MRI and procedures for operating a clinical pTX system

Abstract

Magnetic resonance imaging system (101) with a pulse sequence control unit (110) which is designed - for storing a local coil-specific restriction of the amplitude and phase space of a local coil in the form of a local coil-specifically defined (S2) transmit vector (B1_H), which specifies for a local coil (106) with which amplitude ratios (|A1 / A2|, |A2 / A3|) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) transmitting elements (TX1, TX2, TX3) of the local coil (106) may be controlled by a transmitting device (RFPA), - to start a patient-specific calibration measurement (S6) on a patient (105) to generate patient-specific calibration data (B1_Map) representing a field distribution, - to determine (S7) deviations of the calibration data (B1_Map) representing a field distribution generated in a patient-specific calibration measurement on a patient (105) from the stored local coil-specific transmission vector (B1_H), wherein an imaging MRI measurement (S10) is not permitted (S8) by an interrupt signal from a monitoring device (K_SUP) if deviations exceed a threshold value, and is otherwise carried out (S10) under monitoring with the monitoring device (K_SUP).

Description

[0001] The invention relates to a magnetic resonance imaging system and a method for its operation.

[0002] Magnetic resonance imaging devices (MRIs) for examining objects or patients using magnetic resonance imaging are known, for example, from DE 103 14 215 B4.

[0003] WO 2013 / 105006 A1 discloses a method for MR imaging that provides an improved B1 field distribution. This method determines a B1 distribution.

[0004] It is an object of the present invention to optimize a magnetic resonance imaging system and a method for its operation. This object is achieved by the features of the independent patent claims. Advantageous further developments are specified in the subclaims and the following description.

[0005] Further features and advantages of possible embodiments of the invention will become apparent from the following description of exemplary embodiments with reference to the drawings. In this drawing: Fig. 1 Elements of an MRI, Fig. 2 steps of a method according to the invention, Fig. 3 schematically and simplified an MRI system.

[0006] Fig. 3 shows (among other things, particularly for technical background information) an MRI imaging device 101 (located in a shielded room or Faraday cage F) with a hollow cylinder 102 with a tubular chamber 103 in this case, into which a patient bed 104 with a body, e.g. of an examination object (e.g. a patient) 105 (with or without a local coil arrangement 106) can be moved in the direction of the arrow z in order to generate images of the patient 105 using an imaging method. A local coil arrangement 106 is arranged on the patient here, with which images of a partial area of ​​the body 105 in the FOV can be generated in a local area (also called field of view or FOV) of the MRI. Signals of the local coil arrangement 106 can be evaluated by an evaluation device (168, 115, 117, 119, 120, 121 etc.) of the MRI 101 that can be connected to the local coil arrangement 106, e.g. via coaxial cable or radio (167) etc. (e.g.converted into images, stored or displayed).

[0007] In order to examine a body 105 (an examination object or a patient) using magnetic resonance imaging with an MRI device 101, various magnetic fields, whose temporal and spatial characteristics are precisely coordinated with one another, are radiated onto the body 105. A strong magnet (often a cryomagnet 107) in a measuring cabin with a tunnel-shaped opening 103, generates a static, strong main magnetic field B0, which is, for example, 0.2 Tesla to 3 Tesla or even more. A body 105 to be examined is supported on a patient couch 104 and moved into an approximately homogeneous region of the main magnetic field B0 in the viewing area FoV (also called the "field of view" or "field of view"). An excitation of the nuclear spins of atomic nuclei of the body 105 takes place via magnetic high-frequency excitation pulses B1 (x, y, z, t) which are transmitted via a high-frequency antenna 108 (and / or possiblya local coil arrangement). Radiofrequency excitation pulses are generated, for example, by a pulse generation unit 109, which is controlled by a pulse sequence control unit 110. After amplification by a radiofrequency amplifier 111, they are guided to the radiofrequency antenna 108 and / or local coil 106. The radiofrequency system shown here is merely schematically indicated. More than one pulse generation unit 109, more than one radiofrequency amplifier 111, and multiple radiofrequency antennas 108a, b, c may also be used in a magnetic resonance scanner 101.

