Patent Application
20140184222
The present invention relates generally to a magnetic resonance imaging (MRI) system and in particular to an apparatus for active high order shimming.
Magnetic resonance imaging (MRI) is a medical imaging modality that can create pictures of the inside of a human body without using x-rays or other ionizing radiation. MRI uses a powerful magnet to create a strong, uniform, static magnetic field (i.e., the “main magnetic field”). When a human body, or part of a human body, is placed in the main magnetic field, the nuclear spins that are associated with the hydrogen nuclei in tissue water become polarized. This means that the magnetic moments that are associated with these spins become preferentially aligned along the direction of the main magnetic field, resulting in a small net tissue magnetization along that axis (the “z axis,” by convention). An MRI system also comprises components called gradient coils that produce smaller amplitude, spatially varying magnetic fields when a current is applied to them. Typically, gradient coils are designed to produce a magnetic field component that is aligned along the z axis, and that varies linearly in amplitude with position along one of the x, y or z axes. The effect of a gradient coil is to create a small ramp on the magnetic field strength, and concomitantly on the resonant frequency of the nuclear spins, along a single axis. Three gradient coils with orthogonal axes are used to “spatially encode” the MR signal by creating a signature resonance frequency at each location in the body. Radio frequency (RF) coils are used to create pulses of RF energy at or near the resonance frequency of the hydrogen nuclei. The RF coils are used to add energy to the nuclear spin system in a controlled fashion. As the nuclear spins then relax back to their rest energy state, they give up energy in the form of an RF signal. This signal is detected by the MRI system and is transformed into an image using a computer and known reconstruction algorithms.
The gradient coil assembly used in an MRI system may be a shielded gradient coil assembly that consists of inner and outer gradient coil assemblies bonded together with a material such as epoxy resin. Typically, the inner gradient coil assembly includes inner (or main) coils of X-, Y-, and Z-gradient coil pairs or sets and the outer gradient coil assembly includes the respective outer (or shielding) coils of the X-, Y- and Z-gradient coil pairs or sets. The Z-gradient coils are typically cylindrical with a conductor spirally would around the cylindrical surface. The transverse X- and Y-gradient coils are commonly formed from a copper panel with an insulating backing layer. A conductor turn pattern (e.g., a fingerprint pattern) may be cut in the copper layer of the gradient coil.
MRI systems require a uniform main magnetic field, B0, in the imaging volume, however, inhomogeneities in the magnetic field may be introduced by various factors such as manufacturing tolerances, environmental effects, design restrictions, imperfections in the magnet, ferromagnetic material near the installation site, and so forth. Inhomogeneities in the magnetic field, B0, can adversely affect data acquisition and reconstruction of an MR image. For example, magnetic field inhomogeneities may distort position information in the scan volume and degrade the image quality. A process known as “shimming” may be used to compensate for or remove inhomogeneities from the magnetic field, B0. An MRI magnet may be shimmed using shim or correction coils (active shimming) or passive shims such as pieces of ferromagnetic materials (passive shimming).
Active shimming uses dedicated coils in the magnet to generate a corrective magnetic field. Typically, a current is passed through the shim coils to create the corrective magnetic fields. The current through the shim coils may be adjusted or regulated to provide the appropriate corrective field. Shim coils may be resistive, superconducting or a combination of both. Superconducting shim coils are located inside the magnet and operate in a helium environment. Superconducting shim coils are used to compensate the inhomogeneities (harmonics) caused either by manufacturing tolerances or by the magnetic environment of the scanning room. Typically, the current in the superconducting shim coils is adjusted to a proper value(s) during installation or maintenance of the MRI scanner. Once the current is adjusted to the proper value(s), the current values are fixed and the superconducting coils operate in a persistent mode. To provide static compensation of patient-induced harmonics, which may vary from scan to scan, resistive shim coils (so-called high order shim coils) may be used. The resistive shim coils are often incorporated in the gradient assembly of an MRI scanner and typically include a second order set of shim coils for which the current may be adjusted between scans.
Currently, many MRI systems utilize a wide patient bore which leaves less radial space in the gradient coil system to accommodate high order shim coils. It would be desirable to provide a shim coil design that can be utilized in a small radial space and also allow for second order and higher terms.
In accordance with an embodiment, a matrix shim coil apparatus includes a plurality of coils configured to provide high order shimming, each coil having a figure eight shape and having a first loop and a second loop.
