WO2006088453A1 - Procede d’utilisation d’un petit scanner irm - Google Patents

Procede d’utilisation d’un petit scanner irm Download PDF

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Publication number
WO2006088453A1
WO2006088453A1 PCT/US2005/004920 US2005004920W WO2006088453A1 WO 2006088453 A1 WO2006088453 A1 WO 2006088453A1 US 2005004920 W US2005004920 W US 2005004920W WO 2006088453 A1 WO2006088453 A1 WO 2006088453A1
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imaging
shimming
magnet
gradient
optimal
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PCT/US2005/004920
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English (en)
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Zhao, Lei
Teklemariam, Grum
Lian, Jianyu
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Priority to PCT/US2005/004920 priority Critical patent/WO2006088453A1/fr
Publication of WO2006088453A1 publication Critical patent/WO2006088453A1/fr

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
    • G01R33/48—NMR imaging systems
    • G01R33/54—Signal processing systems, e.g. using pulse sequences ; Generation or control of pulse sequences; Operator console
    • G01R33/56—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution
    • G01R33/563—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution of moving material, e.g. flow contrast angiography
    • G01R33/56375—Intentional motion of the sample during MR, e.g. moving table imaging
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/3806—Open magnet assemblies for improved access to the sample, e.g. C-type or U-type magnets
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/383—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using permanent magnets
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G01R33/3873—Compensation of inhomogeneities using ferromagnetic bodies ; Passive shimming
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
    • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
    • G01R33/387—Compensation of inhomogeneities
    • G01R33/3875—Compensation of inhomogeneities using correction coil assemblies, e.g. active shimming

Definitions

  • MR magnetic resonance
  • the magnet systems for MRI scanners have to accommodate the insertion of a human being and generate a homogeneous region large enough to cover a cylindrical area with a diameter between about 20 to about 50 cm, preferably about 40 cm, spherical volume (DSV) over the subject.
  • the magnets are typically made from permanent magnets in low-field systems ( ⁇ 5,000 gauss; ⁇ 0.5T) and superconducting magnet systems in high field systems (>10,000 gauss; >1T).
  • ppm parts per million
  • the large imaging volumes also mean concomitant higher power consumptions by the gradient and rf systems as well. In superconducting magnet systems, this also translates into higher cryogenic fluid consumption which means higher maintenance and service costs.
  • the gradient and rf magnetic subsystems require higher power.
  • the gradients have to generate sufficiently strong gradient fields, with good linearity, that are switched fast over a large volume requiring higher voltages and currents. This makes the gradients and amplifier designs more complex and costly.
  • the rf coils have to provide sufficiently strong and homogeneous rf fields over large volumes. Once again the power required to achieve this scales strongly with volume and therefore the rf coils and amplifiers become more complex to design and costly as well.
  • the higher power consumption requires special housing with heat ventilation for these amplifiers.
  • MRI magnetic resonance imaging
  • the present invention resolves these problems.
  • One solution to cutting MRI system costs is to reduce its size without reducing the patient bore size or patient gap.
  • the bore length can be reduced.
  • the magnet diameter and volume are reduced.
  • the gradient coil and RF coil sizes will need to be reduced accordingly.
  • the RF and gradient power consumption is then significantly reduced, putting less demand on the amplifiers.
  • an MRI system size is reduced to half its original size, the cost can be reduced by more than 35% of the original cost.
  • reducing MRI system size has presented many problems which have caused image quality to deteriorate to such a great extent that small systems have not substantially penetrated the market.
  • a first consequence of size reduction is the shrinkage of the available imaging region. This means that most of the body may not be accessible for scanning in one session requiring multiple repositioning to scan all necessary regions and increasing scan time that otherwise would take less time in a conventional system.
  • the reduction of magnet size also means a reduction in the gradient and rf sizes. Consequently, the gradients will have a less linear region causing more distortions closer to the center of the magnet than conventional systems. Likewise, the rf coils will have a less homogeneous region than conventional systems. Therefore, the overall field-of-view (FOV) is drastically reduced.
