals detected by the nuclear magnetic resonance signal detector coil 23; a computer unit 7 for controlling operations of the apparatus; a sequencer 8 for controlling operations of the gradient magnetic field power source 4 and the RF pulse transmitter 5; a console 9 including a CRT display for displaying nuclear magnetic resonance images and a keyboard for entering commands such as imaging conditions; a static magnetic field controller 10 for controlling the static magnetic field coil 21, and a power source switching unit 11 containing a plurality of switches for selectively controlling the power source elements unit 42 in order to control the currents supplied to the gradient magnetic field coils 22. In detail, as shown in Fig. 3, the gradient magnetic field coils 22 comprises an X-axis gradient magnetic field coil 22a, a Y-axis gradient magnetic field coil 22b, and a Z-axis gradient magnetic field coil 22c, while the basic power source unit 41 comprises an X-axis basic power source 41a, a Y-axis basic power source 41b, and a Z-axis basic power source 41c, in correspondence with the gradient magnetic filed coils 22. Each of these X-, Y-, and Z-axis basic power sources has 1/10 of a power capacity of a single conventional gradient magnetic field coil power source for one gradient magnetic field coil which is assumed to be 1. Meanwhile, the power source elements unit 42 comprises a plurality of power source elements 42a such as a power source element-1, power source element-2, etc. Each power source element 42a of the power source elements unit 42 has the power capacity of less than 1/10. Combining the basic power source unit 41 and the power source elements unit 42 together, the gradient magnetic field power source 4 as a whole has the power capacity of approximately 1.73 times a maximum required power capacity for each of X-, Y-, and Z-channels. When the imaging conditions are specified at the console 9, in accordance with the specified imaging conditions, the computer unit 7 controls the sequencer 8 which in turn controls the gradient magnetic field power source 4 and the power source switching unit 11, such that the appropriate amounts of currents are supplied to the gradient magnetic field coils 22 by using the basic power source unit 41 and a necessary number of power source elements 42a of the power source elements unit 42. The gradient magnetic field power source 4 further includes a signal switching unit 12 for converting a current C2 outputted by the power source elements unit 42 indicating an amount of current presently supplied by the power source elements unit 42 into a current level signal S2 to be given to the basic power source unit 41, and transmitting a current supply request signal 51 indicating an amount of current needed to be supplied from the power source elements unit 42, from the basic power source unit 41 to the power source elements unit 42. Now, as shown in Fig. 4, the X-axis basic power source 41a contains a feed-back configuration formed by an amplifier A1 and an adder A2. The current level signal 52 from the power source elements unit 42 and an output current level signal S3 indicating an amount of output current C0 of the amplifier A1 are added by the adder A2 and this sum is subtracted from a control signal S0 coming from the sequencer 8 indicating an amount of the current to be supplied to the X-axis gradient magnetic field coil 22a. The difference thus obtained is fed to the power source elements unit 42 as the current supply request signal S1, so that a necessary number of the power source elements 42a of the power source elements unit 42 can be employed to supplement the output current C0 of the X-axis basic power source 41a, and an appropriate amount of current can be supplied to the X-axis gradient magnetic field coil 22a by the gradient magnetic field power source 4 as a whole. The Y-axis basic power source 41b and the Z-axis basic power source 41c are constructed similarly. Thus, this apparatus 1 operates as follows. First, the imaging conditions are entered from the console 9, and under the control of the computer unit 7, the sequencer 8 controls the gradient magnetic field power source 4 and the power source switching unit 11, so as to supply an appropriate amount of currents to the gradient magnetic field coils 22. For example, as shown in Fig. 5, when the gradient magnetic field to be produced is GbO, such an amount of current is supplied to the X-axis gradient magnetic field coil 22a that an X-axis gradient magnetic field Gb X is produced, while such an amount of current is supplied to the Y-axis gradient magnetic field coil 22b that a Y-axis gradient magnetic field Gb Y is produced, and such an amount of current is supplied to the Z-axis gradient magnetic