Open MRI magnet with uniform magnetic field

An open magnetic resonance imaging (MRI) magnet having first and second spaced-apart superconductive coil assemblies each including a toroidal-shaped coil housing containing a superconductive main coil and a superconductive bucking coil. The bucking coil is spaced radially inward and radially apart from the main coil. The bucking coil carries an electric current equal in amperage, and opposite in direction, to the main coil. The bucking coils overcome the gross magnetic field distortions in the imaging volume of the main coils (created by the open space between the magnet's superconductive coil assemblies) to produce a magnetic field of high uniformity within the imaging volume.

BACKGROUND OF THE INVENTION 
The present invention relates generally to a superconductive magnet used to 
generate a high magnetic field as part of a magnetic resonance imaging 
(MRI) diagnostic system, and more particularly to such a magnet having an 
open design and a magnetic field of high uniformity. 
MRI systems employing superconductive or other type magnets are used in 
various fields such as medical diagnostics. Known superconductive magnets 
include liquid-helium cooled and cryocooler-cooled superconductive 
magnets. Typically, for a cryocooler-cooled magnet, the superconductive 
coil assembly includes a superconductive main coil surrounded by a thermal 
shield surrounded by a vacuum enclosure. A cryocooler coldhead is 
externally mounted to the vacuum enclosure, has its first stage in thermal 
contact with the thermal shield, and has its second stage in thermal 
contact with the superconductive main coil. Nb--Ti superconductive coils 
typically operate at a temperature of generally 4 Kelvin, and Nb--Sn 
superconductive coils typically operate at a temperature of generally 10 
Kelvin. 
Known superconductive magnet designs include closed magnets and open 
magnets. Closed magnets typically have a single, tubular-shaped 
superconductive coil assembly having a bore. The superconductive coil 
assembly includes several radially-aligned and longitudinally spaced-apart 
superconductive main coils each carrying a large, identical electric 
current in the same direction. The superconductive main coils are thus 
designed to create a magnetic field of high uniformity within a spherical 
imaging volume centered within the magnet's bore where the object to be 
imaged is placed. Although the magnet is so designed to have a highly 
uniform magnetic field within the imaging volume, manufacturing tolerances 
in the magnet and magnetic field disturbances caused by the environment at 
the field site of the magnet usually require that the magnet be corrected 
at the field site for such minor irregularities in the magnetic field. 
Typically, the magnet is shimmed at the field site by using pieces of 
iron, or, for Nb--Ti superconductive magnets cooled by liquid helium, by 
using numerous Nb--Ti superconductive correction coils. The correction 
coils are placed within the superconductive coil assembly radially near 
and radially inward of the main coils. Each correction coil carries a 
different, but low, electric current in any required direction including a 
direction opposite to the direction of the electric current carried in the 
main coils. Shielding coils may also be used within the superconductive 
coil assembly to prevent the high magnetic field created by and 
surrounding the main coils from adversely interacting with electronic 
equipment in the vicinity of the magnet. Such shielding coils carry 
electric current of generally equal amperage, but in an opposite 
direction, to the electric current carried in the main coils and are 
positioned radially outward of the main coils. 
Open magnets typically employ two spaced-apart superconductive coil 
assemblies with the space between the assemblies allowing for access by 
medical personnel for surgery or other medical procedures during MRI 
imaging. The patient may be positioned in that space or also in the bore 
of the toroidal-shaped coil assemblies. The open space helps the patient 
overcome any feelings of claustrophobia that may be experienced in a 
closed magnet design. Before Applicants' invention, known superconductive 
open magnets existed only in the literature. However, the literature is 
silent on how such magnets can be designed to have a magnetic field of 
high uniformity within the imaging volume when the creation of the open 
space between the superconductive coil assemblies grossly distorts the 
magnetic field creating a magnetic field of low uniformity within the 
imaging volume. Such magnetic field distortion is well beyond that which 
can be overcome by using known magnet shimming technology. 
