The present invention relates to an MRI scanning system and method.

MRI scanning is well known. MRI scanning of a subject body involves the generation of a strong homogenous, uniform, static (DC) magnetic field, B ₀ , within a scanning region in order to align nuclei within the body (when it is placed in the scanning region) in the direction of the B ₀ field, and subsequent irradiation of the scanning region with an RF signal to create a further oscillating (AC) magnetic field, B ₁ , to excite some of the nuclei within the body by causing them to come out of alignment with the static B ₀ field. The B ₁ field is then switched off and the excited nuclei continue to precess around the direction of the B ₀ field. This precessional motion generates its own oscillating magnetic field which is detected and used to determine the location and material of different types of body tissue.

The static field, created by a strong DC magnet, may not be completely homogenous once the subject is placed within it and so shimming is carried out to increase homogeneity. This shimming is carried out by shimming coils which provide further DC magnetic fields which combine with the field from the main magnet to produce the static B ₀ field.

Sensors of a conventional MRI scanning system will detect precession of nuclei of the object being scanned in order to obtain an image of the object. Precession occurs at different frequencies and rates in different materials (for example in water or in fat, both of which are found in body tissue) . Signals from both water and fat areas can be detected.

It can be desirable to attempt to excite nuclei within water only (or within fat only) when taking images, so that the different excitation characteristics existing within different materials does not need to be resolved before a clear image of the object can be formed. In some examples it mat be desirable to excite nuclei exclusively in another material, for example in silicon as in breast implants. Conventionally, the strong magnetic field is adjusted by setting shim currents to make the resonant frequencies of water and fat sufficiently constant across the entire region of interest such that their spectral ranges do not overlap. This condition is often not achieved, particularly for larger regions of interest.

When an object (e.g. a human body) to be scanned is present in the region of interest (scanning region) , the object will have some effect on the both the static and RF magnetic fields - it will disturb these fields.

Aspects of the invention are defined in the independent claims. Preferable features of the invention are defined in the dependent claims, and description. It will be apparent to the skilled person that the specific features mentioned in the description can be used in a broader sense and protection of these is sought in a broad sense.

Advantageously, this invention allows localised disturbances in the magnetic and RF fields due to the presence of the object being scanned to be taken into account (and possibly compensated for) by independently driving each RF signal generator to take into account localised variation in each sub-region.

Optionally, in some embodiments, the shim field is also driven as required in combination with the RF coil drive so that together they achieve the desired excitation properties taking into account the localised disturbances due to the presence of the body in the field (or due to any other field-influencing factor) . Embodiments of the present invention will now be described, by way of example only, with reference to the accompanying drawings, in which:

Figure 1 schematically shows an MRI scanning system according to an embodiment of this invention.

Figure 2 shows a central slice of axial 3D pelvis exam shown for five volunteers, comparing binomial and optimized water-selective excitations. Window and level settings are the same for all images. The difference images show the optimized image subtracted from the binomial one; positive differences in fat indicate improved fat suppression. The improvements are more obvious in larger volunteers (numbers 2, 3 & 5) but in all cases fat suppression quality is improved.

Figure 3 shows slices from volunteer number 5 with distance from central slice indicated. Optimization was performed for the central slice (z = 0mm) . The best image quality is obtained in the central slice after optimisation, however there is an improvement in fat suppression quality in all slices. For z = -50mm the off- resonance effect is severe and so the optimized pulse offers only a partial correction but is better than the binomial pulse.

Figure 4 shows optimized pulse amplitudes and phases compared to standard binomial pulse for volunteer number 5. The traces for the different channels are labeled, with the approximate position of the corresponding RF coil with respect to the magnet bore indicated on the diagram.

Figure 5 shows instantaneous RF field at the centre of each subpulse for volunteer number 5 calculated using measured B ₁ maps and optimized pulses (Fig 4) . The overall amplitude of the RF field approximately follows the 1-3-3- 1 ratio between subpulses but with spatial variation. The phase of the RF field changes through the subpulses with similar spatial variation to the measured off resonance map. Figure 6 shows predicted frequency responses for binomial and optimized pulses in the anterior left hand side of volunteer number five, where there is a strong off-resonance effect (see Fig 2) . The target excitation is only defined in narrow bands of frequencies, as depicted here. The optimized excitation results in reduced flip angle in fat ( ~ -440Hz; from 5° to < 1 °) and increased flip angle in water (OHz; from 7° to 14°) . As well as being shifted with respect to the binomial frequency response, the response of the optimized pulse also has a different shape.

