JP2018183292A - Magnetic resonance imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide an MRI apparatus capable of continuous irradiation of selective saturation pulses while preventing a decrease in the maximum output of a power source. A transmission coil includes a plurality of channels, and a power supply has a plurality of systems. The system is connected to one or more of the channels, and the maximum output is limited according to the duty ratio. By supplying high-frequency power pulses to a plurality of channels in order from such a system and irradiating the high-frequency magnetic field pulses in order from a plurality of channels, a plurality of continuous high-frequency magnetic fields are continuously emitted in a non-irradiation time of a predetermined value or less for the entire transmission coil. The number of systems is set so that the duty ratio of each system becomes equal to or less than a predetermined value when a pulse is applied. [Selection diagram] Fig. 3

Description

  The present invention relates to a measurement technique using a magnetization transfer method or a saturation transfer method in a magnetic resonance imaging apparatus.

  A magnetic resonance imaging (MRI) apparatus is a nuclear magnetic resonance phenomenon in which hydrogen nuclei contained in a subject placed in a strong and stable magnetic field (static magnetic field) resonate with an electromagnetic wave (high frequency magnetic field) of a specific frequency. Is a diagnostic apparatus that images information inside a subject. Most of the signals from hydrogen nuclei that can be detected by living bodies are obtained from water. Since the signal strength of water changes due to various factors such as the water density in the living tissue and the change rate (relaxation time) of the nuclear magnetic resonance state, information such as structural changes and changes in properties of the living tissue caused by diseases, etc. It is possible to measure and image as a change in water signal intensity. Various MR imaging methods have been proposed depending on the type of information to be detected.

  In recent years, an MR imaging method called chemical exchange dependent transfer (CEST) imaging has attracted attention, and it enables mapping of in vivo compounds (metabolites) and detection of in vivo information such as pH. It is expected. The CEST imaging is an imaging method based on a contrast principle that has been known as a magnetization transfer method or a saturation transfer method. Whereas MRS (Magnetic Resonance Spectroscopy) and MRSI (Magnetic Resonance Spectroscopic Imaging), which are conventional metabolite measurement methods, directly detect information on in-vivo metabolites that are low-concentration substances, CEST imaging has a high concentration. This is a method that allows in-vivo metabolites to be observed with high sensitivity indirectly through changes in the water signal.

  As a specific measurement method of CEST imaging, first, as a prepulse, a plurality of subjects are irradiated with a high frequency magnetic field pulse (selective saturation pulse) having a resonance frequency of a chemically exchangeable proton contained in a target compound (narrow band). Irradiation) to saturate chemically exchangeable protons. For example, 40 subjects are irradiated with a Gaussian pulse having a pulse length of 35 ms as a selective saturation pulse. Next, selectively saturated protons are replaced by protons in the surrounding free water by chemical exchange, and the magnetic resonance signal from the free water protons decreases (negative contrast), which is converted into a normal MRI pulse sequence. (For example, a one-shot two-dimensional fast spin echo pulse sequence). By repeating this measurement while continuously changing the irradiation frequency of the selective saturation pulse, two peaks showing a signal decrease at the resonance frequency of the proton that can be chemically converted in the target compound and the resonance frequency of free water. A characteristic (multimodal) signal change curve (Z spectrum) can be obtained.

  For example, when the Amide group proton contained in the egg white of a chicken egg is measured using the CEST sequence, the transmission frequency of the pre-pulse (selective saturation pulse) is “the resonance frequency (average value) of the Amide group proton contained in the egg white + 3. Set to 5 ppm "and measure the decrease in magnetic resonance signal. By repeating this measurement while continuously changing the transmission frequency of the selective saturation pulse, the resonance frequency “+3.5 ppm” of the Amide group proton and the resonance frequency “0 ppm” of the free water proton are asymmetrical indicating a signal drop. A Z spectrum with varying properties is obtained.

In addition, when measuring the Amide group proton contained in the egg white of the egg using the CEST sequence, the difference signal of the magnetic resonance signal of the signal when the transmission frequency of the selective saturation pulse is −3.5 ppm and +3.5 ppm It is also possible to generate an image called an APT weighted image using an APT (Amide Proton Transfer) signal. The APT signal is expressed by Expression (1).
(Equation 1)
APT signal = (S _{sat−3.5 ppm−} S _{sat + 3.5 ppm} ) / S ₀ (1)
Here, S _sat -3.5 _ppm and S _sat +3.5 ppm are magnetic resonance signals measured when the transmission frequency of the saturation pulse is −3.5 ppm and +3.5 ppm, respectively, and S ₀ is saturated. This is a signal imaged with the pulse turned off.

  When CEST imaging is performed on a chicken egg to obtain an APT-weighted image, pixels showing a strong signal value only in the egg white portion are observed in the APT-weighted image, and the signal increase due to APT is observed in the egg white portion including the Amide group proton. It can be confirmed that only has occurred.

  Since the intensity of the magnetic resonance signal obtained by CEST imaging decreases in proportion to the irradiation energy of the selective saturation pulse until the saturation state is reached, many pulses having a relatively long irradiation time are transmitted as the selective saturation pulse. The At this time, since it is desirable not to generate a non-transmission time (dead time) as much as possible in the transmission interval of the selective saturation pulse, Non-Patent Document 1 uses a two-channel independent drive coil as a transmission coil. Are connected to separate power supply devices, and a configuration is disclosed in which alternating transmission is performed from a two-channel coil.

J. et al. Keupp et al. Proc. Intl. Soc. Mag. Reson. Med. , 19, 2011.

  As described above, the signal intensity of chemically exchangeable protons is reduced (saturated) in proportion to the irradiation energy of the selective saturation pulse. Therefore, in order to achieve complete saturation in a short time, a large irradiation energy is selected. It is desirable to irradiate the saturation pulse continuously with almost no dead time (non-irradiation time).

