MRI METHOD FOR BO-MAPPING

Abstract

A B.sub.0-mapping method determines the spatial distribution of a static magnetic field in a pre-selected imaging zone comprising computation of the spatial distribution of a static magnetic field from a spatial distribution of spin-phase accruals between magnetic resonance echo signals from the imaging zone and an estimate of the proton density distribution in the imaging zone. The invention provides the field estimate also in cavities and outside tissue. Also the field estimate of the invention suffers less from so-called phase-wraps.

Claims

1. A B.sub.0-mapping method for determining the spatial distribution of a static magnetic field in a pre-selected imaging zone, the method comprising: computation of the spatial distribution of a static magnetic field from a spatial distribution of spin-phase accruals between magnetic resonance echo signals from the imaging zone and an estimate of the proton density distribution in the imaging zone.

2. The B.sub.0-mapping method of claim 1, further including; computation of a phase-estimate magnetic susceptibility distribution that is consistent with the spin-phase accrual distribution, computation of a proton-estimate magnetic susceptibility distribution that is consistent the estimated proton spin density distribution, fit a final magnetic susceptibility distribution so as to minimise the differences both (i) between the final magnetic susceptibility distribution and the phase-estimate magnetic susceptibility distribution and (ii) between the final magnetic susceptibility distribution and the proton-estimate magnetic susceptibility distribution, compute the spatial distribution of the static magnetic field from the final magnetic susceptibility distribution.

3. The B.sub.0-mapping method of claim 2, wherein the final magnetic susceptibility distribution is computed in an iterative manner, in which iteration is between constraints of minimal difference between the final magnetic susceptibility distribution and the phase-estimate magnetic susceptibly distribution and minimal difference between the final magnetic susceptibility distribution and the proton-estimate magnetic susceptibility distribution.

4. The B.sub.0-mapping method of claim 1, wherein the estimated proton density distribution is formed by a segmentation of at least three components.

5. The B.sub.0-mapping method of claim 4, in which the segmentation involves at least components representing soft-tissue, interstitial voids and air, respectively.

6. The B.sub.0-mapping method of claim 5, in which the segmentation further involves components representing silicone, metal and ceramic, respectively.

7. The B.sub.0-mapping method of claim 1, wherein the estimated proton density distribution is formed by a segmentation of at most three components preferably at most the components representing soft-tissue, interstitial voids and air, respectively.

8. The B.sub.0-mapping method as claimed in of claim 3, wherein the iterative procedure is initialised from an initial estimate of the spatial distribution of the static magnetic field and preferably the accuracy of the spatial distribution of the static magnetic field.

9. The B.sub.0-mapping of claim 3, wherein a self-consistent minimisation procedure minimises the differences both (i) between the final magnetic susceptibility distribution and the phase-estimate magnetic susceptibility distribution and (ii) between the final magnetic susceptibility distribution and the proton-estimate magnetic susceptibility distribution.

10. A computer program including executable instructions stored on a non-transitory medium for determining the spatial distribution of a static magnetic field in a pre-selected imaging zone including instruction to computation of the spatial distribution of a static magnetic field from a spatial distribution of spin-phase accruals between magnetic resonance echo signals from the imaging zone and an estimate of the proton density distribution in the imaging zone.

11. A magnetic resonance examination system configured to execute the computer program of claim 10.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0029] FIG. 1 shows a schematic representation of an implementation of a computation of the spatial distribution of a static magnetic field according to the invention;

[0030] FIG. 2 shows a more detailed schematic representation of an implementation of a computation of the spatial distribution of a static magnetic field according to the invention and

[0031] FIG. 3 shows a diagrammatic representation of an magnetic resonance examination system in which the invention is incorporated.

DETAILED DESCRIPTION OF THE EMBODIMENTS

[0032] FIG. 1 shows a schematic representation of an implementation of a computation of the spatial distribution of a static magnetic field according to the invention. According to this implementation an spatial distribution of spin-phase accruals between magnetic resonance echo signals from the imaging zone is computed 101 from acquired magnetic resonance signals and applied as input to computation 103 of the spatial distribution of a static magnetic field. Also an estimate of the proton density distribution in the imaging zone as accessed 102 and input to an estimate of the proton density distribution in the imaging zone. The proton density distribution may be estimated from magnetic resonance signals acquired during a preparation phase, e.g. collaterally obtained from measurement of spatial coil sensitivity profiles of the magnetic resonance examination system's radio frequency receiver coils.

