The present invention provides an apparatus for reducing acoustic noise in a magnetic resonance imaging device including passive shielding located outside the actively shielded gradient winding elements in order to reduce the magnitude of fields that spread outside the gradient coil assembly in unwanted directions and interact with the magnet cryostat or other metallic magnet parts, inducing eddy currents that cause consequent acoustic noise. The passive shielding elements are conducting layers located on the outer radius of the cylindrical gradient coil assembly in a cylindrical magnet system, conducting layers located at the ends of the gradient coil assembly in a cylindrical magnet system, and conducting layers located inside the actively shielded gradient winding inner elements in a cylindrical magnet system. The passive shielding could also be located on separate structures that are vibrationally isolated from the magnet cryostat. The actively shielded gradient winding can also be extended to portions at the ends of the actively shielded gradient winding and further to portions inside the inner radius of the inner portion of the actively shielded gradient winding. The actively shielded gradient windings and passive shielding should be designed concurrently in order to substantially optimize the gradient linearity and reduce the eddy currents generated in metallic parts of the magnetic resonance imaging system.
William A Edelstein, Tesfaye K. Kidane, Victor Taracilla, Tanvir N. Baig Timothy P. Eagan Robert W. Brown
Classification: Measuring; Testing
This application claims the benefit of: U.S. Provisional Application No. 60/497,267, filed Aug. 25, 2003; U.S. Provisional Application No. 60/500,225, filed Sep. 5, 2003; and U.S. Provisional Application No. 60/570,593, filed May 13, 2004, which are hereby incorporated herein by reference.
BACKGROUND OF THE INVENTION
The present invention relates generally to a magnetic resonance imaging (MRI) scanner and more particularly to a low acoustic noise MRI scanner.
MRI scanners, which are used in various fields such as medical diagnostics, typically create images based on the operation of a magnet, a gradient coil assembly, and a radiofrequency coil(s). The magnet creates a uniform main magnetic field that makes unpaired nuclear spins, such as hydrogen atomic nuclei, responsive to radiofrequency excitation via the process of nuclear magnetic resonance (NMR). The gradient coil assembly imposes a series of pulsed, spatial-gradient magnetic fields upon the main magnetic field to give each point in the imaging volume a spatial identity corresponding to its unique set of magnetic fields during an imaging pulse sequence. The radiofrequency coil applies an excitation rf (radiofrequency) pulse that temporarily creates an oscillating transverse nuclear magnetization in the sample. This sample magnetization is then detected by the excitation rf coil or other rf coils. The resulting electrical signals are used by the computer to create magnetic resonance images. Typically, there is a radiofrequency coil and a gradient coil assembly within the magnet.
Magnets for MRI scanners include superconductive-coil magnets, resistive-coil magnets, and permanent magnets. Known superconductive magnet designs include cylindrical magnets and open magnets. Cylindrical magnets typically have an axially-directed static magnetic field. In MRI systems based on cylindrical magnets, the radiofrequency coil, the gradient coil assembly and the magnet are generally annularly-cylindrically shaped and are generally coaxially aligned, wherein the gradient coil assembly circumferentially surrounds the radiofrequency coil and wherein the magnet circumferentially surrounds the gradient coil assembly. Open magnets typically employ two spaced-apart magnetic assemblies (magnet poles) with the imaging subject inserted into the space between the assemblies. This scanner geometry allows access by medical personnel for surgery or other medical procedures during MRI imaging. The open space also helps the patient overcome feelings of claustrophobia that may be experienced in a cylindrical magnet design.
A gradient coil assembly comprises a set of windings in a support structure that produce the desired gradient fields. Such an assembly for a human-size whole-body MRI scanner typically weighs about 1000 kg. The windings consist of wires or conductors formed by cutting or etching sheets of conducting material (e.g. copper) to form current paths to generate desired field patterns. The wires or conducting coils or plates are themselves typically held in place by fiberglass overwindings plus epoxy resin.
Generally, the various components of the MRI scanner represent sources and pathways of acoustic noise that can be objectionable to the patient being imaged and to the operator of the scanner. For example, the gradient coil assembly generates loud acoustic noises, which many medical patients find objectionable. The acoustic noises occur in the imaging region of the scanner as well as outside of the scanner. Known passive noise control techniques include locating the gradient coil assembly in a vacuum enclosure.
