U.S. patent number RE35,656 [Application Number 08/515,177] was granted by the patent office on 1997-11-11 for ultra-fast multi-section mri using gradient and spin echo (grase) imaging.
This patent grant is currently assigned to Brigham & Women's Hospital, Inc.. Invention is credited to David A. Feinberg, Koichi Oshio.
United States Patent |
RE35,656 |
Feinberg , et al. |
November 11, 1997 |
Ultra-fast multi-section MRI using gradient and spin echo (GRASE)
imaging
Abstract
Fast magnetic resonance imaging uses combined gradient echoes
and spin echoes. In each of one or more TR intervals, after an
initial NMR RF nutation pulse, a sequence of 180.degree. RF
nutation pulses is used to refocus the RF response into
corresponding string of spin echoes. However, in addition, during
the time that such spin echo would normally occur after each such
180.degree. RF nutation pulse, a plurality of alternating polarity
read-out magnetic gradient pulses is utilized so as to very rapidly
form a sub-sequence of gradient echoes. This fast multi-section MRI
sequence utilizes the speed advantages of gradient refocusing while
overcoming the image artifacts arising from static field
homogeneity and chemical shift. Image contrast is still determined
by the T2 contrast in Hahn spin echoes. A novel k-space trajectory
temporally modulates signals and demodulates artifacts. The echo
responses are selectively phase-encoded and time shifted in
occurrence so as to smoothly distribute unwanted phase shift from
field inhomogeneity and/or chemical phase shift effects over the
entire phase encoded dimension in k-space. The technique can also
be extended so as to provide T2-weighted multi-slab
three-dimensional volume images.
Inventors: |
Feinberg; David A. (New York,
NY), Oshio; Koichi (Brookline, MA) |
Assignee: |
Brigham & Women's Hospital,
Inc. (Boston, MA)
|
Family
ID: |
24921843 |
Appl.
No.: |
08/515,177 |
Filed: |
August 15, 1995 |
Related U.S. Patent Documents
|
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
Reissue of: |
727229 |
Jul 5, 1991 |
05270654 |
Dec 14, 1993 |
|
|
Current U.S.
Class: |
324/309;
324/307 |
Current CPC
Class: |
G01R
33/5615 (20130101); G01R 33/5618 (20130101); G01R
33/56554 (20130101) |
Current International
Class: |
G01R
33/561 (20060101); G01R 33/54 (20060101); G01V
003/00 () |
Field of
Search: |
;324/307,309,306,312,300 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
|
|
|
|
|
|
|
0175184 |
|
Aug 1985 |
|
EP |
|
0318212 |
|
May 1989 |
|
EP |
|
WO 91/02263 |
|
Feb 1991 |
|
WO |
|
Other References
Crooks et al, "Nuclear Magnetic Resonance", Apr. 1982, vol. 143,
No. 1, Nuclear Magnetic Resonance Whole-Body Imager Operating at
3.5 K Gauss.sup.1, pp. 169-174. .
Hennig & Friedburg, "Clinical Applications and Methodological
Developments of the Rare Technique", Magnetic Resonance Imaging,
vol. 6, No. 4, 1988, pp. 391-395. .
Hennig et al, "Rare imaging: A Fast Imaging Method for Clinical
MR", Magnetic Resonance in Medicine 3 (1986), pp. 823-833. .
Rzedzian et al, "Instant Images of the Human Heart Using a New,
Whole-Body MR Imaging System", American J. Roentgenol, vol. 149,
Aug. 1987, pp. 245-250. .
Feinberg et al, "Multiple Spin-Echo Magnetic Resonance Imaging",
Radiology, 1985, vol. 155, pp. 237-442. .
Hahn, "Spin Echoes", Physical Review, vol. 80, No. 4, Nov. 15,
1950, pp. 580-594. .
Mansfield, "Multi-Planar Image Formation Using NMR Spin Echoes", J.
Phys. C: Solid State Phys., vol. 10, 1977; pp. L55-L58. .
Feinberg et al, "Echo Planar-Inner Volume Imaging at 0.35T",
Proceedings of Fifth Annual Meeting of The Society of Magnetic
Resonance in Medicine, p. 950. .
Feinberg et al, "Halving MR Imaging Time by Conjugation:
Demonstration at 3.5 kG", Radiology, 1986, vol. 161, pp. 527-531.
.
Ordidge et al, "Snapshot Imaging at 0.5T Using Echo-Planar
Techniques", Magnetic Resonance in Medicine, vol. 10 (1989), pp.
227-240. .
Pykett et al, "Instant Images of the body by Magnetic Resonance",
Magnetic Resonance in Medicine, vol. 5 (1987), pp. 563-571. .
Feinberg et al, "Echo-Planar Imaging with Asymmetric Gradient
Modulation and Inner-Volume Excitation", Magnetic Resonance in
Medicine, vol. 13, (1990), pp. 162-169. .
Feinberg et al, "Tissue Perfusion in Humans Studied by Fourier
Velocity Distribution, Line Scan, and Echo-Planar Imaging",
Magnetic Resonance in Medicine, vol. 16, (1990), pp. 280-293. .
Oshio et al, "A Computer Simulation of T.sub.2 Decay Effects in
Echo Planar Imaging", Magnetic Resonance in Medicine, vol. 11
(1989), pp. 389-397. .
Mansfield et al, "Zonally Magnified EPI in Real Time by NMR", J.
Phys. E:Sci Instrum., vol. 21 (1988), pp. 275-279. .
Mansfield et al, "Planar Spin Imaging by NMR", Journal of Magnetic
Resonance, vol. 27, pp. 101-119..
|
Primary Examiner: Arana; Louis M.
Attorney, Agent or Firm: Nixon & Vanderhye P.C.
Claims
What is claimed is:
1. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
nutating nuclei within a slice-volume to initiate a TR
interval;
repetitively applying 180.degree. NMR RF pulses to further nutate
nuclei within the same said slice-volume by substantially
180.degree. at subsequent .Iadd.equal time .Iaddend.intervals
within the same TR interval .[.and.]..Iadd., each of said equal
time intervals being substantially twice the interval between said
initial nutating step and the first 180.degree. NMR RF pulse,
.Iaddend.thus to generate a train of NMR spin echoes;
.[.between pairs.]. .Iadd.only after each .Iaddend.of said
180.degree. NMR RF pulses, applying a plurality of alternate
polarity read-out magnetic gradient pulses to produce sub-sequences
of plural gradient echoes occurring between said 180.degree. NMR RF
pulses.
2. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
nutating nuclei within a slice-volume to initiate a TR
interval;
repetitively applying 180.degree. NMR RF pulses to further nutate
nuclei within the same said slice-volume by substantially
180.degree. at subsequent intervals within the same TR interval and
thus to generate a train of NMR spin echoes;
between pairs of said 180.degree. NMR RF pulses, applying a
plurality of alternate polarity read-out magnetic gradient pulses
to produce sub-sequences of plural gradient echoes, and
phase-encoding each said gradient echo within each sub-sequence
.[.to traverse a discontinuous trajectory in k-space which is
interleaved with the trajectories of other sub-sequence.]. to
traverse a discontinuous trajectory in k-space which is interleaved
with the trajectories of other sub-sequences so as to more evenly
distribute field inhomogeneity and/or chemical phase shift effects
over the phase-encoded dimension of k-space.
3. A method as in claim 2 wherein the time occurrences of gradient
echoes within different said sub-sequences are relatively shifted
so as to more evenly distribute field inhomogeneity and/or chemical
phase shift effects over the phase-encoded dimension .[.if.].
.Iadd.of .Iaddend.k-space.
4. A method as in claim 2 further comprising, at the conclusion of
each said subsequence, applying a phase-decoding magnetic gradient
pulse to return nuclei phase encoding to the same point in k-space
prior to application of the next 180.degree. NMR RF nutation
pulse.
5. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
generating a train of NMR spin echoes using a sequence of plural
180.degree. NMR RF nutation pulses;
for each such RF refocused spin echo, generating a sub-sequence of
NMR gradient echoes by using a sequence of alternating polarity
read-out magnetic gradient pulses;
phase-encoding each such gradient echo to trace different
trajectories in k-space, the phase-encoding being modulated to
cause gradient echoes of each said sub-sequence to have k-space
trajectories which are interleaved with those of other such
sub-sequences so as to more evenly distribute T2* and/or chemical
phase shift effects over the phase-encoded dimension of k-space;
and
after each said sub-sequence of gradient echoes, applying a
phase-decoding magnetic gradient pulse to return nuclei phase
encoding to the same point in the k-space prior to generation of
the next NMR spin echo.
6. A method as in claim 5 wherein said train of NMR spin echoes is
generated, during each of plural TR intervals, by an initial
90.degree. NMR RF nutation pulse followed by a sequence of plural
180.degree. NMR RF nutation pulses, each said RF nutation pulse
occurring during a slice-select magnetic gradient pulse
G.sub.z.
7. A method as in claim 6 wherein said sub-sequences of NMR
gradient echoes are generated by a sequence of alternating polarity
G.sub.x read-out magnetic gradient pulses occurring after each
180.degree. NMR RF nutation pulse.
8. A method as in claim 7 also including use of a dephasing G.sub.x
read-out magnetic gradient pulse occurring between said initial
90.degree. NMR RF nutation pulse and the first of said 180.degree.
NMR RF nutation pulses.
9. A method as in claim 7 wherein said phase-encoding during each
sub-sequence is achieved by an initial G.sub.y magnetic gradient
pulse of a first polarity and by subsequent G.sub.y magnetic
gradient pulses of a second polarity and wherein said
phase-decoding is achieved by a further G.sub.y magnetic gradient
pulse of said first polarity.
10. A method for generating MRI signals, said method
comprising:
(a) subjecting NMR nuclei within an image volume to a perturbing
NMR RF nutation pulse;
(b) thereafter subjecting said NMR nuclei to
(i) a 180.degree. NMR RF nutation pulse following by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of plural gradient echoes occurring
after said 180.degree. NMR RF nutation pulse and before application
of another 180.degree. NMR RF nutation pulse, and
(iii) repeating steps (i) and (ii) .Iadd.at equal time intervals
within the same TR interval, each of said equal intervals being
substantially twice the interval between the initial said
perturbing pulse and the first 180.degree. NMR RF pulse .Iaddend.to
generate a further sequence of gradient echoes .Iadd.without
applying alternate polarity read-out magnetic gradient pulses
before the first 180.degree. RF nutation pulse.Iaddend..
11. A method as in claim 10 wherein said perturbing NMR RF nutation
pulse is a 90.degree. NMR RF nutation pulse.
12. A method as in claim 10 wherein step (b) (iii) includes plural
repetitions of steps (i) and (ii) to generate plural further
sequences of gradient echoes.
13. A method as in claim 10, 11 or 12 wherein steps (a) and (b) are
repeated in each of plural TR intervals to generate additional
sequences of gradient echoes.
14. A method in claim 10, 11 or 12 wherein .Iadd.all .Iaddend.said
NMR RF nutation pulses occur during a slice volume selecting
magnetic gradient pulse in a multi-slice sequence.
15. A method as in claim 14 wherein each magnetic gradient read-out
pulse is preceded by a phase-encoding magnetic gradient pulse of
predetermined magnitude, different from the magnitude of other such
phase-encoding pulses.
16. A method for generating MRI signals, said method
comprising:
(a) subjecting NMR nuclei within an image volume to a perturbing
NMR RF nutation pulse;
(b) thereafter subjecting said NMR nuclei to
(i) a 180.degree. NMR RF nutation pulse followed by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of gradient echoes, and
(iii) repeating steps (i) and (ii) to generate a further sequence
of gradient echoes.
said NMR RF nutation pulses occurring during a slice volume
selecting magnetic gradient pulse in a multi-slice sequence,
each magnetic gradient read-out pulse being preceded by a
phase-encoding magnetic gradient pulse of predetermined magnitude,
different from the magnitude of other such phase-encoding pulses,
and
wherein the magnitude of phase-encoding magnetic gradient pulses
with each repetition of steps (i) and (ii) generate MRI gradient
echoes respectively corresponding to non-contiguous trajectories in
k-space, the MRI gradient echoes generated from other repetitions
of steps (i) and (ii) respectively filling in the remaining
contiguous trajectories in k-space in an interleaved fashion.