Figure 7 shows (a) Simulated water and fat excitations in subject number 1 for binomial (1-3-3-1) pulse and optimized pulses using different w. The same window settings are used for all water images and all fat images, (b-d) : Plots summarize results of forward simulations for data from all volunteers; all values are normalized to the corresponding values from a 1-3-3-1 binomial pulse and averaged across the volunteer group with error bars depicting variability within the group, (b) Mean square RF drive amplitude (c) Mean flip angle in water band (ΘJ and fat band (θ _f ) . (d) Dynamic range (standard deviation divided by mean) .

Referring to figure 1 , an embodiment of this invention provides an MRI scanning system 100 for providing an image of an object in a region of interest. The object is typically a body of a patient and in the schematic figure 1 , the body is intended to be positioned in scanning region 102. The system comprises a magnet 104 arranged to provide a static magnetic field in the region 102 of interest. The skilled person will understand that figure 1 is purely schematic - in practice, the scanning region is usually between two pole pieces of the magnet or inside a bore of the magnet.

The system also comprises a shimming mechanism 106 comprising a plurality of shimming field generators, in this embodiment shimming coils, arranged to generate a shim magnetic field in the region of interest 102. Similar shimming mechanisms are typically found on existing MRI scanners and are used to try and correct for disruptions/inhomogeneities in the static magnetic field caused by the presence of the body, or by the magnet configuration or by any other factor. In some other embodiments of this invention there may be no shimming mechanism.

The MRI scanning system 100 also includes a plurality, in this embodiment eight, of independent RF field generators. Each RF field generator comprises an RF signal source and amplifier mechanism 110 along with an RF coil element, RFl , RF2, RF3, RF4, RF5, RF6, RF7 and RF8. Each RF coil is arranged to irradiate some or all of the region of interest with a desired RF field made up of one or more RF signals. This is achieved by controlling the currents flowing in each RF coil which oscillate at a desired frequency and are large enough to produce the desired magnetic fields. Each signal comprises one or more pulses and the plurality of desired fields sum to provide an overall desired RF field. The system 100 further comprises a control mechanism 108 arranged to take into account spatial variations in the static magnetic field, such as variations caused by the presence of the body, and thereby to control the amplitude and phase properties of each pulse to control spatially across the image the amount of resonance excitation in a first material, such as water, relative to excitation in a second material, such as fat.

The control mechanism 108 may be a computer arranged to communicate with each RF coil RFl - RF8 and also with the shimming mechanism 106. In this embodiment, each RF coil is also able to function as an RF receiver as well as an RF transmitter. This technology is known in MRI scanning. In other embodiments, separate RF receiver coils may be provided. The control mechanism 108 is arranged to gather data relating to the spatial variation of magnetic field (i.e. both the B ₀ and B ₁ fields) across a region of the body to be imaged, for example across a section of the pelvis, when the body is positioned in place (in a 'ready to scan' position in the scanning region 102. This is achieved by measuring magnetic field across the region of the body whilst the body is in position in this embodiment. In other embodiments, the control mechanism may have access to a data store, such as a database in a memory in a computer, which provides data relating to expected spatial variation across the region of the body.

Once this spatial variation data is available, the control mechanism 108 sets RF pulse amplitude, phase of RF pulses to be transmitted from each RF coil, RFl - RF8, and also, in combination, sets shimming current to be transmitted from the shim coil so that the shimming mechanism 106 will provide a desired effect in combination with the RF coils such that the amount of resonance excitation in water relative to excitation in fat across the region of the body being imaged is controlled accurately. Techniques for setting the degree of control are described in further detail below.

In some embodiments, where the system does not comprise shim coils, the control mechanism 108 controls only the RF pulses.

In some embodiments where the system does comprise shim coils, but the shim coils are used only in the known way for shimming the static magnetic field, the control mechanism 108 controls only the RF pulses and not the shim coils.

