MRI compensated for spurious NMR frequency/phase shifts caused by spurious changes in magnetic fields during NMR data measurement processes

At least one extra NMR measurement cycle is performed without any imposed magnetic gradients during readout and recordation of the NMR RF response. Calibration data derived from this extra measurement cycle or cycles can be used for resetting the RF transmitter frequency and/or for phase shifting other conventionally acquired NMR RF response data to compensate for spurious changes in magnetic fields experienced during the NMR data measuring processes. Some such spurious fields may be due to drifting of the nominally static magnetic field. Another source of spurious fields are due to remnant eddy currents induced in surrounding conductive structures by magnetic gradient pulses employed prior to the occurrence of the NMR RF response signal. Special procedures can be employed to permit the compensation data itself to be substantially unaffected by relatively static inhomogeneities in the magnetic field and/or by differences in NMR spectra of fat and water types of nuclei in imaged volumes containing both.

This invention is generally related to magnetic resonance imaging (MRI) 
utilizing nuclear magnetic resonance (NMR) phenomena. It is more 
particularly directed to apparatus and method for practicing MRI which 
provides compensation for spurious NMR frequency/phase shifts caused by 
spurious changes in magnetic fields during NMR data measurement processes. 
This application is related to the commonly assigned concurrently filed 
application Ser. No. 07/181,386 (refiled as Ser. No. 07/283,059) to Yao 
entitled MRI USING ASYMMETRIC RF NUTATION PULSES AND/OR ASYMMETRIC 
SYNTHESIS OF COMPLEX CONJUGATE SE DATA TO REDUCE TE AND T2 DECAY OF NMR 
SPIN ECHO RESPONSES in that the invention therein described may be 
conveniently employed concurrently with the present invention. 
MRI is now a widely accepted and commercially available technique for 
obtaining digitized visual images representing the internal structure of 
objects (such as the human body) having substantial populations of nuclei 
which are susceptible to NMR phenomena. In general, the MRI process 
depends upon the fact that the NMR frequency of a given nucleus is 
directly proportional to the magnetic field superimposed at the location 
of that nucleus. Accordingly, by arranging to have a known spatial 
distribution of magnetic fields (typically in a predetermined sequence) 
and by suitably analyzing the resulting frequency and phase of NMR RF 
responses (e.g., through multi-dimensional Fourier Transformation 
processes), it is possible to deduce a map or image of relative NMR 
responses as a function of the location of incremental volume elements 
(voxels) in space. By an ordered visual display of this data in a suitable 
raster scan on a CRT, a visual representation of the spatial distribution 
of NMR nuclei across a cross section of an object under examination may be 
produced (e.g., for study by a trained physician). 
Typically, a nominally static magnetic field is assumed to be homogeneous 
within the cross section to be imaged. In addition, typical MRI systems 
also superimpose magnetic gradient fields of the same orientation but with 
intensity gradients which are assumed to vary linearly in predetermined 
directions (e.g., along mutually orthogonal x,y and z axes) while being 
constant and homogeneous in all other dimensions. 
Unfortunately, the "real world" does not always conform exactly to these 
assumptions. In spite of several techniques known to and used by those in 
the art to substantially achieve these assumptions, there are inevitably 
small spurious changes which occur in the magnetic fields during NMR data 
measurement processes. For example, the nominal strength of the static 
field may drift with respect to time (less of a problem with cyrogenic 
super-conducting magnets than with permanent or resistive magnet 
embodiments). In addition, rapid imposition of a sequence of magnetic 
gradient fields produce eddy curren(s in nearby conductive members (e.g., 
the typically conductive containers for super-conducting magnet coils, the 
metal of a permanent magnet, etc.). The magnetic field produced by such 
eddy currents is, of course, directed so as to oppose the magnetic field 
which induced the eddy currents. To compensate for this effect, one 
conventional approach is to initially overdrive the magnetic gradient 
pulse so as to produce thedesired net magnetic gradient within the imaged 
volume. 
Unfortunately, these eddy currents do not instantaneously vanish once the 
magnetic gradient pulses are switched off. Rather they more gradually die 
out and, as a result, there may be remnant magnetic fields still present 
(e.g., when the NMR RF response signal occurs . Typically, the magnetic 
gradient pulses used to achieve slice/volume selective NMR responses are 
of substantially greater intensity/duration than other magnetic gradient 
pulses and, accordingly, are often the principal source of spurious 
magnetic fields due to induced eddy currents. 
The spurious changes in magnetic fields due to drift of the nominal static 
magnetic field are of much less significance with super-conducting 
magnets. For example, in low field permanent magnet MRI systems, the field 
drift over a scan sequence may be equivalent to a 100 Hertz NMR frequency 
shift which may cause partial dislocations of about 10% along the z-azis 
and 5% along the x-axis. However, for medium or high field cryogenic 
magnet MRI systems, the field drift over an entire day may be only 20 
Hertz or so. Accordingly, this problem is of more concern with lower 
strength permanent magnet MRI systems. 
On the other hand, spurious changes in magnetic fields caused by remnant 
eddy currents have been and remain more of a problem with super-conducting 
magnet structures than with permanent magnets (e.g., because more good 
electrically conductive material is generally present in cyrogenic 
super-conducting apparatus). Phase shift errors of as much as about 200 
degrees or so have been observed at the beginning of the typical spin echo 
sampling window time (which errors gradually decay during the sampling 
window as the remnant eddy currents die out). 