What is needed is an open MRI magnet designed to have a highly uniform 
magnetic field within its imaging volume despite the gross magnetic field 
distortions created by the open space between the magnet's superconductive 
coil assemblies. 
SUMMARY OF THE INVENTION 
It is an object of the invention to provide an open superconductive MRI 
magnet designed to have a magnetic field of high uniformity within its 
imaging volume. 
The open MRI magnet of the invention includes a first superconductive coil 
assembly including a generally toroidal-shaped first coil housing, a 
generally annular-shaped first superconductive main coil, and a generally 
annular-shaped first superconductive bucking coil. The first coil housing 
surrounds a first bore and has a generally longitudinal first axis. The 
first main coil and the first bucking coil are each generally coaxially 
aligned with the first axis and disposed within the first coil housing 
with the first bucking coil spaced radially inward and radially apart from 
the first main coil. The first main coil carries a first main electric 
current in a first direction, and the first bucking coil carries a first 
bucking electric current in a direction opposite to the first direction. 
The open MRI magnet of the invention also includes a second 
superconductive coil assembly including a generally toroidal-shaped second 
coil housing, a generally annular-shaped second superconductive main coil, 
and a generally annular-shaped second superconductive bucking coil. The 
second coil housing is longitudinally spaced apart from the first coil 
housing, surrounds a second bore, and has a generally longitudinal second 
axis which is generally coaxially aligned with the first axis. The second 
main coil and the second bucking coil are each generally coaxially aligned 
with the second axis and disposed within the second coil housing with the 
second bucking coil spaced radially inward and radially apart from the 
second main coil. The second main coil carries a second main electric 
current in the first direction, and the second bucking coil carries a 
second bucking electric current in the opposite direction. 
In preferred embodiments, the first bucking coil is spaced radially apart 
from the first main coil a distance equal generally to at least twice the 
radial thickness of the first main coil, all the electric currents are 
generally equal in amperage, and the superconductive coil assemblies are 
generally mirror-image. 
Several benefits and advantages are derived from the invention. With 
Applicant's open MRI magnet design, bucking coils may be chosen by 
magnetic field analysis to overcome the gross magnetic field distortions 
within the imaging volume of the main coils (created by the open space 
between the magnet's superconductive coil assemblies) to produce a 
magnetic field of high uniformity within the imaging volume. Applicant's 
highly uniform magnetic field permits high quality MRI imaging. 
Applicant's open magnet design overcomes any claustrophobia feelings of 
patients. Applicant's design of an open magnet with a highly uniform 
magnetic field gives access to the patient by medical personnel for 
surgery or other medical procedures during high-quality MRI imaging.

DETAILED DESCRIPTION OF THE INVENTION 
Referring now to the drawings, wherein like numerals represent like 
elements throughout, FIGS. 1-2 show the open magnetic resonance imaging 
(MRI) magnet 10 of the present invention. The magnet 10 includes a first 
superconductive coil assembly 12 with a generally toroidal-shaped first 
coil housing 14 which surrounds a first bore 16 and which has a generally 
longitudinal first axis 18. The magnet 10 further includes a second 
superconductive coil assembly 20 with a generally toroidal-shaped second 
coil housing 22 which surrounds a second bore 24 and which has a generally 
longitudinal second axis 26. The second coil housing 22 is longitudinally 
spaced apart from the first coil housing 14 by structural posts 28, and 
the second axis 26 is generally coaxially aligned with the first axis 18. 
Preferably, the second superconductive coil assembly 20 is a generally 
mirror image of the first superconductive coil assembly 12 about a plane 
30 (seen on edge as a dashed line in FIG. 2) oriented perpendicular to the 
first axis 18 and disposed longitudinally midway between the first and 
second coil housings 14 and 22. 
