MR IMAGING WITH T1 COMPENSATED B1 MAPPING

Abstract

The invention relates to a method of MR imaging. It is an object of the invention to provide an improved B.sub.1 mapping method that is less affected by T.sub.1 relaxation. The invention proposes that a first stimulated echo imaging sequence (25) is generated comprising at least two preparation RF pulses (α) radiated during a first preparation period (21) and a sequence of reading RF pulses (β) radiated during a first acquisition period (22) temporally subsequent to the first preparation period (21). A first set of FID signals (I.sub.FID) and a first set of stimulated echo signals (I.sub.STE) are acquired during the first acquisition period (22). A second stimulated echo imaging sequence (27) is generated comprising again at least two preparation RF pulses (α) radiated during a second preparation period (21) and a sequence of reading RF pulses (β) radiated during a second acquisition period (22) temporally subsequent to the second preparation period (21). A second set of FID signals (I.sub.FID) and a second set of stimulated echo signals (I.sub.STE) are acquired during the second acquisition period (22). The first and second sets of FID signals (IFID) have different T.sub.1-weightings and/or the first and second sets of stimulated echo signals (I.sub.STE) have different T.sub.1-weightings. A B.sub.1 map indicating the spatial distribution of the RF field of the RF pulses is derived from the acquired first and second sets of FID (I.sub.FID) and stimulated echo (I.sub.STE) signals, wherein the different T.sub.1-weightings are made use of to compensate for influences on the B.sub.1 map caused by T.sub.1 relaxation. Preferably, either the first or the second preparation period (21) is preceded by an RF inversion pulse to obtain the different T.sub.1-weightings. Moreover, the invention relates to an MR device (1) and to a computer program for an MR device (1).

Claims

1. A method of magnetic resonance (MR) imaging of at least a portion of an object placed in the examination volume of an MR device, the method comprising: subjecting the portion of the object to a first stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated toward the portion of the object during a first preparation period and a sequence of reading RF pulses (β) radiated toward the portion of the object during a first acquisition period temporally subsequent to the first preparation period; acquiring a first set of FID signals (I.sub.FID) and a first set of stimulated echo signals (I.sub.STE) during the first acquisition period; subjecting the portion of the object to a second stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated toward the portion of the object during a second preparation period and a sequence of reading RF pulses (β) radiated toward the portion of the object during an second acquisition period temporally subsequent to the second preparation period; acquiring a second set of FID signals (I.sub.FID) and a second set of stimulated echo signals (I.sub.STE) during the second acquisition period, wherein the first and second sets of FID signals (I.sub.FID) have different T.sub.1-weightings and/or the first and second sets of stimulated echo signals (I.sub.STE) have different T.sub.1-weightings; deriving a B.sub.1 map indicating the spatial distribution of the RF field of the RF pulses within the portion of the body from the acquired first and second sets of FID (I.sub.FID) and stimulated echo (I.sub.STE) signals, wherein the different T.sub.1-weightings are made use of to compensate for influences on the B.sub.1 map caused by T.sub.1 relaxation.

2. The method of claim 1, wherein the first and/or the second stimulated echo sequence is preceded by one or more pre-preparation RF pulses manipulating the longitudinal magnetization at the beginning of the first stimulated echo sequence to differ from the longitudinal magnetization at the beginning of the second stimulated echo sequence.

3. The method of claim 2, wherein the first or the second preparation period is preceded by an RF inversion pulse.

4. The method of claim 3, wherein the RF inversion pulse is an adiabatic RF inversion pulse.

5. The method of claim 2, wherein the one or more pre-preparation RF pulses are saturation RF pulses.

6. The method of claim 1, wherein the portion of the object is subjected to the second stimulated echo imaging sequence with a delay after the first stimulated echo sequence which delay is shorter than T.sub.1.

7. The method of claim 1, wherein the step of deriving the B.sub.1 map involves computing an FID difference image by subtracting MR images reconstructed from the FID signals (I.sub.FID) of the first and second sets respectively as well as computing a stimulated echo difference image by subtracting MR images reconstructed from the stimulated echo signals (I.sub.STE) of the first and second sets respectively, wherein the B.sub.1 map is derived from the voxel-wise intensity ratio of the stimulated echo difference image and the FID difference image.

8. The method of claim 1, wherein the FID (I.sub.FID) and stimulated echo (I.sub.STE) signals are acquired as gradient-recalled echo signals.

