JP5380585B2 - MRI equipment - Google Patents

Description

  The present invention relates to magnetic resonance imaging that images the inside of a subject based on the magnetic resonance phenomenon of spins (nuclear spins), and to obtain an arteriovenous phase image without using a contrast agent. ) Apparatus and MR (magnetic resonance) imaging method.

  Magnetic resonance imaging is an imaging method in which a nuclear spin of a subject placed in a static magnetic field is magnetically excited with a high-frequency signal of its Larmor frequency, and an image is reconstructed from an MR signal generated by the excitation. .

  In the field of magnetic resonance imaging, when an blood flow image of a lung field or abdomen is obtained, clinically, MR angiography in which a contrast medium is administered to a subject and angiography is started. However, this contrast-enhanced MR angiography method involves the administration of a contrast agent, and therefore requires invasive treatment. First of all, the burden on the patient's mental and physical strength is great. Also, the inspection cost is high. Furthermore, the contrast agent may not be administered depending on the patient's constitution.

  On the other hand, time-of-flight (TOF) method, phase contrast (PC) method, and the like are known as alternatives to contrast MR angiography.

  Among these, the time-of-flight method and the phase contrast method are methods that utilize the effect of a flow such as blood flow. The flow effect is caused by one of two properties of the moving spin. One is that the spin simply moves its position, and the second depends on the phase shift of the transverse magnetization caused by the movement of the spin in the gradient magnetic field. Among them, the method based on the former position movement is the TOF method, and the method based on the latter phase shift is the phase contrast method.

  However, since the TOF method and the phase contrast method described above are methods that use the effect of the flow of fluid such as blood, depending on the performance of the MRI apparatus, the speed is generally 2 to 2. Only a blood flow of 3 cm / s or more could be depicted, and a flow at a velocity lower than this could hardly be detected. For example, the flow of peripheral veins, lymphatic vessels, CSF (spinal fluid), pancreatic ducts, etc. of a patient (person) is slow and most speeds are 1 cm / s or less. In addition, since there is an influence of displacement due to pulsation or the like, the flow of these low-speed fluids cannot be detected conventionally.

  Even in the above-described TOF method or phase contrast method, it is necessary to image a slice perpendicular to the blood flow direction. That is, it is necessary to take an axial image with the slice direction aligned with the vertical direction. For this reason, in the case of a two-dimensional slice image, the image does not follow the blood flow. For this reason, when trying to obtain a three-dimensional image, there is a problem that the number of slices increases, and the entire imaging time becomes longer.

  The present invention was made in order to break down the current state of the prior art, and it is possible to reliably depict a low flow rate seen in the blood flow of the lower limbs without administering a contrast agent. The first purpose.

  In addition, the second object of the present invention is to reliably depict a low flow rate flow seen in the blood flow of the lower limbs with high image quality by short-time imaging without administering a contrast medium.

(1) An MRI apparatus according to an embodiment executes a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field, and includes a signal acquisition unit and an image generation unit. . The signal collecting means is configured such that a dephasing pulse for dephasing a magnetization spin of blood in the subject in accordance with a blood flow velocity is applied in a direction in which a read gradient magnetic field pulse for reading an echo signal from the subject is read. The echo signal is collected by executing the scan so that it is applied in step (b). The image generation means generates an arterial and vein separated image of the blood based on the echo signal.

(2) An MRI apparatus according to another embodiment executes a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field, and receives an echo signal from a lower limb of the subject. Echo signals are collected by performing a scan so that a dephasing pulse that dephases the magnetization spin of blood in the lower limb is applied with an intensity according to the blood flow velocity in the direction of application of the readout gradient magnetic field pulse for reading. A signal collecting unit; and an image generating unit configured to generate a blood flow image of the lower limb based on an echo signal, wherein the signal collecting unit sets a cardiac time phase in a systole as a first cardiac time phase of the subject. In addition, the diastolic cardiac time phase is set as the second cardiac time phase of the subject, and the first scan is executed in the first cardiac time phase during imaging of the same slice or slice encoding amount. Action Repeats the operations and for executing the second scan in the second cardiac time phase alternating.
(3) An MRI apparatus according to another embodiment executes a scan based on a pulse sequence including a read gradient magnetic field pulse on a subject placed in a static magnetic field, and receives an echo signal from a lower limb of the subject. Echo signals are collected by performing a scan so that a dephasing pulse that dephases the magnetization spin of blood in the lower limb is applied with an intensity according to the blood flow velocity in the direction of application of the readout gradient magnetic field pulse for reading. A signal collecting unit; and an image generating unit configured to generate a blood flow image of the lower limb based on the echo signal. The signal collecting unit collects echo data for filling a part of the k space by the first scan and The echo data that fills the entire k space is collected by the second scan, and the image generation means arranges the echo data of the first scan in the first k space. Both echo data of the second scan are arranged in the second k space, and are areas that are part of the second k space and correspond to the uncollected areas with respect to the uncollected areas on the first k space. Copy the echo data .

( 4 ) An MRI apparatus according to another embodiment executes a scan based on a pulse sequence including a read gradient magnetic field pulse on a subject placed in a static magnetic field, and reads an echo signal from the subject. The first cardiac phase of the subject is applied such that a dephasing pulse for dephasing the magnetization spin of the fluid in the subject in the direction of application of the readout gradient magnetic field pulse is applied at an intensity corresponding to the fluid velocity. Collecting means for collecting echo signals by executing the first scan at the second cardiac phase of the subject and the second scan at the second cardiac phase of the subject, and image generating means for generating an image of the fluid based on the echo signals With. In the MRI apparatus of this form, the signal collecting means executes the second scan continuously with the first scan.

( 5 ) An MRI apparatus according to another embodiment executes a scan based on a pulse sequence including a read gradient magnetic field pulse on a subject placed in a static magnetic field, and reads an echo signal from the subject. The echo signal is generated by performing a scan so that a dephasing pulse that dephases the magnetization spin of the moving fluid in the subject in the direction in which the readout gradient magnetic field pulse is applied is applied with an intensity corresponding to the fluid velocity. Signal collecting means for collecting, image generating means for generating an image of the fluid based on the echo signal, and means for receiving input of information specifying the intensity of the dephasing pulse or the fluid. In this type of MRI apparatus, the signal acquisition means controls the intensity of the dephasing pulse in accordance with the intensity of the input dephasing pulse or the fluid.

  According to the MRI apparatus of the present invention, for example, an imaging scan is performed in synchronization with different cardiac time phases, or a dephasing pulse or a rephasing pulse is added to a readout gradient magnetic field pulse, so that a contrast medium is administered. In addition, it is possible to reliably depict the low flow rate seen in the blood flow of the lower limbs. In particular, an image obtained by separating the arteries and veins of the lower limbs can be drawn in a short time and with high image quality.

