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Description  |
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This invention relates to apparatus for determining nuclear magnetic
resonance (NMR), particularly to apparatus for imaging biological tissue,
and more particularly to such apparatus wherein the magnetic field is
produced substantially by permanent magnet materials.
BACKGROUND OF THE INVENTION
In the last few years advances in nuclear magnetic resonance (NMR)
techniques have made it possible to form two and three dimensional spin
density images of solids and liquids. A number of novel and sophisticated
variants have also been introduced to the rapidly expanding field of
imaging. An important aspect of all these developments is the ability to
form images of biological tissue in vivo. The NMR method is non-invasive
and has a much lower radiation hazard than the more usual X-ray imaging
methods.
In addition to producing spin density pictures, these new NMR imaging
techniques can all be adapted to measure spatial variations of the
spin-lattice relaxation time in a specimen. The cell water in cancerous
tissue, for example, is known to have longer spin-lattice relaxation time
than that in normal tissue. Thus NMR imaging, though in its infancy, holds
promise as a diagnostic tool for the early detection of tumors.
The NMR imaging techniques to be described all rely on the preparation
and/or observation of the nuclear spin system in the presence of one or
more magnetic field gradients. The field gradients serve to spatially
differentiate regions of the specimen by changing the Larmor resonance
frequency of the spins from one region of the specimen to another.
Individual protons or hydrogen nuclei are found in most organic and
biological material and have a natural isotopic abundance of 99.9844
percent. The other 0.0156 percent of nuclear sites is taken up with the
other naturally occurring heavy hydrogen isotope, deuterium.
Each nucleus has associated with it a small nuclear magnetic moment and a
quantity of angular momentum called spin. Regarded classically, the
combined effect of magnetic moment and spin causes a proton to precess
about the direction of an applied static magnetic field much as a spinning
top precesses about the gravitational field direction if perturbed from
the upright position. For protons, the precessional frequency is
independent of the angle of inclination of the magnetic moment with
respect to the static magnetic field and is called the Larmor angular
frequency .omega..sub.o. However, it does depend directly on the magnitude
of the static magnetic field B.sub.o through the relationship
.omega..sub.o =.gamma.B.sub.o, where the constant is called the
magnetogyric ratio. This relationship is the key to much of what follows.
If B.sub.o is varied then .omega..sub.o will vary. If a linear magnetic
field gradient is superimposed on an otherwise spatially uniform B.sub.o,
then the protons in a specimen placed in these fields would experience a
magnetic field higher than B.sub.o in some places and lower in others. An
account of the development of NMR spin imaging is given by Mansfield,
Contemp. Phys. Vol. 17, No. 6, pp. 553-576 (1976). The basic principles of
NMR necessary to understand imaging are discussed and main methods of
imaging are described and illustrated with examples of images of proton
spin distributions in a number of biological specimens.
NMR imaging of humans for medical diagnostic purposes presents the magnet
designer with formidable problems of an unusual nature Hoult et al., Rev.
Sci. Instrum. 52(9), pp. 1342-51 Sept. (1981) state that a magnet is
required which produces a field of at least 0.1 T with a homogeneity of
preferably 1 ppm over the region of interest of the patient, say the head
or torso In addition, Hoult et al. state that linear field gradients of up
to 10.sup.-2 Tm.sup.-1 in any direction may be required. A short term
field stability of better than 0.1 ppm may be mandatory over a period of a
second in order to avoid phase noise on the NMR signal, while the long
term stability may need to be about 1 ppm. Further, all this must be
accomplished in a hospital environment where it is likely that serious
perturbations will be caused by large amounts of steel (reinforcement,
water pipes, etc.), in the building structure, and by other ferromagnetic
objects which will be moving (elevators, beds, nearby trucks, etc.). Iron
core electromagnets have not been used because of difficulty in achieving
the required field uniformity and because of the excessive weight of these
magnets. Current designs are therefore generally air-cored electromagnets
of either resistive or superconducting design. A spherical shaped
electromagnet for NMR imaging is described by Hoult et al., supra. A
superconductive NMR magnet for in vivo imaging is described by Goldsmith
et al., Physiol. Chem. & Phys., 9, pp. 105-107 (1977). In addition, Hanley
discussed superconducting and resistive magnets in NMR scanning in a paper
presented at the 1981 International Symposium on Nuclear Magnetic
Resonance Imaging held at the Bowman Gray School of Medicine, Wake-Forest
University, Winston-Salem, N.C. A superconducting magnet can attain much
higher fields than a simple electromagnet but its cost will be much
higher.