[0008] Furthermore, the magnetic resonance device 101 has gradient coils 112x, 112y, 112z, with which magnetic gradient fields B G(x, y, z, t) are radiated for selective slice excitation and for spatial encoding of the measurement signal. The gradient coils 112x, 112y, 112z are controlled by a gradient coil control unit 114 (and possibly via amplifiers Vx, Vy, Vz), which, like the pulse generation unit 109, is connected to the pulse sequence control unit 110.

[0009] Signals emitted by the excited nuclear spins (of the atomic nuclei in the object under examination) are received by the body coil and / or at least one local coil arrangement 106, amplified by associated radio-frequency preamplifiers 116, and further processed and digitized by an MR receiver 117. The recorded measurement data are digitized and stored as complex numerical values ​​in a k-space matrix. A corresponding MR image can be reconstructed from the k-space matrix using a multidimensional Fourier transformation.

[0010] For a coil that can be operated in both transmit and receive mode, such as the body coil or a local coil 106, the correct signal transmission is controlled by an upstream transmit-receive switch 118.

[0011] An image processing unit 119 generates an image from the measured data, which is displayed to a user via a control console 120 and / or stored in a storage unit 121. A central computer unit 122 controls the individual system components. In MR imaging, images with a high signal-to-noise ratio (SNR) are typically acquired today using so-called local coil arrays (coils). These are antenna systems that are mounted in close proximity on (anterior) or below (posterior), on or in the body 105. During an MR measurement, the excited nuclei induce a voltage in the individual antennas of the local coil, which is then amplified by a low-noise preamplifier (e.g., LNA, preamp) and finally passed on to the receiving electronics. To improve the signal-to-noise ratio, even with high-resolution images, so-called high-field systems (1.5T-12T or more) are used. If more individual antennas can be connected to an MR receiving system than there are receivers, a switching matrix (sometimes also referred to or implemented as an RCCS) is installed between the receiving antennas and the receiver.This routes the currently active receive channels (usually those currently within the magnet's field of view) to the available receivers. This makes it possible to connect more coil elements than there are receivers, since with full-body coverage, only the coils located within the magnet's field of view, or homogeneity volume, need to be read.

[0012] A local coil arrangement 106, for example, is generally referred to as an antenna system which can consist of one or, as an array coil, several antenna elements (in particular coil elements). These individual antenna elements are designed, for example, as loop antennas (loops), butterfly, flex coils or saddle coils. A local coil arrangement comprises, for example, coil elements, a preamplifier, additional electronics (sheath wave barriers, etc.), a housing, supports and usually a cable with a plug through which it is connected to the MRI system. A receiver 168 attached to the MRI system filters and digitizes a signal received by a local coil 106, e.g. via radio, etc. and transmits the data to a digital signal processing device which usually derives an image or a spectrum from the data obtained by a measurement and makes it available to the user, e.g. for subsequent diagnosis by the user and / or for storage.

[0013] Fig. 1-3 illustrate embodiments according to the invention.

[0014] For signal excitation in the latest generation of MR scanners, multiple transmission channels and elements (e.g., in one or more local coils) are used simultaneously, using methods that are at least known internally. This is also known as "pTX" (= parallel transmit method). Safe operation for the patient can only be guaranteed if the SAR (Specific Absorption Rate) dose is known in terms of time and space.

[0015] A pTX system allows for the free adjustment of the phase and amplitude of each transmit element, making the calculation of SAR limits a complex task. Depending on the phase and amplitude, so-called SAR hotspots can occur in the tissue, which are also influenced by the individual patient's physiology (e.g., cysts or tumors).

[0016] For single-channel transmitter systems known at least internally, a so-called "K-factor" is determined based on a finite element simulation (FE) and validation measurements. This describes the SAR dose occurring per unit of time and amplitude (in the worst case, i.e., in the SAR hotspot of the local coil). Taking into account a safety margin (usually at least a factor of 2), the applicable RF power is limited by the k-factor to the legally prescribed limit (lookahead or online monitoring).