The functional block diagram which shows the structural example of the MRI apparatus which concerns on embodiment. The figure explaining the time-sequential relationship of the ECG-prep scan and two imaging scans in 1st Embodiment. 6 is a schematic flowchart illustrating an ECG-prep scan procedure. The timing chart which illustrates the time-sequential relationship with respect to the ECG signal of an ECG-prep scan. The schematic image figure when delay time is changed dynamically obtained by ECG-prep scan. 4 is a schematic flowchart illustrating an example of first and second imaging scans. 4 is a schematic flowchart illustrating an example of first and second imaging scans. 6 is a timing chart illustrating the timing of the first and second imaging scans based on the electrocardiographic synchronization method in the first embodiment. The figure explaining the dephasing pulse and rephasing pulse added to a read gradient magnetic field. The figure explaining the positional relationship of a three-dimensional imaging region and the blood vessel to image. 4 is a schematic flowchart for explaining echo data calculation and display processing in the first embodiment. The schematic diagram explaining the outline | summary of the difference calculation for obtaining an arterial phase image. The schematic diagram explaining the outline | summary of the difference calculation for obtaining a vein phase image. The figure which illustrates the simultaneous display state of an arterial phase image and a venous phase image. The pulse sequence of two imaging scans implemented by the modification of 1st Embodiment. The figure explaining the time-sequential relationship of the ECG-prep scan which concerns on 2nd Embodiment, and one imaging scan. 3 is a schematic flowchart showing an example of an imaging scan. 3 is a schematic flowchart showing an example of a scanning scan. 6 is a timing chart illustrating the timing of an imaging scan based on an electrocardiogram synchronization method according to the first embodiment. The figure explaining a mode that the echo data each collected in the systole and the diastole are arranged in k space. 10 is a schematic flowchart for explaining echo data calculation and display processing according to the second embodiment.

(First embodiment)
A first embodiment will be described with reference to FIGS.
A schematic configuration of an MRI (magnetic resonance imaging) apparatus according to this embodiment is shown in FIG.

(1.1) Configuration of Apparatus This MRI apparatus includes a bed unit on which a subject P 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 high frequency A transmission / reception unit that transmits and receives signals, a control / calculation unit that controls the entire system and image reconstruction, and an electrocardiogram measurement unit that measures an ECG signal as a signal representing the cardiac time phase of the subject P are provided. .

The static magnetic field generation unit includes, for example, a superconducting magnet 1 and a static magnetic field power supply 2 that supplies current to the magnet 1, and an axial direction of a cylindrical opening (diagnosis space) into which the subject P is loosely inserted. (Z-axis direction) to generate a static magnetic field _{H 0.} In addition, the shim coil 14 is provided in this magnet part. The shim coil 14 is supplied with a current for homogenizing a static magnetic field from a shim coil power supply 15 under the control of a host computer to be described later. The couch portion can removably insert the top plate on which the subject P is placed into the opening of the magnet 1.

  The gradient magnetic field generator includes a gradient magnetic field coil unit 3 incorporated in the magnet 1. The gradient coil unit 3 includes three sets (types) of x, y, and z coils 3x to 3z for generating gradient magnetic fields in the X-axis direction, the Y-axis direction, and the Z-axis direction orthogonal to each other. The gradient magnetic field unit also includes a gradient magnetic field power supply 4 that supplies current to the x, y, and z coils 3x to 3z. The gradient magnetic field power supply 4 supplies a pulse current for generating gradient magnetic fields to the x, y, z coils 3x to 3z under the control of a sequencer 5 described later.

By controlling the pulse current supplied to the x, y, z coils 3x to 3z from the gradient magnetic field power source 4, the gradient magnetic fields in the three axes (X axis, Y axis, Z axis) directions which are physical axes are synthesized. it can be arbitrarily set and change the logical axial consisting slicing direction gradient magnetic field G _S, the phase encode direction gradient magnetic field G _E, and readout direction (frequency encode direction) gradient magnetic field G _R which are orthogonal to each other. Slice direction, phase encoding direction, and gradient magnetic fields in the readout direction are superimposed on the static magnetic field H _0.

  The transmission / reception unit includes an RF coil 7 disposed in the vicinity of the subject P in the imaging space in the magnet 1, and a transmitter 8T and a receiver 8R connected to the coil 7. The transmitter 8T and the receiver 8R operate under the control of the sequencer 5 described later. The transmitter 8T supplies the RF coil 7 with an RF current pulse having a Larmor frequency for exciting nuclear magnetic resonance (NMR). The receiver 8R takes in an MR signal (high frequency signal) such as an echo signal received by the RF coil 7, and performs various signal processing such as preamplification, intermediate frequency conversion, phase detection, low frequency amplification, and filtering. After that, A / D conversion is performed to generate digital data (original data) of the MR signal.

  Further, the control / arithmetic unit includes a sequencer (also called a sequence controller) 5, a host computer 6, an arithmetic unit 10, a storage unit 11, a display device 12, an input device 13, and a sound generator 16. Among these, the host computer 6 has a function of instructing the pulse sequence information to the sequencer 5 according to the stored software procedure (not shown) and supervising the operation of the entire apparatus.

  This MRI apparatus is characterized by performing MR scan based on an electrocardiographic synchronization method based on a synchronization timing (cardiac time phase) of a preselected value. As shown in FIG. 2, the host computer 6 performs a preparation scan (hereinafter referred to as an ECG-prep scan) for executing a preparation pulse sequence for determining synchronization timing in advance, and electrocardiographic synchronization based on the synchronization timing. Two imaging scans (hereinafter referred to as imaging scans) for executing the imaging pulse sequence are performed while executing a main program (not shown). An example of the ECG-prep scan execution routine is shown in FIG. 3, and examples of the first and second imaging scan execution routines based on ECG synchronization are shown in FIGS.

  The sequencer 5 includes a CPU and a memory, stores pulse sequence information sent from the host computer 6, controls the operations of the gradient magnetic field power source 4, the transmitter 8T, and the receiver 8R according to this information, The digital data of the MR signal output from the receiver 8R is once inputted and transferred to the arithmetic unit 10. Here, the pulse sequence information is all information necessary for operating the gradient magnetic field power source 4, the transmitter 8T, and the receiver 8R according to a series of pulse sequences, for example, x, y, z coils 3x to 3z. Includes information on the intensity, application time, application timing, and the like of the pulse current applied to.

  The pulse sequence may be a two-dimensional (2D) scan or a three-dimensional scan (3D) as long as the Fourier transform method is applied. In addition, as a pulse train, the SE method, the fast SE method, the EPI (Echo Planar Imaging) method, the FASE (Fast Asymmetric SE) method (that is, the fast SE method combined with the half Fourier method) SE series pulse trains are suitable.

  Further, the arithmetic unit 10 inputs the digital data (also referred to as original data or raw data) output from the receiver 8R through the sequencer 5, and the digital data is input to the k space (also referred to as Fourier space or frequency space) by its internal memory. Data is arranged, and this data is subjected to two-dimensional or three-dimensional Fourier transform for each set to reconstruct image data in real space. In addition, the arithmetic unit can execute data composition processing and difference calculation processing (including weighted difference processing) as necessary. This synthesis processing includes processing for adding each pixel, maximum value projection (MIP) processing, and the like. As another example of the above synthesis process, the axes of a plurality of frames may be aligned in Fourier space, and the original data may be synthesized into one frame of original data. The addition processing includes simple addition processing, addition averaging processing, weighted addition processing, and the like.

  The storage unit 11 can store not only the reconstructed image data but also the image data that has been subjected to the above-described combining process and difference process. The display device 12 displays an image. Further, parameter information for selecting a synchronization timing desired by the surgeon, scan conditions, pulse sequences, information relating to image synthesis and difference calculation can be input to the host computer 6 via the input unit 13.

  The voice generator 16 can issue a breath holding start message and a breath holding end message as voice when instructed by the host computer 6.

  Further, the electrocardiogram measurement unit attaches to the body surface of the subject and detects an ECG signal as an electrical signal, and performs various processing including digitization processing on the sensor signal to perform the host computer 6 and the sequencer. And an ECG unit 18 for outputting to the PC. The measurement signal from the electrocardiogram measurement unit can be used as required by the sequencer 5 when performing each of the ECG-prep scan and the electrocardiographic synchronization imaging scan. Thereby, the synchronization timing of the electrocardiographic synchronization method can be set appropriately, and data can be collected by performing an electrocardiographic synchronization imaging scan based on the synchronization timing.

