Flow imaging by means of nuclear magnetic resonance

A nuclear magnetic resonance projection image of fluid flow in a vessel is obtained by generating two data sets encoded with phase information indicative of two different velocity profiles of the fluid (taken at different times), combining the data sets, and displaying the resulting data set as an image.

STRUCTURE 
Referring to FIG. 1, blood vessel 10 lies within a matrix of tissue 12 and 
carries blood flowing in the direction indicated by arrow 14. The velocity 
of the blood in direction tends to vary depending on how close it is to 
wall 16 of vessel 10. For example, in certain types of laminar flow the 
velocity of the blood is smallest at wall 16 and increases to a maximum at 
the central axis x of vessel 10. Thus, for a particular imaginary planar 
slice 18 (having a particular location along axis z) the blood velocity 
component in the x direction might vary from small to large to small as 
one traverses the slice from one wall to the opposite wall in a direction 
indicated by axis y. 
Referring to FIG. 2, nuclear magnetic resonance apparatus 30 has a magnet 
32 for imposing a uniform constant magnetic field B.sub.o on vessel 10 and 
tissue 12. A G.sub.z gradient coil 34 (arranged to provide a linear 
magnetic field gradient along the z-axis) is connected to a G.sub.z driver 
36, which can provide selected levels of power at selected times to coil 
34. Likewise, G.sub.x gradient coil 38 and G.sub.y gradient coil 40, which 
can provide linear magnetic field gradients respectively along the x-axis 
and y-axis, are respectively connected to G.sub.x driver 42 and G.sub.y 
driver 44, for providing power to the coils. 
An RF antenna 50 (arranged to transmit and receive an RF signal to and from 
blood vessel 10 and tissue 12) is connected to modulator 52 which is in 
turn connected to an RF signal generator 54 and is arranged to impose an 
RF signal at a selected frequency and at selected times upon vessel 10 and 
tissue 12. RF antenna 50 is also connected via amplifier 57 to demodulator 
58, which is arranged to demodulate received signals (from vessel 10 and 
tissue 12) into quadrature real and imaginary components (indicated by the 
doubling of line 60). 
Computer processor 70 is connected via I/O interface 71 to G.sub.x, 
G.sub.y, and G.sub.z, drivers 42, 44, 36, to control the magnitude and 
timing of the linear gradient pulses imposed along the three axes upon 
vessel 10 and tissue 12, and to signal generator 54 to control the 
frequency, magnitude, and timing of the RF pulses. 
Processor 70 is also connected via I/O interface 71 to demodulator 58 to 
receive the real and imaginary components of the demodulated received RF 
signal. 
Processor 70 is further connected to a memory 72 (which stores software to 
control the operation of the system and data representing the received RF 
signals), to a CRT 74 (which displays images representing the received RF 
signal, as well as information needed by the operator to control the 
system), and to a keyboard 76 (by which the operator can enter data and 
information to control the entire operation of the system). A camera 75 
attached to CRT 74 takes photographs of the displayed images. 
Referring to FIG. 3, processor 70 includes G.sub.z pulse magnitude and 
timing control 90 which is connected to trigger G.sub.z driver 36 to apply 
selected levels of gradient at selected times. Likewise G.sub.x and 
G.sub.y pulse magnitude and timing controls 92, 94 are connected to 
trigger respectively G.sub.x and G.sub.y drivers 42, 44. 
Similarly RF signal generator frequency, magnitude, and timing control 96 
is connected to trigger RF signal generator 54 to provide an RF pulse 
having a selected frequency and magnitude, at selected times. 
Controls 90, 92, 94, 96 are all connected to a sequencer 100 which provides 
the necessary specific magnitude, timing, and frequency parameters at the 
proper times for a given NMR field pulse sequence. 
Sequencer 100 is connected to a user-entered parameter processor 102 which 
receives parameters from the keyboard and converts them to a form usable 
by sequencer 100. 
Sequencer 100 is also connected to control an A-to-D converter 104 which is 
in turn connected to demodulator 58 to digitize the real and imaginary 
components of the received RF signal. A sampler 106 is connected to the 
converter 104 to sample the digitized signal components and to memory 72 
to store the samples for later processing. 
