JP2008017925A - Nuclear magnetic resonance imaging apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an MRI apparatus capable of suppressing a degradation in image quality due to table movement, in an imaging pulse sequence. <P>SOLUTION: A nuclear magnetic resonance imaging apparatus is further provided with a B0 coil for applying magnetic field pulses (B0 pulses) of uniform space distribution to an imaging space. In the pulse sequence of continuously moving a table, the BO pulses are applied to the imaging space in synchronism with the applying of gradient magnetic field pulses to the imaging space for the moving direction of the table. Thus, A change in an amount of a gradient magnetic field applied to an excitation area attending the table movement is corrected by the BO pulses. Thus, acquired nuclear magnetic resonance (NMR) signals are not affected by the change in the amount of the gradient magnetic field due to the table movement and the degradation in the image quality is suppressed. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention measures nuclear magnetic resonance (hereinafter referred to as “NMR”) signals from hydrogen, phosphorus, etc. in a subject, and visualizes nuclear density distribution, relaxation time distribution, etc., and nuclear magnetic resonance imaging (MRI). More particularly, the present invention relates to an MRI apparatus that performs a continuous moving bed imaging method for imaging a wide region by exciting a subject and acquiring NMR data while continuously moving a table on which the subject is mounted.

  In the MRI apparatus, the sequence is repeatedly executed while changing the phase encoding amount, and an echo signal necessary for reconstruction of one image is acquired. Therefore, the number of repetitions greatly affects the image acquisition time. When performing high-speed imaging, a multi-echo type sequence that generates a plurality of echo signals within one iteration is generally used. For example, an FSE sequence or an EPI sequence corresponds to this. A diffusion weighted imaging (DWI) method is used for cerebral infarction and whole body tumor diagnosis.

  An imaging method called whole body MRI is being developed to diagnose whole body tumors. There are various known techniques for whole body MRI. For example, Non-Patent Document 1 is known. In the technique described in Non-Patent Document 1, 2D or 3D imaging is performed by aligning the bed moving direction and the readout gradient magnetic field direction. The imaging region in the apparatus coordinate system is moved along with the bed movement by changing the excitation frequency as the subject moves along with the bed movement. Thereby, imaging is performed while the subject position in the subject coordinate system is fixed.

  Compared with multi-station imaging, the moving bed imaging method does not require time (3-5 s) for moving the bed between stations, and thus has an advantage that imaging can be performed at high speed. It has been reported that image quality degradation accompanying imaging while moving the bed can be suppressed by a method of recovering by signal processing and by lowering the bed moving speed within a range where image quality degradation does not occur. In these imaging methods, in order to avoid folding artifacts in the body axis direction, in addition to the method in which the body axis direction is read and the gradient magnetic field direction is used as in Non-Patent Document 1, the body axis direction is set as the slice direction for two-dimensional imaging. There is a way to do it.

  An outline of an example of conventional continuous table moving imaging will be briefly described. FIG. 9 shows the relationship between the subject and the imaging space. An area 1001 in FIG. 9 represents an imaging space, and the subject 401 is placed on the table 1002. The table 1002 moves freely in the X direction. Since the positional relationship between the imaging space 1001 and the subject 401 is changed by moving the table 1002, images of different parts can be acquired.

  For example, when an imaging section is set horizontally with respect to the table 1002 (that is, when the Z direction is the slice direction), the results are as shown in FIGS. FIG. 10A shows a case where the table 1002 is viewed from directly above, and FIG. In this case, as the table 1002 being imaged moves in the + (−) X direction, the imaging region 501 also moves in the − (+) X direction relative to the subject.

  At this time, the table movement direction and the read gradient magnetic field direction are set to the same direction, and the echo signal is measured while recursively changing the phase encoding amount of the pulse sequence in conjunction with the movement of the table 1002. Then, when the echo signal measured at one repetition time of the pulse sequence is subjected to one-dimensional Fourier transform in the readout direction (Kx), hybrid data as shown in FIG. 11A can be acquired. In FIG. 10A, the horizontal axis represents the position (X), and the vertical axis represents the phase encoding amount (Ky) applied. As the position of the table 1002 moves as the shooting progresses, the position (X) of the acquired data moves. However, since the phase encoding amount is recursively repeated, it is acquired with the same phase encoding amount. If the hybrid data is arranged, the positions in the X direction are continuously connected (that is, 502-1 and 502-9, 502-2 and 502-10, etc.). FIG. 10B shows the result of combining the X-direction positions in this way. In each of the lines 503-1 to 503-8, the phase encoding amount applied at the time of echo signal measurement is different, so that a final image is obtained by performing Fourier transform in the Ky direction.

