JPH05123313A - Magnetic resonance imaging device - Google Patents

Abstract

PURPOSE:To provide a satisfactory blood vessel image having no unevenness while keeping a short measuring period without increasing the imaging time. CONSTITUTION:A pulse sequencer for sending a blood flow to a sequencer is provided. In the measurement of the pulse sequencer, only the first echo signal measurement of an imaging target slice is synchronized with an electrocardiogram, and the measurement repeated thereafter with the same slice is conducted according to a repeating time set non-synchronously to all the electrocardiograms. Thus, the signal strength of blood flow is nearly uniformized between each slice without being affected by the heartbeat, and the two- dimensional projected blood vessel image can be also obtained as a uniform image.

Description

Detailed Description of the Invention

[0001]

The present invention relates to nuclear magnetic resonance (hereinafter referred to as N
The present invention relates to a magnetic resonance imaging apparatus for obtaining a tomographic image of a desired site of a subject by utilizing a phenomenon (abbreviated as MR), and more particularly to a magnetic resonance imaging apparatus having a blood flow drawing means.

[0002]

2. Description of the Related Art A magnetic resonance imaging apparatus utilizes an NMR phenomenon to measure the density distribution, relaxation time distribution, etc. of nuclear spins (hereinafter simply referred to as "spins") at a desired inspection site in a subject, An image of an arbitrary cross section of the subject is displayed as an image from the measurement data. As shown in FIG. 3, the conventional magnetic resonance imaging apparatus includes a static magnetic field generating magnet 2 for applying a static magnetic field to the subject 1, a magnetic field gradient generating system 3 for applying a gradient magnetic field to the subject 1, and A sequencer 7 that repeatedly applies a high-frequency pulse that causes NMR to the atomic nuclei of the atoms that form the biological tissue of the sample 1 in a predetermined pulse sequence, and the atoms that form the biological tissue of the sample 1 by the high-frequency pulse from the sequencer 7. A transmission system 4 for irradiating a high frequency magnetic field to cause NMR to the nuclei of the above, a reception system 5 for detecting an echo signal emitted by the above-mentioned NMR, and an image reproduction by using the echo signal detected by this reception system 5. A signal processing system 6 for performing a configuration calculation was provided, and echo signals emitted by NMR were repeatedly measured to obtain a tomographic image.

In this device, as shown in FIG.
~ Static magnetic field generating magnet 2 for generating a static magnetic field of about 2 tesla
The subject 1 is placed inside. At this time, the spins in the subject 1 perform precession about the direction of the static magnetic field at a frequency determined by the strength H ₀ of the static magnetic field. This frequency is called the Larmor frequency, and the Larmor frequency ν ₀ is Here, H ₀ : static magnetic field strength γ: gyromagnetic ratio has a unique value for each kind of nucleus. When the angular velocity of the Larmor precession is ω ₀ , ω ₀ = 2πν ₀ , and therefore ω ₀ = γ · H ₀ (2)

The high frequency irradiation coil 14 in the transmission system 4
Larmor frequency ν of the nucleus to be measured by a
_When a high frequency magnetic field (electromagnetic wave) having a frequency f ₀ equal to ₀ is applied, spins are excited and transition to a high energy state.
When this high-frequency magnetic field is cut off, the spins return to the original low energy state with a time constant corresponding to each state. The electromagnetic wave emitted at this time is received by the high frequency receiving coil 14b in the receiving system 5, amplified by the amplifier 15, shaped in waveform, and then digitized by the A / D converter 17 to be processed by a central processing unit (hereinafter referred to as CPU). Send to 8. The CPU 8 reconstructs an image based on this data, and displays a tomographic image of the subject 1 on a display (hereinafter referred to as CRT) 20. The above-mentioned high-frequency magnetic field is obtained by amplifying a signal sent from the sequencer 7 controlled by the CPU 8 by a high-frequency transmitting coil power supply (not shown).
Obtained by sending to 4a.

