MR imaging with T1-compensated B1 mapping

Abstract

This invention relates to an MR imaging method. The object of this invention is to provide an improved B1 mapping method less affected by T1 relaxation. This invention proposes generating a first stimulated echo imaging sequence (25), which includes a sequence of at least two preparation RF pulses (α) irradiated during a first preparation period (21) and a readout RF pulse (β) irradiated during a first acquisition period (22) following the first preparation period (21). A first set of FID signals (I1) is acquired during the first acquisition period (22). FID ) and the first group of stimulated echo signals (I STE A second stimulated echo imaging sequence (27) is generated, which again includes a sequence of at least two preparation RF pulses (α) irradiated during the second preparation period (21) and a readout RF pulse (β) irradiated during the second acquisition period (22) after the second preparation period (21). A second set of FID signals (I) is acquired during the second acquisition period (22). FID ) and the second group of stimulated echo signals (I STE The first group of FID signals (I) FID ) and the second group of FID signals (I FID ) have different T1 weights, and / or the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE The RF field of the RF pulse has different T1 weights. The B1 plot indicating the spatial distribution of the RF field is based on the first group of acquired FID signals (I FID ) and the second group of FID signals (I FID ) and the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE The results are derived by using different T1 weights to compensate for the effects on the B1 graph caused by T1 relaxation. Preferably, an RF inversion pulse is applied before the first preparation period (21) or the second preparation period (21) to obtain different T1 weights. Furthermore, the present invention relates to an MR device (1) and a computer program for the MR device (1).

Description

Technical Field

[0001] This invention relates to the field of magnetic resonance (MR) imaging. The invention focuses on a method for performing MR imaging on at least a portion of an object. The invention also relates to MR devices and computer programs to run on the MR devices. Background Technology

[0002] Today, image-forming MR methods that utilize the interaction between magnetic fields and nuclear spins to form two-dimensional or three-dimensional images are widely used, especially in the field of medical diagnostics. This is because these image-forming MR methods are superior to other imaging methods in many ways for imaging soft tissues, do not require ionizing radiation, and are generally non-invasive.

[0003] In a typical MR method, the patient's body is positioned in a strong, uniform magnetic field (B0 field), the direction of which simultaneously defines the axis (typically the z-axis) of the coordinate system upon which the measurement is based. Depending on the magnetic field strength, the field generates different energy levels for individual nuclear spins, which can be excited (spin resonance) by applying an alternating electromagnetic field (RF field, also known as the B1 field) at a defined frequency (so-called Larmor frequency or MR frequency). From a macroscopic perspective, the distribution of individual nuclear spins produces a global magnetization that can be decoupled from equilibrium by applying an electromagnetic pulse (RF pulse) of appropriate frequency perpendicular to the z-axis, causing the magnetization to precess about the z-axis. This precession describes the surface of a cone, the aperture angle of which is called the flip angle. The magnitude of the flip angle depends on the intensity and duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spin deflects from the z-axis to the transverse plane (flip angle 90°).

[0004] After the RF pulse terminates, the magnetization relaxes back to its original equilibrium state, with magnetization in the z-direction re-established with a first time constant T1 (spin lattice or longitudinal relaxation time) and magnetization in the direction perpendicular to the z-direction relaxing with a second time constant T2 (spin-spin or transverse relaxation time). Changes in magnetization can be detected by means of one or more receiving RF coils arranged and oriented within the examination volume of the MR device in a manner that measures changes in magnetization in the direction perpendicular to the z-axis. After applying, for example, a 90° pulse, the decay of transverse magnetization is accompanied by a transition of nuclear spins from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (phase decoupling) (induced by local inhomogeneities in the main magnetic field). Phase decoupling can be compensated for by means of a refocusing pulse (e.g., a 180° pulse). This generates an echo signal (spin echo) in the receiving coil.

[0005] To achieve spatial resolution within the body, linear magnetic field gradients extending along three principal axes are superimposed on a uniform magnetic field, resulting in a linear spatial dependence of the spin resonant frequency. The signal picked up in the receiving coil then contains components of different frequencies that can be correlated with different locations within the body. The MR signal data obtained via the RF coil corresponds to the spatial frequency domain and is referred to as k-space data. k-space data typically comprises multiple lines acquired using different phase codes. Each line is digitized by collecting multiple samples. A set of k-space data is converted into an MR image using a Fourier transform.

[0006] Typically, the generated transmitted RF field (B1) is expected to be generated. + The field (hereinafter referred to as B1) has relatively uniformity to excite magnetic resonance imaging (MRI) across the cross-section and / or volume of the patient's body being imaged, thus preserving important MR contrast characteristics. However, as the MR frequency increases, accompanied by an increase in the strength of the main magnetic field, this becomes more difficult due to conduction losses and wave propagation effects within the patient's body. Therefore, accurate measurement of the spatial distribution of the emitted RF field is important for many MR imaging applications, enabling appropriate prospective (if available) and retrospective correction / compensation. This requires robust and fast B1 mapping techniques. Especially at higher field strengths, the acquisition speed of known B1 mapping methods is often limited by SAR (specific absorption rate), T1 relaxation time, or characteristic sequence properties. Therefore, multi-emission B1 calibration scans still represent a cumbersome procedure, as measurement time is proportional to the number of emission channels involved, increasing the need for more efficient sampling schemes.

