Patent Application
20210373099
This invention generally relates to Magnetic Resonance Imaging and spectroscopy. Embodiments may be particularly advantageous in non-uniform magnetic fields.
Magnetic Resonance Imaging (MRI) exploits the nuclear magnetic resonance (NMR) phenomena by combining NMR with gradient magnetic fields to allow cross-sectional slice-selective excitation of nuclei within a subject under examination. In multi-slice imaging, a pulse-sequence of radio-frequency magnetic fields (RF pulse) and associated magnetic field gradients are used with further two dimensional (2D) encoding of the NMR signals to create a 2D image of a portion of the subject. Each slice has an in-slice resolution of around 0.5 mm to 1 mm and slices are spaced around 2 mm apart. A 3D image of the subject is obtained by combining many slices together.
Ideally, in an MR system the RF pulse should deliver a target rotation (a) of the nuclear magnetization vector to provide uniform signal strength over the dimensions of the sample. However, in practice the RF field might typically vary by as much as 50% causing loss of both signal strength and alteration of image contrast by producing magnetization rotations that are far away from the target excitation angle. This variation is typically caused by local magnetic and electrical field effects in the subject, and can lead to spatial inhomogeneity in the local radio-frequency (RF) transverse magnetic field (B1) the nuclei are exposed to. By increasing the static magnetic field strength (B0) an improved signal-to-noise ratio may be obtained along with improved spatial resolution in the images created. However, the above mentioned inhomogeneity in the B1 field is more problematic at B0 fields above 3 T and can lead to imaging artefacts which, in the worst case, are manifested as zero signal in some regions of the image. B1 inhomogeneity effects may also occur at low or medium B0 fields, and when inhomogeneous RF coils such as surface coils are used.
In some cases (e.g. rapid 3D imaging, setting-up patient positioning before a longer relaxation-weighted scan and “freezing” images where the body is in motion) RF pulses are used in rapid MR sequences such as in FLASH and MPRAGE schemes. In FLASH (Fast Low Angle Single Shot) a low spin flip angle)(<90° is combined with rapid repetition of the sequence. In such cases the repetition time (TR) could be between 5 and 50 ms. During the TR interval, the MR signal relaxes back toward equilibrium along the longitudinal (z) axis with an exponential rate constant defined by the T1 value of the tissue being imaged.
With high repetition rates (low TR) a steady-state signal amplitude is quickly formed. The amount of signal measured thus depends on TR and T1. An optimal flip angle, a can be obtained for a particular T1 value, so that the image has 10 to 30% of the S/N ratio and is T1-weighted in its contrast. The optimum signal for a given TR and T1 obtained at an angle α determined by:
cos α=e(−TR/T1)
Such techniques are also susceptible to field inhomogeneities.
Three dimensional MRI also exists. This can be distinguished from multi-slice imaging by the fact that image resolution is the same along all three axes. This property enables any plane orientation to be extracted from a 3D data set and enables surface rendering methods to be used to visualise 3D surfaces of the object, (e.g. the brain surface), in an interactive manner. The MRI method in 3D imaging used does not contain slice selective RF pulses, but excites the whole of the field of view of the RF coil. The third axis of spatial information is encoded using an additional outer loop of incremented phase encoding using the Gz gradient.
However, in 3D MRI a 3-dimensional Fourier Transform is needed to reconstruct the data, and to obtain a 256×256×256 image matrix requires 256×256 experiments. If this were performed with a spin echo sequence with TR=1 second, it would require approximately 20 hours. Therefore, the FLASH sequence may also be used, but without slice selection. Using a TR of 10 milliseconds, the imaging duration is reduced to a more acceptable, 10 minutes.
Nuclear magnetic resonance spectroscopy, similarly tends not be performed in a slice selective manner, but may be used with similar high repetition rate RF pulses.
Reference to any prior art in the specification is not an acknowledgment or suggestion that this prior art forms part of the common general knowledge in any jurisdiction or that this prior art could reasonably be expected to be understood, regarded as relevant, and/or combined with other pieces of prior art by a skilled person in the art.
In order to address at least some of the drawbacks noted above, the present inventors have developed a composite pulse sequence that causes a series of magnetic moment rotations that, in combination, are equivalent to a pulse sequence that would cause a single rotation having a target desired rotation angle α. The composite pulse sequence involves a plurality of pulses which each individually have a desired rotation (A°, B° etc) that is less than the target desired rotation α°. The pulses each cause a rotation about respective axes. The rotation axes are preferably orthogonal to each other. Slice selection magnetic gradients can be employed to make the component rotations of the composite pulse slice selective. Optionally phase correction (re-phasing) gradients can also be included in the pulse sequence.