[0017] For pTX systems, an analogous procedure is known, at least internally, in which the safety margin scales with the number of channels. This takes into account the risk of a worst-case overlap of the SAR hotspots of the transmitting elements in the patient. The disadvantage is that the potential advantages of the pTX system cannot be fully utilized due to the high safety margins.

[0018] Another method is known, at least internally (the VOP method = virtual observation points), in which the SAR limits are predicted by applying a parameterizable patient model. The model is determined through a series of FE-based simulations. The SAR hotspots in the patient's tissue (i.e., several models with different sizes, genders, ages, etc.) are calculated based on the electromagnetic fields generated by the transmitting elements. The advantage is that the required safety margins can be significantly lower than with the K-factor approach (and do not increase linearly with the number of transmitting elements).Since the influence of each transmitting element can be represented independently, the model can be used to calculate the SAR in the patient for any linear combination of the transmitting amplitudes and phases (before the measurement and also during the measurement by comparison with measured transmitted and reflected RF amplitudes and phases).

[0019] The complexity of the procedure and the associated difficulty of proving safe clinical operation can be a disadvantage.

[0020] Patient movement during measurements, anatomical deviations from the predicted standard model, and thermal effects in the TX / RX path can lead to a change in the actual SAR hotspots compared to the theoretically assumed hotspots. The larger the number of transmitting elements, the more pronounced such effects can be. These aspects (and others) would be analyzed and a safe solution developed as part of a potential clinical approval.

[0021] According to embodiments of the invention, for a clinical application, a local coil-specific restriction of the amplitude and phase space of the transmission elements (antennas TX1, TX2, TX3) of at least one local coil 106 (and / or body coil) is introduced, e.g. here in the form of a transmission vector Bl_H, which indicates for a local coil 106 which amplitude-phase combinations the (pulse sequence control unit 110 of the) local coil 106 may transmit.

[0022] Fig. 1 shows a simplified pulse sequence control unit 110, which supplies several antennas TX1, TX2, TX3 (a local coil) (which may also be used for reception) with transmission signals (with permissible amplitude / phase combinations) via an amplifier 111, e.g. called RFPA, wherein the transmission signals (and / or returning signals) transmitted to the antennas TX1, TX2, TX3 are monitored by a monitoring device K_SUP.

[0023] How Fig. 2, the transmission vector B1_H is first determined for each local coil 106 in a step S2 on the basis of an FE simulation (in a step S1) and / or a calibration measurement (on a patient 105 to be examined in the MRI by MRI “measurement”), which indicates the amplitude ratios and phases permitted for MRI imaging.

[0024] A patient-specific calibration measurement (in a step S6) is carried out on a patient 105 to be examined (before an MRI imaging on him) with a conservatively calculated SAR dose, whereby a wide variety of methods for determining the SAR dose on the patient 105 are conceivable (e.g., microwave thermometry, consideration of phases and amplitudes of forward and backward transmission power to / from antennas of the (at least) one local coil, etc.).

[0025] From calibration data obtained with the calibration measurement S6 (e.g. as 'B1 map' or with reference symbol 'B1_Map' in Fig. 2), the transmission vector B1_H (for MRI imaging on the patient) can be calculated / optimized such that, as far as possible, a maximally homogeneous B1 excitation can be achieved in the coil volume (volume within a local coil 106). In addition, the relative deviations of the transmission vector B1_H from the simulated reference value are limited in order to limit resulting SAR increases (at this point in the process, anatomical deviations would potentially be detectable). The result of these calculations is the vector space V_B1_H.

[0026] Especially for magnetic field strengths >3T, such field homogenization in the imaging volume is an important prerequisite for clinically acceptable operation of the MR scanner 101.