[0007] K. Nehrke and (Magn. Reson. Med., Vol. 86, No. 5, pp. 1517-1526, 2012) A B1 mapping method known as the “DREAM” (Double Refocusing Echo Acquisition Mode) has been proposed, which can cover the entire transmit coil volume in about one second, exceeding the magnitude of other existing fast B1 mapping techniques such as standard flip-angle imaging (AFI). The DREAM method employs a stimulated echo sequence consisting of a preparation period and two preparation RF pulses, and then generates a single snapshot low flip-angle gradient echo queue in time during an acquisition period following the preparation period. The stimulated echo and free inductive attenuation (FID) are refocused into a gradient recalled echo in a fast field echo queue (TFE) manner in close temporal succession, and their ratio is used to derive the actual flip angle of the stimulated echo preparation RF pulse.

[0008] The known DREAM method ignores T1 relaxation and therefore represents approximate results. During the preparation phase, even with low-to-high k spatial sampling, T1 relaxation leads to the recovery of the FID signal and the attenuation of the stimulated signal. Therefore, T1 relaxation results in a systematic underestimation of the flip angle of the preparation RF pulse, and consequently, an underestimation of B1. This is particularly relevant when short T1 and / or long echo queues occur during acquisition. Furthermore, the error increases with the B1-encoded flip angle, thus reducing the accurate operating range of the method. Additionally, during the echo queue, T1 recovery / attenuation enhances / reduces the high spatial frequencies of the MR images reconstructed from the FID and stimulated echo signals (the FID and stimulated echo signals are sorted for low-to-high contours to derive the B1 map), leading to artifacts, especially near the edges of anatomical structures. To mitigate these problems, the echo queue during the acquisition phase must be as short as possible, which limits the resolution achievable in a single snapshot. Furthermore, the maximum B1-encoded flip angle must be limited to approximately 60°, which restricts the dynamic range of the method. Summary of the Invention

[0009] It is easy to see from the foregoing that an improved B1 mapping method is needed that is less affected by T1 relaxation.

[0010] According to the present invention, a method for performing MR imaging on at least a portion of an object placed in the examination volume of an MR device is disclosed. The method includes the following steps:

[0011] The portion of the object is subjected to a first stimulated echo imaging sequence, the first stimulated echo imaging sequence comprising at least two preparation RF pulses irradiating the portion of the object during a first preparation period and a sequence of readout RF pulses irradiating the portion of the object during a first acquisition period after the first preparation period;

[0012] During the first acquisition period, the first set of FID signals and the first set of stimulated echo signals are acquired.

[0013] The portion of the object is subjected to a second stimulated echo imaging sequence, the second stimulated echo imaging sequence comprising at least two preparation RF pulses irradiating the portion of the object during a second preparation period and a sequence of readout RF pulses irradiating the portion of the object during a second acquisition period after the second preparation period;

[0014] During the second acquisition period, a second set of FID signals and a second set of stimulated echo signals are acquired, wherein the first set of FID signals and the second set of FID signals have different T1 weights, and / or the first set of stimulated echo signals and the second set of stimulated echo signals have different T1 weights.

[0015] Based on the first and second sets of FID signals and the first and second sets of stimulated echo signals, a B1 map indicating the spatial distribution of the RF field of the RF pulse within the part of the body is derived, wherein the different T1 weightings are used to compensate for the influence of T1 relaxation on the B1 map.

[0016] The key point of this invention is the application of two stimulated echo imaging sequences, designed to impose different T1 weights on the FID signal and stimulated echo signal acquired during a first acquisition period and a second acquisition period, respectively. The echo queues for the first and second acquisition periods preferably employ the same phase and frequency encoding to provide identical k-space sampling. Different (and primarily known) T1 weights can then be used directly to compensate for the effects of T1 relaxation on the resulting B1 map. According to the invention, B1 maps are acquired using at least two instances of the known DREAM sequence. The method remains fast and maintains the high motion robustness of the DREAM method.

[0017] Generally, a stimulated echo sequence comprises three (e.g., 60° or 90°) RF pulses, where the first two RF pulses are preparation pulses. The first preparation RF pulse excites the magnetic resonance and transforms the longitudinal NMR into transverse NMR. The second preparation RF pulse “stores” a fraction of the dephased transverse NMR along the longitudinal axis. In the case of the 90° RF pulse, this fraction is almost half of the dephased transverse magnetization. A third RF pulse is applied temporally during the acquisition period following the preparation period. The third RF pulse (“read RF pulse”) transforms the stored longitudinal NMR back into transverse NMR, thereby generating the so-called stimulated echo. Other RF refocusing echoes are generated by these three RF pulse sequences, but these echoes are not of interest here and may be suppressed by a suitable gradient switching mechanism operating in parallel with the RF irradiation. According to the invention, the stimulated echo signal is acquired together with the FID signal and used for B1 mapping.