  However, a high-frequency power source of a general MRI apparatus can continuously irradiate a high-frequency pulse with less energy from a transmission coil, but continuously irradiates a high-power high-frequency pulse with little dead time. It is not designed to be used. Therefore, when a high frequency power supply for a general MRI apparatus continuously outputs a high frequency power pulse for a selective saturation pulse for CEST imaging, the output tends to gradually decrease. When the output of the high-frequency power supply is gradually reduced in this manner, it is impossible to irradiate a selective saturation pulse of desired energy, and it is not possible to saturate chemically exchangeable protons within a predetermined time.

  An object of the present invention is to provide an MRI apparatus that enables continuous irradiation of a selective saturation pulse while preventing a decrease in the maximum output of a power source.

  In order to achieve the above object, an MRI apparatus of the present invention includes a transmission coil that irradiates a subject with a high-frequency magnetic field pulse, and a power source that is connected to the transmission coil and supplies a high-frequency power pulse. The transmission coil includes a plurality of channels, the power supply has a plurality of systems, the systems are respectively connected to one or more of the channels, and the maximum output is limited according to the duty ratio. By supplying high-frequency power pulses sequentially to a plurality of channels from such a system and sequentially irradiating the high-frequency magnetic field pulses from the plurality of channels, a plurality of high-frequency magnetic fields are continuously generated in a non-irradiation time equal to or less than a predetermined value as the entire transmission coil The number of systems is set so that the duty ratio of each system is less than or equal to a predetermined value when pulses are irradiated.

  ADVANTAGE OF THE INVENTION According to this invention, the MRI apparatus which enables continuous irradiation of a selective saturation pulse can be provided, preventing the fall of the maximum output of a power supply.

It is a block diagram which shows the structure of the MRI apparatus of one Embodiment of this invention. (A) is an example of the transmission coil of the embodiment, and is a diagram showing a four-channel coil in which four coils constitute four channels, and a power supply system connected to each channel, (b) ) Is a diagram showing a four-channel coil in which four elements operate synchronously as one channel. (A) is a figure of the pulse sequence which shows the output pulse for every channel in the case of performing simultaneous irradiation of the selective saturation pulse from all the channels of the 4-channel coil of a comparative example, (b) is 4 of this embodiment. It is a figure of the pulse sequence which shows the output pulse for every channel at the time of irradiating the selective saturation pulse sequentially from each channel using a channel coil, (c) is 2 using the 4 channel coil of this embodiment. It is a figure of the pulse sequence which shows the output pulse for every channel in the case of transmitting alternately channel by channel. (A) is an external view of the horizontal magnetic field type MRI apparatus of the embodiment, (b) is an external view of the vertical magnetic field type MRI apparatus of the embodiment, and (c) is an open feeling of the embodiment. FIG. It is a figure which shows an example of a CEST pulse sequence. It is a flowchart at the time of implementing this embodiment. It is a figure which shows the example of a display screen with the display apparatus which receives an input from the operator at the time of implementing this embodiment. (a) And (b) is a figure which shows the magnetic field distribution of the selective saturation pulse of a comparative example, respectively (c)-(f) of an Example.

  An embodiment of the present invention will be described.

<< First Embodiment >>
Hereinafter, a first embodiment of the present invention will be described with reference to the drawings.

  As shown in FIG. 1, the MRI apparatus of the first embodiment includes at least a transmission coil 5 that irradiates a subject 1 with a high-frequency magnetic field pulse, and a power source 7 that is connected to the transmission coil and supplies a high-frequency power pulse. I have. The transmission coil 5 includes a plurality of channels 5-1 to 5-4, for example, as shown in FIG. The power supply 7 has a plurality of systems (for example, 7-1 to 7-4). The number of systems 7-1 to 7-4 may be different from the number of channels 5-1 to 5-4 as long as it is plural. The systems 7-1 to 7-4 are connected to one or more of the channels 5-1 to 5-4, respectively. Further, the systems 7-1 to 7-4 are limited in maximum output according to the duty ratio.

  In addition, the systems 7-1 to 7-4 referred to here have a structure in which the respective output values do not affect each other, that is, one system can output the maximum output from another system while the system outputs the maximum output. If it is. Therefore, each system may have a completely electrically independent configuration, or may have a structure in which a part is shared.

The systems 7-1 to 7-4 supply high-frequency power pulses to the connected channels, respectively. In this case, for example, as shown in FIG. 3 (b), the entire transmitting coil 5, so that a plurality onset continuously high frequency magnetic field pulses at a predetermined value or less of non-irradiation time T _D is irradiated, the system 7-1 When the high frequency power pulse is supplied to the channels (5-1 to 5-4) connected from 7-4, the system 7 so that the duty ratios of the systems 7-1 to 7-4 are not more than a predetermined value. Numbers of -1 to 7-4 are set. In other words, the channels (5-1 to 5--5) are set so that the duty ratio of the high-frequency power pulses supplied to the channels (5-1 to 5-4) is not more than a predetermined value as shown in FIG. The number of 4) is set. The above-described duty ratio is desirably 25% or less.

  Thus, in the MRI apparatus of this embodiment, paying attention to the fact that the maximum output of the power supply system changes depending on the duty ratio of the high-frequency power pulse output from the power supply system and input to the channel, The ratio is set to a predetermined value or less to suppress the reduction in maximum output. At this time, in this embodiment, the non-irradiation time of the high-frequency magnetic field pulse is not increased, but the number of channels (number of systems) of the transmission coil is increased. As a result, a high-frequency magnetic field pulse having a large energy, such as a selective saturation pulse for CEST imaging, can be continuously emitted as a whole for the transmission coil 5.