[0027] Due to the overall homogeneous excitation in the imaging volume (FoV), the calculation of the SAR dose is simplified. Essentially, the classic K-factor approach can be used, i.e., the SAR hotspots (calculated in step S3) are calculated based on an FE model (finite element model) of the local coil 106 (with a patient model) in the homogeneous excitation mode ('CP mode'), and then, if necessary (in step S4 in Fig. 2) SAR hotspots for B1_H are calculated for possible deviations such as different possible positions of the patient 105 in the MRI 101 and / or for different possible head sizes of a patient 105, and a K-factor determined thereby is transferred to a SAR monitoring device S5 (S5).

[0028] Before starting an imaging measurement (S10, i.e. MRI imaging on the patient), the (monitoring) vector V_B1_H (which can indicate in which range ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12 angle ϕ and in which ranges |A1 / A2| <a1, |A2 / A3|<a2 sich Amplituden A1, A2, A3 und / oder Amplitudenverhältnisse a1, a2 befinden) an die (in Fig. 1 simplified representation) measurement system independent monitoring component K_SUP is transferred as in Fig. 2 indicated in step S9.

[0029] The monitoring device K_SUP includes, for example, measuring devices for the phase-accurate measurement of transmitted and reflected RF power of all transmitting elements TX1, TX2, TX3, etc. In an advantageous embodiment, the monitoring device K_SUP can be a component of an MR receiver 117 in a controller and / or in a local coil 106, in a further embodiment, an independent measuring unit with directional couplers and the required evaluation circuit or software.

[0030] During, for example, a calibration measurement (S6) and / or an entire imaging measurement SMR (i.e. examination of a patient with a magnetic resonance imaging device) is checked (according to reference symbol S7 in Fig. 2) one (or in case of several, each) monitoring device K_SUP at suitable time intervals (e.g. 10µs (microseconds)), whether the measured (transmitted) amplitude ratios (|A1 / A2|, |A2 / A3| etc.) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) of the transmitting elements TX1, TX2, TX3 are in the permissible range (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12, |A1 / A2| <a1, |A2 / A3|<a2 usw.) liegen. Im durch Schritt S8 in Fig. In the event of an error described in section 2 (e.g. due to strong patient movement or faulty / inadmissible RF pulses), the monitoring unit K_SUP sends an interruption signal to the pulse sequence control unit 110. As an additional error criterion, the reflected power on a connecting line AL1, AL2, AL3 of an antenna TX1, TX2, TX3 can optionally be used.