field coil 22c that an Z-axis gradient magnetic field Gb Z is produced. Here, as shown in Fig. 5, each of the X-axis gradient magnetic field Gb X , Y-axis gradient magnetic field Gb Y , and Z-axis gradient magnetic field Gb Z can be increased in units of an increment ΔG corresponding to the addition of one power source element 42a from the power source elements unit 42, so that any other desired gradient magnetic field can be obtained by using suitable numbers of the power source elements 42a for each gradient magnetic field component. As another example, when the gradient magnetic field to be produced is along the Z-axis and of magnitude 1, as in a case of applying a slicing gradient magnetic field for a slicing plane normal to the Z-axis, since the Z-axis basic power source 41c has the power capacity of only 1/10, nine additional power source elements of power capacity 1/10 are employed from the power source elements unit 42 for the Z-axis gradient magnetic field coil 22c by the power source switching unit 11. Thus, according to this embodiment, the gradient magnetic field power source 4 as a whole can be operated optimally, with a minimum amount of unproductive power capacity, by supplementing the basic power source unit 41 with a necessary number of power source elements 42a of the power source elements unit 42. This enable to limit the maximum total power capacity of the gradient magnetic field power source to be 1.73, in contrast to the conventional configuration which has the maximum total power capacity 3. Also, for this reason, even in a case of using a modern imaging technique such as an echo planer method in which the required power for the reading gradient magnetic field is much larger than the required powers for the slicing gradient magnetic field and the phase encoding gradient magnetic field, there is no need to provide a very powerful power source for each of the X-, Y-, and Z-gradient magnetic field coils. Referring now to Fig. 6, there is shown another embodiment of a nuclear magnetic resonance imaging apparatus according to the present invention. In this embodiment, the nuclear magnetic resonance imaging apparatus 1a comprises a magnet unit 2a containing a static magnetic field coil 21, gradient magnetic field coils 50, a nuclear magnetic resonance (NMR) signal detector coil 23, and an RF pulse transmitter coil 24, all of which are provided on a coil core 20; a gradient magnetic field power source 14 for supplying currents to the gradient magnetic field coils 50, which includes a basic power source unit 43 and a power source elements unit 44 to be described in detail below; an RF pulse transmitter for applying RF pulses to the RF pulse transmitter coil 24; a nuclear magnetic resonance signal receiver 6 for receiving the nuclear magnetic resonance signals detected by the nuclear magnetic resonance signal detector coil 23; a computer unit 37 for controlling operations of the apparatus; a sequencer 38 for controlling operations of the gradient magnetic field power source 14 and the RF pulse transmitter 5; a console 9 including a CRT display for displaying nuclear magnetic resonance images and a keyboard for entering commands such as imaging conditions; a static magnetic field controller 10 for controlling the static magnetic field coil 21, and a power source switching unit 31 containing a plurality of switches for selectively switching the power source elements 42 in order to control the currents supplied to the gradient magnetic field coils 50. Here, those structural elements having the same labels as in the previous embodiment are identical to the structural elements of the previous embodiments, while those structural elements having the same names as in the previous embodiment but with different labels are playing similar roles as the corresponding structural elements of the previous embodiments but with some modifications. In detail, as shown in Fig. 7, the gradient magnetic field coils 50 comprises an X-axis gradient magnetic field coils 51, a Y-axis gradient magnetic field coils 52, and a Z-axis gradient magnetic field coils 53, where each of these X-, Y-, and Z-axis gradient magnetic field coils 51, 52, and 53 contain a plurality of mutually parallel individual coils 51a, 52a, and 53a, respectively, such as an X-axis gradient magnetic field coil-1, Y-axis gradient magnetic field coil-1, 2-axis gradient magnetic field coil-1, etc. On the other hand, the basic power source unit 43 comprises an X-axis basic power source 43a, a Y-axis basic power source 43b, and a Z-axis basic power source 43c, in correspondence with the X-, Y-, and Z-axis gradient magnetic filed coils 51, 52, and 53, respectively. Each of these X-, Y-, and Z-axis basic power sources has 1/10 of a power