  • a moving patient table is normally used to move the subject through the system FOV for scanning and thereby extend the available FOV.
  • the patient table pauses several times during a scan.
  • images are acquired within the limited FOV.
  • This move and pause action requires extra time and may shift the patient causing overlapping artifacts between different scanned images.
  • the table can move continuously and phase errors generated can be compensated; however, this approach has restrictions on table motion and encoding directions, and can not be generally applied.
  • the present invention is directed to surmount all the above and similar problems.
  • This invention is directed to methods of operating a whole body magnetic resonance imaging (MRI) systems that are much more cost-effective than current conventional MRI systems because of reducing the open magnet size and hence the size of RF and gradient coils.
  • the size reduction for these key components results in a significant reduction in cost for magnet materials and for power supplies and amplifiers. Even though cost is lowered with a small scanner size, the imaging quality is preserved by the techniques of this invention. Due to the reduction in scanner size, the present system is more open than all conventional systems (both cylindrical and vertical open) and therefore more suitable for interventional MRI.
  • the systems can either be permanent, electrical or superconducting based magnets
  • a first requirement of the present invention is to reduce the magnet size significantly while keeping the whole body access and the magnetic field strength substantially the same as in conventional whole-body systems.
  • the bore diameter can be kept the same while the bore length is made 20-30% shorter.
  • the gap size is substantially maintained while the magnet diameter and total magnet volume are reduced by about 30 to 60% of a conventional system. This directly translates into a significant cutting of magnet material cost.
  • the gradient and RF coils are also reduced similarly in physical size.
  • the smaller gradient or RF coils need much less power to achieve the same performance.
  • the gradient and RF power amplifiers can be reduced significantly. Since the price of amplifiers is strongly related to the power required, the amplifier costs are also significantly reduced.
  • the gradient and RF amplifier cost another very expensive part of a MRI system, is also reduced, lowering the total system cost considerably.
  • a problem with simply reducing magnet, gradient coil and RF coil sizes is that it results in a reduction of optimal imaging area, i.e., the useful imaging area with sufficient field homogeneity, gradient linearity, and RF field homogeneity.
  • Another problem is a decrease in field stability.
  • a small scanner, especially a low field permanent magnet system, is more susceptible to temperature drifts than a large conventional scanner. This invention provides methods of dealing with these problems while benefiting from a significant cost reduction.
  • a combination of passive shimming and high-order active shimming is utilized to improve magnetic field homogeneity in an open magnet to 10 ppm and less overall variation .
  • passive shimming more attention is focused on avoiding the introduction of high order field inhomogeneity.
  • multi-order active shims are used.
  • the active shims are based on an orthonormal basis set, and more particularly they are based on a spherical harmonic set including zeroth, first, second, and higher order shimming coils which have an excellent ability to provide a further lOOppm or more shimming capability. Since the shims are driven by currents, they are easily controlled.
  • By integrating the shim controls into the MRI console an image based shimming method is developed for the best overall shimming possible.
  • the active shimming is combined with a "Dynamic Shimming while Imaging” (DSWI) technique to dramatically improve imaging plane field homogeneity where active shimming is optimal and the shimmed area is small.
  • DSWI allows different imaging slices or slabs with different shimming coefficients within one scan to be shimmed individually.
  • a different set of shimming parameters are acquired using field mapping pulse sequences for different imaging slices or slabs. These pre-acquired shimming parameter sets are stored in the MRI scanner and the actual imaging pulse sequence is applied.
  • the corresponding optimal shimming parameters are loaded before the play out of any gradient or RF pulses.
  • the shimming parameter changes can be every TR (repetition time) or after multiple TRs, depending on encoding orders. Generally whenever slice position or orientation is changed, shimming coefficients are change accordingly. This procedure results in optimal field homogeneity for whole volume imaging.
