U.S. patent application number 11/196131 was filed with the patent office on 2006-02-16 for rf coil for imaging system.
This patent application is currently assigned to The General Hospital Corporation d/b/a Massachusetts General Hospital, The General Hospital Corporation d/b/a Massachusetts General Hospital. Invention is credited to J. Thomas JR. Vaughan.
Application Number | 20060033501 11/196131 |
Document ID | / |
Family ID | 22467324 |
Filed Date | 2006-02-16 |
United States Patent
Application |
20060033501 |
Kind Code |
A1 |
Vaughan; J. Thomas JR. |
February 16, 2006 |
RF coil for imaging system
Abstract
An RF coil suitable for use in imaging systems is provided which
coil has a dielectric filled cavity formed by a surrounding
conducting enclosure, the conducting enclosure preferably being
patterned to form continuous electrical paths around the cavity,
each of which paths may be tuned to a selected resonant frequency.
The patterning breaks up any currents inducted in the coil and
shortens path lengths to permit higher frequency, and thus higher
field strength operation. The invention also includes improved
mechanisms for tuning the resonant frequency of the paths, for
selectively detuning the paths, for applying signal to the coil,
for shortening the length of the coil and for controlling the field
profile of the coil and the delivery of field to the object to the
image.
Inventors: |
Vaughan; J. Thomas JR.;
(Stillwater, MN) |
Correspondence
Address: |
SCHWEGMAN, LUNDBERG, WOESSNER & KLUTH
1600 TCF TOWER
121 SOUTH EIGHT STREET
MINNEAPOLIS
MN
55402
US
|
Assignee: |
The General Hospital Corporation
d/b/a Massachusetts General Hospital
|
Family ID: |
22467324 |
Appl. No.: |
11/196131 |
Filed: |
August 3, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10750031 |
Dec 29, 2003 |
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11196131 |
Aug 3, 2005 |
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10367489 |
Feb 14, 2003 |
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10750031 |
Dec 29, 2003 |
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09575384 |
May 22, 2000 |
6633161 |
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10367489 |
Feb 14, 2003 |
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60135269 |
May 21, 1999 |
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Current U.S.
Class: |
324/322 ;
324/318 |
Current CPC
Class: |
G01R 33/422 20130101;
G01R 33/34007 20130101; G01R 33/34046 20130101; G01R 33/3657
20130101; G01R 33/3635 20130101; G01R 33/3642 20130101; G01R
33/3453 20130101; G01R 33/5659 20130101; G01R 33/345 20130101; G01R
33/5612 20130101; G01R 33/3628 20130101 |
Class at
Publication: |
324/322 ;
324/318 |
International
Class: |
G01V 3/00 20060101
G01V003/00 |
Claims
1. A magnetic resonance imaging system comprising: (a) a housing
providing a medical diagnostic chamber for a subject therewithin
lying along an axis; (b) a transmit/receive inductor system about
said axis in proximity with said housing; (c) a gradient inductor
system operatively associated with said transmit/receive inductor
system; (d) a static magnetic field inductor system operatively
associated with said transmit/receive inductor system; (e) said
transmit/receive inductor system constituting a coil having an
outer surface about said axis and including a series of electrical
transmission line elements paraxially distributed with respect to
said axis about said subject, each of said transmission line
elements including an outer conductor and an inner conductor, said
inner conductor being spaced from said outer conductor in a
direction perpendicular to said outer surface; (f) said coil
initially transmitting to said subject fields of radio frequency
energy as a transmit signal, and responsively receiving from said
subject fields of magnetic resonance energy as a receive signal;
(g) said gradient inductor system initiating perturbations in said
fields and producing signals derived responsively from said
perturbations; (h) said signals corresponding to spatial indicia
derived from said subject.
2. The magnetic resonance imaging system of claim 1 wherein said
coil establishes concentrations of electromagnetic fields among
said transmission line segments.
3. The magnetic resonance imaging system of claim 2 wherein, by
adjusting the distance between said transmission line segments, the
interaction of the magnetic fields of said transmission line
segments with an external sample can be controlled and optimized
for nuclear magnetic resonance signal generation and/or
detection.
4. The magnetic resonance imaging system of claim 1 wherein said
plural transmission line segments decrease the inductance of each
line segment and minimize the electric fields associated therewith,
whereby dielectric tissue losses is said subject are reduced.
5. The magnetic resonance imaging system of claim 1 wherein said
plural transmission line segments have inherent shielding, whereby
coupling between said transmission line segments is controlled.
6. The magnetic resonance imaging system of claim 1 wherein said
plural line segments are combined to optimize NMR signal generation
and/or reception.
7. The magnetic resonance imaging system of claim 1 wherein signals
form said plural line segments are combined to decode spatial
information derived from the NMR signal, thereby to increase the
sensitivity and speed of data acquisition.
8. The magnetic resonance imaging system of claim 1 wherein said
inductor consists of N transmission line segments arranged in a
geometric pattern in which said line segments are substantially
equidistant from each other.
9. The magnetic resonance imaging system of claim 1 wherein said
geometric pattern is circular or elliptical.
10. The magnetic resonance imaging system of claim 1 wherein said
geometric pattern is flat or curved.
11. The magnetic resonance imaging system of claim 1 wherein each
of said transmission line segments includes at least two individual
conductors together with additional lumped or distributed
capacitive or inductive circuit components.
12. The magnetic resonance imaging system of claim 1 wherein each
transmission line element couples to the others through mutual
inductance and capacitive coupling.
13. The magnetic resonance imaging system of claim 1 wherein
distributed impedance elements are connected between certain of
said transmission line segments to alter the coupling
therebetween.
14. The magnetic resonance imaging system of claim 1 wherein
impedance elements are connected between said transmission line
segments to establish interactions that establish frequency
dependent relations between the currents and voltages present on
certain of said transmission line segments.
15. The magnetic resonance imaging system of claim 1 wherein a
given current distribution is obtained on said transmission line
elements at a given frequency by adjustment of the geometry of said
transmission line elements and circuit components connected among
said transmission line elements.
16. The magnetic resonance imaging system of claim 1 wherein the
fields generated by the currents in said transmission line elements
are superposed to create a given magnetic field configuration for
use in either or both the generation and detection of the NMR
signal.
17. The magnetic resonance imaging system of claim 1 including RF
power amplifiers and/or RF receivers coupled to at least one of
said transmission line elements for transferring energy into said
coil during the generation of said transmit signal and out of said
coil during the reception of said receive signal.
18. The magnetic resonance imaging system of claim 1 including at
least an RF power amplifier reactively coupled to at least one of
said transmission line elements for transferring energy into said
coil during the generation of said transmit signal, the impedance
of said RF power amplifier and the impedance of said one of said
transmission line elements being matched.
19. The magnetic resonance imaging system of claim 1 including at
least an RF receiver reactively coupled to at least one of said
transmission line elements for transferring energy from said coil
during the reception of said receive signal, the impedance of said
RF receiver and the impedance of said one of said transmission line
elements being matched.
20. The magnetic resonance imaging system of claim 1 wherein the
phases of the current in a plurality of said transmission line
segments are offset so as to create an elliptically polarized
magnetic field for generating and/or detecting nuclear magnetic
resonance signals.
21. The magnetic resonance imaging system of claim 17 including a
plurality of diodes operatively connected to a plurality of said
transmission line segments for tuning the coupling between said
transmission line segments and said RF amplifiers and
receivers.
22. The magnetic resonance imaging system of claim 1 including
reactive coupling elements between one or more transmission line
elements to allow the currents on each transmission line element to
be relatively independent.
23. The magnetic resonance imaging system of claim 1 with
individual preamplifiers connected to each transmission line
element with impedance mismatches designed to allow each
transmission line element to operate independently allowing the
signals from each transmission line element to be combined either
before or after image reconstruction for optimal image
reception.
24. The magnetic resonance imaging system of claim 1 with
individual preamplifier/receivers connected to each transmission
line element with the independent information obtained from
individual transmission line elements being used to decode spatial
information regarding said subject.
25. The magnetic resonance imaging system of claim 1, with
individual power amplifiers connected to each transmission line
element with impedance mismatches designed to allow the current of
each transmission line element to be independently controlled
allowing a transmit field of desired spatial intensity and phase to
be generated.
26. A magnetic resonance imaging system comprising: (a) a housing
providing a medical diagnostic chamber with a static homogenous
magnetic field for a subject therewithin lying along an axis; (b) a
plurality of transmit/receive inductor systems about said axis in
proximity with said housing; (c) a gradient inductor system
operatively associated with said transmit/receive inductor systems;
(d) a static magnetic field inductor system operatively associated
with said transmit/receive inductor systems; (e) at least one of
said transmit/receive inductor systems constituting a coil having
an outer surface about said axis and including a series of
electrical transmission line elements paraxially distributed with
respect to said axis about said subject, each of said transmission
line elements including an outer conductor and an inner conductor,
said inner conductor being spaced from said outer conductor in a
direction perpendicular to said outer surface; (f) each said coil
selectively transmitting to said subject fields of radio frequency
energy, and selectively receiving from said subject fields of
magnetic resonance energy; (g) said gradient inductor system
initiating perturbations in said fields and producing signals
derived responsively from said perturbations; (h) said signals
corresponding to spatial indicia derived from said subject.
27. The magnetic resonance imaging system of claim 26, wherein one
of said coils is a conventional loop inductor.
