JP3434816B2 - MRI equipment - Google Patents

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to magnetic resonance imaging in which the inside of a subject is imaged based on the magnetic resonance phenomenon of spins (nuclear spins), and an arteriovenous phase image is obtained without using a contrast agent. And an MR (magnetic resonance) imaging method.

[0002]

2. Description of the Related Art In magnetic resonance imaging, nuclear spins of a subject placed in a static magnetic field are magnetically excited by a high-frequency signal of its Larmor frequency, and an image is reconstructed from an MR signal generated by this excitation. Is an imaging method.

In the field of magnetic resonance imaging, in order to obtain a blood flow image in the lung field or abdomen, clinically, MR angiography for administering an angiography by administering a contrast agent to a subject has begun. However, this contrast-enhanced MR angiography method requires invasive treatment because it involves the administration of a contrast agent, and above all, it imposes a heavy mental and physical burden on the patient. Also, the inspection cost is high. further,
In some cases, contrast media cannot be administered depending on the patient's constitution.

On the other hand, as an alternative method to the contrast MR angiography method, time-of-flig
ht: TOF) method, phase contrast (PC)
The law is known.

Of 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 effect of the flow is that the moving spin has 2
Caused by one of two properties. The first is that the spin simply moves its position, and the second is due to the phase shift of transverse magnetization caused by the movement of the spin in the gradient magnetic field. Among these, the former method based on the position movement is the TOF method, and the latter method based on the phase shift is the phase contrast method.

Further, a technique called a DBI (Direct Bolus Imaging) method is disclosed in Japanese Patent Laid-Open No. 2004-242242. An example of this pulse sequence is described, for example, in FIG. That is, after selectively exciting a certain slice (reference plane) perpendicular to the Z-axis direction, during the echo time TE, a dephase pulse and a rephase pulse are sequentially applied in the Z-axis direction to obtain an echo signal. This makes it possible to obtain an image in which the reference plane viewed from the X-axis direction and the blood flow moving in the Z-axis direction from the reference plane between the selective excitation and the echo generation appear. In other words, this DBI method uses the direction of the blood flow in the subject and the read gradient magnetic field pulse (in the above case, Z
A blood flow image is obtained by matching the direction of the gradient magnetic field pulse Gz) in the axial direction.

[0007]

[Patent Document 1] JP-A-6-114034

[0008]

However, in the above-mentioned method, although it depends on the performance of the MRI apparatus, in general, only the blood flow with a velocity of 2 to 3 cm / s can be visualized, Almost no lower velocity flow could be detected. For example, a patient's peripheral vein, lymph vessels, C
The flow of SF (spinal fluid), pancreatic duct, etc. is slow, and the velocity is 1 cm / s or less in most cases. Moreover, because of the influence of displacement due to pulsation and the like, these low-velocity fluid flows have hitherto been undetectable.

In the above method, it is necessary to image a slice perpendicular to the blood flow direction. Therefore, in the case of the two-dimensional slice image, the image does not follow the blood flow.
Therefore, when trying to obtain a three-dimensional image, the number of slices increases, and there is also a problem that the entire imaging time becomes long.

The present invention has been made in order to break the current state of the art as described above, and it is possible to reliably and in a short time a low-velocity flow found in the blood flow of the lower limbs without administering a contrast agent. The purpose is to draw with.

[0011]

In order to achieve the above-mentioned object, an MRI apparatus according to the present invention has one aspect.
In this way, the flow showing the movement inside the subject placed in the static magnetic field
The application direction of the readout gradient magnetic field pulse is essentially the same as the body movement direction
In a state in which the read gradient magnetic field pulse is included.
Scanning based on the loose sequence
Signal collecting means for collecting echo signals from the
Image for generating an image from echo signals collected by the means
Image generation means. The read gradient magnetic field pulse,
The pulse body for reading the echo signal and this pulse
Phase addition of the magnetizing spins of the fluid
Intensity controllable phase behavior control pulse
This phase control pulse is generated by the magnetization spin of the fluid.
Dephase pal to dephase or rephase
Or a rephase pulse.

As an example, the intensity of the phase behavior control pulse may be changed according to the velocity of the fluid flow.

As another example, the signal collecting means
Are different from each other in the same cardiac cycle of the subject.
In one of the two cardiac phases, the phase-encoding of the first k-space
It is placed in the central region that forms the low frequency region in the code direction.
The echo signal for obtaining the echo data
Means for performing a first scan using a sequence
Note that in the other of the two cardiac phases, the phase error of the second k-space
A central region and a high frequency region that form a low frequency region in the
D is arranged at one and of the opposite end portions forming the wave area Kode
Echo signal to obtain the pulse sequence
And a means for performing a second scan, while
The image generating means is obtained by the first and second scans.
The first and second k-spaces depending on the echo signal
From the echo data placed in each of the specified areas
An error to be placed in the remaining area of each of the first and second k spaces
A means for obtaining the code data and placing it in the remaining area, and
2 of the first and second k-spaces arranged by
Means for obtaining the image from a set of echo data,
It

Further, the MRI apparatus according to the present invention is different from that described above.
According to this aspect, the motion inside the subject placed in the static magnetic field is
Application of readout gradient magnetic field pulse in the direction of fluid movement
The read gradient magnetic field pulse
Scan based on a pulse sequence containing
Signal collecting means for collecting echo signals from the subject, and
Image from the echo signals collected by the signal collection means of
And an image generating means for generating. At this time, the reading
A gradient magnetic field pulse for reading the echo signal
A pulse body and the fluid added to the pulse body and
Intensity controllable phase that controls the phase behavior of the magnetized spin
This phase behavior control pulse is formed by the behavior control pulse and
To change the strength of the fluid depending on the velocity of the fluid flow
It is characterized in that it is configured in.

In this configuration, the signal according to another example described above is used.
It is also possible to apply the configuration of the collection means and the image generation means.
Wear.

[0016] For this reason, the application direction of the readout gradient pulse at the time of Imaging scan, so that substantially coincides with the direction of flow of the low flow rates such as those found in the lower extremity vascular fluid (such as blood flow), the pulse sequence by the operator Is preset. Moreover, a dephase pulse or a rephase pulse is added as a phase behavior control pulse to this read gradient magnetic field pulse.

Therefore, the relative signal value difference between the flowing fluid and the fluid flowing at a lower velocity than that can be increased by the phase behavior control pulse.
For example, when using dephasing pulse, even in the blood vessels of the lower limbs whose flow velocity is lower than that of the abdomen and chest, the arteriovenous can be clearly separated and displayed with high visualization ability from such relative signal value difference. You can

As described above, in order to give a relative difference in signal value between arteries and veins, the operator adjusts the application direction of the read gradient magnetic field pulse to the flow direction, and actively dephases or rephases the magnetization spin. It is possible to provide an MR imaging method of controlling the flow void effect by utilizing the above.