The normal high-frequency power supply 7 has an upper limit of 50 on the duty ratio (= τ / T, where T is the time of one cycle when high-frequency power pulses are continuously transmitted and τ is the continuous transmission time of high-frequency power pulses). There are many specifications of about 90%. However, in practice, this upper limit depends on the output intensity. When a plurality of high-frequency high-frequency magnetic field pulses such as a selective saturation irradiation pulse for CEST imaging are continuously transmitted, the maximum output that can be output is the duty ratio. Determines the upper limit. An example of the relationship between the maximum output possible and the duty ratio is shown in Table 1. As is clear from Table 1, when the duty ratio increases, the maximum output decreases greatly. For this reason, when a high-power high-frequency power pulse is output from the power source with a large duty ratio, the output energy actually gradually decreases. Specifically, in the power supply shown in Table 1, when the maximum output when the duty ratio is 10% is 100%, the maximum output when the duty ratio is 50% is reduced to 10%.

  In the MRI apparatus of the present embodiment, in consideration of this phenomenon, the number of systems (7-1 to 7-4) is set so that the duty ratio of each system (7-1 to 7-4) is not more than a predetermined value. Since it is set, the duty ratio of each system can be suppressed to a predetermined value or less, and the high frequency power pulse of the output required from each system (7-1 to 7-4) to the channel (5-1 to 5-4) Can be output. As a result, a plurality of high frequency magnetic field pulses having desired energy as selective saturation pulses for CEST imaging can be continuously emitted as a whole as shown in FIG. 3B.

  In the MRI apparatus of the present embodiment, the number of channels (5-1 to 5-4) and the number of systems (7-1 to 7-4) may be fixed numbers determined at the time of design, The configuration may be such that the selection unit (computer 9 described later) selects a channel for supplying a high-frequency power pulse from the system so as to satisfy a predetermined duty ratio condition from among a plurality of channels and systems. When the number of channels and the number of systems are fixed numbers determined at the time of design, as an example, the number of systems and the number of channels are each set to 4 or more, and each channel is connected to the system. Further, when the selection unit (computer 9 to be described later) selects a channel for supplying high-frequency power pulses from the system, as an example, the transmission coil 5 includes a plurality of channels, and the power supply 7 includes a plurality of systems. Is configured to be connected to each channel, and the selection unit is configured to set the number of channels for supplying high-frequency power pulses from the system so that the duty ratio of the high-frequency power pulses supplied to the channels is a predetermined value or less. To do.

  Hereinafter, the overall structure of the MRI apparatus of this embodiment will be described.

  FIG. 1 is a block diagram showing the configuration of the MRI apparatus of this embodiment. The MRI apparatus includes, in addition to the transmission coil 5 and the power source 7 described above, a static magnetic field coil 2 that generates a static magnetic field in a space where the subject 1 is placed, and a gradient magnetic field coil for applying gradient magnetic fields in three directions orthogonal to each other. 3, a receiving high-frequency coil 6 that receives a magnetic resonance signal generated from the subject 1 (hereinafter simply referred to as a receiving coil), and a shim coil 4 that adjusts the static magnetic field uniformity. Note that the shim coil 4 is not necessarily provided.

  Various types of static magnetic field coils 2 are employed according to the structure of the MRI apparatus. For example, FIG. 4A shows an example of a horizontal magnetic field type MRI apparatus 100 that generates a static magnetic field using a solenoid coil (tunnel magnet) as the static magnetic field coil 2. FIG. 4B shows an example of a hamburger type (open type) vertical magnetic field type MRI apparatus 200 in which magnets (static magnetic field coil 2) are separated into upper and lower sides in order to enhance the feeling of opening. FIG. 4C shows an MRI apparatus using the same tunnel type magnet (static magnetic field coil 2) as that in FIG. 4A. However, by shortening the depth of the magnet and tilting it obliquely, the feeling of opening is enhanced. This is an example of the MRI apparatus 300. These are merely examples, and the MRI apparatus of the present embodiment is not limited to these forms, and various known forms of MRI apparatuses can be employed regardless of the form and type of the apparatus.

  A gradient magnetic field power supply 12 and a shim power supply 13 are connected to the gradient magnetic field coil 3 and the shim coil 4, respectively, and by supplying electric power thereto, the gradient magnetic field and the static magnetic field equalization are placed in the space where the subject 1 is arranged. Apply a magnetic field to correct once. The transmission coil 5 is supplied with a high-frequency power pulse from a high-frequency power supply (transmitter) 7, and the subject 1 is irradiated with a high-frequency magnetic field pulse from the transmission coil 5. The receiving coil 6 detects a nuclear magnetic resonance signal generated by the subject 1, and the detected signal is sent to the computer 9 through the receiver 8. In the present embodiment, a case where separate transmission coils 5 and reception coils 6 are used will be described as an example. However, the transmission coil 5 and the reception coil 6 are configured as a single coil. Also good.

  The computer 9 performs various arithmetic processes on the magnetic resonance signal obtained via the receiver 8 in accordance with a program registered in advance or an instruction from the user, and generates spectrum information and image information. A display device 10, a storage device 11, a sequence control device 14, an input device 15 and the like are connected to the computer 9. The display device 10 is an interface that displays spectrum information and image information generated by the computer 9 to an operator. The input device 15 is an interface for an operator to input meter measurement conditions, conditions necessary for arithmetic processing, parameters, and the like. In the storage device 11, spectrum information and image information generated by the computer 9, information input via the input device 15, and the like are recorded as necessary.

  The sequence controller 14 controls the operations of the gradient magnetic field power supply 12, the shim power supply 13, the high frequency magnetic field power supply 7 and the receiver 9, and the timing of application of gradient magnetic field pulses, high frequency magnetic field pulses and reception of nuclear magnetic resonance signals. In addition, the application amount is controlled. The timing is controlled according to a time chart called a pulse sequence preset by the imaging method. Detailed conditions such as the selection of the pulse sequence to be used and the amount of each application are registered in advance in the storage device 11 as a program or instructed by the user via the input device 15.