  • an active thermal compensation system is used to stabilize magnet temperature and thus prevent field strength drifting. Due to the smaller amount of materials used in low field small permanent magnet systems as compared to conventional systems, the magnet temperature is more susceptible to ambient temperature changes which in turn affects the stability of the magnetic field strength. To overcome the temperature changes, thermal heating blocks with thermo-sensors are bound onto the magnet and its poles and an inverse feedback circuit is used to compensate for thermal changes at multiple locations. For better active thermal compensation the tolerances are made higher and more sensors and heaters used with faster response times.
  • the sweet spots are in the central part of the scanner.
  • gradient performance is better if the imaging plane (volume) is oriented so that it is always aligned with one of the inherent physical gradient directions (X, Y, Z) thereby avoiding large oblique planes.
  • the desired subject imaging volume can be larger than these regions, and their orientation can be arbitrary. The following methods developed in this invention solve this problem.
  • an automated patient table system with 6 degrees of freedom, is used so that the target subject imaging planes (volumes) are always manipulated to be align within the optimal regions. If a target imaging volume is larger than the optimal imaging region, the volume is divided into several slices (slabs), and each slice (slab) is aligned with the optimal imaging region.
  • the patient table is automatically controlled to rotate around the magnet vertical center, tilt along its long and short axes, move up and down, move in and out of the scanner, and move from left to right (Fig. 2). This requires properly securing the subject on the patient table by either frames or straps to prevent the patient from moving or slipping.
  • imaging planes (volumes) are within the above defined imaging regions, field homogeneity and gradient and RF linearity are all optimized.
  • Imaging planes are aligned with the orthogonal magnet coordinates, logical gradients are aligned with the physical gradients. This approach improves gradient efficiency (high maximal gradient and fast gradient switching time) and lowers the requirement for the gradient amplifier power. Imaging at central parts of the magnet also reduces eddy current induction.
  • a method is developed to obtain large subject volume images using a limited optimal imaging region. Imaging is always performed in the limited region, and objects move through the optimal region to obtain large coverage. This process is achieved by a "continuously no-pause moving table" and an adjustment of gradient fields to ensure the same gradient fields are seen by the subject during this motion. "Continuous movement” is used herein to mean a no-pause table movement, i.e.
  • the table continuously moves.
  • an imaging plane or volume
  • spins experience different gradients. This generates the wrong excitation, encoding, and acquisition information.
  • the zeroth shimming is adjusted during a scan to compensate for the motion of the table. In this situation, a moving spin which moves along with the patient table will always experience the same gradient field as if it were stationary.
  • the Dynamic Shimming while Imaging (DSWI) procedure may also be used in combination with the moving table as follows: before a scan, a patient table with a subject moves along the same path as the scan path. During the table movement, a few sample points are identified based on subject geometry. At a given sample point, the table stops and an optimal shimming coefficient set is determined using field mapping methods specific to the subject on the table at that sample point. Afterwards, all the shim coefficients are determined for all the sample points. The coefficients for each shimming channel are interpolated so that a continuous shimming coefficient set is obtained for different table (subject) positions. During image scanning, at the start of each TR, these sets of shimming coefficients are loaded so that the shim is always optimized at each sample point of the patient table as it moves through the imaging region.
  • An open whole body small scanner provides a large opening enabling the easy use of robotic controlled interventional devices for interventional MRI procedures.
  • the low aspect ratio and high degree access of the design allows implementation of virtually unencumbered interventional techniques.
  • reducing the radius of the magnet while effectively keeping the same patient gap allows closer access to the patient.
  • a typical scanner with a 40cm DSV FOV will have a patient access space of 40cm and magnet pole radius of 55cm.
  • reducing the DSV to 20cm will reduce the radius of the magnet pole by 10cm to 45cm. Consequently, physicians can gain 10cm more access to the patient.
  • its small fringe field allows for non-MRI compatible devices operated in the same scan room, enabling flexible clinical setup, especially for operating room setup. The replacement of MRI compatible devices with non-MRI compatible devices can significantly reduce clinical cost.
  • Fig. 1 is a simplified block diagram showing the elements of an MRI system of this invention.
  • Fig. 2 shows the patient table orientations with 6 degrees of freedom movement. Thick line arrows show the 6 pairs of patient table moving directions.