A permanent magnet system would be superior to prior art superconducting
and resistive electromagnet designs in the following ways:
(a) There is no need for a means of generating the large amounts of power
required to maintain the field as in the resistive magnet systems.
(b) There is no need to provide cooling means to either remove generated
heat as in the resistive magnets or to maintain cryogenic temperatures as
in the superconducting magnets.
(c) The field of the permanent magnet is not subject to power supply drift
like that of resistive magnets or superconducting magnets not operated in
the persistant current mode.
(d) The field of the permanent magnet is not subject to gradual decay like
that of superconducting magnets operating in the persistent current mode.
(e) The material used can be a readily available ferrite magnet material
that is transparent to electromagnetic waves of the frequencies of
interest (5 MHz to 15 MHz).
(f) The external field strength falls off rapidly with distance away from
the magnet, leading to significantly reduced interference with the bias
field from ferromagnetic objects in the vicinity of the apparatus.
NMR permanent magnet designs using a ferromagnetic pole structure do not
make efficient use of the permanent magnet magnetic flux potential. Thus,
new permanent magnet NMR designs making more efficient use of the magnetic
flux potential are desired.
SUMMARY OF THE INVENTION
The present invention provides a permanent magnet NMR imaging apparatus for
imaging of biological tissue. The permanent magnet NMR imaging apparatus
in accord with this invention comprises bias means for generating a bias
field, means for generating gradient fields, and radio frequency means for
applying a pulse of electromagnetic radiation to the nuclear spins
associated with biological tissue at their Larmor frequency and detecting
the resultant signals emitted by them; wherein said bias means comprises a
plurality of dipole ring magnets, each dipole ring magnet comprising a
plurality of segments comprising an oriented, anistropic permanent magnet
material. The NMR imaging apparatus of this invention does not use a
ferromagnetic pole structure and makes efficient use of the magnetic flux
potential.
Preferably each dipole ring magnetic comprises eight segments of permanent
magnet material arranged in a ring so that the easy axis orientation is
determined by the formula
.alpha.=2.theta.-(.pi./2)
where .theta. is the angle between the radial symmetry line of a segment
and the x-axis (which is in the plane of the dipole ring magnet) and
.alpha. is the angle between said radial symmetry line and the x-axis.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is an isometric view of a permanent magnet NMR imaging apparatus in
accord with the present invention
FIG. 2 is a front elevational view partially cut away of the NMR imaging
apparatus illustrated in FIG. 1.
FIG. 3 is a cross-sectional view of the apparatus of FIG. 2 taken along
line 3--3 of FIG. 2.
FIG. 4A is a sketch illustrating an embodiment of an NMR imaging apparatus
of this invention having three ring dipoles.
FIG. 4B is a sketch illustrating an embodiment of an NMR imaging apparatus
of this invention having four ring dipoles.
FIG. 4C is a sketch illustrating an embodiment of an NMR imaging apparatus
of this invention having three ring dipoles wherein the center ring dipole
has a smaller outer radius than the outer ring dipoles.
FIG. 4D is a sketch illustrating an embodiment of an NMR imaging apparatus
of this invention having five ring dipoles wherein two of the ring dipoles
comprise material having a negative residual inductance, B.sub.r.
FIG. 5 is a front elevational view of a segment of permanent magnet
material formed from individual bricks.
FIG. 6 is a front elevational view in cross-section of a form for laying up
individual bricks to form a segment of permanent magnet material of the
desired size and shape as shown in FIG. 5.
FIG. 7 is an exploded isometric view of a segment of permanent magnet
material with a backing plate.
FIG. 8 is an isometric view of a segment as shown in FIG. 7 mounted on a
positioning ram for forming a dipole ring magnet.