  In imaging performed while continuously moving the table as described above, a wide range of information on the table moving direction of the subject can be obtained within one imaging slice. On the other hand, the table movement speed is controlled corresponding to the execution time (for example, repetition time) of the pulse sequence, or conversely, the execution time of the pulse sequence (for example, repetition time TR) is controlled corresponding to the table movement speed. Therefore, it is necessary to be able to combine the hybrid data without disconnecting the shooting area.

In table continuous movement imaging, if the table moving speed is slow, the change in the subject position has little effect on the image, but as the table moving speed increases, the change in the subject position between echo signals increases, and There is a problem that the image cannot be reconstructed. When the imaging section is set as shown in FIG. 10, the moving speed V of the table 1002 uses the sequence repetition time TR, the field of view FOVx in the readout gradient magnetic field direction, the number N of data acquisition points in the phase encoding direction, and the repetition time TR. And
V = FOVx / (N × TR) (Formula 1)
Is calculated. A typical table speed is 1 cm / s to 3 cm / s, and it has been reported that the image quality deteriorates at a speed higher than this.

  For this reason, in the conventional table continuous moving imaging, the moving speed of the table is set to a low speed at which the movement can be substantially ignored in the time interval between the excitation of the pulse sequence and the signal detection. For example, assuming that the read gradient magnetic field application time of a typical pulse sequence (SE, GrE) is 10 ms at a bed moving speed of 2 cm / s, the distance moved to the read gradient magnetic field application time is 0.2 mm. Since the MRI pixel size is typically 1 to 2 mm and the slice thickness is 2 to 5 mm, the movement distance during execution of the pulse sequence of 0.2 mm is about 10 to 20% of the pixel size and ignored. This is a possible variation. Therefore, image quality deterioration accompanying bed movement can be ignored. Therefore, in the conventional example, as in Non-Patent Document 1, image quality degradation is suppressed by changing the excitation frequency only between RF excitation pulses, and table movement at the time interval between excitation and signal detection is considered. Absent.

On the other hand, in Patent Document 1, the phase of k-space data acquired while moving the table in the body axis direction is corrected in the body axis direction, so that it is included in the acquired data due to the change in the table position during k-space data acquisition. A method for removing the phase amount is disclosed.
Yodong Zhu et al. Extended field-of-view imaging with table translation and frequency sweeping, Magnetic Resonance in Medicine 49: 1106-1112,2003. JP 2005-503874 A (WO2003 / 027701)

  When an entire body is imaged with an MRI apparatus, an imaging time of about 250 to 500 s is conventionally required, but high-speed imaging of about 15 to 30 s is potentially desired like X-ray CT. For this purpose, it is necessary to apply a multi-echo sequence and increase the bed moving speed 5-10 times. However, for example, when the bed moving speed is increased to 20 cm / s, the reading gradient magnetic field application time of a typical pulse sequence (SE, GrE) is about 10 ms, and the distance that the bed moves during that time is 2 mm. Since the general pixel size of the MRI apparatus is 1-2 mm, this moving distance during sequence execution is 100% in terms of the pixel size and cannot be ignored. For this reason, conventionally, the image quality deterioration accompanying the movement of the bed is large and has not been put to practical use.

  One of the clinical applications of whole body MRI is whole body DWI for tumor detection, but continuous table movement MRI, especially DWI for high speed table movement, has not been proposed.

  Moreover, although the method of removing the influence of table movement by correcting the acquired data is proposed in Patent Document 1, a technique for acquiring data in which the influence of table movement is suppressed is not proposed. .

  An object of the present invention is to provide an MRI apparatus capable of suppressing image quality deterioration due to table movement in an imaging pulse sequence.

  In order to solve the above-described problems, an MRI apparatus of the present invention continuously moves a table, a static magnetic field generator that generates a static magnetic field in an imaging space, a table on which an object is mounted and placed in the imaging space. A drive unit, a gradient magnetic field generation unit that applies a gradient magnetic field pulse to the imaging space, a high-frequency irradiation unit that irradiates the imaging space with a high-frequency magnetic field pulse, a detection unit that acquires a nuclear magnetic resonance signal generated by the subject, and a detection A signal processing unit that performs arithmetic processing on the detection signal of the unit to generate an image, and a control unit that controls the operations of the driving unit, the high-frequency irradiation unit, and the gradient magnetic field generation unit to execute a predetermined pulse sequence. The MRI apparatus further includes a B0 coil for applying a magnetic field pulse (B0 pulse) having a uniform spatial distribution to the imaging space. In the pulse sequence, high-frequency magnetic field pulses are applied and gradient magnetic field pulses are applied while the table is continuously moved. In the pulse sequence, the control unit applies the B0 pulse in synchronization with the application of the gradient magnetic field pulse in the moving direction of the table. As a result, the change in the amount of the gradient magnetic field applied to the excitation region as the table moves can be corrected with the B0 pulse. Therefore, the acquired nuclear magnetic resonance (NMR) signal is not affected by the change in the gradient magnetic field amount due to table movement, and image quality deterioration can be suppressed.