  • Fig. 3 shows the optimal imaging planes oriented along the physical X, Y, and Z axes for a vertical permanent magnet MRI system
  • Fig. 4 shows top views of the continuous table movement procedure during image scanning, beginning with (a) start of 1 st slab imaging; (b) end of 1 st slab imaging; and (c) end of whole volume imaging.
  • Fig. 1 illustrates a simplified block diagram of a MRI system for producing images in accordance with embodiments of the present invention.
  • the system could have an open, cylindrical shaped electrical, superconducting, or a permanent magnet.
  • An open, cylindrical permanent magnet is shown as an example in Fig. 1.
  • the static magnet field is provided by the magnet 101 as shown in Fig. 1.
  • the field is oriented vertically; however, this invention applies to any other types of magnets and field orientation as well.
  • Passive shimming blocks 102 mounted on the inner surfaces of the magnet, are used for improving the homogeneity of the magnetic field. These passive shimming blocks 102 are usually ferrite or high permeability materials.
  • active shimming coils 103 are also mounted within the gradient coils 104.
  • the shimming coils generate, but are not limited to, 0 th , 1 st , 2 nd or higher order spherical harmonic gradient fields to compensate for the static field non-homogeneity.
  • the current of the shimming coils are supplied by the active shimming amplifiers 110, which are controlled by the host computer 130 through the system control interface unit 120.
  • temperature sensors 141 are placed throughout the magnet 101. Based on the detected temperature changes, the temperature control unit 140 (including control circuits and amplifiers) will adjust the electrical current supplied to the heating plates 142 to stabilize the magnet temperature.
  • a frequency stabilization system can be used.
  • a small center frequency detection coil with a small phantom 109 can be placed in the magnet and be controlled by the center frequency detector 111.
  • the detected center frequency will be used as feedback through the system control interface unit 120 to modulate the 0 th order shimming amplifier 110, so that the static magnetic field is stable.
  • a subject 105 lies on the patient table 106.
  • a RF coil 107 is attached to the patient table (or mounted on the magnet after the gradient coil, typically referred to as a fixed RF body coil) that covers the subject's target imaging area 108.
  • the RF coil excites and receives signals from the subject. These can be two separate RF coils, one for transmission and one for reception.
  • a T/R switch 117 is needed to separate the transmitted and received signals.
  • the RF signals generated by the spectrometer 113 are amplified by the RF amplifier 115 and then fed to the T/R switch 117.
  • the received signal from the T/R switch 117 is sent to the spectrometer 113 through a pre-amplifier 116.
  • the spatial selection and encoding is achieved by the X, Y, Z gradient fields generated by the gradient coils 103 mounted on the upper and lower magnet poles inside the patient gap.
  • the gradient power is supplied by the gradient amplifiers 114, which are controlled by the spectrometer 113.
  • a pulse sequence is first designed based on the desired image contract.
  • a pulse sequence composes of groups of RF and gradient pulses, and acquisition actions ordered at a certain time sequence.
  • a pulse sequence is designed in the host computer 130 and downloaded to the spectrometer 113.
  • the spectrometer 113 generates gradient and RF pulses based the pulse sequence designed which are fed to the gradient amplifiers 114 and RF amplifiers 115, that then drive the gradient coils 103 and the RF coil 107 to generate the designed gradient and RF fields.
  • the spins in the subject will emit certain RF signals with their spatial encoding from the gradient field generated by the gradient coils 103.
  • the subject RF signals are first amplified by the pre-amplifier 116 after passing through a T/R switch 117.
  • the amplified signal is then acquired by the spectrometer 113 and stored in the spectrometer as raw data.
  • the raw data is finally sent to the host computer 130 via the system control interface unit 120.
  • image reconstruction primarily Fourier transform, is performed and subject images are displayed on the monitor 131.
  • An image pulse sequence normally requires a repetition of itself many times with the adjustment of a portion of the pulses.
  • the duration of each repetition is called a TR (repetition time).
  • TR repetition time
  • the total scan time is the multiplication of TR and the total number of the repeats.