(1.2) ECG-prep scan Next, an optimal synchronization timing determination process by ECG-prep scan will be described with reference to FIGS.
The host computer 6 starts the ECG-prep scan shown in FIG. 3 in response to a command from the input device 13 while executing a predetermined main program (not shown).

First, the host computer 6 reads scan conditions and parameter information for executing the ECG-prep scan from the input device 13 (step S1 in the figure). The scan conditions include the type of scan, the pulse sequence, the application direction of the readout gradient magnetic field pulse, and the like. The parameter information includes an initial time T ₀ (here, an elapsed time from the peak value of the R wave in the ECG signal) for determining the synchronization timing (time phase) of ECG synchronization, a time increment in increments of time Δt, and the number of times The upper limit value of the counter CNT and the like are included, and these parameters can be arbitrarily set by the operator.

Next, the host computer 6 clears the number counter CNT that counts the number of executions of the sequence and the time increment parameter T _inc for determining the synchronization timing (CNT = 0, T _inc = 0: step S2). Thereafter, the host computer 6 sends message data to the sound generator 16 to cause the subject (patient) to perform a breath holding command such as “please hold your breath” (step S3). This breath holding is preferably performed in order to suppress the body movement of the subject during the execution of the ECG-prep scan. However, in some cases, the ECG-prep scan is executed without performing the breath holding. Also good.

  When the preparation is completed in this way, the host computer 6 sequentially executes the processes after step S4. Thereby, it shifts to scan execution while changing the synchronization timing of electrocardiogram synchronization.

Specifically, the delay time T _DL from the peak arrival time of the R wave is calculated by T _DL = T ₀ + T _inc (step S4). Next, the ECG signal subjected to signal processing by the ECG unit 18 is read, and it is determined whether or not an R wave peak value appears in the signal (step S5). This determination process is repeated until the R wave appears. When R wave appears (step S5, YES), the delay time T _DL of the time calculated in step S4 whether elapsed since the R-wave peak time is subsequently determined (step S6). This determination process is continued until the delay time T _DL elapses.

When the delay time _TDL elapses from the peak time of the R wave (step S6, YES), the sequencer 5 is instructed to start each pulse sequence (step S7: see FIG. 4). This pulse sequence is preferably set to the same type as an imaging pulse sequence described later, and is, for example, a 2D-FASE (Fast Asymmetric SE) method in which a high-speed SE method is combined with a half Fourier method. Of course, various sequences such as a high-speed SE method and an EPI method can be adopted for this sequence. In response to this command, the sequencer 5 starts executing the type of pulse sequence commanded by the operator, so that the region of the desired part of the subject is scanned. This ECG-prep scan may be performed by a two-dimensional (2D) scan, for example, when an imaging scan (main scan) for collecting image data is a three-dimensional (3D) method, or according to the region of the imaging scan. A three-dimensional scan may be performed. In this embodiment, the imaging scan is executed as a three-dimensional scan, but the ECG-prep scan is executed as a two-dimensional scan from the viewpoint of reducing the scan time. In view of the mission of ECG-prep scanning, a two-dimensional scan is sufficient.

After the sequence execution start command, the number counter CNT = CNT + 1 is calculated (step S8), and the time increment parameter T _inc = ΔT · CNT is calculated (step S9). As a result, the count value of the number counter CNT is incremented by 1 each time the execution of the pulse sequence is commanded, and the increment parameter T _inc for adjusting the synchronization timing is increased in proportion to the count value.

Next, it waits as it is until a predetermined period (for example, about 500 to 1000 msec) necessary for execution of each pulse sequence elapses (step S10). Further, it is determined whether or not the number counter CNT has reached a predetermined upper limit value (step S11). In order to optimize the synchronization timing, the number counter CNT = 5 is set when, for example, five two-dimensional images are taken while changing the delay time _TDL to various time values. If the number counter CNT has not reached the upper limit value (step S11, NO), the process returns to step S5 and the above-described process is repeated. On the other hand, when the number counter CNT reaches the upper limit value (step S11, YES), a command for releasing the breath holding is issued to the sound generator 16 (step S12), and the subsequent processing is returned to the main program. For example, the breath-holding voice message is "You can breathe."

When the above-described processes are sequentially executed, for example, the preparation pulse sequence is executed at the timing shown in FIG. For example, if the initial time T ₀ = 300 msec and time increment ΔT = 100 msec are commanded, the delay time T _DL = 300 msec for the first sequence, the delay time T _DL = 400 msec for the second sequence, the third time the delay time _T DL = 500 msec for it, ... delay time _{T DL} to attain synchronization timing condition is adjusted such.

For this reason, when the first R wave after the breath holding command reaches the peak value, the first scan IMG _prep1 based on, for example, the two-dimensional FASE method is predetermined after the delay time T _DL (= T ₀ ) from the arrival time. Echo signals are collected over time (500-1000 msec). Even when the next R wave appears during the continuation of the sequence, the sequence is continued without being involved in the appearance of the R wave because there is a standby process in step S10 of FIG. That is, the execution process of the sequence started in synchronization with a certain heartbeat is continued over the next heartbeat, and echo signals are collected.

Then, when the number counter CNT has not reached the predetermined value, the processes of steps S5 to S11 are executed again. Therefore, in the example of FIG. 4, when the third R wave appears and reaches the peak value, the second scan is performed when the delay time T _DL = T ₀ + T _inc = 400 msec has elapsed since this arrival time. IMG _prep2 continues for a predetermined time, and echo signals are collected in the same manner. After this scan, the next R wave appears, and when the delay time T _DL = T ₀ + 2 · T _inc = 500 msec, the third scan IMG _prep3 continues for a predetermined time, and echo signals are collected in the same way The Further, after this scan is finished, the next R wave appears, and when the delay time T _DL = T ₀ + 3 · T _inc = 600 msec elapses, the fourth scan IMG _prep4 continues for a predetermined time, and similarly an echo signal is _generated. Collected. This scan is continued a desired number of times, for example, 5 times, and echo data of the same cross section for a total of 5 frames (sheets) is collected.

  The echo data is sequentially sent to the arithmetic unit 10 via the receiver 8R and the sequencer 5. The arithmetic unit 10 reconstructs echo data in k space (frequency space) into real space image data by a two-dimensional Fourier transform method. This image data is stored in the storage unit 11 as blood flow image data. For example, in response to an operation signal from the input unit 13, the host computer 6 sequentially displays the blood flow image in a cine.

That is, as schematically shown in FIG. 5, for example, n two-dimensional coronal images having different time phases of the lower limbs are displayed. In this coronal image, an artery AR and a vein VE that flow in the vertical direction in the lower limb are positioned. However, the imaging timing, that is, “delay time T _DL = initial time T ₀ + Tinc · Δt” from the R wave is different for each image. The operator visually observes these images, and selects an image in which the artery AR and vein VE appear in the highest signal and an image in which only the vein appears in the highest signal. Among them, the systolic synchronization timing T _DL = T _DL1 is determined by the delay time T _DL1 corresponding to an image in which only the vein VE appears in a relatively high signal. The diastolic synchronization timing T _DL = T _DL2 is determined by the delay time T _DL2 corresponding to an image in which the arteries AR and veins VE appear in a relatively high signal.