[0033] FIG. 2 shows a more detailed schematic representation of an implementation of a computation of the spatial distribution of a static magnetic field according to the invention. From the spin phase accruals 101, an initial magnetic susceptibility distribution 201 is estimated that is consistent with the spin phase accruals. From the proton density distribution another initial magnetic susceptibility distribution is estimated 202 that is consistent with the proton density distribution. Then, an iterative approach 203 is applied to fit a magnetic susceptibility distribution that optimally reconciles consistency with both the spin phase accrual distribution and the proton density distribution. From the ultimate fitted magnetic susceptibility distribution, the spatial magnet field distribution is computed 103.

[0034] The input to the envisaged processing is:

[0035] An initial version of a B.sub.0-map (actually: its deviation from a perfectly homogenous field), B.sub.m(x) ‘m’ stands for “measured”. The vector x is 3-dimensional. A rough location-dependent knowledge of the inaccuracy of the above, expressed as σ.sub.B(x).

[0036] An estimate of proton-density of the object, ρ(x).

[0037] Yet, that proton-density map is not of direct use, but it requires some type of pre-processing that can segment the volume e.g. into three categories: “soft”, “interstitial” and “air”.

[0038] The output is an improved estimate of the error field, denoted as {tilde over (B)}(x). An important intermediate result is defined as {tilde over (χ)}(x) denoted as “an improved estimate of the local magnetic susceptibility” (although the word ‘improved’ is somewhat misplaced, since there is no ‘initial’ estimate here).

[0039] The envisaged processing estimates the intermediate {tilde over (χ)}(x), by

[00001]$\tilde{\chi }\left(x\right)=\underset{\chi }{\min }\left[\frac{{.Math.B_0ℱ\left\{\left[\frac{1}{3}-\frac{k_{z,p}^2}{k_{x,p}^2+k_{y,p}^2+k_{z,p}^2}\right]{ℱ}^{-1}\left(\tilde{\chi }\left(x\right)\right)\right\}-B_m\left(x\right)-n{B}_{warp}.Math.}^2}{{\sigma }_B^2\left(x\right)}\right]$

[0040] Subject to setting of the value of {tilde over (χ)}(x) in selected ranges in volumes of which the material content is a priori known. Good results are achieved when setting: [0041] −9.Math.10.sup.−6<{tilde over (χ)}(x)<−7.Math.10.sup.−6 for soft tissue [0042] −9.Math.10.sup.−6<{tilde over (χ)}(x)<0 for interstitial [0043] {tilde over (χ)}(x)=0 for air.

[0044] In these equations,

[0045] B.sub.0 is the value of the main field strength

[0046] In e.g. k.sub.x,p, the ‘p’ stands for “physical”, i.e. in units of l/length (e.g. l/m or l/mm).

[0047] ℑ indicates the Fourier transform.

[0048] n is any integer.

[0049] B.sub.warp corresponds to

[00002]$\frac{1}{\Delta \mathrm{TE}}$

(assuming that B.sub.m(x) has been established by examining phase difference between two echoes).

[0050] In principle, there is a one-to-one correspondence between {tilde over (B)}(x) and {tilde over (χ)}(x):

[00003]$\tilde{B}\left(i;\chi \right)=B_0ℱ\left[\left(\frac{1}{3}-\frac{k_{z,p}^2}{k_{x,p}^2+k_{y,p}^2+k_{z,p}^2}\right){ℱ}^{-1}\left(\tilde{\chi }\left(x\right)\right)\right]$

[0051] So far for the processing.

[0052] All of this could be seen as an advanced way of filtering B.sub.m(x) into {tilde over (B)}(x).

[0053] The central element of the processing, the

[00004]$“ℱ\left[\left(\frac{1}{3}-\frac{k_{z,p}^2}{k_{x,p}^2+k_{y,p}^2+k_{z,p}^2}\right){ℱ}^{-1}\left(\tilde{\chi }\left(x\right)\right)\right]”$

can be reasonably approximated by a multi-resolution decomposition of {tilde over (χ)}(x), combined with a local (small-kernel-)convolution on each of the levels.

[0054] Alternatively:

[00005]$\tilde{\chi }\left(x\right)=\underset{\chi }{\min }\left[\frac{{.Math.\tilde{B}\left(x;\chi \right)-B_m\left(x\right)-n{B}_{warp}.Math.}^2}{{\sigma }_B^2\left(x\right)}\right],$

while for example
−9.Math.10.sup.−6<{tilde over (χ)}(x)<−7.Math.10.sup.−6 for soft tissue, −9.Math.10.sup.−6<{tilde over (χ)}(x)<0 for interstitial, and

[0055] {tilde over (χ)}(x)=0 for air. The magnetic field consistent with the susceptibility distribution is denoted as {tilde over (B)}(x; χ), while the information on the protons density is taken into account in the set value ranges for the susceptibility in the segmented areas. In other words, given the set values ranges for the susceptibility values, the susceptibility is matched to produces the magnetic field distribution n that fits the measured magnetic field distribution.