Achieving uniform fat-suppression is an important diagnostic requirement for body and musculoskeletal MR examinations. The most commonly used approaches distinguish water from fat based on either differences in frequency or Tl or both. Tl based methods (1) use an inversion pulse followed by a short delay so that fat, which has a short Tl , is suppressed. Such approaches are robust with respect to static B ₀ field inhomogeneity and radiofrequency (RF) B ₁ inhomogeneity if adiabatic pulses are used, but tend to be slow and inefficient in terms of signal to noise ratio (SNR) due to the inversion pulses. A number of different frequency based methods exist which rely on the chemical shift between water and fat resonances (Δf~435Hz at 3T) . At different field strengths Δf is different; Δf is directly proportional to the applied field strength, for example at 7T Δf«1015Hz while at IT Δf«145Hz. Methods include frequency selective saturation of the fat signal (2) , water selective imaging using spectral-spatial (spsp) pulses (3) and signal phase based water-fat separation (4,5) . Methods that rely on selective excitation can be efficient and flexible, but require a homogenous B ₀ field in order to work well which is hard to achieve for large fields of view. The phase based methods can be more robust to B ₀ field variation, but generally require extra data, with information from multiple images acquired at offset echo times being combined to determine and correct for local frequency offsets.

Spectral spatial (spsp) RF pulses are commonly used to excite only protons from water or fat within a spatially local region, typically a slice or slab. Slice selective spsp pulses consist of trains of slice selective subpulses with different amplitudes; a commonly used example is the binomial excitation pulse (6) where the relative amplitudes are given by binomial coefficients. Pulses of this type can be described straightforwardly using the small tip angle approximation (STA) (3,7) which establishes a Fourier relationship between the resulting transverse magnetization (m(x)) and the applied RF pulse (b(t)) :

In this formalism x = (x,y,z, ω) and so m(x) is the resulting transverse magnetization as a function of space and (angular) frequency. The transmit sensitivity of the RF coil and static field inhomogeneity are represented by S(x) and B ₀ (x) respectively and γ is the Gyromagnetic ratio. Fourier transform variable k(t) = (k _x ,k _y ,k _z ,k _ω ) has components corresponding to both spatial and temporal frequencies; the spatial components are given by integrals of the

T applied magnetic field gradients ^ (t) = - _γ ΪG (t' ) _d t' where for example G _x is

the x gradient strength and the temporal term is given by k _ω( t) = t -T where T is the total pulse duration. Simultaneous slice and frequency selection (i.e. selection in z-ω) are achieved by applying RF energy while traversing a trajectory in k _z -k _ω (3) . For the slice selective spsp pulses used, the k-space trajectory is such that the spatial and spectral responses are independent to first approximation, as the subpulse duration is much shorter than the interval between them. In this case, spatial selection is determined from the individual subpulses whilst spectral selectivity comes from the pattern of amplitude and phase variation of the pulse train (8, 9) . A binomial pulse train with N subpulses separated in time by period τ has frequency response m( _ω ) = giving strong fat suppression if τ = l/(2Δf) = 1.15ms at 3T. Use of higher N provides broader stop bands at the expense of longer RF pulse durations. Inhomogeneity of the static magnetic field causes the water and fat resonant frequencies to vary in space so that they may move out of the frequency bands defined by the RF pulse's structure. The result can be poor fat suppression or even fat selection with water suppressed in regions where the effect is severe. Pulses with a broader stop band are less vulnerable to this effect, and for this reason at 3T the N = 4 binomial sequence (1-3-3-1) is commonly used, however failure of fat suppression is still often seen. This clinically used N = 4 binomial pulse has been used as an exemplar below.

The recent advent of parallel transmission (the use of a plurality of RF coils to produce a desired overall RF field) has provided more direct control of the RF field in space as well as time; we can describe these extra degrees of freedom by modifying Eq. [1] (10) :

m(x) = />£ S ₀ (x)| b _c (t)e ^ιx ^e'^'-^dt (2)

where S _c and b _c (t) are the spatial transmit sensitivity and RF pulse waveform respectively of coil c and N _c is the number of coils. Much interest has been placed on using these degrees of freedom to improve spatial homogeneity of excitation, particularly at high field strengths (11 ,12) . RF shimming (13) is a straightforward method of achieving this by adjusting the relative amplitudes and phases of each channel such that the overall RF field is more uniform. Alternatively the entire waveform and k-space trajectory can be redesigned to gain more control. In this vein there is a body of work exploring use of composite RF pulses (referred to as 'fast k _z ' or 'spokes' (14,15)) in which subpulses played out along a k _z trajectory are offset in k _x -k _y in order to achieve modulation of in-plane excitation. Recently it has been recognized that spectral properties of these pulses can be optimized in order to improve wideband uniformity (16-18) . There is thus a close link with conventional water selective binomial pulses in which each subpulse can be seen as a spoke at (k _x ,k _y ) = 0. In this invention the potential for using extra degrees of freedom available from parallel transmission to improve performance of binomial style pulses by correcting for B ₀ inhomogeneity is explored. Fat suppressed images of the pelvis acquired in healthy volunteers are presented, demonstrating the application of this method to a large field of view (FOV) .

