Magnetic resonance imaging magnet having minimally symmetric ferromagnetic shield

A magnetic resonance imaging magnet has a solenoidal superconducting winding producing a magnetic field and defining an examination region within a bore thereof. The examination region receives a body to be examined. The solenoidal superconducting winding generates a quasi-static magnetic field for aligning atomic nuclei in the body being examined. A two-fold or minimally symmetric ferromagnetic shield has a pair of magnetic flux return paths positioned on opposite sides of the solenoidal superconducting winding for reducing the magnetic field intensity in a region proximate to and outside of the solenoidal superconducting winding. A magnetic field inhomogeneity compensating system reduces magnetic field inhomogeneities introduced into the magnetic field within the examination region by the minimally symmetric ferromagnetic shield to provide a highly uniform quasi-static magnetic field within the examination region.

BACKGROUND OF THE INVENTION 
The invention is related to a passive, ferromagnetic shield for a magnetic 
resonance imaging magnet. In particular, the invention is related to a 
magnet having a rectangular, ferromagnetic shield with an open top and an 
open bottom and an associated plurality of magnetic field inhomogeneity 
compensating devices for ensuring that a high homogeneity magnetic field 
is generated in a patient examining region within a bore of the magnet. 
It is well known to employ magnetic resonance imaging techniques as a 
diagnostic tool in the treatment of disease. Magnetic resonance imaging 
equipment, however, is notoriously expensive, in part because a 
quasi-static or static highly homogeneous magnetic field must be generated 
within the examination region occupied by a patient's body. Often a 
superconducting magnet of the type disclosed in U.S. Pat. No. 4,782,671 
for Cooling Apparatus for MRI Magnet and Method of Use and assigned to the 
instant assignee, is employed to generate the static magnetic field. If 
the quasi-static magnetic field is not homogeneous to within less than 
twenty parts per million over the diameter spherical volume (DSV), the 
field inhomogeneities can prevent an accurate depiction of the internal 
condition of the patient. 
In addition to the requirement that the static magnetic field be highly 
uniform, it must also be relatively intense. As a consequence, it has been 
found necessary in most commercial embodiments of magnetic resonance 
imaging equipment to employ superconducting magnets for generating the 
magnetic field strengths required to image the requisite detail within the 
human body. As a result of having these high field strengths, it has been 
found that it is necessary to engage in relatively elaborate shielding of 
the magnetic field. The Food and Drug Administration of the United States 
Government has required that the magnetic resonance imaging equipment be 
provided with an area of exclusion bounded by a five gauss flux line or 
surface in order to avoid interference with other hospital diagnostic 
equipment, as well as with devices such as neurostimulators and cardiac 
pacemakers. 
A number of methods have been employed in the past for shielding magnet 
resonance imaging magnets. In one method employing active shielding, an 
additional coil or coils are wound about the superconducting coil for 
generating a magnetic field which when added to the MRI field 
substantially reduces the magnetic field external to the imaging unit and 
thereby reduces the volume bounded by the five gauss surface. The problem 
with such a system is that it is relatively bulky and expensive to 
manufacture and often requires the use of extra superconducting magnets. 
The extra magnets in turn require additional cryogenic refrigeration 
capacity and the like. This can add as much as thirty percent to the cost 
of a magnetic resonance imaging system. 
Other magnetic resonance imaging systems employ passive magnetic shielding. 
In one type of passive shielding the magnetic resonance imaging system may 
be placed within a room having walls consisting of ferromagnetic material 
which provides return paths for the magnetic flux. The room, of course, 
cannot be used for anything other than magnetic resonance imaging due to 
the high flux within the room when the superconducting magnet has current 
circulating in it. In addition, ferromagnetic rooms are relatively 
expensive to build and place high structural demands upon the building in 
which they are situated due to the weight of the ferromagnetic material. 
Another approach to limiting the external magnetic field is to use a 
passive ferromagnetic shield having a plurality of symmetric magnetic 
return paths exterior to the magnetic resonance imaging magnet as is 
disclosed in U.S. Pat. No. 4,743,880 to Breneman, et al. 
The number of flux return paths may be reduced in order to reduce the cost 
of the magnet, provide better access to the internal portions of the 
magnet and provide horizontal plane shielding. Unfortunately, if 
asymmetric instead of symmetric magnetic flux return paths are employed, 
it has been found that inhomogeneities are introduced into the magnetic 
field within the magnet bore which render the magnetic resonance imaging 
system useless for diagnostic purposes. 