9. The method of claim 1, wherein two stimulated echo signals are acquired after each reading RF pulse (β) during each of the first and second acquisition periods.

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

11. The method of claim 9, wherein different T.sub.2-weightings of the two stimulated echo signals are made use of to compensate for influences on the B.sub.1 map caused by T.sub.2 relaxation.

12. A magnetic resonance (MR) device comprising at least one main magnet coil for generating a uniform, steady magnetic field within an examination volume, a number of gradient coils for generating switched 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 an object positioned in the examination volume, a control unit for controlling the temporal succession of RF pulses and switched magnetic field gradients, and a reconstruction unit for reconstructing MR images from the received MR signals, wherein the MR device is arranged to perform a method including: subjecting at least portion of the object to a first stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated toward the portion of the object during a first preparation period and a sequence of reading RF pulses (β) radiated toward the portion of the object during a first acquisition period temporally subsequent to the first preparation period; acquiring a first set of FID signals (I.sub.FID) and a first set of stimulated echo signals (I.sub.STE) during the first acquisition period; subjecting the portion of the object to a second stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated toward the portion of the object during a second preparation period and a sequence of reading RF pulses (β) radiated toward the portion of the object during an second acquisition period temporally subsequent to the second preparation period; acquiring a second set of FID signals (I.sub.FID) and a second set of stimulated echo signals (I.sub.STE) during the second acquisition period, wherein the first and second sets of FID signals (I.sub.FID) have different T.sub.1-weightings and/or the first and second sets of stimulated echo signals (I.sub.STE) have different T.sub.1-weightings; deriving a B.sub.1 map indicating the spatial distribution of the RF field of the RF pulses within the portion of the body from the acquired first and second sets of FID (I.sub.FID) and stimulated echo (I.sub.STE) signals, wherein the different T.sub.1-weightings are made use of to compensate for influences on the B.sub.1 map caused by T.sub.1 relaxation.

13. A computer program to be run on a magnetic resonance (MR) device, which computer program comprises instructions stored on a computer readable medium for: generating a first stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated during a first preparation period and a sequence of reading RF pulses (β) radiated during a first acquisition period temporally subsequent to the first preparation period; acquiring a first set of FID signals (I.sub.FID) and a first set of stimulated echo signals (I.sub.STE) during the first acquisition period; generating a second stimulated echo imaging sequence comprising at least two preparation RF pulses (α) radiated during a second preparation period and a sequence of reading RF pulses (β) radiated during an second acquisition period temporally subsequent to the second preparation period; acquiring a second set of FID signals (I.sub.FID) and a second set of stimulated echo signals (I.sub.STE) during the second acquisition period, wherein the first and second sets of FID signals (I.sub.FID) have different T.sub.1-weightings and/or the first and second sets of stimulated echo signals (I.sub.STE) have different T.sub.1-weightings; deriving a B.sub.1 map indicating the spatial distribution of the RF field of the RF pulses from the acquired first and second sets of FID (I.sub.FID) and stimulated echo (I.sub.STE) signals, wherein the different T.sub.1-weightings are made use of to compensate for influences on the B.sub.1 map caused by T.sub.1 relaxation.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0029] The enclosed drawings disclose preferred embodiments of the present invention. It should be understood, however, that the drawings are designed for the purpose of illustration only and not as a definition of the limits of the invention. In the drawings:

[0030] FIG. 1 schematically shows an MR device for carrying out the methods of the invention;

[0031] FIG. 2 shows a schematic diagram illustrating the imaging sequence according to the invention;

[0032] FIG. 2a schematically illustrates the succession of two stimulated echo imaging sequences in one variant of the invention;

[0033] FIG. 2b schematically illustrates the succession of two stimulated echo imaging sequences in a further variant of the invention;

[0034] FIG. 3 demonstrates the accuracy of the B.sub.1 mapping approach of the invention;

[0035] FIG. 4 illustrates B.sub.1 mapping with and without T.sub.1- and T.sub.2-compensation according to the invention.