Furthermore, since it is not necessary to administer a contrast medium, non-invasive imaging can be performed, and from this point also, the mental and physical load on the patient is significantly reduced.

[0020]

BEST MODE FOR CARRYING OUT THE INVENTION Embodiments of the present invention will be described below.

(First Embodiment) A first embodiment will be described with reference to FIGS.

FIG. 1 shows a schematic configuration of an MRI (magnetic resonance imaging) apparatus according to this embodiment.

(1.1) Configuration of the apparatus This MRI apparatus includes a bed on which the subject P is placed, a static magnetic field generating section for generating a static magnetic field, and a gradient magnetic field generating section for adding position information to the static magnetic field. And a transmission / reception unit that transmits and receives high-frequency signals, a control / calculation unit that controls the entire system and that performs image reconstruction, and an electrocardiographic measurement unit that measures an ECG signal as a signal representing the cardiac phase of the subject P. I have it.

The static magnetic field generator comprises, for example, a superconducting magnet 1 and a static magnetic field power supply 2 for supplying a current to the magnet 1, and has a cylindrical opening (diagnostic space) into which the subject P is loosely inserted. The static magnetic field H ₀ is generated in the axial direction (Z-axis direction).
A shim coil 14 is provided in this magnet portion. A current for homogenizing the static magnetic field is supplied to the shim coil 14 from the shim coil power supply 15 under the control of the host computer described later. In the bed, the top plate on which the subject P is placed can be retractably inserted into the opening of the magnet 1.

The gradient magnetic field generator comprises a gradient magnetic field coil unit 3 incorporated in the magnet 1. The gradient magnetic field coil unit 3 includes three sets (types) for generating gradient magnetic fields in the X-axis direction, the Y-axis direction, and the Z-axis direction that are orthogonal to each other.
X, y, z coils 3x to 3z. The gradient magnetic field unit also includes a gradient magnetic field power supply 4 that supplies current to the x, y, z coils 3x to 3z. This gradient magnetic field power supply 4 is controlled by a sequencer 5 which will be described later, and the x, y, z coils 3x.
A pulse current for generating a gradient magnetic field is supplied to 3z.

Gradient magnetic field power source 4 to x, y, z coils 3x
By controlling the pulse current supplied to ~ 3z,
Gradient magnetic fields in the three-axis (X-axis, Y-axis, and Z-axis) directions, which are physical axes, are combined, and the slice direction gradient magnetic field G _S , the phase encode direction gradient magnetic field G _E , and the read direction (frequency encode direction) that are orthogonal to each other are synthesized. ) The direction of the logical axis composed of the gradient magnetic field G _R can be arbitrarily set and changed. The gradient magnetic fields in the slice direction, the phase encode direction, and the read direction are superposed on the static magnetic field H ₀ .

The transmitting / receiving section is an RF coil 7 arranged near the subject P in the imaging space inside the magnet 1, and this coil 7
And a transmitter 8T and a receiver 8R connected to each other. The transmitter 8T and the receiver 8R operate under the control of the sequencer 5 described later. The transmitter 8T has a nuclear magnetic resonance (N
An RF current pulse of Larmor frequency for exciting (MR) is supplied to the RF coil 7. The receiver 8R takes in MR signals (high frequency signals) such as echo signals received by the RF coil 7 and performs various signal processing such as preamplification, intermediate frequency conversion, phase detection, low frequency amplification, and filtering on the MR signals. 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, and a display 1.
2, an input device 13, and a sound generator 16. Among them, the host computer 6 has a function of commanding the pulse sequence information to the sequencer 5 by a stored software procedure (not shown) and controlling the operation of the entire apparatus.

One of the features of this MRI apparatus is that it performs an MR scan based on an electrocardiographic synchronization method based on the 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 the synchronization timing in advance, and an ECG synchronization based on the synchronization timing. Two imaging scans (hereinafter referred to as imaging scans) that execute the imaging pulse sequence
Is performed during execution of a main program (not shown).
An example of the execution routine of the ECG-prep scan is shown in FIG.
6 and 7 show an example of the first and second imaging scan execution routines based on electrocardiographic synchronization.

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

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.
The form of the pulse train is SE method, high-speed SE
A pulse train of the SE system such as a method, an EPI (Echo Planar Imaging) method, a FASE (Fast Asymmetric SE) method (that is, an imaging method combining a fast SE method and a half Fourier method) is suitable.

The arithmetic unit 10 also inputs the digital data (also called raw data or raw data) output from the receiver 8R through the sequencer 5, and the k space (also called Fourier space or frequency space) by its internal memory. Then, the digital data is placed in the image data, and the data is subjected to a two-dimensional or three-dimensional Fourier transform for each set to reconstruct the image data in the real space. The arithmetic unit is
If necessary, data composition processing for images and difference calculation processing (including weighted difference processing) can also be executed. This combining process includes a process of adding for each pixel, a maximum intensity projection (MIP) process, and the like. Further, as another example of the above-described combining process, the axes of a plurality of frames may be aligned in the Fourier space and the original data may be combined into the original data of one frame. Note that 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 processing and difference processing. The display 12 displays an image. Further, the operator can input the parameter information for selecting the synchronization timing, the scan condition, the pulse sequence, the information about the image synthesis and the calculation of the difference to the host computer 6 through the input device 13.

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

Further, the electrocardiographic measurement unit is an ECG sensor 17 that is attached to the body surface of the subject to detect an ECG signal as an electrical signal, and various processing including digitization processing is performed on this sensor signal to perform host computer processing. 6 and an ECG unit 18 for outputting to the sequencer 5. The measurement signal from the electrocardiogram measurement unit can be used by the sequencer 5 as necessary when executing each of the ECG-prep scan and the electrocardiographically synchronized imaging scan.
Thereby, the synchronization timing of the electrocardiographic synchronization method can be set appropriately, and the electrocardiographic synchronization imaging scan can be performed based on the synchronization timing to collect data.

(1.2) ECG-prep scan Next, the process of determining the optimum synchronization timing by the ECG-prep scan will be described with reference to FIGS.

The host computer 6 executes an 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 executes the ECG-pr
Scan conditions for executing the ep scan and parameter information are read 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 read gradient magnetic field pulse, and the like.
The parameter information includes an initial time T ₀ (here, the elapsed time from the peak value of the R wave in the ECG signal) for determining the synchronization timing (time phase) of the electrocardiographic synchronization, the step width Δt in the time increment, and the number of times. Includes the upper limit value of the counter CNT,
These parameters can be arbitrarily set by the operator.