Large pulsed electrical currents, typically 200 A or more, with risetimes and durations typically in the submillisecond to millisecond range, are applied to the windings. Because these windings are located in strong static magnetic fields (e.g., 1.5 T for a typical clinical imager to much higher values for research systems), the currents interact with the static field and strong Lorentz forces are exerted on different parts of the gradient coil assembly. These forces in turn compress, expand, bend or otherwise distort the gradient coil assembly. It will be readily understood by those skilled in the art that the frequencies of the acoustic noise so generated will be in the audio range. Typically there are strong components of noise from 50 Hz and below to several kHz at the upper end of the frequency range.
The magnet cryostat and other metallic parts of the magnet assembly are also sources of vibration and acoustic noise. For a cylindrical magnet, the actively shielded gradient winding generally comprises three layers, one for each Cartesian direction (x, y and z), each layer typically consisting of an inner, primary winding portion that creates a gradient field in the imaging region plus an outer, concentric shielding winding portion that substantially reduces the fields outside the gradient assembly. Some of the gradient-produced fields leak through or around the shielding. These leakage fields can induce eddy currents in the magnet metallic parts, for example, the cryostat inner bore. These eddy currents in turn produce Lorentz forces on the cryostat inner bore leading to mechanical motion of the cryostat inner bore and consequent acoustic noise according to WA Edelstein et al., Magnetic Resonance Imaging 20, 155–163, 2002.
The shielded gradients described above are constructed with inner gradient windings to produce gradient fields Gx, Gy and Gz of the main Bo field along axes x, y and z. They also have larger diameter shielding windings designed to substantially cancel external fields produce by the inner gradient windings so that the net fields outside the shielding windings are substantially zero. There are generally three separate inner windings, one for each of the directions x, y and z, and three separate outer, shielding windings, one for each of the directions x, y and z. The complete x-gradient winding consists of various parts, including inner and shielding windings, which are generally electrically connected to form a single electrical circuit. The configurations of y- and z-gradient windings are similar. Up to the present, the general practice has been to design the inner and shielding windings so that each axis has a single layer of inner windings and a single layer of shielding windings. Typically, the inner x-winding is electrically connected in series to the outer, shielding x-winding and the inner plus outer x-winding circuit is then powered by a single power supply. The same is true of the y- or z-winding.
FIG. 1A is a cross-sectional side view of a conventional actively shielded gradient. 304 is the metallic inner bore of the magnet cryostat. 20 is the inner gradient windings and 30 is the outer, shielding gradient windings. 20 and 30 are supported in 102, a solid, nonconducting annular support structure.
However, the active gradient shielding is not perfect, as some pulsed magnetic fields leak through and around the shielding windings, interact with the magnet or cryostat structure and cause eddy currents in those structures. As shown in Schenck et al., U.S. Pat. No. 4,617,516, 1986, the z-gradient windings are generally made using a wire wound in one direction from one end of the gradient to the center, and the wire reverses from the center outward. Substantial field can leak through the widely spaced turns in the center. For all windings, field can leak around the ends of the windings to interact with the cryostat or other metallic parts of the magnet structure. It has been shown that eddy currents in the metallic inner bore of the magnet cryostat produced Lorentz forces on the inner bore leading to vibrations of the inner bore and consequent acoustic noise (WA Edelstein et al., Magnetic Resonance Imaging 20, 155–163, 2002.)
Some recent designs to improve gradient active shielding produce increasingly complex gradient current patterns. The length of these windings must be limited to ensure efficient gradient operation. Compromises thereby incurred limit the effectiveness of the additional shielding achieved. The recent quest for shorter length and rapidly changing gradients has also tended to work against effective shielding.
The active gradient shields in cylindrical magnet MRI systems heretofore has been confined to a cylindrical surface disposed concentrically relative to the magnet. This arrangement is not the best configuration to produce the most effective shielding.
Passive shielding alone will not provide adequate shielding and eddy current control because of its finite time constants. Multiple layers of active gradient shielding will have limited effectiveness because active gradient shielding must have discrete current paths. We are proposing the use of passive gradient shielding in conjunction with active gradient shielding. One example of this approach is disclosed in Mulder et al., U.S. Pat. No. 6,326,788, 2001. FIG. 3 is a cross-sectional side view of their design. It is similar to the conventional actively shielded gradient shown in FIG. 1 with the addition of a passive shield 210 consisting of a conducting layer on the outer radius of the solid, nonconducting annular cylindrical support structure 102. The idea of passive shielding acting in combination with active shielding can be extended further with substantially improved efficacy.