A Fourier transform element 108 is connected to memory 72 for performing a 
complex two-dimensional Fourier transform on a family of received signal 
samples to produce real and imaginary components of a two-dimensional 
array of data in the frequency domain. The parameters in accordance with 
which the Fourier transform is performed are received from processor 102 
based on user entered information. 
A display controller 110 is connected to Fourier transform element 108 to 
organize and process the frequency domain data into image information for 
delivery to CRT 74, again in accordance with parameters received from 
processor 102 as provided by the user. 
Sequencer 100 is also connected to gating circuitry 112 which provides 
signals enabling the sequencer to synchronize successive pulse sequences 
to occur at the same point in successive heart beats. 
The invention can be implemented by appropriately connecting and 
configuring available hardware and by specifying operating parameters for 
available related software, in accordance with the foregoing and following 
description (for example, NMR imaging hardware and software available from 
Technicare Corporation, Solon, Ohio, or similar systems available from 
other vendors). 
OPERATION 
Referring to FIGS. 4, 5, in order to generate blood velocity profile images 
corresponding to planar slice 18 (FIG. 1), vessel 10 is oriented as nearly 
as possible to lie in the x-y plane with the predominant direction of its 
axis lying along the x-axis, and the system parameters are configured to 
perform two series of steps, one for data gathering and the other for 
processing and display. In the data gathering series (120), the blood 
pulse timing is first determined (122). At a preselected time (t.sub.1) 
relative to the blood pulse a signal pulse sequence is begun (124) by 
applying a slice selective field gradient pulse G.sub.z (1) (126). While 
the slice selective gradient is being applied, a 90.degree. RF pulse is 
applied during the period between times t.sub.2 and t.sub.3 (126). The 
effect of the G.sub.z and RF pulses is to excite only those nuclei within 
planar slice 18, the slice of interest. During the period between times 
t.sub.3 and t.sub.4 a gradient G.sub.z (2) of opposite polarity to G.sub.z 
(1) is applied. At time t.sub.4, a refocusing gradient pulse G.sub.x (1) 
is applied (128). At time t.sub.5, a 180.degree. RF signal pulse is 
applied (130) to flip the magnetization vector of the nuclei, causing 
their phases to tend to reconverge. 
In the interval between times t.sub.6 and t.sub.7 a linear phase-encoding 
pulse gradient G.sub.y at a selected level (e.g., level G.sub.y (1) on 
FIG. 5) is applied (132) to encode the nuclei along the y-axis with 
different phases. Immediately thereafter and during the interval between 
times t.sub.7 and t.sub.8, a linear frequency-encoding gradient "read" 
pulse G.sub.x (2) is applied (134) which imparts different frequencies to 
different nuclei along the x-axis. Pulse G.sub.x (2) is timed to occur 
over an interval which spans the spin-echo signal 136, whose center point 
occurs at t.sub.e. 
The time durations of the G.sub.x, G.sub.y, and G.sub.z gradient pulses are 
shown only schematically in FIG. 5. Preferably pulse G.sub.y is about two 
to three times as long as pulse G.sub.z, and pulse G.sub.x is about 5 to 
20 times as long as pulse G.sub.y. Thus for G.sub.z between 1 and 2 
milliseconds, G.sub.y would be between 2 and 3 milliseconds, and G.sub.x 
between 10 and 20 milliseconds. It can be shown (as suggested in the Hahn 
article cited above) that the phase shift in a spin echo experiment 
depends on the square of the time duration of the gradient pulse. The 
contribution of the x-axis gradient pulse and hence of the velocity 
component along the x-axis is thus caused to dominate the total phase 
shift. For example, with G.sub.x lasting 10 times longer than G.sub.y, the 
x-axis velocity will be weighted by a factor of 100 in the phase shift. 