In accordance with another embodiment, a matric shim coil apparatus includes a plurality of coils configured to provide high order shimming, each coil having a figure eight shape, and wherein each coil is further folded along an axis of the coil creating a folded coil with an upper portion and a lower portion.
The invention will become more fully understood from the following detailed description, taken in conjunction with the accompanying drawings, wherein like reference numerals refer to like parts, in which:
The system control computer 32 includes a set of modules in communication with each other via electrical and/or data connections 32a. Data connections 32a may be direct wired links, or may be fiber optic connections or wireless communication links or the like, In alternative embodiments, the modules of computer system 20 and system control computer 32 may be implemented on the same computer system or a plurality of computer systems. The modules of system control computer 32 include a CPU module 36 and a pulse generator module 38 that connects to the operator console 12 through a communications link 40. The pulse generator module 38 may alternatively be integrated into the scanner equipment (e.g., resonance assembly 52). It is through link 40 that the system control computer 32 receives commands from the operator to indicate the scan sequence that is to be performed. The pulse generator module 38 operates the system components that play out (i.e., perform) the desired pulse sequence by sending instructions, commands and/or requests describing the timing, strength and shape of the RF pulses and pulse sequences to be produced and the timing and length of the data acquisition window. The pulse generator module 38 connects to a gradient amplifier system 42 and produces data called gradient waveforms that control the timing and shape of the gradient pulses that are to be used during the scan. The pulse generator module 38 may also receive patient data from a physiological acquisition controller 44 that receives signals from a number of different sensors connected to the patient, such as ECG signals from electrodes attached to the patient. The pulse generator module 38 connects to a scan room interface circuit 46 that receives signals from various sensors associated with the condition of the patient and the magnet system. It is also through the scan room interface circuit 46 that a patient positioning system 48 receives commands to move the patient table to the desired position for the scan.
The gradient waveforms produced by the pulse generator module 38 are applied to gradient amplifier system 42 which is comprised of Gx, Gy and Gz amplifiers. Each gradient amplifier excites a corresponding physical gradient coil in a gradient coil assembly generally designated 50 to produce the magnetic field gradient pulses used for spatially encoding acquired signals. The gradient coil assembly 50 forms part of a resonance assembly 52 that includes a polarizing superconducting magnet with superconducting main coils 54. Resonance assembly 52 may include a whole-body RF coil 56, surface or parallel imaging coils 76 or both. The coils 56, 76 of the RF coil assembly may be configured for both transmitting and receiving or for transmit-only or receive-only. A patient or imaging subject 70 may be positioned within a cylindrical patient imaging volume 72 of the resonance assembly 52. A transceiver module 58 in the system control computer 32 produces pulses that are amplified by an RF amplifier 60 and coupled to the RF coils 56, 76 by a transmit/receive switch 62. The resulting signals emitted by the excited nuclei in the patient may be sensed by the same RF coil 56 and coupled through the transmit/receive switch 62 to a preamplifier 64. Alternatively, the signals emitted by the excited nuclei may be sensed by separate receive coils such as parallel coils or surface coils 76. The amplified MR signals are demodulated, filtered and digitized in the receiver section of the transceiver 58. The transmit/receive switch 62 is controlled by a signal from the pulse generator module 38 to electrically connect the RF amplifier 60 to the RF coil 56 during the transmit mode and to connect the preamplifier 64 to the RF coil 56 during the receive mode. The transmit/receive switch 62 can also enable a separate RF coil (for example, a parallel or surface coil 76) to be used in either the transmit or receive mode.
The MR signals sensed by the RF coil 56 or parallel or surface coil 76 are digitized by the transceiver module 58 and transferred to a memory module 66 in the system control computer 32. Typically, frames of data corresponding to MR signals are stored temporarily in the memory module 66 until they are subsequently transformed to create images. An array processor 68 uses a known transformation method, most commonly a Fourier transform, to create images from the MR signals. These images are communicated through the link 34 to the computer system 20 where it is stored in memory. In response to commands received from the operator console 12, this image data may be archived in long-term storage or it may be further processed by the image processor 22 and conveyed to the operator console 12 and presented on display 16.