Thus, although perhaps for several different reasons, an ability to 
compensate for spurious changes in magnetic fields should be of benefit in 
all types of MRI systems. 
Where the FID is recalled (i.e., by reversed gradient pulses rather than 
the production of a true spin echo with a 180.degree. RF pulse) the phase 
errors caused by spurious fields are of enhanced importance. This is 
because the recalled FID immediately follows both the rephasing slice 
gradient and a dephasing readout gradient. It is inherently more sensitive 
to static or transient inhomogeneity of the field since there is no 
180.degree. RF pulse. 
In some super-conducting magnets there is also another effect. The gradient 
pulsing may cause a dynamic or transient change of the field strength. In 
some way (perhaps by exerting forces on the magnet winding that moves it 
around in the cryostat) the pulsing gradients cause the field to change to 
a new value while the scan is running. Once the gradient pulses stop, the 
field returns to the pre existing quiescent value. This dynamic change is 
repeatable for the same gradient pulses and is different when different 
shape, duration or direction pulses are used. Thus, a phase correction 
would be of benefit when scans of the same orientation are used where some 
gradient parameters change and one wants the images from the different 
scans to register with one another. 
The fact that eddy currents can produce phase problems in MRI is known in 
the prior art (see, for example, "NMR Velocity-Selective Excitation 
Composites for Flow and Motion Imaging and Suppression of Static Tissue 
Signal" by Moran et al, IEEE Trans. In Med. Imaging, Vol MI-6, No. 2, June 
1987, pp 141-147 at p 142). However, earlier proposed solutions, to the 
extent there are any, appear to focus primarily upon techniques for better 
achieving the assumed ideal magnetic field distributions. For example, 
U.S. Pat. No. 4,300,096 - Harrison et al is directed towards techniques 
for closed loop control systems aimed at achieving matched shape and 
magnitude of various magnetic gradient pulses in an MRI system. U.S. Pat. 
Nos. 3,495,162 - Nelson and 3,496,454 - Nelson also teach prior NMR 
spectrometer systems which attempt to achieve improved homogeneity and 
constant magnetic field strength (e.g., by tracking a variable reference 
NMR frequency from a special NMR sample via a closed loop control 
channel). 
It should also be noted that spurious phase errors in captured MRI data are 
especially bothersome where data conjugation techniques are utilized to 
synthesize data (based upon assumed symmetry properties) thereby 
materially reducing the overall time required for capturing MRI data 
sufficient to generate an MRI image. Here, if there are spurious phase 
errors in the actually measured and collected data set, then the 
synthesized data set generated therefrom (e.g., by conjugation processes) 
will necessarily also include at least these phase errors. And since the 
phase errors do not exhibit the assumed theoretical symmetry underlying 
the synthesized conjugation process, the resulting synthesized data will 
be even more severely corrupted than would normally be expected due to 
spurious changes in magnetic fields. Accordingly, a technique for avoiding 
spurious frequency/phase errors would be of even greater advantage in MRI 
processes utilizing conjugate synthesized data (e.g., as described in U.S. 
Pat. No. 4,728,893 - Feinberg or in Margosian et al, "Faster MR Imaging 
Imaging With Half The Data", Health Care Instrumentation, Vol. 1, pages 
195-197). 
Those in the art will recognize that a "square" spectrum of RF frequencies 
(obtained by using a sinc-shaped RF pulse envelope in the time domain) may 
be used in conjunction with a slice-selective magnetic gradient to elicit 
RF spin echoes from a slice-volume of NMR nuclei. However, if there is no 
applied gradient during the SE readout, then the NMR frequency spectrum 
will immediately shrink to reflect the only remaining static magnetic 
field. If it is truly homogeneous, then it should have a substantially 
single frequency spectrum. In fact, this type of experiment has long been 
used to initially calibrate the MRI frequency versus slice location data 
initially before the first MRI scan of the day is performed. In this 
manner, the center frequency of the center slice position has been 
determined and used for thereafter establishing the RF transmitter center 
frequency. 
Although techniques for better achieving the assumed ideal magnetic field 
distribution will continue to be of primary importance, we have discovered 
new techniques which permit one to compensate in other ways for spurious 
NMR frequency/phase shifts caused by spurious changes in magnetic fields 
during NMR data measurement processes in MRI. 
In at least some typical MRI systems now commercially available (e.g., 
those available from Diasonics Inc), if an NMR spin echo response is 
recorded in the absence of any magnetic gradient pulse, then ideally the 
response will comprise only a very narrow band of frequencies 
(corresponding to the frequencies in the selected slice) and all of the 
sampled time domain responses will have the same zero relative phase. 
Accordingly, by actually taking such data in one or more extra 
"calibration" measurement cycle(s), and by noting the extent to which the 
detected frequency spectrum differs from the expected, and the extent to 
which the relative measured phase of the sampled RF signals differs from 
zero, one can derive compensation factors (a) to be applied to the already 
recorded data and/or (b) to reset the RF transmitter center frequency for 
subsequent NMR measurement cycles. Compensation to already collected spin 
echo data taken within a given measurement cycle can be made by 
appropriately phase shifting the data (either in the time domain or the 
frequency domain). 