Methods RF pulse design

The inventors have used the image domain STA formalism for multiple transmitter systems proposed in (10) , in which Eq. [2] is discretized in space and time and written in matrix form m(x) = A b(k) . The x and k variables include the spectral response as outlined in the introduction. For a pulse train of N subpulses, the k-space trajectory consists of N points k = (0,0,0,k _ω ) where k _ω = (n-N)τ with n = [1 ,N] so the RF pulse train b(k) is described by N complex values for each of the N _c coils. These were calculated using Magnitude Least Squares (MLS) optimization (19,20) which performs the minimization b = arg _b min{ I | | A b | - m | | _w ² + λ | | b | | ² } where m is a real valued target magnetization, W represents relative weights as a function of space/frequency and λ is a regularization parameter that can be used to control total RF power. Ignoring the phase of the target allows more degrees of freedom for improving the fidelity of the magnitude response. The implementation used follows the local variable exchange method in (19) initialized using a random target phase; no phase smoothness constraints (as in (20)) were used.

The nature of the solution depends on the way in which the problem is posed. Since we are attempting to correct the binomial excitation for off-resonance effects, we might choose to define the target excitation as the binomial frequency response with no off-resonance effect. This has a known analytic form and is symmetric about ω= 0; however, neither of these properties are necessary for good fat suppression and as an optimization target they are overly restrictive. For water selective imaging, only the responses at the water and fat frequencies are important. In keeping with this the target was defined only within narrow frequency bands centred on OHz ("water band") and -435Hz ("fat band") , with the goals of uniform flip angle in the water band and zero in the fat band for all spatial locations. The width of these bands was selected to be 40Hz to accommodate the expected local line widths of the water and fat resonances. Optimal pulse weightings were calculated in a single voxel STA model using this target with no off-resonance effect. The resulting normalized pulse amplitudes and phase offsets are given in table 1.

Optimized Pulses from Single Voxe L Model

N Relative pulse amplitude Relative phase (degrees) Binomial

2 1 .0 1 0 0 0 1 1

3 1.0 2 0 1.0 0 0 0 1 2 1

4 1.0 3 .0 3 0 1.0 0 0 0 0 1 3 3 1

5 1.0 10.3 22 .4 10.3 1.0 179 0 0 0 -179 1 4 6 4 1

1.0 1.6 9. 8 9.8 1.64 1 5 10 10

6 179 0 0 0 0 -179 1 0 5 1

Table 1 Relative pulse amplitude and phase computed from single voxel optimization with no off-resonance compared with equivalent binomial pulses. For N<4 the solutions reproduced binomial weightings. This implies that for the N = 4 pulses presented in this paper, it is reasonable to use the 1-3-3-1 sequence as an unoptimized comparison.

Frequencies outside of the water and fat bands are left unconstrained by excluding them from x; the same approach is taken for spatial locations outside of the object. The two optimization goals - flip angle uniformity for water and zero fat excitation - compete with one another. Since small failures in fat suppression are diagnostically more significant than small deviations of water flip angle, we set the requirement for fat suppression to be stronger than that of water signal uniformity. This was achieved by reducing the relative weight of the water band in W. Experience gained from pilot data suggested that a relative weighting of w = 10 ³ for the water band gives an effective balance between the two goals, and this value was used for all of the presented in vivo experiments. The effect of this parameter was retrospectively investigated using forward simulations from the calibration data obtained for this study; the optimization was carried out for different values of w with the predicted excitations and required RF power evaluated.

All experiments used λ = 16 which was found to yield an effective trade off between RF power and excitation fidelity. Computation grid resolutions were 8mm in space and 20Hz in frequency with optimization performed using Matlab R2008a (Mathworks, Natick, MA) on an IBM x3755 system (four dual core processors, 32 GB RAM) linked via Ethernet connection to the scanner console.

