X-ray or gamma ray systems or devices – Specific application – Computerized tomography
Reexamination Certificate
2002-11-08
2004-11-16
Bruce, David V (Department: 2882)
X-ray or gamma ray systems or devices
Specific application
Computerized tomography
C378S015000, C378S901000
Reexamination Certificate
active
06819734
ABSTRACT:
The present application hereby claims priority under 35 U.S.C. §119 on German patent publication number DE 10155089.8 filed Nov. 9, 2001, the entire contents of which are hereby incorporated herein by reference.
BACKGROUND OF THE INVENTION
Image data of an examined measurement object can be obtained using modern medical diagnosis methods, such as X-ray computed tomography (CT), for example. The examined measurement object is generally a patient.
X-ray computed tomography—abbreviated to CT hereinafter—is a special X-ray recording method which fundamentally differs from the traditional X-ray tomography method in terms of image construction. CT recordings yield transverse sectional images, that is to say images of body layers which are oriented essentially perpendicularly to the body axis. The tissue-specific physical quantity represented in the image is the distribution of the attenuation value of X-ray radiation &mgr;(x, y) in the sectional plane. The CT image is obtained by reconstructing the one-dimensional projections of the two-dimensional distribution of &mgr;(x, y), which projections are supplied by the measurement system used, from numerous different viewing angles.
The projection data are determined from the intensity I of an X-ray after it has passed through the layer to be imaged and its original intensity I
0
at the X-ray source in accordance with the absorption law
ln
I
I
0
=
-
∫
L
μ
(
x
,
y
)
ⅆ
l
(
1
)
The integration distance L represents the path of the X-ray considered through the two-dimensional attenuation distribution &mgr;(x, y). An image projection is then composed of the measured values of the line integrals through the object layer, with the measured values being obtained with the X-rays of a viewing direction.
The projections—characterized by the projection angle &agr;—originating from many different directions are obtained by a combined X-ray tube detector system which rotates about the object in the layer plane. The most common apparatuses at the present time are so-called “fan beam apparatuses”, in which a tube and an array of detectors (a linear arrangement of detectors) rotate in the layer plane jointly about a center of rotation, which is also the center of the circular measurement field. “Parallel beam apparatuses”, which exhibit very long measurement times, are not explained here. It shall be pointed out, however, that a transformation from fan to parallel projections and vice versa is possible, so that the present invention, which will be explained with reference to a fan beam apparatus, can also be used for parallel beam apparatuses without any restriction.
In the case of fan beam geometry, a CT recording comprises line integral measured values—ln(I/I
0
) of arriving beams which are characterized by a two-dimensional linkage of the projection angle &agr;&egr;[0,2&pgr;) and the fan angles &bgr;&egr;[−&bgr;
0
,&bgr;
0
] defining the detector positions (&bgr;
0
is half the fan aperture angle). Since the measurement system only has a finite number k of detector elements and a measurement comprises a finite number y of projections, this linkage is discrete and can be represented by a matrix:
{tilde over (
p
)}(&agr;
y
,&bgr;
k
)[0,2&pgr;)×[−&bgr;
0
,&bgr;
0
] (2)
or
{tilde over (
p
)}(
y,k
): (1,2, . . .
N
P
)×(1,2, . . .
N
S
) (3)
The matrix {tilde over (p)}(y, k) is called a sinogram for fan beam geometry. The projection number y and the channel number k are of the order of magnitude of 1000.
The principle of image reconstruction in CT by calculating the &mgr; value distribution will not be discussed further, for the sake of brevity. This is illustrated in detail, for example, in “Bildgebende Systeme für die medizinische Diagnostik” [“Imaging Systems for Medical Diagnosis”], third edition, Munich: Publicis MCD Verlag, 1995, ed.: Morneburg, Heinz, ISBN 3-89578-002-2.
However, the task of image reconstruction is not yet concluded with the calculation of the &mgr; value distribution of the radiographed layer. In the medical field of application, the distribution of the attenuation coefficient &mgr; represents only an anatomical structure which still has to be represented in the form of an X-ray image.
After a proposal by G. N. Hounsfield, it has become generally customary to transform the values of the linear attenuation coefficient &mgr; (which has the dimension unit cm
−1
) to a dimensionless scale in which water acquires the value 0 and air the value −1000. The conversion formula to this “CT number” reads as follows:
CT
number
=
μ
-
μ
Water
μ
Water
1000
(
4
)
The unit of the CT number is called a “Hounsfield Unit” (HU). This scale is highly suited to the representation of anatomical tissue since the unit HU expresses the deviation in thousandths of &mgr;
water
and the &mgr; values of most substances inherent to the body differ only little from the &mgr; value of water. From the range of numbers (from −1000 for air to approximately 3000), only integers are used as carriers of image information.
However, the representation of the entire scale range of about 4000 values would far surpass the discrimination capability of the human eye. Moreover, the observer is often interested only in a small excerpt from the range of attenuation values, e.g. the differentiation of grey and white brain substance, which differ only by about 10 HU.
For this reason, so-called image windowing is used. In this case, only part of the CT value scale is selected and spread over all available grey shades. Even small attenuation differences within the chosen window thus become perceptible grey tone differences, while all the CT values below the window are represented black and all the CT values above the window white. The image window can be varied arbitrarily but in terms of its central level and in terms of its width.
The image data obtained usually contained not only the desired image information of the examined measurement object but also information which can be attributed to disturbing influences during the measurement operation.
Generally, a distinction is made between two different categories of problems which reduce the quality of the image data obtained: image noise and artifacts. These two problems will be explained in more detail below.
Image noise can in turn be subdivided into a plurality of causes.
The main part of the image noise is brought about by the quantum noise which results from the fact that each radiation comprises a finite number of quanta, so that the number of measured quanta always virtuates normally distributed about an average value.
Further causes of the image noise are the usually not exactly monochromatic quanta of the X-ray tubes that can be realized in practice, and scattered radiation due to interactions between the X-ray radiation used and the electron shell of atoms during transmission through the measurement object.
Artifacts are also subdivided further: aliasing, partial volume artifacts, hardening artifacts and motion artifacts are typical artifacts whose occurrence depends in particular on the geometry or a motion of the measurement object.
Effects corresponding to the above-described image noise and the artifacts can also be found in other imaging systems for medical diagnosis.
Ring artifacts constitute a particular form of artifacts the cause of which is to be sought primarily in the computer tomography imaging system itself that is used:
As already described above, a plurality of detectors (up to 1000) are used in fan beam apparatuses. Therefore, there is the possibility of inadequate calibration of the individual detectors. In other words, identical attenuations of the radiation penetrating through the measurement object are measured differently by different detectors.
In the case of inadequate calibration of the individual detectors of a computer tomograph, the image data obtained have concentric rings or partial ring arcs about the center of rotation, w
Bruce David V
Harness & Dickey & Pierce P.L.C.
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