In addition to the above static magnetic field and high frequency magnetic field, the magnetic resonance imaging apparatus described above is provided with a gradient magnetic field coil 9 for producing a gradient magnetic field for obtaining positional information in space. These gradient coil 9
A gradient magnetic field is generated by being supplied with a current from a gradient magnetic field power source 10 that operates by a signal from the sequencer 7.

Here, the imaging principle of the magnetic resonance imaging apparatus will be described with reference to FIG. First, FIG.
As shown in (a), an atomic nucleus placed in a static magnetic field H ₀ in the Z direction behaves like a single bar magnet in classical physics, and has the Larmor frequency ν ₀ described above in the Z axis. Are doing precession around. This frequency is given by the equation (2) and is proportional to the strength H _{0 of the} static magnetic field. First (1)
Γ in the equation and the equation (2) is called a gyromagnetic ratio and has a value unique to the nucleus. In general, the number of nuclei to be measured is enormous and each rotates in an arbitrary phase. Therefore, when viewed as a whole, the components in the XY plane cancel each other out, and the macroscopic magnetization of only the Z direction component is generated. Remain. In this state, as shown in FIG. 4 (b), the Larmor frequency ν
_When a high frequency magnetic field H ₁ having a frequency equal to ₀ is applied, the macroscopic magnetization begins to fall in the Y direction. This tilting angle is proportional to the product of the amplitude of the high frequency magnetic field H ₁ and the application time, and the high frequency magnetic field H ₁ when tilting 90 ° with respect to the pulse application time is 90 ° pulse, and the high frequency magnetic field H ₁ when tilting 180 ° is 180 °. ° Called a pulse. In addition, X in FIG. 4 (a), (b),
The Y and Z three axes are Cartesian coordinate axes that are orthogonal to each other.

A two-dimensional Fourier imaging method is a method generally used for imaging in such a magnetic resonance imaging apparatus. FIG. 5 is a timing diagram schematically showing a pulse sequence of a typical spin echo method of the above two-dimensional Fourier imaging method. Figure 5
2A, the irradiation timing of the high-frequency magnetic field signal and the envelope for selectively exciting the slice position of the subject are shown in FIG. (B) shows the application timing of the gradient magnetic field Gz in the slice direction, and (c) shows the application timing of the gradient magnetic field Gy in the phase encoding direction and the measurement by changing its amplitude, and (d). The figure shows the timing of application of the gradient magnetic field Gx in the frequency encoding direction. Further, FIG. 7E shows the echo signal (NMR signal) to be measured.

In the pulse sequence shown in FIG. 5, first,
After applying the 90 ° pulse, a 180 ° pulse is applied at Te / 2 when the echo time is Te. After the 90 ° pulse is applied, each spin starts rotating in the XY plane at its own velocity, so that a phase difference occurs between the spins over time. Here, when a 180 ° pulse is applied, each spin reverses symmetrically with respect to the X-axis, and since it continues to rotate at the same speed thereafter, the spin converges again at time Te shown in FIG. 5, and as shown in FIG. To form an echo signal. The signal is measured as described above, but the spatial distribution of the signal must be obtained in order to form a tomographic image. For this purpose, a linear gradient magnetic field is used. A spatial magnetic field gradient can be created by superimposing a gradient magnetic field on a uniform static magnetic field. Since the spin rotation frequency is proportional to the magnetic field strength as described above, the spin rotation frequencies are spatially different in the state in which the gradient magnetic field is applied. Therefore, the position of each spin can be known by examining this frequency. For this purpose, the phase encoding direction gradient magnetic field Gy and the frequency encoding direction gradient magnetic field Gx shown in FIG. 5 are used.

Using the pulse sequence described above as a basic unit, the intensity of the phase-encoding direction gradient magnetic field Gy is changed every time, and the pulse sequence is repeated a predetermined number of times, for example 256 times, every constant repetition time Tr. The spatial distribution of macroscopic magnetization shown in FIG. 4A is obtained by subjecting the measurement signal thus obtained to two-dimensional inverse Fourier transform. In the above explanation,
If the three types of gradient magnetic fields do not overlap with each other, X, Y, Z
Or any combination thereof. Regarding the basic principle of the above magnetic resonance imaging, "NMR Medicine (Basic and Clinical)" (edited by Nuclear Magnetic Resonance Medical Research Group, Maruzen Co., Ltd., Showa)
Issued January 20, 1959).