FIG. 9 is a side elevational view partly in cross-section of the
positioning ram and mounted segment of FIG. 8.
FIG. 10 is a plan view of a plurality of positioning rams mounted in a jig
for forming a dipole ring magnet.
FIG. 11 is an isometric view of a dipole ring magnet positioned with
respect to an imaging volume illustrating the measurement of the magnetic
flux density in the imaging volume contributed by the ring dipole.
FIG. 12 is a front elevational view, partly in cross-section, of a fixture
for adjusting the spacing between dipole ring magnets for tuning an NMR
imaging apparatus wherein one dipole ring magnet has already been placed.
FIG. 13 is a front elevational view, partly in cross-section, of the
fixture of FIG. 12 having four dipole ring magnets placed therein and with
safety stops in place.
FIG. 14 is an isometric view of an assembled NMR imaging magnet in accord
with this invention.
FIG. 15 is a sketch of four ring dipoles in cross-section illustrating
their physical parameters.
FIG. 16 is a flow diagram of a method for optimizing the design of a system
of permanent magnet ring dipoles for NMR imaging.
FIG. 17 is a block diagram of a gradient coil system for NMR imaging.
FIG. 18 is a block diagram of an R. F. system for NMR imaging.
DESCRIPTION OF THE INVENTION
In accord with the present invention an NMR imaging apparatus is provided
wherein the bias magnetic field is generated by permanent magnet dipoles.
The bias magnetic field is substantially stationary and uniform. As used
herein "a substantially stationary and uniform magnetic field" is a
magnetic field that has sufficient uniformity and that is sufficiently
stable to obtain images of biological tissue as desired. Preferably the
bias magnetic field does not vary by more than 5.times.10.sup.-4 in the
design imaging volume or space in which the test specimen is placed, and
more preferably the bias field variance is less than 1.times.10.sup.-5.
Further, the field stability is preferably no worse than 5.times.10.sup.-6
(sec.sup.-1). However, as noted above these parameters can be varied
depending upon the acceptable quality of the image desired.
The invention will be further described with reference to the drawings
wherein FIG. 1 illustrates an NMR imaging apparatus 10, in accord with one
embodiment of the invention, having an opening of sufficient diameter to
accept an adult human for scanning. The NMR imaging apparatus 10 consists
of four collars or rings 20, each ring comprising a dipole magnet made of
permanent magnet material.
Each ring or collar 20 consists of eight segments 22 of permanent magnet
material. More or less segments can be used. However, eight segments
provide quite satisfactory results in the embodiment described. The
permanent magnet material is an oriented, anisotropic permanent magnet
material, such as a rare-earth/cobalt material, or a ferrite ceramic
material, or the like. Preferred materials have a unity permeability.
Suitable materials include, for example, sammarium cobalt, barium ferrite,
strontium ferrite, and the like. Conveniently each segment 22 has a
trapezoidal shape.
Each segment is built up to the desired size from individual bricks 23, as
illustrated in FIG. 5. Each segment 22 could also be made from a solid
block of such permanent magnet material. Bricks 23 positioned along the
surfaces of each segment 22 are suitably cut to provide the desired final
shape. To form each segment 22, the individual bricks 23 are laid up in a
form 25, which is conveniently made of fiberglass. The surfaces of each
brick 23 that will come in contact with the surface of other bricks are
coated with an adhesive prior to laying the bricks in the form 25.
Unmagnetized bricks having a dimension of about 15 cm.times.10 cm.times.2.5
cm are conveniently used to build each segment 22. The bricks are trimmed
to ensure that they have sharp corners so that they fit tightly together
in the form and leave no air gaps. The bricks are thoroughly cleaned of
oil, grease and loose material. A two-part adhesive has been found
convenient. The surface of one-brick is coated with, for example, Loctite
Loquic Primer N or its equivalent. The surface of the second brick that
contacts the coated surface of the first brick is then coated with, for
example, Loctite Superbonder 326, or its equivalent. After the bricks are
laid up in form 25, the bricks are allowed to cure for a sufficient time
so that the segment 22 can be removed from the form 25. Typically about
ten minutes is sufficient for the initial curing. However, the time will
vary depending upon the particular adhesive being used and other
conditions such as temperature. The initially cured segment 22 is
thoroughly cleaned with a degreasing solvent and sufficient time is
allowed to fully cure the adhesive.