17. A method as in claim 16 wherein the phase-encoding magnetic
gradient pulses have magnitudes which generate a sequence of MRI
gradient echoes in k-space having substantially reduced phase
shifts between next-adjacent k-space echoes caused by field
inhomogeneity and/or chemical shift effects occurring during each
repetition of steps (i) and (ii).
18. A method as in claim 17 wherein the phase shifts caused by said
field inhomogeneity and/or chemical shift effects increase
monotonically in approximately equal amount from one gradient echo
to the next throughout the phase-encoded dimension of k-space.
19. A method as in claim 18 wherein prior to each repetition of
steps (i) and (ii), a phase-return magnetic gradient pulse is
applied to said NMR nuclei having polarity and magnitude for
substantially cancelling all prior phase-encoding magnetic gradient
pulses and thus momentarily returning the NMR nuclei to the same
point in k-space.
20. A method for generating MRI signals, said method
comprising:
(a) subjecting NMR nuclei within an image volume to a perturbing
NMR RF nutation pulse;
(b) thereafter subjecting said NMR nuclei to
(i) a 180.degree. NMR RF nutation pulse followed by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of gradient echoes, and
(iii) repeating steps (i) and (ii) to generate a further sequence
of gradient echoes.
said NMR RF nutation pulses occurring during a slice volume
selecting magnetic gradient pulse in a multi-slice sequence,
each magnetic gradient read-out pulse being preceded by a
phase-encoding magnetic gradient pulse of predetermined magnitude,
different from the magnitude of other such phase-encoding pulses,
and
wherein prior to each repetition of steps (i) and (ii), a
phase-return magnetic gradient pulse is applied to said NMR nuclei
having polarity and magnitude for substantially cancelling all
prior phase-encoding magnetic gradient pulses and thus momentarily
returning the NMR nuclei to the origin of k-space.
21. A method as in claim .[.19.]. .Iadd.20 .Iaddend.wherein steps
(a) and (b) are repeated in each of plural TR intervals to generate
additional sequences of gradient echoes.
22. Apparatus for generating MRI signals from NMR nuclei within an
image volume, said apparatus comprising:
means for nutating nuclei within a slice-volume to initiate a TR
interval;
means for repetitively applying 180.degree. NMR RF pulses to
further nutate nuclei within the same said slice-volume by
substantially 180.degree. at subsequent .Iadd.equal time
.Iaddend.intervals within the same TR interval .[.and.]..Iadd.,
each of said equal intervals being substantially twice the interval
between said initial nutating pulse and the first 180.degree. NMR
RF pulse, .Iaddend.thus to generate a train of NMR spin echoes,
means for applying a plurality of alternate polarity read-out
magnetic gradient pulses .[.between pairs.]. .Iadd.after each
.Iaddend.of said 180.degree. NMR RF pulses to produce sub-sequences
of plural gradient echoes occurring between said 180.degree. NMR RF
pulses.
23. Apparatus for generating MRI signals from NMR nuclei within an
image volume, said apparatus comprising:
means for nutating nuclei .Iadd.within .Iaddend.a slice-volume to
initiate a TR interval;
means for repetitively applying 180.degree. NMR RF pulses to
further nutate nuclei within the same said slice-volume by
substantially 180.degree. at subsequent intervals within the same
TR interval and thus to generate a train of NMR spin echoes;
means for applying a plurality of alternate polarity read-out
magnetic gradient pulses between pairs of said 180.degree. NMR RF
pulses to produce sub-sequences of plural gradient echoes; and
means for phase-encoding each said gradient echo within each
sub-sequence to traverse a discontinuous trajectory in k-space
which is interleaved with the trajectories of other sub-sequences
so as to more evenly distribute field inhomogeneity and/or chemical
phase shift effects over the phase-encoded dimension of
k-space.
24. Apparatus as in claim 23 including means for shifting the time
occurrences of gradient echoes within different said sub-sequences
so as to more evenly distribute field inhomogeneity and/or chemical
phase shift effects over the phase-encoded dimension of
k-space.
25. Apparatus as in claim 23 further comprising means for applying
a phase-decoding magnetic gradient pulse at the conclusion of each
said subsequence to return nuclei phase encoding to the origin of
k-space prior to application of the next 180.degree. NMR RF
nutation pulse.
26. Apparatus for generating MRI signals from NMR nuclei within an
image volume, said apparatus comprising:
means for generating a train of NMR spin echoes using a sequence of
plural 180.degree. NMR RF nutation pulses;
means for generating a sub-sequence of NMR gradient echoes for each
such .[.RF refocused.]. .Iadd.NMR .Iaddend.spin echo by using a
sequence of alternating polarity read-out magnetic gradient
pulses;
means for phase-encoding each such gradient echo to trace different
trajectories in k-space, the phase-encoding being modulated to
cause gradient echoes of each said sub-sequence to have k-space
trajectories which are interleaved with those of other such
sub-sequences so as to more evenly distribute field inhomogeneity
and/or chemical phase shift effects over the phase-encoded
dimension or k-space; and
means for applying a phase-decoding magnetic gradient pulse after
each said sub-sequence of gradient echoes to return nuclei phase
encoding to the origin of k-space prior to generation of the next
NMR spin echo.
27. Apparatus as in claim 26 wherein said means for generating a
train of NMR spin echoes generates during each of plural TR
intervals, by an initial 90.degree. NMR RF nutation pulse followed
by a sequence of plural 180.degree. NMR RF nutation pulses, each
said RF nutation pulse occurring during a slice-select magnetic
gradient pulse G.sub.z.
28. Apparatus as in claim 27 wherein said means for generating a
sub-sequence of NMR gradient echoes generates a sequence of
alternating polarity G.sub.x read-out magnetic gradient pulses
occurring after each 180.degree. NMR RF nutation pulse.
29. Apparatus as in claim 28 also including means for generating a
dephasing G.sub.x read-out magnetic gradient pulse occurring
between said initial 90.degree. NMR RF nutation pulse and the first
of said 180.degree. NMR RF nutation pulses.
30. Apparatus as in claim 28 wherein said means for .[.generating a
sub-sequence.]. .Iadd.applying a phase-decoding magnetic gradient
pulse .Iaddend.includes means for generating an initial G.sub.y
magnetic gradient pulse of a first polarity and subsequent G.sub.y
magnetic gradient pulses of a second polarity and a further G.sub.y
magnetic gradient pulse of said first polarity to achieve said
phase-decoding.
31. Apparatus for generating MRI signals, said apparatus
comprising:
(a) means for subjecting NMR nuclei within an image volume to a
perturbing NMR RF nutation pulse;
(b) means for thereafter subjecting said NMR nuclei to
(i) a 180.degree. NMR RF nutation pulse following by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of gradient echoes, and
(iii) repeating steps (i) and (ii) .Iadd.at equal time intervals
within the same TR interval, each of said equal intervals being
substantially twice the interval between the initial said
perturbing pulse and the first 180.degree. NMR RF pulse .Iaddend.to
generate a further sequence of plural gradient echoes occurring
after said 180.degree. NMR RF nutation pulse and before application
of another 180.degree. NMR RF nutation pulse .Iadd.without applying
alternate polarity read-out magnetic gradient pulses before the
first 180.degree. RF nutation pulse.Iaddend..
32. Apparatus as in claim 31 including means for repetitively
operating means (a) and (b) in each of plural TR intervals to
generate additional sequences of gradient echoes.
33. Apparatus as in claim 31 wherein said means (a) and means (b)
include means for generating said NMR RF nutation pulses during a
slice volume selecting magnetic gradient pulse in a multi-slice
sequence.
34. Apparatus as in claim 33 wherein said means (b) includes means
for generating a phase-encoding magnetic gradient pulse of
predetermined magnitude, different from the magnitude of other such
phase-encoding pulses prior to each magnetic gradient read-out
pulse.
35. Apparatus for generating MRI signals, said apparatus
comprising:
(a) means for subjecting NMR nuclei within an image volume to a
perturbing NMR RF nutation pulse;
(b) means for thereafter subjecting said NMR nuclei to
(i) a 180.degree.60 NMR RF nutation pulse followed by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of gradient echoes, and
(iii) repeating steps (i) and (ii) to generate a further sequence
of gradient echoes,
said means (a) and means (b) including means for generating said
NMR RF nutation pulses during a slice volume selecting magnetic
gradient pulse in a multi-slice sequence;
said means (b) includes means for generating a phase-encoding
magnetic gradient pulse of predetermined magnitude, different from
the magnitude of other such phase-encoding pulses prior to each
magnetic gradient read-out pulse; and
wherein means (b) includes means for causing the magnitude or
phase-encoding magnetic gradient pulses within each repetition to
generate MRI gradient echoes respectively corresponding to
non-contiguous trajectories in k-space, the MRI gradient echoes
generated from other repititions .[.of.]. respectively filling in
the remaining contiguous trajectories in k-space in an interleaved
fashion.
36. Apparatus as in claim 35 wherein means (b) includes means for
causing the phase-encoding magnetic gradient pulses to have
magnitudes which generate a sequence or MRI gradient echoes in
k-space having substantially reduced phase shifts between
next-adjacent k-space echoes caused by field inhomogeneity and/or
chemical shift effects occurring during each repetition.
37. Apparatus as in claim 36 wherein the means (b) includes means
for causing the phase shifts of said field inhomogeneity and/or
chemical shift effects to increase monotonically in approximately
equal amount from one gradient echo to the next throughout the
phase-encoded dimension of k-space.
38. .[.A method.]. .Iadd.Apparatus .Iaddend.as in claim 37
including means for generating a phase-return magnetic gradient
pulse applied to said NMR nuclei having polarity and magnitude for
substantially cancelling all prior phase-encoding magnetic gradient
pulses and thus momentarily returning the NMR nuclei to the same
point in k-space prior to each repetition of a 180.degree. RF
pulse.
39. Apparatus for generating MRI signals, said apparatus
comprising:
(a) means for subjecting NMR nuclei within an image volume to a
perturbing NMR RF nutation pulse;
(b) means for thereafter subjecting said NMR nuclei to
(i) a 180.degree.60 NMR RF nutation pulse followed by
(ii) a plurality of alternating polarity magnetic gradient read-out
pulses to generate a sequence of gradient echoes, and
(iii) repeating steps (i) and (ii) to generate a further sequence
of gradient echoes,
said means (a) and means (b) including means for generating said
NMR RF nutation pulses during a slice volume selecting magnetic
gradient pulse in a multi-slice sequence;
said means (b) includes means for generating a phase-encoding
magnetic gradient pulse of predetermined magnitude, different from
the magnitude of other such phase-encoding pulses prior to each
magnetic gradient read-out pulse; and
means for generating a phase-return magnetic gradient pulse applied
to said NMR nuclei having polarity and magnitude for substantially
cancelling all prior phase-encoding magnetic gradient pulses and
thus momentarily returning the NMR nuclei to the same point in
k-space prior to each repetition of a 180.degree. RF pulse.
40. Apparatus as in claim .[.38.]. .Iadd.39 .Iaddend.wherein means
(a) and (b) are repetitively operated to define plural TR intervals
generating additional sequences of gradient echoes.