  • the size of the magnet is significantly smaller, by up to 50% or more, than conventional magnets while preserving the magnetic field strength of a conventional magnet. More specifically, for a horizontal bore cylindrical magnet, the size reduction is in the length of the magnet. For an open, vertical field permanent magnet system, the size reduction is the magnet diameter and its total height. The diameter or length reduction can be as much as a quarter or higher. The total material volume reduced is even higher because the size varies quadratically with the radius.
  • Gradient coils 104 and fixed RF body coils in this embodiment are also reduced in size to fit into the small magnet. Because the power needed for gradient and RF coils is roughly proportional to the fourth and second power of the coil sizes, respectively, the required gradient and RF amplifiers in this embodiment have much smaller power requirements than conventional systems.
  • the patient openness as the ratio of magnet radius 150 over patient gap 151. Due to the reduction in magnet radius 150, the openness in a circular magnet is improved from aspect ratios of 1.3 or greater to 1.1 or less.
  • the magnet is preheated using high power (>1 kW) resistors or plates 142 attached to the two poles of the magnet until a magnet temperature well above ambient is reached (>32°C). Thereafter, smaller resistive heaters 143 (typically not greater than 50W) uniformly distributed along the magnet exterior surfaces are used to heat the magnet. These heaters are controlled by a feedback mechanism where temperature sensors also uniformly distributed over the magnet to monitor the temperature throughout and the necessary amount of differential heat is applied to maintain a stable temperature of the magnet above ambient. Therefore, heat flows from the magnet to the environment and not from the environment to the magnet. To ensure a stable magnetic field strength or center frequency, a frequency stabilization method is used, especially for a small permanent magnet scanner where center frequency drifting is a problem.
  • a small RF coil 109 with a small phantom is placed in the magnet.
  • the center frequency detector 111 a single channel spectrometer, is used to drive the RF coil and acquire magnetic resonance signals. The excitation and acquisition of these signals takes place when the gradients are off. These signals are Fourier transformed and the center frequency is calculated in real-time. The center frequency offset is converted to a DC current which is then fed to the zeroth order active shimming coil 103. The field generated by the zeroth order shimming coil 103 is used to compensate the static magnetic field drifting.
  • a combination of passive shimming using shimming blocks 102 and high-order active shimming 103 is used in an open magnet as shown in Fig. 1.
  • the magnetic field in a region covering the imaging region is mapped and passive shimming blocks 102 are added first at the outer edges of the magnet pole face.
  • the field is then remapped and based on the homogeneity changes either more elements are added or removed, and these steps are repeated several times until no further changes in the homogeneity are observed.
  • shim blocks are added at radii smaller than the previous one and the process of mapping the field, checking for homogeneity changes and adding or removing shim blocks at this smaller radius is repeated several times. When there are no further changes in the homogeneity, shim blocks get added at radii smaller than the previous ones.
  • active shimming can be performed.
  • the advantage of active shimming in this setting is that the shims are based on an orthonormal basis and so each orthogonal component can be shimmed without much interaction with other terms. This is a significant advantage over the passive shimming passes and produces more than lOOppm of shimming capability.
  • the active shimming is done on a spherical phantom using image data.
  • the phantom is filled with copper sulfate doped water.
  • a phase sensitive pulse sequence is used to obtain a 3D phase map of the phantom.
  • This phase map is subsequently converted to a magnetic field map.
  • a least-mean-square fitting is conducted on the field map data to obtain the coefficients of the shims. Thereafter, using a scaling relation between shim current strength and these coefficients that has been pre-determined, the correction is applied.
  • a Dynamic Shimming While Imaging (DSWI) method is applied.
  • a phase-sensitive pulse sequence is first applied and a 3D magnetic field strength map is obtained and stored in the host computer 130.
  • scan parameters including slice (slab) locations, are first chosen by an operator.
  • the active shimming coil 103 coefficients are calculated for each slice (slab) location of a multi-slice (-slab) scan (including one slice (slab) scan). Because the target shimming region for each slice (slab) is much smaller than a whole volume shimming approach, the field homogeneity is improved dramatically for each slice (slab).