In vivo experiments

Experiments were performed using a 3T Achieva MRI system (Philips Healthcare, the Netherlands) equipped with an 8-channel body coil capable of parallel transmission (21) . The coil consists of strip elements arranged around the bore of the magnet, giving maximum variation in transmit sensitivity in the axial plane. When simulating a single body coil, the elements are driven with a fixed pre-calibrated amplitude and phase relationship designed to give quadrature excitation for a wide FOV (referred to as 'quadrature mode') - A 6- element phased array surface coil was used for signal reception. Research Ethics Committee approval was obtained for the study and all participants gave written informed consent prior to enrolment. In total 5 healthy volunteers (2 male, 3 female) underwent fat suppressed pelvic imaging.

A slab selective spsp pulse with N = 4 subpulses and slab thickness 200 mm was used for all experiments. Flyback gradients were used between subpulses in order to avoid side lobes occurring at the gradient oscillation frequency from appearing inside the intended stop band (8,22) . Gradient waveforms were designed with maximum slew rate 30 T/m/s. The sine subpulses had time- bandwidth-product of 8 and time between them was set to τ = l/(2Δf) = 1.15ms, as is generally used in clinical imaging. Experiments compared optimized pulses using parallel transmission with a 1-3-3-1 binomial sequence using the same trajectory but with the coil driven in quadrature mode. Optimized pulses were designed to give a target flip angle of θ _w = 20° in water and θ _f = 0° in fat; binomial pulse amplitudes were set using the scanner's standard power optimization method.

Imaging was performed using a standard abdominal protocol, 3D Fourier encoded RF spoiled gradient echo sequence (TlFFE) with flip angle 20° , TR = 30 ms and TE = 4.6 ms. Acquired images were in axial orientation with a 400mm FOV and resolution 2 x 2 x 5 mm ³ . Standard first order B ₀ shimming was performed using scanner preparation phases, the same shim settings and centre frequency were used for all experiments on a given subject.

The scanner software estimated the SAR load of the sequence incorporating the 1-3-3-1 pulse to be ~ 6% of the local torso limit (10 WVKg) depending on patient weight. Scanner SAR estimates are based on mean square RF drive amplitudes; in the case of transmitting different pulses on each channel such estimates are inaccurate as they do not take into account the effect on the electric field (E-field) of changing phase relationships between channels. As a conservative safety margin the operating procedure adopted was to ensure that the local SAR estimate produced by the scanner did not exceed 10% of the local torso limit. Time-resolved B ₁ field estimates were also produced by combining RF pulse waveforms with the B ₁ maps and peak / mean square values were computed by integrating these through space and time.

Calibration Data Although imaging was performed over a 200mm slab, optimization was only performed for a single slice at the centre of this slab. The effect of the optimization outside the slice of interest was also evaluated. B ₁ mapping was performed using the actual flip angle imaging (AFI) sequence (23) with a slice selective pulse, implemented with modifications proposed in (24) using the "all but one" array mapping method from (25) . AFI used TR ₁ = SOmS, TR ₂ = 150ms, TE = 4.6 ms (in phase) , nominal flip angle 80° with resolution 4 x 4 x 10 mm ³ and FOV varying depending on the size of the subject; acquisition time was approximately lm50s for all 8 channels. B ₀ mapping was performed using a multiecho gradient echo sequence with all echoes in phase (3 echoes, first TE = 2.3ms, ΔTE = 2.3ms) to avoid interference between water and fat; TR = 30 ms, flip angle 20° . Flyback gradients were used to minimize errors from eddy currents. Resolution and field of view settings were the same as the B ₁ maps, acquisition time was approximately 35s (16 signal averages) . The relatively large interecho spacing leads to phase aliasing in the acquired B ₀ map, this was corrected in postprocessing using spatial unwrapping.

Results

Figure 2 contains images from all five volunteers, comparing the standard binomial and optimized excitations. The images from binomial pulse excitation show varying degrees of fat suppression failure depending on the quality of the original shim. The optimized pulses produced universal improvement of fat suppression quality. Although optimization was performed only for a single 10 mm slice, it was observed that fat suppression uniformity improved throughout the 200 mm imaged volume in all volunteers. Figure 3 shows exemplary slices from one subject (volunteer number five) over range ± 50mm out of the central slice. The uniformity of fat suppression is improved for all of these slices; importantly overall image quality is not worsened in other slices by the optimization. More quantitative comparisons can be made by considering the predicted excitation properties from the STA model in each case. For a target flip angle of 20° in water, the mean flip angle in fat as predicted from the model for the binomial pulse was 0.72° ± 0.55° (mean and standard deviation calculated across the volunteer group) . This fell to 0.12° ± 0.04° for subject specific pulses, indicating a six-fold improvement in fat suppression. The dynamic range of flip angles in the water band was quantified by taking the standard deviation divided by the mean flip angle and fell from 0.26 ± 0.05 before optimization to 0.16 ± 0.05 afterwards.

