JP2008093418A - Magnetic resonance imaging diagnostic device, and method for controlling the same - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To show effect of flow or flow and a susceptibility with favorable precision. <P>SOLUTION: A gradient magnetic field generating part generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a lead-out gradient magnetic field respectively along a slice axis, a phase encode axis, and a lead-out axis. A host computer 6 sets dephasing quantity for intensifying signal lowering due to flow of the artery and the vein in an interest range of a subject 200 with relation to at least one axis of the three axes. A sequencer 5 and the host computer 6 control the gradient magnetic field generating part to realize a pulse sequence for a gradient echo system including dephasing gradient magnetic field pulses in accordance with the dephasing quantity with relation to the axis to which the dephasing quantity has been set. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention relates to a magnetic resonance diagnostic imaging apparatus suitable for capturing a magnetic susceptibility enhanced image on a head or the like and a control method thereof.

The T ₂ ^* weighted image by the gradient echo method is widely used as an imaging method for imaging a difference in magnetic susceptibility in a region of interest because it sharply reflects inhomogeneity of the local magnetic field. When T ₂ ^* weighted imaging is applied to the head, rephase is generally performed by gradient moment nulling (GMN) to eliminate the effect of blood flow on image quality. Obtain a T ₂ ^* weighted image. Since the difference in magnetic susceptibility is reflected in the contrast as the echo time is longer, the T ₂ ^* weighted image is generally imaged with a relatively long echo time. Further, Non-Patent Document 1 proposes an imaging method that is more sensitive to changes in magnetic susceptibility than T ₂ ^* weighted imaging. In this imaging method, by multiplying the phase shift enhancement mask by the amplitude image, it is possible to enhance both the effects of amplitude attenuation and phase difference induction by the magnetic susceptibility effect.

FIG. 14 is a diagram illustrating a pulse sequence of a three-dimensional (3D) gradient echo method in which first-order GMN is performed on three axes of slice, phase encoding, and readout. If the phase shift is caused only by the difference in magnetic susceptibility at the time when the Echo peak is observed, the shift amount is proportional to the difference in magnetic susceptibility. On the other hand, when the phase change due to the magnetic susceptibility is not considered, the phase shift amount of the spin by applying the gradient magnetic field is expressed by the following equation (1) with a period corresponding to the echo time (TE) as an integration interval.

  Here, γ is a magnetic rotation ratio of approximately 2π × 42.6 MHz / T. G (t) is a gradient magnetic field waveform vector and corresponds to Gss, Gpe, and Gro in each axis of slice, phase encoding, and readout. r0, v0, and a0 represent spin position, velocity, and acceleration vectors at time t = 0, respectively.

  Each term in the equation (1) represents a phase change according to position, velocity, and acceleration, respectively, and corresponds to 0th, 1st, and 2nd gradient moments, respectively. Although the third and higher moments are omitted in the equation (1), the third and higher moments also contribute to the phase change.

GMN is to determine G (t) so that the gradient moment up to a certain order is as small as possible, for example, 0 in TE, and is also called rephase. However, in the phase encoding, the 0th-order moment changes for each encoding step. For this reason, in the case of GMN on the phase encoding axis, the moment at TE is set to a certain value determined for each encoding step in the 0th order, and the first order or higher is set to a small value such as 0, for example. For a blood flow having a spin with motion, rephasing is not performed in the first-order or higher moment term of the equation (1) with only the 0th-order GMN, and a phase change occurs due to motion. For this reason, the vector sum of the spins of the blood flow becomes small due to the phase dispersion, and the blood flow exhibits a low signal without collecting signals. At this time, the phase dispersion varies depending on the flow velocity, so that depending on the blood flow, the phase dispersion is insufficient and the signal is not sufficiently low, and in some cases, an artifact caused by the blood flow may be caused. Therefore, when obtaining an image in which the influence of a flow such as a blood flow is eliminated by T ₂ ^* weighted imaging on the head, it is necessary to perform at least a first-order GMN. The order in which GMN is performed depends on whether or not G (t) can be rephased in that order in TE, but the influence of the flow can be reduced as higher-order rephasing is performed.

Non-Patent Document 1 proposes a method of performing phase enhancement processing to further enhance magnetic susceptibility on the T ₂ ^* weighted image of the head obtained by a gradient echo method that performs first-order GMN on three axes. The contents are as follows.

The original amplitude image before the phase enhancement process is an image reflecting the difference in magnetic susceptibility, which eliminates the influence of blood flow by the first-order rephase. On the other hand, focusing on the phase data, the second-order or higher moment is not rephased in the equation (1). However, by rephasing up to the first order, the influence of the flow such as blood flow can be largely eliminated, and the phase shift of the spin is almost dominated by the phase change due to the magnetic susceptibility. It can be thought of as showing the difference. A phase mask image that lowers the signal value as the phase shift is larger based on the phase data is created, and this is multiplied by the amplitude image a plurality of times, whereby an image in which the phase change due to the magnetic susceptibility is further enhanced can be obtained. In the image obtained by this method, a difference in magnetic susceptibility between tissues is reflected as a difference in contrast. In the head, since there is a large difference in magnetic susceptibility between venous blood having a high deoxyhemoglobin concentration and surrounding tissue having a high oxyhemoglobin concentration, if phase mask processing is performed on the T ₂ ^* image of the head, the vein An image with improved rendering is obtained. The effectiveness of this method for BOLD venography has been reported.
Magn Reson Med 52: 612-618, 2004.

  First problem: A technique disclosed in Non-Patent Document 1 can obtain a magnetic susceptibility-enhanced image with improved venous rendering ability. However, the arterial rendering ability is smaller in susceptibility heterogeneity than veins. There was a problem that the flow effect could not be reflected in the contrast without improving. In the technique disclosed in Non-Patent Document 1, the phase mask process needs to be performed a plurality of times in order to sufficiently improve the vein rendering ability. Therefore, the phase difference between the vein and its vicinity is further emphasized. As a result, there were problems such as overestimation of the venous vascular cavity and enhancement of artifacts associated with the magnetic susceptibility effect.

  Second problem: With the technique disclosed in Non-Patent Document 1, it is possible to generate a phase mask that reflects a phase change that is dominant only in the magnetic susceptibility. Therefore, there is a problem that it is not possible to prevent the occurrence of artifacts due to second-order or higher moments due to pulsation or the presence of blood vessels with complicated running.

  The present invention has been made in consideration of such circumstances, and the object of the present invention is first to accurately depict the effect of flow or flow and magnetic susceptibility.

  A second object of the present invention is to further reduce the influence of the phase shift exerted by the flow and improve the ability to depict the vein.

  According to a first aspect of the present invention, there is provided a magnetic resonance imaging apparatus for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis; A setting means for setting a dephasing amount for emphasizing a signal decrease due to arterial and venous flow in a region of interest of the subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis; Control means for controlling the generating means with a gradient echo pulse sequence including a dephase gradient magnetic field pulse corresponding to the dephase amount with respect to the axis for which the dephase amount is set by the setting means.