28. The magnetic resonance imaging system of claim 26, wherein one
of said coils is a conventional loop inductor which is detuned
during transmit function, said transmit function being performed by
a transmission line coil which is detuned during receive.
29. The magnetic resonance imaging system of claim 26, wherein one
of said coils is a phased array of conventional loop inductors.
30. The magnetic resonance imaging system of claim 26, wherein one
of said coils is a phased array of conventional loop inductors
which are detuned during transmit function, said transmit function
being performed by a transmission line coil which is detuned during
receive function.
31. The magnetic resonance imaging system of claim 26, wherein one
of said coils is an array of said transmission line elements each
operated independently with individual preamplifiers/receivers.
32. The magnetic resonance imaging system of claim 26, wherein one
of said coils is an array of said transmission line elements each
operated independently with individual preamplifiers/receivers,
said array being detuned during system transmit function.
33. The magnetic resonance imaging system of claim 26, wherein said
system includes at least two coils, one of said coils being a
transmit coil and the other of said coils being a receive coil.
34. A magnetic resonance imaging system comprising: (a) a housing
providing a medical diagnostic chamber for a subject therewithin
lying along an axis; (b) a transmit inductor system about said axis
in proximity with said housing; (c) a gradient inductor system
operatively associated with said transmit inductor system; (d) a
static magnetic field inductor system operatively associated with
said transmit inductor system; (e) said receive inductor system
constituting a coil having an outer surface about said axis and
including a series of electrical transmission fine elements
paraxially distributed with respect to said axis about said
subject, each of said transmission line elements including an outer
conductor and an inner conductor, said inner conductor being spaced
from said outer conductor in a direction perpendicular to said
outer surface, said coil including a means for detuning said coil
to prevent disturbance of the transmit fields generated by a
separate transmit inductor system; (f) said coil initially
transmitting to said subject fields of radio frequency energy as a
transmit signal; (g) said gradient inductor system initiating
perturbations in said fields.
35. The magnetic resonance imaging system of claim 34 wherein said
coil establishes concentrations of transmit electromagnetic fields
among said transmission line elements.
36. The magnetic resonance imaging system of claim 34 wherein, by
adjusting the distance between said transmission line elements, the
interaction of the magnetic fields of said transmission line
elements with an external sample can be controlled and optimized
for nuclear magnetic resonance signal generation excitation.
37. The magnetic resonance imaging system of claim 34 wherein said
series of transmission line elements decrease the inductance of
each line element and minimize the electric fields associated
therewith.
38. The magnetic resonance imaging system of claim 34 wherein said
series of transmission line elements have inherent shielding.
39. The magnetic resonance imaging system of claim 34 wherein said
transmit inductor system consists of N transmission line elements
arranged in a geometric pattern in which each of said transmission
line elements is substantially equidistant from each adjacent
transmission line element.
40. The magnetic resonance imaging system of claim 39 wherein said
geometric pattern is circular or elliptical.
41. The magnetic resonance imaging system of claim 39 wherein said
geometric pattern is flat or curved.
42. The magnetic resonance imaging system of claim 34 wherein said
outer and inner conductors include additional lumped or distributed
capacitive or inductive circuit components.
43. The magnetic resonance imaging system of claim 34 wherein each
of said transmission line elements couples to the other of said
transmission line elements through mutual inductance and capacitive
coupling.
44. The magnetic resonance imaging system of claim 34 wherein
distributed impedance elements are connected between certain of
said transmission line elements to alter the coupling
therebetween.
45. The magnetic resonance imaging system of claim 34 wherein
impedance elements are connected between said transmission line
elements to establish interactions that establish frequency
dependent relations between the currents and voltages present on
certain of said transmission line elements.
46. The magnetic resonance imaging system of claim 34 wherein a
given current distribution is obtained on said transmission line
elements at a given frequency by adjustment of the geometry of said
transmission line elements and circuit components connected among
said transmission line elements.
47. The magnetic resonance imaging system of claim 34 wherein the
fields generated by the currents in said transmission line elements
are superposed to create a given magnetic field configuration for
use the generation of the NMR signal.
48. The magnetic resonance imaging system of claim 34 including RF
power amplifiers coupled to at least one of said transmission line
elements for transferring energy into said coil during the
generation of said transmit signal.
49. The magnetic resonance imaging system of claim 34 including at
least an RF power amplifier reactively coupled to at least one of
said transmission line elements for transferring energy into said
coil during the generation of said transmit signal, the impedance
of said RF power amplifier and the impedance of said one of said
transmission line elements being matched.
50. The magnetic resonance imaging system of claim 34 wherein the
phases of the current in a plurality of said transmission line
elements are offset so as to create an elliptically polarized
magnetic field for generating and/or detecting nuclear magnetic
resonance signals.
51. The magnetic resonance imaging system of claim 34 including a
plurality of diodes operatively connected to a plurality of said
transmission line elements for tuning the coupling between said
transmission line elements.
52. The magnetic resonance imaging system of claim 34 including
coupling components between one or more of said transmission line
elements to allow the currents on each of said transmission line
elements to be independently controlled with separate power
amplifiers connected to one or more of said transmission line
elements allowing a transmit field of desired spatial intensity and
phase to be generated.
53. The magnetic resonance imaging system of claim 34 with
individual power amplifiers connected to each transmission line
element with impedance mismatches designed to allow the current of
each transmission line element to be independently controlled
allowing a transmit field of desired spatial intensity and phase to
be generated.
54. A magnetic resonance imaging system comprising: (a) a housing
providing a medical diagnostic chamber for a subject therewithin
lying along an axis; (b) a receive inductor system about said axis
in proximity with said housing; (c) a gradient inductor system
operatively associated with said receive inductor system; (d) a
field inductor system operatively associated with said receive
inductor system; (e) said receive inductor system constituting a
coil having an outer surface about said axis and including a series
of electrical transmission line elements paraxially distributed
with respect to said axis about said subject, each of said
transmission line elements including an outer conductor and an
inner conductor, said inner conductor being spaced from said outer
conductor in a direction perpendicular to said outer surface, said
coil including a means for detuning said coil to prevent
disturbance of the transmit fields generated by a separate transmit
inductor system; (f) said coil receiving from said subject fields
of magnetic resonance energy; (g) said gradient inductor system
initiating perturbations in said fields and producing signals
derived responsively from said perturbations; (h) said signals
corresponding to spatial indicia derived from said subject.
55. The magnetic resonance imaging system of claim 54 wherein, by
adjusting the distance between said transmission line elements, the
interaction of the magnetic fields of said transmission line
elements with an external sample can be controlled and optimized
for nuclear magnetic resonance signal detection.
56. The magnetic resonance imaging system of claim 54 wherein said
series of transmission line elements decrease the inductance of
each transmission line element and minimize the electric fields
associated therewith.
57. The magnetic resonance imaging system of claim 54 wherein said
series of transmission line elements have inherent shielding.
58. The magnetic resonance imaging system of claim 50 wherein said
series of transmission line elements are combined to optimize NMR
signal reception.
59. The magnetic resonance imaging system of claim 50 wherein
signals from said series of transmission line elements are combined
to decode spatial information derived from the NMR signal.
60. The magnetic resonance imaging system of claim 50 wherein said
receive inductor system consists of N transmission line elements
arranged in a geometric pattern in which each of said transmission
line elements is substantially equidistant from each adjacent
transmission line element.
61. The magnetic resonance imaging system of claim 60 wherein said
geometric pattern is circular or elliptical.
62. The magnetic resonance imaging system of claim 60 wherein said
geometric pattern is flat or curved.
63. The magnetic resonance imaging system of claim 59 wherein said
outer and inner conductors include additional lumped or distributed
capacitive or inductive circuit components.
64. The magnetic resonance imaging system of claim 59 wherein each
of said transmission line elements couples to the other of said
transmission line elements through mutual inductance and capacitive
coupling.
65. The magnetic resonance imaging system of claim 59 wherein
distributed impedance elements are connected between certain of
said transmission line elements alter the coupling
therebetween.
66. The magnetic resonance imaging system of claim 59 wherein
impedance elements are connected between said transmission line
elements to establish interactions that establish frequency
dependent relations between the currents and voltages present on
certain of said transmission line elements.
67. The magnetic resonance imaging system of claim 59 wherein
a=given current distribution is obtained on said transmission line
elements at a given frequency by adjustment of the geometry of said
transmission line elements and circuit components connected among
said transmission line elements.
68. The magnetic resonance imaging system of claim 59 wherein the
fields generated by the currents in said transmission line elements
are superposed to create a given magnetic field configuration for
use in the detection of the NMR signal.
69. The magnetic resonance imaging system of claim 59 including RF
receivers coupled to at least one of said transmission line
elements for transferring energy out of said coil during
receive.
70. The magnetic resonance imaging system of claim 59 wherein the
phases of the current in a plurality of said transmission line
elements are offset so as to create an elliptically polarized
magnetic field for detecting nuclear magnetic resonance
signals.
71. The magnetic resonance imaging system of claim 69 including a
plurality of diodes operatively connected to a plurality of said
transmission line elements for tuning the coupling between said
transmission line elements and said RF receivers.
72. The magnetic resonance imaging system of claim 59 including
coupling elements between one or more of said transmission line
elements in order to make the currents on each of said transmission
line elements relatively independent allowing the signals from two
or more of said transmission line elements to be optimally combined
before or after image reconstruction.
