Magnetic resonance imaging

Details Magnetic resonance imaging Magnetic resonance imaging

2000-11-15

2002-05-28

Arana, Louis (Department: 2862)

Electricity: measuring and testing

Particle precession resonance

Using a nuclear resonance spectrometer system

C324S309000

Reexamination Certificate

active

06396269

ABSTRACT:

BACKGROUND
This invention relates to magnetic resonance (MR) imaging.
The invention is particularly concerned with reduction in the time needed to collect data for imaging a region of interest of a patient.
For example, it is sometimes desired to view the progress of a contrast agent in the bloodstream through a region, in order to help distinguish between benign tumour, malignant tumour and unaffected tissue, and rapid imaging is required to monitor the flow of contrast agent to the various sites. A malignant tumour may require a greater supply of blood and could show up in this way.
A prior art magnetic resonance imaging apparatus is shown in
FIG. 1. A
patient
1
(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
FIG. 1 and
, in the horizontal X and the vertical Y directions as seen in
FIG. 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 a pair of r.f. receive coils
4
,
5
pick 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 read-out 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 (FIG.
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.
On the K-space matrix shown in
FIG. 2
, the columns of data points correspond to r.f. pulses followed by different magnitudes of phase-encode 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 e.g. how large the maximum phase-encode gradient is.
The signals received from the r.f. receiver coils
4
,
5
are subject to a two dimensional fast Fourier Transform in Fourier Transform processors
6
,
7
, to produce pixelated images which are stored in image memories
8
,
9
.
A slice image is shown in FIG.
3
. For the purposes of explanation, the symbols of a circle, a square and a triangle have been illustrated in both the object
2
and the image in FIG.
3
.
FIG. 3
implies that the spacing of data points in the phase-encode gradient direction is sufficient to image the whole of the slice shown in FIG.
1
.
Between each r.f. pulse, there is a certain minimum pulse repetition time, and the collection of data implied by
FIGS. 2 and 3
may therefore take too long to permit flows of contrast agent to be adequately monitored.
One technique used to reduce the data collection time is to cut out, say, half the phase-encode steps e.g. by keeping the same maximum phase-encode gradient but omitting every other column of data. This would then halve the data collection time.
The spacing of the data points in the phase-encode direction would now have doubled, so that the field of view in the corresponding image domain would have halved. (The field of view in the read-out direction would remain the same because the number of data points collected during read-out would remain the same.) The imaged area would now cover little more than the square illustrated in FIG.
1
. This is shown in FIG.
4
. Unfortunately, MR signals from the circle and the triangle would still be picked up by each receive coil
4
,
5
, and the processing is such that these regions would be folded back over the square, with the circle and the triangle being interposed. This is a problem known as aliasing. Separate points P
1
, P
2
in the object would be imaged at the same pixel in FIG.
4
.
It is to enable the aliased image to be unfolded that two receive coils
4
,
5
have been used. (Magnetic Resonance in Medicine 42:952-962 (1999)—SENSE: Sensitivity Encoding For Fast MRI by Klaas P Pruessmann, Markus Weiger, Markus B Scheidegger and Peter Boesiger). With the two receive coils and suitable processing, the benefit of a halved data collection time can be retained, but the image can be unfolded. This technique is referred to as SENSE (sensitivity encoding).
The response of coil
4
to MR signals from point P
1
is perhaps greater than its response to MR signals emanating from point P
2
, and the position is perhaps the other way round with regard to the coil
5
. Knowledge of the relative sensitivities of the coils
4
,
5
can be used in the processor
10
to unfold the aliased image of
FIG. 4
into an image having the pattern of
FIG. 3
, although with a lower signal-to-noise ratio because only half the data has been collected.
For example, coil
4
receives MR signals from both points P
1
and P
2
, and the combined intensity of the resulting pixel stored in image memory
8
(C
1
) depends on the sensitivity of the coil
4
to point P
1
(S
11
) and the sensitivity of the coil
4
to point P
2
(S
12
). The image memory
9
also has an intensity stored (C
2
) corresponding to points P
1
and P
2
which depends on the sensitivity of coil
5
to P
1
(S
21
) and on its sensitivity to P
2
(S
22
). Thus,
C
1
=S
11
P
1
+S
12
P
2
C
2
=S
21
P
1
+S
22
P
2
More generally
(
C
)=(
S
).(
P
)
where (C), (S) and (P) are matrices.
Provided the sensitivity terms are known, the intensity corresponding to each pixel P
1
and P
2
separately can be calculated in processor
10
, for each pair of points of the image which are mapped onto the same pixel in the aliased image. The aliased image can then be unfolded and viewed in full on display
11
.
To calibrate the coils
4
,
5
requires a knowledge of the sensitivity of each coil at each point. Of course, the data collected at each image memory
8
,
9
maps the sensitivity of each coil
4
,
5
onto each pixel, but in a way which is clearly weighted by the intensity produced by each point of the object.
The paper referred to suggests collecting the sensitivity data using a third receive coil, in the form of a body coil, which has a uniform sensitivity over each part of the object.
To suppress the weighting produced by the object being imaged, it turns out that the lines of the images in the image memories
8
,
9
have to be divided by corresponding lines i.e. for the same point in K-space in the read-out direction, produced by the imaging with the body coil.
However, for example, in the case of an object with the weighting shown in
FIG. 5
, the sensitivity profile for the coil
4
could be as shown in FIG.
6
and that for the body coil as shown in FIG.
7
.
FIG. 8
shows the division of one by the other and the presence of noise.
This is because the body coil picks up noise from the whole of the section of the body being imaged, and the resulting sensitivity profile produced by the division, for each line of the image, is masked by noise.
To overcome this, the paper referred to proposes to divide the response of one coil
4
(
FIG. 9
) by the response of the other coil
5
(
FIG. 10

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