Claims

[1] Magnetic resonance imaging system (101) with a pulse sequence control unit (110) which is designed - for storing a local coil-specific restriction of the amplitude and phase space of a local coil in the form of a local coil-specific (S2) transmission vector (B1_H), which specifies, locally coil-specifically for a local coil (106), with which amplitude ratios (|A1 / A2|, |A2 / A3|) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) transmission elements (TX1, TX2, TX3) of the local coil (106) may be controlled by a transmission device (RFPA), - to start a patient-specific calibration measurement (S6) on a patient (105) to generate patient-specific calibration data (B1_Map) representing a field distribution, - for determining (S7) deviations of the calibration data (B1_Map) representing a field distribution generated in a patient-specific calibration measurement on a patient (105) from the stored local coil-specific transmission vector (B1_H), wherein an imaging MRI measurement (S10) is not permitted (S8) by an interruption signal of a monitoring device (K_SUP) if deviations exceed a threshold value, and is otherwise carried out (S10) under monitoring by the monitoring device (K_SUP). [2] Magnetic resonance imaging system (101) according to claim 1, characterized by that the monitoring device (K_SUP) is designed for phase-accurate measurement of transmitted and / or reflected RF power of the transmitting elements (TX1, TX2, TX3). [3] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by , that the monitoring device (K_SUP) is designed to check during (S10) an imaging measurement (SMR), whether measured and / or transmitted amplitude ratios (|A1 / A2| , |A2 / A3|) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) of the transmitting elements (TX1, TX2, TX3) in a permissible (V_B1_H) range (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12; |A1 / A2| <a1, |A2 / A3| <a2 usw.) liegen. [4] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by , that the monitoring device (K_SUP) is designed to check at intervals during (S10) an imaging measurement (SMR), whether measured and / or transmitted amplitude ratios (|A1 / A2|, |A2 / A3|) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) of the transmitting elements (TX1, TX2, TX3) are in a permissible range specified by a monitoring vector (V_B1_H) (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12; |A1 / A2| <a1, |A2 / A3|<a2 usw.) liegen. [5] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by , that the monitoring device (K_SUP) is designed to if a check during (S10) an imaging measurement (SMR) shows that measured and / or transmitted amplitude ratios (|A1 / A2|, |A2 / A3|) and phases (ϕ01<ϕ<ϕ11, ϕ02<ϕ< <ϕ12) of the transmitting elements (TX1, TX2, TX3) are outside a permissible range specified by a monitoring vector (V_B1_H) (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12, |A1 / A2| <a1, |A2 / A3|<a2 usw.) liegen, to send an interruption signal (U) for causing an interruption of the imaging measurement to the pulse sequence control unit (110) of the magnetic resonance imaging system (101). [6] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by , that the monitoring device (K_SUP) is designed, if a check during (S10) an imaging measurement (SMR) shows, that a reflected power at one or more connecting lines (AL1, AL2, AL3) of transmitting devices (TX1, TX2, TX3) of a local coil (106) is above a predetermined threshold value, to send an interruption signal (U) for causing an interruption of the imaging measurement to the pulse sequence control unit (110) of the magnetic resonance imaging system (101). [7] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by that the monitoring device (K_SUP) is designed as a component of an MR receiver (117; 106). [8] Magnetic resonance imaging system (101) according to one of claims 1 to 6, characterized by that the monitoring device (K_SUP) is designed as an independent measuring unit with directional couplers and an evaluation device. [9] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by that the monitoring device (K_SUP) is designed to take into account a monitoring vector (V_B1_H) which indicates in which range (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12) angle (ϕ) and in which ranges (|A1 / A2| <a1, |A2 / A3|<a2) sich Amplituden (A1, A2, A3) und / oder Amplitudenverhältnisse (a1, a2) befinden dürfen. [10] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by , that the monitoring device (K_SUP) is designed to compare phases and / or amplitudes and / or amplitude ratios on at least one connecting line (AL1, AL2, AL3) of an antenna (TX1, TX2, TX3) of the local coil (106) with (ϕ01<ϕ<ϕ11, ϕ02<ϕ<ϕ1<ϕ12, |A1 / A2| <a1, |A2 / A3| <a2, A1, A2, A3) Phasen (ϕ) und / oder Amplituden (A1, A2, A3) und / oder Amplitudenverhältnissen (a1, a2). [11] Magnetic resonance imaging system (101) according to one of the preceding claims, characterized by that it is designed to store at least one fixed (S4) factor (K-factor) representing a maximum permissible SAR dose in the patient (105) occurring per unit of time and amplitude and to transfer it (S5) to a monitoring device (K_SUP). [12] Method for operating a magnetic resonance imaging system (101), characterized by that a local coil-specific restriction of the amplitude and phase space of a local coil is used in the form of a stored, fixed (S2) transmission vector (B1_H), which specifies for a local coil (106) with which amplitudes (A1, A2, A3) and phases (ϕ01 <ϕ<ϕ11, ϕ02<ϕ< <ϕ12) sendeelemente (tx1, tx2, tx3) der lokalspule (106) von einer sendeeinrichtung (rfpa) angesteuert werden dürfen,- that a calibration measurement is started on a patient (105) to generate calibration data (B1_Map) representing a field distribution, - that deviations from calibration data (B1_Map) representing a field distribution generated in a calibration measurement on a patient (105) are determined with the stored transmission vector (B1_H), wherein an imaging MRI measurement (S10) is not permitted (S8) by an interruption signal from a monitoring device (K_SUP) if deviations exceed a threshold value, and is otherwise carried out (S10) under monitoring by the monitoring device (K_SUP).

Images