  The control unit may be configured such that the control unit sets the magnetic field intensity of the B0 pulse to an intensity corresponding to the intensity of the synchronized gradient magnetic field pulse. It is also possible to adopt a configuration in which the magnetic field strength of the B0 pulse is changed linearly with time within one pulse. Thereby, it is possible to cope with table movement within one pulse time. The control unit can also be configured to set the intensity of the B0 pulse to the product of the intensity of the synchronized gradient magnetic field pulse and the table moving distance from the irradiation of the high-frequency magnetic field pulse to the application of the gradient magnetic field pulse.

  The pulse sequence includes a diffusion including a 90 ° excitation high-frequency magnetic field pulse, a 180 ° excitation high-frequency magnetic field pulse, and a diffusion-weighted gradient magnetic field pulse applied in a predetermined direction before and after the 180 ° excitation high-frequency magnetic field pulse. An enhanced image capture sequence can be used. In particular, when the diffusion emphasis gradient magnetic field pulse is applied in the moving direction of the table, it is particularly effective to apply the B0 pulse in synchronization with the diffusion emphasis gradient magnetic field pulse. The diffusion-enhanced gradient magnetic field pulse has a high intensity, so that image quality deterioration easily occurs due to table movement. However, application of the B0 pulse can suppress the image quality deterioration.

Hereinafter, the nuclear magnetic resonance imaging (MRI) apparatus of the present invention will be described in detail with reference to the drawings.
(First embodiment)
First, the configuration of the MRI apparatus according to the first embodiment will be described with reference to FIG. The MRI apparatus generates a magnetic field in the imaging space around the subject 401, a B0 coil 415 that generates a magnetic field with a flat spatial distribution (zeroth order component) in the imaging space, and a gradient magnetic field in the imaging space. A gradient magnetic field coil 403, an RF (high frequency) coil 404 that generates a high frequency magnetic field in the imaging space, and an RF probe 405 that receives an NMR signal generated by the subject 401. In addition, a bed (table) 412 on which the subject 401 is mounted, an RF transmission unit 410, a signal detection unit 406, a gradient magnetic field power source 409, a B0 coil power source 416, a monitor device 415, a bed control unit 414, a signal processing unit 407, A control unit 411 and an image display unit 408 are provided.

  The B0 coil 515 includes a coil configured to generate a magnetic field having a flat spatial distribution (zero-order component), and generates a magnetic field having a uniform intensity distribution in accordance with a signal from the B0 coil power supply 416. The gradient magnetic field coil 403 includes gradient magnetic field coils in three directions of x, y, and z, and each generates a gradient magnetic field in response to a signal from the gradient magnetic field power supply 409. The gradient magnetic field coil 403 is designed so that the magnetic field strength of the gradient magnetic field becomes 0 at the center of the imaging space (x = 0, y = 0, z = 0).

  Since the B0 coil 415 and the gradient magnetic field coil 403 need to suppress eddy currents in order to enable high-speed control (typically pulse control of 1 ms or less), it is desirable that they are active shield types. Specifically, each of the B0 coil 415 and the gradient magnetic field coil 403 is configured to include a main coil that generates a predetermined magnetic field and a shield coil that actively generates a magnetic field that cancels the leakage magnetic field.

  The RF coil 404 generates a high-frequency magnetic field according to the signal from the RF transmitter 410. On the other hand, the NMR signal received by the RF probe 405 is detected by the signal detection unit 406, then signal processed by the signal processing unit 407, and further converted into an image signal by image reconstruction processing. The reconstructed image is displayed on the image display unit 408.

  The B0 coil power supply 416, the gradient magnetic field power supply 409, the RF transmission unit 410, and the signal detection unit 406 are controlled by the control unit 411. In general, the control of the gradient magnetic field power source 409, the RF transmission unit 410, and the signal detection unit 406 is called an imaging pulse sequence. In this embodiment, the control including the control of the B0 coil power source 416 is called an imaging pulse sequence. .

  The bed 412 is for mounting the subject 401 and arranging it in the imaging space. The bed 412 incorporates a drive unit that moves the subject 401 in the HF (top-to-foot) direction (arrow 413 in FIG. 1). The bed control unit 414 drives the bed 412 and moves the position of the subject 401 while matching the execution of the imaging pulse sequence. The moving speed of the bed 412 of this embodiment is a moving speed v = 2.0 cm / s-20.0 cm / s from the top of the head toward the foot.

