GB2374674A

GB2374674A - Magnetic Resonance Imaging

Info

Publication number: GB2374674A
Application number: GB0109793A
Authority: GB
Inventors: Joseph Vilmos Hajnal; Mark Bydder; David James Larkman
Original assignee: Marconi Medical UK Ltd
Current assignee: Philips Design Ltd
Priority date: 2001-04-20
Filing date: 2001-04-20
Publication date: 2002-10-23
Also published as: GB0109793D0

Abstract

Better signal-to-noise ratio is obtained when combining the images from an array coil 6-13 used in magnetic resonance imaging apparatus by using the relative sensitivity of each coil obtained by division of the images from each coil on a pixel by pixel basis.

Description

MAGNETIC RESONANCE IMAGING This invention relates to magnetic resonance (MR) imaging. A prior art magnetic resonance imaging apparatus is shown in Figure 1. A patient I (shown in section) is slid axially into the bore 2 of a superconducting magnet 3, and the main magnetic field is set up along the axis of the bore, termed by convention the Z-direction. Magnetic field gradients are set up, for example, in the Z-direction, to confine the excitation of magnetic resonant (MR) active nuclei (typically hydrogen protons in water and fat tissue) to a particular slice in the Z-direction e. g. that illustrated in Figure 1 and, in the horizontal X and the vertical Y-directions as seen in Figure 1, to encode the resonant MR nuclei in the plane of the slice. An r. f transmit coil (not shown) applies an excitation pulse to excite the protons to resonance, and an r. f. receive coil array consisting of a pair of coils 4,5 picks up relaxation signals emitted by the disturbed protons.

To encode/decode received signals in the Y-direction, the signals are detected in the presence of a magnetic field gradient, termed a frequency encode or read-out (R. O.) gradient, to enable different positions of relaxing nuclei to correspond to different precession frequencies of those nuclei about the direction of the main magnetic field due to the influence of the gradient. The data is digitised, and so for each r. f. excitation pulse, a series of digital data points are collected, and these are mapped into a spatial frequency domain known as k-space (Figure 2). Each r. f. pulse permits at least one column of digital data points to be collected.

To encode/decode the received signals in the X-direction, after each r. f. pulse has been transmitted and before data is collected with the read-out gradient applied, a magnetic field gradient in the X-direction is turned on and off. This is done for a series of magnitudes of magnetic field gradients in the X-direction, one r. f. pulse typically corresponding to a different magnitude of gradient in the X-direction. The series of measurements enable spatial frequencies to be built up in the X-direction.

On the k-space matrix shown in Figure 2, the columns of data points correspond to data collected at different magnitudes of phase-encode (P. E. ) gradients.

The field of view imaged by the magnetic resonance imaging apparatus depends on the spacing of the data points in the phase-encode and read-out directions, and the resolution of the image depends on how far the points extend in each direction i. e. how large the maximum phase-encode gradient is, and on the magnitude of the read-out gradient combined with the duration of data collection.

Conventionally, the data collected by the r. f. receive coil arrangement and depicted in Figure 2 is subject to a two dimensional fast Fourier Transform in a Fourier Transform processor (not shown) to produce a pixelated spatial image.

A slice image is shown in Figure 3. For the purposes of explanation, the symbol of a circle la, has been illustrated in both the patient I shown in Figure 1 and the image shown in Figure 3. Figure 3 implies that the spacing of data points in the phase-encode gradient direction is sufficient to image the whole of the circle shown in Figure 1.

The final image is produced by combining signals from the image shown in Figure 3 corresponding to each of the array coils 4,5.

Another form of array coil is shown in Figure 4. An array of coils 6 to 13 is arranged beneath the spine 14 of a patient 7 (shown schematically). Such a coil arrangement can be used to produce a saggital (vertical longitudinal) section through the spine (plane 15).

The array coil 6-13 may be used with magnetic resonance imaging apparatus as shown in Figure 1, with the patient again extending longitudinally in the bore 2 of the magnet 3. The slice 15 may be selected by a magnetic field gradient along the x axis, and the signals received by the coils may be spatially encoded in the Y-direction by means of a read-out magnetic field gradient in that direction. The slice 15 may be spatially encoded/decoded in the Z-direction by means of a magnetic field gradient applying phase-encode gradients in the Z-direction.

Each coil produces a spatial frequency array of data, termed k-space, as is shown in Figure 2, which is Fourier Transformed to form an image as shown in Figure 3. These images are then combined to produce the final image.

For both the arrangements of Figures 1 and Figure 4, the signals from the array coils are combined by combining the intensity p, of the final image from each coil on a pixel by pixel basis. It has been pointed out (Roemer P B, Edelstein W A, Hayes C E, Souza S P, Mueller 0 M-The NMR Phased Array, Magnetic Resonance in Medicine 16,192-225 (1990) that the optimal array coil performance is achieved by coherently adding on a pixel by pixel

basis the intensities Pi with each weighted by the sensitivity of the coil b, at the pixel location. Optimal signal-to-noise ratio is obtained from the sum

over all coils, where P is the resultant intensity of each pixel of the combined image. Roemer noted that in most practical circumstances b, is not known at each pixel location but that, provided that pi is much greater than the background noise, a good approximation is achieved by replacing bi with p, so that

This works well in strong signal regions, but incurs a noise penalty. This is because coils of the array which for one reason or another do not contribute any signal at all still contribute noise, which would be included in the overall sum.

The invention provides magnetic resonance imaging apparatus comprising means for exciting magnetic resonant (MR) active nuclei in a region of interest, an array of at least two r. f. receive coils for receiving data from the region of interest, and means for producing an image by combining signals from the coils of the array, using the relative sensitivity of each coil.