Referring again to FIGS. 4, 5, spin-echo signal 136 is sensed, A-to-D 
converted and sampled, and the samples (representing a time sequence of 
signal amplitude levels) are stored (138). That completes the data 
gathering steps for one time dependent set of signal samples. (The 
identical steps can be repeated several times with the results being 
averaged to improve the signal-to-noise ratio.) 
A family of such signal sample sets is obtained by repeating the steps a 
number of times, each time using a different magnitude for the 
phase-encoding gradient pulse G.sub.y. Thus, between iterations, the level 
of G.sub.y is reset and a delay period is allowed to pass (140) before the 
next iteration begins. The delay period can be selected to synchronize 
each iteration with the heart beat. 
Once the family of signal sample sets are taken and stored, they are 
processed and displayed (150). A complex two-dimensional Fourier 
transformation is performed (152) to give spatially dependent real and 
imaginary components in the frequency domain for a two-dimensional array 
of pixels, and the real component array is displayed (154) as an image 
representative of the velocity profile of the blood over planar slice 18. 
In the image, the intensity represents the real part of the complex image 
datum at each pixel. The midpoint of the greyscale represents the zero 
value. Data with real parts greater than zero (corresponding to phase 
angles with positive cosines) give an image intensity value greater than 
the zero value. Data with negative real parts give image values less than 
the zero value. 
Referring to FIG. 6, in performing the Fourier transformation, the set of 
signal samples which are used for the computation span a time interval 
whose center time (t.sub.s) is different by a small offset amount 160 from 
the center of the spin-echo signal (t.sub.e). The effect of the offset in 
the Fourier transformation process is that the image data is multiplied by 
a phase factor that depends linearly on the frequency-encoded coordinate, 
x. For each sampling interval (e.g., 30 microseconds) within the offset 
amount there is produced 180.degree. of total background phase variation 
across the image in the x direction. The result is a striping of the image 
which improves readability. 
Information about the motion of the blood through vessel 10 during the 
course of each pulse sequence can be shown to be carried through to the 
Fourier transformed spatially dependent data in the form of apparent phase 
shifts, whose magnitudes depend on the velocities of the nuclei. The real 
component of the Fourier transformed data preserves this phase shift 
information, so that a display of the real component will show variations 
in the phase shift in a pattern which will reflect the velocity profile 
within vessel 10. By orienting the length of the blood vessel along the 
same axis (the x-axis) as the frequency encoding gradient pulse, and by 
making the frequency encoding gradient pulse longer than the phase 
encoding pulse, the resulting Fourier transformed image data is made more 
sensitive to velocity along the x-axis than along the y-axis. Offsetting 
the sample interval relative to the center of the spin-echo signal has the 
effect of adding a linearly increasing phase shift in the x-axis direction 
which produces a highly useful striping of the image as explained below. 
Placing the phase-encoding pulse G.sub.y close in time to the 
frequency-encoding pulse G.sub.x reduces any error which might occur as a 
result of the nuclei changing position between the two pulses. 
In one example, a velocity profile image was formed of water flowing 
through a 7/16" inside diameter tube. Flow was constant, gravity driven, 
and calibrated with a Mettler top-loading scale. The water was doped with 
CuSO.sub.4 to have a relaxation constant T.sub.1 of approximately 300 
milliseconds at 20 megaherz. The tube was placed in the magnetic field 
with its axis aligned with the direction of the frequency-encoding 
gradient (i.e., along the x-axis). Because the flow was constant, rather 
than pulsatile, the pulse sequences were not gated to flow pulses, but 
were simply repeated every 300 milliseconds. The frequency encoding 
gradient strength was 6.times.10.sup.3 Hz/cm. The resulting phase shift as 
a function of velocity of the nuclei can be calculated as 8.2 
radians/cm/sec. The ratio depends on the pulse sequence which is 
preferably arranged so that the ratio will produce an image with striping 
which is useful for the flow velocities of interest. Ratios of at least 
0.2 radians/cm/sec. appear to be useful. The magnet was a 1.44 T. (61.5 
MHz) 8 cm superconducting magnet (fabricated by Technicare, Solon, Ohio). 