73. The magnetic resonance imaging system of claim 59 with
individual preamplifiers connected to each of said transmission
line elements with impedance mismatches designed to allow each of
said transmission line elements to operate independently allowing
the signals from two or more of said transmission line element to
be optimally combined either before or after image
reconstruction.
74. The magnetic resonance imaging system of claim 59 with
individual preamplifier/receivers connected to each transmission
line element with the independent information obtained from
individual transmission line elements being used to decode spatial
information regarding said subject.
Description
RELATED APPLICATIONS
[0001] This application is a continuation of U.S. patent
application Ser. No. 10/750,031, filed Dec. 29, 2003, entitled "RF
Coil for Imaging System," by J. T. Vaughan, which is a continuation
of U.S. patent application Ser. No. 10/367,489, filed Feb. 14,
2003, entitled "RF Coil for Imaging System," by J. T. Vaughan,
which application is a divisional of U.S. patent application Ser.
No. 09/575,384, filed May 22, 2000, now issued U.S. Pat. No.
6,633,161, (Oct. 14, 2003), entitled "RF Coil for Imaging System,"
by J. T. Vaughan, which application claims the benefit of U.S.
Provisional Patent Application Ser. No. 60/135,269, filed May 21,
1999, entitled "RF Coil for Imaging System," by J. T. Vaughan, each
of which is incorporated herein by reference.
FIELD OF THE INVENTION
[0002] This invention relates to imaging systems employing radio
frequency (RF) coils for RF field generation, and more particularly
to RF coils for use in such systems which coils facilitate higher
frequency, higher efficiency, higher energy operation, permit use
of larger coils, facilitate flexibility in coil design to
accommodate a variety of applications and provide enhanced
signal-to-noise performance so as to achieve among other things
improved MRI, fMRI and MR spectroscopic imaging, all the above
being achieved without significant increase in cost. The invention
applies similarly to EPR or ESR.
BACKGROUND OF THE INVENTION
[0003] Nuclear magnetic resonance (NMR) or magnetic resonance
imaging (MRI), functional MRI (fMRI), electron spin resonance (ESR)
or electron paramagnetic resonance (EPR) and other imaging
techniques using RF field generating coils are finding increasing
utility in applications involving imaging of various parts of the
human body, of other organisms, whether living or dead, and of
other materials or objects requiring imaging or spectroscopy. For
purposes of this application, RF shall be considered to include
frequencies from approximately 1 MHz to 100 GHZ, the upper ranges
of which are considered to be microwaves. While existing such
systems are adequate for many applications, there is often a need
for higher signal-to-noise and improved spectral resolution in such
imaging so as to permit higher spatial resolution, higher image
contrast, and faster imaging speed. In fMRI applications for
example, where multiple images may be taken over time and a
difference image generated to permit visualization of small changes
in blood oxygen use over time in the body being imaged, differences
between successive images may be very small, requiring high
signal-to-noise to permit detection. A major limitation to higher
resolution, and/or faster imaging is an insufficient signal to
noise ratio. If the image signal intensity is below the noise
level, an image can not be made. It is therefore important in high
resolution systems to design an RF coil to maximize signal and to
minimize noise. The RF coil of such a system should also be
designed to minimize eddy currents propagating therein which are
induced by time transient currents in gradient coils or by other
causes.
[0004] The signal-to-noise ratio (SNR) and spectral resolution are
increased by increasing the magnetic field strength of the system,
generally expressed in tesla (T). The SNR benefits of image speed,
spatial resolution, and contrast are also increased with the
magnetic field strength. However, the frequency of which the nuclei
of atoms in the body resonates varies as a function of the applied
magnetic field, with each atomic species having a unique magnetic
field dependent resonant frequency referred to as the Larmor
frequency. For the human body which is composed primarily of
hydrogen atoms in water, fat and muscle tissue, these hydrogen
nuclear (proton) frequencies are approximately 64 MHZ for a field
strength of 1.5 T, 170 MHZ (4 T), 175 MHZ (4.1 T), 300 MHZ (7 T),
340 MHZ (8 T) and 400 MHZ (9.4 T). Other species of atomic nuclei
in a body would resonate at other frequencies for a given field
strength. However, while conventional birdcage coils in existing
MRI and related systems might resonate at a frequency of 170 MHZ (4
T) for example, the conventional birdcage coil with lumped elements
(reactance) will operate very inefficiently, radiating much of its
energy like an antenna, rather than conserving its energy like a
"coil". At higher frequencies still, such lumped element coils of
human head or body dimensions will not reach the Larmor resonant
frequency required, limiting the magnetic field strength at which
such MRI or EPR systems can operate. Further, since frequency is a
function of the electrical path lengths (measured in wavelengths)
in the RF coil, higher frequency, and thus higher field strength
operation, has been previously achievable only with very small
coils which are not always useful for imaging a human being or
other larger objects. A need therefore exists for an RF coil design
which provides short electrical path lengths and shields against
radiative losses, while still permitting an RF coil to be
constructed with physical dimensions sufficient to image a human
body and/or other larger objects with high frequency RF energy,
thus permitting high field strength operation. It is also desirable
to be able to tune each path of an RF coil to a precise resonant
frequency, to be able to provide two or more resonant frequencies
for different paths on the coil, and to be able to easily
adjust/retune the resonant frequency of a path or paths.
[0005] Still another potential problem in operating these imaging
systems, especially at high fields, is in driving the RF coil in a
manner so as to achieve a homogeneous RF field, even when a body is
positioned in the field, or to achieve some other desired field
profile. Many factors influence field profile or contours including
the manner in which the coils are driven, the geometric and
frequency dependent electrical properties of the anatomy or object,
and the frequency dependent properties of the coil circuits.
Techniques for controlling these and other factors to achieve a
homogenous or other desired field profile are therefore desirable.
Also, while in many systems the same coil is used for both the
transmitting of RF energy and the receiving thereof, the coils
being switched between transmit and receive circuitry, there are
many applications where the homogeneous excitation of NMR signal is
achieved with a large volume coil and a small local receive coil
having very short path lengths is used for achieving high SNR
operation, such local receive coil being placed as close to the
region of the body being imaged as possible. However, having both
the large transmit coil and the local RF receive coil tuned to the
same frequency results in the coils being destructively coupled (by
Lenz's Law for example), this defeating enhanced operation from the
local receive coil. It is therefore desirable to be able to quickly
detune the large RF transmit coil during a receive operation by a
local RF receive coil and vice versa. Improved ways of achieving
this objective, particularly in an RF coil providing the
characteristics previously indicated, are therefore desirable.
[0006] Finally, some of the advantages of having a local receive
coil, and in particular the ability to place the RF coil closely
adjacent to a region where imaging is desired, could be achieved if
the RF coil were designed so as to localize both the transmission
and reception of RF energy. While coils adapted for performing this
function in certain specialized situations have existed in the
past, a more general purpose design for RF coils to facilitate
their use in producing localized RF fields and the localized
reception of RF (NMR) signal, is desirable in order to achieve the
enhanced SNR benefits of higher image signal, resolution, speed and
contrast.
[0007] While some of the advantages indicated above are achieved by
distributed impedance RF coils disclosed in U.S. Pat. Nos.
5,557,247 and 5,744,957, which patents have the same inventor as
this invention, the systems taught in these patents, and in
particular the RF coils thereof, do not provide optimum performance
in all situations, and improvements are possible on various aspects
of these RF coils, including eddy current suppression, design of
the coil for optimum positioning in a greater number of cases,
improved control of field profile, improved tuning options and
improved detuning in situations where the use of two or more coils
is desired.
SUMMARY OF THE INVENTION
[0008] In accordance with the above, this invention provides an RF
coil for use in an imaging system, which coil has a cavity formed
as a conductive enclosure in which resonant field can be excited,
the enclosure being formed at least in part of an electrical
conductor patterned to form RF conductive paths around the cavity.
At least one tuning mechanism may be provided which determines a
resonant frequency or frequencies for such paths. The tuning
mechanism may be fixed, resulting in a preselected resonant
frequency for the path, or variable to provide a tunable resonant
frequency or frequencies. The tuning mechanism may reactively
adjust the electrical length of each path to tune the path. The
path reactance may also be adjusted to achieve a selected field
profile for the coil. The tuning mechanism may tune all the paths
to the same resonance frequency or may selectively tune the paths
to resonate at two or more different frequencies. In particular,
alternate ones of the paths may be tuned to resonate at a first
frequency and the remaining ones of the paths tuned to resonate at
a second frequency.
[0009] The coil may also include a dielectric which at least
substantially fills the cavity the thickness of the conductor for
at least selected portions of the enclosure may be substantially
greater than one skin depth at the resonant frequency, the
dielectric filling the cavity having a dielectric constant
different from that of air. This results in signals of different
frequencies propagating on the outer and inner surfaces of the
conductor.
[0010] Each of the N paths on the coil may have at least one
non-conductive gap formed therein, and the tuning mechanism may
include a reactance and/or an impedance across at least selected
ones of said gaps. The reactance/impedance for at least selected
ones of the gaps may be variable to control the resonant frequency
for the corresponding path. The reactance for at least some
embodiments includes a capacitor, the capacitance of which may be
varied and/or an inductor the inductance of which may be varied.