Unfortunately, the derived compensation factors may also be contaminated 
with errors caused by relatively static field inhomogeneities and/or by 
interference between the slightly different frequency signals emanating 
from fat-like and water-like nuclei in the imaged volume. Thus, especially 
where more exacting phase corrections are to be made (e.g. in compensation 
for eddy current effects), special inversion recovery and/or other changes 
are made in the NMR measurement cycles used to accumulate the compensation 
data so as to reduce or cancel such potential error sources.

The novel signal processing and control procedures utilized in the 
exemplary embodiment typically can be achieved by suitable alteration of 
stored controlling computer programs in existing MRI apparatus. As one 
example of such typical apparatus, the block diagram of FIG. 1 depicts the 
general architecture which may be employed in such a system. 
Typically, a human or animal subject (or other object) 10 is inserted along 
the z-axis of a static cryogenic magnet which establishes a substantially 
uniform magnetic field directed along the z-axis within the portion of the 
object of interest. Gradients are then imposed within this z-axis directed 
magnetic field along the x,y or z axes by a set of x,y,z gradient 
amplifiers and coils 14. NMR RF signals are transmitted into the body 10 
and NMR RF responses are received from the body 10 via RF coils 16 
connected by a conventional transmit/receive switch 18 to an RF 
transmitter 20 and RF receiver 22. 
All of the prior mentioned elements may be controlled, for example, by a 
control computer 24 which conventionally communicates with a data 
acquisition and display computer 26. The latter computer 26 may also 
receive NMR RF responses via an analog-to-digital converter 28. A CRT 
display and keyboard unit 30 is typically also associated with the data 
acquisition and display computer 26. 
As will be apparent to those in the art, such an arrangement may be 
utilized so as to generate desired sequences of magnetic gradient pulses 
and NMR RF pulses and to measure NMR RF responses in accordance with 
stored computer programs. As depicted in FIG. 1, the MRI system of this 
invention will typically include RAM, ROM and/or other stored program 
media adapted (in accordance with the following descriptions) so as to 
generate phase encoded spin echoes during each of multiple measurement 
cycles within each of the possible succession of MRI data gathering scans 
(sometimes called "studies"). And to process the resulting MRI data into a 
final high resolution NMR image. 
FIG. 2 depicts a typical prior art data acquisition sequence. Here, a 
single MRI data gathering scan or "study" involves a number N (e.g., 128 
or 256) of successive data gathering cycles. In fact, if, as depicted in 
FIG. 2, a multi-slice scan is involved, then each of the N events 
comprising a single scan or "study" may actually comprise M single slice 
MRI data gathering cycles. In any event, for a given single slice data 
gathering cycle p, a slice selection z-axis gradient pulse (and associated 
phase correction pulse) may be employed to selectively address a 
transmitted 90.degree. RF nutation pulse into a slice volume centered 
about the center frequency of the transmitted RF signal and of sufficient 
magnitude and duration so as to nutate a substantial population of nuclei 
within the selected slice-volume by substantially 90.degree. . Thereafter, 
a y-axis gradient pulse is employed (of magnitude .phi..sub.Q for this 
particular cycle and varying between maximum magnitudes of both polarities 
over the N data gathering cycles for that particular slice) to phase 
encode the signal. After a predetermined elapsed time a 180.degree. RF NMR 
nutation pulse will be transmitted to selectively excite the same slice 
volume (via application of the appropriate z-axis gradient) at a time tau. 
In accordance with the "rule of equal times," a true spin echo signal SE 
evolves, reaching a peak after a further elapsed time tau. During the 
recordation of the RF NMR response signal (where amplitude and phase of RF 
is measured at successive sample points), a readout x-axis magnetic 
gradient is employed so as to provide spatial frequency encoding in the 
x-axis dimension. Additional spin echo responses can also be elicited by 
the use of additional 180.degree. nutation pulses or suitable other 
techniques --albeit they will be of decayed amplitude due to the NMR T2 
decay. 
The time-to-echo TE is depicted in FIG. 2 as is the repetition time TR. As 
just explained, TE interacts with the T2 NMR exponential decay parameter 
to reduce signal amplitude. The TR interacts with the T1 NMR exponential 
recovery parameter so that one generally has to wait for substantial 
relaxation of previously excited nuclei before the next measurement cycle 
of the same volume starts. 
It will be noted that during each RF excitation pulse, there is a slice 
selection G.sub.z magnetic gradient pulse switched "on" so as to 
selectively excite only the desired "slice" or "planar volume" (e.g., a 
slice of given relatively small thickness such as 5 or 10 millimeters 
through the object being imaged). During each resulting spin echo NMR RF 
response, x-axis phase encoding is achieved by applying an x-axis magnetic 
gradient during the readout procedure (typically each spin echo pulse is 
sampled every 30 microseconds or so with digitized sample point, complex 
valued, data being stored for later signal processing). 