  Next, a structural example of the transmission coil 5 of the present embodiment will be described. In the present embodiment, as the transmission coil 5, a multi-channel coil provided with a plurality of channels each independently transmitting a high frequency magnetic field (RF) to the subject 1 is used. In the transmission coil 5 in FIG. 2A described above, four coils of channels 5-1, 5-2, 5-3, and 5-4 constitute four channels. When the subject 1 arranged as shown in FIG. 2A is irradiated with a high-frequency magnetic field pulse using the transmission coil 5 having this configuration, the upper right portion is a channel when viewed from the top of the subject 1. The high frequency magnetic field generated by the coil of 5-1 becomes a dominant region, the lower right part is channel 5-2, the lower left part is channel 5-3, and the upper left part is the high frequency magnetic field generated by the coil of channel 5-4. It becomes a typical area. The four coils of the channels 5-1, 5-2, 5-3, and 5-4 are covered with an RF shield 5-s that prevents magnetic field interference between the gradient magnetic field coil 3 and the transmission coil 5. Yes.

  FIG. 2B is another example of the 4-channel transmission coil 5. In the transmission coil 5 of FIG. 2B, the channels 5-1a, 5-1b, 5-1c, and 5-1d are irradiated with the high frequency magnetic field pulse in synchronization with the first channel (5-1), and the channel 5- 2a, 5-2b, 5-2c, and 5-2d are the second channel (5-2), and the channels 5-3a, 5-3b, 5-3c, and 5-3d are the third channel (5-3), the channels A high frequency power pulse is operated from the power supply 7 so that 5-4a, 5-4b, 5-4c, and 5-4d radiate a high frequency magnetic field pulse in synchronization with each other as the fourth channel (5-4). Using the coil of this configuration, for example, the subject 1 was irradiated with a high-frequency magnetic field pulse simultaneously from the coils of the five channels 5-1a, 5-1b, 5-1c, and 5-1d of the first channel (5-1). In this case, high frequency magnetic field pulses generated by the respective channels 5-1a, 5-1b, 5-1c, and 5-1d are evenly irradiated to the entire subject 1.

  In both coils of FIGS. 2A and 2B, the amplitude and phase of the high-frequency magnetic field pulse transmitted from each channel are set independently by the computer 9. The systems 7-1 to 7-4 constituting the power source (transmitter) 7 of the present embodiment are each channel 5-1, 5-2, 5-3 according to control from the computer 9 (via the sequence control device 14). The high frequency power pulse is transmitted independently to 5-4. At this time, the systems 7-1 to 7-4 of the high-frequency magnetic field power supply 7 that feeds the channels 5-1, 5-2, 5-3, and 5-4 have a structure in which at least the maximum output value does not affect each other. Therefore, the duty ratio of each system 7-1 to 7-4 and the maximum output value corresponding to the duty ratio can be set independently for each system, and even when the maximum output value is restricted as shown in Table 1. Each system is restricted independently.

  Next, an example of a CEST pulse sequence executed by the MRI apparatus of this embodiment will be described with reference to FIG. The first half of the pulse sequence of FIG. 5 is a pre-pulse sequence 51 for performing selective saturation. A pre-pulse sequence 51 in FIG. 5 represents an example in which 40 Gaussian pulses (selective saturation pulse) 151 having a pulse length of 35 ms are continuously emitted. The excitation band (half width) of the Gaussian pulse 151 having a pulse length of 35 ms is about 0.5 ppm. Among the protons of the subject 1 irradiated with the pre-pulse sequence 51, protons having a resonance frequency in a bandwidth of 0.5 ppm having a bandwidth of 0.5 ppm with the transmission frequency of the Gaussian pulse 151 as a center frequency are in a resonance state, Absorb energy. This causes a change in the ratio of the upward spin to the downward spin of each proton. When many selective saturation pulses (Gaussian pulses) 151 are irradiated and the amount of energy absorbed by protons exceeds a certain amount, the ratio of the upward spin to the downward spin of each proton is 1, and the magnetic The macroscopic magnetization that generates the resonance signal becomes zero (protons whose resonance frequency is included in the narrow band irradiated with the selective saturation pulse 151 are completely saturated). When the transmission frequency of the Gaussian pulse 151 is set to +3.5 ppm, which is the resonance frequency (average value) of the Amide group (—NH) proton, a large number of “Amid group protons” are irradiated with the selective saturation pulse 151. Saturates. Since proton exchange occurs at a constant rate (rate) at all times between “Amid group protons” and “protons contained in water (free water) that generates most ordinary MRI signals”, the saturation state The protons in the water are replaced with protons contained in the water, and the MRI signal generated from the water is reduced. The reduced MRI signal is measured by the measurement pulse sequence 52 in the latter half of the pulse sequence of FIG. The pulse sequence shown in FIG. 5 is repeated while changing the transmission frequency of the selective saturation pulse 151 from, for example, +6.0 ppm to −6.0 ppm at 0.5 ppm intervals.

  The measurement pulse sequence 52 may be any pulse sequence that can measure a normal MR image, and a basic pulse sequence such as a spin echo method or a gradient echo method may be used. However, in CEST imaging, it is necessary to change the transmission frequency of the selective saturation pulse 151 little by little as described above and to perform the entire pulse sequence of FIG. 5 each time. Therefore, the measurement pulse sequence 52 is as short as possible. A pulse sequence that can obtain an image in time is desirable. For example, a pulse sequence capable of measuring an image with one shot, such as a one-shot two-dimensional fast spin echo method or a one-shot two-dimensional echo planar method, can be used as the measurement pulse sequence 52. Also, a pulse sequence that requires multiple measurements, a multi-slice pulse sequence, or a three-dimensional spatial measurement pulse sequence may be used.

  In addition, as the pre-pulse sequence 51, a specific numerical value and a name such as “the pulse length is 35 ms, the waveform type is a Gauss type selective saturation pulse 151, and the number of irradiations is 40” are given. If the pre-pulse sequence can be approached, the pulse length, waveform type, and number of irradiations can be freely selected and combined.