The variable impedance and/or reactance may be controlled to tune,
retune and/or detune the path in which it is located. Where the
enclosure is formed of an outer wall, inner wall, and side walls,
end conductive lands maybe formed on each of the walls, with
corresponding lands on each wall being connected to form the paths
and the gaps being formed in the conductor for each of the paths
for at least one of the walls. For some embodiments, the gaps are
formed in the outer wall conductor for each path.
[0011] The resonant frequency of the paths may be determined by
distributed capacitance and distributed inductance for the path,
the distributed capacitance being determined by the area of the
electrical conductor for each path, a dielectric fill material for
the cavity and/or the dimensions of the dielectric fill material.
The electrical conductor forming each path may be a thin foil, the
distributed inductance for the path being a function of the path
length. At least one reactance component may be provided in at
least selected ones of the paths, the reactance component being
either distributed or lumped. A distributed or discrete reactance
may be selected to achieve a desired resonant frequency for the
paths. The paths have a cumulative reactance which includes at
least in part the distributed capacitance/inductance, the
cumulative reactance for the paths being tuned to result in D
different resonant frequencies for the coil, every Dth path
symmetrically spaced around the coil being tuned to the same
frequency.
[0012] The coil may include a circuit which applies RF signal to
and/or receives RF signal from M selectively spaced ones of said
paths, where M is an integer and 1.ltoreq.M.ltoreq.N. The RF
signals may be phase shifted corresponding to a phase shift for the
corresponding paths to provide circular or other polarization for
the coil. Each RF signal is preferably reactively coupled to the
corresponding path, the coupling reactance for each path being
variable for some embodiments to independently match/tune the path.
In particular, the coupling reactances may be impedance matched to
different loading conditions for the coil. For some embodiments,
the RF coil may be used to transmit and/or receive RF signals, but
not both simultaneously, and includes a detuning mechanism for the
paths, the detuning mechanism being operative when the RF coil is
not in the one of the transmit/receive modes for which it is being
used. The detuning mechanism may include a mechanism for altering
the path length and/or impedance for each path to be detuned, and
in particular may include for some embodiments a PIN diode circuit
for each path which facilitates rapid switching to a changed
impedance state sufficient to effect the path detuning.
Alternatively, the RF drive signals may be phase-shifted
corresponding to the phase shift for the paths to which they are
applied to provide circular polarization for the coil, the detuning
mechanism including circuitry which reverses the phase of the RF
drive signals.
[0013] For an enclosure which is formed of an outer wall, inner
wall, and side walls, within conductive lands being formed for each
wall, corresponding lands on each wall being connected to form the
paths, at least the outer wall may have two conductive layers
separated by a dielectric, the two conductive layers each being
slotted to form a pattern of lands, with slots on each layer being
overlaid by lands of the adjacent layer. The degree of overlap for
the lands of said layers is at least one factor controlling coil
resonant frequency.
[0014] At least one of the side walls may also have an aperture
through substantially the center thereof through which a body to be
analyzed may be passed to an area inside the inner wall, the
conductive layer on the inner wall being patterned to provide a
selected magnetic flow pattern in said aperture. One of the side
walls may also be closed, the closed side wall being slotted to
form a land pattern covering at least most of the wall.
[0015] The imaging system may also have at least one gradient coil
which induces low frequency eddy currents in the RF coil, the
slotting on at least the outer wall and side walls resulting in the
breaking up of and substantial attenuation of such eddy currents
without substantial attenuation of RF currents and fields. The
electrical conductor for at least the outer wall and side walls may
be a conductive layer which is thin enough to attenuate low
frequency eddy currents while still conducting RF currents. For
such embodiment, the conductor layer has a thickness substantially
equal to one skin depth at the resonance frequency to which the
coil is tuned, which thickness is substantially equal to
approximately 5 microns for an illustrative embodiment.
[0016] For some embodiments, each of the paths has at least one
circumferential/azimuthal slot formed therein to break the path
into smaller paths. A fixed, variable and/or switched reactive
coupling and/or an impedance coupling may be connected across each
of the circumferential slots. Where a reactive coupling is
utilized, such coupling is a capacitive coupling for illustrative
embodiments.
[0017] An RF drive signal input is provided to at least one of the
paths, the path inductively coupling an RF drive signal and a path
to adjacent paths.
[0018] The dielectric material filling the cavity may provide a
selected path capacitance, and thus a selected resonant frequency.
A mechanism may be provided for controlling the dielectric fill of
the cavity and thus the resonant frequency of the coil.
[0019] The electrical conductor may be patterned to form N
conductive lands for the enclosure, each of a selected width, and
the number N of conductive paths and the width of conductive lands
for each path may be selected to achieve a desired resonant
frequency and a desired field contour.
[0020] The enclosure is preferably formed to break induced eddy
currents and/or to shape the RF magnetic field patterns.
[0021] For some embodiments, a lid is mounted to at least one end
of the coil. The lid may be conductive, non-conductive or segmented
to be partially conductive. A plurality of sample spaces may also
be formed in the dielectric at a selected portion of the enclosure,
such portion being one of the side walls for an illustrative
embodiment. The sample spaces extend at least part way into the
dielectric from the side wall. Alternatively, the open center
chamber or aperture of the coil may contain a dielectric which
preferably fills such chamber and a plurality of sample spaces may
penetrate such dielectric. At least a portion of at least selected
ones of the paths may be formed as conductive tubes or coaxial tube
conductors.
[0022] For embodiments having a close end wall, the closed end wall
functions as an RF mirror, the end wall having a radial slotting
pattern covering at least most of the wall for an embodiment where
the electrical conductors for each wall are slotted.
[0023] For some embodiments, field applied to at least one of the
electrically conducting paths causes an alternating magnetic field
in the cavity and at least one aperture is formed in at least the
electrical conductor through which magnetic field may be applied to
an adjacent body. For some such embodiments, the coil is shaped to
conform to the body being imaged, the surrounding walls including
connected top and bottom walls, with the at least one aperture
being formed in only the bottom wall. The coil may be flexible to
conform to a surface of a body being imaged and the at least one
apertures may be arranged to be adjacent the areas to be imaged of
the body being imaged. Where the areas to be imaged are at least
one projection on a body, such projection may extend into the
cavity through an adjacent aperture. For such embodiments, the
dielectric may be conformable to an outer surface of a projection
extending into the cavity so as to minimize discontinuity between
the projection and the dielectric. Where the coil is formed is a
closed loop, an aperture may be formed in only one of an inner or
outer wall of the coil. Apertures may be arranged to be adjacent
areas to be imaged of a body being imaged.
[0024] Various features of the invention, such as the detuning
mechanism, maybe employed independent of other features of the
invention. Another potentially independent feature is the
dielectric material filling the cavity being utilized to control
the resonant frequency for one or more of the paths. In particular,
the dielectric material may be different in different areas of the
cavity so as to selectively shape the coil field. A mechanism may
also be provided for controlling the dielectric fill of the cavity
and thus the resonant frequency of the coil, for example the amount
of fluid in the cavity being controlled where a fluid dielectric is
employed. Acoustic damping material may also be provided as a fill
for at least a portion of the cavity. A dielectric material may
also be selectively positioned between the coil and a body to be
imaged to control and/or shape the field applied to the body from
the coil. The dielectric material preferably substantially fills
the space between the coil and at least a selected area of the
body, the dielectric constant of the dielectric substantially
matching that of the body in such area. Where a selected area is
the area to be imaged, the dielectric concentrates and directs the
field to such area. In its broadest sense, the invention includes a
conductive enclosure which is patterned to suppress low frequency
currents and EMI noise.
[0025] In accordance with still another aspect of the invention,
the RF coil includes a cavity formed by at least an inner and an
outer conductor, a dielectric material filling the cavity and at
least one sample space formed in the coil. The sample space may for
example be formed in the dielectric material, projecting therein
from a wall of the enclosure. For some embodiments, the coil is a
transmission line stub, the inner and outer conductor of the coil
being the inner and outer conductors of the transmission line stub,
respectively. A conductive cap may short one end of the
transmission line stub. The sample space is preferably located at a
distal end of the stub, the sample space extending from such distal
end into the dielectric material or a hollowed-out portion of the
center conductor. The stub is tuned and matched so that maximum
current, and therefore maximum RF field, occur at such distal
end.
[0026] The foregoing and other objects, features and advantages of
the invention will be apparent in the following more particular
description of preferred embodiments of the invention as
illustrated in the accompanying drawings.
[0027] IN THE DRAWINGS
[0028] FIGS. 1A and 1B are perspective views of RF coils in
accordance with illustrative embodiments of the invention.
[0029] FIG. 2A is a sectional view taken generally along the line
2-2 of FIG. 1A.
[0030] FIG. 2B is the same sectional view as FIG. 2A, but for an
alternative illustrative embodiment
[0031] FIGS. 3A and 3B are illustrations of slotted conductor
configurations for an outside wall and an inside wall,
respectively, of an illustrative coil having double-sided conductor
clad dielectric substrate on both walls.
[0032] FIGS. 3C and 3D are illustrations of slotted conductor
configurations for an alternative embodiment.
[0033] FIGS. 4A and 4B are illustrations of slotted conductor
patterns for the back wall and front wall, respectively, of an
illustrative coil embodiment having double-sided conductors on each
of these walls and a closed back wall. The front and back end walls
of FIGS. 4A and 4B are preferably utilized in conjunction with the
inside and outside walls of FIGS. 3A, 3B or of FIGS. 3C, 3D to form
a conductor pattern for an RF coil embodiment.