[0018] According to the present invention, the actual image generation process based on stimulated echo is accelerated by replacing the third RF pulse of the standard stimulated echo sequence with a queue of readout RF pulses with a low flip angle, wherein each readout RF pulse only refocuses a small portion of the nuclear magnetization. Multiple FID signals and stimulated echo signals with appropriate phase and frequency encoding are acquired to generate a complete B1 map. For this purpose, efficient sampling schemes such as EPI, spiral, parallel imaging, or compressed sensing can be applied in conjunction with the present invention. The timing parameters of the imaging sequence can be adjusted so that the sensitivity and chemical shift-induced effects are substantially equal for both the FID signal and the stimulated echo signal.

[0019] It should be noted that the at least two prepared RF pulses used to store the B1-encoded magnetization along the z-axis do not necessarily have to be of the same type or flip angle.

[0020] In a preferred embodiment, one or more pre-prepared RF pulses precede the first stimulated echo sequence and / or the second stimulated echo sequence, said pre-prepared RF pulses manipulating the longitudinal magnetization at the beginning of the first stimulated echo sequence to be different from the longitudinal magnetization at the beginning of the second stimulated echo sequence. Proper pre-preparation of the longitudinal magnetization (i.e., preparation of the longitudinal magnetization prior to the actual preparation period of the respective stimulated echo imaging sequence) enables the first and second sets of FID signals to have different T1 weights, and / or the first and second sets of stimulated echo signals to have different T1 weights. This is a prerequisite for utilizing different T1 weights when deriving the B1 map according to the invention.

[0021] In one possible variation, an RF inversion pulse precedes either the first preparation period or the second preparation period. Therefore, in this case, the pre-prepared RF pulse is an RF inversion pulse. By applying the RF inversion pulse only before one of the two stimulated echo sequences and simultaneously executing the other stimulated echo sequence without magnetization inversion, desired distinct T1 weights of the first and second sets of FID signals and the first and second sets of stimulated echo signals are obtained in a deliberate manner. For optimal inversion, an adiabatic RF inversion pulse can be employed. As is known to those skilled in the art, an adiabatic RF pulse is an amplitude-modulated and frequency-modulated RF pulse that is insensitive to B1 inhomogeneities, thus allowing for complete spin inversion without regard to the typically present B1 inhomogeneities.

[0022] In another variation, the one or more pre-prepared RF pulses are saturated RF pulses. In this case, longitudinal magnetization is prepared by irradiating one or more saturated RF pulses to adjust the longitudinal magnetization to zero. Different longitudinal relaxations are achieved by varying the delay between the saturated RF pulse and the start of the corresponding stimulated echo imaging sequence, such that the longitudinal magnetization at the start of the first stimulated echo sequence differs from the longitudinal magnetization at the start of the second stimulated echo sequence.

[0023] In another preferred embodiment, the portion of the object undergoes a second stimulated echo imaging sequence with a delay following the first stimulated echo sequence, the delay being shorter than T1. This short delay between the two stimulated echo imaging sequences prevents the longitudinal magnetization from fully relaxing to equilibrium, thereby ensuring that the longitudinal magnetization at the beginning of the first stimulated echo sequence differs from the longitudinal magnetization at the beginning of the second stimulated echo sequence.

[0024] In another preferred embodiment, the step of deriving the B1 map involves: calculating an FID differential image by subtracting MR images reconstructed from the first set of FID signals and the second set of FID signals, respectively; and calculating a stimulated echo differential image by subtracting MR images reconstructed from the first set of stimulated echo signals and the second set of stimulated echo signals, respectively. The B1 map is derived based on the voxel-by-voxel intensity ratio of the stimulated echo differential image and the FID differential image. Results show that, for cases where magnetization reversal is applied before one of the two stimulated echo sequences, deriving the B1 map based on the intensity ratio of the differential images cancels out the T1 relaxation effect that occurs during magnetization preparation for readout. Therefore, a B1 map unaffected by T1 relaxation is obtained.

[0025] According to another preferred embodiment of the invention, both of the at least two preparation RF pulses have a flip angle of 45°-90°. In this way, the amplitude of the acquired stimulated echo signal is maximized, which is advantageous in terms of signal-to-noise ratio.

[0026] In another preferred embodiment, the two stimulated echo signals are acquired after each readout RF pulse during each of the first and second acquisition periods. These two stimulated echo signals can be a direct stimulated echo signal and a conjugate (also referred to as a "virtual") stimulated echo signal. Acquisition of the direct and conjugate stimulated echo signals supports a T2-compensated acquisition scheme. In particular, the effects on the B1 plot caused by T2 relaxation can be compensated using different T2 weights of the two stimulated echo signals. The two stimulated echo signals correspond to different coherent paths generated from the stimulated echo sequence. For details, refer to WO2013 / 105006A1.