41. A method for generating MRI signals from NMR nuclei with an
image volume, said method comprising:
nutating nuclei within a slice-volume to initiate a TR
interval;
repetitively applying 180.degree. NMR RF pulses to further nutate
nuclei within the same said slice-volume by substantially
180.degree. at subsequent intervals within the same TR interval and
thus to generate a train of NMR spin echoes;
between pairs of said 180.degree. NMR RF pulses, applying a
plurality of alternate polarity read-out magnetic gradient pulses
to produce sub-sequences of gradient echoes; and
wherein the time occurrences of gradient echoes within different
said sub-sequences are relatively shifted so as to more evenly
distribute field inhomogeneity and/or chemical phase shift effects
over the phase-encoded dimension if k-space.
42. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
nutating nuclei within a slice-volume to initiate a TR
interval;
repetitively applying 180.degree. NMR RF pulses to further nutate
nuclei within the same said slice-volume by substantially
180.degree. at subsequent intervals within the same TR interval and
thus to generate a train of NMR spin echoes;
between pairs of said 180.degree. NMR RF pulses, applying a
plurality of alternate polarity read-out magnetic gradient pulses
to produce sub-sequences of gradient echoes; and
at the conclusion of each said subsequence, applying a
phase-decoding magnetic gradient pulse to return nuclei phase
encoding to the same point in k-space prior to application of the
next 180.degree. NMR RF nutation pulse.
43. Apparatus for generating MRI signals from NMR nuclei within an
image volume, said apparatus comprising:
means for nutating nuclei within a slice-volume to initiate a TR
interval;
means for repetitively applying 180.degree. NMR RF pulses to
further nutate nuclei within the same said slice-volume by
substantially 180.degree. at subsequent intervals within the same
TR interval and thus to generate a train of NMR spin echoes;
means for applying a plurality of alternate polarity read-out
magnetic gradient pulses between pairs of said 180.degree. NMR RF
pulses to produce sub-sequences of gradient echoes; and
means for shifting the time occurrences of gradient echoes within
different said sub-sequences so as to more evenly distribute field
inhomogeneity and/or chemical phase shift effects over the
phase-encoded dimension of k-space.
44. Apparatus for generating MRI signals from NMR nuclei within an
image volume, said apparatus comprising:
means for nutating nuclei within a slice-volume to initiate a TR
interval;
means for repetitively applying 180.degree. NMR RF pulses to
further nutate nuclei within the same said slice-volume by
substantially 180.degree. at subsequent intervals within the same
TR interval and thus to generate a train of NMR spin echoes;
means for applying a plurality of alternate polarity read-out
magnetic gradient pulses between pairs of said 180.degree. NMR RF
pulses to produce sub-sequences of gradient echoes; and
means for applying a phase-decoding magnetic gradient pulse at the
conclusion of each said subsequence to return nuclei phase encoding
to the origin of k-space prior to application of the next
180.degree. NMR RF nutation pulse.
45. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
(i) nutating nuclei within a selected volume to initiate a TR
interval;
(ii) applying a 180.degree. NMR RF pulse to further nutate nuclei
within said volume by substantially 180.degree. at a subsequent
interval to generate a NMR spin echo RF response; and
(iii) applying a plurality of phase-encoding magnetic .[.and.].
gradient pulses and alternate polarity read-out magnetic gradient
pulses to produce a sequence of plural gradient echoes wherein the
time domain positions of at least some or said gradient echoes is
shifted as a result of shifted time domain positions for the
phase-encoding magnetic gradient pulses and alternate polarity
read-out magnetic gradient pulses from the nominally expected
respective Hahn spin echo times to provide controlled increments of
undesired phase shifts in contiguous k-space MRI data.
46. A method as in claim 45 wherein steps (ii) and (iii) are
repeated at least once after one occurrence of step (i).
47. .[.A method.]. .Iadd.Apparatus .Iaddend.for generating MRI
signals from NMR nuclei within an image volume, said method
comprising:
(i) .Iadd.means for .Iaddend.nutating nuclei within a selected
volume to initiate a TR interval;
(ii) .Iadd.means for .Iaddend.applying a 180.degree. NMR RF pulse
to further nutate nuclei within said volume by substantially
180.degree. at a subsequent interval to generate a NMR spin echo RF
response; and
(iii) means for applying a plurality of phase-encoding magnetic
.[.and.]. gradient pulses and alternate polarity read-out magnetic
gradient pulses to produce a sequence of plural gradient echoes
wherein the time domain positions of at least some of said gradient
echoes is shifted as a result of shifted time domain positions for
the phase-encoding magnetic gradient pulses and alternate polarity
read-out magnetic gradient pulses from the nominally expected
respective Hahn spin echo times to provide controlled increments of
undesired phase shifts in contiguous k-space MRI data.
48. .[.A method.]. .Iadd.Apparatus .Iaddend.as in claim 47 wherein
.[.steps.]. .Iadd.means .Iaddend.(ii) and (iii) are operated at
least twice after one operating of means (i). .Iadd.
49. A method for generating MRI signals from NMR nuclei within an
image volume, said method comprising:
nutating nuclei to initiate an MRI data acquisition pulse
sequence;
repetitively applying a plurality of 180.degree. NMR RF pulses only
at subsequent equal intervals to produce a Hahn spin echo
occurrence at equal intervals after each 180.degree. NMR RF pulse;
and
after each 180.degree. NMR RF pulse, applying a plurality of
alternate polarity read-out magnetic gradient pulses to produce
sub-sequences of gradient echoes occurring after each said
180.degree. NMR RF pulse..Iaddend..Iadd.50. A method as in claim 49
wherein a phase-encoding magnetic gradient pulse of the same
polarity but different respective magnitude is applied prior to
each gradient echo thus producing sub-sequences of phase-encoded
gradient echoes..Iaddend..Iadd.51. A method as in claim 50 wherein
prior to each repetition of a 180.degree. NMR RF pulse, a
phase-return magnetic gradient pulse is applied to substantially
cancel all prior phase-encoding magnetic gradient pulses and thus
momentarily return NMR nuclei to the origin of
k-space..Iaddend..Iadd.52. A method as in claim 50 wherein said
phase-encoding magnetic gradient pulses produce phase-encoded
gradient echo sub-sequences that each have non-contiguous k-space
trajectories, contiguously interleaved with those
of other gradient echo sub-sequences..Iaddend..Iadd.53. A method as
in claim 51 wherein said phase-encoding magnetic gradient pulses
produce phase-encoded gradient echo sub-sequences that each have
non-contiguous k-space trajectories, contiguously interleaved with
those of other gradient echo sub-sequences..Iaddend..Iadd.54. A
method for generating MRI signals from NMR nuclei within an image
volume, said method comprising:
nutating nuclei to initiate an MRI data acquisition sequence;
repetitively applying a plurality of 180.degree. NMR RF pulses, at
least some of which 180.degree. pulses are followed by a plurality
of alternate polarity read-out magnetic gradient pulses to produce
sub-sequences of gradient echoes with phase-encoding magnetic
gradient pulses of different respective magnitudes being applied
prior to each echo occurrence thus providing phase-encoded echo
signals; and
prior to each repetition of a 180.degree. NMR RF pulse, applying a
phase-return magnetic gradient pulse to substantially cancel all
prior phase-encoding magnetic gradient pulses and thus momentarily
return NMR nuclei to the origin of k-space..Iaddend..Iadd.55. A
method as in claim 54 wherein said 180.degree. NMR RF pulses are
all spaced at equal time intervals from each other to produce a
Hahn spin echo occurrence at equal intervals after each 180.degree.
NMR RF pulse..Iaddend..Iadd.56. A method as in claim 54 wherein the
phase-encoding magnetic gradient pulses of at least some contiguous
sub-sequences are of the same
polarity..Iaddend..Iadd.57. A method as in claim 54 wherein said
phase-encoding magnetic gradient pulses within a given sub-sequence
produce a sub-sequence of phase-encoded gradient echoes having
non-contiguous k-space trajectories..Iaddend..Iadd.58. A method for
generating MRI signals from NMR nuclei within an image volume, said
method comprising:
nutating nuclei to initiate an MRI data acquisition sequence,
and
repetitively applying a plurality of 180.degree. NMR RF pulses, at
least some of which 180.degree. pulses are followed by a plurality
of alternate polarity read-out magnetic gradient pulses and
phase-encoding magnetic gradient pulses of the same polarity but
different respective magnitudes to produce sub-sequences of
phase-encoded gradient echoes, each sub-sequence having
non-contiguous k-space trajectories..Iaddend..Iadd.59. A method as
in claim 58 wherein said 180.degree. NMR RF pulses are all spaced
at equal time intervals from each other to produce a Hahn spin echo
occurrence at equal intervals after each 180.degree. NMR RF
pulse..Iaddend..Iadd.60. A method as in claim 58 wherein prior to
each repetition of a 180.degree. NMR RF pulse, a phase-return
magnetic gradient pulse is applied to substantially cancel all
prior phase-encoding magnetic gradient pulses and thus momentarily
return NMR nuclei to the origin of k-space..Iaddend.
Description
BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention relates generally to magnetic resonance imaging
(MRI) utilizing nuclear magnetic resonance (NMR) phenomena
associated with selected NMR nuclei of a patient image volume
within an MRI apparatus. It is more particularly directed to method
and apparatus for achieving an MRI NMR pulse sequence which
combines gradient and spin echo (GRASE) MRI techniques in
advantageous ways.
2. Related Prior Art
Over the last ten years or so, commercial MRI systems have become
readily available. Some magnetic resonance spectroscopic imaging
(MRSI) apparatuses are also now fairly well-known in the art and in
use in at least laboratory environments. Similar MRI techniques are
utilized in both MRI and MRSI and the term MRI will be collectively
used hereinafter to refer to either or both of such techniques and
apparatuses.
In conventional MRI apparatus, the relevant patient anatomy is
positioned within a predetermined patient imaging volume where a
large magnet (e.g, cryogenic, resistive and/or permanent magnet)
structure creates a substantially constant and homogeneous magnetic
field B.sub.0. Conventional gradient coil structures of various
types are also included in the MRI apparatus so as to permit rapid
superposition of magnetic gradients with the base magnetic field
B.sub.0 in the image volume. Typically, these magnetic gradients
are labeled G.sub.x, G.sub.y and G.sub.z --indicating gradients
oriented along the usual x,y,z Cartesian coordinate system (the
B.sub.0 field typically being aligned with the z-axis of the same
coordinate system). Radio frequency (RF) coils are also tightly RF
coupled to the image volume for both transmitting and receiving RF
signals to and from the patient tissue nuclei located
therewithin.
As is well-known by those in the art, nuclei having an odd number
of protons (e.g., hydrogen nuclei) will tend to align their
rotating net magnetic moments with the quiescent background
magnetic field B.sub.0. However, when subjected to a suitable RF
signal at the proper Larmor frequency (proportional to the magnetic
field at the site of the nucleus), the rotating net magnetic
moments of a substantial proportion of such nuclei may be tilted or
nutated away from the quiescent orientation. If subsequently
released from such electromagnetic nutation forces, the nuclei will
tend to again revert to the quiescent orientation--and will emit
characteristic RF signals which can be detected with suitable MRI
RF receiving circuits. By subjecting NMR nuclei in a selected image
volume to particular sequences of RF nutation pulses and magnetic
gradient pulses, NMR RF responses can be detected and processed
(e.g., via multi-dimensional Fourier Transformation) so as to yield
data representing the spatial distribution of NMR nuclei within the
imaged volume. Such data can then be displayed visually where the
intensity or color of each pixel or group of pixels in a
two-dimensional display represents the NMR nuclei density at a
respectively corresponding spatial location within the imaged
volume.
Commercially available MRI systems incorporate sophisticated
computer control systems for effecting preprogrammed NMR sequences
of RF and magnetic gradient pulses for particular types of MRI
effects. Virtually any desired NMR sequence can be programmed
within the operational limits of the RF and magnetic gradient
drivers (e.g., as to magnitudes, rise and fall times, maximum duty
cycles, etc). This permits virtually an infinite variety of
combinations and permutations of RF and magnetic gradient pulses
and many of these possibilities have yet to be explored.