  Next, in the pre-pulse sequence 51 of the CEST pulse sequence, the output for each channel when a large number of selective saturation pulses 151 are irradiated will be further described with reference to FIGS. 3A to 3C show pulse sequences when the pre-pulse sequence 51 is executed using the coils of the four channels 5-1 to 5-4 as shown in FIGS. 2A and 2B. Show. FIG. 3A is a comparative example, and shows a pulse sequence in the case where a single power source is connected to four channels, high-frequency power pulses are simultaneously supplied to all channels, and simultaneous irradiation of selective saturation pulses is performed. Yes. FIG. 3B shows a pre-pulse sequence in which the duty ratio of this embodiment is reduced to about 20%. FIG. 3C shows a pre-pulse sequence in which the duty ratio is reduced to 50% or less and the uniformity of the magnetic field distribution of the selective saturation pulse 151 is improved. 3A to 3C show an example in which a rectangular wave is used instead of a Gaussian wave as the waveform of the selective saturation pulse 151 for easy understanding. In addition, since the maximum output of each system 7-1 to 7-4 of the power supply 7 is limited as shown in Table 1 above depending on the duty ratio, the output of the selective saturation pulse 151 is actually also with time. Although it is gradually reduced to the limit value (maximum output value in Table 1), in FIGS. 3A to 3C, it is represented as a pulse having a constant height.

  The pulse sequence shown in FIG. 3A of the comparative example has a configuration in which continuous irradiation is always performed simultaneously from all four channels 5-1 to 5-4. For this reason, the duty ratio of the high-frequency magnetic field power supply 7 is a high ratio such as 70%, and the maximum output value is limited according to the duty ratio when continuously irradiated. For example, when the performance of the power supply 7 is as shown in Table 1 above, the duty ratio is 70%, so that only 7% of the maximum output when the duty ratio is 10% can be obtained.

  On the other hand, in the pulse sequence of FIG. 3B of the present embodiment, the systems 7-1 to 7-4 are connected to the coils of the four channels 5-1 to 5-4, respectively, and the channels are selected sequentially. Saturation pulse 151 is irradiated. Thereby, since the duty ratio of each system 7-1 to 7-4 of the power supply 7 can be a relatively low ratio of about 20%, the restriction due to the maximum output value is relaxed. For example, when the performance of each of the systems 7-1 to 7-4 of the power supply 7 is as shown in Table 1 above, the duty ratio is 20%, so 50% of the maximum output when the duty ratio is 10%. Output can be obtained.

  In the case of the pulse sequence of FIG. 3B, since the selective saturation pulse 151 is sequentially irradiated from the four channels 5-1 to 5-4, the selective saturation pulse 151 is irradiated to the entire subject 1 by one irradiation. However, the selective saturation pulse 151 is irradiated to the entire subject 1 only after four irradiations, but there is no problem as the CEST pre-pulse sequence 51. The protons in the region irradiated with the selective saturation pulse 151 are saturated enough to saturate all the protons before the saturated protons return to their original state because the protons in an amount corresponding to the irradiated energy amount are saturated. It is only necessary to irradiate an appropriate number (energy) of selective saturation pulses 151. Therefore, even if the selective saturation pulse 151 is sequentially irradiated from each channel, all protons can be saturated during the pre-pulse sequence 51.

  In normal MRI measurement, in order to improve the spatial uniformity of the proton spin excitation high frequency magnetic field and inversion high frequency magnetic field, the standard deviation (or maximum value) of the result of adding the high frequency magnetic field distribution with positive and negative signed values is used. Using the minimum value as a uniformity index, the amplitude and phase of the high-frequency magnetic field pulse are set so that a uniformity greater than a predetermined value can be obtained. In the case of the CEST selective saturation pulse 151 of this embodiment, the absolute value (square root of the sum of squares) of the high-frequency magnetic field distribution of each selective saturation pulse 151 is added, and the standard deviation (or the maximum value) of the obtained absolute value sum is obtained. And the minimum value) are used as the uniformity index, and the irradiation condition (at least one of amplitude and phase) of the selective saturation pulse 151 may be set so that a uniformity equal to or higher than a predetermined value can be obtained.

  In this embodiment, since the selective saturation pulse 151 is sequentially transmitted from the four channels as shown in FIG. 3B, the four times of irradiation is considered as one set, and the absolute value of the high-frequency magnetic field distribution at each irradiation is determined. By adding the standard deviation of the sum of absolute values (or the maximum and minimum values) as a uniformity index, and optimizing the irradiation condition (at least one of amplitude and phase) of the selective saturation pulse 151 It becomes possible to further improve the uniformity.

  In the example of FIG. 3B, when performing sequential transmission from the four channels, the channel is “channel 1 transmission → channel 2 transmission → channel 3 transmission → channel 4 transmission → channel 1 transmission →... Although the case where transmission is repeated in the order of numbers has been described, transmission may be performed sequentially in an optimal order according to coil characteristics regardless of the channel number.

  On the other hand, FIG. 3 (c) uses the coils of four channels 5-1 to 5-4 to alternate between set 1 of “channel 1 + channel 3” and set 2 of “channel 2 + channel 4” with a duty ratio of 50% or less. This is a pulse sequence to be transmitted. Thus, RF shimming can be performed by adjusting the amplitude and phase between the two channels in each set while setting the systems 7-1 to 7-4 to a duty ratio of 50% or less. For example, the absolute value of the high-frequency magnetic field distribution irradiated from two channels in each set is added, and the standard deviation (or maximum and minimum values, etc.) of the obtained absolute value is used as the uniformity index. By performing the optimization (adjustment of at least one of the amplitude and phase of the selective saturation pulse), the uniformity can be further improved. This makes it possible to improve the spatial uniformity of the high-frequency magnetic field while preventing the reduction of the maximum output, so that many high static magnetic field MRI apparatuses (for example, a static magnetic field intensity of 3.0 T) on which CEST imaging is performed. It is possible to suppress the CEST signal distribution resulting from the spatial distribution of the magnetic field of the high-frequency magnetic field pulse that is likely to occur in