Referring to FIG. 7, the upper half 170 is a display of the real component 
of the Fourier transformed data resulting from fluid flow in the tube, 
while the lower half 172 resulted from an identical tube with the fluid 
not flowing. 
The lower half image 172 shows a stripe pattern with the stripes orthogonal 
to the x-axis. The stripes represent a background phase which increases 
linearly with distances along the x-axis. Each black or white stripe 
represents a background phase shift of 180.degree. (.pi. radians). In 
upper half image 170, the phase shift due to the motion of the fluid is 
superimposed over the linear background phase shift. Thus the image gives 
an easily seen representation that the velocity along the central axis of 
the tube is higher than along the wall (because at the central axis the 
phase shift per unit length along the x-axis is greater). Further, because 
each stripe represents 180.degree. of phase shift, it is possible to 
measure the difference between the phases at the wall and at the central 
axis at one position (174) along the x-axis by counting the number of 
stripes which must be traversed along the central x-axis in order to reach 
the stripe which begins at line 174 at the wall of the tube. Here there 
are 3 stripes between points 176, 178, which amounts to a 3.pi. radians 
phase shift which translates to a maximum flow velocity of 1.2 cm/sec 
(3.pi. radians= 1.2 cm/sec) or an average flow velocity of 8.2 
radians/cm/sec 0.6 cm/sec, which is within 15% of the mechanically 
calibrated average velocity. 
The display format enables direct inference from striping of data phase 
with a precision of .+-.90.degree., which is acceptable if small compared 
with typical phase shifts being studied. By using pulse sequences whose 
magnitude and duration produce relatively high phase shift/velocity 
ratios, phase shifts of many times 360.degree. can be obtained. 
Referring to FIG. 8, in another example, the phase image produced by 
non-moving fluid in a bifurcating tube is shown in the lower left (180), 
and produces a set of parallel stripes representing the background phase 
offset. When the fluid is moving (182), the stripe pattern is shifted to 
reflect the velocity profile of the fluid. An image of fluid moving 
through tubes which recombine is shown in the upper right of FIG. 8 (184). 
In FIG. 8, the tube diameter is 1/2", and flow rate is 100 cc/min. The 
maximum phase shift discernible on the images is 10.pi. radians 
corresponding to a maximum velocity of 3.83 cm/sec. The maximum expect 
velocity (based on fluid mechanics) is 4.0 cm/sec., within about 2% of the 
measured figure. 
Referring to FIG. 9 in another example, images of a 3/16" inside diameter 
tube with a 3/32 inch stenosis (constriction) are shown for non-moving 
(190) and moving fluid (192). 
In the images of moving fluid, velocity is inferred from the displacement 
of the stripes left or right, with the slopes of the stripes representing 
changes in flow velocity. Phase stripes which show forward concavity thus 
do not imply retrograde flow, but rather the existence of higher shear 
rates near the central axis compared with the tube wall (as in laminar 
flow). 
Pixels where the phase becomes indistinct correspond to points where the 
velocity gradient is high enough that different velocity values coexist 
within one pixel. Interference of the resulting phases causes loss of 
signal, an effect which can be reduced by increasing the spatial 
resolution. Alternately, as explained below, such interference can be 
exploited to produce high contrast projection images of blood flow. 
Introduction of a background phase offset into the imaging phase enhances 
the readability of phase shifts for four reasons. First, phase shifts can 
be calculated in two ways: stripe count or X displacement. Accuracy and 
precision are improved by this redundancy. Second, two-dimensional Fourier 
transform images have better resolution in X than in Y which gives the X 
displacement method the advantage. Third, background offset exposes any 
imperfections in the background phase, enabling correction by the reader. 
Fourth, in consequence of the first three advantages, phase offset enables 
the implementation of pulse sequences with higher characteristic 
velocity/phase shift ratios, reducing the importance of any residual phase 
ambiguity. 
The system can also be used to generate so-called projection images, in 
which data from a number of stacked planar slices are effectively 
accumulated into one array. For example, a projection image of FIG. 1 
would represent not only slice 18 but slices above and below it along the 
z-axis. 