Over the years, many different MRI pulse sequences have been
developed and used to successfully image various types of patient
tissues. A few of the more well-known MRI pulse sequences are
briefly described below:
Spin Echo (SE) MRI (FIG. 2)
Traditional SE typically employs, during each repetition time
interval or "TR" intervals an initial 90.degree. RF nutation pulse
followed by one or more 180.degree. RF nutation pulses to form spin
echo RF responses. Only one echo per 180.degree. RF nutation pulse
is acquired for a given image (if multiple echoes are acquired,
they are each used for respectively different echo images). Each SE
is differently phase-encoded in one dimension (e.g., the y-axis)
from all other SE responses so as to trace out a different part of
phase-encoded "k"-space. Two-dimensional Fourier Transformation of
the acquired k-space data ultimately provides data that can be used
to directly display a meaningful visual image on a CRT or the like.
This technique therefore typically requires a long MRI data
acquisition time (e.g., several minutes) to generate the required
number (e.g., 256 or 512) of phase-encoded spin echoes for a
complete image.
Echo Planar Imaging (EPI) MRI (FIG. 3)
Traditional EPI typically uses an initial 90.degree. RF nutation
pulse (and an optional 180.degree. RF nutation pulse) to generate a
spin echo which is thereafter repetitively refocused by a read-out
magnetic gradient that is of quickly alternating polarity to form a
train of multiple "gradient" echoes (GE). Each of these GE are
typically differently phase encoded by either a small constant
phase-encoding magnetic gradient or by small magnetic gradient
pulses occurring between the echoes. While an EPI sequence can
collect a full image set in a very short time (e.g., tens of
milliseconds), it requires comparatively very high performance MRI
system hardware.
MBEST, ABEST, and Instascan, (Oridge et al, Magn Reson. Med., Vol
10, p227 (1989); Feinberg et al, Magn. Reson. Med., Vol 13, p162
(1990); Rzedzian et al, Amer. J. Roentgenol, Vol 149, p245 (1987))
are variants of the original EPI k-space trajectory proposed by
Mansfield et al. These methods continuously displace the signal
trajectory along the phase axis of k-space during the echo
train.
Rapid Acquisition With Relaxation Enhancement (RARE) MRI (FIG.
4)
Traditional RARE sequences typically also start with a 90.degree.
RF nutation pulse. Then a train of 180.degree. RF nutation pulses
are applied to generate multiple spin echoes. Each spin echo is
differently phase-encoded to consecutively trace out k-space.
However, here, after each echo a phase-decode pulse of opposite
polarity is used to return to the origin in k-space. This helps
suppress stimulated echo artifacts that might otherwise result from
imperfections in the 180.degree. RF nutation pulses. Since a
180.degree. RF nutation pulse is required to generate each echo,
the required image acquisition time is much longer than with EPI.
RARE also typically requires multiple excitations (i.e., multiple
TR intervals) to collect a full image data set whereas EPI
typically may be performed in "one shot" (i.e,, in one TR
interval).
Small Flip Angle Methods (FIG. 5)
Traditional small flip angle (e.g., GRASS, FLASH, FISP, etc.) MRI
uses an initial RF nutation pulse of reduced magnitude (i.e., less
then 90.degree., e.g., 45.degree. or even less) to maintain a
relatively high signal to noise ratio (S/N) within a shortened
repetition time TR. Perhaps the main disadvantage of this method is
that resultant image contrast is different from that obtained using
SE MRI--which is the currently accepted standard in clinical
MRI.
EPI (e.g., Mansfield et al, J. Magn. Reson., Vol 27, p101 (1977))
is known to place major hardware demands on MRI systems; including
static field B.sub.0 homogeneity, magnetic gradient power and
magnetic gradient switching time. Although the RF refocused variant
of echo planar imaging, RARE, e.g., Henning et al, Magn. Reson.
Med., Vol 3, p.823 (1986)) is not so hampered by chemical shift,
image distortions nor other field inhomogeneity effects, it is
considerably slower. The use of slice selective 180.degree. RF
nutation pulses requires more time than gradient polarity switching
(in the range of milliseconds as compared to microseconds). A
second disadvantage of RARE is its much higher RF energy
deposition, SAR, which can exceed presently acceptable safety
limits for the human body.
One earlier proposed approach to minimizing the limitations of EPI
and RARE is to alternate between gradient and RE refocusing within
the echo train, as suggested in earlier experiments for single shot
inner volume imaging (Feinberg et al, Proceedings Fifth Annual
Meeting of the Society of Magnetic Resonance in Medicine P.950
(1986)). Not only is this Abstract not coherently understandable,
it is also only disclosed for use with a single subsection of a
slice (defined by transversely intersecting slices). Furthermore,
it does not disclose any order for combining GE with SE. Nor does
this early published abstract recognize or deal with the various
potential sources of artifacts and errors which arise when one
actually does combine GE with SE.
While MRI has improved rapidly over the last decade, it is
remarkable that multi-section 2DFT spin echo imaging (Crooks,
Radiology (1982)) has remained the most commonly accepted standard
for routine clinical MRI studies of the body and head. Since the
earliest studies in 1982, there has been a progressive increase in
the image signal-to-noise ratio (S/N) due to hardware improvements
and pulse sequence improvements (Feinberg et al, Radiology (1985)).
The resulting higher system S/N has permitted faster imaging times
by reducing the number of excitation (NEX), otherwise used to raise
S/N. Conjugate synthesis of data (Half Fourier or NEX=1/2) yields
nearly another factor of two reduction in imaging time by using to
advantage a natural symmetry in the spin echo signals for computer
synthesis of half of the phase encoded signals. The conjugate
synthesis method has identical tissue contrast, chemical shift and
spatial resolution as regular spin echo imaging but with an
expected reduction of S/N by about 30 to 40%, acceptable for many
T2 weighted screening exams.
An alternative approach for faster imaging, RARE, (Henning et al,
Magn. Reson, Med. Vol 3, p.823 (1986); Henning et al Magn, Recon.
Imag., Vol 6, p391 (1988)) reduces imaging time by performing phase
encoding during multiple cycles (TR) of signal excitation. A
reduction in the number of imaging sections can be a penalty.
However, these images have similar contrast to conventional spin
echo imaging, with an 8 to 16 fold reduction in imaging time from
NEX=1 spin echo techniques. The imaging speed is currently limited
by RARE's increased RF energy deposition in the human body (SAR)
due to the rapid application of a large number of 180.degree. RF
nutation pulses. Ultimately, the speed of RARE imaging is
physically limited by the fairly large total time period required
for multiple slice selective 180.degree. RF pulses during which
time NMR signals cannot be read out.
In EPI (Manfield et al, J. Magn. Reson., Vol 27, p.101 (1977))
signal refocusing by rapid gradient polarity switching can be used
instead of the slower RF refocusing of RARE. In this way, both EPI
and its modern variants, MBEST and Instascan can make images in
imaging times less than 100 msec.
These EPI methods produce a larger amount of chemical shift on the
image phase axis (rather than on the frequency axis), typically a
10 pixel mis-registration between water and fat. This problem can
be circumvented using fat suppression methods. EPI images typically
have had lower spatial resolution and S/N than spin echo imaging,
which can be improved with multiple excitation cycles (multiple TR)
and longer imaging times than the original single shot technique.
Gradient hardware with high maximal gradient strength and short
gradient rise times are required for EPI which to date has
effectively limited EPI applications to a handful of research
centers.
Several issued patents are also known which appear to relate to MRI
techniques in which slice selective RF nutation pulses (directed to
the same spatial volume) are associated with the generation of both
spin echoes and gradient echoes:
U.S. Pat. No. 4,796,635--Dumoulin (1989);
U.S. Pat. No. 4,818,942--Rzedzian (1989);
U.S. Pat. No. 4,833,407--Holland et al (1989);
U.S. Pat. No. 4,896,113--Pelc (1990); and
U.S. Pat. No. 4,901,020 Ladebeck et al (1990).
Dumoulin appears primarily to use conventional echo sequences
generated by gradient reversals (i.e. "gradient echoes"). There is
no suggestion of using plural 180.degree. RF nutation pulses in
each TR interval.
Rzedzian appears to use plural gradient echoes after each 90
degree-180 degree RF pulse pair. Special traversals of k-space are
also involved. However, there does not appear to be any suggestion
in this document for repeating a 180 degree RF pulse and plural
gradient echo sub-sequences after an initial 90 degree RF pulse in
each TR interval--nor for a non-sequential traversal of
k-space.
Holland et al generate a single spin echo and at least one gradient
echo each TR and map each into a common k-space for constructing a
single image. Phase encoding pulses within each TR are off-set by
about .[.174.]. .Iadd.1/4 .Iaddend.of k-space so that the upper two
quarters of k-space are each gradually completed as successive TR
intervals occur. Thereafter, complex conjugation is used to
symmetrically construct the lower 1/2 of k-space. There is no
intention of using plural spin echoes in a single echo train, this
is why the pulse sequence diagram shows redundancy of the sequence
after one spin echo group of gradient echoes. There is no
suggestion of multiple 180.degree. RF pulses in one TR interval.
Holland uses "multiple" TR excitations with a single spin echo
during each TR excitation--which is further encoded by two gradient
echoes, which are further complex conjugated to produce two
additional synthesized signals, thus a factor of four reduction in
imaging time from spin echo imaging.
Pelc uses plural spin-echoes and plural gradient echoes during each
TR (i.e., after each 90.degree. RF pulse). However, it appears that
such are all with the same phase encoding and used for averaging to
map into the same line of k-space. In other words, Pelc does not
combine data from multiple spin echoes to improve imaging speed.
Instead, Pelc uses information from gradient echoes adjacent the
second spin echo only for various error corrections.
Ladebeck et al appears to use at least one spin echo and a
plurality of gradient echoes within a single TR. However, it
appears that the spin echoes and gradient echoes are separately
mapped into k-space for generating separate respective images. Nor
is there any apparent suggestion of taking several gradient echoes
after a 180 degree RF pulse.
Also called to applicants' attention by a prior art search are the
following which presently appear to be even less relevant:
U.S. Pat. No. 4,792,758--Sattin (1988)
U.S. Pat. No. 4,800,889--Dumoulin et al (1989)
U.S. Pat. No. 4,871,967--Rotem et al (1989)
U.S. Pat. No. 4,893,081--Zur (1990)
U.S. Pat. No. 4,896,112--Ratzel et al (1990)
U.S. Pat. No. 4,959,611--Brovost etal (1990)
Perhaps the closest prior art known to us at this time is:
Rzedzian and Pykett IL, Amer. J. Roentgenol. 49,245 (1987);
Feinberg DA, U.S. Pat. No. 4,684,891 (August 1987);
Hennig J, Frieburg H, Magn Reson Imag: 6:391 (1988);
Mansfield P, Ordidge R. J. and Coxan, J. Phys E21, 278 (1988);
and
Feinberg et al, 1986 SMRM Abstract.
The Rzedzian et al, Henning et al and Feinberg et al papers have
already been briefly referenced above. The Feinberg 1981 patent is
an example using slice selective 90.degree. and 180.degree. RF
pulses each TR interval to generate a train of phase-encoded spin
echoes which after T2 correction, are used to fill in k-space for a
common image.
BRIEF SUMMARY OF THE INVENTION
Although the earlier limited attempts by Feinberg et al (Supra.
Soc. Magn. Res. in Med., p950, (1986)) to combine gradient echoes
and spin echoes in each TR interval were flawed, we have continued
to work on the problem and have now succeeded. That is to say, we
have now discovered a workable class of MRI NMR pulse sequences
which can effectively combine GE (e.g., as used in EPI) and SE to
obtain many of the advantages enjoyed separately by each prior MRI
technique--while yet avoiding some of the relative disadvantages
earlier perceived for these individual techniques. By creating
multiple short gradient echo trains between successive 180.degree.