As depicted in FIG. 2, it will be understood that, in practice, the number 
of measurement cycles typically is equal to the number of desired lines of 
resolution along the y-axis in the final image (assuming that there is no 
data synthesis via conjugation processing). After a measurement cycle is 
terminated with respect to a given "slice," it is allowed to relax for a 
TR interval (usually on the order of the relaxation time T1) while other 
"slices" are similarly addressed so as to obtain their spin echo 
responses. Typically, on the order of hundreds of such measurement cycles 
are utilized so as to obtain enough data to provide hundreds of lines of 
resolution along the y-axis dimension. A sequence of N such y-axis phase 
encoded spin echo signals is then typically subjected to a two-dimensional 
Fourier Transformation process so as to result in an N.times.N array of 
pixel values for a resulting NMR image in a manner that is by now well 
understood in the art. 
Such prior art MRI techniques are based on the assumption (a) that the 
static magnetic field does not drift over the several minutes required for 
an entire MRI data gathering scan or "study" (or perhaps even a sequence 
of plural such scans) and (b) that the magnetic fields induced by remnant 
eddy currents do not introduce any phase error in the recorded spin echo 
signals (an assumption which is especially important where synthesized 
complex conjugate data are utilized). 
In a super conducting magnet, static field drift compensation is probably 
not as important as compensation for spurious perturbations in recorded 
signal phase caused by remnant eddy currents induced in the cyrostat by 
pulsed magnetic gradients occurring prior to the recorded NMR response. 
The most significant remnant eddy current disturbance is caused by the 
strongest magnetic gradient pulses --namely, the slice selective ones. 
On the other hand, in the typically lower field, permanent magnet MRI 
systems, while compensation for remnant eddy current effects can still be 
useful, compensation for drifting static magnetic field intensity (e.g., 
with respect to temperature) is probably more important. 
The relatively weaker phase encoding G.sub.y and readout G.sub.x magnetic 
gradients produce less of an adverse phase perturbation effect and, if 
employed would tend to mask the phase and frequency of a desired 
compensation signal due to their normal phase and frequency encoding 
roles. Accordingly, in a preferred embodiment, at least one (preferably an 
extra first and an extra last) measurement cycle or "template" is provided 
in which at least the readout G.sub.x gradient (and preferably also the 
G.sub.y phase encoding gradient) is left "off" at least during the actual 
readout and recordation of the NMR response signal (and preferably at all 
times). One possible such calibration cycle is schematically depicted in 
FIG. 3. 
In this way, a compensation or "calibration" NMR response "template" is 
obtained from each slice and recorded under influence only of the actual 
then-existing static field intensity and remnant fields related to eddy 
currents caused by the slice selective G.sub.z magnetic gradient pulses. 
If the static field strength has not changed from the initial set-up or 
calibration arrangement, and if there are no adverse fields due to remnant 
eddy currents, then the measured NMR response should have an expected 
frequency spectrum and phase. To the extent that this differs from the 
expected value, then correction or compensation data may be generated for 
use either in subsequent data gathering cycles and/or in the processing of 
already recorded NMR data so as to compensate for any spurious changes in 
magnetic fields during the just-completed earlier data measurement cycles. 
When a drifting field strength is the main concern, the frequency spectrum 
of the spin echo during the "calibration" cycle where no phase encoding or 
readout gradient is employed (e.g., as in FIG. 3), is a measurement of the 
actual quiescent field strength. For example, the NMR spin echo elicited 
from a slice volume at the nominal center of the magnet effectively 
measures the field at this location. Spin echoes elicited during the 
"calibration" cycle from other slices located at distances from the center 
of the magnet will also vary in frequency due to inhomogeneities in the 
static magnetic field and due to fields caused by remnant eddy currents. 
If, for whatever reason, the actual magnetic field distribution changes 
from the assumed ideal during the course of a scan sequence, then the 
frequency of the spin echo in the first extra "calibration" measurement 
cycle (e.g., see FIG. 3) will be different from that of the last extra 
calibration measurement cycle. If the frequency drift is assumed to be 
linear over the entire scan interval, then a corresponding incremental 
correction (proportional to the elapsed time from the beginning of the 
scan sequence and to the overall incremental change during the entire scan 
sequence) may be applied to all of the recorded data during its 
processing. Furthermore, if successive or multiple scans or "studies" are 
employed, a slow frequency drift (caused by field drift) will cause a 
given imaged object to effectively slide across the image space with 
respect to time such that given slice volumes effectively move with 
respect to a fixed spatial coordinate system as the NMR frequency changes 
in response to spurious magnetic field changes. This can be corrected by 
periodically resetting the center frequency of the RF transmitter during a 
sequence of scans. 
The following equation expresses the resonant NMR frequency f in terms of 
nominal static magnetic field strength H.sub.o, the assumed linear 
gradient magnitude G and position x: 
EQU f=.gamma.H.sub.o =.gamma.Gx [Equation 1] 
Inverting this equation gives position x as a function of the nominal field 
strength H.sub.o and NMR frequency f: 
EQU x =f/.gamma.G -H.sub.o /G [Equation 2] 
At the center position where x =0, the NMR resonant frequency f 
=.gamma.H.sub.o. If field drift increases the actual field at the center 
of the magnet to value H.sub.o +.delta., then the same NMR resonant 
frequency f actually corresponds to a position: 
##EQU1## 
Accordingly, as the magnetic field strength drifts to higher values, the 
object effectively "slides" towards the negative x direction (assuming a 
positive x-axis gradient and negative frequencies in the minus x-axis 
dimension as measured from the center of the image space) and a slice 
volume corresponding to a given frequency comes from further back (in the 
negative z-axis direction) in the object. 