  The magnetic resonance imaging apparatus according to the second aspect of the present invention includes a generating unit that generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis; Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject And control means for controlling the generating means so as to realize a gradient echo pulse sequence having a length suitable for canceling the phase shift of the vein.

  According to a third aspect of the present invention, there is provided a magnetic resonance imaging apparatus for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis; A setting means for setting a dephasing amount for emphasizing a signal decrease due to arterial and venous flow in a region of interest of the subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis; Control means for controlling the generating means with a pulse sequence of an asymmetric spin echo system including a gradient magnetic field pulse corresponding to the dephasing amount with respect to the axis for which the dephasing amount is set by the setting means.

  According to a fourth aspect of the present invention, there is provided a magnetic resonance imaging apparatus for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis; Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject Control means for controlling the generating means so as to realize a pulse sequence of an asymmetric spin echo system having a length suitable for canceling the phase shift of the vein.

  A control method according to a fifth aspect of the present invention is a magnetic resonance comprising a generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis. A method for controlling an image diagnostic apparatus, the method for emphasizing signal degradation due to arterial and venous flow of a region of interest of a subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis A dephasing amount is set, and the generating unit is controlled by a gradient echo pulse sequence including a dephasing gradient magnetic field pulse corresponding to the dephasing amount with respect to the axis on which the dephasing amount is set by the setting unit.

  According to a sixth aspect of the present invention, there is provided a control method comprising a generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis. A method for controlling an image diagnostic apparatus, wherein re-phase is performed from first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encode axis, and the readout axis. The generating means is controlled so as to realize a gradient echo pulse sequence in which the echo time is set to a length suitable for canceling the phase shift of the vein of the region of interest of the subject.

  A control method according to a seventh aspect of the present invention is a magnetic resonance comprising a generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis. A method for controlling an image diagnostic apparatus, the method for emphasizing signal degradation due to arterial and venous flow of a region of interest of a subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis A dephasing amount is set, and the generating unit is controlled by an asymmetric spin echo pulse sequence including a gradient magnetic field pulse corresponding to the dephasing amount with respect to an axis on which the dephasing amount is set by the setting unit.

  A control method according to an eighth aspect of the present invention is a magnetic resonance comprising a generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis. A method for controlling an image diagnostic apparatus, wherein re-phase is performed from first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encode axis, and the readout axis. The generating means is controlled so as to realize an asymmetric spin echo pulse sequence in which the echo time is set to a length suitable for canceling the phase shift of the vein of the region of interest of the subject.

  According to the first, third, fifth and seventh aspects, the flow or the effect of flow and magnetic susceptibility can be accurately depicted.

  According to the second, fourth, sixth, and eighth aspects, it is possible to further reduce the influence of the phase shift exerted by the flow and to improve the ability to depict the vein.

  Hereinafter, first and second embodiments of the present invention will be described with reference to the drawings.

  FIG. 1 is a diagram showing a configuration of a magnetic resonance imaging apparatus (hereinafter referred to as an MRI apparatus) 100 according to the first and second embodiments.

  The MRI apparatus 100 includes a bed unit on which a subject 200 is placed, a static magnetic field generation unit that generates a static magnetic field, a gradient magnetic field generation unit for adding position information to the static magnetic field, and a transmission / reception unit that transmits and receives high-frequency signals. And a control / arithmetic unit responsible for overall system control and image reconstruction. The MRI apparatus 100 includes a magnet 1, a static magnetic field power supply (static power supply) 2, a gradient magnetic field coil unit 3, a gradient magnetic field power supply (static power supply) 4, a sequencer (sequence controller) 5, and a host as components of these components. The computer 6, the RF coil unit 7, the transmitter 8T, the receiver 8R, the arithmetic unit 10, the storage unit 11, the display device 12, the input device 13, the shim coil 14, and the shim coil power supply 15 are included. The MRI apparatus 100 also includes an electrocardiogram measurement unit that measures an ECG signal as a signal representing the cardiac time phase of the subject 200 and a breath-hold command unit that commands the subject 200 to hold the breath. . Constituent elements of the electrocardiogram measurement unit and the breath holding command unit include a sound generator 16, an ECG sensor 17, and an ECG unit 18.

The static magnetic field generation unit includes a magnet 1 and a static magnetic field power supply 2. For example, a superconducting magnet or a normal conducting magnet can be used as the magnet 1. The static magnetic field power supply 2 supplies a current to the magnet 1. Thus, the static magnetic field generator generates a static magnetic field B ₀ in a cylindrical space (diagnostic space) into which the subject 200 is sent. The magnetic field direction of the static magnetic field B ₀ substantially coincides with the axial direction (Z-axis direction) of the diagnostic space. Note that a shim coil 14 is further provided in the static magnetic field generation unit. The shim coil 14 generates a correction magnetic field for homogenizing the static magnetic field by supplying current from the shim coil power supply 15 under the control of the host computer 6. The bed unit uses the top plate on which the subject 200 is placed for diagnosis. Send it to the space or pull it out of the diagnostic space.

The gradient magnetic field generation unit includes a gradient magnetic field coil unit 3 and a gradient magnetic field power supply 4. The gradient coil unit 3 is arranged inside the magnet 1. The gradient coil unit 3 includes three sets of coils 3x, 3y, and 3z for generating respective gradient magnetic fields in the X-axis direction, the Y-axis direction, and the Z-axis direction that are orthogonal to each other. The gradient magnetic field power supply 4 supplies a pulse current for generating a gradient magnetic field to the coils 3x, 3y, and 3z under the control of the sequencer 5, respectively. Thus, the gradient magnetic field generation unit controls the pulse current supplied from the gradient magnetic field amplifier 7 to the coils 6x, 6y, 6z, thereby generating gradient magnetic fields in the three axes (X axis, Y axis, Z axis) directions which are physical axes. By combining these, the respective gradient magnetic fields in the respective logical axis directions composed of the slice direction gradient magnetic field Gss, the phase encode direction gradient magnetic field Gpe, and the readout direction (frequency encode direction) gradient magnetic field Gro that are orthogonal to each other are arbitrarily set. Slice direction, phase encoding direction and the readout direction of the gradient magnetic fields Gss, Gpe, Gro is superimposed on the static magnetic field B _0.

  The transmission / reception unit includes an RF coil unit 7, a transmitter 8T, and a receiver 8R. The RF coil unit 7 is disposed in the vicinity of the subject 200 in the diagnostic space. The transmitter 8T and the receiver 8R operate under the control of the sequencer 5. The transmitter 8T supplies an RF current pulse having a Larmor frequency for causing nuclear magnetic resonance (NMR) to the RF coil unit 7. The receiver 8R takes in an MR signal (high frequency signal) such as an echo signal received by the RF coil unit 7 and performs various signal processing such as preamplification, intermediate frequency conversion, phase detection, low frequency amplification, or filtering. After that, A / D conversion is performed to generate digital amount of echo data (original data) corresponding to the echo signal.