RF nutation pulses in a single TR interval, field inhomogeneity and
chemical shift effects evolve over the relatively short time period
between adjacent 180.degree. RF pulses--instead of the longer time
of the total echo train as in EPI. Advantages over RARE, include
reduced RF power deposition (SAR) and potentially much faster
imaging times. This combined excitation approach, defined here as
gradient and spin echo (GRASE) imaging, maintains a very fast
imaging speed intermediate between EPI and RARE.
The exemplary embodiment of GRASE uses more than one 180.degree. RF
nutation pulse per TR interval and combines all the resulting echo
data into a common k-space image plane for increased speed. EPI
techniques (even with modern variations called MBEST and
Instascan), utilize only one 180.degree. RF pulse per TR interval
in the sequence and combine data from a region around a single Hahn
echo time. By using plural 180.degree. RF pulses per TR interval,
GRASE limits the amount of chemical shift and distortion errors due
to field inhomogeneity.
The novel GRASE technique accomplishes high speed imaging of the
body and head with tissue contrast similar to traditional spin echo
imaging and with no significant increase in SAR. In the current
exemplary implementation, GRASE is about 25 times faster than spin
echo imaging, while maintaining high spatial resolution and image
quality.
Multi-sectional body imaging in 18 seconds at a TR of 2 sec.,
overcomes image degradation due to respiratory motion and reduces
peristalsis related artifacts. Advantages of the exemplary GRASE
technique include reduced chemical shift, reduced image distortion
due to field inhomogeneity and its demonstrated performance on
clinical MR systems without gradient hardware modification. There
also may be a major improvement in T2 weighted abdominal MRI.
Slice selective 90.degree. and plural slice-selective 180.degree.
RF pulses produce a train of spin echo signals each TR. Each spin
echo signal also is repeatedly refocused with plural read out
gradients to produce multiple gradient echo signals from each spin
echo. Combining this process in a standard "multi-section
excitation scheme permits large improvements over current imaging
methods by increasing net signal to noise ratio per imaging time
and by greatly reducing imaging time, as much as 24 fold.
This invention produces multiple gradient echoes from spin echoes
while the earlier work of others typically used either gradient
echoes or spin echoes alone to produce images. The GRASE advantage
over RARE (multiple spin echo technique) includes production of
more signals per 90.degree. RF excitation (faster imaging), higher
signal to noise per imaging time and lower RF power deposition to
human body (SAR). Advantages over echo planar imaging (EPI or
gradient echo techniques), include decreased T2* decay, higher S/N
per image, less image distortion from field inhomogeneity and
chemical shift artifact. Unlike EPI, GRASE can be implemented on
standard commercially available MRI imagers.
GRASE permits T2 weighted images to be made in about 18 seconds,
the time of a single breath hold, thus eliminating respiratory
motion blurring in MRI of the human abdomen and thorax. These
multi-slice T2 weighted images are obtained in approximately 1/24th
the time of current images, and so permit rapid imaging of patients
likely to move in longer time intervals, for example pediatric
patients or emergency trauma patients. Diffusion sensitivity of
images permits thermal imaging for interventional laser therapy.
GRASE also can be used to reduce imaging time of more complicated
methods, 3D multi-slab images, high spatial resolution
(512.times.512 pixel arrays), and flow imaging. Near "real-time" or
cine-MR methods can also be an extension of GRASE and permit
changes in MRI as a diagnostic tool, with such advantages as are
currently employed in ultrasound imaging.
The GRASE sequence can be incorporated into virtually any
commercially available MRI imager. Magnetic gradient performance
(maximal strength, rise time and stability) determine the largest
number of signals possible to obtain in each excitation cycle. More
than two gradient echoes per spin echo preferably include added
data correction for variable T2* effect, T2 effect and phase
shifting between gradient echoes.
While the subsecond imaging time of EPI largely overcomes cardiac
motion artifacts, imaging in under 20 seconds using GRASE MRI would
be sufficiently fast to allow for a patient's breath hold and to
substantially eliminate respiratory motion artifacts which
currently limit the clinical utility of T2 weighted multi-sectional
body imaging. With this goal in mind, we have combined gradient
refocusing and RF refocusing techniques. Gradient refocusing
methods are used to produce several signals from each of multiple
RF refocused spin echoes, to obtain significant time improvements
over RARE and reduced SAR. Similarly, image quality can be
maintained at a very high level and with much less chemical shift
than in EPI. As described further below, attempts to achieve GRASE
involve what initially appear as significant obstacles due to
inherent modulation of field inhomogeneity and chemical shift
effects in the GRASE echo train. This has led to a totally new
k-space image trajectory in the exemplary GRASE embodiment
differing significantly from the k-space image trajectories
typically used in RARE and EPI techniques. In comparison, the
exemplary GRASE k-space trajectory sweeps through multiple
discontinuous and purposefully modulated paths on the k-space phase
axis as a function of time in the echo train in order to reduce
image chemical shift and field inhomogeneity errors.
The exemplary GRASE technique uses multiple gradient refocusings
between plural 180.degree. RF refocusing pulses and combines all
the resulting signals into one image, with a reduction in imaging
time being proportional to the number of acquired signals. It is
not believed possible to perform this complex method using any
straightforward combination of currently known imaging methods such
as combining known methods like echo planar gradient echo imaging
with multiple 180.degree. RF RARE.
In this regard, GRASE is not a method of simply generating multiple
gradient echoes from a single spin echo (which several others have
pursued) of the echo planar type sequences. One disadvantage of
such prior attempts is the field inhomogeneity errors, T2* decay,
chemical shift, and low signal to noise, all of which are known
problems of EPI.
Our GRASE MRI technique is believed to be most successful, at least
in part, because of four preferred features of the exemplary image
sequence (which may be employed in various sub-combinations and
permutations).
I. Echo Generation
First, as previously mentioned, the gradient echoes and spin echoes
are sequentially combined using two or more 180.degree. RF pulses
and two or more gradient echoes per spin echo or per each
180.degree. RF nutation pulse. This first step reduces SAR,
chemical shift, field inhomogeneity errors, and T2* decay of
signals and it increases signal amplitude for higher signal to
noise ratios.
II. k-space Trajectory
Secondly, our exemplary GRASE technique demodulates the remaining
small field inhomogeneity errors (which inherently otherwise would
become modulated when these gradient echoes and spin echoes are
combined--i.e., juxtaposed in interleaved fashion in k-space and
jointly Fourier Transformed to make a single image). Without this
second step, the advantage of reducing field inhomogeneity errors
and the above mentioned problems, by the use of multiple
180.degree. pulses may not be adequately fulfilled since severe
image artifacts may result from the discontinuous pattern of the
remaining small phase errors disposed throughout the total echo
train. Therefore, once the signals are acquired with multiple
temporally sequential magnetic gradient and 180.degree. RF
refocusings, only one specific complex order of phase encoding is
preferably used to reduce or substantially eliminate this periodic
pattern (modulation) of the remaining small field inhomogeneity and
chemical shift phase errors on the phase axis of 2D and 3D
k-space.
III. Phase And Magnitude Corrections Using the Initial
Excitation
Since an initial excitation (to establish a steady state TR) is
preferable anyway, we use this opportunity to acquire a template
data set from which phase error and T2, T2* decay can be determined
for the entire subsequent image data acquisition sequence. Once so
acquired, it is preferably used to correct all such image data for
phase array and for T2 and T2* decay.
IV. Echo Shift
This special GRASE phase encode trajectory may itself be sufficient
to permit successful Fourier Transform image formation. However,
some slight discontinuity in phase error may even then remain in
the data set (because the signals are all recorded at the same
respective temporal position with respect to the Hahn echo time),
leading to a discontinuous or stepped pattern of phase errors (with
the number of steps equal to the number of gradient echoes between
each 180.degree. RF refocusing pulse). By time shifting differently
the temporal position of each group of gradient echoes (with
respect to the Hahn echo time) in a specific pattern, this stepped
pattern can be subdivided into multiple smaller steps. In this
manner, the data set can be forced to have a substantially
continuous linear variation of phase error (throughout k-space) so
that the Fourier Transform has essentially linear data as an input.
Our experience shows that without this step, areas in the body with
fat adjacent to water density tissue may have a band of artifactual
signal parallel to this tissue boundary. This step of echo time
shifting thus finally makes for a substantially completely artifact
free image.
These steps preferably are all utilized together in the exemplary
GRASE technique so as to produce extremely high quality images with
ultra fast imaging speeds. The other known ultra-fast methods of
imaging with dissimilar phase-encoded gradient echoes in long echo
trains (e.g., EPI, MBEST and Instascan) are performed with huge
image artifacts, chemical shift, and signal-to-noise loss which can
only be offset with expensive improvements in gradient system
hardware to reduce these errors. However, these latter techniques
must pay a huge price in image signal-to-noise ratio since they
necessitate very large bandwidth signals and nevertheless do not
overcome field inhomogeneity error and chemical shift problems.
It may be argued that the first exemplary GRASE step was suggested
in the prior 1986 SMRM abstract by Feinberg et al. However, a
different pulse sequence called "inner volume imaging" was actually
there being discussed. In any case, no specific way to combine
gradient refocusing and 180.degree. RF refocusing was described. It
was only stated that the two were somehow to be combined and no
pulse sequence diagram was shown. Furthermore, expert workers in
MRI have not been able to understand the method being described in
this abstract. In fact, the 1986 abstract described a failed
experiment which did not successfully produce human or biological
images. Only images of homogeneous oil or water phantoms could be
made due to the huge errors resulting when oil and water components
were attempted to be imaged together. The Feinberg et al 1986
abstract also only described an imaging method called "inner
volume-echo planar imaging" in which orthogonal planes of
90.degree. RF excitation and 180.degree. RF refocusing (not
parallel, i.e., substantially congruent, volume selective NMR RF
excitations as in exemplary multi-slice GRASE sequences) were used
so that only a smaller subsection image could be produced--and not
an entire cross-sectional image as produced in the exemplary GRASE
technique.
3DFT imaging methods are well-known to be highly demanding on data
acquisition times with the independent encoding of three spatial
dimensions.
The incorporation of 3D spatial encoding with multi-sectional
acquisition has been previously evaluated (Crooks et al) where
lengthy imaging times were required
(T=TR.times.PEy.times.PEz.times.NEX) for 256.times.8 slices per
slab. Thus, for T2 weighted images
T=(2sec).times.(256).times.(8).times.(1) requires 64 minutes. It
has been found possible to obtain contiguous 3D T2 weighted images
in more acceptable imaging times by merging multi-cycle GRASE
imaging methods.
