Electricity: measuring and testing – Particle precession resonance – Spectrometer components
Reexamination Certificate
2002-10-21
2004-04-06
Shrivastav, Brij B. (Department: 2862)
Electricity: measuring and testing
Particle precession resonance
Spectrometer components
C324S320000
Reexamination Certificate
active
06717409
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates in general to a gradient system for use in magnetic resonance tomography (MRT) for examining patients. The present invention relates, in particular, to a switchable gradient system and to a method for calculating such a system, in the case of which so-called booster coils are used.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear spin resonance and has been used successfully as an imaging method for over 15 years in medicine and in biophysics. In this method of examination, the object is exposed to a strong, constant magnetic field. This aligns the nuclear spins of the atoms in the object, which were previously oriented randomly. Radio-frequency energy can excite these “ordered” nuclear spins to a specific oscillation (resonant frequency). In MRT, this oscillation generates the actual measurement signal (RF response signal), which is picked up by suitable receiving coils.
Having exact information relating to the respective origination location of the RF response signal (location information or location coding) is a precondition for the image reconstruction. This location information is obtained by means of additional magnetic fields (magnetic gradient fields) relative to the static magnetic field along the three spatial directions. These gradient fields are small by comparison with the main field and are generated by additional resistance coils in the patient opening of the magnet. Due to the gradient fields, the overall magnetic field differs in each volumetric element and therefore so is the resonant frequency. If a defined resonant frequency is irradiated, it is therefore possible to excite only the atomic nuclei that are situated at a location at which the magnetic field fulfills the resonance condition. A suitable change in the gradient fields allows the location of such a volumetric element where the resonance condition is fulfilled to be displaced in a defined fashion, and thus for the desired region to be scanned. The gradient fields are therefore switched repeatedly in an MR sequence for excitation (coding) and reading out (detection) of the nuclear resonance signals.
This allows a free choice of the layer to be imaged, as a result of which it is possible to obtain tomographic images of the human body in all directions.
The basic design of an MRT apparatus is illustrated in
FIG. 4. A
basic field magnet
23
(for example an axial superconducting air-coil magnet with active stray field screening) which generates a homogeneous magnetic basic field in an interior space. The superconducting magnet
23
contains superconducting coils which are located in liquid helium. The basic field magnet
23
is surrounded by a double-shell tank which is made of stainless steel, as a rule. The inner tank, which contains the liquid helium and serves in part as winding body for the magnet coils, is suspended at the outer tank, which is at room temperature, via fiber-glass-reinforced plastic rods which are poor conductors of heat. A vacuum prevails between inner and outer tanks.
The cylindrical gradient coil
25
in the inner space of the basic field magnet
23
is inserted concentrically into the interior of a support tube by means of support elements
24
. The support tube is delimited externally by an outer shell
26
, and internally by an inner shell
27
.
The gradient coil
25
has three windings which generate respective gradient fields, each field being proportional to the current in each coil and being orthogonal to one another. As illustrated in
FIG. 5
, the gradient coil
25
has an x-gradient coil
28
, a y-gradient coil
29
and a z-gradient coil
30
, which are respectively wound around the coil core
31
and thus generate a gradient field, respectively in the directions of the Cartesian coordinates x, y and z. Each of these coils is fitted with a dedicated power supply unit in order to generate independent current pulses with accurate amplitudes and timing in accordance with the sequence programmed in the pulse sequence controller
20
. The required currents are at approximately 250 A. Since the gradient switching times are to be as short as possible, current rise rates and current fall rates (slew rates) on the order of magnitude of 250 kA/s are necessary. A high gradient power is required in order to obtain images with high spatial resolution and a short measuring time. According to the current state of the art, the gradient intensities are 30-60 mT/m for switching times of 100-500 &mgr;s.
However, these high rates of change in the magnetic field in the body of the patient, at times, can cause painful peripheral nerve stimulations. The threshold for peripheral nerve stimulations scales with the magnitude of the homogeneity volume (DSV=Diameter of Spherical Volume, or field mode), which is fixed by the gradient system.
For this reason, and in order to do justice to different applications in MRT—in particular in functional imaging—it is necessary to lend the MRT machine DSV flexibility. The technically possible system power, in conjunction with avoidance of peripheral nerve stimulations, can be fully exploited according to the current state of the art by using a gradient system that has a number of field characteristics (field modes). The field characteristic, or the field mode, describes a generally spherical region in the interior of the homogeneous basic field in which the gradient deviates by less than 5% from the reference value at the coil center. The radius and quality of the corresponding homogeneity region definitively determine the essential system properties of the gradient system such as switching time, maximum gradient intensity and stimulation threshold. According to the current state of the art, they can be changed in discrete steps by using switchable gradient coils.
According to the current state of the art, a switchable gradient system with a number of field characteristics can be implemented in various ways respectively having different advantages and disadvantages:
A) by combining or integrating a number of (completely shielded) coil sections.
B) Modular conductor bundling, by combining suitable conductor bundles within a coil plane for a discrete number of field characteristics.
In order, for example, to generate two different field characteristics, in method A two different (actively shielded) coils are interleaved. Different field modes can be obtained by appropriate electric interconnection of the two coils.
FIG. 6
a
schematically illustrates the idealized z-direction field pattern of two whole-body gradient fields DSV
1
, DSV
2
with different homogeneity radii. The two fields are generated by a gradient system of a whole-body tomography apparatus with two independent coil sections in the z-direction.
FIG. 6
b
shows a transverse section (x-y plane) through the whole-body tomography apparatus.
The gradient system is shielded from the outside and from the superconducting basic field magnet by a cryoshield
32
(referred to as the tank above). The system uses a large whole-body coil
33
, which produces a correspondingly large spherical homogeneity volume (DSV
1
)
34
.
Disadvantages of this large-volume whole-body coil
33
are a high inductance and a high stimulation effect. These two disadvantages can be compensated by the use of a second smaller coil—the so-called insert coil
35
. By energizing the insert coil
35
, a small, elliptical homogeneity volume (DSV
2
)
36
inside the large homogeneity volume of the basic field magnet is typically obtained.
As can be seen from
FIG. 6
b
, each of the two coil sections
33
,
35
occupies a hollow cylinder of a certain thickness. The radii of the respective hollow cylinders are different as a rule, and the coil sections are therefore located on different winding planes. As already indicated and as can be seen in
FIG. 4
b
—this leads to a reduction in the inside diameter of the gradient tube and to a reduction in the patient space.
Such a design is therefore compatible with th
Kimmlingen Ralph
Schuster Johann
Schiff & Hardin LLP
Shrivastav Brij B.
Siemens Aktiengesellschaft
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