DETAILED DESCRIPTION OF THE EMBODIMENTS

[0036] With reference to FIG. 1, an MR device 1 is shown. The device comprises superconducting or resistive main magnet coils 2 such that a substantially uniform, temporally constant main magnetic field B.sub.0 is created along a z-axis through an examination volume. The device further comprises a set of (1.sup.st, 2.sup.nd and—where applicable—3.sup.rd order) shimming coils 2′, wherein the current flow through the individual shimming coils of the set 2′ is controllable for the purpose of minimizing B.sub.0 deviations within the examination volume.

[0037] A magnetic resonance generation and manipulation system applies a series of RF pulses and switched magnetic field gradients to invert or excite nuclear magnetic spins, induce magnetic resonance, refocus magnetic resonance, manipulate magnetic resonance, spatially and otherwise encode the magnetic resonance, saturate spins, and the like to perform MR imaging.

[0038] Most specifically, a gradient pulse amplifier 3 applies current pulses to selected ones of whole-body gradient coils 4, 5 and 6 along x, y and z-axes of the examination volume. A digital RF frequency transmitter 7 transmits RF pulses or pulse packets, via a send-/receive switch 8, to a body RF coil 9 to transmit RF pulses into the examination volume. A typical MR imaging sequence is composed of a packet of RF pulse segments of short duration which taken together with each other and any applied magnetic field gradients achieve a selected manipulation of nuclear magnetic resonance. The RF pulses are used to saturate, excite resonance, invert magnetization, refocus resonance, or manipulate resonance and select a portion of a body 10 positioned in the examination volume. The MR signals are also picked up by the body RF coil 9.

[0039] For generation of MR images of limited regions of the body 10 by means of parallel imaging, a set of local array RF coils 11, 12, 13 are placed contiguous to the region selected for imaging. The array coils 11, 12, 13 can be used to receive MR signals induced by body-coil RF transmissions. In parallel transmit applications, the array RF coils 11, 12, 13 may also be used for RF transmission, for example for the purpose of RF shimming. The resultant MR signals are picked up by the body RF coil 9 and/or by the array RF coils 11, 12, 13 and demodulated by a receiver 14 preferably including a preamplifier (not shown). The receiver 14 is connected to the RF coils 9, 11, 12 and 13 via send-/receive switch 8.

[0040] A host computer 15 controls the current flow through the shimming coils 2′ as well as the gradient pulse amplifier 3 and the transmitter 7 to generate any of a plurality of MR imaging sequences, such as echo planar imaging (EPI), echo volume imaging, gradient and spin echo imaging, fast spin echo imaging, and the like. For the selected sequence, the receiver 14 receives a single or a plurality of MR data lines in rapid succession following each RF excitation pulse. A data acquisition system 16 performs analog-to-digital conversion of the received signals and converts each MR data line to a digital format suitable for further processing. In modem MR devices the data acquisition system 16 is a separate computer which is specialized in acquisition of raw image data.

[0041] Ultimately, the digital raw image data is reconstructed into an image representation by a reconstruction processor 17 which applies a Fourier transform or other appropriate reconstruction algorithms, such like SENSE or SMASH. The MR image may represent a planar slice through the patient, an array of parallel planar slices, a three-dimensional volume, or the like. The image is then stored in an image memory where it may be accessed for converting slices, projections, or other portions of the image representation into appropriate format for visualization, for example via a video monitor 18 which provides a man-readable display of the resultant MR image.

[0042] FIG. 2 shows a diagram illustrating an imaging sequence according to the invention. The depicted imaging sequence is a stimulated echo sequence which is subdivided into a preparation period 21 and an acquisition period 22. Two preparation RF pulses having a flip angle of a are applied during the preparation period 21. The two preparation RF pulses are separated by a time interval T.sub.E. A de-phaser magnetic field gradient G.sub.m2 is applied between the two preparation RF pulses. A sequence of reading RF pulses having flip-angle R are generated during the acquisition period 22, which is temporally subsequent to the preparation period 21. An FID signal I.sub.FID and a stimulated echo signal I.sub.STE are acquired after each reading pulse as gradient-recalled echoes in the presence of refocusing/readout magnetic field gradients G.sub.m1/G.sub.m. Spoiler gradients are applied to suppress undesired residual transverse magnetization. According to the invention, a phase-cycling scheme is implemented, wherein the depicted imaging sequence is played out twice, once with an adiabatic 180° RF inversion pulse preceding the preparation period 21, and once without preceding inversion of the longitudinal magnetization (indicated by 180°/0° in FIG. 2). For clarity, the employed slice-selection and phase-encoding magnetic field gradients are omitted in FIG. 2. The adiabatic 180° RF inversion pulse may be slice selective in the presence of a slice selection gradient (not depicted).