The segment 22 is then placed in the magnetizing coil of a magnetizer
capable of producing a peak pulsed field of at least 8 kilo-oersteds
throughout the volume of the segment to fully magnetize the permanent
magnet material. The segment is clamped in the correct orientation in
accord with the formula .alpha.=2.theta.-(.pi./2) as aforesaid depending
upon the predetermined position of the segment in the ring 20. A suitable
fixture made of non-ferromagnetic, non-conducting material is used to
clamp the segment. After pulsing the magnetizer to magnetize the permanent
magnet material, the segment is removed and an aluminum backing plate 27
is bonded to the segment 22 using an epoxy resin 28 or the like as
illustrated in FIG. 7.
Each of eight segments 22 is then clamped in a positioning ram 30 as
illustrated in FIGS. 8 and 9, with the positioning ram 30 at its outer
stop. The fine adjustment screw 32 on each ram 30 is set near the middle
of its travel as illustrated in FIG. 9. When all eight segments are
positioned in their rams 30, in accord with the predetermined alignment of
the easy axis of each segment, the rams are moved forward to their
innermost stops as illustrated in FIG. 10.
The initial design and optimization of an NMR imaging apparatus in accord
with the present invention will now be illustrated with respect to the
configuration shown in FIG. 15. Consider four ring dipoles placed along an
axis as shown in FIG. 15. The field B.sub.y (x, y, w) is defined as:
##EQU1##
wherein x, y and w are the spatial coordinates of any point with the w
axis being the axis of the ring dipoles, N is the number of ring dipoles,
and g is the remanent field. The variables a, b and z are physical
parameters of the ring dipole configuration as illustrated for four (4)
ring dipoles in FIG. 15.
The design of the magnet for NMR imaging is optimized in accord with the
flow diagram illustrated in FIG. 16. Values for U.sub.o, U.sub.3 and b/a
are selected: e.g. U.sub.o =0.05, 0.10, 0.15, 0.20; U.sub.3 =0.9; b/a=1.8,
2.0, 2.2, 2.5.
The expansion of the field on the axis near the center of the configuration
will be of the form.
B.sub.y =B.sub.0 +B.sub.2 (z/a).sup.2 +B.sub.4 (z/a).sup.4 +B.sub.6
(z/a).sup.6 + . . .
Adjust the geometric parameters so that B.sub.2 =B.sub.4 =0. When B.sub.2
and B.sub.4 are approximately 0, calculate B.sub.o and B.sub.6. Also
calculate the volume V of permanent magnet material. Vary the values of
U.sub.o, U.sub.3 and b/a to obtain the desired B.sub.o and minimum volume
with acceptable distortion (non-uniformity of field), approximately equal
to B.sub.6.
Check to see that the distortion in the radial direction is within
acceptable limits by analyzing the field distortion at points off the
center axis. Introduce small amounts of B.sub.2 and B.sub.4 terms so that
the distortion at the extremities of the working volume is reduced at the
expense of a small increase in distortion in the interior region of the
working volume, thus lowering the overall distortion in the working
volume.
Although the procedures described herein are typically used to obtain the
optimum field uniformity in the imaging field, the procedures can be used
to obtain less than optimum field uniformity, if conditions do not require
the optimum.
After the design is optimized by the above procedure, the individual ring
dipoles are assembled using the fixtures illustrated in FIGS. 8 and 9 and
are tuned by the following procedure.