Referring to FIG. 10, the pulse sequence for taking a set of samples for a 
projection image begins at time q.sub.1 with the start of a long 
frequency-encoding gradient pulse 200. No slice-selective gradient pulse 
G.sub.z is used since the image is not meant to be slice-selective. At 
time q.sub.2 (while the G.sub.x pulse continues) a 90.degree. RF pulse is 
imposed, followed by a 180.degree. pulse at time q.sub.3, and a phase 
encoding gradient pulse 202 beginning at time q.sub.4. The spin-echo 
signal is centered at time q.sub.e. Gradient pulse 200 ends at time 
q.sub.5, after which a delay occurs before the next pulse sequence begins. 
Thus, the projection images are obtained without either slice selection or 
z-axis encoding. 
The Fourier transformation of the sets of data samples is performed without 
imposing the background phase offset used for the slice-selective images. 
Instead, the data component 90.degree. away from the phase of the 
stationary nuclei in the object being imaged is the one used to form the 
image. This in effect suppresses the contribution of the stationary nuclei 
to the final image, while emphasizing the contribution of the moving 
nuclei. A greater than 90% reduction in stationary nuclei signal intensity 
has been achieved, permitting the imaging of flow velocities greater than 
10 cm/sec with vessel diameter to total diameter ratios greater than 1/20. 
Referring to FIG. 11, the upper portion 210 shows a projection image of a 
bifurcating and recombining tube through which fluid is flowing at 300 
cc/min. The 1/2" inside diameter tubing lies within an 
8".times.10".times.12" cavity of stationary water. The measured maximum 
phase shift is 5.pi. radians which implies a maximum velocity of 
V.sub.max, =6.44 cm/sec. The expected value of V.sub.max is 12.0 cm/sec, 
illustrating that projection imaging underestimates velocities. In this 
case, however, it would be more reasonable to expect projection images to 
represent the average velocity, V.sub.avg, rather than V.sub.max since the 
entire tube diameter contributes to the observed phase shift. In this 
experiment V.sub.avg =6.0 cm/sec, close to the observed velocity value. 
FIG. 12 is another example of projection imaging, this time of a rotating 
disc, with the axis of rotation aligned with the z-axis. The disc is 
composed of water-saturated towels inside a plastic container 20 cm in 
diameter and 1.5 cm thick. The disk is rotating at 30 rpm corresponding to 
a maximum tangential velocity of 10.pi. cm/sec. A rigid body rotating in 
the x-y plane at frequency w has the property that at any point (x,y), the 
velocity V(x,y)=2.pi./w/ (y,-x). Therefore the x velocity component is 
proportional to y, and points of equal phase shift lie on horizontal lines 
(constant y) because they sustain equal x velocities. The pulse sequence 
was repeated every 300 msec, t.sub.e =10.0 msec, the frequency encoding 
gradient was G.sub.x =3.times.10.sup.3 Hz/cm and the calculated ratio of 
phase to velocity was 
EQU P(t.sub.e)/V.sub.x =0.28 radians/cm/sec. 
At 30 cm/sec, this corresponds to a total phase shift of 3.pi. radians 
which is in reasonable accord with the experiment. This demonstrates that 
high velocity can yield good signal intensity without spatial distortion. 
Projection imaging is highly efficient, enabling three-dimensional volume 
to be surveyed in times characteristic of two-dimensional imaging 
experiments. 
In another technique for generating projection images, two distinct sets of 
data are accumulated, then subtracted to form a resulting image. The two 
data sets are derived in such a manner that, for static portions of 
tissue, the data are identical and cancel, while for moving portions (such 
as blood) the two data sets differ. One data set is taken during systole, 
the other during diastole. In the resulting image the blood, and 
implicitly the vessels in which it flows, are seen clearly, while the 
static tissue is suppressed. 
