NMR imaging method of low flow rate fluid

An NMR imaging method of low flow rate fluid by measuring a time-reversed FID signal in a steady state free precession (SSFP) state comprises the steps of imaging in a dephase sequence, a rephase sequence or both of them, and processing the resulting image data by a maximum intensity projection method, a minimum intensity projection method or an image subtraction method to extract image data of fluid.

BACKGROUND OF THE INVENTION 
The present invention relates to an imaging method which is effective in 
rapidly imaging low flow rate fluid in a human body by a dislocation 
imaging device which utilizes nuclear magnetic resonance phenomenon 
(hereinafter referred to as MRI). 
The following articles 1) to 9) disclose the technologies relating to the 
present invention. 
1) IEEE Trans. on Medical Imaging, MI-5, No. 3, pp. 140-151, 1986. 
2) Magnetic Resonance in Medicine 6, 274-234, 1988. 
3) Magnetic Resonance in Medicine 7, 35-42, 1988. 
4) Journal of Magnetic Resonance 62, 12-18, 1985. 
5) Magnetic Resonance in Medicine 3, 140-145, 1986. 
6) Magnetic Resonance in Medicine 4, 9-23, 1987. 
7) Magnetic Resonance in Medicine 10, 324-337, 1989. 
8) U.S. Pat. No. 4,788,500. 
9) Magnetic Resonance in Medicine 14, 222-229, 1990. 
A prior art method for imaging fluid in the NMR imaging is discussed in 
detail in the reference 1). A principle of selection of blood flow 
utilizes a gradient magnetic field pulse which causes a change in a phase 
by flow or movement, which is called a flow encode pulse. Where the flow 
encode pulse is present along the direction of flow of the blood flow, a 
change of phase is produced in an excited spin in the blood flow in 
accordance with a flow rate. By subtracting between two images 
reconfigured based on NMR signals (hereinafter referred to as signals) 
detected in a dephase sequence which includes the flow encode pulse and a 
rephase sequence which does not include the flow encode pulse, image data 
of only the blood flow can be detected. A principle thereof is as follows. 
The blood flow in a blood vessel is a laminar flow which is of high flow 
rate at a center and of low flow rate at a periphery. Accordingly, when it 
is imaged by the dephase sequence, the excited spin has different phase 
change depending on a distance from the center of the blood vessel. As a 
result, signal data projected on a plane parallel to the blood flow has 
random orientation of the phases of the spins and resultant vectors cancel 
each other so that no signal is produced from the blood vessel or a signal 
amplitude is very small. On the other hand, when it is imaged by the 
rephase sequence, the phases of the spins which have once disturbed are 
reordered as the blood flows and if the signal is measured at a specific 
timing, the signal without phase change can be detected. Accordingly, when 
it is imaged by the rephase process, a signal may be detected even from 
the blood flow which includes a laminar flow. A signal may be detected 
from a static area by either one of the sequences, but the signals of the 
static area are cancelled out by subtracting the images produced by the 
two sequences so that the signal of only the blood vessel is produced. 
This method is generally called a subtraction method. 
A method for imaging a signal produced in a steady-state free precession is 
discussed in the references 2) to 4). When an RF pusle is irradiated at an 
extremely short repetition cycle compared to a relaxation time of the 
excited spin, the steady state free precession occurs so that the NMR 
signal is periodically produced in the imaging area in a stable manner, 
and a free induction decay (FID) signal is produced immediately after the 
RF pulse and a time-reversed free induction decay signal is produced 
immediately before the next RF pulse. It is known that the time-reverse 
FID signal has a similar property to that of an echo signal produced by 
the RF pulse of the two-preceding stage and the RF pulse of the preceding 
stage, that is, an echo signal having an echo time (TE) which is double of 
TR. 
For the time-reversed FID signal, the two sequentially applied RF pulses 
function as a 90.degree. pulse and a 180.degree. pulse in the spin echo 
pulse sequence. As a result, the RF pulse in the preceding stage functions 
to invert the phase of the NMR signal produced by the RF pulse of the 
preceding stage thereof. In this case, a gradient magnetic field is 
applied such that the phase-inverted signal converges, then diverses and 
is reconverged. 