[0027] The method of the present invention described to date can be performed by means of an MR device comprising: at least one main magnet coil for generating a uniform, stable magnetic field within an examination volume; a plurality of gradient coils for generating switching magnetic field gradients in different spatial directions within the examination volume; at least one RF coil for generating RF pulses within the examination volume and / or for receiving MR signals from the body of a patient positioned within the examination volume; a control unit for controlling the temporal succession of the RF pulses and the switching magnetic field gradients; and a reconstruction unit for reconstructing MR images based on the received MR signals. The method of the present invention is preferably implemented by corresponding programming of the reconstruction unit and / or control unit of the MR device.

[0028] The method of the present invention can be advantageously performed in most MR devices currently used in clinical practice. Therefore, only a computer program is needed to control the MR device to perform the steps of the method described above. The computer program can reside on a data carrier or in a data network, and can be downloaded and installed in the control unit of the MR device. Attached Figure Description

[0029] The accompanying drawings disclose preferred embodiments of the invention. However, it should be understood that the drawings are for illustrative purposes only and not as a limitation of the invention. In the drawings:

[0030] Figure 1 An MR apparatus for performing the method of the present invention is schematically shown;

[0031] Figure 2 A schematic diagram illustrating the imaging sequence according to the present invention is shown;

[0032] Figure 2a The succession of two stimulated echo imaging sequences in a variant of the present invention is illustrated schematically.

[0033] Figure 2b The succession of two stimulated echo imaging sequences in another variation of the invention is illustrated schematically.

[0034] Figure 3 This demonstrates the accuracy of the B1 mapping method of the present invention;

[0035] Figure 4 The diagram illustrates the B1 mapping with and without T1 compensation and T2 compensation according to the present invention. Detailed Implementation

[0036] refer to Figure 1The diagram illustrates an MR device 1. The device includes a superconducting or resistive main magnet coil 2, which creates a substantially uniform, time-constant main magnetic field B0 along the z-axis through the examination volume. The device also includes a set of first-order, second-order, and third-order (where applicable) shimming coils 2', wherein the current through each individual shimming coil in this set of shimming coils 2' is controllable for the purpose of minimizing deviations in B0 within the examination volume.

[0037] The magnetic resonance generation and manipulation system applies a series of RF pulses and switched magnetic field gradients to reverse or excite nuclear magnetic spin, induce magnetic resonance, refocus magnetic resonance, manipulate magnetic resonance, encode magnetic resonance in a spatial manner and other ways, saturate spin, etc., thereby performing MR imaging.

[0038] Most specifically, gradient pulse amplifier 3 applies current pulses to selected coils in the whole-body gradient coils 4, 5, and 6 along the x, y, and z axes of the examination volume. Digital RF frequency transmitter 7 transmits RF pulses or pulse packets to RF body coil 9 via transmit / receive switch 8 to transmit RF pulses into the examination volume. Typical MR imaging sequences consist of packets of short-duration RF pulse segments used together, and any applied magnetic field gradient achieves selected manipulation of the MRI. RF pulses are used to saturate, excite, reverse magnetization, refocus, or manipulate the resonance, and to select portions of the body 10 located within the examination volume. The MR signal is also picked up by RF body coil 9.

[0039] To generate MR images of a limited region of body 10 using parallel imaging, a set of local array RF coils 11, 12, and 13 are placed adjacent to the selected region for imaging. The array coils 11, 12, and 13 can be used to receive MR signals generated by the body coil RF emission. In parallel emission applications, the array RF coils 11, 12, and 13 can also be used for RF emission, for example, for RF shimming purposes.

[0040] The resulting MR signal is picked up by RF body coil 9 and / or array RF coils 11, 12, 13 and demodulated by receiver 14, which preferably includes a preamplifier (not shown). Receiver 14 is connected to RF coils 9, 11, 12 and 13 via transmit / receive switch 8.

[0041] The host computer 15 controls the current flowing through the shimming coil 2', the gradient pulse amplifier 3, and the transmitter 7 to generate any MR imaging sequence from a variety of MR imaging sequences (e.g., echo plane imaging (EPI), echo volume imaging, gradient and spin echo imaging, fast spin echo imaging, etc.). For a selected sequence, the receiver 14 rapidly and alternately receives one or more MR data lines after each RF excitation pulse. The data acquisition system 16 performs analog-to-digital conversion on the received signals and converts each MR data line into a digital format suitable for further processing. In modern MR devices, the data acquisition system 16 is a separate computer dedicated to acquiring raw image data.

[0042] Finally, reconstruction processor 17 reconstructs the original digital image data into an image representation by applying a Fourier transform or other suitable reconstruction algorithm, such as SENSE or SMASH. The MR image can be represented by planar slices of the patient, an array of parallel planar slices, a three-dimensional volume, etc. The image is then stored in an image memory, which can be accessed to convert slices, projections, or other portions of the image representation into a suitable format for visualization (e.g., via video monitor 18), which provides a human-readable display of the resulting MR image.