Referring to FIG. 13, the pulse sequences (a two-dimensional Fourier 
transform sequence) begins at a time s.sub.1 when the QRS complex occurs 
in the electrocardiogram (ECG). At time s.sub.2, after an appropriate gate 
delay whose duration is based (in a manner described below) on whether the 
data is being taken for systole or diastole, a 90.degree. RF pulse is 
imposed. Thereafter, in the interval between s.sub.3 and s.sub.4, a 
compensating gradient pulse G.sub.x is applied simultaneously with a phase 
encoding pulse G.sub.y at a selected one of 256 different levels. Next, at 
time s.sub.5 (i.e., 4.5 milliseconds after s.sub.2 or half the 9 
millisecond interval between s.sub.2 and the echo signal--s.sub.e) a 
180.degree. RF pulse is imposed. The spin echo signal is spanned by a 
frequency encoding readout gradient pulse G.sub.x (equivalent to 1,000 
Hz/cm) in the interval between S.sub.6 and S.sub.7. The spin echo signal 
is centered at time s.sub.e. The readout period occurs between s.sub.8 and 
s.sub.9. 
The combined effect of the compensating G.sub.x pulse and the readout 
G.sub.x pulse, which straddle the 180.degree. RF pulse is to cause no net 
phase component to be imparted to static protons but to cause a net phase 
component to be imparted to each moving proton. The net component depends 
on the fact that because the proton moves to a new x position in the time 
between the occurrences of the two G.sub.x pulses, it is subjected to 
different magnitudes of phase shifting by the two G.sub.x pulses; those 
magnitudes (which in static protons would be equal and therefore cancel) 
do not cancel. 
Following the readout period with respect to one pulse sequence, a new 
pulse sequence (using a different value for the G.sub.y gradient) is begun 
at time S.sub.10 upon the occurrence of the next QRS complex. A succession 
of 256 pulse sequences one for each different value of G.sub.y is used 
with a gate delay appropriate to diastole to obtain a first data set. A 
second succession of 256 pulse sequences is used with a gate delay 
appropriate to systole to obtain a second data set. The pulse sequences 
used to obtain the two data sets are identical. Referring to FIG. 27, the 
respective gating delays for systole and diastole are both provided by 
gating circuitry 112 to a switch 113 that passes the appropriate delay to 
sequencer 100. Referring to FIG. 28, processor 70 performs a 
two-dimensional complex Fourier analysis (in element 108) of the resulting 
two arrays, producing two corresponding images of the same subject 
respectively gated to diastole and systole 109, 111. The images are 
subtracted from each other (in element 115) to obtain a clear resulting 
blood flow image of high contrast and high resolution in which the static 
tissue is suppressed. In practice, it has been found necessary to weight 
the two data sets before subtraction in order to maximize the cancellation 
of the static proton signals and hence the image contrast. The weighting 
has been done empirically by applying different weights to a background 
region of the data sets until the minimum background image intensity is 
obtained. 
The pulse sequence of FIG. 13 is designed to generate velocity-dependent 
proton phase shifts in the blood of 1 cycle per meter per second. This is 
accomplished, in particular, by arranging the echo time (i.e., the 
interval between s.sub.2 and s.sub.e) to be no greater than 15 
milliseconds, preferably no greater than 10 milliseconds. This value 
results in the relative preservation of the blood proton signal for 
diastolic flow velocities (which are, e.g., typically less than 0.1 meters 
per second) because the velocity imposed phase shifts are quite small and 
the blood proton phases reinforce each other in the projected image. 
Conversely, the value of 1 cycle per meter per second produces a relative 
loss of the blood proton signal for systolic flow velocities (which are 
typically between 0.5 and 1.5 meter per second) because the velocity 
imposed phase shifts are large enough to cause a randomization of and 
hence cancellation of phases in the projection image. When the systole and 
diastole images are subtracted, the static tissue phases (which are 
identical in the two images) cancel while the blood phase in diastole 
remains highly visible as an indicator of blood flow. 
Referring to FIG. 14, each projected image lies in an x-y plane 200 which 
includes a grid of 256 by 256 pixels (e.g., pixel 202, shown out of scale) 
in an area of 50 cm by 50 cm. Each pixel represents a projection of the 
proton signal from all protons in a voxel (volume element) 204 located at 
the same x and y coordinates as the resulting pixel 202 but spanning all z 
coordinate values in a sample that includes tissue 206, an artery 208, and 
moving blood 210 within the artery. Voxel 204 thus includes portions 212, 
214, 216, 218, 220, that lie respectively in tissue, vessel, blood, 
vessel, and tissue. 