To avoid doubt, the term “subject” is used in the present specification to mean any biological or non-biological entity which is the subject of the MR investigation. In the illustrative embodiments the subject is described in the context of a human patient or an animal subject. However in other embodiments the subject could be a biological or non-biological sample.
In a third aspect the present invention provides a magnetic resonance system configured to perform a method according to an embodiment of the first or second aspects described above. Such a system may include:
magnetic field producing means for producing a magnetic field (B0);
radio-frequency magnetic field generating means configured to produce radio-frequency magnetic fields (B1a and B1b); and
positioning means for positioning at least part of a subject to be exposed to the effective magnetic field.
In other aspects of the present invention, there are provided magnetic resonance (MR) pulse sequences to be used with a magnetic resonance system. The pulse sequences may be used by any one of the methods disclosed herein.
The appended claims define additional embodiments and aspects of the present invention.
As used herein, except where the context requires otherwise, the term “comprise” and variations of the term, such as “comprising”, “comprises” and “comprised”, are not intended to exclude further additives, components, integers or steps.
Further aspects of the present invention and further embodiments of the aspects described in the preceding paragraphs will become apparent from the following description, given by way of example and with reference to the accompanying drawings.
Illustrative embodiments will now by described by way of example only. The examples described will be adapted for two dimensional MRI and thus are slice selective. Some embodiments described herein can be considered as special cases of the methods described in PCT/AU2016/050068 also in the name of The University of Melbourne, the contents of which are incorporated herein by reference for all purposes.
By way of overview, the illustrative embodiments of the pulse sequence can be used on a FLASH sequence and are described for a fixed repetition time (TR) and spin lattice relaxation time (T1), and include two-component composite RF pulses. Each of the two components of the RF pulse may be chosen to cause separate excitation angles (A° and B°) with the magnetization excited first along different, and preferably transverse planes for each rotation (called herein Mx-transverse plane, and the My-transverse plane). In the preferred embodiments, the ratio of the excitation angles)(A°:B° are chosen to produce two complimentary slice shapes that when added together, produce a slice profile defined within the desired region of space, but preferably with a greater tolerance to the overall RF amplitude in that region of space.
The temporal shape (and hence slice shape profile), and gradient magnitudes, for each of the two pulses in the composite pulse may be different to each other. The exact temporal waveforms would be tailored to the T1 values of the tissues being imaged and the TR values using a simple computer algorithm. The computer algorithm simulates the steady-state signal achieved across the selected slice profile by using rotation matrices that describe the evolution of the sample magnetization in the rotating reference frame as defined by the Bloch equations.
Turning now to the figures,
The magnetic field producing means 20 is configured to produce a static uniform magnetic field B0,s 22 aligned to a longitudinal direction along the z-axis (
The magnetic field gradient producing means 30 is configured to produce a magnetic field gradient G. This can be thought of an additional magnetic field that alters the magnetic field B0,s 22 to produce a modified magnetic field B0. The gradient is not strong enough to vary the direction of the field, so B0 is always parallel with B0,s 22 in the longitudinal axis. Therefore it suffices to define B0 in terms of the component in the longitudinal direction and it is unnecessary to refer to it as a vector quantity. It will therefore be referred to as a scalar quantity B0 without loss of generality. As will be discussed further below, the gradient is used for slice selection, but could be omitted if spectroscopy or three dimensional MRI is being performed.
The radio-frequency (RF) magnetic field generating means 40 is configured to produce transversely oriented RF magnetic fields B1a and B1b, i.e. oriented such that they lie in the x-y plane, that oscillate at a radio-frequency corresponding to the Larmor frequency of a nuclei of interest for MRI (typically protons or carbon-13) exposed to the magnetic field B0. The RF magnetic fields may be linearly or circularly polarised depending on the type of RF magnetic field generating means 40 used and have a phase defined by the operator.
The positioning means 50 is for positioning at least part of a subject 60 in the magnetic field B0.
The system also includes a RF receiver 46, such as RF receiver coils, for receiving an MRI signal. In some embodiments, the RF receiver is part of the RF magnetic field generating means 40. The RF receiver is typically only sensitive to RF magnetic fields oriented in the transverse plane.