Next, the host computer 6 counts the number of times the sequence is executed and a count counter CNT and a time increment parameter T for determining the synchronization timing.
_{The inc} is cleared (CNT = 0, _Tinc = 0: step S2). After this, the host computer 6 sends the sound generator 16
Message data is sent to the subject, and the subject (patient) is instructed to perform a breath-holding command such as "Hold your breath" (step S3). This breath-hold is ECG-pr
Although it is preferable to carry out in order to suppress the body movement of the subject during the execution of the ep scan, the ECG-prep scan may be executed in a state where the breath holding is not carried out in some cases.

When the preparation is completed in this way, the host computer 6
Executes the processing of step S4 and thereafter sequentially. As a result, the process shifts to scan execution while changing the synchronization timing of ECG 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 unit 18
The ECG signal subjected to the signal processing in step S1 is read, and it is determined whether or not the peak value of the R wave in the signal appears (step S5). This determination process is repeated until the R wave appears. When the R wave appears (step S5, YES), it is subsequently determined whether or not the delay time T _DL calculated at step S4 has elapsed from the R wave peak time (step S6). This determination process is also continued until the delay time T _DL elapses.

When the delay time T _DL elapses from the peak time of the R wave (step S6, YES), the sequencer 5 is instructed to start each pulse sequence (step S).
7: see FIG. 4). This pulse sequence is preferably set to the same type as the pulse sequence for imaging described later, and for example, 2D-FASE (Fast Asymmetric S / F) which is a combination of the fast SE method and the half Fourier method.
E) Law. Of course, this sequence uses the high-speed SE method,
Various types such as the EPI method can be adopted. In response to this command, the sequencer 5 starts executing the pulse sequence of the type commanded by the operator, so that the region of the desired region of the subject is scanned. The ECG-prep scan may be performed by a two-dimensional (2D) scan when the imaging scan (main scan) for image data acquisition is a three-dimensional (3D) method, or may be performed according to the area of the imaging scan. You may perform by a three-dimensional scan.
In the present 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 shortening the scan time. Considering the mission of the ECG-prep scan, the two-dimensional scan is sufficient.

After the instruction to start the sequence execution, the number counter CNT = CNT + 1 is calculated (step S
8) Furthermore, the time increment parameter T _inc = ΔT ·
CNT calculation is performed (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 instructed, and the incremental parameter T _inc for adjusting the synchronization timing is increased in proportion to the count value.

Next, a predetermined predetermined period (for example, 500 to 1000 m) required for executing each pulse sequence.
It waits as it is until about (sec) (step S10). Further, it is determined whether or not the number counter CNT reaches a predetermined upper limit value (step S11).
In order to optimize the synchronization timing, the delay time T _DL
When five two-dimensional images are to be taken while changing the to various time values, the counter CNT = 5 is set. When the number of times counter CNT = upper limit value has not been reached (step S11, NO), the process returns to the process of step S5 and the above-described process is repeated. On the contrary, when the counter CNT = upper limit value is reached (step S11, YE
S), a breath-hold release command is issued to the voice generator 16 (step S12), and the subsequent processing is returned to the main program. The breath-hold voice message is, for example, "It is okay to breathe."

When the above processes are sequentially executed, the preparation pulse sequence is executed at the timing shown in FIG. 4, for example. For example, the initial time T ₀ = 30
0 msec, when to have been commanded time increment [Delta] T = 100 msec, the delay time _T DL = 300 msec for the first round of the sequence, the second delay time _T DL = 400 msec thereto, the third delay time to it _{T DL} = 500 msec, etc., the delay time T _DL that determines the synchronization timing is adjusted.

Therefore, when the first R wave after the breath-holding command reaches the peak value, the delay time T is reached from the arrival time.
_{After DL} (= T ₀ ), for example, the first scan IMG _prep1 based on the two-dimensional FASE method is _performed for a predetermined time (50
(0 to 1000 msec), and echo signals are collected. Even if the next R wave appears during the continuation of this sequence, the sequence is continued without being involved in the appearance of this R wave because of the waiting process of step S10 in FIG. That is, the execution process of the sequence started in synchronization with a certain heartbeat is continued over the next heartbeat, and the echo signal is collected.

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 elapses from the arrival time. IMG _prep2 continues for a predetermined time, and echo signals are similarly acquired. After this scan ends, the next R wave appears and the delay time T _DL = T ₀
When + 2 · T _inc = 500 msec has elapsed, the third scan IMG _prep3 continues for a predetermined time, and echo signals are similarly collected. Further, after this scan ends, the next R wave appears, and the delay time T _DL = T ₀ + 3 ·
When T _inc = 600 msec has elapsed, the fourth scan IMG _prep4 continues for a predetermined time, and echo signals are similarly acquired. This scan continues for a desired number of times, for example, 5 times, and a total of 5 frames (sheets) of echo data of the same cross section are 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 image data in real space by the two-dimensional Fourier transform method. This image data is stored in the storage unit 11 as blood flow image data. The host computer 6 responds to an operation signal from the input device 13, for example, and sequentially displays the blood flow images in 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, arteries AR and veins VE that flow in the lower limbs almost vertically are located. However, "delay time T _DL =
The initial time T ₀ + Tinc · Δt ”differs for each image. The operator visually observes these images to see the arterial AR and vein V
An image in which E appears in the highest signal and an image in which only veins appear in the highest signal are selected. Among these, due to the delay time T _DL1 corresponding to the image in which only the vein VE appears in a relatively high signal, the systolic synchronization timing T _DL = T
_DL1 is determined. Further, the delay time T corresponding to the image in which the artery AR and the vein VE appear in a relatively high signal.
_{By DL2} , diastolic synchronization timing T _DL = T
_DL2 is decided.

Therefore, the operator visually observes a plurality of blood flow images imaged by dynamically changing the delay time T _DL in this way, and selects two optimal cardiac phases in systole and diastole, respectively. Delay time T _DL = T _DL1 , T
_DL2 (synchronization timing) is determined, and this delay time T _DL
For example, the process of reflecting the information in the imaging scan that is continuously performed is manually performed.

When an image determined by visual observation is designated, the delay time T _DL given to the designated image is automatically stored as the optimum synchronization timing, and this timing T _DL is automatically read during the imaging scan. You may build and install the software like this. This enables automatic designation processing of the ECG synchronization timing.

(1.3) Imaging Scan Next, the operation of two imaging scans (that is, two imagings) in this embodiment will be described with reference to FIGS.

The host computer 6 executes a predetermined main program (not shown) and, as a part thereof, executes the processing of each imaging scan shown in FIG. 6 in response to the operation information from the input device 13.

Now, it is assumed that the first imaging scan (imaging) is assigned to systole. In this case, the host computer 6 first uses the ECG-pr described above.
Optimal delay time T _DL (= T _DL1 or T _DL2 > T _DL1 ) for systole determined by operator through ep scan
Is input, for example, via the input device 13 (step S2
0).