capacity of a single conventional gradient magnetic field coil power source for one gradient magnetic field coil which is assumed to be 1. Meanwhile, the power source elements unit 44 comprises a plurality of power source elements 44a such as a power source element-1, power source element-2, etc. Each power source element 44a of the power source elements unit 44 has the power capacity of less than 1/10. Combining the basic power source unit 43 and the power source elements unit 44 together, the gradient magnetic field power source 14 has the power capacity of approximately 1.73. Also, as shown in Fig. 7, the X-, Y-, and Z-axis basic power source 43a, 43b, and 43c are directly connected to the first individual coil of the X-, Y-, and Z-axis gradient magnetic field coils 51, 52, and 53, respectively, i.e., to the X-axis gradient magnetic field coil-1, Y-axis gradient magnetic field coil-1, and Z-axis gradient magnetic field coil-1, whereas the remaining individual coils, such as the X-axis gradient magnetic field coil-2, Y-axis gradient magnetic field coil-2, Z-axis gradient magnetic field coil-2, etc., are directly and separately connected to the power source switching unit 31. Consequently, in this embodiment, each individual coil of the gradient magnetic field coils 50 is separately incorporated in an independent closed circuit configuration. The operation of the apparatus 1a of this embodiment is substantially the same as that of the previous embodiment, where each structural elements are functioning in the manner similar to the corresponding structural elements of the previous embodiment. Namely, when the imaging conditions are specified at the console 9, in accordance with the specified imaging conditions, the computer unit 37 controls the sequencer 38 which in turn controls the gradient magnetic field power source 14 and the power source switching unit 31, such that the appropriate amounts of currents are supplied to the gradient magnetic field coils 50 by using the basic power source unit 43 and a necessary number of power source elements 44a of the power source elements unit 44. The gradient magnetic field power source 14 further includes a signal switching unit 32 for converting a current C2 outputted by the power source elements unit 44 indicating an amount of current presently supplied by the power source elements unit 44 into a current level signal S2 to be given to the basic power source unit 43, and transmitting a current supply request signal S1 indicating an amount of current needed to be supplied from the power source elements unit 44, from the basic power source unit 43 to the power source elements unit 44. The X-, Y-, and Z-axis basic power sources 43a, 43b, and 43c has a configuration similar to that shown in Fig. 4 for the previous embodiment. It is obvious that, as in the previous embodiment, each of the X-axis gradient magnetic field, Y-axis gradient magnetic field, and Z-axis gradient magnetic field can be increased in units of an increment ΔG corresponding to the addition of one power source element from the power source elements unit 42, so that any other desired gradient magnetic field can be obtained by using suitable numbers of the power source elements for each gradient magnetic field component. Thus, as in the previous embodiment, according to this embodiment, the gradient magnetic field power source 14 as a whole can be operated optimally, with a minimum amount of unproductive power capacity, by supplementing the basic power source unit 43 with a necessary number of power source elements of the power source elements unit 44. This enable to limit the maximum total power capacity of the gradient magnetic field power source to be 1.73, in contrast to the conventional configuration which has the maximum total power capacity 3. Also, for this reason, even in a case of using a modern imaging technique such as an echo planer method in which the required power for the reading gradient magnetic field is much larger than the required powers for the slicing gradient magnetic field and the phase encoding gradient magnetic field, there is no need to provide a very powerful power source for each of the X-, Y-, and Z-gradient magnetic field coils. In addition, in this embodiment, because each individual coil of the gradient magnetic field coils 50 is separately incorporated in an independent closed circuit configuration, an interference among current supply lines from the basic power source unit 43 and the power source elements unit 44 can be avoided so that the gradient magnetic field can be produced stably, and a practical problem of making a stable joint between the current supply lines from the basic power source unit 43 and the power source elements unit 44 can be circumvented.