Therefore, the surgeon visually observes a plurality of blood flow images picked up by dynamically changing the delay time T _DL in this way, and the optimal delay time T in each systole and diastole as two cardiac phases. _DL = _TDL1 and _TDL2 (synchronization timing) are determined, and the process of reflecting this delay time _TDL in the subsequent imaging scan is performed manually, for example.

When an image determined by visual observation is designated, the delay time _TDL given to the designated image is automatically stored as the optimum synchronization timing, and the timing _TDL is automatically read at the time of the imaging scan. You may build and install the software. As a result, the ECG synchronization timing automatic designation processing can be performed.

(1.3) Imaging Scan Next, the operation of two imaging scans (that is, two imaging operations) in this embodiment will be described with reference to FIGS.
The host computer 6 executes a predetermined main program (not shown), and as part of that, executes the imaging scan process shown in FIG. 6 in response to the operation information from the input unit 13.

Assume that the first imaging scan (imaging) is assigned to the systole. In this case, the host computer 6 first sets the optimal delay time T _DL (= T _DL1 or T _DL2 > T _DL1 ) for the systole determined by the operator through the above-described ECG-prep scan, for example, the input device 13. (Step S20).

Next, the host computer 6 scans the scanning conditions specified by the operator from the input device 13 (application direction of readout gradient magnetic field pulse, image size, number of scans, standby time between scans, pulse sequence corresponding to the scan region, etc.) and image processing Information on the method (MIP processing, difference processing, etc. In the case of difference processing, any of simple difference, weighted difference processing, addition processing, etc.) is input, and the information including the delay time _TDL is used as control data. Then, the control data is output to the sequencer 5 and the arithmetic unit 10 (step S21).

  Next, when it is determined that there is a notice of completion of preparation before scanning (step S22), a command to start breath holding is output to the sound generator 14 in step S23 (step S23). As a result, the voice generator 14 issues a voice message such as “please hold your breath” in the same manner as in the ECG-prep scan, so that the patient who hears it stops breathing (see FIG. 9). .

  Thereafter, the host computer 6 instructs the sequencer 5 to start the first (or second) imaging scan (step S24).

When the sequencer 5 receives this imaging scan start command (FIG. 7, step S24-1), the sequencer 5 starts reading the ECG signal (step S24-2), and determines the peak value of the R wave (reference waveform) in the ECG signal. The occurrence of the predetermined nth time is determined from the ECG trigger signal synchronized with the peak value (step S24-3). Here, the reason for waiting for the appearance of the R wave n times (for example, twice) is to estimate the time when the shift to breath holding is surely made. When a predetermined n-th R wave appears, it performs a process of waiting for the delay time _{T DL1} set (step S24-4).

The sequencer 5 executes the first imaging scan on the assumption that the time when the optimum delay time T _DL1 (or T _DL2 ) has passed is the optimum electrocardiographic synchronization timing (step S24-5). Specifically, the transmitter 8T and the gradient magnetic field power source 4 are driven in accordance with the pulse sequence information already stored, and a first imaging scan (imaging) based on, for example, a three-dimensional FASE pulse sequence is shown in FIG. As shown in (a) and (c), it is executed in synchronization with the electrocardiogram (the phase encoding direction gradient magnetic field is not shown in FIG. 10 (c)).

According to this pulse sequence, the application direction RO of the readout gradient pulse G _R, for example, as shown in FIG. 10, is set to match substantially in the direction of flow of the blood flow imaging purposes (artery AR, vein VE) ing.

Further, readout gradient pulses G _R contained in the pulse sequence as shown in FIG. 8 (c) and FIG. 9 (a) ~ (c) , a pulse body P _body for frequency encoding acquiring echo signals, the It consists of two dephasing pulses P _dephase as control pulses added continuously before and after the time of the pulse body P _body . This _dephase pulse P phase has the same polarity as the pulse body P _body for frequency encoding, and thereby has a function of promoting dephasing of the moving magnetic spin.

It should be noted that the _dephase pulse P phase hardly exhibits the dephase function for the magnetized spin that hardly moves. Thus, read-out gradient pulse G _R It is important that the applied substantially matches the direction of motion of the imaging subject fluid (blood or lymph).

Preferably, the intensity of the dephasing pulse P _dephase can be changed or controlled according to the speed of lymph or blood flow as a fluid to be imaged. FIGS. 9A to 9C illustrate examples of decreasing the intensity of the _dephase pulse P dephasing in this order. Generally, the blood pressure is changed or controlled so as to decrease the intensity of the _dephase pulse P phase as the blood flow velocity increases.

When the velocity of the fluid to be imaged (blood flow or the like) is relatively high, as shown in FIG. 9 (d), a total of two re-sequential control pulses continuously before and after the pulse body P _body are obtained. A phase pulse _Prephase is added. The _rephase pulse P _rephase is opposite in polarity to the frequency encoding pulse body P _body and has a function of rephasing the magnetized spin to suppress artifacts in order to suppress excessive dephase. The intensity of the _rephase pulse _Prephase is also preferably changed according to the flow rate.

In the first embodiment, therefore, both in the imaging scan of the first time and the second time (to be described later), dephasing pulse _{P dephase} or rephase _{P Rephase} is added to the readout gradient pulse _{G R.}

For this reason, by executing the above-described three-dimensional FASE pulse sequence, echo signals energized by the excitation 90 ° RF pulse and the refocus 180 ° RF pulse are collected for each slice encoding and each phase encoding. The This echo signal rephasing effect of the phase of the magnetization spin due to dephasing effect or rephasing _{P Rephase} magnetization spins in phase due to dephasing pulse _{P dephase} is reflected.
This will be described in detail together with the display operation described later, and the outline thereof will be described as follows.

In other words, for a fluid flowing along the application direction of the readout gradient magnetic field pulse, dephasing effect due to dephasing pulse _{P dephase} leads to the promotion of flow voids (flow void) effect. For this reason, the intensity of the echo signal is reduced by the dephasing pulse. On the other hand, in the case of fluid that hardly flows in that direction, the degree of promotion of the flow void effect due to the _dephase pulse P _dephase is low, and the intensity of the echo signal does not decrease so much.

In the case of the _rephase pulse _Prephase , the effect of dephasing is suppressed according to the flow of the fluid due to the _rephase action.

  The echo interval in the above pulse sequence is shortened to about 5 msec. Thus, echo signals are collected from the three-dimensional imaging region Rima set on the lower limb, for example, as shown in FIG. 10 with a scan time of about 600 msec based on the first slice encoding amount SE1.

  When the scan based on the first slice encoding is completed, the sequencer 5 determines whether or not the scan for the final slice encoding is completed (step S24-6), and this determination is NO (the scan based on the final slice encoding is performed). In the case of not completed), while monitoring the ECG signal, for example, from the R wave used in the previous imaging scan, for example, 2 heartbeats (2R-R) and wait until a shorter set period elapses (step S24- 7). The repetition time TR is set to 4 heartbeats (4R-R) or less.

Thus, for example, after waiting for a period corresponding to 2R-R, for example, when the third R wave appears (step S24-7, YES), the sequencer 5 returns the process to step S24-4 described above. As a result, when the specified delay time _TDL1 has elapsed from the ECG trigger signal synchronized with the third R-wave peak value, a scan based on the next slice encoding amount SE2 is executed in the same manner as described above, and the three-dimensional imaging region Echo signals are collected from Rima (steps S24-4, 5). Similarly, echo signals are collected up to the final slice encoding amount SEn (for example, n = 8).

  When the final scan based on the slice encoding amount SEn is completed, the determination in step S24-6 is YES, and the completion notification of the first (or second) imaging scan is output from the sequencer 5 to the host computer 6 (step S24-6). S24-8). As a result, the processing is returned to the host computer 6.