Next, the host computer 6 scan conditions specified by the operator from the input device 13 (application direction of read gradient magnetic field pulse, image size, number of scans, waiting time between scans, pulse sequence according to scan region, etc.). And information about the image processing method (MIP processing, difference processing, etc., in the case of difference processing, any one of simple difference, weighted difference processing, addition processing, etc.) is input, and the delay time T
The information including _DL is processed into control data, and the control data is output to the sequencer 5 and the arithmetic unit 10 (step S21).

Next, if it can be judged that the preparation completion notice before the scan has been given (step S22), step S2
In step 3, a command to start breath holding is output to the voice generator 14 (step S23). As a result, the voice generator 14 is
"Hold your breath" as with G-prep scan
The patient hears this and stops breathing because the voice message with such contents is issued (see FIG. 9).

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

The sequencer 5 receives the command to start the imaging scan (FIG. 7, step S24--
1) Start reading the ECG signal (step S24)
-2), E in which the predetermined n-th appearance of the peak value of the R wave (reference waveform) in the ECG signal is synchronized with the peak value
It is judged from the CG trigger signal (step S24-3).
Here, waiting for the appearance of the R wave n times (for example, twice) is for the purpose of predicting the time when the breath wave is surely shifted. When the predetermined n-th R wave appears, the set delay time T _{D L1}
The process of waiting only is performed (step S24-4).

This optimum delay time T _DL1 (or T
_The sequencer 5 executes the first imaging scan, assuming that the time when _DL2 ) has elapsed is the optimal electrocardiographic synchronization timing (step S24-5). Specifically, the transmitter 8T and the gradient magnetic field power source 4 are driven according to the already stored pulse sequence information, and the first imaging scan (imaging) based on the pulse sequence of the three-dimensional FASE method is performed as shown in FIG. As shown in (a) and (c), it is executed in synchronism with the electrocardiogram (the gradient magnetic field in the phase encoding direction is not shown in FIG.

According to this pulse sequence, the application direction RO of the readout gradient magnetic field pulse G _R is almost in the direction of the blood flow (artery AR, vein VE) for imaging as shown in FIG. 10 by the operator. It is preset to match.

[0061] Further, readout gradient pulses _{G R} contained in the pulse sequence Figure 8 (c) and FIG. 9 (a) ~
As shown in (c), a frequency encoding pulse body P _body for collecting an echo signal and this pulse body P body
It is composed of two _dephase pulses P _dephase as a control pulse continuously added before and after the _body . This dephase pulse P
_dephase is a pulse body P for frequency encoding.
It has the same polarity as the _body, and thus has a function of promoting the dephasing of the moving magnetized spins.

The dephase pulse P
_{The dephase} hardly exerts a dephase function on the magnetized spin that is hardly moved. Therefore, the read gradient magnetic field pulse G _R is set by the operator so as to be applied in substantially coincidence with the direction of movement of the fluid (blood or lymph) for imaging.

Preferably, the dephase pulse P
_The intensity of the _dephase can be changed or controlled independently of the pulse body P _body according to the velocity of the lymph or the blood flow as the fluid to be imaged. FIG. 9 (a)
(C) shows the dephase pulse P in this order.
An example of reducing the strength of the _dephase is illustrated. Generally, as the blood flow velocity increases, the intensity of the _dephase pulse P _dephase is changed or controlled so as to decrease.

When the velocity of the fluid to be imaged (such as blood flow) is relatively high, as shown in FIG. 9 (d), a total of control pulses are continuously generated before and after the pulse body P _body. Two rephase pulses P
_rephase is added. This rephase pulse P
_rephase is the pulse body P for frequency encoding.
_The polarity is opposite to that of the _body , and it has a function of rephasing the magnetization spins to suppress artifacts in order to suppress excessive dephase. The intensity of the _rephase pulse P _rephase is also preferably changed or controlled independently of the pulse body P _body according to the flow velocity.

[0065] In the first embodiment, therefore, both in the imaging scan of the first time and the second time (to be described later), dephasing pulses to the readout gradient pulse _{G R P dephase} or rephase P
_rephase is added.

Therefore, by executing the above-mentioned pulse sequence of the three-dimensional FASE method, the echo signals urged by the excitation 90 ° RF pulse and the refocus 180 ° RF pulse are generated for each slice encode and each phase encode. To be collected. This echo signal reflects the _dephasing action of the phase of the magnetized spin due to the _dephase pulse P _dephase or the rephasing action of the phase of the magnetized spin _due to the _rephase P _rephase .

This will be described in detail together with the display operation to be described later, but the outline thereof is as follows.

That is, for the fluid flowing along the application direction of the read gradient magnetic field pulse, the dephasing effect due to the dephasing pulse P _dephase leads to promotion of the flow void effect. Therefore, the intensity of the echo signal is reduced by the dephase pulse. On the contrary, in the case of a fluid that hardly flows in that direction, the degree of promotion of the flow void effect due to the dephasing pulse P _dephase is low, and the intensity of the echo signal does not decrease so much.

In the case of the _rephase pulse P _rephase , the rephasing action suppresses the effect of dephasing in accordance with the flow of fluid.

The echo interval in the above pulse sequence is shortened to about 5 msec. As a result, about 600 msec based on the first slice encoding amount SE1
Echo signals are collected from the three-dimensional imaging region Rima set in the lower limb as shown in FIG.

When the scan based on the first slice encoding is completed, the sequencer 5 determines whether or not the scan of the final slice encoding is completed (step S24-6), and this determination is NO (final slice encoding is determined). ECG)
While monitoring the signal, for example, the R wave used in the previous imaging scan is awaited until a short set period e.g. 2 heartbeats (2R-R) elapses (step S
24-7). The repetition time TR is 4 heartbeats (4R-
R) Set below.

In this way, after waiting for a period corresponding to, for example, 2R-R, for example, when the third R wave appears (step S24-7, YES), the sequencer 5 performs the processing in step S24-4 described above. return. As a result, when the designated delay time T _DL1 has elapsed from the ECG trigger signal synchronized with the third R wave peak value, the scan based on the next slice encode amount SE2 is executed in the same manner as described above, and the three-dimensional imaging region Echo signals are collected from R _ima (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 encode amount SEn is completed, the determination in step S24-6 becomes YES, and the sequencer 5 outputs a notice of completion of the first (or second) imaging scan to the host computer 6. (Step S24-8). This allows
The processing is returned to the host computer 6.

Upon receipt of the scan completion notice from the sequencer 5 (step S25), the host computer 6 outputs a breath-hold cancellation command to the voice generator 16 (step S2).
6). Then, the voice generator 16 sends a voice message to the patient, such as "Breathe fine".
The breath-hold period is over (see Figure 8).

As a result, the first (or second) imaging scan (imaging) by electrocardiographic synchronization is performed every 2R-R based on, for example, the 3D-FASE method.