[0043] Figure 2 A diagram illustrating an imaging sequence according to the present invention is shown. The depicted imaging sequence is a stimulated echo sequence, which is subdivided into a preparation period 21 and an acquisition period 22. During the preparation period 21, two preparation RF pulses with a flip angle α are applied. The two preparation RF pulses are separated by a time interval T. E Separated. A phase depletion magnetic field gradient G is applied between the two prepared RF pulses. m2 A sequence of readout RF pulses with a flip angle β is generated during the acquisition phase 22, following the preparation phase 21. This occurs in the presence of a refocusing / readout magnetic field gradient G. m1 / G m In this case, the FID signal I is acquired after each read pulse. FID Signal and stimulated echo signal I STE As a gradient refocusing echo, a phase scrambler gradient is applied to suppress unwanted residual transverse magnetization. According to the invention, a phase cycling scheme is implemented, wherein the depicted imaging sequence is performed twice, once before preparation period 21 with an adiabatic 180° RF inversion pulse, and once without a preceding (…). Figure 2 (Indicated by 180° / 0°) Longitudinal magnetization reversal. For clarity, in Figure 2The slice selection and phase-encoded magnetic field gradients used in this method are omitted. In the presence of a slice selection gradient (not depicted), the adiabatic 180° RF inversion pulse can exhibit slice selectivity.

[0044] Without an RF inversion pulse and neglecting T1 relaxation, longitudinal magnetization occurs directly after preparation sequence 21, as given by the following equation:

[0045] M z,FID =cos 2 (α)·M0

[0046] M z,STE =sin 2 (α) / 2·M0

[0047] Among them, M z,FID and M z,STE These refer to unprepared (i.e., in-phase) longitudinal magnetization and stimulated echo prepared (i.e., out-of-phase) longitudinal magnetization, respectively. According to M... z,FID The generated FID signal I FID and according to M z,STE The generated stimulated echo signal I STE They are at different time points T E1 and T E1 The data was acquired at +ΔT. Two echoes I FID I STE The delay ΔT between them is determined by the following relationship:

[0048] ΔT=A mc2 / G m

[0049] Among them, A mc2 Refers to the gradient G of the phase depletor m2 The gradient-time area, and G m This refers to the intensity of the readout magnetic field gradient. Ignoring the T1 and T2 effects, the two acquired echo signals I... FID and I STE It is given by the following equation:

[0050] I FID =sin(β)·cos 2 (α)·M0

[0051] I STE =sin(β)·sin 2 (α) / 2·M0

[0052] Where β is the nominal flip angle of the read RF pulse. Combining the above equations will produce:

[0053] α = tan -1 (2ISTE / I FID )

[0054] Therefore, based on the following equation, the unknown flip angle α for preparing the stimulated echo RF pulse can be derived from the ratio of the acquired echo signals:

[0055] α = tan -1 (2I STE / I FID )

[0056] The equations above are approximate results because they do not consider T1 relaxation. T1 relaxation causes the recovery of the FID signal and the attenuation of the stimulated signal according to the following equation:

[0057] I FID (k)=sin(β)(E1·cos 2 (α)+(1-E1))·M0

[0058] I STE (k)=sin(β)·E1·sin 2 (α) / 2·M0

[0059] in,

[0060] Here, k and TR refer to the index and repetition time of the gradient echo in the sequence of read RF pulses, respectively, and ΔT is the time interval between the preparation period 21 and the first gradient echo. Therefore, E1 refers to the T1 relaxation term responsible for the recovery of the FID signal and the attenuation of the stimulated echo signal. It is therefore easy to recognize that T1 relaxation leads to a systematic underestimation of the flip angle α, especially for short T1 and / or long echo queues. The error increases with increasing α, thus reducing the accurate operating range of the method.

[0061] This invention proposes a second example of a stimulated echo imaging sequence with an adiabatic 180° RF reversal pulse occurring immediately preceding preparation period 21. Utilizing the preceding longitudinal magnetization reversal, the acquired FID signal and stimulated echo signal are given by the following equation:

[0062] I FID,inv (k)=sin(β)(-E1·cos 2 (α)+(1-E1))·M0

[0063] I STE,inv (k)=-sin(β)·E1·sin 2 (α) / 2·M0

[0064] The following is generated by subtracting the inverted FID and stimulated echo MR images (reconstructed from the first FID and first stimulated echo signals according to the invention) from the corresponding conventional, non-inverted prepared MR images (reconstructed from the second set of FID signals and the second set of stimulated echo signals according to the invention):

[0065] I FID,Δ (k)=I FID (k)-I FID,inv (k)=2sin(β)·E1·cos 2 (α)·M0

[0066] I STE,Δ (k)=I STE (k)-I STE,inv (k)=sin(β)·E1·sin 2 (α)·M0

[0067] Therefore, when these two equations are divided, the relaxation term E1 cancels out. The flipped angle diagram (forming diagram B1) is then given by the following equation:

[0068] α = tan -1 (2I STE,Δ / I FID,Δ )

[0069] Therefore, the equation for calculating the flip angle remains unchanged, but the improved method of the present invention is fully T1 compensated.