Referring to FIG. 15, in each projective pixel 202, the resulting proton 
signal 230 has two constituents: a larger coherent part originating in the 
stationary background material 212, 214, 218, 220, and a smaller part 
originating in the blood protons 216 whose coherence depends on the flow 
velocity. At diastolic flow velocities the imaging pulse sequence produces 
phase shifts smaller than 0.1 cycle. In that case, the blood proton signal 
behaves coherently (as represented by the phase arrows in section 216 of 
FIG. 15, which point in generally the same direction) and adds to the 
background signal to form an additive resulting proton signal 230. 
Referring to FIG. 16, in systole, peak velocities generate phase shifts of 
0.5 to 1.5 cycles in moving blood protons 216. Each voxel 204, however, 
intersects the blood in vessel 208 along a chord of points (section 216 in 
FIG. 16) and so all velocities between zero and the maximum are sampled by 
partial volume averaging. This range of velocities yields a blood signal 
which behaves incoherently (as represented by the randomly directed phase 
arrows in segment 216 of FIG. 16) and makes little or no contribution to 
the net proton signal 232 for the pixel 202. Systole signal 232 is thus 
considerably smaller than diastole signal 230 (FIG. 15). 
This subtractive projective imaging technique is performed in a 0.6 T (25 
MHz, proton) superconducting imaging system (Technicare Corporation, 
Solon, Ohio). The brief duration of the data acquisition window (6 msec 
between S.sub.8 and S.sub.9 in FIG. 13) requires a comparatively large 
readout gradient G.sub.x of 0.25 Gauss per centimeter (corresponding to 1 
kHz per centimeter, proton) and the use of a correspondingly broadband 
time-domain signal filter (50 kHz). To produce the two images for diastole 
and systole, data acquisition times for both data sets average 8 minutes 
(512 cardiac cycles in subjects with a normal pulse). 
Artifacts caused by patient movement between the two data sets could be 
reduced by interleaving the systole and diastole pulse sequences, but the 
time between successive pulse sequences would have to be increased to 
accommodate the time required to switch the gating, thus increasing the 
aggregate time required for data acquisition. Respiratory artifacts could 
be removed by timing the pulse sequences to occur at the same times 
relative to the respiration cycles. 
For example, FIGS. 17, 18, and 19 show respectively the diastolic, 
systolic, and resulting subtraction flow images of a human chest in a 
45.degree. right oblique projection in which the x coordinate (FIG. 14) 
appears vertically. The diastolic and systolic gate delays (S.sub.1 to 
S.sub.2 in FIG. 13) were respectively 10 and 150 milliseconds. Blood 
vessels seen in the flow image (FIG. 19) are labeled aa (ascending aorta), 
da (descending aorta), pa (pulmonary artery), pas (right pulmonary 
segmental branches), rca and lca (right and left carotid arteries), rsc 
and lsc (right and left subclavian arteries). 
In another example, FIGS. 20, 21, 22, 23 show systolic (FIGS. 20, 22) and 
resulting flow images (FIGS. 21, 23) of thighs (FIGS. 20, 21) and knees 
(FIGS. 22, 23) of a human subject, projective to the coronal plane. The x 
coordinate is vertical. Diastolic gate delays of 10 msec were used for 
both knees and thighs; systolic delays of 250 and 300 msec were used at 
the thighs and knees respectively in this individual (height 180 cm). 
Arrows mark the superficial (sf) and deep (df) femoral arteries of the 
thigh and the popliteal (pop), anterior tibial (ta), posterior tibial 
(tp), and peroneal (per) of the knee and calf. 