[0043] Without the RF inversion pulse and ignoring T.sub.1 relaxation, the longitudinal magnetization directly after the preparation sequence 21, is given by

M.sub.z,FID=cos.sup.2(α).Math.M.sub.0,

M.sub.z,STE=sin.sup.2(α)/2.Math.M.sub.0,

[0044] wherein M.sub.z,FID and M.sub.z,STE denote the un-prepared (i.e. in-phase) and the stimulated echo-prepared (i.e. de-phased) longitudinal magnetization, respectively. Both the FID signal I.sub.FID generated from M.sub.z,FID and the stimulated echo signal I.sub.STE generated from M.sub.z,STE are acquired at different points in time T.sub.E1 and T.sub.E1+ΔT, respectively. The delay ΔT between the two echoes I.sub.FID, I.sub.STE is determined by the relation:

ΔT=A.sub.mc2/G.sub.m,

[0045] wherein Amc2 denotes the gradient-time area of the de-phaser gradient Gm2 and Gm denotes the strength of the readout magnetic field gradient. Neglecting T1- and T2-effects, the two acquired echo signals IFID and ISTE are given by:

I.sub.FID=sin(β).Math.cos.sup.2(α).Math.M.sub.0,

I.sub.STE=sin(β).Math.sin.sup.2(α)/2.Math.M.sub.0,

[0046] wherein β is the nominal flip angle of the reading RF pulses. Combining the above equations yields:

α=tan.sup.−1(2I.sub.STE/I.sub.FID)

[0047] Thus, the unknown flip angle α of the stimulated echo preparation RF pulses can be derived from the ratio of the acquired echo signals according to:

α=tan.sup.−(2I.sub.STE/I.sub.FID)

[0048] The above equations are an approximation because they do not consider T.sub.1 relaxation. T.sub.1 relaxation results in a recovery of the FID signal and a decay of the stimulated signal according to:

I.sub.FID(k)=sin(β)(E.sub.1.Math.cos.sup.2(α)+(1−E.sub.1)).Math.M.sub.0,

I.sub.STE(k)=sin(β).Math.E.sub.1.Math.sin.sup.2(α)/2.Math.M.sub.0,

with E.sub.1=e.sup.−(k.Math.TR+ΔT)/T.sup.1,

[0049] wherein k and TR denote the index and the repetition time of the gradient echoes in the sequence of reading RF pulses, respectively, and ΔT is the time interval between the preparation period 21 and the first gradient echo. Thus, E.sub.1 denotes the T.sub.1 relaxation term responsible for the recovery of the FID signal and a decay of the stimulated echo signal. Thus, it is readily realized that T.sub.1 relaxation will lead to a systematic underestimation of the flip angle α, especially for short T.sub.1 and/or long echo trains. The error increases with increasing a thus reducing the accurate working range of the method.

[0050] The invention proposes to add a second instance of the stimulated echo imaging sequence with the adiabatic 180° RF inversion pulse being played out immediately before the preparation period 21. With the preceding inversion of the longitudinal magnetization, the acquired FID and stimulated echo signals are given by:

I.sub.FID,inv(k)=sin(β)(−E.sub.1.Math.cos.sup.2(α)+(1−E.sub.1)).Math.M.sub.0,

I.sub.STE,inv(k)=−sin(β).Math.E.sub.1.Math.sin.sup.2(α)/2.Math.M.sub.0.

[0051] Subtracting the inversion-prepared FID and stimulated echo MR images (reconstructed from the first set of FID and stimulated echo signals within the meaning of the invention) from the corresponding conventional, non-inversion-prepared MR images (reconstructed from the second set of FID and stimulated echo signals within the meaning of the invention) yields:

I.sub.FID,Δ(k)=I.sub.FID(k)−I.sub.FID,inv(k)=2 sin(β).Math.E.sub.1.Math.cos.sup.2(α).Math.M.sub.0,

I.sub.STE,Δ(k)=I.sub.STE(k)−I.sub.STE,inv(k)=sin(β).Math.E.sub.1.Math.sin.sup.2(α).Math.M.sub.0.