  • the calculated shimming coil coefficients are stored in an array with the same slice (slab) scan order of the imaging scan.
  • shimming amplifiers 110 are adjusted first to insure the corresponding imaging slice (slab) has the optimal field homogeneity.
  • the adjustment of the shimming amplifiers 110 is achieved by retrieving the previously stored shimming coil coefficients triggered at the start of each TR.
  • the optimal regions 301 are defined as those regions where field homogeneity, gradient linearity, RF field linearity and gradient performance are optimal. These regions are aligned with the X, Y, and Z axes typically covering 10 or more centimeters thickness.
  • the optimal imaging regions are the three orthogonal slabs shown as the shaded regions 301 of Figure 3.
  • gradient performance maximal gradient amplitude and slew rate is optimal.
  • the patient is oriented to the oblique angle rather than generating oblique gradients.
  • image scanning is limited to the optimal imaging regions 301.
  • a flexible table design is used to ensure this requirement is met in an open magnet configuration as shown in Figure 2.
  • This patient table can be oriented into a wide range of positions and orientations.
  • the patient table is automatically controlled and has 6 degrees of movement, including moving left/right, up/down, in/out of the scanner, rotating around the scanner vertical axis 201 shown in Fig. 2, and tilt along the patient table's long and short axes.
  • a patient table moving path is calculated to ensure that imaging planes (volumes) pass through one of the optimal regions of the scanner.
  • Image scanning is then synchronized with the movement of the patient table, so that scanning is only applied to planes (volumes) passing through the optimal imaging regions.
  • frames or bands are used to safely secure the patient to the patient table.
  • the patient table 106 moves without any pause until the target subject imaging volume 404 has passed through the optimal imaging region 405, where image scanning is performed (see Fig. 4).
  • a magnetic field adjustment scheme is employed as explained below.
  • Figure 4 shows the scanning procedure to image a large target imaging volume 404 in a limited optimal imaging region 405.
  • the target imaging volume 404 is divided into N slabs (or slices).
  • the table position and orientation are adjusted by the table control unit 112 such that the imaging slabs are parallel to one of the optimal imaging regions 301.
  • the optimal imaging region is shown as 405.
  • the slab thickness 403 is smaller than half the width of the optimal imaging region 402.
  • the leading edge of the target imaging volume 404 also the leading edge of slab 1 is aligned with the center line of the optimal imaging region 401 (one of the central planar axes of the magnet) (see Fig. 4a).
  • the image scan starts.
  • the direction of table motion is preferably orthogonal to 401.
  • the speed is preferably constant and equal to 2*SLTH/T, where SLTH is the slab thickness and T is the total imaging time of a slab. With this speed, when a slab scan is finished, the trailing edge of the scanned slab aligns with the center of the optimal imaging region 401 (see Fig. 4b). The remaining slabs are imaged similarly with the same table velocity until the N" 1 slab is scanned.
  • the slab when the table is moving, the slab does not move through the same gradient fields. Therefore, yet in another invention, a method is developed to compensate for this problem.
  • the magnetic field strength is adjusted continuously by the zeroth order shimming coil 103. Consequently, if the combined gradient vector (physical X, Y and Z gradients) has a projection along the patient table moving direction, the magnetic field strength will be adjusted so that all moving spins experience the permanent static field plus a continuously varying zeroth order shimming field plus the gradient field with the net result being exactly the same as if the spins were stationary and they were experiencing just the permanent static field plus the gradient field.
  • the zeroth order shimming field changing rate at time t is set to be the negative product, -G(t) *V(t).
  • the minus sign reflects the fact that this is used to compensate for the field changes due to the motion of the patient table.
  • the zeroth order shimming coil current is reset back to its value before the start of this slab scan. Then, at the end of this first scan slab, the leading edge of the second slab is located at the central line of the optimal imaging region 401 as shown in Fig.4b. As the table continues to move, the scanning as in the first slab scan continues. If the slab sections have gaps, the table will move until the leading edge of the second slab aligns with 401 before the scanning starts. Using this method, there is no change in the pulse sequence itself, except that the patient table motion is synchronized with the slab scanning compensation scheme described above.