  The control / arithmetic unit includes a sequencer 5, a host computer 6, an arithmetic unit 10, a storage unit 11, a display device 12, and an input device 13.

  The sequencer 5 includes a CPU and a memory. The sequencer 5 stores the pulse sequence information sent from the host computer 6 in a memory. The CPU of the sequencer 5 controls the operations of the gradient magnetic field power source 4, the transmitter 8T and the receiver 8R according to the sequence information stored in the memory, and once inputs the echo data output from the receiver 8R, and calculates this Transfer to unit 10. Here, the sequence information is all information necessary for operating the gradient magnetic field power supply 4, the transmitter 8T, and the receiver 8R in accordance with a series of pulse sequences. For example, pulses applied to the coils 6x, 6y, and 6z. Information on current intensity, application time, application timing, and the like is included. The sequencer 5 has a function of controlling the gradient magnetic field generator with a gradient echo pulse sequence including a dephase gradient magnetic field pulse corresponding to a dephase amount set by the host computer 6 under the control of the host computer 6. Have. Furthermore, the sequencer 5 has a function of controlling the gradient magnetic field generation unit so as to generate a plurality of sets of magnetic resonance signals related to the same slice while varying the dephase amount under the control of the host computer 6.

  The host computer 6 has various functions realized by executing a predetermined software procedure. One of these functions is to command the pulse sequence information to the sequencer 5 and control the overall operation of the apparatus. One of the functions described above is to provide a dephasing amount for emphasizing signal degradation due to arterial and venous flow in the region of interest of the subject 200 with respect to at least one of the slice axis, the phase encode axis, and the readout axis. Set. One of the functions described above is to enable the sequencer 5 to collect magnetic resonance signals in a gradient echo system pulse sequence including a dephasing gradient magnetic field pulse corresponding to the dephasing amount with respect to the axis where the dephasing amount is set. Control. One of the functions described above determines a part to be emphasized for signal degradation. One of the above functions controls the sequencer 5 so as to generate a plurality of sets of magnetic resonance signals related to the same slice while varying the amount of dephase. One of the above functions controls the arithmetic unit 10 so as to generate a plurality of prepared images related to the region of interest based on the plurality of magnetic resonance signals, respectively. One of the above functions controls the display 12 to display a plurality of preparation images. One of the functions is to determine one prepared image desired by the operator among the plurality of prepared images. One of the functions is to obtain a phase shift according to the tissue flow by subtracting a phase shift determined from the known magnetic susceptibility of the tissue from the phase shift obtained based on the magnetic resonance signal.

  The host computer 6 performs an imaging scan following a preparatory work such as a positioning scan. The imaging scan is a scan that collects a set of echo data necessary for image reconstruction, and is set to a two-dimensional scan here. The imaging scan can be performed using an ECG gate method based on an ECG signal. This ECG gate method may not be used in some cases.

  The arithmetic unit 10 inputs the echo data output from the receiver 8 </ b> R through the sequencer 5. The arithmetic unit 10 arranges the input echo data in a Fourier space (also referred to as k-space or frequency space) set in the internal memory. The arithmetic unit 10 reconstructs real space image data by subjecting the echo data arranged in the Fourier space to a two-dimensional or three-dimensional Fourier transform. In addition, the arithmetic unit 10 can perform data composition processing, difference arithmetic processing, and the like as necessary.

  In the synthesis process, an addition process that adds image data of two-dimensional plural frames for each corresponding pixel, a maximum value projection (MIP) process that selects a maximum value or a minimum value in the line-of-sight direction for the three-dimensional data, or a minimum Value projection (minIP) processing and the like are included. As another example of the synthesis process, the axes of a plurality of frames may be matched in the Fourier space and synthesized into one frame of echo data as it is. The addition processing includes simple addition processing, addition averaging processing, weighted addition processing, and the like.

  The storage unit 11 stores the reconstructed image data and the image data that has been subjected to the above-described combining process and difference process.

  The display device 12 displays various images to be presented to the user under the control of the host computer 6. A display device such as a liquid crystal display can be used as the display 13.

  The input device 13 inputs various information such as imaging conditions desired by the surgeon, a pulse sequence, information relating to image synthesis and difference calculation, and the like. The input unit 13 sends the input information to the host computer 6. The input device 14 is appropriately provided with a pointing device such as a mouse or a trackball, a selection device such as a mode change switch, or an input device such as a keyboard.

  The breath holding command unit includes a sound generator 16. The voice generator 16 issues a breath holding start message and a breath holding end message as voices in response to a command from the host computer 6.

  The electrocardiograph unit includes an ECG sensor 17 and an ECG unit 18. The ECG sensor 17 is attached to the body surface of the subject 200 and detects the ECG signal of the subject 200 as an electrical signal (hereinafter referred to as a sensor signal). The ECG unit 18 performs various processes including a digitization process on the sensor signal, and then outputs them to the sequencer 5 and the host computer 6. The sensor signal is used by the sequencer 5 when performing an imaging scan. Thereby, the synchronization timing by the ECG gate method (electrocardiogram synchronization method) can be set appropriately, and the ECG gate method imaging scan based on this synchronization timing can be performed to collect data.

  Next, the operation of the MRI apparatus 100 configured as described above will be described in detail.

(First embodiment)
The first embodiment will be described below. This first embodiment corresponds to the first object.

  FIG. 2 is a flowchart of processing of the host computer 6 for imaging in the first embodiment.

  In step Sa1, the host computer 6 instructs the sequencer 5 to collect data. In response to this instruction, the sequencer 5 collects data as described below.

  FIG. 3 is a diagram showing a pulse sequence in the first embodiment. The waveforms shown in FIG. 3 are, in order from the top, a radio frequency pulse (RF) applied to the imaging target, a gradient magnetic field waveform (Gss) in the slice direction, a gradient magnetic field waveform (Gpe) in the phase encoding direction, and a gradient magnetic field in the readout direction. A waveform (Gro) and an echo signal (Echo) are shown.

  As can be seen from FIG. 3, the pulse sequence in the first embodiment conforms to the gradient echo method and is similar to the conventional pulse sequence shown in FIG. However, in the first embodiment, imaging by the gradient echo method is performed after rephasing so that only the magnetic susceptibility is dominant in the phase shift by eliminating the influence of the flow. The difference is that an image with enhanced flow is obtained by performing a phase. That is, each gradient magnetic field is different from that in FIG. 14 so as to cause dephase.

  FIG. 4 and FIG. 5 are diagrams for simply explaining the difference in G (t) between rephase and dephase as representing the difference between rephase and dephase sequences. For simplicity, G (t) is shown as a rectangular wave.