[0056] FIG. 3 shows diagrammatically a magnetic resonance imaging system in which the invention is used. The magnetic resonance imaging system includes a set of main coils 10 whereby the steady, uniform magnetic field is generated. The main coils are constructed, for example in such a manner that they enclose a tunnel-shaped examination space. The patient to be examined is placed on a patient carrier which is slid into this tunnel-shaped examination space. The magnetic resonance imaging system also includes a number of gradient coils 11, 12 whereby magnetic fields exhibiting spatial variations, notably in the form of temporary gradients in individual directions, are generated so as to be superposed on the uniform magnetic field. The gradient coils 11, 12 are connected to a gradient control 21 which includes one or more gradient amplifier and a controllable power supply unit. The gradient coils 11, 12 are energised by application of an electric current by means of the power supply unit 21; to this end the power supply unit is fitted with electronic gradient amplification circuit that applies the electric current to the gradient coils so as to generate gradient pulses (also termed ‘gradient waveforms’) of appropriate temporal shape The strength, direction and duration of the gradients are controlled by control of the power supply unit. The magnetic resonance imaging system also includes transmission and receiving coils 13, 16 for generating the RF excitation pulses and for picking up the magnetic resonance signals, respectively. The transmission coil 13 is preferably constructed as a body coil 13 whereby (a part of) the object to be examined can be enclosed. The body coil is usually arranged in the magnetic resonance imaging system in such a manner that the patient 30 to be examined is enclosed by the body coil 13 when he or she is arranged in the magnetic resonance imaging system. The body coil 13 acts as a transmission antenna for the transmission of the RF excitation pulses and RF refocusing pulses. Preferably, the body coil 13 involves a spatially uniform intensity distribution of the transmitted RF pulses (RFS). The same coil or antenna is usually used alternately as the transmission coil and the receiving coil. Furthermore, the transmission and receiving coil is usually shaped as a coil, but other geometries where the transmission and receiving coil acts as a transmission and receiving antenna for RF electromagnetic signals are also feasible. The transmission and receiving coil 13 is connected to an electronic transmission and receiving circuit 15.

[0057] It is to be noted that it is alternatively possible to use separate receiving and/or transmission coils 16. For example, surface coils 16 can be used as receiving and/or transmission coils. Such surface coils have a high sensitivity in a comparatively small volume. The receiving coils, such as the surface coils, are connected to a demodulator 24 and the received magnetic resonance signals (MS) are demodulated by means of the demodulator 24. The demodulated magnetic resonance signals (DMS) are applied to a reconstruction unit. The receiving coil is connected to a preamplifier 23. The preamplifier 23 amplifies the RF resonance signal (MS) received by the receiving coil 16 and the amplified RF resonance signal is applied to a demodulator 24. The demodulator 24 demodulates the amplified RF resonance signal. The demodulated resonance signal contains the actual information concerning the local spin densities in the part of the object to be imaged. Furthermore, the transmission and receiving circuit 15 is connected to a modulator 22. The modulator 22 and the transmission and receiving circuit 15 activate the transmission coil 13 so as to transmit the RF excitation and refocusing pulses. In particular the surface receive coils 16 are coupled to the transmission and receive circuit by way of a wireless link. Magnetic resonance signal data received by the surface coils 16 are transmitted to the transmission and receiving circuit 15 and control signals (e.g. to tune and detune the surface coils) are sent to the surface coils over the wireless link.

[0058] The reconstruction unit derives one or more image signals from the demodulated magnetic resonance signals (DMS), which image signals represent the image information of the imaged part of the object to be examined. The reconstruction unit 25 in practice is constructed preferably as a digital image processing unit 25 which is programmed so as to derive from the demodulated magnetic resonance signals the image signals which represent the image information of the part of the object to be imaged. The signal on the output of the reconstruction monitor 26, so that the monitor can display the magnetic resonance image. It is alternatively possible to store the signal from the reconstruction unit 25 in a buffer unit 27 while awaiting further processing.

[0059] The magnetic resonance imaging system according to the invention is also provided with a control unit 20, for example in the form of a computer which includes a (micro)processor. The control unit 20 controls the execution of the RF excitations and the application of the temporary gradient fields. To this end, the computer program according to the invention is loaded, for example, into the control unit 20 and the reconstruction unit 25. The B.sub.0-mapping may be computed by the processor of the control unit and then used in the reconstruction unit 25 to correct the reconstructed magnetic resonance image for encoding errors due to the B.sub.0-inhomogeneities. Also, or alternatively, the B.sub.0-mapping may be applied to the gradient control unit to control the gradient coils for active shimming to compensate for the B.sub.0-inhomogeneities. The B.sub.0-mapping may be applied as well to specially designed shim coils (not shown) to compensate for the B.sub.0-inhomogeneities.