What is needed is a method of drastically reducing gradient-induced eddy currents in the magnet cryostat and other magnet parts, and to mitigate consequent vibrations, in order to substantially alleviate one of the principal sources of acoustic noise in MRI scanners.
SUMMARY OF THE INVENTION
The present invention provides a method and apparatus for generating magnetic resonance images with reduced acoustic noise. In one embodiment, such an apparatus includes a gradient coil assembly passive shielding element including at least one conducting element outside the actively shielded gradient coil winding. The passive shielding element is located outside the gradient winding elements and is positioned to substantially block leakage fields that can otherwise produce eddy currents in the magnet cryostat and other magnet structures.
A first aspect of the invention is directed toward a method for reducing acoustic noise in a magnetic resonance imaging device, the method comprising the steps of providing a passive shielding element adjacent an actively shielded gradient coil winding of the device and ensuring an active gradient shielding winding is insulated from the passive shielding element, wherein a first portion of the passive shielding element is located adjacent an outer surface of the actively shielded gradient coil winding and a second portion is located adjacent an end of the actively shielded gradient coil winding.
A second aspect of the invention is directed toward apparatus for reducing acoustic noise in a magnetic resonance imaging device, the apparatus comprising an actively shielded gradient coil winding, an annular support structure substantially enclosing the actively shielded gradient coil winding, and a passive shielding element adjacent the actively shielded gradient coil winding, wherein a first portion of the passive shielding element is adjacent an outer surface of the actively shielded gradient coil winding and a second portion is adjacent an end of the actively shielded gradient coil winding.
A third aspect of the invention is directed toward a passive shielding element comprising a tubular member and an annular disk member residing adjacent an end of the tubular member, wherein the element is sized for application to an annular support structurefor the actively shielded gradient coil winding of a magnetic resonance imaging device.
A fourth aspect of the invention is directed toward apparatus for reducing acoustic noise in a magnetic resonance imaging device, the apparatus comprising an actively shielded gradient coil winding and an annular support structure substantially enclosing the actively shielded gradient coil winding, wherein the actively shielded gradient coil winding includes a first gradient winding portion, a second gradient winding portion having a radius greater than the first gradient winding portion as measured from the center axis of the device, and a third gradient winding residing beyond an end of and disposed at an angle relative to the first and second gradient windings.
The foregoing and other features of the invention will be apparent from the following more particular description of embodiments of the invention.
BRIEF DESCRIPTION OF THE DRAWINGS
The features and advantages of the present invention will become apparent from the following detailed description of the invention when read with the accompanying drawings in which:
FIG. 1 is a cross-sectional side view of a prior art actively shielded gradient assembly without passive shielding;
FIG. 2 is a cross-sectional side view of an entire magnetic resonance imaging system showing magnet and a prior art gradient coil assembly;
FIG. 3 is a cross-sectional side view of a prior art actively shielded gradient assembly with prior art passive shielding;
FIG. 4 is a cross-sectional side view of an actively shielded gradient assembly with passive shielding according to the invention;
FIG. 5 is a cross-sectional side view of an actively shielded gradient assembly with etended active shielding according to the invention;
FIG. 6 is a cross-sectional side view of an actively shielded gradient assembly with passive shielding according to the invention and extended active shielding according to the invention.
DETAILED DESCRIPTION OF INVENTION
Referring to FIG. 2 there is shown an illustrative MRI device 90 to which embodiments of the present invention are applicable. MRI device 90 is of a type useful in producing magnetic resonance (MR) images of a patient or subject. Throughout the figures, like numerals represent like elements. FIGS. 1–6 show MRI device 90 based on a closed, cylindrical superconducting magnet assembly 200. It is to be appreciated by one skilled in the art that the functions and descriptions of the present invention are equally applicable to an open magnet configuration.