  Assuming that the RF application time is 0 and the echo time (TE) is 3, FIG. 4 shows gradient magnetic field strengths of 1, -2, and 1 at intervals of 1, and FIG. Each of the gradient magnetic field strengths is applied. At this time, the 0th-order moment, that is, ∫G (t) dt is both 0. On the other hand, when the first moment, that is, ∫G (t) tdt is obtained, it becomes 0 in FIG. 4 and G (t) is rephased, whereas it becomes 2.25 in FIG. 5 and G (t) becomes dephased. . Therefore, the difference between rephase and dephase in the diagram as shown in FIG. 3 is that the gradient magnetic field strength in each trapezoidal wave except the first trapezoidal wave of Gss and the final trapezoidal wave of Gro determined by the imaging condition, The gradient magnetic field strength is different in the portion to which the gradient magnetic fields Gss, Gpe, and Gro are applied simultaneously.

  In the rephase, the gradient magnetic field is controlled so as to cancel the signal fluctuation due to the influence of the flow. When rephased, the blood flow phase change is greatly reduced and the blood flow signal is collected as a high signal without being degraded. On the other hand, the dephase actively causes signal fluctuation due to the influence of the flow. In the dephase, the phase dispersion of the spin of the flow such as a blood flow having movement is further advanced by the gradient magnetic field. In the dephase, since the vector sum of the spins of the flow becomes small, that is, the attenuation of the amplitude component becomes larger in the flow signal, the flow signal is suppressed and collected as a low signal.

  In FIG. 3, only a period corresponding to 1TE is shown. Any of the spin warp method, the echo planar imaging method, the echo shift method, and the multi-echo method based on the sequence shown in FIG. 3 can be used. Incidentally, in the spin warp method, the RF excitation and the collection of echo signals are repeated every repetition period TR longer than TE. The echo planar imaging method collects echo signals for a plurality of lines in k-space for one RF excitation for each repetition period TR. In the echo shift method, RF excitation and echo signal collection processing are repeated at a repetition period TR shorter than TE. Echo signals generated during RF excitation performed in a certain repetition period TR are not collected in the same period but are collected in the next period. In the multi-echo method, RF excitation is performed once for each repetition period TR, and echo signals relating to the same line in k space for a plurality of images are collected for the one RF excitation.

  In step Sa2, the host computer 6 instructs the arithmetic unit 10 to perform image reconstruction based on the data collected as described above. In response to this instruction, the arithmetic unit 10 performs image reconstruction by a known method, for example. By this image reconstruction, an amplitude image, a phase image, and 3D (three-dimensional) volume data are obtained.

  In step Sa3, the host computer 6 instructs the arithmetic unit 10 to perform an interpolation process on the reconstructed image. In response to this instruction, the arithmetic unit 10 performs an interpolation process by a known method, for example. This interpolation process can be omitted.

  In step Sa4, the host computer 6 instructs the arithmetic unit 10 to generate a composite image. In response to this instruction, the arithmetic processing unit 10 generates a composite image of the amplitude image and the phase image, for example, by a known method. For example, the arithmetic unit 11 generates a phase mask image representing a portion with a large phase shift and a portion with a small phase shift based on the phase image. This phase mask image is an image in which the pixel of the portion where the phase is advanced (or delayed) on the phase image is zero and the pixel of the portion where there is no phase difference is 1. That is, the phase image is an image representing a portion where the magnetic susceptibility increases (a portion having a different phase) and a portion where the magnetic susceptibility decreases. This phase mask image is a phase image in which the phase disturbance of the low frequency component is removed by filtering the phase image. Filter processing includes processing for passing a phase image through a high-pass filter, processing for differentiating phase images before and after passing through a low-frequency filter (low-pass filter), and the like. The arithmetic unit 11 multiplies the phase mask image generated in this way and the amplitude image a predetermined number of times (one to a plurality of times) to further enhance the phase change due to the magnetic susceptibility (SWI image: Susceptibility- Weighted Imaging image) can be obtained. This process can be omitted.

  In step Sa5, the host computer 6 instructs the arithmetic unit 10 to generate a display image. In response to this instruction, the arithmetic processing unit 10 generates a display image by a known method, for example. This process is performed to represent the blood vessel as a continuous tube. As a method of this processing, for example, projection processing such as maximum value projection processing (MIP), minimum value projection processing (minIP), or addition projection can be applied. In the first embodiment, in the amplitude image, both arteries and veins have low image values with respect to surrounding tissues, and thus minIP is optimal. It is also possible to apply volume rendering or surface rendering after extracting the surface. Alternatively, in order to enable observation of the original image signal, cross-sectional transformation (MPR) can also be applied.

  In step Sa6, the host computer 6 causes the display 12 to display the display image, amplitude image, phase image, or the like generated in step Sa5.

  FIG. 6A is a diagram illustrating an example of a display image generated by applying minIP in step Sa5 without performing phase enhancement processing in step Sa4. FIG. 6B is a diagram illustrating an example of an image obtained by the first conventional technique (Non-Patent Document 1) that performs imaging by the gradient echo method including rephase. FIG. 6C is a diagram illustrating an example of an image obtained by the second conventional technique for performing phase enhancement processing on an image obtained by the first conventional technique. As can be seen from FIGS. 6A to 6C, the first conventional technique, that is, the image of FIG. 6B has a very low blood vessel rendering ability. On the other hand, the second conventional technique, that is, the image of FIG. 6C, has improved vein rendering ability compared to the first conventional technique, but the arterial rendering ability without phase shift is still low. . Furthermore, in the second conventional method, that is, the image of FIG. 6C, the S / N of the brain parenchyma is deteriorated by repeating the phase enhancement process a plurality of times. On the other hand, the image obtained by the first embodiment, that is, the image of FIG. 6A, the veins and the arteries are both clearly and faithfully depicted without performing the phase enhancement process. The drawing ability is improved. Moreover, since post-processing such as mask processing is not added, the brain parenchyma can also be drawn with a high S / N.

  Thus, according to the first embodiment, the flow can be accurately depicted without causing problems such as overestimation of the venous vascular cavity and enhancement of artifacts associated with the magnetic susceptibility effect.

  By the way, in the first embodiment, an artifact may occur depending on the amount of dephase. Therefore, it is desirable to adjust the dephase amount to an appropriate value that does not cause artifacts while sufficiently enhancing the blood vessels.

  The dephase amount can be defined by the moment offset amount used in the description of FIGS. 4 and 5. Alternatively, the dephase amount can be defined by a b value. The concepts of the moment offset amount and the b value are different from each other, and the two are not simply in a unit-convertible relationship. However, in explaining the adjustment of the dephase amount, the moment offset amount and the b value Anyway, it is substantially the same. Therefore, in the following description, the dephasing amount is defined by the b value. The reason why the moment offset amount is used in the description of FIGS. 4 and 5 is that it is suitable for comparing rephase and dephase under the same unit.