The mean square RF drive amplitude (often taken as a surrogate measure for the SAR) was reduced by 13% ± 6% across the volunteer group, while peak drive amplitude on a single channel increased by 36% ± 9%. The B ₁ map based RF field model predicts that mean square B ₁ was reduced by 17% ± 9% across the volunteers with the peak value (in space and time) showing no significant change (change of -5% ± 10%) .

The optimized RF subpulse amplitudes and phases for volunteer five are plotted in Figure 4. The amplitudes of the middle two subpulses are higher than the others for all channels, but do not follow the 1-3-3-1 pattern. For most channels the phase changes monotonically in time, suggesting a frequency offset. Figure

5 shows the predicted B ₁ field at the middle of each subpulse. The total field amplitudes more closely follow a 1-3-3-1 pattern but there is spatial variation. Figure 6 shows the predicted frequency responses from the standard and optimized pulses in a voxel on the anterior left hand side of volunteer five, where there is a failure of fat suppression with the binomial pulse visible on

Figs 2 and 3. The target excitation is only defined in narrow bands, and the optimized excitation is not constrained outside of these. As expected there is a frequency shift in the response, however it also has a different shape to the binomial version, losing the symmetry that the latter possesses.

Relative weighting of water vs fat Figure 7a shows the predicted excitations in water and fat using calibration data from volunteer number one for different values of w and Figs 7b-d shows plots that summarize results for all five volunteers. Using w < 10 ⁵ leads to the trivial solution of no excitation in water or fat, satisfying the optimization goal for the fat frequency but not for water. Increasing this weighting improves the properties of the water excitation but at the same time increases the mean flip angle in fat to the point where w = 1 results in a mean fat flip angle that for some volunteers is greater than that from the unoptimized binomial pulse. While there is some variability between the volunteers, the overall trends are the same and the choice of iv = 10 ³ gave both a large improvement in fat suppression and reduced dynamic range of flip angles in water.

Discussion The inventors have optimized commonly used spectral spatial RF pulses for use with parallel transmission and have demonstrated improved fat suppression quality at 3T over a large field of view. An improvement in image quality was obtained in all subjects (Figure 2) and is corroborated by predictions from the STA model.

The optimization used a relative weighting of water with respect to fat of w = 10 ³ and simulations (Fig 7) confirmed that this was a reasonable choice for this study and that results were similar for all of the volunteers. In general it is likely that the optimal choice of this parameter would be different depending on the geometry of the object and the coil. The fact that the optimal choice is stable for the fixed geometry used here, with a range of volunteer sizes suggests that clinical use of this method could perhaps proceed using predetermined optimal values for given examinations.

The pulse design algorithm was tailored to maximize the effectiveness of the available degrees of freedom for fat suppression. At a given location in space, the frequency response of a spsp pulse with N subpulses is determined by the instantaneous RF field at the centre of each subpulse. This in general can be any spectral shape determined by N parameters. By only defining the target excitation in narrow frequency bands around the water and fat resonances, we implicitly state that the symmetry produced by a binomial pulse is not required. In the absence of field imperfections, single voxel optimization (Table 1) shows that for N<4 the binomial sequence is still optimal. For N>5 the methods diverge because higher orders of the binomial sequence give a wider stop band but a narrower pass band; the solutions found by the presented approach instead aim to generate pass and stop bands with flat regions of 40 Hz. These pulses are still however symmetric in time (and in frequency) . Moving to the more demanding problem of optimization over spatially varying fields with the relatively small set of parameters (N x N _c ) results in this symmetry disappearing, as the example response in Fig 6 demonstrates.

Links to frequency correction

It has been proposed that the extra degrees of freedom afforded by parallel transmission may be used to improve performance by individually tuning the centre frequency of each element to match the local resonance frequency (26) . This on its own however does not take into account interactions between the coil elements; neglecting this aspect results in unpredictable performance that can compromise image quality severely. This invention provides a spatially varying RF phase profile which progresses through time (Fig 5) ; in the limit that the subpulses have zero duration a linear phase progression is equivalent to a spatially varying RF frequency distribution. As Fig 5 shows, this phase progression is correlated with the strength and sign of the off-resonance field but it does not follow a strict linear form, demonstrating that correction applied is more complex than localized frequency correction alone.