Next, the principle of blood flow drawing in the magnetic resonance imaging apparatus will be described. As a blood flow drawing method in a magnetic resonance imaging device, firstly, a time-of-flight method using an inflow effect on a slice plane, and secondly, a phase-sensit method for performing a difference by using phase diffusion due to blood flow
Three types of methods are mainly used: the ive method, and thirdly, the phase-contrast method that inverts the polarity of phase diffusion due to blood flow and makes a difference.

First, the first time-of-flight method will be briefly described. In a magnetic resonance imaging device,
When the same region is continuously excited by a high-frequency magnetic field for a short period of time, for example, several tens of ms, the spins contained in the tissue in the region are saturated and the emitted signal becomes low. On the other hand, the spins contained in the blood flow always flow out of the region, and new unsaturated spins flow in, so that relatively high signals are emitted from other tissues. By utilizing this inflow effect, imaging is performed on a plurality of slices, and the obtained images are superimposed and projected, whereby blood flow can be drawn. The increase in blood flow due to this inflow effect is due to "Magnetic Resonance Imaging.
Stark DD et al. Edited, The CVMosby Company, pp1
Details of "08 ~ 137,1988".

Next, the second Phase-sensitive method will be briefly described. In the magnetic resonance imaging apparatus, several kinds of gradient magnetic fields are applied as described above when measuring an echo signal (NMR signal), but the spins excited by the application of these gradient magnetic fields depend on the position and the moving speed. It undergoes phase rotation. That is, as shown in FIG. 6, for example, at time t ₁ , two spins S ₁ and S ₂ exist at the position of X ₀ , one spin S ₁ is stationary, and the other spin S ₂ is When moving in the X direction at the velocity v, as shown in FIG. 7, by applying the gradient magnetic field Gx in the frequency encoding direction from time t ₁ to t ₂ , the phase changes Φs and Φf shown in the following equations, respectively. Receive. From the expressions (3) and (4), Becomes From this equation (5), the stationary spin Φs
It can be seen that the phase difference between (S ¹ ) and the moving spin Φf (S ² ) is proportional to the moving speed v.

Therefore, in the standard spin echo sequence shown in FIGS. 8A to 8E, when the echo signal E is measured, as shown in FIG.
The phases of the moving spin Φf (shown by the solid curve) are not aligned with each other (shown by the broken curve). Here, with respect to the gradient magnetic field in the standard spin echo sequence, as shown in FIG. 9B, the gradient magnetic fields A and B in the negative direction are used.
By adding, the stationary spin and the moving spin are aligned in phase in agreement with the peak of the echo signal E (see FIG. 9 (a)) as shown in FIG. 9 (e). Below, FIG.
The pulse sequences shown in (a) to (e) are called a phase sensitive sequence, and are shown in FIGS.
The pulse sequence shown in (e) is phase insensitive (Phase in
sensitive) sequence. In the phase-insensitive sequence, a signal having the same signal strength as that obtained in the phase-sensitive sequence is obtained in the stationary portion, and the loss of the signal due to phase diffusion is suppressed in the region where the moving magnetization exists, which is higher than that in the phase-sensitive sequence. The signal is obtained. Therefore, FIG.
As shown in, the difference image I ³ is obtained by taking the difference between the phase-sensitive image I ¹ measured by the phase-sensitive sequence and the phase-insensitive image I ² measured by the phase-insensitive sequence. Only moving parts such as blood flow can be imaged. For a method of obtaining a blood vessel image from the difference between images obtained by such a phase-sensitive sequence and a phase-insensitive sequence, see “Cerebral MR A
Research on ngioimaging (cerebral vascular magnetic resonance imaging) -First report- "(Keiji Fukui et al., CT Study 10 (2) 1988), 133.
It is detailed from page to page 142.