The b value is defined as an inner product of integrals of gradient magnetic field waveform vectors as shown in the following equations (2) and (3).

  The b value is related to the amount of dephasing because it represents a signal decrease accompanying dephasing. If the signal without dephasing is S (0) and the signal with dephasing is S (b), there is a coefficient D for each tissue, and S (b) = S (0) · exp (− bD) is defined as a signal drop corresponding to the b value.

FIG. 7 is a diagram showing some examples of minIP images generated as described above by changing the b value. Figure 7 (a) ~ image each b value of imaging of ^{(f), 0.1sec / mm 2} , 1sec / mm 2, 4sec / mm 2, 16sec / mm 2, 32sec / mm 2, 64sec / mm ² .

As can be seen from FIG. 7, veins and arteries are depicted regardless of the b value of 0.1 to 64 sec / mm ² . However, as is apparent from a comparison between the image of FIG. 7E and the image of FIG. 7F, a large artifact occurs when the b value is 64 sec / mm ² . For this reason, the b value is preferably set within a range of about 0.1 to 50 sec / mm ² . Furthermore, as apparent from a comparison between the image of FIG. 7A and the image of FIG. 7B, when the b value is 0.1 sec / mm ² , the blood vessel rendering ability is slightly low. Further, as apparent from comparison between the image of FIG. 7D and the image of FIG. 7E, when the b value is 32 sec / mm ² , the S / N of the brain parenchyma is slightly low. Therefore, a more desirable setting range of the b value is 1 to 20 sec / mm ² .

  However, there are individual differences in the relationship between the amount of dephase and the ability to depict blood vessels. Therefore, the host computer 6 has a function of arbitrarily setting the dephase amount. The host computer 6 may set the dephasing amount based on some numerical value input by the operator through the input device 13, or any one of information (a plurality of candidates) related to the dephasing amount instead of the numerical value. The amount of dephase may be set by selecting this, or the amount of dephase may be automatically set.

  Numerical values that the operator inputs to set the dephase amount are, for example, b value, VENC (velocity encoding), flow velocity, and the like. In order to assist the operator in determining the numerical value to be input, the result of the preparation imaging may be displayed on the display device 12, for example. For example, a plurality of images obtained by performing imaging a plurality of times with different b values are displayed. That is, images as shown in FIGS. 7A to 7F are displayed. Alternatively, for example, MRA (magnetic resonance angiography) images obtained by a plurality of phase contrast methods obtained while varying VENC are displayed. When displaying these images, if the operator can refer to the b value or VENC used to obtain each image, the operator can obtain the b value or VENC can be easily determined as a numerical value to be input. Alternatively, the host computer 6 accepts designation of any one image by the operator. It is also possible that the host computer 6 receives the b value and VENC used to obtain the selected image as the above numerical values. In addition, for each of a plurality of parts for which the signal decrease should be emphasized, information describing a dephase amount suitable for enhancing the signal decrease at the part is stored in, for example, the storage unit 11. The host computer 6 can also set a dephase amount corresponding to a part designated by the operator based on the above information, for example.

  When the dephase amount is automatically set, for example, the CNR (contrast to noise ratio) or SNR (signal to noise ratio) is within a predetermined allowable range among the images obtained by the above-described preparatory imaging. The dephasing amount may be set based on the b value and VENC used to obtain the image. In addition, when there are a plurality of images that meet the above conditions, any image may be selected. However, the larger the dephase amount is, the higher the blood vessel rendering ability is. Therefore, it is preferable to select an image obtained using a larger b value or a smaller VENC.

  On the other hand, in the first embodiment, the flow rendering ability is not changed by TE. For this reason, TE can be changed arbitrarily.

  TE is related to the degree of influence of the local magnetic field non-uniformity on the image contrast. That is, as TE is shortened, the influence of the local magnetic field non-uniformity on the image contrast decreases. Therefore, by setting TE to such a small value that the non-uniformity of the local magnetic field is not reflected in the contrast of the image, an image representing the flow effect more dominantly can be obtained. Conversely, by setting TE to such a large value that the nonuniformity of the local magnetic field is reflected in the contrast of the image, it is possible to obtain an image representing both the effects of flow and magnetic susceptibility.

  TE is the time from when RF excitation is performed until a peak appears in the echo signal. A peak appears in the echo signal when the integral value of the gradient magnetic field Gro becomes zero. Therefore, for example, as shown in FIG. 8, without changing the pattern of the gradient magnetic field Gro, the TE is adjusted by changing the time TA from when RF excitation is performed until the application of the gradient magnetic field Gro is started. Can do. Of course, the TE can be adjusted by changing the pattern of the gradient magnetic field Gro or changing both the time TA and the pattern of the gradient magnetic field Gro.

  An image obtained by a relatively short TE, that is, an image that more dominantly represents the effect of the flow, is suitable for observing the state of the blood vessel because the state of the blood vessel is expressed accurately and easily. In the first embodiment, since the flow is drawn using signal attenuation or phase change due to dephasing, a fine artery, a collateral blood circulation that wraps around from above, and the like can be drawn well. It is possible to obtain an image having clinically useful features such as diagnosis of cerebral infarction. In the case of cerebral infarction, in particular, in the acute phase, a time within several hours from the onset is considered as a time that can be treated with a thrombolytic drug or the like. As time elapses, necrosis of the brain tissue progresses, so it is important to diagnose and connect to treatment while the brain tissue disorder is reversible. The scan time can be shortened by reducing the number of phase encodes, that is, by reducing the resolution. However, thrombus and the like need to be taken under conditions of sufficient S / N and high spatial resolution for diagnosis, and imaging with a scan time as short as possible is desired. Therefore, in the first embodiment, the resolution is lowered by using a parallel imaging (PI) method after using a multi-array coil having a plurality of coil elements as the RF coil unit 7 shown in FIG. Therefore, the imaging time can be shortened, which is extremely effective in the diagnosis of cerebral infarction. In principle, the S / N ratio decreases in inverse proportion to the double speed ratio in the PI method, but the S / N is higher in 3D imaging than in 2D imaging, so the application of the PI method is effective in shortening the time while maintaining spatial resolution. Is. The PI method is a technique in which phase encoding steps are thinned out and filled to generate a folded image from incomplete raw data, and then the image is reconstructed from the raw data. In the PI method, SENSE (sensitivity encoding) and SMASH (simultaneous acquisition of spatial harmonics) are known depending on the reconstruction method, and either may be used. Note that SENSE expands the folds that protrude from the field of view (FOV) using the sensitivity distribution of the plurality of coil elements. SMASH generates complete raw data from the raw data obtained from each of the plurality of coil elements by using the sensitivity distribution of the plurality of coil elements, and reconstructs an image therefrom.