  Upon receiving the scan completion notification from the sequencer 5 (step S25), the host computer 6 outputs a breath holding release command to the voice generator 16 (step S26). Therefore, the voice generator 16 issues a voice message such as “It is fine to breathe” to the patient, and the breath holding period ends (see FIG. 8).

  Thereby, the 1st (or 2nd) imaging scan (imaging) by an electrocardiogram synchronization is performed for every 2R-R based on 3D-FASE method, for example.

Echo signal generated from the patient P, for each slice encoding subjecting the slice gradient magnetic field pulse G _s, is received by the RF coil 7, and sent to the receiver 8R. The receiver 8R performs various preprocessing on the echo signal and converts it into a digital quantity. This digital amount of echo data is sent to the arithmetic unit 10 through the sequencer 5 and is arranged at a position corresponding to the encoded amount of the three-dimensional k-space formed by the memory.

Next, as shown in FIG. 2, the second imaging scan (imaging) for the diastole is performed in the same manner as the first time with an appropriate time interval. However, in the second time, the optimum delay time _TDL2 that determines the predetermined time phase of the expansion period set in advance through the ECG-prep scan described above is read (FIG. 6, steps S20 and S21), and this delay time T ECG synchronization based on _DL2 is taken (FIG. 7, step S24-4).

For this reason, in the case of the second imaging scan, as shown in FIGS. 8B and 8C, it depends on each phase encoding amount SE at the diastolic synchronization timing delayed by the delay time _TDL2 from the R wave peak. A three-dimensional FASE scan is performed. Again, the application direction of the readout gradient pulse G _R almost match the direction of motion of the imaging fluid such as blood flow. Moreover, the readout gradient pulses _{G R,} the control pulses for controlling the behavior of the magnetization spin (dephasing or rephasing) (dephasing pulse _{P dephase} or rephasing pulse _{P Rephase)} is added.

Accordingly, the second time imaging scan, like the first time, the read gradient pulse G dephasing pulses added to the _R P _dephase or image data rephasing pulse P diastolic reflecting the spin control functions _rephase Is obtained.

(1.4) Data processing and image display When the collection of echo data is completed in this way, the host computer 6 causes the arithmetic unit 10 to execute the processing shown in FIG.
As shown in the figure, in response to the command from the host computer 6, the arithmetic unit 6 causes both the systolic k-space and the diastole k-space to calculate echo data based on the half Fourier method (step S31). . That is, the data of the remaining area of the k space for which the echo data has not been collected is calculated according to the complex conjugate relation and arranged. As a result, both the k spaces are filled with echo data.

Thereafter, the arithmetic unit 10 performs image reconstruction by performing three-dimensional Fourier transform on the echo data of the systolic k-space and the diastole k-space, respectively (steps S32 and S33). As a result, as shown in FIGS. 12A and 12B, a time phase image (systolic image) IM _sys in the delay time T _{DL1 in} the systole and a time phase image (expansion) in the delay time T _{DL2 in} the diastole Period image) Three-dimensional data of IM _dia is obtained.

According to this image data, only the vein VE is reflected in the systolic image IM _sys , and the artery AR is hardly reflected. On the other hand, the arterial AR and the vein VE are reflected in the diastolic image _IMdia to some extent.

Here, the principle of such a systolic image _{IM sys} and diastolic image _{IM dia} is obtained, described in detail below from the function of the applied direction and the dephasing pulses _{P dephase} the readout gradient _{G R} described above.

  The phase of magnetization spins of components such as blood flow flowing in the direction in which the read gradient magnetic field pulse is applied is more easily dispersed by the dephase pulse. In other words, for the flowing component, this is equivalent to the promotion of the flow void effect due to the flow itself. On the contrary, the rephasing function acts on the phase of the magnetized spin such as blood flow due to the rephasing pulse.

For example, the lower limb of the subject is taken as an example. In the case of the lower limb, the artery in the systole usually has a low flow rate of 1 cm / s or less, and the flow velocity is low enough to be considered that the vein in the systole and the artery and vein in the diastole are hardly moving. For this leg, as shown in FIG. 8, the imaging scan at a desired time phase each systolic and diastolic s using the read gradient pulse G _R by adding a dephasing pulse P _dephase (imaging) is performed.

  These imaging scans excite the arteriovenous magnetization spins and collect echo signals. At this time, since the arterial and venous flow velocities are slightly different from each other, the difference in the flow velocities is reflected in the promotion of the flow void effect due to the rephasing pulse, and appears as a relative change in the signal value of the echo signal.

  Specifically, the systole is as follows. Since the vein flows very slowly, the echo signal is slightly reduced by the dephasing pulse, but the flow void effect is small, and the blood is depicted in a bright blood with a relatively high signal value. On the other hand, the systolic artery flows at a higher flow rate than the vein, so that the flow void effect promoted by the dephasing pulse is greater than that of the vein. As a result, the signal value of the artery is greatly reduced, and is rendered in black blood. This state is schematically shown in FIG. In the figure, the hatched portion is bright blood and the dotted line portion is black blood.

  On the other hand, in the diastole, both the arteries and veins move only at a very low flow rate, so that both the arteries and veins are depicted in bright blood, although there is a slight signal value decrease due to the dephasing pulse. This state is schematically shown in FIG.

Returning to the description of FIG. 11, the arithmetic unit 10 performs the difference calculation “IM _dia −β · IM _sys ” for each pixel for the systolic image IM _sys and the diastolic image IM _dia to obtain the arterial phase image IM _AR. This is performed (step S34). Here, β is a weighting coefficient. As a result, as shown in FIG. 12, by appropriately setting the weighting coefficient β, the image data of the vein VE becomes almost zero, and three-dimensional image data of the arterial phase image IM _{AR in} which only the artery AR is shown is obtained. It is done.

Further, in order to obtain the vein phase image IM _VE , the difference calculation “IM _dia −IM _AR ” is performed for each pixel (step S35). Image data IM _AR is image data calculated by the above-described weighted difference. As a result, as shown in FIG. 13, the image data of the artery AR becomes almost zero, and the three-dimensional image data of the vein phase image IM _VE showing only the vein VE is obtained. Note that this difference calculation may also be performed using a weighted difference.

When the difference calculation is thus completed, the calculation unit 10 performs MIP (maximum value projection) processing on both the arterial phase image IM _AR and the venous phase image IM _VE and observes those blood vessels from the desired direction. The data of the time two-dimensional image (for example, coronal image) is created (step S36).

The two-dimensional images IM _AR and IM _VE of the arterial phase and the venous phase are displayed on the display unit 12 as shown in FIG. 14, for example, and the image data is stored in the storage unit 11 (step S39).

In this display, in addition to the arterial phase image IM _AR and the venous phase image IM _VE , the systolic image IM _sys and the diastolic image IM _dia may be displayed on the same screen or on the screen of a separate monitor. .

As (1.5) Effects In the above description, in the MRI apparatus of this embodiment, during an imaging scan, a readout gradient pulse G _R in the direction of fluid flow of the low flow rates such as those found in the lower extremity vascular (such as blood flow) They are applied almost together. Moreover, it appended simultaneously, a dephasing pulse _{P dephase} or rephasing pulse _{P Rephase} the gradient pulses _{G R.}

Thereby, the relative signal value difference between the flowing fluid and the fluid flowing only at a lower velocity can be increased by the _dephasing pulse P _dephase or the _rephasing pulse P _rephase . Therefore, for example, when the dephasing pulse is used, even in the lower limb blood vessel whose flow velocity is lower than that of the abdomen or chest, the arteriovenous vein is clearly separated from the relative signal value difference as shown in FIG. It can be displayed with high rendering ability.