FIG. 1 is a cross-sectional side view of a prior art actively shielded gradient assembly. The actively shielded gradient coil winding consists of an inner active gradient winding portion 20 and an outer gradient winding portion 30. Generally these are electrically connected in series and powered by a single power amplifier. They could be powered by two separate power amplifiers. The active gradient coil winding 20 plus 30 is embedded in an insulating support matrix 102, typically comprised of fiberglass-epoxy or other filler materials.
Referring to FIG. 2, this type of cylindrical magnet assembly, with center axis 250, comprises an inner surface referred to as a magnet cryostat inner bore 304 and a cryostat shell 100 disposed radially around the outer surface. The magnet assembly further comprises end cap seals 212. When end cap seals 212 are secured against rubber gaskets 220 positioned between end cap seals 212 and cryostat shell 100, and secured against other rubber gaskets 220 positioned between end cap seals 212 and patient tube 104, an airtight region 106 containing the gradient coil assembly 190 is created.
Typically, cryostat shell 100 encloses a superconductive magnet (not shown) that, as is well-known, includes several radially-aligned and longitudinally spaced-apart superconductive coils, each capable of carrying a large electric current. The superconductive coils produce a homogeneous, main static magnetic field, known as B0, typically in the range from 0.5 T to 8 T, aligned along the center axis 250. Cryostat shell 100 is generally metallic, typically steel or stainless steel.
A patient or imaging subject (not shown) is positioned within a cylindrical imaging volume 201 surrounded by patient bore tube 104. Bore tube 304 is typically made of electrically conducting material such as stainless steel. Gradient coil assembly 102 is disposed around in a spaced apart coaxial relationship therewith and generates time-dependent gradient magnetic field pulses in a known manner. Radially disposed around gradient coil assembly 102 is cryostat shell 100 including warm bore 304. Cryostat shell 100 contains the magnet that produces the static magnetic field necessary for producing MRI images, as described above.
Also shown in FIG. 2 is a schematic view of a vibration isolation suspension arrangement consisting of bracket 112 solidly attached to the magnet cryostat 200, bracket 108 solidly attached to gradient assembly 190. Elastomeric layer 110 supports the weight of gradient assembly 190 and reduces the amplitude of vibrations transmitted from gradient assembly 190 to the magnet cryostat 200.
Leakage pulsed magnetic fields from the gradient assembly 190 create eddy currents in the cryostat inner bore 304. These eddy currents subject the cryostat bore 304 to Lorentz forces which produce vibrations and consequent MRI system acoustic noise as shown by WA Edelstein et al., Magnetic Resonance Imaging 20, 155–163, 2002.
FIG. 3 illustrates an approach by Mulder et al., U.S. Pat. No. 6,326,788, 2001 to apply a passive shield in a position at a radius beyond the active gradient winding 20 plus 30. FIG. 3 is a cross-sectional side view of their design. It is similar to the conventional actively shielded gradient shown in FIG. 1 with the addition of a passive shield 210 consisting of a conducting layer positioned outside the active gradient winding 20 plus 30.
FIG. 4 shows a gradient assembly 190 with extended passive electromagnetic shield comprised of conductive portions 210, 220 and 230. Portions 210, 220 and 230 should be a good electrical conductor, for example, copper. For the purpose of clarity, gaps are shown between portions 210 and 220 and gaps between conductive portions 220 and 230. However, in practice, 210 and 220 will generally be continuously electrically connected, as will 220 and 230, so the entire passive shield would consist of a single conducting layer consisting of portions 210, 220 and 230. Fields that escape from the active shielding windings 20 plus 30 would cause eddy currents in the passive shield 210 plus 220 plus 230. These eddy currents would help contain fields inside the passive shield and prevent those fields from interacting with the cryostat bore 304 or other metallic magnet system parts. It may be advantageous in some situations, however, to have gaps as shown here or in other positions to better optimize the operation of the gradient assembly 190.
The additional conductive portions improve shielding efficacy. This is demonstrated by WA Edelstein et al., Proc. Intl. Soc. Mag. Reson. Med. 11, 747 (2004), where a comparison is made between the efficacy of a passive shield such as portion 210 only and a passive shield which consists of portions 210 and 220. For a 1 mm passive shield and a series of repeated trapezoidal pulses, and relative to a configuration with no passive shield, energy deposited in the cryostat bore 304 is reduced 14.7 dB with a passive shield consisting of portion 210 only and by 19.9 dB with a passive shield consisting of portions 210 and 220.