[0070] One might assume that the method of the present invention relies on a suitable magnetization reversal capability in one of the two instances of the stimulated echo sequence. However, it has been shown that this is not the case. By assuming that the longitudinal magnetization at the beginning of the first stimulated echo sequence differs from that at the beginning of the second stimulated echo sequence, the above equations can be formulated in a more general manner:

[0071] I FID (k)=sin(β)(E1·m1·cos 2 (α)+(1-E1))·M0

[0072] I STE (k)=sin(β)·E1·m1·sin 2 (α) / 2·M0

[0073] I FID,inv (k)=sin(β)(E1·m2·cos 2 (α)+(1-E1))·M0

[0074] I STE,inv(k)=sin(β)·E1·m2·sin 2 (α) / 2·M0

[0075] Where m1 and m2 indicate the fractions M0 of longitudinal magnetization available at the beginning of the first and second stimulated echo sequences, respectively. Note that in the case of perfect reversal, m1 = 1 and m2 = -1 are applied to this particular notation. Then, subtracting the FID and stimulated echo MR images prepared with different longitudinal magnetizations yields:

[0076] I FID,Δ (k)=sin(β)·(m1-m2)·E1·cos 2 (α)·M0

[0077] I STE,Δ (k)=sin(β)·(m1-m2)·E1·sin 2 (α)·M0 / 2

[0078] When these two equations are divided, the relaxation term E1 cancels out the difference (m1-m2). The flipped-angle diagram (forming the B1 diagram) is again given by the following equation:

[0079] α = tan -1 (2I STE,Δ / I FID,Δ )

[0080] Therefore, it can be concluded that the method of the present invention for determining the B1 diagram for T1 compensation does indeed not depend on proper magnetization reversal. The only prerequisite for eliminating the T1 effect is that the longitudinal magnetization at the beginning of the first stimulated echo sequence is different from the longitudinal magnetization at the beginning of the second stimulated echo sequence.

[0081] However, as can be seen from the formula above, the quality of the inversion affects the signal-to-noise ratio in the resulting B1 plot. When the difference (m1-m2) is at its maximum, the difference signal I... FID,Δ (k) and I STE,Δ (k) Maximum. This corresponds to the case of optimal reversal. For this purpose, for example, an adiabatic secant RF pulse can be used, which can be significantly overdriven to a flip angle of approximately 270° or 360°, thereby ensuring perfect magnetization reversal across the entire imaging volume.

[0082] exist Figure 2aIn one embodiment, a preparation module 24 precedes the first stimulated echo sequence 25, and a preparation module 26 precedes the second stimulated echo sequence 27. The RF pulses of preparation modules 24 and 26 manipulate the longitudinal magnetization at the beginning of the first stimulated echo sequence 25 to differ from the longitudinal magnetization at the beginning of the second stimulated echo sequence 27. This pre-preparation of the longitudinal magnetization (i.e., preparation of the longitudinal magnetization prior to the actual preparation period of the respective stimulated echo imaging sequences 25 and 27) enables the first and second sets of FID signals and / or the first and second sets of stimulated echo signals to have different T1 weights. According to the formula above, this can be used to derive the T1-compensated B1 mapping. Preparation modules 24 and 26 may include saturated RF pulses to adjust the longitudinal magnetization to zero. By applying different delays in modules 24 and 26 between the saturated RF pulse and the start of the corresponding stimulated echo imaging sequences 25 and 27, longitudinal relaxation occurs to different degrees, such that the longitudinal magnetization (m1) at the start of the first stimulated echo sequence 25 is different from the longitudinal magnetization (m2) at the start of the second stimulated echo sequence 27.

[0083] exist Figure 2b In the variant, a second stimulated echo imaging sequence 27 is applied after the first stimulated echo imaging sequence 25 with a delay of 28. The delay 28 is shorter than T1. The short delay between the two stimulated echo imaging sequences 25 and 27 prevents the longitudinal magnetization from fully relaxing to equilibrium, thus achieving a difference between the longitudinal magnetization (m1) at the beginning of the first stimulated echo sequence 25 and the longitudinal magnetization (m2) at the beginning of the second stimulated echo sequence 27.

[0084] Figure 3 The experimental application of the method of the present invention is illustrated. B1 mapping was performed on a phantom (a bottle filled with mineral oil) in a 1.5T MR imaging system. The proposed stimulated echo sequence with T1 compensation using adiabatic inversion pulses was employed (scan matrix = 120 × 90, pixel size = 2.5 mm, slice thickness = 7 mm, TR = 7.6 ms, echo queue duration = 660 ms). Two averaging operations were performed with adiabatic RF pulses enabled and disabled, and imaging pulse phases of 0° and 180°, respectively. The phase cycling of the imaging RF pulses and the averaging of the resulting MR images correspond to subtracting the MR image according to the equation above. Alternatively, the receiver phase may also be switched accordingly. A 3-second delay was used between the two stimulated echo imaging sequences to allow for full T1 recovery in the phantom.