The first coil housing 14 includes a first generally-circumferential 
outside surface 32 facing generally towards the first axis 18 and a second 
generally-circumferential outside surface 34 radially spaced apart from 
said first circumferential outside surface 32 and facing generally away 
from said first axis 18. The first coil housing 14 also includes a first 
generally-annular outside surface 36 facing generally towards the plane 30 
and a second generally-annular outside surface 38 longitudinally spaced 
apart from the first annular outside surface 36 and facing generally away 
from the plane 30. 
The first superconductive coil assembly 12 also includes a generally 
annular-shaped first superconductive main coil 40, preferably a generally 
annular-shaped first additional superconductive main coil 42, and a 
generally annular-shaped first superconductive bucking coil 44. The first 
superconductive coils 40, 42 and 44 are conventionally supported on coil 
forms (not shown in the figures). 
The first main coil 40 is generally coaxially aligned with the first axis 
18, is disposed within the first coil housing 14, and carries a first main 
electric current in a first direction. The first direction is defined to 
be either a clockwise or a counterclockwise circumferential direction 
about the first axis 18 with any slight longitudinal component of current 
direction being ignored. The first additional main coil 42 is generally 
coaxially aligned with the first axis 18, is disposed within the first 
coil housing 14, is longitudinally spaced apart from the first main coil 
40, and carries a first additional main electric current in the first 
direction (i.e., the direction of the electric current carried by the 
first main coil 40). Extra additional main coils may be needed to achieve 
a high magnetic field strength without exceeding the critical current 
density of the superconductor being used in the coils, as is known to 
those skilled in the art. 
The first bucking coil 44 is generally coaxially aligned with the first 
axis 18, is disposed within the first coil housing 14, and carries a first 
bucking electric current in a direction opposite to the first direction 
(i.e., in a direction opposite to the direction of the electric current 
carried by the first main coil 40). The first bucking coil 44 is spaced 
radially inward and radially apart from the first main coil 40. 
Preferably, the first main coil 40 is disposed radially towards the second 
circumferential outside surface 34, and the first bucking coil 44 is 
disposed radially towards the first circumferential outside surface 32. A 
coil is said to be disposed towards one of two spaced apart surfaces when 
it is disposed closer to the one surface than the other surface. Thus, the 
first main coil 40 is disposed radially closer to the second 
circumferential outside surface 34 than to the first circumferential 
outside surface 32. Likewise, the first bucking coil 44 is disposed 
radially closer to the first circumferential outside surface 32 than to 
the second circumferential outside surface 34. In an exemplary embodiment, 
the first additional main coil 42 is disposed radially towards the second 
circumferential outside surface 34 at a radial distance from the first 
axis 18 which is smaller than that of the first main coil 40. It is 
preferred that the longitudinal distance of the first main coil 40 from 
the plane 30 is generally equal to the longitudinal distance of the first 
bucking coil 44 from the plane 30. 
Preferably, the first main coil 40 is disposed longitudinally towards the 
first annular outside surface 36, and the first bucking coil 44 is 
disposed longitudinally towards the first annular outside surface 36. In 
an exemplary embodiment, the first additional main coil 42 is disposed 
longitudinally towards the second annular outside surface 38. 
It is preferred that the first main electric current of the first main coil 
40, the electric current of the first additional main coil 42, and the 
first bucking electric current of the first bucking coil 44 are generally 
equal in amperage. 
The first main coil 40 typically would be a superconductive wire or 
superconductive tape wound such that the first main coil 40 has a 
longitudinal extension and a radial extension (i.e., radial thickness) far 
greater than the corresponding dimensions of the superconductive wire or 
superconductive tape. Preferably, the first bucking coil 44 is spaced 
radially apart from the first main coil 40 a distance equal generally to 
at least twice the radial thickness of the first main coil 40. In an 
exemplary embodiment, the first main coil 40 has a longitudinal extension 
and is disposed such that the first main coil 40 totally longitudinally 
overlaps the first bucking coil 44. 