[0034] FIG. 5A is a semi-schematic diagram of a coil for an
illustrative embodiment illustrating a novel drive mechanism for
such coil and a novel detuning mechanism for the coil.
[0035] FIGS. 5B is an enlarged schematic view of a detuning circuit
for an embodiment.
[0036] FIG. 6 is a top perspective view of a receiver coil
illustrating detuning mechanism.
[0037] FIGS. 7A and 7B are a top perspective and bottom perspective
view, respectively, of an RF coil for another embodiment of the
invention.
[0038] FIG. 8 is a planar rear view of a coil in accordance with
the teachings of this invention with a patient therein and a field
"focusing" dielectric pillow.
[0039] FIGS. 9A, 9C and 9E are perspective views of three
additional embodiments of the invention and FIGS. 9B, 9D and 9F are
sectional views of FIGS. 9A, 9C and 9E, respectively.
[0040] FIGS. 10A-10C are perspective views of three coaxial stub
embodiments of the invention.
[0041] FIG. 10D is a cross-sectional view of a generalized coaxial
stub embodiment.
DETAILED DESCRIPTION
[0042] FIGS. 1A and 2A show an illustrative embodiment of the
invention which overcomes many of the problems discussed earlier.
In particular, the RF coil 10 shown in these figures has a
conducting cavity formed as a conductive enclosure 12 in which
resonant field can be excited, the enclosure being formed by a
surrounding, conducting wall 16, which wall may be supported by a
non-conducting wall 14. Conducting wall 16 may be a whole wall
which is at least selectively patterned as described later, or may
be formed of conducting tubes, coaxial tubes as in the U.S. Pat.
No. 5,557,247 or other appropriate spaced components.
Cavity/enclosure 12 is filled with air or another dielectric
material and supporting wall 14 may also be formed of a dielectric
material having a dielectric constant which substantially matches
that of the material in cavity 12. Cavity 12 may also be defined by
a solid piece of dielectric material of the appropriate shape.
[0043] The enclosure or cavity 12 may have a circular (FIG. 1B) or
elipsoidal FIG. 1A) transaxial cross-section for many applications,
and an elongated axial cross-section, shown for example in FIG. 2A
or FIG. 2B, so that the cavity has an outer conducting wall 16O, an
inner conducting wall 16I, a front conducting wall 16F, and a rear
conducting wall 16R. (1a,b5a,b Disclosure) As indicated above, each
of the conducting walls 16 maybe supported by a corresponding
non-conducting wall 14, or only selected ones of the conducting
walls may be so supported. In accordance with the teachings of this
invention, each of the conductor walls 16O, 16I, 16F, 16R is
slotted or otherwise patterned to form a plurality of lands 18
separated by nonconducting slots 20. While in the figures, the
slots 20 are shown as being substantially parallel to each other
and to the axis 21 (FIG. 2A) of the coil in each of the inner and
outer walls, and continuous with substantially radial slots in the
front and rear walls, this is not a limitation on the invention,
and other patterns are possible. The exact pattern of slots and
lands formed for a given coil may vary with application and with
the desired field profile for the coil. While there is no limit to
the width of the slots and lands, narrow slots and wide lands
provide better Faraday shielding. Wide slots and narrower lands
allow RF magnetic flux to pass and allow for visual, auditory,
physical, and other access to/from the coil's interior through it's
inner and/or outer walls. Thus, for one embodiment, conducting
cavity walls 160, 16F, and 16R have narrow slots and wide lands,
and conducting wall 16I has narrow lands and wide slots. For this
embodiment, conducting wall 16I may for example be formed of
coaxial tube conductors.
[0044] One of the objectives of the coils shown in FIGS. 1A, 1B,
2A, and 2B is to suppress eddy currents in the coil, and in
particular in the outer wall thereof, caused by the proximity of
the RF coil to gradient coils for the imaging system, and to
suppress other low frequency noise in the conductor, such eddy
currents/low frequency noise causing image blurring and therefore
adversely affecting the signal-to-noise ratio and resolution
achievable from a system employing the coil. One way in which such
eddy currents have been suppressed in the past is to have the
thickness of the conductor for at least outer wall 16O of the coil,
and preferably for all walls of the coil, thin enough so as to
attenuate the low frequency eddy currents induced by the gradient
coil, while still conducting RF currents. This is possible since
the skin depth required to conduct signal decreases with increasing
frequency, so that if the thickness of the conductor 16 is
substantially equal to one skin depth at the resonant frequency to
which the coil is tuned, this frequency being a relatively high RF
frequency, then the conductor will not pass, and will suppress or
attenuate the low frequency gradient field induced signals or other
low frequency noise signals.
[0045] However, while this mechanism preserves the coil's RF
efficiency while attenuating switched gradient induced eddy
currents, it alone is not sufficient to fully suppress gradient
and/or other low frequency noise for some applications such as
fMRI. This objective is facilitated by the slotting or dividing of
at least outer wall conducting 16O and preferably by the slotting
of this wall and at least end conducting walls 16F and 16R. The
slotting of the front and rear conducting walls is desirable to
prevent switched gradient induced eddy current flow through or
around the ends of the coil. The narrower this slotting (i.e. the
greater number there are of nonconducting slots 20, and therefore
the narrower the width of each land 18), the more effective this
eddy current suppression becomes. The combination of the conductor
thickness being substantially equal to one skin depth at the
resonant frequency and the slotting of conductor 16, preferably for
at least outer wall 16O and end walls 16F and 16R, provides a
substantial elimination of all eddy current induced/low frequency
noise in RF coil 10, and thus far clearer images and/or faster
imaging, then can otherwise be obtained.
[0046] Further, in order to achieve increased field strength to 4
T, 7 T, 9.4 T or even higher, it is necessary to be able to operate
RF coil 10 with increasingly high frequencies. For example, as
previously indicated, for a field strength of 4 T, coil 10, when
used in an MRI embodiment on the human body, must have a resonant
frequency of about 170 MHZ, and this frequency goes to 400 MHZ for
a 9.4 T field strength. However, for a coil to resonate at these
higher frequencies, the reactance of the coil (i.e. its inductance
and capacitance) must be relatively low. Such low reactances are
either not achievable, or are achievable only for coils so small as
to have limited practical application, when lumped inductors and
capacitors are used in conventional lumped element circuit designs
for the coil. Therefore, distributed capacitance and inductance has
been used in distributed element circuit designs to facilitate
desired lower reactances. However, while a coil 10 such as that
shown in FIG. 1 with distributed reactance can offer higher
frequency performance than coils operating with lumped reactance
components, even such head and body sized coils have difficulty
operating with maximum efficiency at Larmor frequencies
corresponding to field strengths above 4 or 5 T. The coil 10
reduces this problem by breaking the conducting wall to form a
large number of continuous paths; the lands 18 of the walls being
connected or formed into N such conductive paths. Breaking the
conducting wall to form N conductive paths also improves the
homogeneity of the field, the higher the value of N, the more
homogeneous the field. N may for example be 16 or 24. FIG. 1
further shows a circumferential or azimuthal slot 22 being formed
in for example the center of outer conducting wall 16O, which slot
is covered by a collar 24 of a conductive material (the collar 24
being shown in dotted lines in FIG. 1 so that the structure
thereunder may be more easily viewed). The slot 22 further breaks
up the paths, thus shortening the individual paths and reducing the
inductance thereof. The greater number of breaks in each of N paths
formed around coil 10, the lower the inductance, and thus the
higher the resonant frequency for such coil. Collar 24 may be moved
to vary the capacitance formed by the collar, slot 24 and lands 18
of conductor 16, to vary the capacitance of each path, and thus
fine tune the resonant frequency thereof. Tuning could similarly be
performed by variable or fixed capacitors bridging the gap or gaps
in each path. Capacitance for the paths is also determined by the
thickness of dielectric filled cavity 12, by the dielectric
material in this cavity, by the area and thickness of conductive
wall 16, etc.
[0047] Distributed inductance is determined primarily by the
uninterrupted conductor length and by the width of each path. Thus,
assuming that all paths are to operate at the same resonant
frequency, the slotting of conductor 16 is selected so that all of
the lands 20 are of equal width; however, in applications where
different paths are to have different resonant frequencies, for
example every other path having a first resonant frequency and the
remaining paths a second resonant frequency, every other path could
be of a first width to provide the first resonant frequency and the
intervening paths of a second width to provide the second resonant
frequency. Various parameters of the paths may also be selected or
adjusted, including capacitance, inductance, phase, and conductor
thickness of at least selected walls, to control relative current
carrying or otherwise control field contours or profiles within the
coil.
[0048] While only a single circumferential/azimuthal slot 22 on the
outer wall 16O is shown in FIGS. 1A and 2A, this is not a
limitation on the invention, and higher frequency operation can be
achieved by having a plurality of circumferential slots 22 for each
path. For example, gaps might be provided at selected one or more
of points A, B, C in FIG. 1B. Further, there are some advantages in
having slot 22 on outer wall 16O of the coil, including the
operation of the tuning ring or capacitors 24 being easier from
this location, and that the E-field applied to a patient/body in
the recess 26 formed by inner wall 16I is significantly reduced,
thus reducing detuning caused by the presence of the patient(body
in the coil interior and reducing E-field induced noise and heating
in the patient or load. This is a problem particularly at higher
frequencies since more EM energy is lost (radiated) at higher
frequencies for a given coil. However, slots 22 may be at a
plurality of locations on outer wall 16O and/or inner wall 16I, on
sidewalls 16F and/or 16R and/or at the intersections of these
various walls. (See points B, C, FIG. 1B). In some applications
these azimuthal slots may be bridged by fixed, variable, or
switched discrete inductance and/or capacitance components (see A
in FIG. 1B), or by distributed components similar to collar 24, but
fixed rather than moveable.