[0085] The B1 map of a single slice with T1 compensation was acquired in steps of 5 degrees for a nominal flip angle α between 5 and 90 degrees. The 10×10 pixel region of interest at the center of the phantom (in...) Figure 3The square marker (represented in the image) was used to average the stimulated echo signal and the FID signal, and the flip angle α was derived from the equation above. Figure 3 The flip angle derived from the nominal angle plot in the graph can be considered an accurate reference. Experiments and analyses were repeated without prior inverted RF pulses for comparison. The data obtained in the graph show that the non-T1 compensated B1 plot begins to degrade above a flip angle of 50°. This is because the oil used in the phantom has a short T1 (approximately 100 ms) and the high-resolution protocol has a long acquisition time, which was intentionally chosen. In contrast, the T1 compensated B1 plot has an accuracy of up to 85°. Therefore, the T1 compensated method covers a much larger portion of the theoretical working range than the non-T1 compensated method.

[0086] Figure 3 The method of the present invention is illustrated in vivo (brain) using a 3T MR imaging system. Transverse slices traversing the brain (intersecting the ventricles) are selected, and B1 maps with and without T1 compensation are acquired. Additionally, combined T1 and T2 compensations are applied using conjugate stimulated echoes. Without T1 compensation (left slice image, designated "normal"), strong T1 contrast is visible between the ventricles, gray matter, and white matter and the head margins. With T1 compensation (middle slice image, designated "T1-comp."), the contrast between the margins, gray matter, and white matter disappears, with only slight contrast remaining in the ventricles. With both T1 and T2 compensations, this small residual contrast disappears (right slice image, designated "T1-,T2-comp."), resulting in a perfectly smooth B1 map. A B1 contour plot along the lines drawn in the slice images shows that T1 weighting of the uncompensated B1 map (contour 41) results in an underestimation of B1 in gray and white matter by approximately 5%. The deviations of contours 42 and 43 are less than 2%, while contour 42 (without T2 compensation) still tends to underestimate B1.

[0087] Figure 3 The example shown illustrates the improved robustness of the inverse-prepared stimulated echo method of the present invention. This may also be beneficial in avoiding user errors when defining appropriate B1 mapping protocols. While reducing the double-echo β-pulse queue length (e.g., by scan segmentation or sacrificing spatial resolution) can reduce the T1 effect by minimizing the FID established during the queue, it cannot completely eliminate the T1 effect. However, the inventors have observed numerous errors in user-defined protocols that can lead to significant anatomical transmission, such as… Figure 3 The left slice image is shown.

[0088] Modifications / extensions to the method of this invention can be conceived. As shown, using magnetization reversal allows for the complete elimination of the T1 effect. However, it is also possible to mitigate these effects when T1 is known. For this purpose, after acquiring the first FID dataset and the stimulated echo dataset, it is possible to acquire a second dataset starting with a different initial magnetization after stimulated echo preparation. For the FID signals of these two echo queues, T1 can be derived by using a best-fit method as a saturation recovery model. This T1 information can then be used as input to apply T1 compensation according to the equations above.

Claims

1. A method for performing MR imaging on at least a portion of an object (10) placed in an examination volume of an MR device (1), the method comprising the steps of: The portion of the object (10) is subjected to a first stimulated echo imaging sequence (25), the first stimulated echo imaging sequence comprising a sequence of at least two preparation RF pulses (α) irradiating the portion of the object (10) during a first preparation period (21) and a readout RF pulse (β) irradiating the portion of the object (10) during a first acquisition period (22) after the first preparation period (21); During the first acquisition period (22), the first group of FID signals (I) were acquired. FID ) and the first group of stimulated echo signals (I STE ); The portion of the object (10) is subjected to a second stimulated echo imaging sequence (27), the second stimulated echo imaging sequence comprising a sequence of at least two preparation RF pulses (α) irradiating the portion of the object (10) during a second preparation period (21) and a readout RF pulse (β) irradiating the portion of the object (10) during a second acquisition period (22) after the second preparation period (21); During the second acquisition period (22), the second group of FID signals (I) were acquired. FID ) and the second group of stimulated echo signals (I STE ), wherein the first group of FID signals (I FID ) and the second group of FID signals (I FID ) have different T1 weights, and / or the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE ) have different T1 weights; Based on the first set of FID signals (I) collected FID ) and the second group of FID signals (I FID ) and the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE To derive a B1 map indicating the spatial distribution of the RF field of the RF pulse within the portion of the object (10), wherein the different T1 weightings are used to compensate for the effects on the B1 map caused by T1 relaxation.

2. The method according to claim 1, wherein, One or more pre-prepared RF pulses precede the first stimulated echo imaging sequence (25) and / or the second stimulated echo imaging sequence (27), the one or more pre-prepared RF pulses manipulating the longitudinal magnetization at the beginning of the first stimulated echo imaging sequence to be different from the longitudinal magnetization at the beginning of the second stimulated echo imaging sequence.