In a multiple scan study, such effects can, in part, be reduced by 
appropriately changing the NMR system's center frequency (from which other 
RF transmit frequencies are relatively measured) before each new scan 
sequence is begun. For example, one may use a frequency measurement taken 
from the first and/or last "calibration" measurement cycles of one scan 
sequence so as to compute a new center frequency for use by the MRI system 
in the next subsequent scan sequence. Accordingly, each successive scan is 
at least recalibrated so as to effectively start at a common location thus 
"resetting" the imaged object back to a correct starting position. 
As to the effects of drift or changes occurring during a given measurement 
cycle, one may assume a simple linear model as depicted in FIG. 4. Here, 
it is assumed that the field strength at a given location changes linearly 
with respect to time by the amount .delta. over the entire sequence. The 
relative frequency/phase change then associated with each measurement 
cycle within the scan sequence may be approximated by the frequency/phase 
difference between the first and last calibration cycle divided by the 
number of cycles and added to the starting frequency. In this process, it 
is, of course, assumed that the drift or other spurious changes in the 
magnetic fields occurring during the approximately 10 to 100 milliseconds 
during which a spin echo NMR RF response is being recorded will be small 
compared to the changes which occur during the approximately 0.5 to 2 
seconds TR time between measurement cycles. The recorded complex-valued RF 
sample signals can be multiplied by a suitable complex exponential of the 
appropriate frequency/phase so as to shift the effective frequency back to 
that corresponding to the first measurement cycle. 
For example, consider the correction for cycle n out of N cycles where a 
total field drift is assumed to be .delta.. The relative frequency of the 
nth cycle is n.delta./N. The measured signal in the time domain is donated 
as S.sub.n (t) and the corrected signal is S.sub.n '(t): 
EQU S.sub.n '(t) =S.sub.n (t) exp (j[2.pi.n.delta./N]t) [Equation 4] 
A simple frequency shift of the Fourier Transform (i.e., in the frequency 
spectrum) of the measured signal is also possible. It may be done exactly 
if .delta./N is exactly equal to a multiple of the frequency resolution 
employed in the system. If, as will probably be the case, it is not 
exactly related as an integer, then an approximate shift in the frequency 
domain can still be effected (with some acknowledged inaccuracy caused by 
this lack of precision). 
In practice, the process can be done in a number of ways. The special 
"calibration" cycles such as depicted in FIG. 3 (i.e., those omitting at 
least some and preferably all of the magnetic gradient pulses --except for 
the slice selective gradient pulses which remain) may be interspersed 
within the normal scan sequence or may be tacked on to the beginning 
and/or end of such a sequence as should be apparent to those in the art in 
view of the above description. For example, if one desires 128 
projections, then 130 measurement cycles may actually be employed so as to 
collect an additional first and last measurement cycle without x and y 
gradient pulses. These extra "calibration" measurement cycle data are 
stored on disc along with the other conventional 128 measurement cycle 
data and used by the MRI control computer to change the transmit 
frequencies used for subsequent scans. 
The "calibration" data may also be used by the MRI array processor during 
image reconstruction so as to provide corrections. For example, the array 
processor may use them as soon as possible to produce a conventional 128 
line image --thus, avoiding any requirement for the long term storage of 
extra image data. In another approach, the first and last "calibration" 
measurement cycles can be cycle numbers 1 and 128 so that only 126 phase 
encoded projections are actually acquired and used in the image 
reconstruction process. In this alternative, the array processor may fill 
the usual first and last data sets with zeros so that it will still 
conventionally process 128 projections --albeit there is meaningful 
information in only 126 of those projections. This will slightly reduce 
image resolution --but may be preferable for retrofitting existing MRI 
systems where disc files and the like are already formatted so as to 
accept a certain size data file (e.g., corresponding to 128 projections). 
Although there are many ways to implement a suitable revision of the 
computer programming in a conventional MRI system so as to practice the 
invention, one exemplary embodiment may be realized using the "P command" 
syntax of the disclosed in commonly assigned U.S. Pat. No. 4,707,661 - 
Hoenninger, III et al. In this exemplary embodiment, the presence of phase 
encoding and readout gradients is controlled by suitable look-up tables. 
These tables are level dependent in the preexisting scheme as described in 
the just referenced patent. By adding a second level specification to the 
P command, the first and last measurement cycles of a given scan sequence 
may be caused to use a different set of tables (which happen to contain 
zero values) than for the remaining ones of the measurement cycles. For 
example, a new P command might be implemented as follows: 
EQU (P10:12)S5P13/P20(P21)S1:5P22;C1:128;L1,L3 
In this example, L3 specifies a table pre-filled with zero values. The 
linker operating upon this new P command detects the double level (L1,L3) 
and links one cycle of the command with the table specified by the second 
level. This cycle is then loaded and executed. The sequencer stops and is 
loaded then with as many cycles at the first level as will fit. This 
continues until all of the first level cycles are done. Finally, the last 
measurement cycle with the second level tables is loaded and executed. It 
should be noted that the second level code should not be mixed with the 
first level code because the sub-routines may contain tables that would be 
different for each level and all the cycles in each load share the same 
subroutine code. For example, when the subroutine is loaded with the 
readout gradient turned off, all measurement cycles for this load will 
have the readout magnetic gradient turned off. Accordingly, this may only 
be a solitary first or last cycle in this example. 