  Next, in the MRI apparatus of this embodiment, the operation of each unit when performing CEST imaging will be described with reference to the flowchart of FIG. The computer 9 controls each unit so as to realize the operation of the flow of FIG. 6 by reading and executing a program stored in a built-in memory in advance. First, the computer 9 prompts an operator who performs MRI measurement to set the subject 1 in the MRI apparatus (step 601). For example, a display that prompts the user to set the subject 1 is displayed on the display device 10. Thereby, when the operator sets the subject 1, the computer 9 receives the setting of measurement conditions from the operator via the input device 15. For example, the computer 9 displays a display screen as shown in FIG. 7 on the display device 10 and receives the type of measurement sequence from the operator. Specifically, acceptable sequences such as SE sequence, FSE sequence, CEST sequence, etc. are displayed, and the computer 9 accepts selection by the operator. When the operator selects CEST as the measurement sequence, an input of general measurement parameters is accepted. Further, the computer 9 receives from the operator the selection of the total number of channels that transmit the CEST selective saturation pulse 151 and the necessity of RF shimming when the selective saturation pulse 151 is transmitted, as necessary (step 602). .

  Next, when the computer 9 accepts that the operator has clicked the measurement button (step 603), the computer 9 causes the sequence control device 14 to perform pre-measurement (step 604). In this pre-measurement, measurement of the high-frequency magnetic field distribution necessary to determine the transmission conditions (amplitude / phase) of each channel at the time of CEST saturation selection pulse transmission is performed. Based on the high-frequency magnetic field distribution data obtained in the previous measurement, the total number of channels that transmit the selective saturation pulse 151, the necessity of RF shimming, and the like, the computer 9 determines the transmission condition (amplitude) of each channel when transmitting the CEST saturation pulse. (Phase) is calculated (step 605). The computer 9 sets the transmission conditions after the adjustment in the sequence control device 14, and the sequence control device executes the CEST imaging main measurement (steps 606 and 607). When performing RF shimming, the computer (calculation unit) 9 calculates a uniformity index from the high-frequency magnetic field distribution of the high-frequency magnetic field pulse, and the high-frequency magnetic field pulse is calculated so that the uniformity is equal to or greater than a predetermined value based on the uniformity index. At least one of the amplitude and the phase is calculated (optimized) by a known method.

  Further, in step 602, the computer 9 may receive a duty ratio value or a maximum output value setting desired by the operator from the operator. In this case, when the computer (selection unit) 9 receives the value of the duty ratio in step 602, in step 605, the duty ratio of the high frequency power pulse supplied to the channel is equal to or less than the received value. Sets (selects) the number of channels for supplying high-frequency power pulses. Further, when the computer 9 accepts the setting of the maximum output value from the operator in step 602, the computer 9 refers to Table 1 stored in advance in the built-in memory and corresponds to the set maximum output value. The number of channels that supply high-frequency power pulses may be set (selected) so that the duty ratio is obtained and the duty ratio of the high-frequency power pulses supplied to the channels is equal to or less than the accepted value. Thereby, CEST imaging can be performed at a duty ratio or maximum output desired by the operator.

  In the above example, an example in which normal MRI measurement is performed as the measurement sequence 52 after the pre-pulse sequence 51 for the selective saturation pulse 151 has been described. However, as the measurement sequence 52, MRS (Magnetic Resonance Spectroscopy) or MRSI (Magnetic Measurement such as Resonance Spectroscopic Imaging) or angiography may be performed.

  In the above example, the configuration in which the selective saturation pulse 151 is irradiated by the transmission coil having a plurality of channels has been described. However, for RF shimming or the like of four-channel alternating transmission, an imaging pulse sequence (in which a pseudo-saturation high-frequency magnetic field pulse is irradiated) For example, the present invention can also be applied to Magnetic Resonance Spectroscopy method, Magnetic Resonance Spectroscopic Imaging method, Magnetization Transfer Magnetic Resonance Angiography method, Magnetic Resonance Decoupling method).

  Further, in the above example, when irradiation is performed with a transmission coil having a plurality of channels, regarding the uniformity index of the magnetic field, during sequential irradiation (irradiating the selective saturation pulse 151 sequentially from the four channels 5-1 to 5-4). Although the case where the same number of high frequency magnetic field distributions as the number of channels is added has been described, the number of high frequency magnetic field distributions different from the number of channels may be added. For example, when performing four-channel sequential irradiation, the sequential irradiation repeated six times is considered as one set, and the absolute value of the high-frequency magnetic field distribution at each irradiation is added, and the standard deviation of the sum of the absolute values obtained (or Uniformity can be further improved by setting at least one of the amplitude and phase of the selective saturation pulse so that a uniformity equal to or greater than a predetermined value can be obtained using the maximum and minimum values as the uniformity index. It becomes.

  As described above, according to the present embodiment, in an MRI apparatus using an independent drive type coil having a plurality of channels as a transmission coil, when a plurality of high-frequency magnetic field pulses are continuously transmitted, the maximum output is a predetermined value or less. It is possible to perform substantially continuous irradiation without lowering to a low level.

  In addition, according to the present embodiment, when transmitting a plurality of high-frequency magnetic field pulses continuously in an MRI apparatus including a coil having four or more channels as a transmission coil, two or more channels are transmitted as shown in FIG. It is possible to improve the spatial uniformity of the high-frequency magnetic field by performing irradiation alternately or sequentially as a set and adjusting the amplitude and phase of the high-frequency magnetic field between multiple channels of each set (RF shimming). .

  Furthermore, according to the present embodiment, when improving the spatial uniformity of the selective saturation pulse, the absolute value (square root of the sum of squares) of the high-frequency magnetic field distribution of each channel is added, and the standard deviation of the obtained absolute value sum ( Alternatively, the uniformity can be further improved by setting at least one of the amplitude and the phase using the maximum value and the minimum value as the uniformity index. When transmitting sequentially from the four channels, considering four irradiations as one set, the absolute values of the high-frequency magnetic field distribution at each irradiation are added together, and the standard deviation of the sum of absolute values obtained (or the maximum) It is possible to further improve the uniformity by performing optimization using the value and the minimum value as the uniformity index.