BRIEF DESCRIPTION OF THE DRAWINGS
These as well as other advantages and objects of this invention
will become more fully understood by study of the following
detailed disclosure of exemplary classes of GRASE MRI NMR pulse
sequences, taken in conjunction with the accompanying drawings, of
which:
FIG. 1 is a generalized and simplified schematic block diagram or
an MRI system adapted to perform exemplary GRASE MRI pulse
sequences;
FIG. 2 through FIG. 5 schematically depict various prior art MRI
pulse sequences;
FIG. 6 graphically illustrates the magnitude changes due to T2 and
T2* decay from echo for data obtained without phase decoding;
FIG. 7A schematically depicts a presently preferred type of GRASE
MRI pulse sequence wherein an RF refocused echo train is formed
with the CPMG scheme, and within each RF echo, and multiple lines
of data are acquired using gradient reversals while each data line
is phase encoded differently by G.sub.y gradient pulses (the number
of each signal corresponding to the k-space location in FIG. 7B)
and the entire sequence being repeated with the phase encoding
lobes slightly changed (arrows);
FIG. 7B schematically depicts the k-space trajectory for the
exemplary GRASE MRI pulse sequence of FIG. 7A (with numbers on the
left hand side corresponding to the signal numbers in FIG. 7A) and
showing the trajectory to scan over a substantial portion of
k-space within every RF refocused echo, with a slightly different
(i.e., offset) starting position each time and with multiple
successive excitation cycles thus being interleaved to fill
k-space;
FIG. 7C schematically depicts the filled-in data structure of
k-space for the exemplary GRASE sequence of FIG. 7A (the right hand
side representing the k-space data set, arrows showing the
direction of the data sampling, the left hand side illustrating the
T2 and T2* decay, phase shift due to "chemical shift" and field
inhomogeneity on the phase axis);
FIGS. 8A-8C compare the k-space trajectories of a) the exemplary
FIG. 7A GRASE sequence, b) the RARE sequence, and c) the EPI
sequence (numbers on the left hand side corresponding to the signal
number order as a function of time in the echo train showing that
in EPI and RARE the k-space trajectories are continuously displaced
on the phase axis while in GRASE, the temporal trajectory is
discontinuous and scans over almost the entirety of k-space within
every RF refocused echo, with slightly different offset starting
positions after each 180.degree. RF pulse and with multiple
excitation cycles thereafter filling k-space in an interleaved
manner in both GRASE and multi-cycle RARE albeit both RARE and EPI
move continuously in time along the ky phase axis so that in EPI
the chemical shift and field inhomogeneity errors accumulate
throughout the entire k-space trajectory);
FIG. 9 schematically depicts an exemplary 3D GRASE MRI
sequence;
FIG. 10 schematically depicts an exemplary diffusion weighted GRASE
MRI sequence;
FIG. 11 schematically depicts an exemplary echo time shifting
technique preferably used in all GRASE MRI sequences to further
improve image quality; and
FIGS. 12A-12D further schematically depict the exemplary echo time
shifting technique of FIG. 11 and illustrate the possible impact of
same in k-space and the accumulated phase-shift artifacts both with
and without such echo time shifting.
DETAILED DESCRIPTION OF PREFERRED EXEMPLARY EMBODIMENTS OF THE
INVENTION
A typical MRI system is schematically depicted in block form at
FIG. 1. Here a large main magnet 50 (e.g., a cryogenic or resistive
electromagnet or combination thereof or, especially in lower field
MRI systems, a permanent magnet structure) is used to constantly
generate a substantially homogeneous background magnetic field
B.sub.0 throughout the patient image volume 52. Typically, B.sub.0
is aligned parallel to the z-axis of the usual Cartesian coordinate
system as is also depicted in FIG. 1. Conventional magnetic
gradient coils 54 are typically located inside the main magnet
structure and are independently controllable so as to rapidly
produce one or more desired magnetic gradients in the background
magnetic field within the patient image volume 52 along any of the
orthogonal coordinate directions (e.g.. typically referred to as
magnetic gradients G.sub.x, G.sub.y, G.sub.z). Closely coupled to
the patient image volume 52 are RF coils 56 (perhaps a single coil
for both transmit and receive operations or perhaps different coils
for different RF signaling purposes) which permit the transmission
and/or reception of NMR RF signals to and from NMR nuclei within
the patient image volume 52.
These components are conventionally communicated to an MRI RF and
system control 60 (most of which is typically located in a separate
room outside the shielded gantry room of the main magnet 50).
Either as a part of the MRI RF and system control 60 or as a
separate subset of the MRI computer systems, image reconstruction
and processing apparatus 62 is also provided to process the raw
acquired RF NMR signal responses (typically after suitable RF
signal conditioning and digital sampling) so as to produce visual
images on the CRT screen of a control console 64 (or otherwise to
create visualizable digitized images in magnetic, silicon,
photographic film, or other visual display media). An operator may
typically control the whole MRI system operation (including choice
of particularly desired MRI pulse sequences) from the keyboard of
console 64.
Also schematically depicted in FIG. 1 is the control program memory
including a suitable GRASE control program for effecting the
exemplary GRASE MRI pulse sequences hereinafter described. Those
ordinarily skilled in the complex art of MRI should be able to
routinely reduce the exemplary GRASE MRI pulse sequence(s) to a
suitable computer control program for any particular MRI system
without the need for explicit disclosure of such exemplary
programming in the form of program code, flowcharts, or the like.
The memory 66 in FIG. 1 may comprise any conventional program
storage media such as magnetic disc, magnetic tape, silicon storage
media and the like. Upon selection via operator control at console
64, the GRASE control program in memory 66 will be called up and
executed by the MRI RF and system control 60 so as to produce the
desired sequence of RF NMR nutation pulses and magnetic gradient
pulses for the exemplary GRASE pulse sequence(s) described
hereinafter.
The typical prior art MRI pulse sequences depicted in FIGS. 2-5
have already been described above and thus need no further
description here. Suffice it to say that they are considerably
different from the exemplary GRASE pulse sequences depicted in
subsequent figures and text of this patent application.
One exemplary GRASE pulse sequence in accordance with this
invention is shown in FIG. 7A. The selected slice is first excited
by a 90.degree. RF nutation pulse, and a train of spin echoes
thereafter is generated by two or more 180.degree. RF pulses. For
each RF refocused spin echo, multiple gradient echoes are generated
by rapidly switching the polarity of the readout gradient. Each
echo that has been generated this way is phase encoded differently
by phase encoding and decoding pulses at the beginning and the end
of each gradient echo train and short blips between each gradient
echo. The k-plane representation of this particular exemplary phase
encoding scheme is illustrated in FIG. 7A and 7B. To implement in
current commercial scanners, the sequence should be repeated with
the phase encoding gradients slightly offset in each excitation
cycle, so that the entire k-plane is eventually covered in
interleaved fashion. However, if faster hardware is available, it
may also be possible to scan a full image in one-shot (i.e., within
one TR interval involving only one initial 90.degree. NMR
excitation).
The GRASE pulse sequence takes advantage of both RF refocused
echoes and gradient refocused echoes. Some advantages of using
gradient echoes (e.g., advantages over a RARE sequence) are
It is faster. Since there is less overhead of refocusing RF pulses,
the required time per echo is shorter than for a RARE sequence.
There is lower power deposition. One of the problems with a RARE
sequence is the high RF power deposition to the patient. Our GRASE
sequence uses fewer RF pulses than RARE, and the SAR value
therefore can be reduced to the current FDA guideline.
Some advantages of using RF refocused echoes (e.g., advantages over
an echo planar sequence are
There are fewer artifacts. In echo planar imaging, field
inhomogeneity causes a spatial distortion in the acquired image. By
using refocusing RF pulses, our GRASE method is less sensitive to
field inhomogeneity.
In the exemplary GRASE sequence, the number of the RF refocused
echoes per excitation and the number of the gradient echoes per
refocusing RF pulse are adjustable to optimize the sequence to a
particular application. For example, with 8 RF echoes per
excitation and 4 gradient echoes per refocusing RF pulse (G4-R8),
32 echoes can be collected per excitation using RF echo interval of
20 ms and the sampling window of 2 ms, which is achievable with
current commercial scanners. With a TR of 2000 msec, 256.times.256
multi-slice images can be obtained in 16 seconds.
A small timing error or a gradient error can cause a phase error in
the reconstructed images. Usually magnitude images are used and
this type of phase error does not therefore affect the image.
However, if the amount of phase error is different for each echo
(which is possible in a GRASE sequence), a ghost artifact will
result even if only a magnitude image is obtained. Also, the
magnitude of each echo varies due to the T2 decay and the T2*
decay. This also causes a ghost-like artifact.
These errors can be at least partly corrected by following steps
using a probe data set which is obtained with the phase encoding
gradient pulse turned off. This can be done with the first
excitation, which is used to establish a steady state, without
increasing the scan time.
Correction steps
1. Template data collection
Collect NGE.times.NSE "template" data set with phase encoding
pulses turned off. This can be done with the first 90.degree. RF
nutation excitation of the imaging sequence. Since an additional
excitation at the beginning is necessary anyway for establishing a
steady state prior to the first measurement of image data, this
template data collection does not require additional scan time.
From this template file, three kinds of information can be
obtained, i.e. phase error, T2 and T2* decay.
2. Phase correction
Phase error mainly comes from various hardware timing errors. Two
major components of this are a constant phase error within an echo
and an echo position error. After Fourier transformation, or in the
reconstructed image, these become zeroth and 1st order (one which
is constant over the image, and one which linearly varies with
spatial position). Zeroth order error can be estimated by
where si is the complex value of the i-th pixel of data after 1D
transformation. First order error can be estimated by
where s is the complex value of the i-th pixel of data after 1D
transformation and s.sub.i * is the complex conjugate of s See C.
B. Ahn and Z. H. Cho, IEEE Trans. Med. Imag. MI-6, 32 (1989).
Zeroth and first order phase errors can be corrected by
This step is done after 1DFT. This correction is done for each of
NGE=NSE echoes. Echoes with the same relative timing with respect
to the 90.degree. RF pulse but in different excitations can be
corrected using the same template data. Additional 1DFT in the
other direction will give reconstructed images.
3. T2 and T2* correction
Since the template data are obtained without phase encoding, each
echo in the data set corresponds to the same DC line in k-space,
and should have the same amplitude. However, as shown in FIG. 6,
the magnitude changes due to T2 and T2* decay from echo.
The three curves in FIG. 6 correspond to three gradient echo
groups. The central group corresponds to Hahn spin echoes and the
curve reflects T2 decay. The difference between the echo groups
comes from T2* decay. This magnitude modulation causes artifacts in
the image, seen as blurring, ghosts and ringing depending on the T2
and T2* values. This can be mostly corrected by scaling the raw
data inversely (K. Oshio and M. Singh, magn. Reson. Med. 11, 389
(1989). However, this amplifies noise since the procedure includes
dividing by small numbers.
It is also possible to use only one polarity of the read out
gradient for signal recording to eliminate such correction steps.
By using two gradient echoes on either side of TE, the T2* and T2
decay will be similar and the gradients will be of the same
polarity so that time reversal is not necessary. This is at the
expense of not using one of the three potential signals after each
180.degree. RF pulse, see FIG. 7A.
Numerous phase encoding schemes can be used with our imaging
method. For example, the data can be phase encoded differently on
each successive signal so that k-space has a substantially
continuous T2 decay across it. Alternatively, two or more images
can be generated with different T2 weighting, equivalent of a first
and second echo image, by making redundant phase encoding during
the first and second half of the echo train, and grouping the data
separately for 2DFT reconstruction.
A different phase-encoding scheme is used in the preferred
exemplary GRASE pulse sequence of FIG. 7A. An RF refocused echo
train is formed using the CPMG scheme
(90.degree.-180.degree.-180.degree.- . . . ), with a number of spin
echoes (NSE) and three or more (i.e., a number NGR) gradient
recalled echoes being created, centered about each Hahn spin echo
time. Therefore the total number of signals per echo train is the
product of the number of RF refocused echoes NSE and the number of
gradient recalled echoes per RF echo NGR. In our exemplary
implementation, NSE is eight and NGR is three, giving the total of
twenty-four signals per excitation. Eight excitations cover
256.times.192 data points of k-space. The standard multi-section
excitation scheme was performed using slice selective RF
excitations and frequency offsets for multi-slice imaging.
In EPI, without the use of multiple 180.degree. RF pulses, the
chemical shift and field inhomogeneity errors become much larger as
they evolve over the entire echo train.
Each signal is phase encoded differently to scan twenty-four lines
in k-space by phase encoding pulses that precede each signal. The
phase encoding is returned to zero before the next RF refocusing
pulse in a way similar to that used in the RARE sequence.
It is apparent that the effect of T2* and chemical shift differ
throughout each gradient echo train. That is, these differences
reoccur with the periodicity of each 180.degree. RF pulse interval.
In short, moving temporally from echo to echo, the T2* and chemical
shift changes are modulated. Therefore, if this echo train were to
trace out a continuous incremental path through the phase encoded
axis of k-space, there would be a rapid modulation of chemical
shift and T2* changes along the phase axis of k-space. After 2D
Fourier Transform, this would result in ghosting artifacts in the
image. In EPI, the phase encoding is continuously incremented
during the echo train, resulting in an incremented displacement
along the phase axis k-space as depicted in FIG. 8B.