In some embodiments, the system 10 includes a control unit 70. Control unit 70 is communicatively coupled with the other components (20, 30, 40, 50) of the system 10. Control unit 70 may include a storage means 72 for storing instructions that determine how the control unit 70 controls the other components (20, 30, 40, 50). Instructions include programs for generating MRI pulse sequences that vary the RF magnetic fields B1 and the magnetic field gradient G to selectively excite nuclei in a cross-sectional slice of the subject exposed to the magnetic field B0. By varying the gradients over two dimensions in k-space, the MRI signals can be spatially encoded to produce a 2D raw image (phase encoding, frequency encoding). Using known Fourier transform MRI techniques, the 2D raw image can be converted or transformed into a 2D image of a cross-sectional slice of the subject. Careful selection of pulse sequence parameters can be used to improve image contrast between various compounds or materials within the subject. By taking many 2D images a 3D image of the subject can be obtained.
The magnetic field producing means 20 may either be controlled by the control unit 70 or it may be persistently producing field B0 (as is usually the case for a superconducting magnet system). The magnetic field producing means 20 and magnetic field gradient producing means 30 may also be in communication with the control unit 70 such that the control unit can monitor their status and/or functionality. For example, the control unit 70 may monitor whether the correct magnetic field strength is being produced, either directly through measuring the proton frequency of the signal from water or indirectly by monitoring an electrical characteristic of the field producing means 20 such as power output.
The subject 60 contains an ensemble of nuclei each with a magnetic moment. When at least a portion of the subject 60 (therefore the ensemble of nuclei within the portion) is exposed to the magnetic field B0 it is considered that, statistically, a greater proportion of the nuclei's magnetic moments become aligned with the magnetic field B0. The time-averaged magnetisation of the portion exposed to the magnetic field B0 is, at equilibrium, described by a net magnetisation vector M, 24 parallel to the direction of the magnetic field B0 (
As will be appreciated by the person skilled in the art, exposure of a subject to a magnetic field is not intended to be limited to mean exposure of a surface of the subject, or the near sub-surface, and is intended to include exposing the nuclei within and throughout the subject to said magnetic field. The use of the term is also intended to include the situation where the MRI system has a persistent magnetic field B0 and the subject is introduced into the field.
Rotation of Magnetisation Vector by RF Magnetic Fields
As is known in the art, a transverse RF magnetic field (B1) that is orthogonal to the main magnetic field B0 is typically used to cause rotation of the net magnetisation M, 24 away from the longitudinal axis (z-axis) so that a component of magnetization is created in the transverse plane. This is necessary for the RF receivers to measure a MRI signal. Typically, in a flash sequence a low angle rotation, say between 10° and 30°, is desired.
As illustrated in
The desired first angle of rotation θ1 can be set by choosing an appropriate combination of duration and amplitude of a pulsed RF magnetic field B1a. As noted above, parts of the subject being scanned may affect the local strength of the RF magnetic fields (B1) at particular locations (spatial inhomogeneity) and cause the corresponding rotation angle at said locations to also be affected. This may result in up to a 50% variation in the actual rotation angle compared to the set angle, i.e. for a desired 15° rotation angle, this could result in an actual rotation between 7° and 23°.
The present inventor has identified that by exposing the subject to a second slice-selective RF magnetic field B1b that is configured to rotate the magnetisation about an orthogonal axis in the rotating reference frame (or in the case of circularly polarised RF magnetic fields, that is 90° out of phase with the first RF field B1a), portions of the subject where the rotation angle deviates from the desired angle α° can be further rotated closer to it. This is further explained in an exemplary embodiment with regard to
As shown in
The second rotation θ2 can be considered as only rotating the residual component of M that remains along the Z axis after the first rotation, i.e Mz 25b. The transverse component Mt 25a is aligned with the y-axis and thus is not displaced by the second rotation.
The second angle θ2 can be selected in the same manner as the first angle. In a preferred embodiment, the second angle θ2 is chosen to match the particular tissue or substance to be imaged. Importantly, the spatial inhomogeneity of the first RF magnetic field does not vary greatly with direction of the applied RF field and therefore will have the same effect on the second RF magnetic field B1b and therefore the corresponding rotation angle.
An advantage of some embodiments is that the resultant rotation caused by the multi-part part rotation is more uniformly close to the desired a rotation over a larger range of non-uniform B1 field conditions, than if only one rotation is performed. In this way, the two part rotation may be seen as being less sensitive to inhomogeneity in the RF magnetic field B 1.