The third Phase-contrast method will be described. As described above, the spin in the bloodstream undergoes phase rotation due to the application of the gradient magnetic field. Therefore, when a pair of positive and negative gradient magnetic field pulses as shown in FIG. 11 is applied, the spins in the blood flow receive the phase rotation Φf− or Φ + depending on the flow velocity depending on the application order. This pair of positive and negative gradient magnetic field pulses is called a flow encode pulse. Reverse the polarity of the flow encode pulse (that is, reverse the order of +-)
If so, since the polarity of the phase rotation is also reversed as described above,
If these signals are applied alternately and the complex difference of the obtained signals is taken, the signal of the stationary portion which is not subjected to the phase rotation is removed and only the blood flow signal is detected. The obtained signal intensity changes depending on the flow velocity, and the signal intensity becomes maximum when the flow velocity is such that the phase difference between Φf− and Φf + is π. This Phase Contras
For the t method, see “Magnetic resonance angiography.
Dumoulin CL, et al. Radiology 161: 717 ~ 720, 1986 ".

[0015]

In the above-mentioned prior art, particularly in the two-dimensional time-of-flight method, when the measurement of each slice is performed by ignoring the beat cycle of the heart of the subject,
Since the blood flow velocity is different depending on the cardiac phase, the flow velocity is different at each phase encoding step, the inflow effect is changed, and the signal derived from the blood flow is also changed. In this case, in each slice, a signal intensity depending on the flow velocity near 0 phase encoding is exhibited. Since the flow velocity in the vicinity of 0 phase encoding is indefinite for each slice, the signal intensity in the blood vessel changes for each slice, especially in a blood vessel whose speed changes drastically.
As shown in Fig. 3, the image is nonuniformly drawn. On the other hand, in order to avoid this, when the ECG gated imaging shown in FIG. 14 is performed, the measurement period (hereinafter referred to as TR) becomes the same as the heartbeat cycle or becomes an integral multiple, which is required for the time-of-flight method. Short TR becomes impossible. In addition, the imaging time becomes enormous.

An object of the present invention is to provide a good blood vessel image free from nonuniformity while maintaining such a short TR required for the time-of-flight method without increasing the imaging time.

[0017]

In order to achieve the above object, a static magnetic field generating means for giving a static magnetic field to a subject, a gradient magnetic field generating means for giving a gradient magnetic field to the subject, and a living tissue of the subject. A sequencer that repeatedly applies a high-frequency pulse for causing NMR to the atomic nuclei of the atoms in a predetermined pulse sequence, and a high-frequency magnetic field for causing NMR to the atomic nuclei of the biological tissue of the subject by the high-frequency pulse from this sequencer. Transmitting system to irradiate, and the above N
A receiving system for detecting an echo signal emitted by the MR,
A signal processing system for performing an image reconstruction calculation using the echo signal detected by this receiving system and a means for displaying the obtained image are provided, and the echo signal emitted by NMR is repeatedly measured to obtain a tomographic image. In the magnetic resonance imaging apparatus to be obtained, the sequencer comprises a pulse sequence for depicting the blood flow in the subject, and when measuring using the pulse sequence, only the first echo signal measurement of the slice to be imaged is synchronized with the electrocardiogram. Measurement shall be started,
All measurements repeated thereafter in the same slice enable electrocardiographic approximately synchronous blood vessel imaging by performing measurements according to the repetition time set asynchronously in the electrocardiogram.

Further, the above-mentioned blood flow drawing pulse sequence is
It is a pulse sequence based on the two-dimensional Fourier transform method using the inflow effect, and it is assumed that multiple adjacent thin slices are sequentially imaged, and an echo signal is synchronized with the electrocardiogram only at the start of imaging each slice. It is a pulse sequence based on the two-dimensional Fourier transform method that collects blood and images the blood flow by the difference between the images by the rephase method and the dephase method, and it is assumed that a plurality of adjacent thin slices are sequentially imaged. The echo signals are collected in synchronization with the electrocardiogram only when the imaging of each slice is started.