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

Then, as shown in FIG. 2, the second imaging scan (imaging) for the diastole is performed at an appropriate time in the same manner as the first imaging scan. However, in the case of the second time, the optimum delay time T _DL2 that determines the predetermined time phase of the diastole set in advance through the ECG-prep scan described above is read (FIG. 6, steps S20 and S2).
1), electrocardiographic synchronization based on this delay time T _DL2 is established (FIG. 7, step S24-4).

Therefore, in the case of the second imaging scan, as shown in FIGS. 8 (b) and 8 (c), each phase encoding amount is obtained at the diastolic synchronization timing delayed by the delay time T _DL2 from the R wave peak. 3D FASE based on SE
A legal scan is performed. Also in this case, the application direction of the read gradient magnetic field pulse G _R is set so as to almost coincide with the direction of movement 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 pulses P
_rephase ) is added. This control pulse corresponds to the phase behavior control pulse of the present invention.

Therefore, by the second imaging scan, the read gradient magnetic field pulse G is read as in the first scanning.
Image data in diastole reflecting the spin control function of the _dephase pulse P _dephase or the _rephase pulse P _rephase added to _R can be 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, the arithmetic unit 6 responds to a command from the host computer 6 to calculate echo data based on the half Fourier method in both the systolic k space and the diastolic k space ( 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 by the complex conjugate relationship and arranged. This fills both k-spaces with echo data.

After that, the arithmetic unit 10 determines k for systole.
Image reconstruction is performed by performing three-dimensional Fourier transform on the echo data in the space and the echo data in the k-space for diastole (step S32,
S33). As a result, as shown in FIGS. 12A and 12B, an image (systolic image) IM _sys of the time phase of the delay time T _{DL1 in} the systole and the delay time T in the diastole.
Three-dimensional data of a temporal image (diastolic image) _IMdia of _DL2 is obtained.

According to this image data, the systolic image IM
Only the vein VE is reflected in the _sys , and the artery AR is hardly reflected. Meanwhile, diastolic image IM
_The arteries AR and veins VE are reflected in the _dia to some extent.

Here, such a systolic image IM _sys.
And the principle of 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 the magnetization spin of the component such as the blood flow flowing in the application direction of the read gradient magnetic field pulse is more likely to disperse due to the dephase pulse. In other words, for the flowing component, it is equivalent to the promotion of the flow void effect due to the flowing component itself. On the contrary, due to the rephasing pulse, the rephasing function acts on the phase of the magnetized spin such as the blood flow.

For example, the lower limb of the subject is taken as an example. In the case of the lower limbs, even the arteries in the systole usually have a low flow rate of 1 cm / s or less, and the veins in the systole and the arteries and veins in the diastole have a low flow rate that can be considered to be 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 magnetized spins and collect echo signals. At this time, since the flow velocities of the arteries and veins 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 veins flow very slowly, the echo signal is slightly reduced by the dephasing pulse, but the flow void effect is small, and it is visualized in bright blood with a relatively high signal value. On the other hand, since the systolic artery flows at a higher flow rate than the vein, the degree of promotion of the flow void effect due to the dephasing pulse is larger than that of the vein. As a result, the decrease in the signal value of the artery is large, and the black blood (black)
It is depicted in the blood). This state is shown in FIG.
It is schematically represented in (a). In the figure, the hatched portion is bright blood and the dotted portion is black blood.

On the other hand, in the diastole, both the artery and the vein move at an extremely low flow rate, so that both the artery and the vein are visualized in bright blood, although there is a slight decrease in the signal value due to the dephase pulse. . This state is schematically shown in FIG. 12 (b) described above.

Returning to the explanation of FIG. 11, the arithmetic unit 10
To obtain the arterial phase image IM _AR
The difference calculation “ _IMdia− β · _IMsys ” is performed for each pixel for the _sys and the diastolic image _IMdia (step S34). Here, β is a weighting coefficient. As a result, by appropriately setting the weighting coefficient β as shown in FIG. 12, the image data of the vein VE becomes almost zero,
Three-dimensional image data of the arterial phase image IM _AR showing only the artery AR is obtained.

Further, to obtain the vein phase image IM _VE ,
The difference calculates the _"IM dia _{-IM AR"} performed for each pixel (step S35). The image data IM _AR is image data calculated by the above-mentioned 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 by the weighted difference.

When the difference calculation is completed in this way, the calculation unit 10 performs MIP (maximum intensity projection) processing on both the arterial phase image IM _AR and the venous phase image IM _VE to obtain the blood vessels from the desired direction. Data of a two-dimensional image (for example, a coronal image) at the time of observation is created (step S36).

Two-dimensional image IM of this arterial phase and venous phase
_AR and IM _VE are, for example, as shown in FIG.
2 and the image data are stored in the storage unit 11 (step S39).

At the time of this display, the arterial phase image IM
_In addition to _AR and venous phase image IM _VE , systolic image IM
_{The sys} and the diastolic image _IMdia may be displayed on the same screen or on the screen of a separate monitor.

[0095] (1.5) As described effect above, the MRI apparatus of this embodiment, during an imaging scan, a readout gradient pulses G _R of the low flow rates such as those found in the lower limbs vascular fluid (such as blood flow) The voltage is applied almost in line with the flow direction. Moreover, at the same time,
Gradient magnetic field pulse G _R and dephase pulse P
_dephase or _rephase pulse P _rephase
Is added.

As a result, the relative signal value difference between the flowing fluid and the fluid flowing at a lower velocity than that can be increased by the _dephase pulse P _dephase or the _rephase pulse P _rephase. . So, for example, when using dephasing pulse,
Even in a blood vessel of the lower limb whose flow velocity is lower than that of the abdomen or chest, the arteriovenous can be clearly separated and displayed with high visualization ability from the relative signal value difference as shown in FIG.

As described above, in order to give a relative difference between signal values between arteries and veins, the application direction of the read gradient magnetic field pulse is adjusted to the flow direction, and the dephasing or rephasing of the magnetization spin is positively performed. The method of controlling the flow void effect by utilizing the above is a novel method first developed by the present inventor.

In this embodiment, the ECG-prep is used.
Optimal EC for systole and diastole by scanning
Since the G synchronization timing is set in advance, it is possible to reliably capture the target blood flow in each time phase of systole and diastole. Furthermore, since the optimal setting of the electrocardiographic synchronization timing in advance eliminates the need for re-imaging, the operational burden on the operator and the physical and mental burden on the patient are reduced.

Further, since the slice direction or slice encode direction can be set to a direction other than the vertical direction of the patient, the whole image can be compared with the method of imaging in the vertical direction perpendicular to the blood flow such as the TOF method. Scan time is short. This reduces the burden on the patient and increases the patient throughput.