In order to provide a high-resolution image using nuclear magnetic 
resonance equipment, it is important to control precisely the magnitude 
and direction of the quasi-static magnetic field. The quasi-static 
magnetic vector field determines in part the frequency at which the 
hydrogen nuclei precessing within the magnetic field will undergo spin 
flips evidenced by absorption of radio frequency energy of a pre-selected 
frequency injected into the examination space. If the field varies in 
magnitude or in direction, and if a pair of gradient fields are added to 
it in order to provide spacial localization for the energy absorption 
signal, the inhomogenieties in the primary quasi-static field will reduce 
the resolution of the magnetic resonance imaging apparatus to the point at 
which it is impossible to obtain images of adequate resolution. 
Another problem with which users of magnetic resonance imaging magnets are 
faced is compliance with Food and Drug Administration standards requiring 
that areas of the hospital, clinic or trailer in which the magnetic 
resonance magnet is located are not subject to a magnetic field intensity 
greater than five gauss. As a result, most magnetic resonance imaging 
superconducting magnets are shielded in order to reduce the volume bounded 
by the five gauss surface. 
The shielding may take the form of a room constructed about the magnet of 
the type disclosed in U.S. Pat. No. 4,646,046 to Vavrek, et al. for 
Shielded Room Construction for Containment of Fringe Magnetic Fields. 
Other shields may be cylinders built about the magnet with closely spaced 
flux return bars of the type disclosed in U.S. Pat. No. 4,646,045 to 
Chari, et al. for Aperture Size Disc Shaped End Caps of a Ferromagnetic 
Shield for Magnetic Resonance Magnets. Still other shielding devices 
employ multiple flat plates, which provide flux return paths as taught in 
the octagonal structure disclosed in U. S. Pat. No. 4,590,452 to Ries, et 
al. for Magnetic Device of Apparatus in Nuclear Spin Tomography With a 
Shielding Device. Some prior magnets employ ferromagnetic cylindrical 
shells of the type disclosed in U. S. Pat. No. 4,590,428 to Muller, et al. 
for Electromagnet for NMR Tomography for shielding. 
Other workers in the art have provided shielded magnetic structures wherein 
the superconducting coil wound therein is not wound on a helix, but rather 
is wound in a variable fashion in order to compensate for perturbations of 
the magnetic field by the shield, however, tesseral or off-axis components 
of the magnetic field cannot be compensated by variations in a 
substantially helically wound coil. Unfortunately, all of these prior art 
approaches suffer from one or more drawbacks. 
The Burnett, et al. approach in U. S. Pat. No. 4,694,269 for a Magnet 
System and Method of Its Manufacture requires that the magnet coil be 
precisely wound in a shape other than a helix so that field perturbations 
may be compensated for. In some cases, however, customers using magnetic 
resonance imaging equipment in nonmedical environments may find it 
unnecessary to provide the type of shielding required by the FDA for use 
in a medical environment. As a result, if the customer elects to leave the 
shielding off the magnet in order to reduce cost, the pre-wound corrective 
coils of Burnett, et al. will introduce perturbations into the internal 
field in the examination space. 
Complete shields of the type disclosed in Muller, et al., U. S. Pat. No. 
4,590,428 are difficult to work with, since complete shields are 
relatively heavy, due to the weight of the ferromagnetic material, such as 
cold-rolled or hot-rolled steel having a thickness ranging from 1 inch to 
21/2 inches. The Muller shield must be removed from the magnet before 
access can be had to the chambers containing the superconducting coil or 
the liquid helium or liquid nitrogen The system taught by Ries, et al., U. 
S. Pat. No. 4,590,452, renders the magnet larger than necessary, which 
would require that the floor of the building in which the magnet is to be 
located be reinforced to carry the weight of the magnet. The structure 
taught in U. S. Pat. No. 4,612,505 to Zijlstra for Nuclear Magnetic 
Resonance Apparatus employing the extremely long cylindrical bars ranged 
about the magnet, consumes a great deal of space which would make it 
undesirable to use the magnet in portable or mobile applications. The 
magnet of Chari, et al., U.S. Pat. No. 4,646,045 would be relatively 
expensive to build due to its cylindrically arranged flux return bars. In 
addition, the Chari magnet itself is completely enclosed, which prevents 
convenient access to the interior. 
