Magnetic resonance imaging apparatus and method

Electricity: measuring and testing – Particle precession resonance – Spectrometer components

Reexamination Certificate

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C324S309000

Reexamination Certificate

active

06806712

ABSTRACT:

The present invention generally relates to a magnetic resonance (MR) imaging apparatus and method for generating a substantially homogenous static magnetic field within a desired spatial region. More particularly the present invention relates to the magnetic resonance imaging of a desired region of a human body (“anatomy of interest”).
Magnetic resonance imaging relies on the availability of a uniform, homogenous, static magnetic field. In order to image an “anatomy-of-interest” (e.g. chest, head or female breast etc.) it needs to be placed within the substantially homogenous zone of the static magnetic field generated by the magnet assembly of the overall magnetic resonance imaging apparatus. This zone of the magnetic field which is used in the magnetic resonance imaging will be referred to as the “imaging-field-zone”.
Since the quality of the MR image is very dependent upon the degree of homogeneity achieved within the imaging-field-zone, a great deal of time and effort is devoted to the design and optimisation of MR imaging magnets to achieve this. As schematically depicted in
FIG. 1
a
, the main objective of the magnet designer is to design a magnet which generates a static magnetic flux density vector 10;
B
(
x,y,z
)=
B
x
i+B
y
J+B
z
k
  (Eq. 1)
over the imaging-field-zone which is directed substantially along a single axis only (e.g. the z axis in
FIG. 1
a
where B (0,0, z)=B
z
k). In Equation 1 above, the vectoral quantities are printed in bold letters and the i, j and k represent ‘unit vector’ quantities along the x-axis, y-axis and z-axis, respectively. Furthermore, it can be shown that the B
z
component (magnitude of the B vector) can be expressed in the form of;
B
z
=B
0
+B
1
+B+
2
+B
3
+ . . . +B
n
  (Eq.2)
where the terms “B
1
+B+
2
+B
3
+ . . . +B
n
” are collectively referred to as “error terms” or “residuals”. The ultimate objective of the magnet designer, during the design process, is to get rid of as many of the “error terms” as possible in order to achieve the ultimate goal of;
B
z
=B
0
  (Eq.3)
However, in practice it is not always possible to obtain the same degree of homogeneity within every point of a given imaging field-zone. This will be explained with reference to
FIG. 1
a
which is a schematic, longitudinal section of a prior-art magnet (conventional, six-coil, cylindrical, narrow-bore, axially-long) assembly. A lateral view of the patient is also included. Conventionally, the patient
1
is arranged to lie on a patient-positioner
7
which is parallel to the z-axis of the magnet. The conventional whole body scanner magnet assembly comprises a low temperature chamber
21
operating at around 4° K (i.e. liquid helium temperature). Superconductive low temperature coils
20
are arranged within the chamber
21
. A thermal radiation shield
22
is arranged around the chamber
21
and around the radiation shield
22
is arranged a high temperature chamber
23
operating at 77K (i.e. liquid nitrogen temperature). Surrounding the high temperature chamber
23
is a vacuum chamber
24
. Within the bore of the magnet of the assembly there is provided an arrangement
26
of room temperature shims, gradient coils and radio frequency (RF) coils.
The general purpose, conventional, whole-body MR imaging systems are designed with the objective of being able to image effectively all parts of the human anatomy by using the centre of the thorax as reference-point
28
(i.e. x=0, y=0 and z=0). They are designed as axially symmetric structures. The magnet design process involves the selection of the number of superconductive coils
20
, their axial and radial locations, the magnitude of the current flowing through the coils, the number of turns, etc. When all these parameters are ‘optimally’ selected, then a substantially homogenous magnetic field within an imaging-field-zone, which is centred at the geometrical centre
28
of the magnet assembly, may be generated. However, although the magnet designer can achieve theoretically possible maximum homogeneity (i.e. B
z
=B
0
) at the geometrical centre
28
of the magnet, inherently the same maximum homogeneity can not be achieved at other spatial locations. In other words, as schematically illustrated in
FIGS. 2
a
and
2
b
, although the field is “pure” at the geometrical centre
28
of the magnet, the degree of homogeneity decreases gradually at increasing distances away from the geometrical centre
28
. Figuratively speaking, for example the homogeneity calculated in a, say, 10 cm diameter imaging-field-zone is much better that the homogeneity calculated for much larger imaging-field-zones (e.g. say 20 cm, 30 cm etc, diameter zones). This is schematically depicted in
FIGS. 2
a
and
2
b
where the most homogenous zone within the magnet's bore
27
is zone
12
a
(i.e. zone with the smallest diameter). The homogeneity calculated for the zones
12
b
,
12
c
and
12
d
is increasingly much lower that the homogeneity calculated for zone
12
a.
Now, let us assume the anatomy-of-interest which needs to be imaged, with a prior art, conventional whole body MR scanner, is female breast. More specifically, as depicted in
FIGS. 1
a
,
1
b
and
1
c
, the anatomy-of-interest
2
comprises the left-breast
2
a
and the right breast
2
b
of the patient
1
. It can be seen that the centre
6
of the anatomy-of-interest is not at the geometrical centre
28
of the magnet assembly. This means that the area
11
of the anatomy-of-interest does not lie within the available most homogenous zones of the magnet's imaging-field. Referring now to
FIGS. 1
b
and
1
c
, if one attempts to position the centre of the breast
6
at the geometrical centre of the whole body magnet
28
by lifting the patient upwards, in majority of the cases, this is not permitted by the narrower bore size of the magnet. As illustrated in
FIG. 1
c
, this would not be possible since the arms
4
a
and
4
b
of the patient
1
would be required to lie within the regions provided for the room temperature shim coils, gradient coils and RF coils
26
in a conventional whole body scanner. For this reason, one could argue that the breast images obtained by a whole body scanner is not always optimal, since they are not always imaged by making use of the most homogeneous magnetic field available.
Unfortunately, at present, the clinical application of breast MR imaging is restricted to conventional, general purpose, whole body scanners (e.g.
figure 1
a
), and the breast referral cases have to compete for scanner time with many other cases of referrals. The design of the conventional ‘whole body’ scanners tend to encapsulate the patient and only permit limited access to patient by the health care professionals and make the task of performing interventional procedures more complicated and difficult. Interventional radiologists and surgeons require:
(a) totally ‘open-access’ to breast and patient;
(b) easy ‘image-guided’ diagnostic procedure (fine needle aspiration, core biopsy etc.);
(c) easy ‘image-guided’ surgical treatment (e.g. lumpectomy) and other therapeutic procedures (e.g. tumour ablation using laser, rf, microwave, focused ultrasound energies, etc.).
The whole body cylindrical magnet based MR scanners (see
FIG. 1
a
) does not allow the direct interactive placements of interventional devices (e.g. biopsy needles and optical/laser fibres used for thermal ablation of tumours) into the breast. They require the withdrawal of the patient from the bore of the magnet in order to gain access to the breast. This together with the possible movement of the patient may result in the imprecise insertion and localisation of the interventional devices.
Furthermore, in addition to the problem of patient inaccessibility within the bore of the conventional whole body scanner, the claustrophobic nature of the design due to patient encapsulation is also a factor affecting patient-acceptance rate of whole body scann

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