A method for drawing slow flow rate fluid by utilizing the signal produced 
in the SSFP state is discussed in the references 5) to 8). Particularly in 
the reference 8), the method is discussed in detail. It describes that 
when the time-reversed FID signal produced in the SSFP state is to be 
detected, the interval of application of the RF pulses, that is, the 
repetition time is changed or the magnitude of the gradient magnetic field 
applied to detect the signal is changed to rotate the phase of the excited 
spin of the fluid so that two images are formed in the two dephase and 
rephase sequences having different gradient magnetic field magnitudes and 
the slow flow rate fluid is drawn by the subtraction between those images. 
In this case, depending on the degree of the phase change of the spin, a 
large signal of the fluid may not be produced by the subtraction. This is 
explained with reference to FIGS. 1 and 2. 
FIG. 1 shows a flow rate distribution of the blood flow. Numeral 1 denotes 
blood and numeral 2 denotes a direction of flow. The flow rate is high at 
points a and d, and low at points b and c. 
FIG. 2 show the phase rotations at the points a to d of FIG. 1 and combined 
signals of the signals at the points a to d in two different phase 
sequences. Mo represents a magnitude of signal in the SSFP state. 
As shown in FIG. 5, when the phases of the resultant signals are in the 
substantially same direction, a differential signal is very small. In the 
example shown in FIG. 2, 
EQU .vertline.Sd.sub.1 -Sd.sub.2 .vertline..apprxeq.2.8 Mo 
Depending on the phase angles of Sd.sub.1 and Sd.sub.2, the signal is 
smaller. 
In the reference 8), no consideration is paid to a method for drawing slow 
flow rate fluid based on data derived from measurement in sole dephase 
sequence or rephase sequence. 
A maximum intensity projection method and a minimum intensity projection 
method in the image processing are discussed in detail in the reference 
9). In the maximum intensity projection method, when a two-dimension image 
projected in any direction is to be produced based on three-dimensionally 
measured data and data which is three dimensionally measured by 
two-dimension multi-slice, a projection image is formed by using a maximum 
intensity pixel of a plurality of pixels of the three-dimension source 
data as a target pixel. In the minimum intensity projection method, a 
minimum intensity pixel is used as a target pixel. 
SUMMARY OF THE INVENTION 
It is an object of the present invention to image low flow rate fluid by 
using a high speed imaging method. 
A most important feature of the present invention resides in the use of the 
rephase sequence in detecting the time-reversed FID signal produced in the 
SSFP state, which rephase sequence has not been utilized in the prior art. 
In order to achieve the above object, in accordance with the present 
invention, the time-reversed FID signal is detected in the dephase 
sequence, the rephase sequence or both of them, and the fluid is imaged in 
the following three steps. 
(a) The minimum intensity projection method is applied to the image data 
produced in the dephase sequence to sample a low intensity signal as a 
fluid signal for imaging. 
(b) The maximum intensity projection method is applied to the image data 
produced in the rephase sequence to sample a high intensity signal as a 
fluid signal for imaging. 
(c) Images produced in the dephase sequence and the rephase sequence are 
subtracted to cancel the static area and sample the fluid for imaging.

DESCRIPTION OF THE PREFERRED EMBODIMENTS 
The preferred embodiments of the present invention will now be explained in 
detail with reference to the accompanying drawings. FIG. 3 shows a block 
diagram of an MRI device which implements the imaging method of the 
present invention. In FIG. 3, numeral 11 denotes a magnet which generates 
a uniform static magnetic field, numeral 12 denotes an exciting system for 
generating an RF magnetic field to cause nuclear magnetic resonance in an 
object, numeral 13 denotes a receiving system for receiving, detecting and 
A/D-converting an electromagnetic wave generated by the object, numeral 14 
denotes a gradient magnetic field generating system capable of 
independently and linearly changing the intensities of magnetic field in 
X, Y and Z directions, numeral 15 denotes an image processing system for 
processing image based on measurement data from a measuring system in 
order to reproduce an image, numeral 16 denotes a sequence control system 
for controlling the operation timings of the respective systems, numeral 
17 denotes a probe for transmitting and receiving the RF wave, and numeral 
18 denotes a console for the operation. 