[0049] FIG. 2B illustrates a number of variations which are
possible in practicing the teachings of this invention. First, the
cavity 12 for FIG. 2A is assumed to be filled with a dielectric,
the thickness and other dimensions, including volume, of the cavity
and the dielectric constant of the dielectric being two of the
factors which determine a resonant frequency for each path. FIG. 2B
illustrates the cavity 12'' as being filled with a fluid, for
example a liquid or a gas, having a known dielectric constant
However, a tube 30 connected to a suitable pump may increase or
decrease the quantity of fluid dielectric in cavity 12'' and/or the
fluid pressure in the cavity and/or may alter the spacing between
at least semi-elastic cavity walls. Any of these changes alter the
dielectric constant or volume of the dielectric in cavity 12'', and
thus can be used to control the resonant frequency or to tune the
resonant frequency of the coil.
[0050] Second, the embodiment of FIG. 2B has more clearly defined
walls which may for example be separately formed as discussed in
conjunction with FIGS. 3A-4B. These separate conductive layers may
envelop a solid dielectric cavity core or a fluid filled, for
example air filled, cavity core. Corresponding lands on adjacent
walls may then be electrically connected, for example directly,
capacitively, or inductively, to form the N continuous electrical
paths around the coil. While for reasons indicated above,
conducting wall 16 would generally be far too thin to provide
structural integrity for the coil without a supporting dielectric
substrate 14, this is not a limitation on the invention and
electrically, all that is required to define cavity 12 is a
surrounding conducting wall 16.
[0051] While one objective of the invention is to provide
distributed capacitances and inductances to achieve higher
frequency and thus higher field strength operation, in some
applications it may be desired to operate a coil of this invention
at a lower resonant frequency to, for example, permit operation at
a lower field strength for a given application, while still
achieving the other advantages of this invention. Adding discrete
reactance, for example the added lumped (fixed, variable, or
switched) capacitance (or inductance) elements 32 shown in FIG. 2B,
may be utilized both to achieve a desired reduced frequency
operation and to tune the paths to a desired lower resonant
frequency. A capacitor 32 could for example be provided for each of
the N paths or, where operation at more than one frequency is
provided, the capacitor 32 might only be in the paths to be
operated at the lower of the frequencies, or the capactance might
be one value for one frequency, and another value for another
frequency. For example, if every other path is to have a resonant
frequency at a first higher frequency and the remaining alternate
paths to be resonant at a second lower frequency, capacitors 32
might appear only for the alternate paths to be operated at the
second lower frequency, or capacitors of alternating values may
appear on alternating elements to tune two frequencies for the
coil. Further, while the embodiment of FIG. 2B has only a single
azimuthal slot 34 for each path and a single lumped capacitance or
inductance element 32, additional azimuthal slots could also be
provided for this embodiment to for example shorten path lengths,
and these slots or gaps could be bridged by appropriate lumped
reactance elements, lumped capacitance elements being shown for
bridging gaps 34, or could be bridged by distributed reactance
elements such as for the gaps of FIGS. 2A, 3B. Except for the
differences discussed above, the embodiments of FIGS. 2A and 2B are
otherwise substantially the same and operate in substantially the
same manner to achieve the benefits of this invention.
[0052] FIG. 3A illustrates a conducting wall 16 which is typically
used as an outer wall 16O, but sometimes used as an inner wall 16I
for a coil 10. FIG. 3B illustrates a transmission line conductor
16I which is most often used at the inside wall for the volume
coils 10 and 10'. These conductors differ from the conductors 16
discussed earlier in that each is made up of an inner layer and an
outer layer separated by a dielectric to form transmission line
elements. For the wall 16 of FIG. 3A, the slots for the outer layer
are shown in solid lines while the slots for the bottom or inner
layer are shown in dotted lines. The figure therefore shows these
layers as being staggered so that the slot of one layer is overlaid
by a land of the other layer. This provides superior RF efficiency
and Faraday shielding for the confinement of fields generated by
the RF coil. Similarly, for FIG. 3B which shows a conductor
preferably used for the inside wall of the coil to form conducting
wall 16I, the solid outer conductor is slotted both longitudinally
and circumferentially, while the dotted bottom conductor is slotted
only longitudinally, with lands on each side of the conductor
overlying slots on the opposite side. The double-sided conductors,
with lands on the conductor for one side overlying slots for the
conductor on the other side, also applies for the two end walls as
illustrated in FIGS. 4A and 4B, FIG. 4A being for example
conducting rear wall 16R, and FIG. 4B showing conducting front wall
16F of a coil suitable for example for head imaging. While not
shown, slots may also be provided in various of the radial lands
for the end walls to shorten path lengths. The degree of overlap
for lands on opposite sides of each double-sided conductor controls
capacitance, and thus resonant frequency, for the conductive paths.
The degree of overlap on the outside and end walls of the cavity
improve the RF conduction, efficiency and shielding of the cavity.
In addition to, or in lieu of, overlapping conductive lands on
double sided cavity walls, capacitors can be used to bridge the
slots or gaps between the lands to provide additional RF conduction
or alternate RF paths across these slots.
[0053] FIGS. 3C and 3D show a conducting outside wall 16O' and a
conducting inside wall 16I' for an alternative embodiment of the
invention. Wall 16O' is slotted into multiple overlapping layers,
as for wall 16O (FIG. 3A) but with additional azimuthal slotting.
Bridging capacitance can be added across the horizontal azimuthal
gaps at for example points A, B, and/or C, or anywhere else such
gaps appear, and can function to tune the coil/path to a desired
frequency. Placing lumped or distributed capacitance at these
outside wall positions has the added benefit of moving capacitive
sections of the coil away from the patient's head or body. Wall
16I' is shown as single-sided printed line segments. This inside
wall embodiment places the more inductive inside wall surface of
the coil nearer to the coil load (for example the patient), while
the more capacitive outside and/or end walls of the coil are away
from the load. The end walls of FIGS. 4A, 4B may be used with the
outer and inner walls 16O' and 16I' and the walls 16O, 16I may be
used together or with other inner/outer wall embodiments depending
on application.
[0054] FIG. 4A also illustrates another optional feature of the
invention in that the conducting rear wall 16R can be an open wall
like FIG. 4B, or a closed wall, this closed wall acting as an RF
mirror which permits the length of the cavity to be shorter. The
shortened coil facilitates decreased electrical path lengths for
all circuits, thus facilitating higher efficiency, higher frequency
operation and thus higher field strength operation. In this case, a
physically shortened coil is also an electrically shortened coil
(in wavelengths) and therefore a more efficient coil for a given
higher frequency and larger dimension such as a human head sized
coil operation at 128 MHz or higher.
[0055] Particularly where the coil is being utilized to image the
head or brain of a patient, the coil being shortened in this way
provides ergonomic benefits in that it permits at least the
patients mouth and sometimes nose/eyes, to be outside of the coil,
reducing the claustrophobic feeling sometimes experienced by
patients in such imaging machines, and also facilitating easier
breathing by the patient through the mouth or nose if exposed. This
design also permits an optical mirror to be mounted within visual
range of the patient to permit visual stimuli to be provided to the
patient, something which is required for various brain imaging
applications, without requiring a transparent section in the coil.
Since it is necessary to have the coil's current pass through a
conductive film or screen covering these transparent regions, the
patient is not afforded an unobstructed view of visual stimuli for
example in some fMRI studies The shortened coil is thus highly
advantageous. However, if the coil design is such as to extend over
the patient's eyes, patient visibility maybe enhanced by providing
thin conductors for the conducting wall 16 in areas over the
patient's eyes; the conducting wall 16 in this area for example
being formed by thin, substantially parallel tubes or coaxial
conductors. Alternatively a transparent section such as a view port
over the face of a human can be provided in various ways, including
1) through widened slots (thin conductors) for the inside and
outside walls over the face, 2) through widened slots or gaps
between elements in the inside wall and a continuous or slotted
conductive screen window in the outside wall, or 3) through
transparent conductive elements, continuous or slotted, for inside
and outside walls of the cavity.
[0056] As seen in FIG. 4A, the rear wall has a unique slotting
pattern, with the slotting extending over substantially the entire
end wall. The rear wall with this unique slotting pattern
contributes to the resonant frequency of each path. The coil with
the closed end wall is also more efficient in that it limits
radiation loss which might otherwise occur through the end wall,
the closed end wall also enhancing coil symmetry and thus
facilitating tuning of the paths. This follows from the fact that
all of the coil elements or paths are symmetrically referenced to
the same ground plane. Finally, since the patient's mouth can
extend beyond the coil, a bite bar can be provided to reduce
patient head movement, something which facilitates signal averaging
and diminishes motion artifacts. Without the ability to maintain
the patient's head absolutely still, applications involving
multiple images, such as for example fMRI, can provide erroneous
results.