What is needed is a magnetic resonance imaging magnet having an easily 
constructed ferromagnetic return path which is relatively light and a 
system for compensating for magnetic field inhomogeneities introduced into 
the examining area of the associated superconducting magnet. The 
superconducting magnet should generate a solenoidal magnetic field in an 
examination region which is homogeneous to within less than twenty parts 
per million to provide a quasi-static field for the production of high 
resolution images by a magnetic resonance imaging apparatus. 
SUMMARY OF THE INVENTION 
Among the various aspects and features of the invention may be noted the 
provision of a shielded magnetic resonance imaging magnet and a method of 
shielding a magnetic field generating device. 
Briefly, the magnet includes a superconducting coil for generating an 
intense quasi-static magnetic field. A minimally symmetric or two-fold 
symmetric shield partially surrounds the superconducting coils and reduces 
the DSV in which external magnetic field intensity exceeds five gauss. The 
shield covers the ends of the superconducting coil and two of the sides, 
but not the top and the bottom. The minimally symmetric is so named 
because its side plates and end plates are symmetric about a plane through 
a bore defining an examination region. The plane of symmetry is parallel 
to the side plates and transverse to the end plates. Although the shield 
is composed of cold-rolled or hot-rolled steel in thicknesses of one inch 
or more, the shield is considerably less massive than those previously 
used because selective horizontal shielding need only be provided. While 
the external field is adequately compensated by the shield, the shield 
introduces inhomogeneities into the magnetic field produced within an 
examination region within a bore of the superconducting coil. A 
compensating system consisting of ferromagnetic rings for compensating for 
zonal inhomogeneities and compensating bars for compensating for off-axis 
or tesseral inhomogeneities is also included. The compensating system 
reduces inhomogeneities in the magnetic field expressed in Legendre 
coefficients up to the sixth order to less than one to two parts per 
million over the DSV of interest. 
It is a principal aspect of the present invention to provide a magnetic 
resonance imaging magnet having a reduced weight and size ferromagnetic 
shield. 
It is another aspect of the present invention to provide a magnetic 
resonance imaging magnet having a compensating system which can reduce 
shield induced magnetic field inhomogeneities within an examination 
region. 
It is a still further aspect of the present invention to provide a shielded 
magnetic resonance magnet having off-axis or tesseral compensating devices 
for reducing off-axis magnetic field inhomogeneities caused by a minimally 
symmetric ferromagnetic shield. 
Other aspects and advantages of the present invention will become apparent 
to one skilled in the art upon a perusal of the following specification 
and claims in light of the accompanying drawings.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
Referring now to the drawings and especially to FIG. 1, a shielded magnet 
embodying the present invention and generally identified by numeral 10 is 
shown therein. The superconducting magnet 10 has means for producing a 
magnetic field comprising a solenoidal superconducting coil 12 defining a 
bore 13. A two-fold or minimally symmetric passive ferromagnetic shield 14 
partially surrounds the solenoidal superconducting coil 12 for reducing 
the exterior magnetic flux in a region about the magnetic adjacent 
thereto. Means for compensating for magnetic field inhomogeneities 16 is 
positioned within an examination region 18 defined by the solenoidal bore 
13 for receiving a patient for imaging purposes. 
The solenoidal superconducting coil 12 is composed of niobium-titanium wire 
or other conventional superconducting materials and is wound to generate a 
solenoidal magnetic field within the examination region 18. The solenoidal 
superconducting coil 12, without more, would also generate an intense 
external field, this is in part due to the fact that the field within the 
examination region usually has an intensity of 0.3 to 1.5 tesla or 3,000 
to 15,000 gauss. Since the divergence of any magnetic field must be zero, 
the magnetic flux exiting the examination region and entering at the 
opposite end of the examination region follows flux return paths within 
the shield 14 this creates magnetic dipoles within the compensating means 
whose own magnetic dipole flux return paths are superposed over the free 
space flux return paths and greatly reduce the external field. This lowers 
or eliminates the magnetic field outside the magnet where electronic and 
magnetic devices, such as neurostimulators and the like, otherwise might 
be disturbed. 