  CEST pulse sequence shown in FIG. 5 (TR = 5000 ms, echo time = 6 ms, matrix number = 128 × 128, echo number: 128 echoes on the human head on an MRI apparatus having a static magnetic field strength of 3.0 Tesla / Shot), a high-frequency magnetic field distribution map of the selective saturation pulse 151 obtained when executing (/ shot) was obtained by simulation. The result is shown in FIG.

FIG. 8A is a comparative example, and shows a high-frequency magnetic field distribution when a two-channel type coil is used as the transmission coil 5 and a plurality of selective saturation pulses 151 are transmitted continuously and simultaneously irradiated from both channels. It is an image. When the standard deviation of “sum considering phase” is used as the uniformity index when adjusting the uniformity of the high-frequency magnetic field distribution by adjusting at least one of the amplitude and phase of the selective saturation pulse 151, the measurement result The standard deviation (U _SD ) of the high-frequency magnetic field distribution calculated from the above was 0.1302.

  FIG. 8B is a comparative example. When the 4-channel type coil shown in FIG. 2A is used as the transmission coil 5 and all the selected saturation pulses 151 are transmitted continuously multiple times, all channels are shown in FIG. As in a), it is an image showing a high-frequency magnetic field distribution when simultaneously irradiated. As a uniformity index when adjusting the uniformity of the high-frequency magnetic field distribution by adjusting at least one of the amplitude and the phase of the selective saturation pulse 151, a standard deviation of “sum considering phase” was used. The standard deviation of the high-frequency magnetic field distribution calculated from the measurement result is 0.1162, which indicates that the uniformity is improved as compared with the case where a two-channel coil is used.

  8 (c) uses the 4-channel type coil of FIG. 2 (a) as the transmission coil 5, and when each of the two channels shown in FIG. It is an image showing the high frequency magnetic field distribution when performing alternate irradiation (referred to as four-channel alternate irradiation). As a uniformity index when adjusting the uniformity of the high-frequency magnetic field distribution by adjusting at least one of the amplitude and the phase of the selective saturation pulse 151, a standard deviation of “sum considering phase” was used. The standard deviation of the high-frequency magnetic field distribution calculated from the measurement result is 0.1031, indicating that the uniformity is improved as compared to the case where four channels are simultaneously irradiated.

  8 (d) uses the 4-channel type coil of FIG. 2 (a) as the transmission coil 5, and performs the sequential irradiation shown in FIG. 3 (b) when a plurality of selective saturation pulses 151 are transmitted continuously. It is the image showing the high frequency magnetic field distribution at the time. As a uniformity index when adjusting the uniformity of the high-frequency magnetic field distribution by adjusting at least one of the amplitude and the phase of the selective saturation pulse 151, a standard deviation of “sum considering phase” was used. The standard deviation of the high-frequency magnetic field distribution calculated from the measurement result is 0.0774, which indicates that the uniformity is improved as compared with the case where the four-channel alternate irradiation in FIG. 8C is performed.

  FIG. 8 (e) uses the 4-channel type coil of FIG. 2 (a) as the transmission coil 5, and each of the two channels shown in FIG. 3 (c) when transmitting a plurality of selective saturation pulses 151 continuously. It is an image showing the high frequency magnetic field distribution when performing alternate irradiation (4-channel alternate irradiation). Unlike the case of FIG. 8C, “sum in absolute value” is used as a uniformity index when the uniformity of the high-frequency magnetic field distribution is optimized by adjusting at least one of the amplitude and phase of the selective saturation pulse 151. Standard deviations were used. The standard deviation of the high-frequency magnetic field distribution calculated from the measurement result is 0.1012, and the uniformity is slightly smaller than in the case of FIG. 8C using the standard deviation of “sum considering phase” as the uniformity index. It has been shown to improve.

  FIG. 8 (f) uses the 4-channel type coil of FIG. 2 (a) as the transmission coil, and the sequential irradiation shown in FIG. 3 (b) was performed when a plurality of selective saturation pulses 151 were transmitted continuously. It is an image showing the high frequency magnetic field distribution at the time. Unlike the case of FIG. 8D, as a uniformity index when the uniformity of the high-frequency magnetic field distribution is optimized by adjusting at least one of the amplitude and phase of the selective saturation pulse 151, “sum in absolute value” is used. The standard deviation is used. When the standard deviation of the high-frequency magnetic field distribution calculated from the measurement result is 0.0629 and 4-channel sequential irradiation is performed using the standard deviation of “sum considering phase” as the uniformity index (FIG. 8D) It is shown that the degree of uniformity is somewhat improved compared to.

  As described above, a CEST pulse sequence (TR = 5000 ms, echo time = 6 ms, number of matrices = 128 × 128, number of echoes: 128 echoes / shot) is performed using sequential irradiation with a 4-channel type coil. The actual output of the high-frequency magnetic field in the cases of 8 (d) and (f) (duty ratio 25%) is as shown in FIG. 8 (duty ratio) in FIG. 50%) can be increased about three times (that is, about 10% to about 32%). For this reason, it is possible to increase the CEST signal (APT signal) proportional to the irradiation energy (time integration value) of the selective saturation pulse.

1: subject, 2: static coil, 3: gradient coil, 4: shim coil, 5: transmission coil, 5-1 to 5-4: channel, 5-s: shield, 6: reception coil, 7: power supply 7-1 to 7-4: System, 8: Receiver, 9: Computer, 10: Display device, 11: Storage device, 12: Power supply for gradient magnetic field, 13: Power supply for shim, 14: Sequence control device, 15 : Input device, 100: MRI device, 151: selective saturation pulse, 51: pre-pulse sequence, 52: pulse sequence for measurement, 200: MRI device, 300: MRI device, RF: high-frequency magnetic field, Gz: gradient magnetic field in Z-axis direction , Gy: gradient magnetic field in the Y-axis direction, Gx: gradient magnetic field in the X-axis direction, TR: repetition time, A / D: time for receiving a magnetic resonance signal and performing A / D conversion.