[0019]

In the magnetic resonance imaging apparatus configured as described above, after the predetermined pre-measurement is completed, the measurement of each slice is started. In the measurement, the sequencer receives the R wave of the electrocardiogram sent from the electrocardiograph or a gate signal generated from the R wave, and controls the measurement. More specifically with reference to FIG. 1, the sequencer receiving the gate signal immediately measures the signal in the first encoding step of the first sheet. Then, after a predetermined TR, the signal measurement of the second encoding step is performed asynchronously with the electrocardiogram. When this process is repeated and, for example, data collection for 256 encodings is completed, the sequence shifts to the measurement of the second slice. Even when measuring the second slice,
The signal measurement in the first encoding step is performed in synchronization with the electrocardiogram. In this way, in each slice, only the first encoding is always performed in synchronization with the electrocardiogram.

[0020]

Embodiments of the present invention will now be described in detail with reference to the accompanying drawings. FIG. 3 is a block diagram showing the overall configuration of the magnetic resonance imaging apparatus according to the present invention. This magnetic resonance imaging apparatus obtains a tomographic image of a subject by utilizing the NMR phenomenon, and as shown in FIG. 3, a static magnetic field generating magnet 2, a magnetic field gradient generating system 3, a transmitting system 4, and a receiving system 4. System 5
A signal processing system 6, a sequencer 7, and a CPU 8. The static magnetic field generating magnet 2 is for generating a uniform static magnetic field around the subject 1 in the body axis direction or in a direction orthogonal to the body axis, and the static magnetic field generating magnet 2 is provided in a space with a certain extent around the subject 1. Permanent magnet type, normal conducting type or superconducting type magnetic field generating means is arranged. The magnetic field gradient generation system 3 is composed of a gradient magnetic field coil 9 wound in three directions of X, Y, and Z, and a gradient magnetic field power supply 10 for driving the respective gradient magnetic field coils, and according to an instruction from a sequencer 7 described later. By driving the gradient magnetic field power supply 10 of each coil, the gradient magnetic field G in the three axis directions of X, Y, and Z is obtained.
x, Gy, and Gz are applied to the subject 1. The slice plane for the subject 1 can be set by the method of applying this gradient magnetic field.

The sequencer 7 repeatedly applies a high frequency magnetic field pulse for causing NMR to the atomic nuclei of the atoms constituting the biological tissue of the subject 1 in a predetermined pulse sequence, and operates under the control of the CPU 8 to Various commands necessary for collecting data of the tomographic image 1 are sent to the transmission system 4, the magnetic field gradient generation system 3 and the reception system 5. The transmission system 4 irradiates a high frequency magnetic field in order to cause NMR to atomic nuclei of the atoms constituting the biological tissue of the subject 1 by the high frequency pulse sent from the sequencer 7, and includes a high frequency oscillator 11, a modulator 12 and a high frequency wave. The high frequency pulse output from the high frequency oscillator 11 is amplitude-modulated by the modulator 12 according to the instruction of the sequencer 7, and the high-frequency pulse is modulated by the high-frequency amplifier 13. After amplification, the electromagnetic waves are radiated to the subject 1 by supplying the high-frequency coil 14a arranged close to the subject 1 with the electromagnetic waves.

The receiving system 5 detects an echo signal (NMR signal) emitted by NMR of atomic nuclei of the living tissue of the subject 1, and includes a high frequency coil 14b on the receiving side, an amplifier 15 and a quadrature detector 16. , A / D converter 17, and the electromagnetic wave (NMR signal) of the response of the subject 1 due to the electromagnetic wave emitted from the high-frequency coil 14a on the transmitting side is generated by the high-frequency coil 14b arranged close to the subject 1. Detected and passed through amplifier 15 and quadrature detector 16 to A
It is input to the D / D converter 17, converted into a digital amount, and further converted into two series of collected data sampled by the quadrature phase detector 16 at the timing according to the instruction from the sequencer 7, and the signal is sent to the signal processing system 6. It is designed to be used. The signal processing system 6 includes a CPU 8, a recording device such as a magnetic disk 18 and a magnetic tape 19, and a CRT.
The CPU 8 performs processing such as Fourier transform, correction coefficient calculation image reconstruction, and the like, and the signal intensity distribution of an arbitrary cross section or a distribution obtained by performing an appropriate calculation on a plurality of signals is imaged to the CRT 20. It is designed to be displayed as a tomographic image.