  When the PI method (especially SENSE) is applied, phase data is obtained for each coil element. Therefore, after synthesizing the plurality of phase data by weighted addition, for example, a phase mask image used in the phase shift emphasis process in step Sa4 is created from the synthesized data obtained by this, so that the PI method is more effective. Phase shift emphasis processing can be performed.

Note that the phase change due to the magnetic susceptibility effect is also influenced by the static magnetic field strength in addition to the TE, so that the appropriate TE value in the first embodiment changes according to the static magnetic field strength. For example, if the echo time TE is set to about 40 ms for 1.5T and about 20 ms for 3T, a blood vessel becomes a flow void due to phase dispersion due to dephase, and an image having a contrast sufficiently reflecting the magnetic susceptibility effect is obtained. It is done. Further, in order to cause a positive phase dispersion of the flow due to dephase and reduce the signal, a T ₂ ^* weighted image can be captured with a short TE. This makes it possible to shorten the imaging time and improve the S / N in the T ₂ ^* weighted image capturing.

Note that a T ₂ ^* weighted image in which the spin of blood flow becomes a low signal due to phase dispersion is obtained by performing GMN with only the zeroth moment without performing rephasing over the first moment as in the prior art. However, in the 0th-order GMN, the phase dispersion of the blood flow is insufficient, so that an image in which a low signal portion and a high signal portion with insufficient phase dispersion are mixed in the blood vessel is displayed. However, it is insufficient as compared with the first embodiment.

  According to an image obtained by relatively long TE, that is, an image showing both the effect of flow and magnetic susceptibility, information such as blood clots and bleeding can be obtained at the same time in addition to information on the state of blood vessels. It is possible to confirm not only the presence or absence of blood clots and bleeding, but also the position thereof, and important information for making a treatment plan can be obtained.

Even in the case of rephasing as in the prior art, the longer the TE, the more the local magnetic field inhomogeneity is emphasized. Therefore, the phase dispersion due to the magnetic susceptibility becomes larger in T ₂ ^* weighted imaging with a longer echo time. It is possible to obtain a T ₂ ^* weighted image with improved blood vessel rendering. However, in this case, the magnetic susceptibility is emphasized not only in the blood flow but also in the substantial part, and unevenness and distortion due to the inhomogeneity of the local magnetic field are conspicuous. Incurs extra time.

Further, the head T ₂ ^* weighted image obtained by a known general gradient echo method has good sensitivity to an image having a large magnetic susceptibility fluctuation such as bleeding. However, if the effect of blood flow is eliminated by performing rephasing, the blood flow is imaged as a high signal and the contrast cannot be clearly distinguished from the brain parenchyma and blood vessels. I do n’t know.

  By the way, since the phase shift data obtained from the echo obtained in the first embodiment is generated by the magnetic susceptibility and the flow, the magnetic susceptibility and the flow cannot be separated as they are. However, by subtracting the phase shift due to the difference in magnetic susceptibility for each tissue from the difference in magnetic susceptibility known in advance, it is possible to obtain phase shift data in which only the flow is dominant.

  As described above, the image obtained in the first embodiment is a non-contrast MRA (magnetic resonance angiography) image. That is, according to the first embodiment, a useful image as described above can be obtained without using a contrast agent, so that the feature of magnetic resonance imaging that is non-invasive can be maximized. is there.

  By the way, in SWI shown in Non-Patent Document 1, it is known that the polarity of the phase changes inside and outside the vein due to the influence of the magic angle effect depending on the angle between the static magnetic field direction and the blood vessel traveling direction. For this reason, in conventional SWI, when imaging is performed with minIP for a cross section orthogonal to the direction of the static magnetic field, the signal is reduced in the portion outside the vein in contact with the vein, so that one vein is parallel as shown in FIG. It will be drawn like two blood vessels. Therefore, when phase shift emphasis ignoring the polarity is performed, both the inside and outside of the vein are reduced in signal, so that the venous vascular cavity is rendered thicker than actual as shown in FIG. However, in the image obtained in the first embodiment, veins are faithfully depicted as shown in FIG. That is, in the first embodiment, an accurate and highly accurate image of the blood vessel cavity can be obtained even in a projection image of a cross section orthogonal to the static magnetic field direction typified by a coronal plane or a sagittal plane.

(Second Embodiment)
Hereinafter, the second embodiment will be described. This second embodiment corresponds to the second object.

  FIG. 12 is a diagram showing a pulse sequence in the second embodiment. The waveforms shown in FIG. 12 are, in order from the top, a radio frequency pulse (RF) applied to the imaging target, a gradient magnetic field waveform (Gss) in the slice direction, a gradient magnetic field waveform (Gpe) in the phase encoding direction, and a gradient magnetic field in the readout direction. A waveform (Gro) and an echo signal (Echo) are shown.

  As can be seen from FIG. 12, the pulse sequence in the second embodiment conforms to the gradient echo method. And unlike 2nd Embodiment, 2nd Embodiment rephases during echo time similarly to the nonpatent literature 1. FIG. However, in Non-Patent Document 1, the gradient magnetic field is applied so that only the first-order moment is zero, whereas in the second embodiment, both the first-order and second-order moments are zero. The difference is that a gradient magnetic field is applied. TE is set to a length that cancels the phase shift of the vein.

  Thus, according to the second embodiment, only the magnetic susceptibility becomes more dominant on the phase fluctuation, and the vein rendering ability in the phase data can be further improved.

  Note that the first and second moments do not have to be strictly zero, but may be values sufficient to cancel amplitude attenuation or phase shift due to flow.

  This embodiment can be variously modified as follows.

  In the first embodiment, it is not necessary to make the amount and polarity of dephase the same. Further, the rephasing may be performed on the remaining axes by setting the axes to be dephased as two axes or one axis instead of three axes. Further, de-phase may be performed only for 0 encoding instead of all encoding. Furthermore, the axis and the encoding step for performing dephase may be limited.

  In the second embodiment, the rephasing axis and the encoding step may be arbitrarily changed. Further, the third and higher moments may be set to a value sufficient to cancel the amplitude attenuation or phase shift due to the flow.

In the first and second embodiments, the pulse sequence may be two-dimensional (2D) instead of 3D. It is also possible to apply the pulse sequence of the asymmetric spin echo system in place of the pulse sequence of the gradient echo method. As shown in FIG. 13, the pulse sequence of the asymmetric spin echo system is the time Ta from the irradiation of the excitation pulse (usually 90 ° pulse) to the irradiation of the inversion pulse (usually 180 ° pulse) and the echo peak from the irradiation of the inversion pulse. The time Tb until is different from TE / 2. In the case of this asymmetric spin echo pulse sequence, the larger the time difference between the time Ta and the time Tb, the more the non-uniformity of the local magnetic field is reflected in the image contrast. Therefore, when the asymmetric spin echo pulse sequence is employed, the time difference between the time Ta and the time Tb is adjusted instead of adjusting the TE as in the first embodiment. If a pulse sequence of an asymmetric spin echo system is employed in this way, it is not necessary to change TE, so that it is possible to capture images with different effects of local magnetic field inhomogeneity while keeping the effect of T ₂ constant. Is possible.