Claims: A nuclear magnetic resonance imaging apparatus, comprising: means (10, 21) for generating a static magnetic field; X-, Y-, and Z-gradient magnetic field coil means (22) for generating gradient magnetic fields to be superposed on the static magnetic field in X, Y, and Z directions; power source means (4) for supplying a required amount of current to the gradient magnetic field coil means (22); means (5, 24) for applying RF pulses onto an object to be examined placed in the static magnetic field; means (6, 23) for collecting nuclear magnetic resonance signals from the object to be examined resulting from the RF pulses, and means (7) for processing the collected nuclear magnetic resonance signals; wherein said power source means (4) includes: a basic power source (41, 43), provided for each of the gradient magnetic field coil means (22), for supplying the current up to a predetermined amount to the gradient magnetic field coil means (22); and a plurality of power source elements (42, 44), each capable of supplying a prescribed amount of current, for supplementing the basic power sources (41, 43) by a necessary number of the power source elements (42, 44) such that a total amount of current supplied by each basic power source (41, 43) and the necessary number of power source elements (42, 44) becomes equal to the required amount of current as a sum of currents supplied by the power source elements (42, 44) as added to a current due to the basic power source (41, 43), each basic power source (41, 43) containing a feed-back configuration (A1, A2) comprising an amplifier (A1) and an adder (A2) which adds an output current level (S3) indicating an amount of output current (C 0 ) of the amplifier (A1) and a current level signal (S2) indicating the amount of current presently supplied by the power source elements, the result of addition being subtracted from a control signal (S0) indicating the amount of current to be supplied to the associated gradient magnetic field coil means (22), the result of the subtraction being applied to the input of the amplifier (A1). The apparatus of claim 1, wherein the power source means (4) has a total power capacity equal to an amount of current capable of making the X-, Y-, and Z-gradient magnetic field coil means (22) to generate a total gradient magnetic field of maximum magnitude. The apparatus of claim 1 or 2, wherein each of the X-, Y-, and Z-gradient magnetic field coil means (22) includes a plurality of mutually parallel individual coils (51a, 52a, 53a) each individual coil being separately incorporated in an independent closed circuit configuration by being directly connected with one of the basic power source (43) and the power source elements. A method of nuclear magnetic resonance imaging, comprising the steps of: generating a static magnetic field; generating gradient magnetic fields to be superposed on the static magnetic field in X, Y, and Z directions by X-, Y-, and Z-gradient magnetic field coil means (22); applying RF pulses onto an object to be examined placed in the static magnetic field; collecting nuclear magnetic resonance signals from the object to be examined resulting from the RF pulses; and processing the collected nuclear magnetic resonance signals; wherein a required amount of current is supplied to the gradient magnetic field coil means (22) by the steps of: supplying the current up to a predetermined amount to the gradient magnetic field coil means (22) by a basic power source (41, 43) provided for each of the gradient magnetic field coil means (22); and supplementing the basic power sources (41, 43) by a necessary number of power source elements (42, 44), each capable of supplying a prescribed amount of current, such that a total amount of current supplied by each basic power source (41, 43) and the necessary number of power source elements (42, 44) becomes equal to the required amount of current as a sum of currents supplied by the power source elements (42, 44) as added to a current due to the basic power source (41, 43), each basic power source (41, 43) containing a feed-back configuration (A1, A2) comprising an amplifier (A1) and an adder (A2), wherein an output current level (S3) indicating an amount of output current (C 0 ) of the amplifier (A1) is added to a current level signal (S2) indicating the amount of current presently supplied by the power source elements, the result of addition being subtracted from a control signal (S0) indicating the amount of current to be supplied to the associated gradient magnetic field coil means (22), the result of the subtraction being applied to the input of the amplifier (A1). The method of claim 4, wherein the basic power source (41, 43) and the power source elements (42, 44) together have a total power capacity equal to an amount of current capable of making the X-, Y-, and Z-gradient magnetic field coil means (22) to generate a total gradient magnetic field of maximum magnitude. The method of claim 4 or 5, wherein each of the X-, Y-, and Z-gradient magnetic field coil means (22) includes a plurality of mutually parallel individual coils (51a, 52a, 53a), each individual coil being separately incorporated in an independent closed circuit configuration by being directly connected with one of the basic power source (43) and the power source elements (44a).

Inventors: YOKOI MOTOHISA, YOKOI, MOTOHISA
Applicants: TOSHIBA KK, KABUSHIKI KAISHA TOSHIBA
Application year: 1990
Main classification: G01R33
IPC codes: A61B5, G01R33