The radial position of segments 22 of the ring dipole is adjusted to
eliminate non-uniformities in the dipole magnetic field in the designed
imaging volume 65 by measuring the magnetic flux density in the imaging
volume using a Hall effect probe 60 as illustrated in FIG. 11. First, the
harmonic content of the field for each of the first eight (8) harmonics
(corresponding to the eight segments) of the magnetic field in the
designed imaging volume is determined by measuring the flux density at a
series of points. Then, the first segment is moved radially a small
distance by means of the fine adjustment screw 32 of the positioning ram
30. The harmonic content of the field in the designed imaging volume 65 is
remeasured. Then the segment 22 is returned to its initial position. Each
segment in succession is displaced radially a small distance and the
harmonic content of the field in the designed imaging volume 65 is
measured. After all eight segments have been displaced and the harmonic
content measured, an 8.times.8 sensitivity matrix is calculated. This
sensitivity matrix shows the sensitivity of the harmonic content of the
magnetic field in the design imaging volume to each segment of the ring
dipole. The elements of the matrix are defined by the following formula:
##EQU2##
where .delta.A.sub.1,n is the amount of change in harmonic content of the
"n"th harmonic due to the change in position of segment "1" and
.delta.r.sub.1 is the amount of change in position of segment "1".
After the sensitivity matrix has been calculated, the inverse of the
sensitivity matrix, or the correction matrix, is calculated. The harmonic
content of the field in the design imaging volume is remeasured. The
harmonic content for each harmonic is then subtracted from the ideal
harmonic content (or designed harmonic content) for each harmonic to
obtain a difference vector. The difference vector is then multiplied by
the correction matrix to obtain the tuning corrections, i.e. the distance
and direction each segment 22 must be moved to more closely approach the
ideal or design harmonic content of the design imaging volume 65
contributed by the ring dipole being tuned. Each segment is then moved the
calculated amount and the process is repeated until the harmonic content
is within specifications, i.e. the magnetic field uniformity is within the
design specification.
A collar structure (not shown) is then attached to the segments of the ring
dipole by mechanically attaching the collar to the backing plates of each
segment with fasteners or by means of adhesives. The dipole ring with its
segments fixed by the collar assembly is then removed from the positioning
rams and placed in the assembly fixture 40 as illustrated in FIG. 12 by
lowering it along rib guides 41.
Each successive ring dipole is tuned as described above and placed in the
assembly fixture 40 until the design number of ring dipoles, four (4) in
this case, are placed in the assembly fixture 40 as shown in FIG. 13. At
this point shims 42 have been placed between each of the ring dipoles as
each ring dipole is drawn into position adjacent the previous ring dipole
by guide arm actuating mechanisms 45. Due to the repelling forces between
each ring dipole the ring dipoles must be mechanically locked in position
before the next ring dipole is placed in the assembly fixture 40. After
the four ring dipoles are in place, safety stops 46 are bolted in place.
The apparatus 10 must now be tuned in the axial direction. The field
strength along the axis of the apparatus 10 can be expressed as a power
series
##EQU3##
where B.sub.z is the field at point z on the axis and C.sub.n are axial
coefficients. The first three axial coefficients are determined by
measuring the magnetic flux density at various positions along the axis
using Hall effect probe 60. Then the position of the first ring dipole is
changed relative to the designed imaging volume by changing the thickness
of the tuning shim 42 to change the separation between the first and
second dipole. The axial coefficients are then redetermined. The initial
tuning shim is replaced between the first ring dipole and the second ring
dipole and the distance between the second and third ring dipoles is
changed. After determining the axial coefficients for that change, the
initial shims are replaced and the process repeated for the separation
between the third and fourth ring dipoles. A sensitivity matrix is
calculated, similar to the radial tuning, wherein the elements of the
matrix are the change in axial coefficient divided by the change in
separation. The inverse of the sensitivity matrix, i.e. the correction
matrix, is calculated and multiplied with the difference vector calculated
from the measured axial coefficients and design axial coefficients. The
above multiplication provides the corrections to be made to the
separations between the ring dipoles. The process is repeated until the
magnetic field with the design imaging volume is within the design
specification.
When the magnetic field within the design imaging volume is within the
design specification for uniformity, permanent shims are machined from a
suitable non-ferromagnetic material to maintain the desired separation
between the ring dipoles. After the permanent shims are positioned between
the ring dipoles, the collars of the ring dipoles are mechanically fixed
together by bolting or welding structural beams 47 to the collars of the
ring dipoles.