As previously mentioned and as shown in the figures, the second 
superconductive coil assembly 20 is a generally mirror image of the first 
superconductive coil assembly 12 about the plane 30. Therefore, in 
addition to the second coil housing 22, the second superconductive coil 
assembly 20 also includes a generally annular-shaped second 
superconductive main coil 46, preferably a generally annular-shaped second 
additional superconductive main coil 48, and a generally annular-shaped 
second superconductive bucking coil 50. The second superconductive coils 
46, 48 and 50 are conventionally supported on coil forms (not shown in the 
figures). 
The second main coil 46 is generally coaxially aligned with the second axis 
26, is disposed within the second coil housing 22, and carries a second 
main electric current in the first direction (i.e., in the same direction 
as the electric current in the first main coil 40). The second additional 
main coil 48 is generally coaxially aligned with the second axis 26, is 
disposed within the second coil housing 22, is longitudinally spaced apart 
from the second main coil 46, and carries a second additional main 
electric current in the first direction. Extra additional main coils may 
be needed by the second superconductive coil assembly 20 to balance any 
extra additional main coils of the first superconductive coil assembly 12, 
as can be appreciated by those skilled in the art. The open MRI magnet 10 
has its superconductive coil assemblies 12 and 20 include main coils 40 
and 46 either alone or in combination with one or more additional main 
coils 42 and 48 such that the magnet 10 has a field strength of generally 
at least 0.3 Tesla within an imaging volume having a volume size at least 
equal to that of a spherical imaging volume having a diameter of generally 
7 inches. 
The second bucking coil 50 is generally coaxially aligned with the second 
axis 26, is disposed within the second coil housing 22, and carries a 
second bucking electric current in a direction opposite to the first 
direction (i.e., in a direction opposite to the direction of the electric 
current carried by the first main coil 40). The second bucking coil 50 is 
spaced radially inward and radially apart from the second main coil 46. 
The magnet 10 has its first and second superconductive coil assemblies 12 
and 20 include bucking coils 44 and 50 alone or in combination with 
additional bucking coils (not shown) such that the design peak-to-peak 
magnetic field inhomogeneity within the imaging volume is less than 
generally 25 parts per million. 
Applicants constructed an exemplary embodiment of the open MRI magnet 10 of 
their invention with mirror-image first and second superconductive coil 
assemblies 12 and 20 using a conventional cryocooler (i.e., a 
Gifford-McMahon cryocooler) for cooling all of the superconductive coils 
to a temperature below their critical temperature to achieve and sustain 
superconductivity. Therefore, the first and second coil housings 14 and 22 
were made to be first and second vacuum enclosures, and a conventional 
first or second thermal shield 52 or 54 was interposed between the 
superconductive coils and the vacuum enclosure of the corresponding 
superconductive coil assembly 12 or 20 as shown in FIG. 2. The coil 
housings 14 and 22 and the hollow structural posts 28 provided a single 
vacuum structure. Conventional thermal insulators (not shown) were used to 
support the superconductive coils within the thermal shields and to 
support the thermal shields within the vacuum enclosures. The cryocooler 
coldhead 56 was attached to the second coil housing 22, and the coldhead's 
first stage 58 was thermally connected to the second thermal shield 54. 
The coldhead's second stage 60 was thermally connected to the second main 
coil 46, the second additional main coil 48, and the second bucking coil 
50 (such thermal connections being straightforward and not shown in the 
figures). The superconductive coils 46, 48 and 50 of the second 
superconductive coil assembly 20 were thermally connected to those 40, 42 
and 44 of the first superconductive coil assembly 12, and the second 
thermal shield 54 was thermally connected to the first thermal shield 52 
by appropriate thermal connectors within the hollow structural posts 28 
(such thermal connections being straightforward and not shown in the 
figures). The magnet 10 was supported on a conventional magnet floor mount 
62. 
