Electricity: magnetically operated switches – magnets – and electr – Magnets and electromagnets – Superconductive type
Reexamination Certificate
2000-11-20
2003-05-27
Barrera, Ramon M. (Department: 2832)
Electricity: magnetically operated switches, magnets, and electr
Magnets and electromagnets
Superconductive type
C335S299000, C335S301000, C324S319000, C324S320000
Reexamination Certificate
active
06570475
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates in general to split type or open superconducting magnets for magnetic resonance imaging (MRI), and in particular to such magnets having large diameter shield coils remotely spaced from the magnet poles.
2. Description of Prior Developments
Split type or “open” superconducting magnets are used in magnetic resonance imaging (MRI) scanners to produce magnetic fields required for patient imaging. Superconducting shield coils are typically used in each half of the split magnets to reduce stray electromagnetic fields. This type of shielding is referred to as active shielding.
A typical design of an actively shielded superconducting split type open MRI magnet assembly includes two generally cylindrical enclosures. A lower cylindrical enclosure and an upper cylindrical enclosure are interconnected by structural supports. Cryogenic and electrical connections are also provided between these enclosures or “halves”.
The opposing cylindrical enclosures form two magnetic poles separated by a gap which contains an imaging region where a patient is imaged. Each enclosure contains several superconducting coils placed inside a liquid helium vessel. The helium vessel is located within a vacuum vessel and maintains the magnet at an operating temperature of about 4 K. The coils contained in the helium vessel are arranged in location as well as in magnetic polarity via the direction of the current they carry, so that they produce a substantially uniform field in a portion of the gap formed between the two enclosures while limiting the stray field outside the device to an acceptable level.
Each enclosure also includes one or more thermal shields as well as thermal insulation, located between the helium vessel and the vacuum vessel to keep heat leak to the 4 K environment within acceptable levels. The uniform internal field is produced by a main superconducting coil and several field shaping coils. In addition, one or more shield coils, are spaced away from the gap to reduce stray field released outside of the MRI scanner.
Sizing and positioning of the coils is done by those skilled in the art by using numerical codes based on static magnetic field equations, with the objective to achieve both the targeted field uniformity inside the imaging region, e.g. through minimizing terms in the Legendre polynomial series, and restricted stray field, e.g. through minimizing external field moments. Although the location and size of the coils cannot be defined arbitrarily, there is some freedom in positioning the coils, and different positions result in different amounts of conductor required for given uniformity and stray field requirements.
A significant advantage of a split type MRI magnet and scanner is the openness of the gap which is formed between the magnet poles. The open gap provides an enhanced view of the patient in the imaging region and allows medical personnel to directly access the patient as the patient is positioned within the gap.
Efforts are being made to increase the gap size to improve patient comfort, visibility and accessibility and to reduce the diameter of the magnet poles to further increase patient comfort, visibility and accessibility. Different magnet designs with different gap configurations can be produced within different envelopes yet provide the same field strength and field uniformity.
Such designs may require different aggregate amounts of superconducting material or “conductor” contained in the various coil designs. The conductor is usually the largest single cost item in an MRI magnet assembly, so it is desirable to minimize the amount of conductor required to produce a given field strength and uniformity.
Moreover, the designs with a larger conductor volume usually result in a higher peak field, B
peak
, and greater accompanying mechanical stress in the coils. These two parameters gradually increase with conductor volume until the risk of design failure becomes unacceptable. Peak field and stresses have a major impact on the progress of future open MRI designs with a target of high field exceeding 1T.
Accordingly, each split-type magnet design represents a compromise between its openness, which is related to gap size and pole diameter, its cost and its operating structural safety margins. Generally, for a given field strength, uniformity and stray field, the larger the maximum diameter of the magnet coils and/or the smaller the patient imaging gap, the lower is the amount of conductor required and the lower is the cost and structural risk of the magnet design. However, these potential cost reducing approaches can result in decreased patient comfort, visibility and accessibility.
Accordingly, a continuing need exists for a split type MRI magnet which provides improved openness through a large split and small diameter pole, yet which requires less superconducting material and which reduces the associated field and stress in a relatively lightweight superconducting magnet, while satisfying stringent uniformity and stray field requirements.
SUMMARY OF THE INVENTION
The present invention has been developed to fulfill the needs noted above and therefore has as an object the provision of an open or split type MRI magnet which provides improved openness or gap size, yet which reduces the total or aggregate amount of superconducting material required, and which also reduces the field and stress within the magnet coils under stringent requirements of uniformity and stray field.
A further object is to provide such a magnet with flattened or axially shortened main coils to reduce the amount of conductor material required and to reduce the amount of stress produced in the main coils.
Yet another object of the invention is to provide a split type MRI magnet with axially stepped field shaping coils to facilitate coil winding and to allow for the use of a single stepped coil support form.
Another object of the invention is to provide a split type MRI magnet with a series of field shaping coils located closely adjacent to the magnet poles and imaging gap.
Still a further object of the invention is to provide a split type MRI magnet having field shaping coils positioned radially outside of a recess formed in the magnet pole of a cryostat enclosure.
The above and additional objects are met by the present invention which is directed to an open split type superconducting magnet for an MRI scanner which uses superconducting shielding coils, with optional external iron shielding, to reduce stray magnetic fields. In accordance with the invention, the primary shield coil, i.e., that shield coil typically with the largest diameter having its current flowing opposite or negative with respect to the current in the main coil, can have a substantially greater axial distance from the imaging gap and advantageously has a substantially larger diameter than the main coil. The larger diameter of the shield coil does not affect the patient's perception of openness as it has no effect on the patient's field of view from the patient's bed.
The main purpose of the negative primary shield coil is to compensate the external stray field produced outside of the magnet by the positive main coil and other coils. By doing so, the primary shield coil inevitably reduces the inner field in the imaging region, so the other coils have to grow in order to compensate this negative change. Contribution of the shield coil to the external field is determined by its magnetic moment, which grows as the square of its diameter. The shield coil of a larger diameter produces similar stray field with less ampere-turns and a reduction of the central field cancellation, hence it requires less conductor in other coils and in the whole magnet. More distant axial and radial positioning of the shield coil from the imaging gap additionally results in a smaller negative contribution in the imaging region, which further reduces the amount of conductor required for other coils.
This increased diameter of the primary shield coil can be carried out
Blecher Leo
Kalafala A. Kamal
Lvovsky Yuri
Wilcox Robert E.
Barrera Ramon M.
Intermagnetics General Corp.
Katten Muchin Zavis & Rosenman
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