In another example, FIGS. 24, 25, 26 show x-ray (FIG. 24) and NMR systolic 
and flow (FIGS. 25, 26) images of atherosclerotic occlusions of the 
superficial femoral arteries, using gate delays of 10 (diastolic) and 300 
(systolic) milliseconds. Arterial segments are marked as for FIGS. 17, 18, 
19. Proximal and distal points of occluded segments are marked "0. prox" 
and "0. dist" (FIG. 24). The appearance of the popliteal arteries 
reconstituted by collateral flow images implies they are pulsatile (as was 
confirmed by Doppler ultrasound examination). The poor appearance of the 
right proximal superficial femoral artery is consistent with 
angiographically proven poor runoff in this vessel. 
In these examples, one image parameter was tailored for each application: 
the systolic gate delay. Apart from individuals with ventricular 
dysrhythmias, the QRS complex itself coincides with arterial diastole so 
the diastole gating delay is always set at 10 milliseconds regardless of 
the location of the artery within the body. However, the arrival time of 
peak systolic flow is variable. In normal individuals arrival times 
increase with distance from the heart. Disease processes may either retard 
the pulse wave (e.g., aneurysm, occlusion) or accelerate it (e.g., 
nonocclusive atherosclerosis). The systolic gate delay was selected 
empirically in each case by performing between 1 and 4 brief (1 minute) 
low resolution (64.times.256 pixels) localization images. Typically the 
gate delay is between 100 and 300 milliseconds. Best results were obtained 
by exploring the likely range of gate delays in 50 msec increments. 
Several factors will affect the contrast of the subtraction image, 
including the following. First, because the fraction of velocity which 
produces phase shifts (and hence contrast) varies as the cosine of the 
angle between the direction of blood flow and the x axis, vessels oriented 
at large angles to the x axis may be unobserved in the flow image due to 
an undiminished systolic signal. Two resulting flow images could be 
acquired, separated by a 90.degree. rotation in the x-y plane. Each vessel 
will have a satisfactory orientation in at least one of these images. 
Second, flow contrast may be undercut by substantial diastolic flow 
velocities (greater than 0.1 meter per second) which reduce the diastolic 
blood signal. Doppler ultrasound data have shown that such diastolic flows 
occur at certain anatomic locations, notably in the arterial supplies of 
the brain and the visceral organs. Maximum contrast could be recouped by a 
diastolic acquisition with reduced phase shift. Third, a portion of the 
arterial blood protons are replaced during each interpulse interval 
(between s.sub.7 and s.sub.10 in FIG. 13) by unsaturated protons formerly 
outside the RF coil (proton refreshment). The saturated proton signal is 
less intense than the unsaturated proton signal by a factor [1-exp(s.sub.7 
-s.sub.10)],/T1.sub.blood, (where T1 denotes the longitudinal relaxation 
time). Affected arterial segments will have proportionally enhanced 
intensity in the flow image. Whether gating is systolic or diastolic, the 
interval between successive pulse sequences represents one cardiac cycle. 
Therefore, the physical locations in the vessel that are subject to proton 
replacement are well-defined. Image subtraction will subtract saturated 
protons in one data set from saturated protons in the other data set, and 
will likewise subtract unsaturated from unsaturated. Relative to image 
subtraction, proton refreshment thus resembles a static variable such as 
proton density more than a dynamic variable such as phase contrast. 
Vascular imaging in the projective format efficient. Neither tomography nor 
existing 3-dimensional imaging techniques can present such detailed 
anatomy from such large territories in so compact and accessible a form. 
Projective imaging also yields purely technical benefits. Magnetic 
resonance imaging times grow exponentially with the dimensionality of the 
experiment. Though projective images are sensitive throughout a 
3-dimensional volume, their data acquisition and processing requirements 
are those of 2-dimensional imaging. A related economy of this method is 
its reliance on completely conventional imaging pulse sequences and 
reconstruction. From a clinical standpoint, the non-invasiveness of this 
method may allow its application outside the compass of current 
indications for angiography. An example might be the periodic assessment 
of arterial bypass grafts. Pulsatility, the ultimate source of contrast, 
tells something more specific about functional capability than simple 
patency. 
Other embodiments are within the following claims. Veins in which blood 
flow is pulsatile (or any vessel containing pulsatile flow of a fluid) can 
also be imaged.