  In this way, in order to give a relative difference in the signal value between the arteriovenous vein, the read gradient magnetic field pulse is applied in the same direction as the flow direction, and the dephasing and rephasing of the magnetized spin are actively used. The technique for controlling the flow void effect is a novel technique first developed by the present inventors.

  In this embodiment, since the optimal ECG synchronization timing for the systole and the diastole is set in advance by the ECG-prep scan, the target blood flow is reliably captured in each time phase of the systole and the diastole. be able to. Furthermore, the optimal setting of the electrocardiogram synchronization timing eliminates the need for re-imaging, thereby reducing the operational burden on the operator and the physical and mental burden on the patient.

  In addition, since the slice direction or slice encode direction can be taken in directions other than the patient's vertical direction, the entire scan time is compared with a method of imaging in the vertical direction perpendicular to the blood flow, such as the TOF method. It's short. This reduces the burden on the patient and increases patient throughput.

  Furthermore, since it is not necessary to administer a contrast medium, non-invasive imaging can be performed, and the mental and physical burden on the patient is remarkably reduced from this point. At the same time, it is freed from the troublesomeness inherent in contrast methods, such as the need to measure the timing of contrast effects, and unlike contrast methods, it enables repeated imaging as required.

(Modification of the first embodiment)
In the above embodiment, as shown in FIG. 8, the first time and the second time imaging scan both added dephasing pulses _{P dephase} or rephasing pulse _{P Rephase} to the read gradient pulse _{G R} Described in the aspect.

As a modified example of this mode, a _dephase pulse P phase is added to the first imaging scan performed in the diastolic time phase as shown in FIG. 15 (a), while the second time is performed in the systolic time phase. A _rephase pulse P _rephase shown in FIG. 4B may be added to the imaging scan.

  That is, the type of control pulse that additionally controls the behavior of the magnetized spin is changed according to the systole and the diastole. Thereby, the signal value can be increased by reflecting the effect of rephase (flow compensation) in the expansion period, and the S / N can be improved.

(Second Embodiment)
Next, a second embodiment according to the present invention will be described with reference to FIGS. Note that the hardware configuration of the MRI apparatus used in this embodiment is the same as or equivalent to that of the first embodiment.

  In the second embodiment, the first imaging scan and the second imaging scan executed in the first embodiment are executed in one imaging scan, and the systolic and diastolic phases of the cardiac cycle are executed. In addition, the present invention relates to a configuration in which the aforementioned dephasing pulse and rephasing pulse are used properly.

Assume that a separated image of the arteries and veins of the lower limb is obtained as a low-speed fluid. As shown in FIG. 16, an ECG-prep scan is first performed, and then one imaging scan is performed under the electrocardiographic synchronization method. The ECG-prep scan is performed by the method described in the first embodiment, and thereby, delay times T _DL1 and T _DL2 from the R wave that provide the most visualization capability in the systole and the diastole are set, respectively. .

Next, an imaging scan based on the electrocardiographic synchronization method based on the delay times T _DL1 and T _DL2 is executed as one imaging. The procedure of this imaging scan is illustrated in FIGS. 17 and 18, and the pulse sequence used for this scan is illustrated in 19.

(2.1) Imaging Scan While executing a predetermined main program (not shown), the host computer 6 executes the processing shown in FIGS. 17 and 18 in response to operation information from the input device 13 as part of it. To do.

In detail, the host computer 6 firstly determines the optimal two delay times T _DL determined by the operator through the above-described ECG-prep scan (that is, the optimal delay time T _{DL1 of} systole and the optimal delay time of diastole). The delay time T _DL2 (> T _DL1 )) is input, for example, via the input device 13 (step S120). Information on the optimum delay times T _DL1 and T _DL2 may be stored in advance in the storage unit 11, for example.

Next, the host computer 6 inputs information on the scanning conditions and the image processing method, processes the information including the delay times T _DL1 and T _DL2 into control data, and the control data is necessary for the sequencer 5 and the arithmetic unit 10. In response to (step S121).

  Next, as in the first embodiment, when the host computer 6 determines that preparation before scanning is completed, the host computer 6 is instructed to start breath holding, and is instructed to start imaging scanning (steps S123 to S124).

  When the sequencer 5 receives a command to start an imaging scan (FIG. 18: Step S124-1), the sequencer 5 starts reading the ECG signal (Step S124-2), and determines a predetermined peak value of the R wave (reference waveform) in the ECG signal. The appearance of the nth time is determined from the ECG trigger signal synchronized with the peak value (step S124-3).

When a predetermined n-th R wave appears, first, the process to wait for the delay time T _DL1 set for a specific time phase of systole (step S124-4).

The sequencer 5 executes a scan for the systole, assuming that the optimal delay time _TDL1 has passed is the optimal ECG synchronization timing (step S124-5).

Specifically, the transmitter 8T and the gradient magnetic field power source 4 are driven according to the pulse sequence information that has already been stored, and for example, the first slice encode amount SE1 based on the pulse sequence of the first three-dimensional FASE method is used. The scan SN _sys1 is executed by the ECG synchronization method as shown in FIG.

In the first scan SN _sys1 , the read gradient magnetic field pulse GR is applied in the body axis direction substantially along the arteriovenous veins of the lower limbs. Further, this read-out gradient pulse G _R, dephasing pulse P _dephase dispersing the magnetic spins of the phase as shown, are added before and after sequentially in time. Further, the echo interval in this pulse sequence is shortened to about 5 msec.

In the pulse sequence used for the first scan SN _sysn for this systole, the number of echoes is set short as shown in FIG. 19, and ends in a short time within one heartbeat from the start of the scan. As schematically shown in FIG. 20, the number of echoes is set so as to be sufficient to collect the echo data arranged only in the central portion (low frequency region) in the phase encoding ke direction of the k space for each slice encoding amount. . For this reason, the second scan SN _dian for the _diastole can be started within the same heartbeat as the first scan SN _sysn for the systole, as shown in FIGS. In addition, echo data that is insufficient in the systolic k space (first k space) K _sys is obtained by copying from an expansion k space (second k space) K _dia described later and an operation based on the half Fourier method. (See FIG. 20).

  As a result, echo signals are collected from the three-dimensional imaging region Rima (see FIG. 10) set on the lower limbs with a short scan time of about several hundreds msec under the initial slice encoding amount SE1.

Next, the sequencer 5 shifts to scan control in the expansion period. Specifically, a process to wait for the delay time _{T DL2} set for a specific time phase of diastole (step S124-6).

The sequencer 5 executes the second scan for the diastole, assuming that the time when the optimum delay time _TDL2 has elapsed is the optimum ECG synchronization timing (step S124-7). Specifically, the transmitter 8T and the gradient magnetic field power source 4 are driven according to the pulse sequence information that has already been stored, and for example, the second slice encoding amount SE1 based on a pulse sequence of the first three-dimensional FASE method is used. The scan SN _dia1 is executed by the ECG synchronization method as shown in FIG.

Also in the second scan _SNdia1 , the readout gradient magnetic field pulse GR is applied in the body axis direction substantially along the arteriovenous veins of the lower limbs. Further, this read-out gradient pulse G _R, rephasing pulse P _Rephase to rephase the magnetization spin as shown, are added before and after sequentially in time. The echo interval in this pulse sequence is also shortened to about 5 msec.