Furthermore, since it is not necessary to administer a contrast medium, non-invasive imaging can be performed, and from this point also, the mental and physical load on the patient is significantly reduced. At the same time, it is possible to eliminate the inconvenience peculiar to the contrast method such as the timing of the contrast effect, and unlike the contrast method, it is possible to perform repeated imaging as necessary.

(Modification of First Embodiment) In the above-described embodiment, as shown in FIG.
In time th and the second imaging scan both dephasing pulses P to the readout gradient pulse G _R
dephase or _rephase pulse P _rephase
Has been described.

As a modification of this embodiment, the first imaging scan performed in the diastole time phase is shown in FIG.
As shown in (a), the _dephase pulse P _dephase
On the other hand, on the other hand, in the second imaging scan performed in the systolic time phase, the rephase pulse P shown in FIG.
_{A rephase} may be added.

That is, according to the systole and the diastole,
The type of control pulse that additionally controls the behavior of the magnetization spin is changed. As a result, the effect of rephase (flow compensation) can be reflected in the diastole to increase the signal value and improve the S / N.

(Second Embodiment) Next, a second embodiment according to the present invention will be described with reference to FIGS. The MRI used in this embodiment
The hardware configuration of the device 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 performed in the first embodiment are executed in one imaging scan, and the systole of the cardiac cycle and The present invention relates to a configuration in which the dephase pulse and the rephase pulse described above are selectively used according to the diastole.

Now, it is assumed that a separated image of the arteries and veins of the lower limb is obtained as the low-velocity 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. E
The CG-prep scan is performed by the method described in the first embodiment, whereby the delay times T _DL1 and T from the R wave that provide the most imaging ability in the systole and the diastole are provided.
_DL2 is set respectively.

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

(2.1) Imaging Scan The host computer 6 is executing a predetermined main program (not shown), and as a part thereof, responds to the operation information from the input device 13 and is shown in FIGS. Execute the process.

More specifically, the host computer 6 firstly determines the optimum two delay times T _DL determined by the operator through the above-mentioned ECG-prep scan (that is, the optimum delay time T _{DL1 in} systole and the expansion). Optimal delay time T
_DL2 (> T _DL1 ) is input, for example, via the input device 13 (step S120). This optimum delay time T
Information of _DL1 and T _DL2 is stored in advance, for example, in the storage unit 1.
It may be stored in 1.

Next, the host computer 6 inputs the information of the scan condition 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 processed by the sequencer 5 and the arithmetic unit. It is output to 10 as needed (step S121).

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

The sequencer 5 receives the command to start the imaging scan (FIG. 18: step S124-
1) Start reading the ECG signal (step S12)
4-2), the appearance of the peak value of the R wave (reference waveform) in the ECG signal for the predetermined nth time is determined from the ECG trigger signal synchronized with the peak value (step S124-).
3).

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

The sequencer 5 executes the scan for the systole, assuming that the time when the optimum delay time T _DL1 has passed is the optimum ECG synchronization timing (step S).
124-5).

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

In this first scan SN _sys1 , the read gradient magnetic field pulse GR is applied in the body axis direction substantially along the arteries and veins of the lower limb. 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. 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 to be short as shown in FIG. 19 so that it ends within a short time within one heartbeat from the start of scanning. ing. As schematically shown in FIG. 20, the number of echoes is set to be sufficient to collect the echo data arranged only in the central portion (low frequency region) in the phase encode ke direction of the k space for each slice encoding amount. . Therefore, 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. Also, the echo data lacking in the systolic k space (first k space) K _sys is the expansion k space (second k space) K described later.
_It is obtained by copying from _dia and the 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 limb in a short scan time of about several hundred msec based on the first slice encoding amount SE1.

Next, the sequencer 5 shifts to scan control in the diastole. Specifically, a process of waiting for the delay time T _DL2 set for the specific time phase of the diastole is performed (step S124-6).

The sequencer 5 executes the second scan for the diastole, assuming that the time when the optimum delay time T _DL2 has passed 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 the second slice encoding amount SE1 based on the pulse sequence of the three-dimensional FASE method is used for the second time, for example. Scan SN _dia1 is executed by the ECG _gated method as shown in FIG.

Also in this second scan SN _dia1 , the read gradient magnetic field pulse GR is applied in the body axis direction substantially along the arteriovenous of the lower limb. Further, a _rephase pulse P _rephase for rephasing the magnetization spins is added to the read gradient magnetic field pulse G _R in succession before and after the time, as shown in the figure. Also, the echo interval in this pulse sequence is shortened to about 5 msec.

As shown in FIG. 19, the pulse sequence used for the second scan SN _dian in the diastole is larger than that in the systole, but the half-Fourier method is used in combination, so the echo filling the entire k-space is increased. It is set up to collect fewer than the number of echoes. As the number of echoes, as schematically shown in FIG. 20, echo data to be arranged only in the central portion (low frequency region) in the phase encode ke direction of the k space and one end portion (high frequency) thereof is collected for each slice encoding amount. It is set to be sufficient.
In the k-space _Kdia for diastole, the insufficient echo data is calculated by the half Fourier method as will be described later. Scan SN in this diastole
_As shown in FIGS. 19 and 20, _dia1 is normally scanned over the next heartbeat.

As a result, based on the first slice encode amount SE1, the scan time is about 600 msec.
Three-dimensional imaging area Rima set on the lower limb (see FIG. 10)
An echo signal is collected from.

When these first imaging scans are completed, the sequencer 5 judges whether or not the final scan is completed (step S124-8), and if this judgment is NO (the final scan is not completed). , EC
While monitoring the G signal, for example, two heartbeats (2R-R) are waited for, for example, two heartbeats (2R-R) from the R wave used for the imaging scan, and the longitudinal magnetization of spins of the stationary parenchyma is stopped. Recovery is positively suppressed (step S124-9).

In this way, after waiting for a period corresponding to, for example, 2R-R, for example, when the third R wave appears from the start of scanning (step S124-9, YES), the sequencer 5 proceeds to step S124-4 described above. The process is returned.

As a result, when the delay time T _DL1 has elapsed from the third R wave peak value, the first scan SN _sys2 for the second systole according to the next slice encoding amount SE2 is the same as that described above. Then, echo signals are collected from the three-dimensional imaging region R _ima (steps S124-4,5). 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 the three-dimensional imaging is performed. Echo signals are collected from the region R _ima (steps S124-6, 7).

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

When the final scans SN _sys and SN _dian with the slice encoding amount SEn are completed, the determination in step S124-8 becomes YES, and the sequencer 5 outputs a notice of the completion of the imaging scan 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 notice from the sequencer 5 (FIG. 17: step S12).
5) Then, a command to release the breath hold is output to the voice generator 16 (step S126).