In FIG. 3, the high frequency coils 14a and 14b on the transmitting side and the receiving side and the gradient magnetic field coil 9 are installed in the magnetic field space of the static magnetic field generating magnet 2 arranged in the space around the subject 1. There is. Further, the electrocardiographic detection system 31 detects the electrocardiogram of the subject 1 and sends it to the electrocardiograph 30 as an electric signal. The electrocardiograph 30 displays the received electrocardiogram signal on the monitor and sends the gate signal generated from the electrocardiogram R wave to the sequencer 7.

Here, in the present invention, the sequencer 7 starts the blood vessel drawing sequence substantially in synchronization with the gate signal. The blood vessel drawing sequence at this time is the same as the above
It may be based on any of the types of blood vessel delineation methods. The two-dimensional time-of-flight method will be described below as an example. FIG. 1 is a timing diagram schematically showing the relationship between the echo signal measurement process of each slice and the electrocardiogram in the electrocardiogram substantially synchronous imaging of the present invention, and FIG. 2 shows it at a detailed level for each phase encoding. FIG. In the two-dimensional time-of-flight method, short TR (about several tens of ms) imaging is repeatedly performed while moving the position by the number of slices to obtain three-dimensional data.

In FIG. 1, (a) shows an electrocardiographic waveform,
(B) shows the timing of signal measurement including all excitation by RF pulse, application of gradient magnetic field, and echo signal measurement,
(C) shows the number of the slice to be imaged at that time. In FIG. 2, (a) shows an electrocardiographic waveform,
(B) shows the irradiation timing of the RF pulse, (c) shows the application timing and intensity of the phase encoding gradient magnetic field, (d) shows the timing of appearance of the echo signal,
(E) shows the phase encode step number of each TR. In FIG. 1, with the appearance of the gate signal, the first
Start slice measurement. The start of measurement is the first in FIG.
This is the same as the signal measurement in the encode step. When the measurement of the first encoding step is completed (from the first excitation to T
After R time), measurement after the second encoding step is continuously performed. At this time, the sequence is completely unrelated to the gate signal from the electrocardiograph. This continues until the signal measurement for the predetermined number of phase encodes of the first slice is completed.

After the signal measurement of the first slice is completed, the sequencer waits for a gate signal for measuring the second slice. This waiting time is the gate waiting time in FIG. When the gate signal appears and the measurement of the second slice is started, the measurement is performed in the same procedure as that of the first slice. By repeating this for the number of slices, two-dimensional images for the number of slices can be obtained. These two-dimensional images are images in which the blood flow is depicted as a relatively high signal, and the three-dimensional data of blood vessels are obtained by sequentially stacking these. This three-dimensional blood vessel data is converted into an arbitrary two-dimensional projection image by the projection processing shown in FIG.

[0027]

Since the present invention is constructed as described above,
By controlling the pulse sequence by the sequencer, the signal intensity of the blood flow is substantially uniform between slices without being affected by heartbeat, and a two-dimensional projected blood vessel image can be obtained as a uniform image.

[Brief description of drawings]

FIG. 1 is a timing diagram schematically showing the entire electrocardiogram-schematic synchronized blood vessel imaging pulse sequence in the magnetic resonance imaging apparatus of the present invention.

FIG. 2 is a timing diagram schematically showing the relationship between an electrocardiogram-schematic synchronized blood vessel imaging pulse sequence and an electrocardiogram in the magnetic resonance imaging apparatus of the present invention.