The adverse effects of drift in the nominally static magnetic field are 
perhaps most noticeable when a number of complete scan sequences are 
performed in a complex sequential "study" of what are supposed to be the 
same slice volumes from the patient. For example, one scan could be 
performed so as to obtain an image of given slice volumes with one setting 
for the MRI TE and TR parameters while other subsequent scans of the same 
slice volumes would be obtained with different settings for the TR and TE 
parameters. By studying a sequence of R such scans of the same slice, it 
may be possible to perceive additional information concerning a potential 
anomaly in the image. However, if the static field intensity drifts over 
the course of these R scans, then the actual center slice location may 
also shift as depicted in FIG. 5 if the RF transmitter continues to use 
the same center frequency for all of the scans. As can be appreciated, if 
the drift is sufficiently severe, the results could be a significant form 
of artifact in the multiple scan study of a given slice. 
However, if the center frequency of the RF transmitter is reajusted or 
reset at the beginning of each scan (e.g., based upon calibration data 
taken from the just-preceding scan or from an initial "calibration" 
measurement cycle), then the actual location of the slice volumes 
throughout the R scans may be maintained with more accuracy and 
consistency as depicted in the lower portion of FIG. 5. 
As already mentioned, there are numerous ways to implement various aspects 
of this invention. For example, one exemplary implementation is depicted 
in the flowchart of FIG. 6. Here, the program of a conventional MRI system 
is entered at 600 and a complete multi-slice MRI data gathering scan is 
performed at 602 in accordance with conventional practice --but also 
including at least one extra measurement cycle (see FIG. 3) with at least 
one omitted magnetic gradient pulse so as to produce extra calibration 
data. 
At block 604, the calibration data may be used to adjust the nominal center 
frequency f.sub.c for the RF transmitter of the MRI system to use during 
subsequent measurement cycles (e.g., the next scan to be performed in a 
sequence of scans). Thereafter, (or alternatively via bypass 606), the 
calibration data may be used at block 608 to adjust already recorded scan 
data within the individual measurement cycles of a given scan for drift 
and/or remnant eddy current effects occurring during the scan. This 
adjustment may be performed either in the time domain or the spatial 
domain --and may be performed prior to the adjustment of the nominal 
center frequency if desired. 
Thereafter, the conventional multi-dimensional Fourier Transformation 
process is completed at block 610 (if the scan data has been adjusted at 
block 608 in the spatial domain, at least one dimension of Fourier 
Transformation may have already taken place) followed by other 
conventional MRI processing so as to produce final images on a CRT screen 
(or for digital storage and later display). Thereafter, a return to the 
normal program processes of the MRI system is taken at block 612. 
The extra "calibration" measurement cycle(s) with omitted readout and phase 
encoding G.sub.x and G.sub.y magnetic gradients may be thought of as a 
sort of template. As previously explained, templates may be acquired, for 
example, at both the beginning and end of a given scan sequence and 
respectively termed "first template" and "last template." In the course of 
a multi-scan study, after steady state conditions have been achieved, it 
may be preferable to use both the first and last templates to calculate 
the change over a given scan (see FIG. 4) based on a linear or other more 
complex model. 
However, it is also been noted that the first template in the first scan of 
a given multi-scan study may provide somewhat misleading initial 
calibration points. Although the exact cause of this phenomenon is 
difficult to describe precisely, it may be because of some slow thermal 
response in the gradient power supplies or because the eddy currents 
themselves take some time to stabilize into a steady state repetitive 
regimen. 
Accordingly, in the presently preferred exemplary embodiment, it has been 
discovered that one may use only the last template (in conjunction with 
known initialization data at the beginning of the scan) to provide 
significant calibration data for at least the first scan of a multi-scan 
study. By the time the last template is taken, the operation of the 
gradient power supplies and/or the effects of the remnant eddy currents 
appear to have stabilized. 
FIG. 7 provides a schematic pictorial depiction of presently preferred 
exemplary embodiments for achieving calibration using the last template 
calibration data. Here, the time domain spin echo data for the last 
template is depicted at 700. As will be appreciated, the envelope 700 is 
actually represented by a succession of stored digital data representing 
the amplitude A and relative phase 8 of each of successive sampling points 
measured by suitable analog-to-digital converter apparatus during the 
actual occurrence of the time domain spin echo of the last template. 
Through a one dimensional Fourier Transform at 702, the discrete time 
domain data 700 can be transformed o the frequency domain and "binned" at 
locations corresponding to 256 discrete frequencies. Bin number 129 has 
been selected as the "center" of this frequency domain. The frequency 
corresponding to bin number 129 is, at the beginning of a scan sequence, 
defined so as to be at the middle of the x-axis field of view (based upon 
the spectrum of frequencies elicited in the initial NMR RF response from 
this planar volume). When a calibration measurement cycle occurs (without 
readout gradient), the NMR RF response should consist of a very narrow 
band of frequencies and substantially zero relative phase shift. However, 
due to spurious changes in the magnetic fields over the course of a 
complete scan sequence, the last template will likely have a peak in the 
spatial frequency domain that is offset from the center bin 129 by some 
measurable amount .DELTA. as depicted in FIG. 7. To recenter this system 
in preparation for subsequent measurement cycles, the RF transmitter 
center frequency f.sub.c can then be reset by the measured amount .DELTA.. 