Note that the pulse sequence used for the second scan SN _dian in this diastole is larger than that in the systole, as shown in FIG. It is set to collect a small number of echoes. As schematically shown in FIG. 20, the number of echoes is collected for each slice encode amount by collecting echo data arranged only in the center (low frequency region) and one end (high frequency) of the phase encoding ke direction in the k space. It is set to be sufficient. In the diastole k-space _Kdia , the missing echo data is obtained by calculation based on the half Fourier method, as will be described later. As shown in FIGS. 19 and 20, the scan SN _dia1 in the diastole is usually scanned over the next heartbeat.

  As a result, echo signals are collected from the three-dimensional imaging region Rima (see FIG. 10) set on the lower limbs with a scan time of about 600 msec based on the first slice encoding amount SE1.

  When these first imaging scans are completed, the sequencer 5 determines whether or not the final scan has been completed (step S124-8). If this determination is NO (the final scan has not been completed), the ECG signal Monitoring, for example, from the R wave used in the imaging scan, for example, 2 heartbeats (2R-R), and wait for a short period of time to elapse, and recover the longitudinal magnetization of the stationary part of the spin. It actively suppresses (step S124-9).

  Thus, for example, after waiting for a period corresponding to 2R-R, for example, when the third R wave appears from the start of scanning (YES in step S124-9), the sequencer 5 performs the process in step S124-4 described above. return.

As a result, when the delay time _TDL1 has elapsed from the third R-wave peak value, the first scan SN _sys2 for the second systole is executed in the same manner as described above according to the next slice encoding amount SE2. Echo signals are collected from the three-dimensional imaging region Rima (steps S124-4 and 5). Further, when the delay time T _DL2 has elapsed from the third R-wave peak value, the second scan SN _dia2 for the second _diastole is executed in the same manner as described above according to the slice encoding amount SE2, and three-dimensional imaging is performed. Echo signals are collected from the region Rima (steps S124-6, 7).

  Similarly, echo signals are collected for the systole and the diastole until the final slice encoding amount SEn (for example, n = 8).

When the final scans SN _sysn and SN _dian with the slice encoding amount SEn are completed, the determination in step S124-8 is YES, and the sequencer 5 outputs an imaging scan completion notification to the host computer 6 (step S124-10). ). As a result, the processing is returned to the host computer 6.

  When the host computer 6 receives the scan completion notification from the sequencer 5 (FIG. 17: step S125), it outputs a breath holding release command to the voice generator 16 (step S126).

  Therefore, as schematically shown in FIG. 19, in one imaging scan (imaging), for example, every 2R-R, the electrocardiographic scan for the systole and the diastole is performed by n 3D-FASE method, for example. It is executed for the slice encoding amount.

The echo signal generated from the patient P is converted into digital amount of echo data as in the first embodiment. This echo data is sent to the arithmetic unit 10 through the sequencer 5, and is arranged corresponding to the phase encoding amount and the slice encoding amount in the three-dimensional k-spaces K _sys and K _dia for the systole and diastole formed by the memory. The

(2.2) Data Processing and Image Display When the echo data collection is thus completed, the host computer 6 instructs the arithmetic unit 10 to execute the processing shown in FIG.

As shown in FIG. 21, in response to a command from the host computer 6, the arithmetic unit 6 completes all data arrangement in the systolic k-space K _sys and the diastole k-space K _dia (steps S131 and S132). Specifically, in step S131, the as shown in FIG. 20, the echo data of one of the high-frequency region Re of the phase encoding direction in the k-space _{K dia} for diastolic (In FIG. 20, the echo up number H-N) is contracted Copied to the corresponding position in the intended k-space _Ksys . This echo data is data of an area that was not collected by the systolic scan. Next, the process proceeds to step S132, and the data of the remaining area where echo data has not been collected by individually applying the half Fourier method to both the systolic k-space K _sys and the diastolic k-space K _dia. Is calculated according to the complex conjugate relation and arranged. Accordingly, both the k-spaces K _sys and K _dia are all filled with data through the processing of steps S131 and S132.

Thereafter, the arithmetic unit 10 performs image reconstruction by three-dimensional Fourier transform on the systolic k-space K _sys and the diastolic k-space K _dia (steps S133 and S134). As a result, as shown in FIGS. 12A and 12B, the image of the delay time T _{DL1 in} the systole (systolic image) IM _sys and the image of the delay time T _{DL2 in} the diastole (diastolic image) IM _Dia three-dimensional data can be obtained. According to this image data, only the vein VE is reflected in the systolic image IM _sys , and the artery AR is hardly reflected. On the other hand, the arterial AR and the vein VE are reflected in the diastolic image _IMdia to some extent.

Therefore, as in the first embodiment (see FIGS. 12 to 14), the arithmetic unit 10 generates the arterial phase image IM _AR by the difference calculation “IM _dia −β · IM _sys ” (step S <b> 135), the difference calculation “ Generation of venous phase image IM _VE by “IM _dia -IM _AR ” (step S136), MIP (maximum value projection) processing of each of arterial phase image IM _AR and venous phase image IM _VE (step S137), and arterial phase and vein Phase two-dimensional image display and image data storage (step S138) are sequentially executed.

(2.3) As described operational effects above, in the MRI apparatus of this embodiment, during the imaging scan, and applies substantially combined readout gradient pulse G _R in the direction of flow of the lower limb blood vessels, it is applied to the systole the gradient pulses _{G R} added dephasing pulses _{P dephase,} also are added respectively rephasing pulse _{P rephase} thereto to be applied to diastole.

  As a result, similar to the behavior control of the magnetization spin described in the first embodiment, the signal value can be lowered by giving a flow void promoting effect to the artery, particularly in the systole, while the diastole. The flow compensation effect can be positively given to the arteries and the arterial flow.

  Therefore, the relative signal value difference between the flowing blood flow and the blood flow flowing only at a lower speed is made significant by the dephasing pulse and the rephasing pulse, and the flow velocity is higher than that of the abdomen and the chest. Even in a low limb blood vessel, the arteriovenous vein can be clearly separated and displayed with high rendering ability based on the relative signal value difference.

  Further, according to the MRI apparatus of the present embodiment, the optimum scan start timing (delay time from the R wave) is set for each systole and diastole within one cardiac cycle. Then, the two-shot scans of the systole and the diastole for one slice encoding are alternately executed sequentially in one imaging scan. Moreover, the first systolic scan within one cardiac cycle shortens the data acquisition time (number of echoes) so that subsequent diastolic scans do not take time, and the collected echo data is used for the systolic phase. It is arranged in a low frequency region that is most important in terms of improving the contrast in the k space. Data deficient in systolic k-space is supplemented by copying data obtained in subsequent diastolic scans that can perform echo acquisition relatively long. In addition, the half-Fourier method is adopted for each scan for systole and diastole, and the scan time is set as short as possible.

  For this reason, the two-shot scan for the systole and the diastole for one-slice encoding usually falls within about two heartbeats. Thus, by repeating these scans alternately one after another, echo data of systolic and diastolic blood flows is collected within one breath holding continuation period in one imaging. That is, three-dimensional data of systolic and diastolic blood flows are collected individually and at optimum timing by one imaging.

Therefore, it is not necessary to individually perform imaging scans for the systole and the diastole (that is, to perform imaging twice in total), and only one imaging is necessary. Therefore, imaging time is significantly reduced and patient throughput is increased. In particular, the effect of shortening the imaging time becomes significant when performing three-dimensional imaging. In addition, since misregistration due to patient movement and the like can be greatly reduced, the quality of the presented image is improved. Further, since a blood flow image (MRA image) obtained by separating the arterial phase and the venous phase can be obtained from the two-phase echo data collected by one imaging, the imaging efficiency is good and the blood to be provided is provided. The flow information is also abundant.
Moreover, according to this 2nd Embodiment, the other effect obtained in 1st Embodiment can be enjoyed similarly.