[0052] Hence, the relaxation term E.sub.1 cancels out, when the two equations are divided. The flip angle map (constituting the B.sub.1 map) is then given by:

α=tan.sup.−1(2I.sub.STE,Δ/I.sub.FIDΔ).

[0053] The equation for computing the flip angle thus remains the same, but the refined approach of the invention is entirely T.sub.1-compensated.

[0054] One may suppose that the approach of the invention relies on proper magnetization inversion capabilities in one of the two instances of the stimulated echo sequence. It turns out, however, that this is not the case. The above equations can be formulated in a more general fashion under the assumption 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:

I.sub.FID(k)=sin(β)(E.sub.1.Math.m.sub.1.Math.cos.sup.2(α)+(1−E.sub.1)).Math.M.sub.0,

I.sub.STE(k)=sin(β).Math.E.sub.1.Math.m.sub.1.Math.sin.sup.2(α)/2.Math.M.sub.0,

I.sub.FID,inv(k)=sin(β)(E.sub.1.Math.m.sub.2.Math.cos.sup.2(α)+(1−E.sub.1)).Math.M.sub.0,

I.sub.STE,inv(k)=sin(β).Math.E.sub.1.Math.m.sub.2.Math.sin.sup.2(α)/2.Math.M.sub.0,

[0055] wherein m.sub.1 and m.sub.2 indicate the fractions of the longitudinal magnetization M.sub.0 available at the beginning of first and second stimulated echo sequences respectively. Note that in case of perfect inversion m.sub.1=1 and m.sub.2=−1 applies in this special notation. Subtracting the differently longitudinal magnetization-prepared FID and stimulated echo MR images then yields:

I.sub.FID,Δ(k)=sin(β).Math.(m.sub.1−m.sub.2).Math.E.sub.1.Math.cos.sup.2(α).Math.M.sub.0,

I.sub.STE,Δ(k)=sin(β).Math.(m.sub.1−m.sub.2).Math.E.sub.1.Math.sin.sup.2(α).Math.M.sub.0/2.

[0056] The relaxation term E.sub.1 and also the difference (m.sub.1−m.sub.2) cancel out, when the two equations are divided. The flip angle map (constituting the B.sub.1 map) is again given by:

α=tan.sup.−1(2I.sub.STE,Δ/I.sub.FID,Δ).

[0057] It can thus be concluded that the approach of the invention for determining a B.sub.1 map that is T.sub.1-compensated does indeed not depend on a proper magnetization inversion. The only prerequisite for eliminating the T.sub.1 influence is that the longitudinal magnetization at the beginning of the first stimulated echo sequence and the longitudinal magnetization at the beginning of the second stimulated echo sequence differ from each other.

[0058] However, as can be seen from the above formulas, the quality of the inversion has an influence on the signal-to-noise ratio in the resulting B.sub.1 map. The difference signals I.sub.FID,Δ(k) and I.sub.STE,Δ(k) are maximum when the difference (m.sub.1−m.sub.2) is maximum. This corresponds to the case of an optimum inversion. To this end, e.g., adiabatic secant RF pulses can be employed which can be significantly overdriven to flip angles of roughly 270° or 3600 degrees ensuring perfect magnetization inversion over the entire imaged volume.

[0059] In the embodiment of FIG. 2a, the first stimulated echo sequence 25 is preceded by a preparation module 24 and the second stimulated echo sequence 27 is preceded by a preparation module 26. RF pulses of the preparation modules 24, 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, 27, achieves that the first and second sets of FID signals and/or the first and second sets of stimulated echo signals have different T.sub.1-weightings. According to the above formulas, this can be made use of for deriving a T.sub.1-compensated B.sub.1 map. The preparation modules 24, 26 may comprise saturation RF pulses to adjust the longitudinal magnetization to zero. By applying different delays in modules 24, 26 between the saturation RF pulses and the beginning of the respective stimulated echo imaging sequence 25, 27, longitudinal relaxation occurs to different extents so that the longitudinal magnetization (m.sub.1) at the beginning of the first stimulated echo sequence 25 differs from the longitudinal magnetization (m.sub.2) at the beginning of the second stimulated echo sequence 27.