Flux return paths are provided in the form of ferromagnetic material, 
specifically cold-rolled or hot-rolled steel. In some applications where 
medical imaging is to be done, the FDA has required that an exclusion area 
be demarcated around a magnet resonance imaging magnet to identify the 
region within which the field intensity exceeds 5 gauss. This is ten times 
the field strengths of the earth's magnetic field (0.5 gauss). In the 
instant invention the two-fold symmetric ferromagnetic shield 14 provides 
the flux return paths for reducing the DSV outside the magnet having a 
flux greater than 5 gauss. The portions of the flux return paths which are 
parallel with the examination region 18 are disposed on opposite sides of 
the solenoidal superconducting coil 12. The flux return paths of the 
two-fold symmetric ferromagnetic shield comprise a rectangular cold-rolled 
steel end plate 20 comprising an outer one inch thick plate 20a and an 
inner one inch thick plate 20b both of cold-rolled steel, and having a 
circular aperture 22 formed therein for access to the examination region. 
The plates 20a and 20b are held together by interrupted tungsten inert gas 
welds along their edges. A rectangular cold-rolled steel side plate 26 
comprising an outer plate 26a having a thickness of one inch, an 
intermediate plate 26b having a thickness of one-half inch and an inner 
plate 26c having a thickness of one inch all of cold-rolled steel, is 
tungsten inert gas welded to an edge 28 of the end plate 20 as may best be 
seen in FIG. 6. The plates 26a, 26b and 26c are tungsten inert gas welded 
together by a plurality of interrupted welds along their edges. A 
multi-layer rectangular cold-rolled steel side plate 30 is connected to an 
edge 32 of the end plate 20 so that the end plate 20 and the side plates 
26 and 30 provide a low reluctance path for the magnetic flux exiting the 
examination region 30. A multi-layer rectangular cold-rolled steel end 
plate 40 is connected at its edges to the side plates 26 and 30. It may be 
appreciated that a top portion 50 and a bottom portion 52 of the magnetic 
resonance imaging magnetic are left open so that a liquid nitrogen fill 
port 54 or a power lead access port 56 connected to the superconducting 
coil 12 may be accessed without having to disassemble the magnet 10 or 
remove a portion of the shield 14. 
The introduction of the two-fold symmetric shield 14 however, perturbs the 
magnetic field within the examination region 18. In order to insure that 
the solenoidal field within the examination region 18 is homogeneous to 
within less than twenty parts per million, the compensating means 16 is 
employed. The compensating means 16 comprises a plurality of symmetric 
compensating rings 60, as may best be seen in FIGS. 1 and 3, which are 
positioned on opposite sides of the examination region 18 and attached by 
gluing or tungsten inert gas welding to an inner wall 58 defining the bore 
13. The symmetric rings 60 compensate for magnetic field perturbations 
which cause the field intensity to change as the z-axis 19 is traversed. 
The symmetric compensating 60 rings are comprised of ferromagnetic 
material, specifically cold-rolled steel having a thickness of 1/32 inch 
to 1/2 inch and a width of 2 inches. They are disposed on opposite sides 
of the examination region 18. The plurality of compensating rings 60 
comprises a first compensating ring 62 having an eddy current reduction 
gap 63, a second compensating ring 64 having an eddy current reduction gap 
65, a third compensating ring 66 having an eddy current reduction gap 67, 
as shown in FIG. 5, and a fourth compensating ring 68 having an eddy 
current reduction gap 69. The compensating rings 60 each have a magnetic 
field induced therein by the magnetic field produced by the solenoidal 
superconducting coil 12. The magnetic fields induced in the compensating 
rings are aligned with the inducing field at locations within the 
compensating rings 60. Outside the compensating rings 60, the induced 
magnetic flux lines close in loops which are oriented substantially 
oppositely to the solenoidal field within the bore 13 to reduce magnetic 
field inhomogeneities within the examination region 18. 
Gradient coils, which are not shown in the drawings must be used to 
generate well resolved images. As may best be seen in FIG. 5, an eddy 
current such as eddy current 70, is induced by time-varying magnetic 
fields in each of the compensating rings 62, 64, 66 and 68 when the 
gradient coils are pulsed. The compensating rings 62, 64, 66 and 68 are 
interrupted by their respective eddy current reduction gaps 63, 65, 67 and 
69 each having a gap width of about one-sixteenth to one-eighth of an 
inch. The eddy current reduction gaps 63, 65, 67 and 69, by introducing 
additional electrical impedance into their respective compensating rings 
62, 64, 66 and 68, reduce the eddy currents induced therein. Additionally, 
the gaps are short enough that the magnetic dipoles induced in the 
compensating rings are not effected materially. 