[0057] FIG. 5A illustrates a number of features of the invention,
including preferred ways for applying RF drive signals and for
outputting received RF signals from the coil, and techniques for
switched detuning of the coil in applications where the coil is
being used for either transmit or receive but not both
simultaneously. In particular, instead of applying drive signal to
only one of the paths 18 and relying on inductive coupling to
couple the RF signal to the remaining paths, FIG. 5A shows a
technique for applying RF drive signal simultaneously to multiple
paths. In particular, an RF drive (transmit) signal on line 40 is
passed through transmit/receive switches 42 to a 180.degree.
splitter 44 which divides the signal into two signals 180.degree.
out of phase with each other. The outputs from splitter 44 are
split again into quadrature hybrids 46A and 46B to drive four line
elements or paths 18 separated by 90.degree. azimuth angles. Since
the electrical phase difference for the signals applied to each of
the paths corresponds to the angular separation between the paths,
the transmit mode is circularly polarized. A second 180.degree.
splitter 48 combines the quadrature combined receive channels
passing through the quadrature hybrids 46 to a common receive
channel 50 which is passed through switches 42 to receive output
line 52. Transmit and receive lines 40, 52 are decoupled from the
power amplifier and the preamplifier respectively by
transmit/receive switches 42. In a body coil application for
example, these switches will have the attributes of low loss, high
speed and very high power ratings, while requiring low switch bias
voltages. All of the components of the drive circuit should be
non-magnetic and can be mounted close to the back of coil 10.
[0058] While for the illustrative embodiment shown in FIG. 5A,
circularly polarized drive signals have been applied to four evenly
spaced paths on coil 10 for improved homogeneity, all that is
required for circular polarization, is that the paths to which
signal is applied be in relative quadrature phase. The number of
paths coupled electrically to the power amplifier or power
amplifiers and the receiver or receivers may be any number between
1 and N where N is the number of elements or paths 18 in FIG. 5A.
Thus, assuming 16 paths for an illustrative embodiment, signal
could be applied to 2, 4, 8, or 16 paths for balanced quadrature
operation, or could be applied to another number of coils. Transmit
signal amplitude, phase angle, and drive paths can be selected for
maximum homogeneity, or for targeting a desired region of interest
in the body or other test object. Similarly, the receive paths and
phase angles can be chosen for overall homogeneity, or for highest
sensitivity reception from a specific region of interest. In
addition, while signals are shown applied to a sidewall of the
paths, signal can be applied at various points on the paths, it
being currently preferred that they be applied to inner conducting
wall 16I at a point near its junctions with a sidewall.
[0059] Further, the signal on each of the lines is shown as being
capacitively coupled to the corresponding path through a variable
capacitor 54. While capacitive coupling is shown, any reactive
coupling (capacitive or inductive) can be used. Operation at two or
more frequencies can be achieved for the coils in FIGS. 1, 5a, by
changing the electrical path lengths for alternating paths for each
frequency desired. For example, for a double tuned coil, all odd
numbered paths would be adjusted or tuned to one electrical length
or frequency. All even numbered paths would be tuned to a second
frequency. This tuning can be achieved by adding or subtracting,
inductance or capacitance in the respective paths. Operation at two
different frequencies may also be achieved by having at least one
wall of the resonant cavity of greater than one skin depth, and by
having a dielectric constant within the cavity different than that
of air outside of the cavity. This results in a signal wavelength
on the inner surface of the conductive cavity wall facing the inner
dielectric to be different (in frequency) to the signal wavelength
on the cavity wall facing the outer air dielectric. This results in
the cavity being resonant at two different frequencies.
[0060] In the discussion so far, it has been assumed that coil 10
is being used both as a transmit coil and as a receive coil.
However, in some applications, particularly where homogeneous
excitation of, and high sensitivity detection from, localized
regions of interest (ROI) is required, separate coils may be
utilized for transmit and for receive. The coil 10 would more
typically be used as a transmit coil, with for example a phased
array or other appropriate receive coil such as is shown in FIG. 6
being placed adjacent the area being imaged on the body. The
receive coil may have varying numbers of coil loops of varying
shapes depending on application. In some applications a coil 10
might only be used as a receive coil.
[0061] One problem when separate transmit and receive coils are
used together is that destructive reactive coupling may occur
between the two coils which can interfere with the imaging and
eliminate the sensitivity benefits achievable form having a
separate receive coil. It is therefore necessary to RF field
decouple the transmit and receive coils from each other. This field
decoupling can be accomplished by orienting the spatial position of
one coil relative to the other, by manipulating the electrical
phase relations of one coil relative to the other, by changing the
field amplitude of one coil relative to the other, by changing the
resonant frequency of one coil relative to the other and/or by
temporal separation of the field of one coil relative to the other
by any combination of the above techniques. While mechanical means,
including relative spatial manipulations of the two coils or
mechanical switching or reorienting of the phase, amplitude and/or
frequency of the coil, current, voltage and RF fields might be
utilized to effect the field decoupling of the two coils, for
preferred embodiments the decoupling is accomplished electrically
or electronically. The actuation or control of such decoupling may
be by PIN diodes, solid state switches such as transistors, and
semiconductor relays, tube switches, electromechanical relays,
varistors, etc. In addition to the "active" electronic components
indicated above, "passive" components may also be used, including
small signal diodes, limiter diodes, rectifier diodes, etc., these
components often being used together with quarter-wave
circuits.
[0062] Further, by the general methods above, coil coupling can
alternatively be maximized for some applications. For example, it
may be desirable for coil 10 to be strongly coupled to a remote or
implanted surface coil where transmission-line coupling may be
impractical.
[0063] Because of its speed, power handling, compactness and
non-magnetic packaging, the PIN diode is a good choice for many
decoupling circuit implementations, including ones involving a coil
10. Such PIN diode circuits can be used to change the electrical
length of a coil or its individual paths, and to thus change the
resonant frequency of one coil relative to the other coil,
decoupling in this case being effected by frequency shifting. A PIN
diode circuit can also be used to open circuit or short circuit a
coil or individual paths thereof to effectively switch the coil on
or off, thereby decoupling it from the other coil. Similarly, PIN
diodes may be used to shift the phase of coil currents to minimize
the coupling between two coils.
[0064] FIG. 5A shows one way in which a PIN diode 56 may be
utilized to detune the paths 18 of coil 10, a separate PIN diode
switched circuit 56 being provided for each path 18 for this
embodiment. Each PIN diode shorts two points on the corresponding
path when conducting, for example a point on an inner wall to a
point on a side wall or outer wall, thereby altering the effective
length, phase or impedance of the path and thus its resonant
frequency.
[0065] FIG. 5B shows that this detuning technique can be used with
a coil of the type shown in U.S. Pat. No. 5,557,247, each
transmission line element 57 having an inner conductor 59 and an
outer conductor 61 separated by a dielectric 63. Each PIN diode 56
for this embodiment is connected through a solder post 60 to short
outer conductor 61 through choke coil 62 to a conductor 64 on rear
wall 68 of the coil. The diodes are current forward biased to short
the path, thus dramatically altering the coil's resonant frequency,
and thus decoupling the coil from another coil, a transmit or
receive coil of the same operational frequency. The resonant
frequency of each path may be quickly restored by voltage back
biasing the diode to disconnect the conductor 61 from the cavity
wall. This PIN diode switching approach is effectively changing the
impedance across an equivalent gap or azimuthal slot located in the
paths at position "C" in FIG. 1B. This or a similar approach using
a PIN diode circuit to change the impedance (higher or lower) in
some or all of the paths can be affected at other gap positions in
the paths such as A or B.
[0066] While in FIGS. 5A and 5B, PIN diodes 56 are used to short
paths 18 for detuning, PIN diodes could also be used to effect
detuning by placing the diodes in the path, for example two walls
being connected through a PIN diode for each path, the path being
open circuited for detuning. The PIN diodes could be used to
quickly switch reactance into or out of each path to change its
resonant frequency, or the PIN diodes could be utilized to effect
detuning in other ways. Further, for the embodiment shown in FIG.
5A, detuning may be effected by reversing the phase for the RF
drive signals applied to the various paths 18 so that electrical
phase is out of phase rather than in phase with the azimuthal
separation of the paths.
[0067] In the discussion so far, coil 10 has been assumed to have a
closed tubular configuration with an RF field mode M=1 or greater,
so that field is applied to a body position within the coil.
However, this is not a limitation on the invention and the coil
could be designed to operate in an M=0 mode for example, as taught
in U.S. Pat. No. 5,744,957. In particular, by, for example, not
slotting the inner wall 16I of the coil, or by having a two layer
overlapping inner conductor as shown in FIG. 3A, the RF field can
be confined in the cavity 12 to circulate therein, field not
exiting the cavity except where an opening is provided in one of
the walls of the cavity through which the field may exit or though
which a body to be imaged may be inserted into the cavity to be
exposed to the field circulating therein. Referring to FIGS. 7A and
7B, the coil 70 may be flat as shown, may have a slight curvature,
may be flat with a circular, elliptical or other appropriate shape
rather than a rectangular shape as shown or may have some other
appropriate shape so as to fit on the body being imaged with
minimal spacing, thereby achieving optimal coupling between the
coil and the body to be imaged, eliminating losses resulting from
dielectric constant mismatches and spaces between the coil and the
body. Coil 10 may for example be flattened to achieve a flat shape.