If a passive shield 210, 220 and 230 is positioned outside the gradient assembly, then any forces on this final shield may cause motion of the shield. However, if the passive shield is firmly mounted on the gradient assembly, its motion will be limited as compared to the motion of a thin cylinder such as the cryostat bore. Vibrational isolation of the gradient assembly 90 will prevent vibrations from the gradient assembly or passive shield from being conveyed to the rest of the MRI system and then causing acoustic noise. Finally, if a vacuum is created around the gradient assembly in airtight region 106, then motion of the gradient assembly or passive shield cannot cause acoustic noise is conveyed through air.
The presence of passive shielding elements 210, 220 and 230 will alter the field distributions when active gradient winding 20 plus 30 is pulsed. Therefore it is important to design the active gradient winding 20 plus 30, and the passive shielding elements 210, 220 and 230 concurrently, in order to optimize the disposition of pulsed gradient fields in imaging region 201 and minimize the eddy currents created in cryostat inner bore 304 and other metallic magnet system parts.
The active shielding can also be extended to enhance its function as shown in FIGS. 5 and 6. FIG. 5 shows actively shielded gradient coil winding portion 30 extended to end portions 32 and, if advantageous, extended further to cylindrical portions 34. For the purpose of clarity, gaps are shown between portions 30 and 32 and gaps between conductive portions 32 and 34. However, in practice, 30 and 32 will generally be continuously electrically connected, as will 32 and 34, so the actively shielded gradient coil winding portion for each axis comprises a single conducting layer consisting of: portions 30 and 32; or portions 30 and 32 and 34. It may be advantageous in some situations, however, to have gaps as shown here or in other positions, so that different active gradient portions have separate power supplies, to better optimize the operation of the gradient assembly 190.
FIG. 6 shows a combination of extended actively shielded gradient coil winding and extended passive shield.
Acoustic noise will also be reduced if the passive shield is mounted on a separate structure, also vibrationally isolated and contained in a vacuum, outside the gradient assembly but separate from the cryostat/magnet structure. Reduction of eddy currents in the cryostat inner bore effected by the passive shield will reduce bore vibrations and consequent acoustic noise.
There may be additional undesirable eddy currents generated by the active outer shield or passive shield that lead to undesirable fields in the imaging region. For example, misalignment of the inner and outer gradient windings tends to produce extra, pulsed, uniform BO magnetic fields (as opposed to gradient fields) synchronized with the pulsed gradient fields. This already happens in typical MRI systems because of interaction of misaligned gradient windings with the cryostat inner bore. It may therefore be desirable to include in this gradient package additional windings (uniform field, second order, third order, fourth order, etc.) that can be pulsed in order to cancel unwanted fields. These extra windings could be mounted on the inner actively shielded gradient winding portion 20, the outer actively shielded gradient winding portion 30, or anywhere at a radius less than that of the passive shielding 210 but at a radius greater than that of the gradient assembly inner bore. If necessary, a primary B0 (uniform field) winding, for example, could have an associated shielding B0 winding on the inner or outer windings. Higher-order windings and corresponding shielding windings could also be installed.
The initial magnitude of the eddy currents and consequent fields in the imaging region 201 will not change dramatically when supplementary passive shields 210, 220 and 230 are introduced. This is because the passive supplementary shields are at only a slightly smaller radius than the cryostat inner bore 304. However, using a copper passive shield, for example, will substantially increase the time constants of the eddy currents as compared to the time constants of eddy currents on the stainless steel bore. This is advantageous because longer time constants, and consequent slower variations of interfering fields, are easier to compensate than rapidly changing interfering fields.
An additional benefit of reducing leakage of pulsed magnetic fields is that eddy currents within the magnet structure will be reduced, thereby lessening power dissipation and undesirable heating of the magnet elements. Typically, the magnet elements are kept at very low temperatures by means of refrigeration or cryogenic fluids. Additional heating undesirably reduces efficiency of the magnet and increases operating costs.
While this invention has been described in conjunction with the specific embodiments outlined above, it is evident that many alternatives, modifications and variations will be apparent to those skilled in the art. Accordingly, the embodiments of the invention as set forth above are intended to be illustrative, not limiting. Various changes may be made without departing from the spirit and scope of the invention as defined in the following claims.