[0060] In the variant of FIG. 2b, the second stimulated echo imaging sequence 27 is simply applied with a delay 28 after the first stimulated echo imaging sequence 25. The delay 28 is shorter than T.sub.1. The short delay between the two stimulated echo imaging sequences 25, 27 prevents the longitudinal magnetization from full relaxation to equilibrium, thereby achieving that the longitudinal magnetization (m.sub.1) at the beginning of the first stimulated echo sequence 25 differs from the longitudinal magnetization (m.sub.2) at the beginning of the second stimulated echo sequence 27.

[0061] FIG. 3 illustrates an experimental application of the method of the invention. B.sub.1-mapping was performed on a phantom (a bottle filled with mineral oil) in a 1.5 T MR imaging system. The proposed T.sub.1-compensated stimulated echo sequence with adiabatic inversion pulse was employed (scan matrix=120×90, pixel size=2.5 mm, slice thickness=7 mm, TR=7.6 ms, echo train duration=660 ms). Two averages 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 corresponds to a subtraction of the MR images according to the above equations. Alternatively, it would also be possible to toggle the receiver phase correspondingly. A delay of 3 s was used between the two stimulated echo imaging sequences to allow full T.sub.1 recovery in the phantom.

[0062] Single-slice T.sub.1-compensated B.sub.1 maps where acquired for nominal flip angles α between 5 and 90 degrees in steps of 5 degrees. A 10×10 pixel region of interest in the centre of the phantom (depicted as square marking in FIG. 3) was used to average the stimulated echo and FID signals, and to derive the flip angle α according to the above equation. The derived flip angles are plotted against the nominal angles in the diagram of FIG. 3, which can be regarded as an accurate reference. The experiments and the analysis where repeated without preceding inversion RF pulse for comparison. The resulting data in the diagram show that the non-T.sub.1-compensated B.sub.1 map starts to degrade above a 50° flip angle. This is due to the short T.sub.1 of the oil (about 100 ms) used in the phantom and the long acquisition duration of the high-resolution protocol, which was chosen on purpose. In contrast, the T.sub.1-compensated B.sub.1 maps are accurate up to 85°. Thus, the T.sub.1-compensated method covers a much larger portion of the theoretical working range than the non-T.sub.1-compensated method.

[0063] FIG. 3 shows an in-vivo (brain) application of the method of the invention in a 3 T MR imaging system. A transversal slice through the brain (intersecting the ventricle) was chosen and B.sub.1 maps with and without T.sub.1 compensation were acquired. In addition, also combined T.sub.1- and T.sub.2-compensation was applied by using also the conjugate stimulated echo. Without T.sub.1 compensation (left slice image, designated “normal”), a strong T.sub.1 contrast between ventricle, grey and white matter and the rim of the head is visible. In case of T.sub.1 compensation (middle slice image, designated “T.sub.1-comp.”), the contrast between rim, grey and white matter disappears, and only slight contrast of the ventricle remains. This tiny remaining contrast disappears in case of both T.sub.1 and T.sub.2 compensation (right slice image, designated “T.sub.1-, T.sub.2-comp.”), yielding a perfectly smooth B.sub.1 map. The B.sub.1 profile plots along the lines depicted in the slice images show that the T.sub.1 weighting of non-T.sub.1-compensated B.sub.1 mapping (profile 41) results in an underestimation of B.sub.1 by about 5% in grey and white matter. Profiles 42 and 43 deviate by less than 2%, while profile 42 (without T.sub.2 compensation) still tends to underestimate B.sub.1.

[0064] The example shown in FIG. 3 shows the improved robustness of the inversion prepared stimulated echo approach of the invention. This might also be beneficial to avoid mistakes of users when defining a suitable B.sub.1 mapping protocol. Although the T.sub.1 effect can be reduced, but not fully excluded by reducing the dual echo p-pulse train length (e.g. by scan segmentation or sacrificing spatial resolution), minimizing FID build up during the train, the inventors have seen many mistakes in user defined protocols that lead to significant anatomy shine through as shown in the left slice image of FIG. 3.

[0065] Modifications/extensions of the approach of the invention are conceivable. As shown, the use of the magnetization inversion allows to exclude the T.sub.1 effects completely. However, those could potentially also be mitigated when T.sub.1 would be known. For this purpose, after acquiring the first FID and stimulated echo data set a second data set can be acquired that starts with a different initial magnetization after the stimulated echo preparation. For the FID signals of both echo trains T.sub.1 can be derived by a best fit method using as saturation recovery model. This T.sub.1 information can then serve as an input to apply a T.sub.1 compensation according to the above equations.