To demodulate and effectively remove such chemical shift
modulation, the k-space trajectory is purposefully further
modulated with the same periodicity (i.e., as the chemical shift
modulation periodicity). By reordering these signals to be
continuous in phase order (rather than in time of acquisition
order), the chemical shift variations may be mapped into a
continuous variation across the entire phase axis of k-space. On
the orthogonal frequency axis of k-space, chemical shift occurs
across only the sample window time, which is small in comparison to
the phase axis time intervals.
Specifically what occurs in the exemplary GRASE sequence of FIG. 7A
is that the phase encoding sweeps through each of three large
increments of k-space during the three respectively corresponding
gradient recalled signals as depicted in FIG. 7B or 7C. The next
set of three signals (i.e., in the next RF refocused echo and still
within the same FID) have an identical k-space trajectory except
for being displaced by a much smaller increment of phase encoding.
Similarly, the phase of each subsequent group of three echoes is
displaced by a small additional increment of gradient offset, so as
to eventually fill in all adjacent lines of k-space.
After one complete excitation cycle, 24 equally spaced lines in
k-space are obtained. The entire k-space is covered by repeating
the excitation cycle with incremental one pixel offsets
accumulating in the phase encoding. The phase encoding amount,
k.sub.y for 1-th excitation, m-th RF echo and n-th gradient echo
can thus be expressed as
where L and M are the total number of 90.degree. RF excitation
cycles and number of 180.degree. RF refocusing pulses,
respectively.
After reordering the signals to have continuous phase increments in
k-space, the modulation pattern for T2, T2* and the phase shifts
due to chemical shift and field inhomogeneity becomes as shown in
FIG. 7C. The right hand side shows the reordered k-space
trajectories, while the left hand side illustrates (a) the
modulation pattern of the T2 and T2* decay, and (b) the phase shift
due to "chemical shift" and field inhomogeneity. In summary,
chemical shift and field inhomogeneity errors are reduced in
proportion to the shorter time of each gradient echo train, while
the novel k-space trajectory eliminates or reduces modulation of
these remaining errors.
The three large bands depicted in k-space correspond to the three
gradient echoes as indicated. The T2* decay and phase shifts occur
only between these groups of signals, and not within them. However,
the amount of T2 decay varies depending on the position of the RF
echo within the entire echo train. The T2 decay is really an
exponential decay but effectively becomes a stepped decay due to
the multi-excitation scheme, since the T2 decay and phase shifts of
an echo are identical for each of the excitations.
The imaging time of GRASE technique can be directly expressed as
T=TR.times.(NL.times.NEX)/NGR.times.NSE) where NL is number of
image lines, NEX=number of excitations, NGR=number of gradient
refocused signals per spin echo, NSE=number of RF refocused spin
echoes. For a TR of 2 seconds, 192 image lines, 3 NGR and 8 NSE,
for T of 16 seconds. One initial excitation is used for
establishing steady state of TR 2 second, giving total imaging time
of 18 seconds.
The exemplary GRASE sequence of FIGS. 7A-7C was implemented on a
1.5 T system (G.E. Signa) using maximum gradient strength of 1
gauss/cm, 4 msec read-out period and 3.2 msec selective RF pulse.
The interval between 180.degree. pulses was 23 msec. The effective
TE (the time at which the origin of the k-space is sampled) was 104
msec. The data acquisition time was 9.times.TR where the initial
excitation establishes steady state. By producing the first
excitation without phase encoding, a template correction file
results. This correction file is used for odd/even gradient echo
phase shifts, and T2 and T2* magnitude filtering so as to correct
for the modulation of T2 and T2* decay (e.g., see Oshio et al,
Magn. Reson. Med. 11, 389 (1989).
A typical GRASE image of a human brain had a field of view of 24
cm. No fat suppression pulses were used. The image of fat within
the skin was centered symmetrically about the brain with no
relative displacement from the water component. The image had
essentially the same tissue contrast as a spin echo image. Many
small vessels, dark curved linear structures were resolved in the
image. There was also a noted absence of ghost artifacts, spatial
distortion due to the field inhomogeneity, or chemical shift in the
fat signal of the skin.
It is significant that this implementation of GRASE imaging did not
require gradient hardware improvements nor static magnetic field
improvements over currently available commercial imaging systems.
The total data acquisition time of 36 seconds for TR of 4000 msec,
in this particular application, produced 22 multi-sectional 5 mm
thick images with high resolution and T2 weighting of clinical
utility.
It is also important to note that except for signal bandwidth
differences, the high speed of GRASE imaging does not result in
signal to noise loss, unlike the method of conjugate synthesis
(Half Fourier, NEX=1/2), which nevertheless could be combined with
GRASE imaging. Unlike RARE, the GRASE sequence uses a fewer
relative number of time consuming selective 180.degree. RF pulses.
This more efficient utilization of the total possible signal period
gives higher S/N per imaging time as well as a lower SAR. A modest
increase in the relative number of gradient echoes to RF pulses,
for example 5 or 7 instead of 3, in principle would permit further
improvement in imaging speed or alternatively used to increase S/N
in images. The effect of such longer gradient echo trains on T2
contrast is not known at this time.
The tissue contrast in GRASE is essentially the same as that of
spin echo imaging, as demonstrated by a human head image. This can
be understood from the pulse sequence diagram and FIGS. 7B and 7C
which show that the second gradient recalled echo is at the Hahn
spin echo position. These Hahn spin echoes cover the center of
k-space which substantially determines the main image contrast. The
effective TE is the center of the entire echo train in this
particular experiment.
With respect to clinical body imaging, breath holding during 18
second imaging times for TR of 2 seconds will permit T2 weighted
imaging in absence of respiratory motion artifacts. Several
modifications of GRASE imaging, for double echo imaging, multi-slab
3D volume imaging and 512.times.512 images can be performed in
clinically acceptable imaging times. Our results show that by
interposing 180.degree. RF pulses between gradient echo trains and
by performing a novel k-space trajectory, very fast MR imaging with
high spatial resolution is achieved without significant
artifacts.
Head and body images of normal volunteers have been made in several
imaging sessions during recent development of the GRASE technique.
Two patients with known radiologic and clinical diagnosis of
multiple sclerosis were imaged with GRASE during their routine T2
weighted spin echo studies.
The GRASE imaging was again implemented on a 1.5 Tesla MR system
(G.E. Signa) using maximum gradient strength of 1 gauss/cm, 4 or 2
msec readout periods with 23 or 18 msec interval between
180.degree. RF pulses, respectively. The effective TGE (the time at
which the origin of the k-space is sampled) was varied from 80 msec
to 104 msec. The data acquisition time was 9.times.TR where the
initial excitation establishes steady state.
A typical GRASE head study using TR=4 sec., NEX=1 in 36 seconds had
image quality sufficient to demonstrate the optic nerve as well as
many small radial vessels in the white matter of the neocortex. The
CSF, gray and white matter had high contrast due to the long TR and
long TE, 4 sec and 104 msec, respectively.
In a different study, a comparison of tissue contrast was made in
the brain between GRASE spin echo, and RARE images with similar
TR=2.5 second and TE of 104-108 msec. There was similar grey and
white matter and CSF contrast as seen on two representative levels.
In the RARE image, the signal from fat in the skin is much higher
than the similar signals in GRASE and spin echo imaging. The flow
artifact in the NEX=178 spin echo image was not present in GRASE or
RARE at the same level.
One representative GRASE image of a patient with multiple sclerosis
had MS plaque well demonstrated in the right frontal lobe white
matter.
Coronal GRASE images of the abdomen were obtained in 18 seconds, TR
2 seconds, and TE 80 msec for 11 sections. No breathing respiratory
motion artifact was seen in these images since imaging was
performed in a single breath hold. Renal arteries branching from
the aorta were demonstrated without the motion artifact often
present in spin echo images. The contrast between liver, spleen and
kidney were as expected for spin echo T2 weighted images.
Sagital GRASE images of the lumbar sacral spine demonstrate the
myelographic effect of CSF when TR was 3 seconds. There was central
disc bulge, greatest at L2-3 disc space and decreased signal from
the L4-5 disc likely from fat loss, all typical features of
degenerative disc disease.
The development of CT technology for abdominal imaging, not unlike
MR, required a large leap in speed to move from the early head
images of 3 minute acquisition times to the present scan times of 3
to 4 seconds. CT body imaging is now of great clinically utility
while MRI has been to date unable to overcome the hurdle or
respiratory motion and organ peristalsis which have significantly
degraded T2 weighted spin echo imaging which otherwise has great
sensitivity to abdominal pathology. T2 weighted GRASE images were
acquired without respiratory motion using a breath hold of 18
seconds, which most patients can perform. This has required
reduction in imaging time by a factor of 25 from spin echo imaging,
NEX=1.
The tissue contrast of GRASE in theory and practice is very close
to that of traditional spin echo imaging since it is determined
predominantly by the central gradient echoes, which in fact are
Hahn spin echoes. These Hahn spin echoes cover the central portion
of k-space where the strong Rst image signals occur. Therefore,
unlike low flip angle images of the brain which have decreased
sensitivity to pathology, MS plaques can be detected using GRASE
similar to spin echo imaging. Two patients were imaged with both
spin echo and GRASE and the number of MS plaques were compared at
identical levels in the brain by a neuroradiologist who found that
GRASE imaging detected all of the plaques shown in the spin echo
image.
It is significant that GRASE imaging uses fewer 180 RF pulses per
TR than RARE, with a large corresponding reduction in SAR. Further
improvements in GRASE imaging time and reduction in SAR are being
investigated using NGR=7, yielding 12 second imaging studies,
holding constant signal bandwidth and S/N. RARE imaging methods
provide a maximum factor of 16 reduction in image time from spin
echo imaging. The current SAR limitation prevents excitation of the
maximal permissible number of slices with RARE. The speed advantage
of GRASE over RARE can be expressed in terms of average signal time
using exemplary time values of selective 180.degree. RF excitation
including FID spoiler gradients (Trf)=8 msec, phase encode and
rephasing pulses (Tpe)=4 msec and readout gradient (Tro)=4 msec, a
simple expression can be calculated,
For RARE with NGR=1, the average time (8+4+4)/ is 14 msec per
signal equivalent to the time between 180.degree. RF pulses, For
GRASE with NGR=3, the average time (8+4+12)/3 is 8 msec per signal.
And for NGR=7, the average time per signal is 5.8 msec. The effect
of increased NGR on T2 contrast is not known at this time. However,
such changes would lead to an increase in chemical shift, which is
currently less than one pixel in both head and body GRASE
images.
It is important to realize that GRASE imaging is a form of multiple
spin echo imaging and therefore does not have the artifacts of
gradient echo images. Gradient echo imaging does not use
180.degree. RF pulses so low S/N occurs due to non-cancellation of
field inhomogeneity errors particularly in body regions far from
the magnet center. These field inhomogeneity errors continuously
increase with time between the RF excitation (90.degree. or less
than 90.degree. in GRASS) and the signal refocusing time, TE. In
spin echo imaging the 180.degree. RF pulses reverse the position of
the spins, leading to cancellation of field inhomogeneity errors at
the Hahn spin echo time.
In comparison, GRASE imaging produces gradient echoes in the time
envelope of each RF refocused spin echo. Field inhomogeneity errors
evolve during the relatively short time between the gradient
recalled echo and the center of the respective spin echo envelope,
not the total time from the 90.degree. RF excitation pulse as in
gradient echo imaging. In the current implementation of GRASE when
NGR is 3 and readout gradient is 4 msec, the field inhomogeneity
error accumulates during a 15 msec period, while the total echo
train extends to 208 msec. In effect, GRASE imaging, like EPI,
utilizes the speed advantage of gradient refocusing techniques, but
without paying the large penalty in chemical shift and field
inhomogeneity errors.