  • the zeroth order shimming values for the entire scan are pre-calculated based on the designed gradient projection along the direction of the patient table motion and the velocity of the table motion using -G(t)*V(t).
  • the values are stored in the system control interface unit 120 or the host computer 130. There are several methods to store and retrieve the zeroth order shim compensation values.
  • a group of continuous zeroth order shimming values are calculated and stored in the system control interface unit 120 or the host computer 130.
  • a trigger is generated by the spectrometer 113 programmed in the imaging pulse sequence. This trigger triggers the system control interface unit or the host computer to play out the corresponding group of zeroth order shimming values.
  • a large FOV as desired can be imaged as long the fixed RF body coil is utilized.
  • the frequency-stabilization method described earlier will use the pre-calculated adjusted magnetic field values as reference instead of keeping it constant.
  • the Dynamic Shimming While Imaging (DSWI) method is applied to scan in the optimal imaging region.
  • the field homogeneity can vary due to the susceptibility effects generated by the interaction of the different parts of the subject with the central part of the magnetic field.
  • shimming is performed dynamically during the table motion as follows. After the scan path is planned, the patient table will move along the same path as the imaging path. During the table movement, the table stops at a select predefined sample points. The sample points are picked according the geometric variation of the subject. If the subject exhibits minimal geometric variations, fewer sample points are needed. If the subject exhibits many geometric variations more sample points are preferred.
  • the number of sample points can be 4-5.
  • the field homogeneity of the optimal imaging region is acquired using field-mapping sequences, such as a gradient echo sequence with different echo times.
  • the field map is used to calculate the shimming coefficients of the targeted optimal imaging region.
  • the coefficients are then converted to shimming coil current values as described previously.
  • the coefficients are for all shimming coils. For this embodiment, it corresponds to 0 th , 1 st and 2 nd order shim fields. After the fields for all the sample points have been mapped, each shimming coil has several coefficients corresponding to different sample points of different table positions.
  • each shim coil at the sample points are interpolated to obtain continuous shimming values.
  • the corresponding shimming values are loaded to maintain optimal field homogeneity at the optimal target region. Because the target shimming region is restricted to a limited region at the central part of the magnet, the shimming result is optimal.

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  • General Physics & Mathematics (AREA)
  • Health & Medical Sciences (AREA)
  • Nuclear Medicine, Radiotherapy & Molecular Imaging (AREA)
  • Vascular Medicine (AREA)
  • General Health & Medical Sciences (AREA)
  • Radiology & Medical Imaging (AREA)
  • Engineering & Computer Science (AREA)
  • Signal Processing (AREA)
  • High Energy & Nuclear Physics (AREA)
  • Magnetic Resonance Imaging Apparatus (AREA)

Abstract

La présente invention décrit de multiples procédés pour effectuer des examens par balayage du corps en utilisant un petit système d’imagerie par résonance magnétique (IRM) efficace en termes de coûts. Une homogénéité du champ magnétique élevé d’un petit I.R.M. ouvert (DSWI) est obtenue par une combinaison d’un ajustement passif et d’un ajustement actif d’ordre élevé. Un procédé d’ajustement dynamique pendant la formation d’image (DSWI) est prévu pour optimiser dynamiquement l’homogénéité de champ pour chaque plaque (tranche) balayée lors de la formation d’image. L’invention concerne également un procédé qui balaye un grand volume sujet en utilisant uniquement une région de formation d’image optimale limitée d’un aimant en ajustant de manière continue la position et les orientations du patient avec une table pour patient ayant 6 degrés de liberté.