The pulse sequence 300 of
a first rotation generated by first RF pulse 51 (B1a), having an amplitude to cause a desired rotation of A° about the x axis. The first rotation is slice selective and thus includes a corresponding first magnetic field gradient 52;
a second rotation, generated by a second RF pulse 55 (B1b) having an amplitude to cause a desired rotation of B° about the y axis. Again the second rotation is slice selective and thus includes a corresponding second magnetic field gradient 54;
one or more phase adjustments; in this case being, a first re-phasing gradient 53 and a second re-phasing magnetic field gradient 56.
In this example the gradients (52, 54) applied at the time of the B1a and B1b fields have the same amplitude and B1a and B1b overlap in frequencies covered, the same selected slice of the ensemble of nuclei in the subject is excited by both B1a and B1b. As will be seen
In this example the second rotation angle B° y is twice that of the first angle A° x. This could be achieved if the pulse length of B1b is twice that of B1a, or the amplitude of B1b is doubled that of B1a, or a suitable combination of pulse length and amplitude adjustment is used provided that the same slice is selected. In other embodiments, B1a and B1b are either identical or any other desirable ratio. To avoid doubt B° y could be smaller than A° x.
In practical embodiments, the RF magnetic fields are limited in time, commonly referred to as RF pulses. In preferred embodiments, the RF magnetic field is modulated as a time-limited sinc function. This can be considered a sinc function multiplied by a window function such as a Hamming, rectangular function or any known window function. However any shaped pulse could be selected.
This MRI pulse sequence 400 begins with a first radio-frequency magnetic field pulse (51) and a corresponding first magnetic field gradient 52 that are used to excite nuclei within a part of a subject to perform a first slice-selective rotation. As noted above this first radio-frequency magnetic field pulse rotates a net magnetisation vector, about a first axis (e.g. the x axis) such that a portion of the magnetisation now lies in along the y axis. As with the previous example the first slice selection gradient 52 is a magnetic field that has a magnitude that increases along direction that is transverse to the slice being imaged. For convenience this is deemed to be a positive gradient.
Next a second radio-frequency magnetic field pulse (55A) and corresponding second magnetic field gradient 54A is used to cause a second slice-selective rotation. As with the previous embodiment this pulse and slice selection gradient cooperate to rotate the net magnetisation about a second axis (the y axis in this example). Where this embodiment differs from the previous embodiment, is that the second slice selection magnetic field gradient 54A has a negative gradient compared to the first slice selection gradient 52. That is, the magnetic field caused by the second slice selection gradient 54A decreases along the direction in which the first slice selection gradient 52 increases. This means that as well as enabling slice selection, the gradient 54A causes at least partial re-phasing of the magnetisation vectors that were de-phased by the first slice selective rotation process.
As will be appreciated the first and second positive and negative gradients will need to be created so that the slices formed by each gradient are in registration with each other. This may require the second RF pulse to have a negative frequency offset applied to so that the slice centres align along direction of the B0 field. This allows slices offset from the centre of the magnet to be excited.
Finally, the pulse sequence (400) of
All individual pulses (i.e the component pulses of the exemplary composite pulses, and the single α pulse) are sinc pulses.
As can be seen, the signal amplitude for both two component composite pulses varies less over a wide range of RF amplitudes than the single Sinc pulse. The signal strength realised by the composite pulses is postulated to be because of improved slice definition. In the embodiments illustrated, the first pulse in the composite pulse sequence, excites areas with high B1 amplitude (i.e. areas where field inhomogeneity causes a locally high field strength). The second pulse which (in these examples) is stronger (e.g. 2 or 4 times in the examples) targets areas where there is low B1 (i.e. areas where field inhomogeneity causes a locally low field strength) but has a lesser effect on the spins excited by the first pulse (because they have a relatively reduced remaining Mz component). Hence the second pulse can be seen as “filling in” the areas missed (of least affected) by the first pulse.
In order to allow ease of use of the pulse sequences described herein in a FLASH sequence or similar high repetition rate imaging strategy, the inventors have also disclosed a method of determining the operating pulse sequence parameters for use in certain imaging situations.
cos α=e(−TR/T1)
Once the parameters have been determined the MR system can be configured to use these parameters in the conventional manner to perform an imaging sequence using the determined parameters. The method described in connection with the present aspect of the invention can be implemented in a variety of ways, for example it may be implemented in software running on the control unit 70 of the MR system. It could be implemented by a separate computer system and the parameters either manually transferred to the control unit 70, or transferred thereto via a communications network or other data transfer interface.
| Number | Date | Country | Kind |
|---|---|---|---|
| 2016903038 | Aug 2016 | AU | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/AU2017/050810 | 8/2/2017 | WO | 00 |