With respect to clinical body imaging, T2 weighted imaging in
absence of respiratory motion artifacts is possible given imaging
times under 20 seconds. There are several apparent modifications of
GRASE imaging such as different TE and TR for desired image
contrast as well as double echo imaging, multi-slab 3D volume
imaging and 512.times.512 images in clinically acceptable imaging
times.
Thus, GRASE imaging can accomplish a T2 weighted imaging with
similar contrast as spin echo. These images have high spatial
resolution, are obtained in times fast enough to overcome
respiratory motion artifacts, and do not produce an excessively
high SAR in the body. Currently we are extending body imaging with
GRASE to use 5 gradient refocusings per 180 pulse, NGR is 5, to
reduce acquisition time from 18 seconds down to 12 seconds with no
change in image quality or S/N except for an acceptable increase in
chemical shift. It is possible that by combining the GRASE
techniques with high performance gradient systems, similar
excellent image quality will be obtained in even faster imaging
time, using 3 or 4 excitations.
A multi-slab 3D GRASE pulse sequence is depicted in FIG. 9.
Selective excitations of M multiple slabs are performed in the
standard multi-section scheme during each TR cycle. Each slab is
phase encoded for inplane resolution G.sub.y by varying the phase
encoding pulses during each echo train (GRASE). Each TR cycle then
varies the phase encoding of the inplane axis and additional TR
cycles phase encode the G.sub.z slice select axis. This permits the
acquisition of N slices/slab.times.M slabs. Alternatively, both
G.sub.y and G.sub.z could be varied during each GRASE excitation as
well as during the TR cycles so that the number of slices/slab can
be varied.
In most 3DFT image methods, the outer slices are of lower image
quality, due to imperfections in the perfectly rectangular). The
outer slices can be overlapped spatially and discarded to obtain a
total of 60 continuous 1.5 mm sections. Application of 0.5 mm
inplane resolution with 512.times.512 matrix display can also be
incorporated in this imaging scheme. The application of phase
encoding during trains of GRASE refocused echoes reduce the total
imaging time by the same factor, 64/NGR.times.NSE min or
64/24=approximately 3 minutes for 60 thin T2 weighted images. The
contiguous 1.5 mm sections of the body are not deteriorated in S/N
given the intrinsic increased signal averaging of 3DFT. However, we
have also found a higher S/N imaging time for the more efficient
signal production in echo train phase encoding as compared to
single spin echo encoding. This more than offsets the losses of S/N
with our wider signal bandwidths.
Diffusion weighted GRASE is depicted in FIG. 10. The use of
multiple pairs of gradient pulses, the Stecjkel-Tanner sequence,
can be incorporated into GRASE to great advantage for diffusion
weighting. This is useful for measuring the thermal changes in
tissue when laser surgery is being guided by MRI. (See Perfusion
Imaging by Feinberg and Jacob for further explanation and relevant
mathematical equations).
Echo Time Shifting is depicted in FIG. 11 which shows how the use
of NGR=3 gives three discrete sample points on the assumed chemical
shift phase variation curve. If, instead, the phase axis is
continuously sampled throughout the k-space data set, the image
will give a standard spatial shift between the water and fat
components (chemical shift). Note that the amount of chemical shift
is still greatly reduced from that of echo planar methods since the
phase errors evolve over a shorter time interval of NGR and are
then refocused by the next 180.degree. RF nutation degree pulse and
repeated. In EPI, without the 180.degree. RF refocusing pulses,
field inhomogeneity errors continue to accumulate to give a
proportionally larger chemical shift.
FIG. 11 shows diagrammatically how each group of NGR (e.g., the
group of 3 gradient echoes between each pair of 180.degree. RF
pulses) is shifted in time by a small time increment. Each such
group of gradient echoes then occur at different points on the
phase error cure of periodically recurring chemical shifts. These
groups of signals are then interleaved in k-space, as shown
diagrammatically for three spin echoes in FIG. 12c (each spin echo
is really a spin echo envelope of NGR signals). This method of
shifting the relative temporal position of each group or NGR
signals results in a more complete incremental sampling of the
chemical shift phase curve so that the Fourier Transform can be
provided with substantially linearly changing data and give simply
a chemical shift greatly shortened from EPI. Without this echo time
shift process, the sharp interfaces between water and fat may
produce artifacts seen as a regional spread of the fat signal onto
the tissues of more water composition.
It should be noted that this small time shifting of echoes between
180.degree. RF pulses does not change any of the tissue
characteristics in the image--it does not change the T2, spin
density of T1 contrast in the image.
FIGS. 12A-12C depict how three different echo time shifts may be
employed to evenly distribute the unavoidable (but relatively
small) chemical shift occurring between pairs of 180.degree. RF
pulses over the entire phase-encoded dimension of k-space (FIG.
12c) as opposed to incurring three discontinuous jumps in chemical
shift (FIG. 12D) if such time shifting is not employed. In effect,
to avoid large phase jumps (which can cause artifacts) one can
effectively sample continuously on the phase shift function of
chemical shift by moving the sample windows and echo signals on the
time scale. This may require one extra readout period per
180.degree. RF pulse to prevent overlap with other gradient
pulses.
In effect, the echo shift described in FIGS. 11 and 12A-12D varies
the position or gradient echo trains with respect to Hahn echo
times in the CPMG sequence to eliminate phase errors due to
magnetic field inhomogeneity and chemical shift. While GRASE image
studies of the abdomen, similar to brain studies, have minimal
spatial distortions and high resolution, abdominal studies show
artifactual image modulations at tissue interfaces with chemical
shift. This novel method similarly improves image quality in
multiple excitation echo-planar imaging.
The basic GRASE technique by itself substantially reduces both
chemical shift and the effect of static magnetic field
inhomogeneity including spatial distortions and T2* dependent
signal loss in images. The reduced chemical shift has permitted
simultaneous imaging of both the water and fat tissue components in
T2 weighted GRASE imaging, similar to spin echo imaging.
However, in GRASE, the remaining smaller field inhomogeneity
effects are manifested as a periodicity during the echo train.
Given that there is no change in the position of the gradient echo
trains with respect to the 180 refocusing or Hahn echo position,
the phase shifts resulting from field inhomogeneity effects are
repeatedly experienced at each respective gradient echo time. When
these gradient echo signals are interleaved on the phase axis of
k-space, their field inhomogeneity phase shifts produce a stepped
pattern, with the number of steps equal to the number of gradient
echoes.
This small non-linearity across the phase axis of k-space is
produced in lieu of a more destructive periodic modulation of phase
shifts without interleaving the signals. Such periodic modulations
are transformed by FT into a convolution of the image with a
periodic function resulting in severe ghosting or replications of
the object structure. A much less severe an artifact is produced by
this residual stepped modulation, seen in body images as bands
parallel to the sharp interface of lipid and water tissues. This
artifact is not present in images of brain tissue having no such
sharp interfaces.
The GRASE sequence timing diagram is shown in FIG. 12A for
simplicity without the phase encode gradients or slice selective
gradients. The k-space index os each signal is shown numerically,
1..9, for this representative sequence of three gradient echoes and
three spin echoes. Using gradient echoes symmetrically centered on
the Hahn echo time, the field inhomogeneity phase is sampled at
three time positions, leading to the three steps of phase shift in
the k-space data. The signals can be described in terms of
M(r)-magnetization, G(t)-magnetic field gradients and r-spatial
position as,
where E(r) is the field component due to homogeneities including
chemical shift and static field. Rewriting eq. (1) in terms of 01
phase encoding and 02 phase error, ##EQU1## where i is the phase
gradient echo index, .DELTA.t is the duration of each gradient echo
period as defined by T-total time of the gradient echo train and
N.sub.GE --total number of gradient echoes.
The time of the gradient echoes with respect to the field
inhomogeneity phase shift curve, can be changed simply by shifting
the position of the gradient refocusing to be slightly off-center
to central the Hahn echo time, dashed lines in FIG. 12A. By
imposing the appropriate increments of gradient echo time shifting,
the phase shift curve is continuously sampled by the subsequent
spin echo envelopes in the echo train, FIG. 11. Similarly, gradient
echo time shifting is performed during multiple excitations of the
GRASE sequence, for 1-th excitation, m-th RF echo and n-th gradient
echo, where L and M are the total number of excitations and RF
refocusing, respectively. The phase shift and phase error are
rewritten for the gradient echo shifted sequence, ##EQU2##
Using these redefined phase shifts, the signal at each ky position
can be expressed as
In this final form E(x,y) is continuously sampled along the ky
axis, resulting in a linear phase shift prior to FT. As a result,
an expected spatial displacement or normal chemical shift will
occur between water and lipid on the phase axis of the image,
rather than the modulated fat-water interface as described
above.
In the GRASE sequence, the index i for 1-th excitation, m-th RF
echo and n-th gradient echo is expressed as
where L and M are the total number of excitations and RF
refocusing, respectively.
While our experiment was designed specifically for the GRASE
sequence, this method can be applied to less complicated image
sequences which do not use multiple spin echoes. Gradient echo
trains which are repeated with multiple excitations to fill out
k-space will experience field inhomogeneity phase shifts similar to
GRASE.
A multiple excitation or "partial" echo planar imaging method
results with interleaving of the signals on the phase axis when M
is set to zero and m=1. The gradient echo time shift becomes,
where M=m=1.
To study the effect of gradient echo time shift in biological
tissues, GRASE imaging was performed in the abdomen of a normal
volunteer. The banding artifact at the boundary of the kidney and
adjacent fat is replaced by chemical shift. The absence of
respiratory motion artifacts with breath holding during the
relatively short 16 second data acquisition time is not possible
with the twenty-four times longer scan time of spin echo
imaging.
The field inhomogeneity errors and chemical shift in GRASE imaging
are reduced in three steps. First, the CPMG sequence is applied to
refocus the phase shifts due to field inhomogeneity, reducing the
net phase errors of gradient echoes by a factor equal to the number
of Hahn spin echo envelopes encoded. Secondly, the temporal
periodicity of phase errors is eliminated by similarly imposing
periodicity in the spatial phase encoding trajectory and reordering
the signals in k-space. Thirdly, the gradient echo time shifting
further demodulates the remaining phase errors to be continuous on
the phase axis. The resulting images of the human body are obtained
with minimal, 2 pixel, chemical shift between fat and water
components. Finally, cardiac gating, synchronizing the pulse
sequence to the heart cycle, further eliminates artifacts. The
small displacements of the right kidney and liver due to cardiac
driven pulsations are therefore removed from the image.
The partial echo planar methods similarly are improved with use of
the gradient echo time shifting. Others have suggested shifting the
90.degree. RF pulse timing, to similarly linearize the field
inhomogeneity effects in partial echo planar imaging. However,
these later methods would change the Hahn echo time and destroy the
CPMG sequence in GRASE.
The additional time required for gradient echo shifting equals time
of one gradient echo plus the gradient rise times. In the GRASE
sequence, 3.2 msec, is expended during each of eight spin echo
periods, for a net trade-off of 12.8 msec longer TE in these
experiments. In our experience, GRASE images of the brain do not
require gradient echo time shifting, given there predominant water
tissue composition. The free lipid components in the surrounding
scalp produces no detectable artifact (1). In body imaging, the
fat-water tissue interface is accurately defined with linear
sampling of field inhomogeneity factors, made possible with these
new methods.
Although the exemplary drawings depict conventional substantially
rectangular gradient pulses, other pulse shapes could be used
instead. For example, sinusoidally shaped gradient pulses (with
appropriate interpolation in k-space) could be used in accordance
with the teaching of Rzedzian (U.S. Pat. No. 4,818,942).
While only a few exemplary embodiments of this invention have been
discussed in detail, those skilled in the art will recognize that
many variations and modifications may be made in these exemplary
embodiments while yet retaining novel features and advantages of
this invention. Thus, it is intended that the appended claims cover
all such variations and modifications.
* * * * *