A patient or imaging subject 112 may be inserted into the resonance assembly 100 along a center axis 114 (e.g., a Z-axis) on a patient table or cradle 116. The patient table or cradle 116 is inserted into the resonance assembly at a “patient end” 142 of the resonance assembly and the opposing end of the cylindrical resonance assembly is a “service end” 144. Center axis 114 is aligned along the tube axis of the resonance assembly 100 parallel to the direction of a main magnetic field, B0, generated by the magnet 102. RF coil 106 may be used to apply a radio frequency pulse (or plurality of pulses) to a patient or subject 112 and may be used to receive MR information back from the subject 112 as is well known in the field of MR imaging. RF shield 105 is used to shield the RF coil 106 from external sources of RF radiation. RF shield 105 may be fabricated from any suitable conducting material, for example, sheet copper, circuit boards with conducting copper traces, copper mesh, stainless steel mesh, other conducting mesh, etc. Gradient coil assembly 104 generates time dependent gradient magnetic pulses that are used to spatially encode points in the imaging volume in a known manner.
Superconducting magnet 102 may include, for example, several radially aligned and longitudinally spaced apart superconductive coils 118, each capable of carrying a large current. The superconductive coils 118 are designed to create a magnetic field, B0, within the patient volume 108. The superconductive coils 118 are enclosed in a cryogen environment within a cryogenic envelope 122. The cryogenic environment is designed to maintain the temperature of the superconducting coils 118 below the appropriate critical temperature so that the superconducting coils 118 are in a superconducting state with zero resistance. Cryogenic envelope 122 may include, for example, a helium vessel (not shown) and thermal or cold shields (not shown) for containing and cooling magnet windings in a known manner. Superconducting magnet 102 is enclosed by a magnet vessel 120, e.g., a cryostat vessel. Magnet vessel 120 is configured to maintain a vacuum and to prevent heat from being transferred to the cryogenic envelope 122.
Gradient coil assembly 104 is a self-shielded gradient coil assembly. Gradient coil assembly 104 comprises a cylindrical inner gradient coil assembly or winding 124 and a cylindrical outer gradient coil assembly or winding 126 disposed in a concentric arrangement with respect to a common axis 114. Inner gradient coil assembly 124 includes inner (or main) X-, Y- and Z-gradient coils and outer gradient coil assembly 126 includes the respective outer (or shielding) X-, Y-, and Z-gradient coils. The coils of the gradient coil assembly 104 may be activated by passing an electric current through the coils to generate a gradient field in the patient volume 108 as required in MR imaging. A volume 138 or space between inner gradient coil assembly 124 and outer gradient coil assembly 126 may be filled with a bonding material, e.g., epoxy resin, visco-elastic resin, polyurethane, etc. Alternatively, an epoxy resin with filler material such as glass beads, silica and alumina may be used as the bonding material. It should be understood that magnet and gradient topologies other than the cylindrical assemblies described above with respect to
A high order matrix shim coil 130 is located at a first radius inside the magnet assembly 100. In
The matrix shim coil 130 includes a plurality of shim coils disposed on a cylindrical surface around the inner gradient coil assembly 124. The shim coils may be placed or mounted on a cylindrical surface with the appropriate dimensions to be placed inside the gradient coil assembly 104 (e.g., in the volume 138 between the inner gradient coil assembly 124 and the outer gradient coil assembly 126). For example, the coils may be etched on a circuit board or the coils may be fabricated from a continuous length on insulated copper wire wound into the desired pattern. In one embodiment, the circuit board may contain an FR4 backing and etched copper. In another embodiment, the circuit board may be a multi-layer Kapton circuit board with interleaving layers of etched copper and Kapton insulation. Preferably, the matrix shim coil 130 has dimensions that allow placement in small (e.g., <3 mm) radial spaces. The matrix shim coil includes n rows of coils along the z direction and m coils around the circumference of the cylindrical structure to form an n×m array of coils. In another embodiment, the matrix shim coil 130 may include an n×n array of coils.
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This written description uses examples to disclose the invention, including the best mode, and also to enable any person skilled in the art to make and use the invention. The patentable scope of the invention is defined by the claims, and may include other examples that occur to those skilled in the art. Such other examples are intended to be within the scope of the claims if they have structural elements that do not differ from the literal language of the claims, or if they include equivalent structural elements with insubstantial differences from the literal language of the claims. The order and sequence of any process or method steps may be varied or re-sequenced according to alternative embodiments.
Many other changes and modifications may be made to the present invention without departing from the spirit thereof. The scope of these and other changes will become apparent from the appended claims.