Claims (10)

A transmission coil that irradiates a subject with a high-frequency magnetic field pulse, and a power source that is connected to the transmission coil and supplies a high-frequency power pulse;
The transmission coil includes a plurality of channels,
The power supply has a plurality of systems, the systems are respectively connected to one or more of the channels, and the maximum output is limited according to the duty ratio,
The high-frequency power pulse is sequentially supplied to the plurality of channels from the system, and the high-frequency magnetic field pulse is sequentially irradiated from the plurality of channels, thereby continuously generating a plurality of shots with a non-irradiation time equal to or less than a predetermined value as the entire transmission coil. The number of the systems is set so that the duty ratio of each system becomes a predetermined value or less when the high frequency magnetic field pulse is irradiated to the magnetic resonance imaging apparatus. A transmission coil that irradiates a subject with a high-frequency magnetic field pulse, and a power source that is connected to the transmission coil and supplies a high-frequency power pulse;
The transmission coil includes a plurality of channels,
The power source has a plurality of systems, the systems are respectively connected to the channels, and the maximum output is limited according to the duty ratio,
The high-frequency power pulse is sequentially supplied to the plurality of channels from the system, and the high-frequency magnetic field pulse is sequentially irradiated from the plurality of channels, thereby continuously generating a plurality of shots with a non-irradiation time equal to or less than a predetermined value as the entire transmission coil. The number of channels is set so that the duty ratio of the high-frequency power pulse supplied to the channel is less than or equal to a predetermined value when the high-frequency magnetic field pulse is irradiated to the magnetic resonance imaging device. apparatus.   3. The magnetic resonance imaging apparatus according to claim 1, wherein the duty ratio is 25% or less. 4. A transmission coil that irradiates a subject with a high-frequency magnetic field pulse, and a power source that is connected to the transmission coil and supplies a high-frequency power pulse;
The power source has four or more systems,
The transmission coil includes four or more channels, and is connected to the system for each channel,
A plurality of the high-frequency magnetic field pulses are continuously irradiated as the entire transmission coil by irradiating the high-frequency magnetic field pulses sequentially from each of the four or more channels with a non-irradiation time of a predetermined value or less. Magnetic resonance imaging apparatus. A transmission coil that irradiates a subject with a high-frequency magnetic field pulse, a power source that is connected to the transmission coil and supplies a high-frequency power pulse, and a selection unit,
The transmission coil includes a plurality of channels,
The power source has a plurality of systems, the systems are respectively connected to the channels, and the maximum output is limited according to the duty ratio,
The high-frequency power pulse is sequentially supplied to the plurality of channels from the system, and the high-frequency magnetic field pulse is sequentially irradiated from the plurality of channels, thereby continuously generating a plurality of shots with a non-irradiation time equal to or less than a predetermined value as the entire transmission coil. When the high-frequency magnetic field pulse is irradiated to the channel, the selection unit sets the number of channels for supplying the high-frequency power pulse so that a duty ratio of the high-frequency power pulse supplied to the channel is equal to or less than a predetermined value. A magnetic resonance imaging apparatus.   The magnetic resonance imaging apparatus according to claim 1, wherein the high-frequency magnetic field pulse selectively saturates protons having a predetermined resonance frequency contained in the subject. A magnetic resonance imaging apparatus characterized by being a selective saturation pulse irradiated on the magnetic field. The magnetic resonance imaging apparatus according to any one of claims 1, 2, 4, and 5,
In order to increase the homogeneity of the high-frequency magnetic field distribution of the high-frequency magnetic field pulse, it further has an arithmetic unit for adjusting at least one of the amplitude and phase of the high-frequency magnetic field pulse,
The arithmetic unit, when transmitting one channel at a time from the plurality of channels, takes four or more predetermined numbers of the high-frequency magnetic field pulses that are continuously irradiated as one set, and the one set of high-frequency magnetic field pulses A magnetic resonance imaging apparatus, wherein a uniformity index is calculated from a high-frequency magnetic field distribution, and at least one of an amplitude and a phase of the high-frequency magnetic field pulse is adjusted based on the uniformity index.   The magnetic resonance imaging apparatus according to claim 7, wherein the calculation unit adds the absolute values of the high-frequency magnetic field distributions of the respective high-frequency magnetic field pulses as the uniformity index, and a standard deviation of the obtained absolute value sum. Alternatively, a magnetic resonance imaging apparatus characterized by obtaining a maximum value and a minimum value. A transmission coil that irradiates a subject with a high-frequency magnetic field pulse, and a power source that is connected to the transmission coil and supplies a high-frequency power pulse;
The transmission coil includes a plurality of sets including four or more channels and one set of two or more channels.
The power supply has a plurality of systems, the systems are respectively connected to one or more of the channels, and the maximum output is limited according to the duty ratio,
The transmission is performed by sequentially performing the operation of supplying the high-frequency power pulse from the system to the plurality of channels of the one set and irradiating the high-frequency magnetic field pulse from the plurality of channels of the one set at the same time for each set. When the simultaneous irradiation of the high-frequency magnetic field pulse is continuously performed a plurality of times with a non-irradiation time equal to or less than a predetermined value as a whole coil, the number of the systems is set so that the duty ratio of each system is equal to or less than a predetermined value. A magnetic resonance imaging apparatus characterized by being set. The magnetic resonance imaging apparatus according to claim 9,
In order to increase the homogeneity of the high-frequency magnetic field distribution of the high-frequency magnetic field pulse, it further has an arithmetic unit for adjusting at least one of the amplitude and phase of the high-frequency magnetic field pulse,
The calculation unit calculates a uniformity index from the high-frequency magnetic field distribution of the high-frequency magnetic field pulses irradiated from the plurality of channels in the set, and uses the high-frequency magnetic field pulse based on the uniformity index and A magnetic resonance imaging apparatus characterized by adjusting at least one of phases.

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