Therefore, as schematically shown in FIG.
In one imaging scan (imaging), for example, 2
For each R-R, an electrocardiographic synchronized scan for systole and diastole is executed for n slice encode amounts by, for example, the 3D-FASE method.

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

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

As shown in FIG. 21, the arithmetic unit 6 responds to a command from the host computer 6 in response to the systolic k space K.
All the data arrangements in the _sys and diastolic k space _Kdia are completed (steps S131 and S132). Specifically, in step S131, as shown in FIG. 20, echo data of one high-frequency region Re in the phase encoding direction in the diastolic k space K _dia (number h in FIG. 20).
Echoes up to n) are copied to corresponding positions in the systolic k-space K _sys . This echo data is data of a region that was not acquired by the systolic scan. Then, the process proceeds to step S132, and the systolic k
The half Fourier method is individually applied to both the space K _sys and the diastolic k space K _dia , and the data of the remaining region where the echo data is not collected is calculated by the complex conjugate relation and arranged. Therefore, step S13
Through the processing of 1 and S132, both k spaces K _sys and K _dia are completely filled with data.

Thereafter, the arithmetic unit 10 determines the systolic k
Image reconstruction by three-dimensional Fourier transform is performed on each of the space K _sys and the diastolic k space K _dia (step S1).
33, S134). As a result, as shown in FIG.
As shown in (b), three-dimensional data of an image (systolic image) IM _sys of the delay time T _{DL1 in} the systole and an image (diastolic image) IM _dia of the delay time T _{DL2 in} the diastole are 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, in the diastolic image _IMdia , both the artery AR and the vein VE are reflected to some extent.

Therefore, the arithmetic unit 10 performs the difference calculation "IM" as in the first embodiment (see FIGS. 12 to 14).
generation of the arterial phase image IM _AR by “ _dia −β · IM _sys ” (step S135), difference calculation “IM _dia − ·
Generation of Vein Phase Image IM _VE by "IM _AR " (Step S136), Arterial Phase Image IM _AR and Vein Phase Image IM
MIP (maximum intensity projection) processing of each _VE (step S13
7), the two-dimensional image display of the arterial phase and the venous phase, and the storage of the image data (step S138) are sequentially executed.

(2.3) Functions and Effects As described above, in the MRI apparatus of this embodiment, the readout gradient magnetic field pulse G _R is applied almost at the same time as the blood flow direction of the lower limb during the imaging scan, and
A _dephase pulse P _dephase is added to the gradient magnetic field pulse G _R applied during _systole, and a _rephase pulse P _rephase is added thereto during _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 the effect of promoting the flow void to the artery flowing in the systole, particularly the artery. In the diastolic arteries and arterial flow, the effect of flow compensation can be positively applied.

Therefore, the relative signal value difference between the flowing blood flow and the blood flow flowing only at a lower velocity than that is made remarkable by the dephase pulse and the rephase pulse, and the flow velocity is increased in the abdomen or Even in the blood vessels of the lower limbs lower than the chest, the arteries and veins can be clearly separated based on the relative signal value difference and can be displayed with high visualization ability.

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

Therefore, normally, the two-shot scan for systole and diastole for one slice encoding is within about two heartbeats. Therefore, by repeating these scans alternately in sequence, echo data of blood flow in systole and diastole is collected within one breath-holding continuous period in one imaging. That is, three-dimensional blood flow data during systole and diastole are collected separately and at the optimum timing in one imaging.

Therefore, it is not necessary to separately perform the imaging scan for the systole and the diastole (that is, the imaging is performed twice in total), and the imaging can be performed once. Therefore, imaging time is significantly reduced and patient throughput is increased. In particular, the effect of shortening the imaging time becomes remarkable in the three-dimensional imaging. In addition, since the misregistration due to the body movement of the patient can be significantly reduced, the quality of the presented image is improved. Furthermore, since a blood flow image (MRA image) in which the arterial phase and the venous phase are separated can be obtained from the echo data of the two time phases collected by one imaging,
The imaging efficiency is good, and the blood flow information provided is rich.

Further, according to this second embodiment, the first
Other operational effects obtained in the above embodiment can be similarly enjoyed.

(Modification of Second Embodiment) In the first and second imaging scans of the above-described second embodiment, as shown in FIG. 19, the readout gradient magnetic field pulse for systole is used. A dephasing pulse was added to the pulse, and a rephase pulse was added to the readout gradient magnetic field pulse for diastole. On the other hand, only the dephasing pulse may be added to the readout gradient magnetic field pulse for systole and diastole. As a result, similarly to the case of the first embodiment (see FIG. 8), the degree of promotion of the flow void effect due to the blood flow velocity that differs for each time phase can be reflected in the strength of the signal value. Can be reliably separated.

(Modified Examples Common to the Embodiments) The present invention is not limited to the structures described in the above-described embodiments, and various modified structures and applications are possible.

For example, in the above-described embodiment, both the arterial phase image and the venous phase image are presented, but for this, only the arterial phase image may be subjected to difference calculation and displayed. That is, step S136 in FIG.
It is possible to omit the difference calculation for the vein phase image in. On the contrary, although the difference calculation of the images of the arterial phase and the venous phase is performed together, the image to be displayed may be only the arterial motion image.

Further, in each of the above-mentioned embodiments, the scan method in which the half Fourier method is applied to each of the systolic scan and the diastolic scan is adopted, but the half Fourier method is not necessarily adopted. You don't have to. In this case, it is advisable to fully collect data in the k space by the diastolic scan and copy the echo data of the high frequency regions at both ends in the slice encoding direction to the corresponding regions of the systolic k space.

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

Further, in the post-processing of the echo data of the above-mentioned embodiment, the echo data is once converted into the image data of the real space, and thereafter the difference calculation is performed to obtain the images of the arterial phase and the venous phase. Although configured, such difference calculation is performed in the k spaces Ksys and Kdia having the same matrix size.
The blood flow image may be obtained by performing the echo data as it is and reconstructing the echo data as the difference result.

Further, as the configuration for detecting the signal indicating the heartbeat of the subject, instead of the above-mentioned one for detecting the ECG signal, for example, a PPG for detecting the pulse wave of the fingertip by an optical signal.
A detection signal called (peripheral gating) may be used.

Furthermore, the MRI apparatus according to each of the above-described embodiments and its modification has one image data of two cardiac phases.
Although the configuration is such that one image data is created, the present invention is not necessarily limited to this according to another aspect of the present invention. For example, a readout gradient magnetic field pulse added with a dephasing pulse or a rephasing pulse is applied so as to approximately match the flow direction of a fluid (blood, lymph, etc.), and one imaging scan is performed regardless of the cardiac phase. You may go and obtain a single image. In this image, since the fluid reflecting the degree of promotion of the flow void effect is reflected in bright or black, it is possible to provide flow information regarding the fluid.