FIG. 3 is a block diagram showing an overall configuration of the present invention and a conventional magnetic resonance imaging apparatus.

FIG. 4 is an explanatory view showing a behavior of spins for explaining an imaging principle of a magnetic resonance imaging apparatus.

FIG. 5 is a timing diagram schematically showing a pulse sequence of a typical spin echo method among two-dimensional Fourier imaging methods in a general magnetic resonance imaging apparatus.

FIG. 6 is an explanatory diagram showing how spins move in a gradient magnetic field.

FIG. 7 is an explanatory diagram showing phase rotations of a moving spin and a stationary spin in a gradient magnetic field.

FIG. 8 is an explanatory diagram showing a state of phase rotation of moving spins in a phase sensitive sequence.

FIG. 9 is an explanatory diagram showing a state of phase rotation of moving spins in a phase insensitive sequence.

FIG. 10 is an explanatory diagram showing a difference between images in the Phase-sensitive method.

FIG. 11 is an explanatory diagram showing the principle of phase difference extraction of flow encode pulses in the Phase-contrast method.

FIG. 12 is an explanatory diagram showing a projection process for creating a two-dimensional blood vessel image with an arbitrary projection angle from three-dimensional blood vessel data.

FIG. 13 is an explanatory view showing a manner of uniformly drawing blood vessels according to the present invention.

FIG. 14 is an explanatory view schematically showing an electrocardiographically synchronized imaging pulse sequence in the conventional method.

[Explanation of symbols]

 1 subject 2 magnetic field generator 3 magnetic field gradient generation system 4 transmission system 5 reception system 6 signal processing system 7 sequencer 8 CPU 9 gradient magnetic field coil 10 gradient magnetic field power supply 14a transmission side high frequency coil 14b reception side high frequency coil 30 electrocardiograph 31 ECG detection system

─────────────────────────────────────────────────── ─── Continuation of front page (51) Int.Cl. ⁵ Identification code Internal reference number FI Technical display location 7831-4C A61B 5/05 312 7831-4C 377 9118-2J G01N 24/08 Y

Claims (3)

[Claims] 1. A static magnetic field generating means for applying a static magnetic field to a subject, and a gradient magnetic field generating means for applying a gradient magnetic field to the subject.
A sequencer that repeatedly applies a high-frequency pulse that causes nuclear magnetic resonance to the atomic nuclei of the biological tissue of the subject in a predetermined pulse sequence, and a nucleus in the atomic nucleus of the biological tissue of the subject by the high-frequency pulse from the sequencer. A transmission system that radiates a high-frequency magnetic field to cause magnetic resonance, a reception system that detects an echo signal emitted by the above-mentioned nuclear magnetic resonance, and an image reconstruction calculation using the echo signal detected by this reception system. In a magnetic resonance imaging apparatus, which comprises a signal processing system for performing and a means for displaying the obtained image, in which the echo signal emitted by nuclear magnetic resonance is repeatedly measured to obtain a tomographic image, the sequencer is in the subject. It is equipped with a pulse sequence that visualizes the blood flow of the Only echo signal measurement shall be started in synchronism with the electrocardiogram, and subsequent repeated measurements in the same slice are all performed asynchronously with the electrocardiogram, so that electrocardiogram-simultaneous vascular imaging is possible. A magnetic resonance imaging apparatus characterized by the above. 2. The blood flow depiction pulse sequence is a pulse sequence based on a two-dimensional Fourier transform method using an inflow effect, and a plurality of adjacent thin slices are sequentially and continuously imaged, and The magnetic resonance imaging apparatus according to claim 1, wherein echo signals are collected in synchronization with an electrocardiogram only at the start of imaging. 3. The pulse sequence for drawing a blood flow is a pulse sequence based on a two-dimensional Fourier transform method for imaging a blood flow by a difference between images by a rephase method and a dephase method, and a plurality of adjacent thin slices. 2. The magnetic resonance imaging apparatus according to claim 1, wherein the echo signals are collected in synchronism with the electrocardiogram only when the imaging of each slice is started.