(Modification of the second embodiment)
In the first and second imaging scans of the second embodiment described above, as shown in FIG. 19, a dephasing pulse is added to the readout gradient magnetic field pulse for systole, and the readout for diastole. A rephase pulse was added to the gradient magnetic field pulse. On the other hand, only the dephasing pulse may be added to the readout gradient magnetic field pulses for systole and diastole. As a result, as in the first embodiment (see FIG. 8), it is possible to reflect the degree of promotion of the flow void effect due to different blood flow velocities for each time phase in the intensity of the signal value. Can be reliably separated.

(Modification common to each embodiment)
In addition, this invention is not limited to the structure as described in each embodiment mentioned above, Furthermore, various deformation | transformation structures and applications are possible.
For example, in the above-described embodiment, both the arterial phase image and the venous phase image are presented. However, for this, only the arterial phase image may be calculated and displayed. That is, the difference calculation for the vein phase image in step S136 in FIG. 21 can be omitted. On the contrary, although the difference calculation of the images of the arterial phase and the venous phase is performed together, the displayed image may be only the arterial moving image.

  Further, in each of the above-described embodiments, the scan method using the half Fourier method is adopted for each of the scans for the systole and the diastole, but this half Fourier method is not necessarily adopted. Good. In this case, it is preferable that the k space is fully collected by the diastolic scan, and the echo data in the high frequency region at both ends of the slice encoding direction is copied to the corresponding region in the systolic k space.

  Furthermore, although the above-described embodiment has been described with respect to the case of performing a three-dimensional scan, this can be similarly applied to the imaging of a two-dimensional scan. The employed pulse sequence is not limited to the FASE method, and a sequence of FSE method or FASE method using an inversion recovery (IR) pulse may be employed.

  Further, the post-processing of the echo data according to the above-described embodiment is configured to convert the echo data into real-space image data once, and then perform a difference calculation to obtain images of the arterial phase and the venous phase. However, such a difference calculation may be performed with the echo data on the k-spaces Ksys and Kdia having the same matrix size, and the echo data as the difference result may be reconstructed to obtain a blood flow image.

  Further, as a configuration for detecting a signal representing the heartbeat of the subject, for example, a detection signal called PPG (peripheral gating) for detecting a pulse wave of a fingertip with an optical signal is used instead of the above-described ECG signal detection. You may do it.

  Furthermore, the MRI apparatus according to each of the above-described embodiments and modifications thereof is configured to create one image data from two cardiac phase image data. However, according to another aspect of the present invention, this is not necessarily the case. It is not limited to. For example, a readout gradient magnetic field pulse to which a dephasing pulse or a rephasing pulse is added is applied so that it substantially coincides with the flow direction of the fluid (blood, lymph, etc.), and one imaging scan is performed regardless of the cardiac phase. A single image may be obtained by going. In this image, the fluid reflecting the degree of promotion of the flow void effect is reflected in bright or black, so that flow information regarding the fluid can be given.

  Furthermore, means for controlling the intensity of the aforementioned dephasing pulse or rephasing pulse according to the flow velocity of the fluid to be imaged can be provided. This means comprises, for example, the input device 13, the host computer 6, and / or the storage unit 11. When the operator inputs information specifying the imaging region and the fluid from the input device 13, the host computer 6 refers to a table (stores the pulse intensity for each fluid) stored in the storage unit 11 in advance, and according to the reference result. Thus, the intensity of the dephase pulse or rephase pulse may be output to the sequencer 5. In addition, the operator can directly give the pulse intensity via the input device 13.

  The description of the embodiment is as described above, but the present invention is not limited to the configuration described in the embodiment, and those skilled in the art can appropriately change the configuration without departing from the gist described in the claims. These can be modified, and their configurations are also included in the present invention.

DESCRIPTION OF SYMBOLS 1 Magnet 2 Static magnetic field power supply 3 Gradient magnetic field coil unit 4 Gradient magnetic field power supply 5 Sequencer 6 Host computer 7 RF coil 8T Transmitter 8R Receiver 10 Arithmetic unit 11 Storage unit 12 Display 13 Input device 17 ECG sensor 18 ECG unit

Claims (5)

In an MRI apparatus for performing a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field,
A dephasing pulse for dephasing the magnetization spin of blood in the subject is applied at an intensity corresponding to the blood flow velocity in the application direction of a readout gradient magnetic field pulse for reading an echo signal from the subject. A signal collecting means for collecting the echo signal by executing the scan;
An MRI apparatus comprising: an image generation unit configured to generate an arteriovenous separation image of the blood based on the echo signal. In an MRI apparatus for performing a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field,
A dephasing pulse for dephasing the magnetic spin of the lower limb blood is applied at an intensity corresponding to the blood flow velocity in the application direction of the readout gradient magnetic field pulse for reading the echo signal from the lower limb of the subject. A signal collecting means for collecting the echo signal by executing the scan;
An image generating means for generating a blood flow image of the lower limb based on the echo signal ,
The signal collecting means sets a systolic cardiac phase as the first cardiac phase of the subject and sets a diastole cardiac phase as the second cardiac phase of the subject. An operation of executing the first scan in the first cardiac phase and an operation of executing the second scan in the second cardiac phase during imaging of slices or slice encoding amounts An MRI apparatus characterized by repeating the above. In an MRI apparatus for performing a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field,
A dephasing pulse for dephasing the magnetic spin of the lower limb blood is applied at an intensity corresponding to the blood flow velocity in the application direction of the readout gradient magnetic field pulse for reading the echo signal from the lower limb of the subject. A signal collecting means for collecting the echo signal by executing the scan;
Image generating means for generating a blood flow image of the lower limb based on the echo signal;
With
The signal collecting means collects the echo data that fills a part of the k space by the first scan, and collects the echo data that fills the entire k space by the second scan,
The image generation means arranges the echo data of the first scan in a first k space and arranges the echo data of the second scan in a second k space, and the first k space. An MRI apparatus , wherein the echo data of an area that is a part of the second k space and that corresponds to the uncollected area is copied to the upper uncollected area . In an MRI apparatus for performing a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field,
A dephasing pulse for dephasing the magnetization spin of a fluid in the subject in motion is applied at an intensity corresponding to the fluid velocity in the application direction of a readout gradient magnetic field pulse for reading an echo signal from the subject. And a signal collecting means for collecting the echo signals by executing a first scan in the first cardiac phase of the subject and a second scan in the second cardiac phase of the subject, ,
Image generating means for generating an image of the fluid based on the echo signal;
The MRI apparatus, wherein the signal collecting unit executes the second scan continuously with the first scan. In an MRI apparatus for performing a scan based on a pulse sequence including a readout gradient magnetic field pulse on a subject placed in a static magnetic field,
A dephasing pulse for dephasing the magnetization spin of a fluid in the subject in motion is applied at an intensity corresponding to the fluid velocity in the application direction of a readout gradient magnetic field pulse for reading an echo signal from the subject. A signal collecting means for collecting the echo signal by executing the scan as described above,
Image generating means for generating an image of the fluid based on the echo signal;
Means for receiving input of information specifying the intensity of the dephasing pulse or the fluid;
The MRI apparatus characterized in that the signal collecting means controls the intensity of the dephasing pulse in accordance with the intensity of the inputted dephasing pulse or the fluid.

Images