Almost any type of passive ferromagnetic shield tends to introduce zonal 
inhomogeneities into the magnetic field within the examination region. 
However, the use of the two-fold symmetric shield of the type disclosed 
herein, also tends to introduce tesseral inhomogeneities into the magnetic 
field. Tesseral or off-axis inhomogeneities are perturbations in the 
magnetic field which are encountered as the angle theta changes. In order 
to compensate for the tesseral or off-axis inhomogeneities, a plurality of 
compensating bars 80, comprising compensating bars 81, 82, 83 and 84, is 
attached to the interior wall 58 by glue or TIG welding. The compensating 
bars 80 remove the X&gt;-Y&gt; and/or XY tesseral or off-axis magnetic field 
inhomogeneities from the magnetic field by causing a portion of the field 
within the examination region to be confined within the compensating bars 
80. These off-axis inhomogeneities vary with the angle theta. It may be 
appreciated that the two-fold symmetric shield has a tendency to "flatten" 
the entire magnetic field, that is to increase the field strength as theta 
approaches zero and decrease the field strength as theta approaches D/2 or 
3D/2. The compensating bars 80 are composed of rectangular cold-rolled 
steel members or of bundles of rods Each of the compensating members is 28 
to 36 inches long. The diameter of the rods may vary from 1/32 inch in 
diameter and having a length of 12 inches, to 1/2 inch diameter having a 
length of 28 to 36 inches. Unfortunately, the compensating bars 80 also 
introduce zonal perturbations or z-axis perturbations within the magnetic 
field which must be removed. 
In order to compensate for the zonal perturbations introduced when the 
tesseral perturbations are removed, a pair of dipole rings 90 and 92, as 
best shown in FIG. 3, positioned between the compensating rings 64 and 66, 
are included. The dipole ring 90 is made up of a plurality of rectangular 
ferromagnetic members 96, 98. The dipole ring 92 is composed of a 
plurality of regularly spaced identical members 98. The members 96 and 98 
are spaced at from 15 to 45 degree intervals about the inside wall 58 of 
the magnet bore 13 with a preferred angular spacing of 30 degrees, for 
reducing the zonal inhomogeneities in the magnetic field. Since the 
members 96 are spaced apart but are spaced uniformly, the field reduction 
effect is angularly substantially uniform and limited by the reduced 
ferromagnetic mass. Thus, only a small field reduction or fine tuning of 
the magnetic field is effected by the dipole rings 90 and 92. Each of the 
members 96 of the dipole rings 90 and 92 is made of cold-rolled or 
hot-rolled steel and is a rectangle having a width of 2 inches, a length 
of 1 inch and a thickness which may range from 1/32 of an inch to 1/4 of 
an inch. Although space limitations prevent solid dipole rings from being 
used in place of the spaced element dipole rings 90 and 92, it may be 
appreciated that in other embodiments solid dipole compensating may be 
employed instead of the spaced element rings. 
The combination of the symmetric compensating rings 60, the tesseral 
compensating bars 80 and the dipole rings 90 and 92 allows the two-fold 
symmetric shield to adequately reduce the magnetic flux outside the magnet 
while maintaining the field homogeneity within the examining region 18, as 
expressed in Legendre coefficients, to within 1 or 2 parts per million. 
This allows high resolution magnetic resonance imaging to be carried out. 
It may be appreciated that since the shield is added after the 
superconducting coil 12 is wound, it is necessary to provide the field 
inhomogeneity compensation within the examination region 18. The use of 
this system also allows the shield 14 to be left off the magnet 10 without 
having to rewind the superconducting coil 12. When the shield 14 is left 
off, the symmetric compensating rings 60, the compensating bars 80 and the 
dipole rings 90 and 92 are also be left out, thereby reducing 
significantly the cost of the resulting unshielded magnetic resonance 
imaging magnet. Furthermore, the two-fold symmetric shield is considerably 
less massive than other designs since ferromagnetic material does not 
cover approximately one-half of the angular area of the magnet surface. 
Although large portions of the coil are left open, providing a relatively 
low mass but effective shield in the horizontal plane, the perturbations 
which would normally be induced are adequately removed by the use the 
longitudinal tesseral compensating bars. 
While there has been illustrated and described a particular embodiment of 
the present invention, it will be appreciated that numerous changes and 
modifications will occur to those skilled in the art and it is intended in 
the appended claims to cover all those changes and modifications which 
fall within the true spirit and scope of the present invention.