  The living body monitoring device 415 monitors the pulsation or pulse wave, cardiac radio wave, and respiratory movement of the living body, which are converted into an electric signal or an optical signal and sent to the control unit 411 in real time.

  A part of the RF probe 405, the signal detection unit 406, and the signal processing unit 407 is shown in FIG. As shown in FIG. 2, the RF probe 405 includes an RF receiving coil 301 and a preamplifier 302 connected thereto, and a signal received by the RF receiving coil 301 is amplified by the preamplifier 302. The signal detection unit 406 includes an AD conversion / quadrature detection circuit 303, and performs quadrature detection after AD conversion of the output of the preamplifier 302. The quadrature detection performed by the AD conversion / quadrature detection circuit 303 is preferably changed for each channel in accordance with the frequency of the imaged signal. The signal processing unit 407 includes a Fourier transform unit 304 and a calculation unit 305, and Fourier transforms the signal after quadrature detection by the Fourier transform unit 304, thereby reconstructing the MRI image detected by the RF receiving coil 301. If necessary, the image signal is further subjected to calculation processing (for example, image composition calculation when signals are detected in parallel from a plurality of channels). The reconstructed image and the image after the arithmetic processing are displayed on the image display unit 408.

  It should be noted that the morphological imaging target of MRI is the main constituent substance of the subject 401, proton, which is widely used clinically. By imaging the spatial distribution of the proton density and the relaxation phenomenon of the excited state, the shape of the human head, abdomen, limbs, etc., or the function of blood vessels (blood, blood flow), etc., is two-dimensional or three-dimensional. Images.

  Next, the imaging pulse sequence will be described. Various general pulse sequences can be used as the imaging pulse sequence used in the present embodiment. In this embodiment, a case where a 3D gradient echo sequence is used will be described. First, a general 3D gradient echo sequence procedure will be described with reference to FIG. As shown in FIG. 3, simultaneously with the application of the slice selective gradient magnetic field pulse 602, after the irradiation with the RF pulse 601, the slice encode gradient magnetic field pulse 603 and the phase encode gradient magnetic field pulse 604 are applied. Thereafter, a dephase pulse 605 and a readout gradient magnetic field pulse 606 are applied, and thereby an echo signal 607 generated at a predetermined echo time TE608 is detected. These pulse sequences are repeated at a repetition time (TR) (interval of the high frequency pulse 601) 609. Note that the slice encoding and phase encoding gradient magnetic field pulses 603 and 604 are changed for each repetition time TR609 to give different slice / phase encoding, and an echo signal 607 obtained by each encoding is detected.

  The above operation is repeated by the number of encodings, and an echo signal necessary for reconstruction of one 3D image is acquired in a predetermined image acquisition time. As the number of slice encodes and phase encodes, a combination of values such as 32, 64, 128, 256, 512, etc. is usually selected per one 3D image. Each echo signal is usually obtained as a time-series signal composed of 128, 256, 512, and 1024 sampling data. These data are subjected to three-dimensional Fourier transform to create one 3D MR image. When specialized for high-speed imaging, TE can be set to 1 ms, TR is set to 2 ms, the number of frequency encoding of one echo is 128, and the number of phase encoding and slice encoding is 40 (half scan application) × 16. The time required to acquire all encoded data is 1.24 s. When the moving speed of the bed 412 is set to 10 cm / s, the bed 412 moves 12.4 cm while acquiring all encoded data.

  In this embodiment, since the amount of movement of the bed 412 is large, an imaging pulse sequence in which application of a B0 pulse is added at a predetermined timing as shown in FIG. In the following description, (x, y, z) represents the position in each region, and as shown in FIG. ), Z is the HF direction. The moving direction of the bed 412 is the z direction. Here, the bed moving direction (z) is made to coincide with the lead-out axis (Gr). The x and y directions may be set to either the Gs direction or the Gp direction.

  The difference between the imaging pulse sequence of FIG. 4 and the general imaging pulse sequence of FIG. 3 is that a pulse magnetic field is applied from the B0 coil 415 in conjunction with Gr. The application timing and waveform of the B0 pulse magnetic field will be described below.

  Considering the time of irradiation of the RF pulse 601 as the origin of time (t = 0), the bed 412 moves from the irradiation of the RF pulse 601 until the read gradient magnetic field pulse Gr606 is applied. When the moving speed of the bed 412 is v and the time is t, the moving distance A (t) = v · t. Note that the z-direction coil of the gradient magnetic field coil 403 is designed to apply a gradient magnetic field with a magnetic field strength of 0 at the center of the imaging space (z = 0), and therefore, the gradient magnetic field pulse Gr (t) at t = 0. Is applied, a gradient magnetic field having an intensity of Gr (t) × z is applied to each point (x, y, z) in the imaging field (x, y, z is a position in the apparatus coordinate system).