Limitations and further degrees of freedom

The design of the RF body coil used gives maximum sensitivity variation in the axial plane used for pelvic imaging, with slow variation through slice. Images for a single volunteer at different slice locations (Fig 3) demonstrate that for this configuration the optimization results in better fat suppression also outside the optimized slice. Less favourable coil and/or slice geometries may require optimization to be performed through multiple slices however extra calibration data would be required to achieve this. In this work we focused on the extra control afforded by parallel transmission. Where there are highly local or large B ₀ offsets this may fail, an example of this can be seen in the images from volunteers two and three (Fig 3) where fat suppression failure in the right hand side of the subject has been reduced but not completely eradicated. The close connection between slice selective spsp pulses and the more general spokes method was mentioned in the introduction, and in order to tackle these more difficult situations, it is likely that moving to a solution where the k-space trajectory can explore the k _x -k _y plane will introduce the extra control required. Indeed preliminary work suggests that this is also the case for single channel transmission. If optimizing the k-space trajectory, then the timing of the subpulses could also be considered. In general we use the minimum τ required to resolve frequency difference Δf; τ = l/2Δf since this minimizes pulse duration. If this can be relaxed then τ > l/2Δf gives a wider bandwidth and therefore more control over the frequency response. Alternatively shorter pulses can be achieved by using τ < l/2Δf which results in aliasing that can be shifted using phase modulation (27,28) at the expense of reducing the efficiency of water excitation.

B ₀ and B ₁ homogeneity are usually controlled using different sub-systems; B ₀ using the gradients and dedicated shim coils and B ₁ more recently using parallel RF transmitters. In this work, a standard linear B ₀ shimming approach was used. The quality of the B ₀ shim might have been improved in some cases by using a different algorithm to calculate the required offsets or by using higher order shimming, this would then impact on the fat suppression quality achieved using the binomial pulse. We have demonstrated that the RF system can be used in a complementary manner which is true regardless of the starting point. The ability to create a spatially varying RF phase progression to counteract off- resonance comes directly from extra degrees of freedom offered by parallel transmission. A still more integrated approach would be to optimize B ₀ shim settings and the RF pulse simultaneously, allowing full advantage to be taken from all degrees of freedom. SAR

The ability of parallel transmission to create a spatially and temporally varying RF field leads to concerns over the nature of the electric fields involved and their impact on RF power deposition. All experiments led to a reduced mean square RF drive amplitude (mean reduction 13%) , which suggests a reduction of the overall power deposition. The peak drive amplitude however increased in all cases (mean increase 36%) . The peak B ₁ field (in space and time) estimated from the calibration data did not however increase by this amount suggesting that cancellation of fields occurs. In this work the RF power is controlled by the scalar parameter λ which penalizes solutions with high mean square RF drive. If required this could be replaced by a diagonal matrix containing independent regularization parameters for each of the RF subpulses / coils (as suggested in (10)) which can be used to avoid large peak amplitudes for coils with low sensitivity in the region of interest. As demonstrated by Fig 7b, there is also a trade-off between water excitation fidelity and mean square RF drive mediated by choice of w; for example using w = 10 ² rather than 10 ³ as used in this work would result in better water flip angle homogeneity but also higher RF power.

Conclusions

Using the extra degrees of freedom available from an eight channel parallel transmission system, the inventors demonstrate improved fat suppression quality over a large field of view at 3T. Results from a pelvis imaging study with five healthy volunteers showed an improvement in image quality in all cases. Spectral spatial RF pulses were designed online with necessary B ₀ and B ₁ field data acquired within 3 minutes and pulse calculation taking approx 15s making the approach highly suitable for future adoption in a clinical setting when parallel transmit technology becomes more widely available.

In prior systems, a static magnetic field, B ₀ is made as homogenous as possible before placing a subject in the field. As discussed above, placing the subject in the field results in disruption of the field which can be at least partially corrected using shimming coils. A transmission RF field, B ₁ , is applied, whose variation across the subject being scanned, i.e. across a slice of the subject, such as across the pelvis of the subject, is not controlled.