The use of the relative sensitivity of each coil reduces the noise penalty inherent in the previous square root of the sum of the squares of the intensity of each pixel.

The relative sensitivity of each coil may be obtained by dividing the sensitivity of each coil with that of one particular coil, or alternatively the sensitivity of each coil may be divided by the square root of the sum of the squares of the sensitivity of each coil. This may be done on a pixel by pixel basis from the images produced by each coil, or the calculation may be done in k-space. It is advantageous to filter the relative sensitivity before the computation to produce the image is carried out.

Magnetic resonance imaging apparatus in accordance with the invention will now be described in detail, by way of example, with reference to the accompanying drawings in which: Figure I is a schematic axial sectional view of known magnetic resonance imaging apparatus; Figure 2 is a representation of data in k-space resulting from the signal picked up by a receive coil of the apparatus; Figure 3 is a representation of the shape in the image domain represented by the data in kspace; and Figure 4 is a perspective schematic view of an alternative form of known array coils.

The magnetic resonance imaging apparatus of the invention may take the general form shown in Figures I and 4.

According to the invention, the image is produced by combining signals from the array coils 4,5 or 6-13 by using the relative sensitivity of each coil. Thus, for either the array of Figure 1 or the array of Figure 4, a standard full field of view image is obtained for each coil. MR signals are received by the coils, suitably encoded by magnetic field gradients, and respective k-space matrices are built up from the signals received by each coil. These are respectively Fourier Transformed to produce real images. The images from the individual coils (the uncombined images) are saved in complex on magnitude form.

Estimates are then made of) {J by taking pixel by pixel ratios of the uncombined images.

This divides out information about the subject being imaged, which is common to all uncombined images, leaving a ratio of sensitivities.

Now that we have a sensitivity, in this case a relative sensitivity, known at each pixel, it is possible to calculate the intensity of each pixel of the resultant image P as follows:

Thus, each pixel of the final image has an optimal signal-to-noise ratio as defined by Roemer, except that each sensitivity is divided by that for one particular coil, which has the effect of setting the sensitivity of the selected coil to unity i. e. normalising it in the calculation of P.

Instead of dividing by an individual coil signal, any combination of coils that is linearly proportional to the anatomical signal can be used, for example the square root of the sum

of the squares of the coil signals. This would be produced by computing, on a pixel by pixel basis the square of intensities of each of the images produced by the coils, summing those squares, and then taking the square root. This gives an image with exactly the same signal properties as the standard sum of squares image, but with improved signal-to-noise ratio.

In order to reduce noise in the estimate of, the uncombined images can be low pass filtered before division and/or the ratio smoothed after division. This low pass filtering blurs the anatomical content, but provided the smoothing kernel is small compared to the spatial variations of the coil sensitivity patterns (which usually only vary slowly across the field of view) this has negligible effect on the ratio obtained.

Complex uncombined image appear to be superior to magnitude images where the signal from a coil is not larger than the noise level because the low pass filtered versions can asymptote to zero with complex (signed) data.

Improved signal-to-noise ratio and reduction of phase-encoded motion artefacts arise where, because a coil has low sensitivity to a given pixel, its contribution is attenuated. Where some of the coils in the array have high local sensitivity patterns, such as with intracavitary coils, the final images can combine all coil information but be presented with a sensitivity profile of only the coils that have slowly varying spatial patterns. For example, if a prostate coil is combined with an external phased array, final images can have signal modulation characteristic of the external phased array, but with locally improved signal-to-

noise ratio where the prostate coil has its high sensitivity region. This is easy to view with a dynamic range of a standard display, while retaining the local sensitivity improvement associated with the internal coil.

The particular examples given show two or eight array coils. However, the invention is not limited to any particular number of array coils. Nor is it limited to any particular arrangement of the coils, nor to any particular form of magnet i. e. electromagnet, resistive or superconducting, or permanent, or to any configuration such as solenoidal or open.

In addition, the invention has been described in terms of the processing being carried out in the image domain. However, it may be performed instead in the Fourier domain for either the final signal combination or the calculation of relative coil intensities.

Claims

CLAIMS 1. Magnetic resonance imaging apparatus comprising means for exciting magnetic resonant (MR) active nuclei in a region of interest, an array of at least two r. f. receive coils for receiving data from the region of interest, and means for producing an image by combining signals from the coils of the array, using the relative sensitivity of each coil.
2. Magnetic resonance imaging apparatus as claimed in claim 1, wherein the image producing means is arranged to produce the relative sensitivity of each coil by dividing the intensity of the image from each coil by the intensity of an image derived from at least one other coil, on a pixel by pixel basis.
3. Magnetic resonance imaging apparatus as claimed in claim 2, in which the division is relative to the intensity of the image obtained by one particular coil of the array.
4. Magnetic resonance imaging apparatus as claimed in claim 2, in which the division is relative to the square root of the sum of the squares of the intensities of the images produced by all the coils of the array.
5. Magnetic resonance imaging apparatus as claimed in claim 1, in which the relative sensitivity is calculated in k-space.
6. Magnetic resonance imaging apparatus as claimed in any one of claims 1 to 5, in which the relative sensitivity of each coil is low pass filtered.

<Desc/Clms Page number 10>
7. A method of magnetic resonance imaging comprising exciting magnetic resonant (MR) active nuclei in a region of interest, receiving r. f. data from the region of interest using an array of at least two r. f. receive coils, and producing an image by combining signals from the coils of the array using the relative sensitivity of each coil.