FIG. 4 shows a signal produced in the SSFP state. Numeral 21 denotes an FID 
signal which is produced immediately after the RF pulse, and numeral 22 
denotes a time-reversed FID signal which is produced immediately before 
the RF pulse. It is known that the time-reversed FID signal has a similar 
property to that of an echo signal produced by a two-preceding RF pulse 23 
and a preceding RF pulse 24, that is, the echo signal having an echo time 
TE which is double of TR. The phase of the NMR signal excited by the RF 
pulse 23 is inverted by the RF pulse 24 so that the phase is again in 
phase and it becomes the time-reversed FID signal 22. 
FIGS. 5A, 5B and 6 show application patterns of the gradient magnetic field 
which is the principle of the present invention. 
FIGS. 5A and 5B show the application patterns of the dephase gradient 
magnetic field which causes the phase rotation in the spin of the fluid. 
They are called flow encode pulses. Numerals 31 and 32, and numerals 33 
and 34 denote gradient magnetic fields of equal application duration and 
opposite output values. The phase rotation .theta. of the fluid to which 
the gradient magnetic field shown in FIG. 5A is applied is represented by 
##EQU1## 
where V is the velocity of the fluid, .gamma. is a magnetic rotation 
ratio, Gx is a gradient magnetic field intensity and X(t)=X.sub.0 +Vt. 
The phase rotation .theta. of the spin to which the gradient magnetic field 
shown in FIG. 5B is applied is represented by 
##EQU2## 
FIG. 6 shows an application pattern of the rephase gradient magnetic field 
which compensates for the phase rotation of the spin. It takes a form of 
continuation of the two pulses, the flow encode pulse of FIG. 5A and the 
reversal of the flow encode pulse of FIG. 5A. The phase rotation of the 
fluid to which the above gradient magnetic field is applied is zero. 
##EQU3## 
FIG. 7 shows an example of change of phase of the spin at a position in the 
blood vessel when the time-reversed FID signal is imaged in the dephase 
sequence and the rephase sequence. In FIG. 7, a to d correspond to a-d in 
FIG. 1, and a+b+c+d represents a combined signal. 
Comparing it to FIG. 2, the differentiation of the image data produced by 
imaging in the dephase sequence and the rephase sequence enables the 
sampling of a stable and large signal. 
In the example of FIG. 7, the subtraction between the two images results in 
EQU S.sub.R =4Mo 
where Mo is same as that in FIG. 2. 
Embodiments of the gradient magnetic field pulse sequence to implement the 
present invention is shown in FIGS. 8, 9 and 10. 
FIG. 8 shows the two-dimension rephase sequence for correcting the phase 
rotation of the fluid based on the signal which is immediately preceding 
to the RF pulse produced in the SSFP state and draw the fluid as a high 
signal intensity area. The repetition times TR1 and TR2 are equal 
(TR1=TR2) and the pulse sequences in the TR1 and TR2 the same. 
An area to be imaged is selected by slicing 42 and it is excited by an RF 
pulse of any angle 41. It is excited with extremely short repetition times 
TR1 and TR2 compared to a relaxation time of the area to be imaged, and a 
signal produced immediately before the RF pulse is sampled as an echo 
signal by gradient magnetic fields 49, 50, 51, 52 and 53 and it is sampled 
by a gate pulse 54. Those are processed in the same phase encode 
projection 47 and 48. Pulses 43 and 44, and pulses 45 and 46 are applied 
to recover the phase disturbed along the slice direction in the slice 
selection, and pulses 42, 43 and 44 and pulses 45, 46 and 42 of TR2 form a 
rephase gradient magnetic field application pattern which renders the 
phase rotation of the fluid to zero. Pulses 49, 50, 51, 52 and 53 are 
applied to frequency-encode the signal and collect the each signal. Those 
pulses form a rephase gradient magnetic field pattern. A phase encode 
pulse 48 which is equal in the absolute value and opposite in the polarity 
to the pulse 47 is applied to reset the phase rotation of the spin after 
the sampling. The signal is measured while the phase encode projection of 
the pulses 47 and 48 is varied to reconfigure the image so that the fluid 
can be drawn as a high intensity signal area. 