As also previously discussed, this invention recognises that the presence of the subject in the magnetic field causes local disruptions in the field. The invention uses measurements of these disruptions or inhomogeneities in order to individually set or control the RF pulses emitted. For example, in prior systems simple binomial pulses are used (e.g 1-3-3-1) . In this invention, each individual pulse is tailored to excite different frequencies across the subject, e.g. across the pelvis. The invention recognises that across the pelvis the level of the magnetic field varies (see for example Figure 5 - although the plot actually shows frequency of excitation, this is proportional to the magnetic field level) due to the presence of the subject. A controller mechanism according to this invention takes into account this variation in order to produce bespoke pulses within each signal (i.e. not simply 1-3-3-1) to specifically excite resonance in one material (in this case water) and to specifically suppress resonance in another material ( in this case fat) across each part of the subject. Prior systems do not do this - instead, as the field varies across the subject, the RF signals vary in their effectiveness in promoting/suppressing excitation in different materials since they are not taking into account the influence of the specific subject on the local magnetic field.

In some embodiments, the magnetic spatial variation due to the subject is measured on each occasion before an MRI scan in order to calibrate the scanning system. For example, a measurement system comprises one or more measurement RF coils in addition to the magnetic field transmission apparatus. These measurement system RF coils are the same coils used for normal MRI scan measurement in this embodiment.

In further embodiments, to save time, the scanning system may have access to an information store which provides an indication of expected magnetic spatial variation for a particular subject, e.g. depending upon their size, shape, orientation (whether they are being put head first or feet first into the scanner body) , body fat index etc. In yet further embodiments, the scanning system may be able to have both measurement of spatial variation due to the subject functionality as well as the ability to look up this information. The control mechanism may be intelligent in order to assess whether look up is suitable for a particular object or whether measurement is suitable - e.g. based on extraordinary circumstances, such as extraordinarily small/large size.

As described previously calculating optimal phase and amplitude of RF pulses for all coil simultaneously can be advantageous, e.g. relative to varying frequency independently for each coil.

In another embodiment, the invention uses a control mechanism which is arranged to control the level of the DC shim magnetic field in combination with the phase and amplitude of each RF pulse in order to produce the desired spatial resonance effects. In some such embodiments, the shim field may be used to produce a static magnetic field which is intentionally not completely homogenous, but which can be corrected locally by the controlled RF pulses. This places an extra information processing burden on the control mechanism but also offers it more degrees of freedom and is therefore potentially a powerful tool in obtaining accurate promotion of resonance in water and suppression in fat.

In general, it is more important to suppress resonance excitation in fat than to promote it in water. In some embodiments this may change and it may become particularly important to promote water excitation even if higher than normal fat excitation is introduced. In such embodiments, it will be clear to the skilled person how to adjust the control mechanism parameters in view of the description above, especially with reference to figure 6.

The spatial variation in magnetic field for which this invention provides correction may be due to the presence of the object (e.g. a body) in the MRI scanner, or it may be due to the magnet configuration itself, or it may be a combination of both of these factors, or it may be influenced by any other factor. In any event the present invention provides a system and method for taking this spatial variation into account and controlling RF pulses, and optionally additionally shimming field, accordingly to provide a desired level of excitation in one material or suppression in another material, or a combination of both.

The RF field generated in the region of interest comprises a complex field which is the sum of the individual RF fields generated by each RF field generator.

Variations and modifications within the scope of the claims will be apparent to the skilled person.

In some embodiments of this invention the RF pulse properties are set such that resonance is excited in water only. In other embodiments, it may be in fat only, or in any other suitable material or tissue.

In some embodiments, the pulses are calculated to excite resonance in one material (e.g. water at 0 Hz) , and suppress excitation in another material (e.g. fat at -435Hz) , with no explicit control in other spectral regions, especially between these frequencies. This is because in such embodiments, the invention is not concerned with the effects outside the frequency regions in which resonance signals are produced.

An embodiment using a 4 pulse signal has been described. The invention can be used with any other suitable signal format. For example using a 3 pulse signal, such as a 3 pulse binomial signal (1-2-1) .

In some embodiments, the RF transmitter coils are used as RF receiver coils, and in other embodiments they are distinct from each other.

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List of symbols lowercase Greek "lambda" w Italic lowercase w lowercase Greek "omega" lowercase Greek "gamma" lowercase Greek "tau" k _x ,k _y ,k _z "kay" with subscript x,y,z k "kay" with subscript lowercase Greek "omega" B ₀ capital B, subscript zero

N _c capital N, subscript lower case c

In equations 1&2, vectors are written using an arrow over the character (in these equations, x and k are written in this way) whereas in the text the same characters are written using bold lowercase. It was not possible to use bold in our equation typesetting software, but the equations should be written using bold rather than arrow notation - the terminology is interchangeable.