PCT/US2005/004920 2005-02-15 2005-02-15 Procede d’utilisation d’un petit scanner irm WO2006088453A1 (fr)

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PCT/US2005/004920 WO2006088453A1 (fr) 2005-02-15 2005-02-15 Procede d’utilisation d’un petit scanner irm

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GB2448479A (en) * 2007-04-18 2008-10-22 Siemens Magnet Technology Ltd Shims for producing toroidal MRI volume of interest
WO2010001285A1 (fr) 2008-07-04 2010-01-07 Koninklijke Philips Electronics N.V. Imagerie par rm avec un champ de visualisation étendu
CN101266289B (zh) * 2008-04-25 2010-06-09 浙江大学 开放式mri系统中横向梯度线圈的变形空间设计方法
US8320647B2 (en) 2007-11-20 2012-11-27 Olea Medical Method and system for processing multiple series of biological images obtained from a patient
WO2013142459A3 (fr) * 2012-03-21 2013-12-19 Clear-Cut Medical Ltd. Système d'imagerie par résonance magnétique (irm) pour l'évaluation de marge d'un échantillon ex-vivo
CN111157931A (zh) * 2020-01-17 2020-05-15 奥泰医疗系统有限责任公司 一种磁共振动态匀场方法
EP3739352A1 (fr) * 2019-05-16 2020-11-18 Siemens Healthcare GmbH Installation de résonance magnétique et procédé de compensation d'inhomogénités du champ magnétique de base du premier ordre dans une zone d'examen de l'installation de résonance magnétique
WO2021081144A1 (fr) * 2019-10-22 2021-04-29 Schlumberger Technology Corporation Systèmes et procédés de mesure numérique de cornue

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Cited By (16)

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GB2448479B (en) * 2007-04-18 2009-06-03 Siemens Magnet Technology Ltd Improved shim for imaging magnets
US7714580B2 (en) 2007-04-18 2010-05-11 Siemens Plc Shim for imaging magnets
GB2448479A (en) * 2007-04-18 2008-10-22 Siemens Magnet Technology Ltd Shims for producing toroidal MRI volume of interest
US8320647B2 (en) 2007-11-20 2012-11-27 Olea Medical Method and system for processing multiple series of biological images obtained from a patient
US9123100B2 (en) 2007-11-20 2015-09-01 Olea Medical Method and system for processing multiple series of biological images obtained from a patient
CN101266289B (zh) * 2008-04-25 2010-06-09 浙江大学 开放式mri系统中横向梯度线圈的变形空间设计方法
WO2010001285A1 (fr) 2008-07-04 2010-01-07 Koninklijke Philips Electronics N.V. Imagerie par rm avec un champ de visualisation étendu
US9797973B2 (en) 2008-07-04 2017-10-24 Koninklijke Philips Electronics N.V. MR imaging with extended field of view
WO2013142459A3 (fr) * 2012-03-21 2013-12-19 Clear-Cut Medical Ltd. Système d'imagerie par résonance magnétique (irm) pour l'évaluation de marge d'un échantillon ex-vivo
CN104303067A (zh) * 2012-03-21 2015-01-21 明确医疗有限公司 用于离体样本的边缘估计的mri系统
US9689817B2 (en) 2012-03-21 2017-06-27 Clear-Cut Medical Ltd. MRI system for margin assessment of ex-vivo sample
EP3739352A1 (fr) * 2019-05-16 2020-11-18 Siemens Healthcare GmbH Installation de résonance magnétique et procédé de compensation d'inhomogénités du champ magnétique de base du premier ordre dans une zone d'examen de l'installation de résonance magnétique
US11105874B2 (en) 2019-05-16 2021-08-31 Siemens Healthcare Gmbh Magnetic resonance unit and method for compensating for basic magnetic field inhomogeneities of the first order in an examination region of the magnetic resonance unit
WO2021081144A1 (fr) * 2019-10-22 2021-04-29 Schlumberger Technology Corporation Systèmes et procédés de mesure numérique de cornue
US11994480B2 (en) 2019-10-22 2024-05-28 Schlumberger Technology Corporation Digital retort measurement systems and methods
CN111157931A (zh) * 2020-01-17 2020-05-15 奥泰医疗系统有限责任公司 一种磁共振动态匀场方法

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