  On the other hand, a specific part (excitation part) in the subject 401 that moves at the speed v moves momentarily by the movement of the bed 412, and the position is described as (x, y, z + vt) in the apparatus coordinate system. The Therefore, a gradient magnetic field of Gr (t) × (z + vt) in the z direction is applied to a specific part (x, y, z + vt) in the subject 401 at time t. When the moving speed v (v = 2.0 cm / s−20.0 cm / s) of the bed 412 is high as in the present embodiment, the gradient magnetic field applied to a specific part by the movement of the head 412. The change in size, that is, the difference (Gr (t) × vt) between Gr (t) × z and (Gr (t) × (z + vt)) cannot be ignored. Therefore, in this embodiment, when a Gr (t) pulse having a linear intensity distribution with respect to z is applied, the effect of high-speed movement is obtained by applying a magnetic field of (−Gr (t) × vt) with the B0 coil. Remove.

  Specific waveforms of the B0 pulses 701 and 702 are shown in FIG. The B0 pulse is a waveform obtained by multiplying the waveform of the gradient magnetic field pulses Gr605 and 606 in the z direction by (-vt). Therefore, the B0 pulses 701 and 702 are applied in synchronization with the gradient magnetic field pulses Gr605 and 606, and the pulse waveform is inclined depending on the change of (−vt). Since the origin of time t is when the RF pulse 601 is irradiated, (−vt) is reset to zero every time the RF pulse 601 is irradiated.

  In this manner, by applying the B0 pulses 701 and 702, the gradient magnetic field Gr (= Gz) = Gr (t) × (z + vt) applied to the specific part is set to B0 = (− Gr (t) × vt). ) Are applied in a superimposed manner, and the term (Gr (t) × (vt)) can be canceled. Therefore, a magnetic field of (Gr (t) × z) that should be originally applied can be applied to the specific part, and the change in the magnitude of the gradient magnetic field due to the movement of the bed 412 (Gr (t) × (vt)) ) Can be corrected. Thereby, it can suppress that the error is contained in the acquired data resulting from the change of the magnitude | size of the applied gradient magnetic field by bed movement.

  As shown in FIG. 4, the pulse waveforms of the B0 pulses 701 and 702 are inclined depending on the change in vt. However, when the width and amplitude of the phase pulse 605 are sufficiently small, the pulse 701 It is also possible to apply a rectangular pulse 701 by approximating that vt is constant.

Next, the relationship between the position of the bed 412 and the gradient magnetic field strength is schematically shown using FIG. In the state of t = 0 immediately after the excitation of the RF pulse 601, the gradient magnetic field Gz (= Gr) is not applied, so that Gz is zero in the entire space as in the graph 901 described in FIG. The time dephase pulse 605 is applied _{t a,} the time readout gradient magnetic field pulse 606 is applied to _t b (= TE), the nuclear magnetization excited by the RF pulse 601, the time t = _{t a} in, since v · _{t a} movement, the nuclear magnetization undergoes gradient _{(Gz '(t a) ·} z) 902 _is expressed as Gz (t a) (z- v · t a). (The gradient Gz (t _a ) is equal to the magnitude Gr (t _a ) of the phase pulse 605.) Therefore, from the B 0 coil 415, B 0 (t _a ) = corresponding to the value of the Gz intercept of the graph 902 gz (t a) _- by applying a (vt _a) of the spatial distribution of sizes is uniform B0 magnetic field pulse 701, not a not move substantially even if moving the nuclear magnetization by the movement of the bed 412 Can receive the same gradient magnetic field. That is,
_{Gz '(t a) · z} -B0 (t a) = Gz (t a) · z
It is.

Next, at time t = t _b , the nuclear magnetization excited by the RF pulse 601 moves by v · t _b, so that the gradient magnetic field Gz ′ (t _b ) 903 received by this nuclear magnetization becomes Gz (t _b ) (z-vt _b ). (The gradient Gz (t _b ) is the magnitude Gr (t _b ) of the readout gradient magnetic field pulse 606.) Therefore, B 0 (t _b ) = Gz (t corresponding to the value of the Gz intercept in the graph 903 _b ) If the B0 magnetic field pulse 702 having a constant spatial distribution of (−vt _b ) is applied, even if the nuclear magnetization is moved by the movement of the bed 412, it is substantially the same as if it has not moved. Can receive a gradient magnetic field. That is,
Gz ′ (t _b ) · z−B 0 (t _b ) = Gz (t _b ) · z
It is.