Among other possible improvements, such resetting also tends to keep the 
z-axis location of selected slice volumes relatively constant for 
subsequent scan sequences. 
The original time domain spin echo data 706 is also depicted in FIG. 7 for 
a typical measurement cycle n of the N measurement cycles of a complete 
single scan sequence. This original time domain data may be compensated to 
provide compensated data 708 by performing suitable phase corrections 
based upon the last template data (assuming a linear change over all N 
measurement cycles) as depicted in FIG. 7 before performing the usual 
multi-dimensional Fourier Transformation at 710 to produce NMR image data 
and a display at 30. 
Alternatively, so as to avoid the need for a lot of complex number 
multiplications (which may be unduly time consuming), the original time 
domain spin echo data 706 may undergo one dimension of Fourier 
Transformation at 712 to produce a corresponding set of spin echo data 714 
in the spatial domain. Here, the necessary phase correction may be 
achieved by a simple shift or translation in the spatial domain to produce 
compensated spin echo data 716. After this same sort of correction has 
been effected for the requisite 256 measurement cycles (assuming 256 lines 
of y-axis resolution and no synthesized data), an additional dimension of 
Fourier Transformation is performed at 718 to produce NMR image data which 
can again be reused to produce essentially the same image at 30. 
To summarize, in the just described exemplary embodiment, the last template 
is used (in conjunction with the known nominal center frequency set at the 
beginning of the scan) to readjust the center frequency of the RF 
transmitter for subsequent data measurement cycles. One dimension of 
Fourier Transformation is performed on the raw time domain template data, 
the peak magnitude is located in the spatial frequency domain and its 
offset from the center bin (e.g., 129) is measured as the required 
adjustment in the then existing RF transmitter center frequency. This 
required offset may be stored or otherwise communicated to the control 
computer 24 before the next measurement cycle sequence is undertaken. 
The measured frequency shift .DELTA. of the Fourier Transformed last 
template is also equivalent to a pixel shift in the image. One or more 
templates may be used (e.g., with linear or higher order interpolation of 
drift calculated from the measured peaks in the spatial domain). Once the 
offset from the center bin in the spatial domain has been determined, the 
lDFT pixel data (actually representing a phase encoded column along the 
y-axis at a given x-axis location) is merely "scrolled" (i.e., translated 
or shifted) in the proper direction by the number of pixels needed to 
recenter the template to within the nearest one pixel dimension (i.e., the 
shifted position is rounded to the nearest pixel location). 
A more complex and accurate phase correction can be achieved in the time 
domain. As depicted in FIG. 7, this involves complex number multiplication 
which is slower than scrolling --but it provides an added benefit. While 
scrolling in the frequency domain may adequately correct for phase error 
due to field drift, the more accurate phase corrections in the time domain 
will more accurately compensate for remnant eddy current effects as well. 
Accordingly, by performing compensation in the time domain, artifact from 
phase errors of all types may be more accurately compensated. This may be 
especially useful, for example, where some of the spin echo data is 
synthesized based upon assumed complex conjugate symmetries --and which 
process is thus very sensitive to phase errors. Here, as depicted in FIG. 
7, a vector is stored containing the complex conjugate of the measured 
template phase angle .theta.. Each data point to be compensated must then 
be multiplied in the time domain by this compensation vector. Such precise 
phase corrections in the time domain may also be useful for reducing 
artifact when flowing substances are being imaged or when water and fat or 
other substances are being separated by use of the phase changes that can 
be produced due to their slightly different NMR frequencies (e.g., 3.5 ppm 
difference). 
It should also be noted that in prior art techniques for synthesizing data 
based upon complex conjugate symmetry, another type of phase correction is 
conventionally utilized. Here, the zero G.sub.y phase encoded measurement 
cycle is used and the signal at the peak of the spin echo is examined. All 
samples are shifted so that the peak is located at the center of the time 
domain sampling window. Then, the measured relative phase of all the 
samples are shifted by similar amounts so as to force the phase at the 
center to be zero (i.e., all real and no imaginary part). The same time 
and phase shifts then are applied to all other measured data for a given 
scan of a given slice. As will be appreciated, the calibration phase 
shifts and other corrections employed in this invention are achieved in a 
different fashion, based upon different data and used for different 
purposes. 
The phase correction template may be especially useful for gradient 
reversal "echoes" (really a recalled FID). In this case, it may be 
important since the phase of the echo is not fully "cleaned up" by a 
180.degree. RF pulse and any magnet inhomogeneities distort the echo. This 
type "echo" is sometimes better called an FID. 
In some of the superconducting magnets there also may be another effect. 
Gradient pulsing causes a dynamic or transient change of the field 
strength. In some way (perhaps by exerting forces on the magnet winding 
that moves it around in the cryostat) the pulsing gradients cause the 
field to change to a new value while the scan is running. Once the 
gradient pulses stop the field returns to the pre-existing quiescent 
value. This dynamic change is repeatable for the same gradient pulses and 
is different when different shape, duration or direction pulses are used. 
Thus, a correction is needed when scans of the same orientation are used 
where some gradient parameters change and one wants the images from the 
different scans to register with one another. 