The constructed magnet 10 (to be hereinafter described) was designed by 
Applicants to have a generally spherical imaging volume 64 (shown as a 
dotted circle in FIG. 2) centered generally at the intersection of the 
plane 30 and the first axis 18 with a magnetic field of generally 0.5 
Tesla, a design peak-to-peak magnetic field inhomogeneity of less than 3 
parts per million (ppm), and a diameter of generally 12 inches. Such 
design was made by Applicants using the principles of the present 
invention, previously disclosed herein, together with conventional 
magnetic field analysis, as is within the skill of the artisan. 
The open MRI magnet 10 constructed by Applicants had its first coil housing 
14 (i.e., first vacuum enclosure) longitudinally disposed generally 11 
inches from the plane 30, had an inner radius of generally 19 inches 
(which is the radius of the first bore 16), an outer radius of generally 
36.5 inches, and a longitudinal thickness of generally 18 inches. The 
first main coil 40 extended longitudinally generally 6 inches, extended 
radially generally 1.5 inches, was longitudinally disposed generally 12.6 
inches from the plane 30, was radially disposed generally 32.5 inches 
from the first axis 18, with the first main coil 40 having generally 
83,000 feet of generally 0.12-inch wide and generally 0.01-inch thick 
Nb--Sn superconductive tape at a temperature of generally 10 Kelvin with 
the first main electric current having an amperage of generally 65 
amperes. The first bucking coil 44 extended longitudinally generally 2 
inches, extended radially generally 1 inch, was longitudinally disposed 
generally 13 inches from the plane 30, was radially disposed generally 21 
inches from the first axis 18, with the first bucking coil 44 having 
generally 11,000 feet of generally 0.12-inch wide and generally 0.01-inch 
thick Nb--Sn superconductive tape at a temperature of generally 10 Kelvin 
with the first bucking electric current having an amperage of generally 65 
amperes. The first additional main coil 42 extended longitudinally 
generally 7 inches, extended radially generally 1.2 inches, was 
longitudinally disposed generally 19.2 inches from the plane 30, was 
radially disposed generally 31.8 inches from the first axis 18, with the 
first additional main coil 42 having generally 79,000 feet of generally 
0.12-inch wide and generally 0.01-inch thick Nb--Sn superconductive tape 
at a temperature of generally 10 Kelvin with said first main electric 
current having an amperage of generally 65 amperes. Such magnet was thus 
designed to have its generally spherical imaging volume 64 centered 
generally at the intersection of the plane 30 and the first axis 18 with a 
magnetic field of generally 0.5 Tesla, a design peak-to-peak magnetic 
field inhomogeneity of less than 3 parts per million (ppm), and a diameter 
of generally 12 inches. Because of manufacturing tolerances and magnetic 
field site disturbances, the measured inhomogeneity was 1000 ppm (which is 
considered to be a good unshimmed level) which Applicants brought down to 
10 ppm (which is considered to be an excellent level for high quality MRI 
imaging) using conventional iron shims, as is within the skill of the 
artisan. It is noted that without the bucking coils 44 and 50, the open 
magnet 10 would have had a design peak-to-peak magnetic field 
inhomogeneity of more than 1000 ppm which can not be reduced by shimming 
because a magnet can never be shimmed below its design level of magnetic 
field inhomogeneity! 
The magnet 10 of the invention provides an open MRI magnet design which 
increases patient comfort and physician access while maintaining high 
quality imaging through the use of bucking coils which shape the magnetic 
field of the main coils (and any additional main coils) to overcome the 
gross magnetic field distortions created by the magnet's open space 66. It 
was Applicants who first discovered that high strength magnetic fields in 
open MRI magnets could be made to be magnetic fields of high uniformity. 
The foregoing description of a preferred embodiment of the invention has 
been presented for purposes of illustration. It is not intended to be 
exhaustive or to limit the invention to the precise form disclosed, and 
obviously many modifications and variations are possible in light of the 
above teaching. For example, the superconductive coil assemblies 12 and 20 
of the magnet 10 of the invention are not limited to being 
cryocooler-cooled, and may be liquid-helium (or other cryogenic) cooled. 
It is intended that the scope of the invention be defined by the claims 
appended hereto.