  Note that the present invention is not limited to the above-described embodiment as it is, and can be embodied by modifying the constituent elements without departing from the scope of the invention in the implementation stage. In addition, various inventions can be formed by appropriately combining a plurality of components disclosed in the embodiment. For example, some components may be deleted from all the components shown in the embodiment. Furthermore, constituent elements over different embodiments may be appropriately combined.

The figure which shows the structure of the magnetic resonance imaging apparatus which concerns on the 1st and 2nd embodiment of this invention. The flowchart of the process in 1st Embodiment by the host computer 6 in FIG. The figure which shows the pulse sequence in 1st Embodiment. The figure showing the difference of the sequence of a rephase and a dephase with FIG. The figure showing the difference of the sequence of a rephase and a dephase with FIG. The figure which shows an example of the image obtained by the minIP image obtained by 1st Embodiment, and the conventional method. The figure which shows some examples of the minIP image produced | generated by changing b value. The figure which shows the example of adjustment of echo time. The figure which shows an example of the image acquired by minIP regarding the cross section orthogonal to the static magnetic field direction in the conventional SWI. The figure which shows an example of the image obtained by performing minIP regarding the cross section orthogonal to a static magnetic field direction for the image obtained by the phase shift emphasis which ignored the polarity in the conventional SWI. The figure which shows an example of the image obtained by 1st Embodiment. The figure which shows the pulse sequence in 2nd Embodiment. The figure which shows an example of the pulse sequence of an asymmetric spin echo system. The figure which shows the pulse sequence of a prior art example.

Explanation of symbols

  DESCRIPTION OF SYMBOLS 1 ... Magnet, 2 ... Static magnetic field power supply, 3 ... Gradient magnetic field coil unit, 3x, 3y, 3z ... Coil, 4 ... Gradient magnetic field power supply, 5 ... Sequencer, 6 ... Host computer, 7 ... RF coil, 8R ... Receiver, 8T: Transmitter, 10: Arithmetic unit, 11: Storage unit, 12: Display, 13: Input device, 14: Shim coil, 15: Shim coil power supply, 16 ... Sound generator, 17 ... ECG sensor, 18 ... ECG unit.

Claims (40)

Generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field and a readout gradient magnetic field along each of the slice axis, the phase encode axis and the readout axis;
A setting means for setting a dephasing amount for emphasizing a signal decrease due to arterial and venous flow in a region of interest of the subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis;
Control means for controlling the generating means with a gradient echo pulse sequence including a dephase gradient magnetic field pulse corresponding to the dephase amount with respect to the axis for which the dephase amount is set by the setting means. A magnetic resonance diagnostic imaging apparatus.   2. The dephasing amount is determined so as to emphasize a signal decrease due to an arterial and venous flow and to suppress a background signal decrease with respect to the artery and vein of the region of interest of the subject. The magnetic resonance diagnostic imaging apparatus described in 1.   The magnetic resonance imaging apparatus according to claim 1, further comprising a generating unit that generates an image related to the region of interest based on a magnetic resonance signal generated by nuclear magnetic resonance from the region of interest.   The magnetic resonance imaging apparatus according to claim 3, wherein the generation unit generates a cross-sectional image related to the region of interest by a projection process from a direction orthogonal to a static magnetic field direction.   4. The magnetic resonance imaging apparatus according to claim 3, wherein the image generated by the generating means is a non-contrast MRA (magnetic resonance angiography) image.   4. The magnetism according to claim 3, wherein the control unit controls the generation unit so as to have an echo time in which a difference in magnetic susceptibility is emphasized and reflected in a contrast of an image generated by the generation unit. Resonance diagnostic imaging device. A plurality of coil elements each receiving the magnetic resonance signal;
The said control means controls the said generation means to perform the parallel imaging which thins out the encoding step regarding the said phase encoding axis | shaft using the difference in sensitivity of these coil elements. Magnetic resonance imaging apparatus.   A plurality of amplitude images and a plurality of phase images are generated based on magnetic resonance signals received by each of the plurality of coil elements, a combined phase image is generated by combining the plurality of phase images, and the plurality of amplitude images The magnetic resonance diagnostic imaging apparatus according to claim 7, further comprising: a generating unit that generates an image in which a difference in phase shift is emphasized and reflected in contrast using the composite phase image.   2. The magnetic resonance imaging apparatus according to claim 1, wherein the control unit controls the generation unit so that a dephasing amount by the dephasing gradient magnetic field pulse is different for each encoding.   2. The magnetic resonance imaging apparatus according to claim 1, wherein the control unit controls the generation unit so that a dephasing amount by the dephasing gradient magnetic field pulse is different only in a part of encoding. Storage means for storing information representing a relationship between a plurality of portions to be emphasized for signal decrease and a dephasing amount for emphasizing signal decrease for each of the plurality of portions;
And a part determination means for determining a part to be targeted for emphasizing the signal drop,
2. The magnetic resonance imaging apparatus according to claim 1, wherein the setting unit sets a dephasing amount corresponding to a part determined by the part determination unit based on information stored in the storage unit. . Means for controlling the generating means to generate a plurality of sets of magnetic resonance signals for the same slice while varying the amount of dephase;
Means for generating a plurality of preparation images for the region of interest based on the plurality of magnetic resonance signals, respectively.
The magnetic resonance imaging apparatus according to claim 1, wherein the setting unit sets the dephase amount based on the plurality of preparation images. Means for controlling the generating means to generate a plurality of sets of magnetic resonance signals for the same slice while varying the amount of dephase;
Means for generating a plurality of preparation images for the region of interest based on the plurality of magnetic resonance signals, respectively.
Means for displaying the plurality of preparation images;
Image determination means for determining one preparation image desired by an operator among the plurality of preparation images;
2. The setting unit according to claim 1, wherein the setting unit sets a dephase amount corresponding to the one preparation image determined by the image determination unit as a dephase amount for emphasizing the signal decrease. Magnetic resonance imaging apparatus. ^2. The magnetic resonance imaging apparatus according to claim 1, wherein the setting unit sets the dephasing amount as a b value within a range of 0.1 sec / mm ^{2 to} 50 sec / mm ² . ^2. The magnetic resonance imaging apparatus according to claim 1, wherein the setting unit sets the dephase amount within a range of 1 sec / mm ^{2 to} 20 sec / mm ² as a b value.   4. The magnetic resonance imaging diagnosis according to claim 3, wherein the control unit performs the control so that an echo time that does not reflect inhomogeneity of a local magnetic field is applied to a contrast of an image generated by the generation unit. 5. apparatus.   4. The magnetic resonance imaging diagnosis according to claim 3, wherein the control unit performs the control so that an echo time that reflects a non-uniformity of a local magnetic field is reflected in a contrast of an image generated by the generation unit. 5. apparatus.   The generation unit generates at least one of an amplitude image based on an amplitude of the magnetic resonance signal, a phase image based on a phase of the magnetic resonance signal, or a combined image of the amplitude image and the phase image. The magnetic resonance imaging apparatus according to claim 3.   The magnetic resonance imaging apparatus according to claim 18, wherein the generation unit generates a three-dimensional image by performing a three-dimensional process on the amplitude image, the phase image, or the composite image.   The said generating means performs projection processing such as minimum value projection, maximum value projection or addition projection, volume rendering, surface rendering, or cross-sectional transformation as the three-dimensional processing. Magnetic resonance imaging apparatus.   Means for obtaining a phase shift corresponding to the flow of the tissue by subtracting a phase shift determined from a known magnetic susceptibility of the tissue from a phase shift obtained based on a magnetic resonance signal generated from the subject under the pulse sequence; The magnetic resonance imaging apparatus according to claim 1, further comprising:   The magnetic resonance imaging apparatus according to claim 1, wherein the pulse sequence is a spin warp method.   The magnetic resonance imaging apparatus according to claim 1, wherein the pulse sequence is an echo planar imaging method.   The magnetic resonance imaging apparatus according to claim 1, wherein the pulse sequence is an echo shift method.   The magnetic resonance imaging apparatus according to claim 1, wherein the pulse sequence is a multi-echo method. Generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field and a readout gradient magnetic field along each of the slice axis, the phase encode axis and the readout axis;
Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject A magnetic resonance diagnostic imaging apparatus comprising: control means for controlling the generating means so as to realize a gradient echo pulse sequence having a length suitable for canceling a phase shift of a vein.   27. The magnetic resonance imaging apparatus according to claim 26, wherein the pulse sequence is a spin warp method.   27. The magnetic resonance imaging apparatus according to claim 26, wherein the pulse sequence is an echo planar imaging method.   27. The magnetic resonance imaging apparatus according to claim 26, wherein the pulse sequence is an echo shift method.   27. The magnetic resonance imaging apparatus according to claim 26, wherein the pulse sequence is a multi-echo method. Generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field and a readout gradient magnetic field along each of the slice axis, the phase encode axis and the readout axis;
A setting means for setting a dephasing amount for emphasizing a signal decrease due to arterial and venous flow in a region of interest of the subject with respect to at least one of the slice axis, the phase encoding axis, and the readout axis;
Control means for controlling the generating means with a pulse sequence of an asymmetric spin echo system including a gradient magnetic field pulse corresponding to the dephase amount with respect to the axis where the dephase amount is set by the setting means. Magnetic resonance imaging diagnostic apparatus. Generating means for generating an image of the region of interest based on a magnetic resonance signal generated by nuclear magnetic resonance from the region of interest;
The control means calculates the time difference between the time from the irradiation of the excitation pulse to the irradiation of the inversion pulse and the time from the irradiation of the inversion pulse to the occurrence of the echo peak, as the contrast of the image generated by the generation means. 30. The magnetic resonance imaging apparatus according to claim 29, wherein the control is performed so that a value that does not reflect the inhomogeneity of the local magnetic field is applied to the image. Generating means for generating an image relating to the region of interest based on a magnetic resonance signal generated by nuclear magnetic resonance from the above;
The control means calculates the time difference between the time from the irradiation of the excitation pulse to the irradiation of the inversion pulse and the time from the irradiation of the inversion pulse to the occurrence of the echo peak, as the contrast of the image generated by the generation means. 30. The magnetic resonance imaging apparatus according to claim 29, wherein the control is performed so that the value reflects a non-uniformity of a local magnetic field. Generating means for generating a slice gradient magnetic field, a phase encode gradient magnetic field and a readout gradient magnetic field along each of the slice axis, the phase encode axis and the readout axis;
Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject A magnetic resonance diagnostic imaging apparatus comprising: control means for controlling the generating means so as to realize a pulse sequence of an asymmetric spin echo system having a length suitable for canceling a phase shift of a vein. Generating means for generating an image relating to the region of interest based on a magnetic resonance signal generated by nuclear magnetic resonance from the above;
The control means calculates the time difference between the time from the irradiation of the excitation pulse to the irradiation of the inversion pulse and the time from the irradiation of the inversion pulse to the occurrence of the echo peak, as the contrast of the image generated by the generation means. 35. The magnetic resonance imaging apparatus according to claim 34, wherein the control is performed so that the non-uniformity of the local magnetic field is not reflected on the magnetic field. Generating means for generating an image relating to the region of interest based on a magnetic resonance signal generated by nuclear magnetic resonance from the above;
The control means calculates the time difference between the time from the irradiation of the excitation pulse to the irradiation of the inversion pulse and the time from the irradiation of the inversion pulse to the occurrence of the echo peak, as the contrast of the image generated by the generation means. 35. The magnetic resonance imaging apparatus according to claim 34, wherein the control is performed so that the value reflects a non-uniformity of a local magnetic field. In a control method of a magnetic resonance imaging apparatus including a generation unit that generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis,
Setting at least one of the slice axis, the phase encoding axis, and the readout axis to set a dephase amount for emphasizing a signal decrease due to arterial and venous flow of the region of interest of the subject;
A control method characterized by controlling the generating means with a gradient echo pulse sequence including a dephasing gradient magnetic field pulse corresponding to the dephasing amount with respect to an axis where the dephasing amount is set. In a control method of a magnetic resonance imaging apparatus including a generation unit that generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis,
Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject A control method comprising controlling the generating means so as to realize a gradient echo pulse sequence having a length suitable for canceling a phase shift of a vein. In a control method of a magnetic resonance imaging apparatus including a generation unit that generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis,
Setting at least one of the slice axis, the phase encoding axis, and the readout axis to set a dephase amount for emphasizing a signal decrease due to arterial and venous flow of the region of interest of the subject;
A control method characterized in that the generating means is controlled by an asymmetric spin echo pulse sequence including a gradient magnetic field pulse corresponding to the dephase amount with respect to an axis on which the dephase amount is set. In a control method of a magnetic resonance imaging apparatus including a generation unit that generates a slice gradient magnetic field, a phase encode gradient magnetic field, and a readout gradient magnetic field along each of a slice axis, a phase encode axis, and a readout axis,
Re-spinning the first to nth (n is an integer of 2 or more) spins with respect to at least one of the slice axis, the phase encoding axis, and the readout axis, and the echo time of the region of interest of the subject A control method characterized by controlling the generating means so as to realize a pulse sequence of an asymmetric spin echo system having a length suitable for canceling a phase shift of a vein.

Images