Further, it is possible to provide means for controlling the intensity of the dephase pulse and the rephase pulse described above according to the flow velocity of the fluid to be imaged. 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 the table (stores the pulse intensity for each fluid) stored in the storage unit 11 in advance, and responds to the reference result. The intensity of the dephase pulse or the rephase pulse may be output to the sequencer 5. Alternatively, the operator can directly apply the pulse intensity via the input device 13.

The embodiments have been described above, but the present invention is not limited to the configurations described in the embodiments.
Those skilled in the art can appropriately change and modify the present invention without departing from the scope of the claims, and the configurations thereof are also included in the present invention.

[0153]

As described above, according to the MRI apparatus of the present invention, the application direction of the read gradient magnetic field pulse included in the pulse sequence is made to substantially coincide with the fluid flow direction .
At the same time, the phase behavior control added to the readout gradient magnetic field pulse
Control pulse with dephase pulse or rephase pulse
Forming and controlling the phase behavior of the pulse intensity of the fluid
Since the flow rate can be changed according to the degree, it is possible to reliably visualize the low-velocity flow seen in the blood flow of the lower limbs and the like without administering a contrast agent. As an example, it is possible to render an image in which the arteries and veins of the lower limbs are separated in high quality in a short time.

[Brief description of drawings]

FIG. 1 is a functional block diagram showing a configuration example of an MRI apparatus according to an embodiment of the present invention.

FIG. 2 is a diagram illustrating a time series relationship between an ECG-prep scan and two imaging scans according to the first embodiment.

FIG. 3 is a schematic flowchart illustrating the procedure of an ECG-prep scan.

FIG. 4 is a timing chart exemplifying a time series relationship with an ECG signal in an ECG-prep scan.

FIG. 5 is a schematic image diagram obtained by ECG-prep scanning when the delay time is dynamically changed.

FIG. 6 is a schematic flowchart showing an example of first and second imaging scans.

FIG. 7 is a schematic flowchart showing an example of first and second imaging scans.

FIG. 8 is a timing chart exemplifying timings of first and second imaging scans based on an electrocardiographic synchronization method according to the first embodiment.

FIG. 9 is a diagram illustrating a dephase pulse and a rephase pulse applied to a read gradient magnetic field.

FIG. 10 is a diagram illustrating a positional relationship between a three-dimensional imaged region and a blood vessel to be imaged.

FIG. 11 is a schematic flowchart illustrating a process of calculating and displaying echo data according to the first embodiment.

FIG. 12 is a schematic diagram illustrating an outline of difference calculation for obtaining an arterial phase image.

FIG. 13 is a schematic diagram illustrating an outline of difference calculation for obtaining a vein phase image.

FIG. 14 is a diagram exemplifying a simultaneous display state of an arterial phase image and a venous phase image.

FIG. 15 is a pulse sequence of two imaging scans performed in the modified example of the first embodiment.

FIG. 16 is an ECG-pr according to a second embodiment of the present invention.
The figure explaining the time series relationship of an ep scan and one imaging scan.

FIG. 17 is a schematic flowchart showing an example of an imaging scan.

FIG. 18 is a schematic flowchart showing an example of an imaging scan.

FIG. 19 is a timing chart illustrating the timing of an imaging scan based on an electrocardiographic synchronization method according to the first embodiment.

FIG. 20 is a diagram illustrating a state in which echo data collected during systole and diastole are arranged in k space.

FIG. 21 is a schematic flowchart illustrating a process of calculating and displaying echo data according to the second embodiment.

[Explanation 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 units 11 Memory unit 12 Display 13 input device 17 ECG sensor 18 ECG unit

Claims (5)

(57) [Claims] 1. Movement inside a subject placed in a static magnetic field
To apply readout gradient magnetic field pulse in the direction of fluid movement
The readout gradient magnetic field pulse with the orientation substantially aligned.
A scan based on a pulse sequence including
A signal acquisition means for collecting echo signals from the subject, the image from the echo signals collected by the signal collection means
And a pulse body for reading out the echo signal, and an intensity changeable for controlling the phase behavior of the magnetization spin of the fluid, which is added to the pulse body. And a phase behavior control pulse, the phase control pulse being formed by a dephase pulse or a rephase pulse for dephasing or rephasing the magnetized spins of the fluid. 2. The MR apparatus according to claim 1, wherein the intensity of the phase behavior control pulse is changed according to the velocity of the fluid flow.
I device. 3. The MRI apparatus according to claim 1 or 2.
In addition, the signal collecting means may be arranged in a central region forming a low frequency region in a phase encoding direction of the first k-space in one of two mutually different cardiac time phases in the same cardiac cycle of the subject. Means for performing a first scan of the echo signal for obtaining the echo data to be arranged using the pulse sequence, and in the other of the two cardiac phases, in the phase encoding direction of the second k-space And a means for performing a second scan using the pulse sequence for an echo signal for obtaining echo data to be arranged in a center region forming a low frequency region and one of both ends forming a high frequency region, The image generation means extracts the echo data arranged in the designated area in each of the first and second k spaces according to the echo signals obtained in the first and second scans. Means for obtaining echo data to be arranged in the remaining areas of the first and second k spaces and arranging the echo data in the remaining areas, and two sets of echoes of the first and second k spaces arranged by this means Means for obtaining the image from data, and an MRI apparatus. 4. Movement inside a subject placed in a static magnetic field
To apply readout gradient magnetic field pulse in the direction of fluid movement
The readout gradient magnetic field pulse with the orientation substantially aligned.
A scan based on a pulse sequence including
A signal acquisition means for collecting echo signals from the subject, the image from the echo signals collected by the signal collection means
And a pulse body for reading out the echo signal, and an intensity changeable for controlling the phase behavior of the magnetization spin of the fluid, which is added to the pulse body. An MRI apparatus characterized in that it is formed by a phase behavior control pulse and the intensity of this phase behavior control pulse is changed according to the velocity of the fluid flow. 5. The MRI apparatus according to claim 4, wherein the signal acquisition unit is provided with two different signals in the same cardiac cycle of the subject.
The phase encoding of the first k-space in one of the cardiac phases
It is placed in the central region that forms the low frequency region in the
The echo signal for obtaining co-data is converted to the pulse sequence
Means for performing a first scan and a second k-space phase in the other of the two cardiac phases.
A central region and a high region that form a low frequency region in the encoding direction.
Echo data to be placed on one of the two ends of the frequency domain
The echo signal for obtaining the
Means for performing a second scan using the image generation means, and the image generation means specifies the first and second k-spaces in accordance with echo signals obtained in the first and second scans. Means for obtaining the echo data to be arranged in the remaining areas of the first and second k spaces from the echo data arranged in the area, and arranging the echo data in the remaining area, and the first and second portions arranged by this means Of k space
Means for obtaining the image from two sets of echo data,
An MRI apparatus characterized in that