The phase correction for eddy currents requires more accuracy and precision 
than correction for field drift. The previously described embodiments work 
best (insofar as eddy currents are concerned) on relatively small image 
volumes having an NMR spectrum consisting of a single absorption line. 
Where there are more substantial field inhomogeneities and/or chemical 
shift effects, such influences are best removed (or at least reduced) 
before making phase shift corrections to compensate for eddy current 
effects. 
One approach is to acquire two sets of calibration data --one with slice 
gradients in normal polarity (e.g., as in FIG. 3) and a second with all 
polarities inverted. The adverse effects caused by field inhomogeneity 
remain the same while those caused by remnant eddy currents change in 
sign. Thus, by subtracting one data set's phase from the phase of the 
other data set, the inhomogeneity errors cancel and, if the result is 
divided by two, the result is the sought-after phase error due to eddy 
currents alone. This gradient reversal approach is depicted in FIG. 9. 
Chemical shift between the water and the fat of body tissue is invariant to 
gradient polarity changes (and thus the correction process depicted in 
FIG. 9 removes chemical shift errors). However, it differs from field 
inhomogeneity in at least one respect. The shift causes water and fatty 
tissue to be selected from two different slices, slightly shifted with 
respect to each other. The direction of the slice's shift in position also 
changes sign when gradient polarities are inverted. So, if the two 
calibration data sets are therefore not collected from a single, 
well-defined slice of the object, the difference will then also contain 
any mismatch between them, which may invalidate the calibration data. 
Therefor, we propose, in one preferred exemplary embodiment, to null out 
the fatty component by a selective presaturation pulse, before acquiring 
each of the calibration data sets. A 180.degree. pulse with a suitabe 
delay is a good candidate for presatuation. This procedure is depicted in 
FIG. 8. 
As shown in FIG. 9, one may make two calibration cycle measurements, first 
with normal slice selection gradient polarities and then with all 
polarities reversed. The effects of the eddy currents have the same 
reversible polarity as the gradient excitation, but the frequency relation 
of fat and water is the same in both cases. Thus, a combination of the 
phase of the two signals will provide cancellation that leaves one with a 
consistent calibration template. 
For example, let SE(+) and SE(-) be the calibration data collected with 
positive and negative gradient polarities respectively. Then the eddy 
current correction template is given by: 
EQU Magnitude of template =1 
EQU Template phase =[phase SE(+) -phase SE(-)]/2 
Multiplying each recorded data point for each "line" of the conventional 
spin echo data set by the complex conjugate of the corresponding data 
point of the template then reduces the phase distortion caused by eddy 
currents. 
When one wants to accurately correct for remnant eddy currents, then an 
inversion recovery calibration cycle may be run as depicted in FIG. 8. 
Here we seek to establish a template for phase correction from human 
tissue having two different frequency sources: fat and water. At 15 MHz 
fat and water have NMR frequencies which differ by about 50 Hz. The result 
is that the spectrum of a single template is not from a single resonance 
line and as one moves away from the peak of the spin echo, signal phase is 
confused by mixture of the signals having these different frequencies. 
One approach to eliminate such confusion is to use an inversion recovery 
sequence as depicted in FIG. 8 where the timing is such that one of the 
two interfering signals is zero when the other is being recorded. This 
depends on a consistent value of T1 for one of the two tissues. Using a 
time-to-inversion TI of 118 msec is known to null fat NMR signals at 15 
MHz. 
For example, as depicted in FIG. 10, if the quiescent spins are inverted by 
a 180.degree. RF nutation pulse, fat nuclei relax back toward the 
quiescent with a shorter time constant than do those having a more 
water-like constitution. Experience has shown that the fat nuclei spins 
will have a zero z magnetization (i.e., zero in FIG. 10) about 118 
milliseconds after inversion (i.e. T1 for fat multiplied by ln 2). Thus, 
if a TI parameter of 118 ms is used in an inversion recovery sequence, the 
fat nuclei NMR signal sources are effectively masked out. 
These FIGS. 8 and 9 sequences have a common feature in that they typically 
may be different from the actual data acquisition cycle by more than just 
the elimination of the readout and phase encoding gradients. The inversion 
recovery (FIG. 8) is a different excitation sequence that depends on the 
different T1 values for fat and water for signal zeroing of the fat 
component. If P commands are used, it would require its own unique P 
command to run. It may be done as a separate calibration data acquisition 
cycle possibly separated further in time from the usual image data 
acquisition scan sequence (e.g. than shown in the FIG. 3 embodiment) 
--especially with a superconducting magnet where eddy current compensation 
is the principal goal. 
The reversed gradient scheme (FIG. 9) acquires two different data sets and 
also may be effected via a unique P command. It has an additional 
complexity since when the slice gradient polarity is reversed the 
excitation frequencies for off center slices also have to be reversed. 
Thus it is easier to do if one already has established a precisely correct 
center slice frequency. Otherwise the plus and minus gradient-defined 
slices may not quite match. 
While only a few presently preferred exemplary embodiments of this 
invention have been described in detail, those skilled in the art will 
appreciate that there are many possible variations and modifications of 
these embodiments which still retain many of their novel features and 
advantages. Accordingly, all such modifications and variations are 
intended to be included within the scope of the appended claims.