Note that the magnitude B0 (t) of the magnetic field in the B0 pulse 702 changes linearly with respect to the time t because the bed moves continuously even in the B0 pulse 702 as shown in FIG. This will be specifically described with reference to FIG. 5 (b) is a graph 903 of FIG. 5 (a), before and after the time _{t =} t b -α of the time _{t b} (However, alpha is small time) gradient magnetic field Gz in and _t = t _b + α Graphs 904 and 905 are shown. From the graph 904 of the gradient magnetic field Gz (t _b −α) at t = t _b −α, the magnitude of B 0 (t _b −α) corresponding to the value of the Gz intercept in the graph 904 is larger than B0 (t _b ). I understand that it is small. On the contrary, the magnitude of B0 (t _b + α) corresponding to the value of the Gz intercept in the graph 905 is larger than B0 (t _b ). Thus, the magnetic field magnitude B 0 (t) in the B 0 pulse 702 changes linearly with respect to time even in the B 0 pulse 702.

  When the time width of the B0 pulse 702 is small and the gradient magnetic field strength Gr is not so large, as a simple idea, an average value can be used in the pulse.

  As described above, in the first embodiment, by applying a B0 pulse having a predetermined waveform at a predetermined timing from the B0 coil 415, an error generated in the gradient magnetic field applied to the excitation region as the table moves. And a gradient magnetic field equivalent to the case where there is no table movement can be applied to the excitation region. Therefore, since the acquired NMR signal does not include an error component associated with table movement, image quality deterioration associated with table movement can be suppressed in the imaging pulse sequence.

(Second Embodiment)
Next, as an MRI apparatus according to the second embodiment, an MRI apparatus that executes an imaging pulse sequence in which the same B0 pulse as that in the first embodiment is applied to a typical DWI (diffusion-weighted image) imaging method. explain. A pulse sequence such as DWI that applies a high-intensity gradient magnetic field has a large change in gradient magnetic field application intensity associated with bed movement not only when moving the bed at high speed but also when moving the bed at low speed. Therefore, the effect of suppressing image quality deterioration can be obtained by applying the B0 pulse.

  The imaging pulse sequence executed by the MRI apparatus according to the present embodiment uses a pulse sequence obtained by applying a diffusion gradient magnetic field to general 2D-SE-EPI as shown in FIG. In the pulse sequence of FIG. 6, MPG (motion probing gradient) pulses 105 and 107 are applied in the Gs direction in order to disperse the phase of the diffusing protons so as not to generate a signal. Here, the Gs direction is the bed 412 moving direction (z direction). That is, the imaging section is a direction (TRS) orthogonal to the body axis, and the MPG direction is the body axis (z) direction.

  The pulse sequence will be specifically described below. The 90-degree pulse 101 is applied together with the slice gradient magnetic field pulse 103, and the rephasing pulse 104 is followed by the first half portion 105 of the MPG pulse. Thereafter, the slice gradient magnetic field pulse 106 and the 180-degree pulse 102 are applied simultaneously, and then the latter half 107 of the MPG pulse is applied. When the application of the MPG pulse is finished, the pre-pulse 108 of the phase encode pulse and the phase pulse 111 of the readout gradient magnetic field pulse are applied. Next, the readout gradient magnetic field pulse 112, the blip gradient magnetic field 109 of the phase encode pulse, the inverted readout gradient magnetic field pulse 113, the blip gradient magnetic field 109, the readout gradient magnetic field pulse 114, and so on are repeated. The echo signals 117 to 121 are detected simultaneously with the application of the readout gradient magnetic field pulse.

  The width of typical MPG pulses 105 and 107 is 20 ms, the interval between two MPG pulses is 30 ms, the echo time (TE) 124 is 60 ms, and the repetition time (TR) 125 is 100 to 150 ms. The inter-echo time 110 is 1 ms. The number of blip pulses 109 is determined from the number of phase encodings and is about 40 to 80.

  When the bed speed is 10 cm / s, when TR = 100 ms, the distance that the bed 412 moves between 1TR is 1 cm, so the bed moving distance between 1TR cannot be ignored. Therefore, in this embodiment, the B0 pulses 126 to 130 are applied in correspondence with the gradient magnetic field pulses Gs103 to 107 to correct the gradient magnetic field applied to the excitation region. Specifically, the intensity (waveform) of the B0 pulses 126 to 130 is obtained by multiplying the Gs pulses 103 to 107 waveform by the bed movement amount (−vt) 131 as shown in FIG. 6 (Gs (t) × (−vt )).