3. The method according to claim 2, wherein, There is an RF inversion pulse before the first preparation period (21) or the second preparation period (21).

4. The method according to claim 3, wherein, The RF inversion pulse is an adiabatic RF inversion pulse.

5. The method according to claim 2, wherein, The one or more pre-prepared RF pulses are saturated RF pulses.

6. The method according to claim 1, wherein, The portion of the object (10) undergoes the second stimulated echo imaging sequence (27) with a delay following the first stimulated echo imaging sequence (25), the delay being shorter than T1.

7. The method according to any one of claims 1-6, wherein, The steps of deriving the B1 diagram involve: by respectively based on the first group of FID signals (I FID ) and the second group of FID signals (I FID The FID difference image is calculated by subtracting the reconstructed MR images, and by calculating the FID difference image based on the first set of stimulated echo signals (I...). STE ) and the second group of stimulated echo signals (I STE The stimulated echo differential image is calculated by subtracting the reconstructed MR image, wherein the B1 image is derived from the voxel intensity ratio of the stimulated echo differential image and the FID differential image.

8. The method according to any one of claims 1-6, wherein, The FID signal (I FID ) and the stimulated echo signal (I STE The signal was collected as a gradient refocusing echo signal.

9. The method according to any one of claims 1-6, wherein, The two stimulated echo signals were acquired after each readout RF pulse (β) during each of the first acquisition period (22) and the second acquisition period (22).

10. The method according to claim 9, wherein, The two stimulated echo signals are the direct stimulated echo signal and the conjugate stimulated echo signal.

11. The method according to claim 9 or 10, wherein, The different T2 weightings of the two stimulated echo signals are used to compensate for the influence on the B1 diagram caused by T2 relaxation.

12. An MR device, comprising: At least one main magnet coil (2) for generating a uniform and stable magnetic field within the examination volume; multiple gradient coils (4, 5, 6) for generating switching magnetic field gradients in different spatial directions within the examination volume; at least one RF coil (9) for generating RF pulses within the examination volume and / or for receiving MR signals from an object (10) positioned in the examination volume; a control unit (15) for controlling the temporal succession of the RF pulses and the switching magnetic field gradients; and a reconstruction unit (17) for reconstructing an MR image based on the received MR signals; wherein the MR device (1) is arranged to perform the following steps: At least a portion of the object (10) is subjected to a first stimulated echo imaging sequence (25), the first stimulated echo imaging sequence comprising a sequence of at least two preparation RF pulses (α) irradiating the portion of the object (10) during a first preparation period (21) and a readout RF pulse (β) irradiating the portion of the object (10) during a first acquisition period (22) after the first preparation period (21); During the first acquisition period (22), the first group of FID signals (I) were acquired. FID ) and the first group of stimulated echo signals (I STE ); The portion of the object (10) is subjected to a second stimulated echo imaging sequence (27), the second stimulated echo imaging sequence comprising a sequence of at least two preparation RF pulses (α) irradiating the portion of the object (10) during a second preparation period (21) and a readout RF pulse (β) irradiating the portion of the object (10) during a second acquisition period (22) after the second preparation period (21); During the second acquisition period (22), the second group of FID signals (I) were acquired. FID ) and the second group of stimulated echo signals (I STE ), wherein the first group of FID signals (I FID ) and the second group of FID signals (I FID ) have different T1 weights, and / or the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE ) have different T1 weights; Based on the first set of FID signals (I) collected FID ) and the second group of FID signals (I FID ) and the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE To derive a B1 map indicating the spatial distribution of the RF field of the RF pulse within the portion of the object (10), wherein the different T1 weightings are used to compensate for the effects on the B1 map caused by T1 relaxation.

13. A computer program to run on an MR device (1), the computer program comprising instructions for performing the following operations: A first stimulated echo imaging sequence (25) is generated, comprising a sequence of at least two preparation RF pulses (α) irradiated during a first preparation period (21) and a readout RF pulse (β) irradiated during a first acquisition period (22) after the first preparation period (21); During the first acquisition period (22), the first group of FID signals (I) were acquired. FID ) and the first group of stimulated echo signals (I STE ); A second stimulated echo imaging sequence (27) is generated, which includes a sequence of at least two preparation RF pulses (α) irradiated during a second preparation period (21) and a readout RF pulse (β) irradiated during a second acquisition period (22) after the second preparation period (21); During the second acquisition period (22), the second group of FID signals (I) were acquired. FID ) and the second group of stimulated echo signals (I STE ), wherein the first group of FID signals (I FID ) and the second group of FID signals (I FID ) have different T1 weights, and / or the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE ) have different T1 weights; Based on the first set of FID signals (I) collected FID ) and the second group of FID signals (I FID ) and the first group of stimulated echo signals (I STE ) and the second group of stimulated echo signals (I STE To derive a B1 map indicating the spatial distribution of the RF field of the RF pulse, wherein the different T1 weights are used to compensate for the effects of T1 relaxation on the B1 map.

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