The apparatus 10 is removed from the axial assembly fixture and placed on a
base 50 as shown in FIG. 14. The permanent magnet material has a
reversible variation of its magnetization with temperature changes. If the
temperature varies by more than approximately 1.degree. C. from place to
place in the magnet, the field will suffer a temporary distortion. This
can be prevented by placing 3 cm of thermal insulation (e.g.
urea-formaldehyde foam) over the entire outside surface of the finished
magnet as shown in FIG. 14. End covers 54 and side panels 55 are attached
for a finished look
NMR imaging apparatus having configurations of three, four, and five ring
dipoles such as illustrated in FIGS. 4A, 4B, 4C and 4D can be optimized
and tuned in a similar manner
An NMR imaging system in accord with the invention also has a gradient
field superimposed on the bias field provided by the ring dipole apparatus
described above. The gradient field can be provided by any means
previously used for providing the gradient field for previous NMR imaging
systems wherein the bias field was provided by electromagnets. An "air
core" current gradient coil system is preferred. The gradient coil can be
located outside the bias magnet as in the NMR imaging system of Lauterbur
et al. at the State University of New York at Stony Brook or inside the
bias magnet adjacent the imaging volume. The power supply to the gradient
coil is supplied as illustrated by the block diagram of FIG. 17. A fixed
D.C. system power supply (SPS) 70 is coupled to the coil system 72 by a
power amplifier 71. The actual field gradients are controlled by a
microprocessor 75. The microprocessor is programmed to control pulse
polarity, pulse height, pulse width, pulse shape, and duty cycle.
The radio frequency (R.F.) coil detects the nuclear magnetic moment of the
hydrogen atoms in the biological tissue. The R. F. System is designed
using well known techniques such as those described by Hoult in his paper
on "Radio Frequency Coil Technology in NMR Scanning" presented at the 1981
International Symposium on Nuclear Magnetic Resonance Imaging held at the
Bowman Gray School of Medicine, Wake-Forest University, Winston-Salem,
N.C.
FIG. 18 is a block diagram for the R. F. System. It is divided into eight
subsystems. The function of subsystem 1, the programmer, is to take
instructions from the microprocessor computer, such as an LSI-11
microprocessor controller, and translate these into the proper voltage
signals needed to operate the various gates, phase shifters, etc. in the
system. Subsystem 2, the transmitter, provides, under control of the
programmer, R. F. pulses of the proper frequency, phase and envelope shape
to excite the spin system. Subsystem 3, the power amplifier, amplifies the
R. F. pulses provided by the transmitter, and matches their impedance to
the transmitting antenna. Subsystem 4 includes the T/R switches for the
transmitter and receiver and the transmitting and receiving antennas
(whether or not they are the same or separate structures). The calibrator,
subsystem 5, provides a signal of the right frequency and a known strength
to be injected into the receiver periodically to check the receiver
sensitivity and prevent gain drift problems. Subsystem 6 is the receiver
which senses the NMR signal and converts it to a useable analog signal.
Finally, subsystem 7 is the output interface, which converts the analog
signal from the receiver to a digital signal that can be fed to the
imaging computer.
The imaging computer reconstructs two or three dimensional images from the
data obtained as a function of changing magnetic field gradients in accord
with known techniques, such as those using Fourier transformations.
Although the invention has been described in detail for an NMR imaging
apparatus comprised of four (4) ring dipoles each having eight segments of
permanent magnet material forming a substantially continuous ring, the
methods described are equally applicable to such systems having more or
less ring dipoles and to ring dipoles made of more or less segments. In
fact, the most preferred axial arrangement of ring dipoles consists of a
positive ring dipole in the center surrounded alternately by two small
negative ring dipoles and then two larger ring dipoles as illustrated in
FIG. 4D. However, for practical reasons an arrangement as illustrated in
FIG. 4B is preferred for an actual apparatus.
The invention has been described in detail including the preferred
embodiments thereof. However, it will be appreciated that those skilled in
the art, upon consideration of the present disclosure, may make
modifications and improvements within the spirit and scope of this
invention.
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