  Thus, in the second embodiment, since the moving direction (z direction) of the bed 412 and the direction of the gradient magnetic field Gs coincide with each other, the B0 pulses 126 to 130 correspond to the gradient magnetic field pulses Gs103 to 107, respectively. Is applied, the error (Gs (t) × (vt)) of the gradient magnetic field Gs caused by the movement of the bed 412 is canceled. Thereby, the error which arises in the data acquired by bed movement can be suppressed. Since the DWI pulse sequence applies a high-strength gradient magnetic field, not only when the bed is moved at a high speed, but also when the bed is moved at a low speed, the change in the gradient magnetic field application intensity accompanying the bed movement is large. However, by applying the B0 pulse as in the present embodiment, an effect of suppressing image quality deterioration can be obtained.

(Third embodiment)
Next, an MRI apparatus according to a third embodiment of the present invention will be described with reference to FIG.
Similar to the second embodiment, the MRI apparatus of the third embodiment executes an imaging pulse sequence in which the B0 pulse is applied to the DWI imaging method, but unlike the second embodiment, The bed moving direction (z direction) and the direction of the readout gradient magnetic field Gr are set to coincide. That is, the imaging section is SAG (sagittal section) or COR (crown section), and MPGs 105 and 107 are applied in the RL (left and right) direction (x direction in FIG. 1) or AP (front and back) direction (y direction). To do.

  In this case, as shown in FIG. 7, the B0 pulses 201 to 206 are applied in correspondence with the gradient magnetic field pulses 111 to 116 in the Gr direction which is the bed moving direction. The gradient magnetic field pulses Gr111 to 116 are the phase pulse 111 of the read gradient magnetic field pulse and the read gradient magnetic field pulses 112 to 116, respectively. The intensity (waveform) of each of the B0 pulses 201 to 206 is (Gr (t) × (−vt)) obtained by multiplying the intensity (Gr) of the Gr pulses 111 to 116 by the bed movement amount (−vt) 207.

  As described above, in the third embodiment, the B0 pulses 201 to 206 for correcting the gradient magnetic field pulses Gr111 to 107 are applied even when the moving direction (z direction) of the bed 412 is the readout gradient magnetic field Gr direction. Thus, the error (Gr (t) × (vt)) of the gradient magnetic field Gr applied to the excitation region due to the movement of the bed 412 can be canceled. Thereby, the error which arises in the data acquired by bed movement can be suppressed.

(Fourth embodiment)
Next, an MRI apparatus according to a fourth embodiment of the present invention will be described with reference to FIG.
Similar to the second embodiment, the MRI apparatus of the fourth embodiment executes an imaging pulse sequence in which the B0 pulse is applied to the DWI imaging method, but unlike the second embodiment, The bed moving direction (z direction) and the direction of the phase encoding gradient magnetic field Gp are set to coincide with each other. That is, the imaging section is SAG (sagittal section) or COR (crown section), and MPGs 105 and 107 are applied in the AP (front-rear) direction (y direction) or RL (left-right) direction (x direction in FIG. 1). To do.

  In this case, as shown in FIG. 8, the B0 pulses 1101 to 1102 are applied corresponding to the gradient magnetic field pulses 108 to 109 in the Gp direction which is the bed moving direction. The gradient magnetic field pulses 108 to 109 in the Gp direction are a pre-pulse 108 of a phase encode pulse and a plurality of blip gradient magnetic fields 109. The intensity (waveform) of each of the B0 pulses 1101 to 1102 is (Gp (t) × (−vt)) obtained by multiplying the intensity (Gr) of the Gp pulses 108 to 109 by the bed movement amount (−vt) 1103.

  Thus, in the fourth embodiment, the B0 pulses 1101 to 1102 for correcting the gradient magnetic field pulses Gp108 to 109 are applied even when the moving direction (z direction) of the bed 412 is the phase encode gradient magnetic field Gp direction. Thus, the error (Gp (t) × (vt)) of the gradient magnetic field Gp applied to the excitation region due to the movement of the bed 412 can be canceled out. Thereby, the error which arises in the data acquired by bed movement can be suppressed.

  In the first to fourth embodiments, the application of the B0 pulse is equivalent to a control for advancing or delaying the phase of all the magnetizations in the field of view. Therefore, the operation shown in this embodiment is performed at the RF irradiation frequency. By controlling the RF irradiation phase and the RF reception phase (including partial application), this is equivalent to the obtained effect. However, the present embodiment can be implemented by operating the B0 coil 415 in synchronization with the gradient magnetic field coil 403, and thus has a great feature that the design of the pulse sequence is simple. The present invention can also be applied to oblique imaging, and can also be applied to multi-echo sequences such as EPI and diffusion-weighted imaging using high-intensity MPG pulses.