FIG. 9 shows a two-dimension dephase sequence which causes the phase 
rotation of the spin of the fluid from the signal which is immediately 
preceding to the RF pulse produced in the SSFP state and draws the fluid 
as a low intensity signal area. The repetition times TR1 and TR2 are same 
as those of FIG. 8. 
Pulses 82, 83, 84 and 82 and 83 of TR2 form a dephase gradient magnetic 
field application pattern along the slice direction, and pulses 87 and 88 
form a dephase gradient magnetic field application pattern along the 
frequency encode direction. S1 and S2 of the pulse 87 represent pulse 
area. S2 is equal to the area of the pulse 88. Since an RF pulse 81 of TR2 
functions to invert the phase of the spin excited by the RF pulse 81 of 
TR1, the pulse 87 of TR2 functions as the opposite polarity pulse to the 
pulse 87 of TR1 for the excited spin. The pulse 87 of TR1 corresponds to 
the pulse 33 of FIG. 5B, and the pulse 87 of TR2 corresponds to the pulse 
34 of FIG. 5B. As a result, the echo signal has a maximum peak when S1 of 
the pulse 87 of TR1 and S1 of the pulse 87 of TR2 are equal. It is sampled 
by a gate pulse 89. In general, the fluid having a low flow rate has small 
V in the formulas (1) and (2) and hence has small .theta. and a change in 
the signal intensity by a change in the phase for the static area is 
small. As a result, the sampling of the fluid is difficult to attain. 
However, since the time-reversed FID signal has TE which is double of TR, 
.tau. and .tau..sub.p in FIG. 5B are large accordingly. 
In FIG. 9, .tau. and .tau..sub.p in the formula (2) are large in order to 
form the dephase gradient magnetic field application pattern in which the 
pulse 87 of TR1 and the pulse 87 of TR2 have large .tau. and .tau..sub.p. 
As a result, the phase rotation .theta. is large and the change in the 
signal intensity for the static area is large. Thus, the sampling of the 
fluid having a low flow rate is easier. 
FIG. 10 shows a two-dimension dephase sequence in which an area 
corresponding to the area S.sub.1 of the pulse 87 of FIG. 9 is extended in 
time and a gradient magnetic field along the frequency encode direction is 
applied to maximize the phase rotation of the time-reversed FID signal. 
The phase rotation of the time-reversed FID signal can be increased by 
increasing the integrated application amount of the pulse 90. 
A principle of the embodiment of FIG. 10 is explained in FIGS. 11A and 11B. 
In FIG. 11A, the application time of the pulse 90 is short. In this case, 
the phase rotation .theta. of the spin is given by 
##EQU4## 
The definitions of the parameters are identical to those in the formula 
(1). 
In FIG. 11B, the application time of the pulse 90 is long. The phase 
rotation .theta. of the spin is given by 
##EQU5## 
As seen from the formulas (4) and (5), the longer the application time of 
the pulse 90 is, the larger is the phase rotation. The pulse 90 may be 
extended up to the end time of the application of the RF pulse 81. 
FIGS. 12 and 13 illustrate a maximum intensity projection method and a 
minimum intensity projection method applied to the present invention. 
When a two-dimension image is to be formed by projecting to any plane based 
on three-dimensionally measured data and three dimension image data by 
two-dimension multi-slice measurement, if the signal intensities A, B and 
C of the image data on the projection line of the source three-dimension 
image data 101 have a relationship of A&gt;B&gt;C, the signal of the fluid 
produced in the rephase sequence has a large intensity. Thus, the image 
data projected on the projection plane 102 uses the maximum intensity data 
"A" to prepare the projection image of the fluid. On the other hand, since 
the signal of the fluid produced in the dephase sequence is low, the image 
data projected on the projection plane 103 uses the minimum intensity data 
"C" to prepare the projection image of the fluid.