The resonant cavity of this invention may thus have a wide variety
of sizes and shapes, modes of operation, conductor patterning,
apertures, etc. Any cavity coil geometry is allowable provided that
an RF field can be generated therein which can be made useful for
MRI or EPR imaging applications. One or more openings 72 may be
provided in a wall of coil 70 which wall is to be adjacent to or in
contact with the body being imaged, each opening 72 being adjacent
a portion of such body on which imaging is desired. Holes are
preferably at B1 magnetic field nodes of the cavity wall. Where the
portion of the body on which imaging is desired is a projection on
the body, for example a woman's breasts, opening 72 may be
positioned as shown in FIG. 7B to permit such projections to enter
cavity 12 through the openings so as to be in the field path in the
cavity. The dielectric material in the cavity may be shaped or
deformable to fit projections extending into the cavity, minimizing
dielectric (impedance) mismatch Openings 72 are strategically
located and dimensioned to both encompass the body portions to be
imaged and to be properly phased. Openings 72 might also be used on
an inside or outside wall of a coil 10 designed to operate in M=0
mode or in a side wall.
[0068] In one application, apertures in for example a side wall of
a coil are each dimensioned to hold an experimental mouse, or the
mouse's head only, to permit a plurality of mice to be
batch/simultaneously imaged. In particular, referring to FIGS. 1A
and 1B, an embodiment of the invention is shown which is suitable
for batch nuclear magnetic resonance (NMR) study of multiple
laboratory samples, which samples may for example be held in test
tubes, or of lab animals such as mice. The coil 100 is of the type
previously described and has a cavity which is filled with a
dielectric material, which material can be a gas such as air, a
fluid or a solid. The dielectric material is preferably matched to
the sample to minimize the electrical impedance boundary between
the sample and the cavity for improved performance. A solid
dielectric can serve as the support wall for sample spaces 102. As
for previous embodiments, 16O, 16L 16F, and 16R are the outer wall,
inner wall, front wall and rear wall, respectively, of the cavity.
For this embodiment, these walls may be slotted to break up eddy
currents and shield the sample, as for the prior embodiments, or
may be continuous. The space or recess 26 inside wall 16I is not
used for this embodiment. While end wall 16F may be a conducting
wall, for example having slotted conductors between the sample
spaces 102 as shown in FIG. 9B, either in addition to or instead of
wall 16F being conductive, a conductive lid 104 may be provided
which is mounted to wall 16F in a manner to provide a good
electrical connection, for example being press-fitted. The lid aids
the coil's performance by completely enclosing the coil cavity 12
and thus more efficiently sealing energy in the coil. Where front
wall 16F is conductive, the lid could be non-conductive. The lid
could also be segmented, but this would require greater care in
mounting the lid to assure the lid segments align with the coil
segments.
[0069] The embodiment of FIGS. 9C and 9D is substantially the same
as the embodiments previously described, for example in conjunction
with FIGS. 1-4B, and in particular has an inner wall configuration
which is substantially the same as that of FIG. 3C. For this
embodiment, the useful space where a sample would be placed is the
center space 26, requiring inner wall 16I to be constructed of
conductive lands, tubes or transmission line elements 18 separated
by slots 20 wide enough to allow magnetic flux to efficiently fill
space 26. Outer wall 16O and end walls 16F and 16R could be any of
the configurations previously described, depending on application.
FIGS. 9C and 9D further illustrate the use of a lid 104 with these
embodiments of the invention, which lid may be conductive or
non-conductive and serves the same functions as those performed by
the lid 104 of FIG. 9A. FIG. 9D also illustrates an optional solid
back wall 106 which also may be continuously conductive or slotted
for the reasons previously discussed.
[0070] FIGS. 9E and 9F illustrate still another embodiment of the
invention. For this embodiment, center space 26 is filled with a
dielectric 108 having sample spaces 110 formed therein. Sample
spaces 110 may extend completely through dielectric 108 or may
extend only partly through the dielectric as for spaces 102. While
for the embodiment shown, dielectric 108 is solid, the gas/air or
liquid dielectric is utilized, sample spaces 110 can be formed of
non-conducting walls or the sample can be immersed in the gas or
liquid dielectric. For some embodiments, sample spaces 110 may be
tubes through which sample passes in a continuous NMR monitoring
process of a gas, liquid or solid sample flow or conveyance. The
cavity walls 16 for the embodiment of FIGS. 9E, 9F would be
substantially the same as those for the embodiment of FIGS. 9C, 9D
and, as for this prior embodiment, will vary with application.
Also, as for the prior embodiments shown in FIG. 9, the lid 104 is
optional, performing the same functions if utilized as for the
prior embodiments.
[0071] FIG. 10 illustrates several transmission line stub
embodiments of the invention which may for example be used for NMR
microprobe applications, permitting small samples to be efficiently
measured. Each of the transmission line stub embodiments 120,
120A-120C, may be capped with a conductive cap 104 to short-circuit
its center conductor 122 with its outer conductor 124.
Alternatively, the cap may be spaced or constructed in a manner
such that the center conductor and outer conductor of the
transmission line are left open-circuited or the cap may be
eliminated completely. The stub is preferably tuned and matched
such that the maximum current, and therefore maximum RF magnetic
field, is located at the sample end of the stub. For the current to
peak at the end, the length of the stub should be approximately a
one half wavelength or be electrically adjusted to be, a full
wavelength increment of the resonant frequency for an open stub and
one quarter wavelength or three quarter wavelength for a shorted
stub. The stub is typically connected to a transmission line to or
from the NMR system by a coaxial connector 126 which is best seen
in cross-sectional view 10D. A variety of coaxial connectors known
in the art might be utilized.
[0072] The sample spaces may be located within the dielectric space
128 between the conductors as shown in FIGS. 10A and 10B, or may be
located in a hollow, slotted center conductor as shown in FIG. 10C.
More particularly, outer wall 124 is the outer conductor or shield
of a typical coaxial transmission line. The inner wall is a center
conductor 122 of a typical coaxial line. The dielectric may be, as
for prior embodiments, gas, liquid or solid and impedance matched
to the sample. In a typical coaxial line, the dielectric is solid
and the sample space would be cut or drilled into the dielectric
material. The sample space 130A for the embodiment of FIG. 10A is
the space between the inner and outer conductors at the end of the
coil. The sample space 130B for the embodiment of FIG. 10B are
multiple sample spaces formed in the dielectric around the center
conductor. For the embodiment of FIG. 10C, the center conductor 122
is a slotted or ______ element conductor, the sample space 130C
being within the hollow end of the center conductor.
[0073] While in the discussion above it has been assumed that the
number of lands on each wall of the coil is the same so that N
continuous RF electrical paths are formed around the coil, this is
not a limitation on the invention. In particular, the number of
lands formed on each wall of the coil may not be the same. Thus,
the outer wall may have a first number of lands N.sub.1 and the
inner wall may have a second number of lands N.sub.2. The side
walls may also have N.sub.1 lands for reasons previously indicated.
N.sub.1 may, for example, be selected at least in part to
effectively breakup up low frequency eddy and other currents
induced in the coil, while N.sub.2 is generally selected to achieve
a desired magnetic field pattern. Even where N.sub.1 and N.sub.2
are not equal, adjacent paths on the walls are still connected to
form a plurality of continuous electric paths around the coil,
these paths providing various ones of the advantages previously
indicated.
[0074] As has been indicated earlier, one advantage of a coil 10 in
accordance with the teachings of this invention is that it can
provide a uniform, homogeneous field inside the coil for imaging
purposes. While such a homogeneous field is advantageous in many
applications, there are applications where some other field pattern
is desirable. Achieving such a patterned field through use of
spacing and polarization of the paths to which signals are applied
and the phasing of such signals has been discussed earlier. The
field may also be patterned by the choice, positioning, and control
of the dielectric in cavity 12 to obtain a desired field pattern.
Still another way of controlling field pattern is illustrated in
FIG. 8 where a dielectric "pillow" 80 is shown inside of coil 10
which dielectric is not part of cavity 12 and is selected to
provide a good dielectric match with body/patient 82 and/or with
the dielectric in cavity 12. The effect of dielectric pillow 80 is
to concentrate or otherwise manipulate the RF magnetic flux in a
region of interest in the patient's head where, for this
embodiment, imaging is desired. Dielectric inserts could be
otherwise positioned between coil 10 and the portion of a body on
which imaging is to be performed, or even within cavity 12, to
concentrate or manipulate field in such areas, thereby enhancing
measurement sensitivity in these regions of imaging and/or to
minimize field coupling to areas which are not to be imaged. The
shape of the insert, as well as its dielectric constant, is a
factor in achieving the desired control and shaping of the Rf
field, the shape of the insert also being useful to restrain body
motion, which motion, as previously indicated, can adversely affect
imaging.
[0075] Another potential problem with MRI and other imaging systems
utilizing RF coils is that rapidly switched currents in the field
gradients can generate intense acoustical noise. Such noise is
often annoying to a patient or even painful. One way in which such
noise can be reduced is by utilizing an acoustic damping material
in cavity 12 as at least part of the dielectric therein, such
acoustic dampening material either forming the entire dielectric,
or being used in conjunction with other dielectric material in
order to achieve a desired dielectric constant or pattern of
dielectric constants in the cavity so as to provide a desired
resonant frequency, field pattern and/or other features of the
invention.
[0076] Thus, while the invention has been particularly shown and
described above with reference to illustrative and preferred
embodiments, the foregoing and other changes of form and detail may
be made therein by one skilled in the art while still remaining
within the spirit and scope of the invention, which is to be
defined only by the appended claims.
* * * * *