  As described above, the present invention is an MRI apparatus that moves a table at high speed in a moving bed whole-body imaging method that images a wide region by exciting a subject and acquiring NMR data while moving the table continuously. Alternatively, there is an effect of suppressing image quality deterioration in a pulse sequence that applies a high-intensity gradient magnetic field such as DWI.

1 is a block diagram showing a configuration of an MRI apparatus according to a first embodiment of the present invention. The block diagram which shows the structure of a part of MRI apparatus of FIG. Explanatory drawing which shows the typical example of the 3D gradient echo sequence implemented in 1st Embodiment. Explanatory drawing which shows the pulse sequence of this invention which implemented B0 pulse in the sequence of FIG. 3 implemented in 1st Embodiment. (A) In the first embodiment, a graph showing that the gradient magnetic field applied to the excitation region moves as the bed moves. (B) Before and after the time tb in the B0 pulse 702, the gradient magnetic field is changed. The graph which shows that B0 which should move and should be applied changes. Explanatory drawing which shows the pulse sequence which applied B0 pulse to the DWI imaging sequence whose bed moving direction is Gs direction of 2nd Embodiment. Explanatory drawing which shows the pulse sequence which applied B0 pulse to the DWI imaging sequence whose bed moving direction is Gr direction of 3rd Embodiment The figure which shows the other Example of this invention. Explanatory drawing which shows the pulse sequence which applied B0 pulse to the DWI imaging sequence whose bed moving direction is Gp direction of 4th Embodiment. Explanatory drawing which shows the outline | summary of the conventional moving bed imaging method. In the conventional moving bed imaging method, (a) a top view and (b) a side view for showing movement of an imaging region when an imaging section is set horizontally with respect to a table. (A) In the conventional moving bed imaging method, explanatory drawing which shows the relationship between the position of a bed (imaging area) and acquired data, (b) explanatory drawing which shows the state which connected acquired data continuously.

Explanation of symbols

  126-130 ... B0 pulse, 401 ... subject, 402, magnet, 403 ... gradient coil, 404 ... RF coil, 405 ... RF probe, 406 ... signal detection unit, 407 ... signal processing unit, 408 ... image display unit, 409 ... Gradient magnetic field power supply, 410 ... RF transmission unit, 411 ... control unit, 412 ... bed, 415 ... B0 coil, 416 ... B0 coil power supply, 701, 702 ... B0 pulse.

Claims (6)

A static magnetic field generator for generating a static magnetic field in the imaging space, a table on which an object is mounted and placed in the imaging space, a driving unit for continuously moving the table, and a gradient magnetic field pulse applied to the imaging space A gradient magnetic field generating unit, a high-frequency irradiation unit that irradiates the imaging space with a high-frequency magnetic field pulse, a detection unit that acquires a nuclear magnetic resonance signal generated by the subject, and a detection signal of the detection unit is processed and processed. A signal processing unit that generates a control unit, and a control unit that controls operations of the driving unit, the high-frequency irradiation unit, and the gradient magnetic field generation unit to execute a predetermined pulse sequence,
A B0 coil for applying a magnetic field pulse (B0 pulse) having a uniform spatial distribution to the imaging space;
The pulse sequence is to perform irradiation of the high-frequency magnetic field pulse and application of the gradient magnetic field pulse while continuously moving the table,
The nuclear magnetic resonance imaging apparatus, wherein the controller applies the B0 pulse in synchronization with applying the gradient magnetic field pulse in the moving direction of the table in the pulse sequence.   2. The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the control unit sets the magnetic field intensity of the B0 pulse to an intensity corresponding to the intensity of the gradient magnetic field pulse to be synchronized. apparatus.   3. The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the controller changes the magnetic field intensity of the B 0 pulse linearly with time within one pulse. 4.   4. The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the control unit applies the gradient magnetic field pulse based on the intensity of the gradient magnetic field pulse synchronized with the intensity of the B 0 pulse and the irradiation of the high frequency magnetic field pulse. A nuclear magnetic resonance imaging apparatus, characterized in that the product is set to the product of the table moving distance up to.   5. The nuclear magnetic resonance imaging apparatus according to claim 1, wherein the pulse sequence includes a 90 ° excitation high-frequency magnetic field pulse, a 180 ° excitation high-frequency magnetic field pulse, and before and after the 180 ° excitation high-frequency magnetic field pulse, A nuclear magnetic resonance imaging apparatus, characterized by a diffusion-weighted imaging sequence including a diffusion-weighted gradient magnetic field pulse applied in a predetermined direction. 6. The nuclear magnetic resonance imaging apparatus according to claim 5, wherein the diffusion-weighted gradient magnetic field pulse is applied in the moving direction of the table, and the B0 pulse is applied in synchronization with the diffusion-weighted gradient magnetic field pulse. A nuclear magnetic resonance imaging apparatus.

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