U.S. patent application number 09/745704 was filed with the patent office on 2001-12-20 for densitometry adapter for compact x-ray fluoroscopy machine.
Invention is credited to Bisek, Joseph P., Ergun, David L., Mazess, Richard B..
Application Number | 20010053202 09/745704 |
Document ID | / |
Family ID | 27617734 |
Filed Date | 2001-12-20 |
United States Patent
Application |
20010053202 |
Kind Code |
A1 |
Mazess, Richard B. ; et
al. |
December 20, 2001 |
Densitometry adapter for compact x-ray fluoroscopy machine
Abstract
An accessory for fluoroscopy equipment is provided to support
the x-ray tube and detector on a pedestal with respect to a patient
limb for quantitative bone densitometry measurement. Software
loaded into the associated digital imaging fluoroscopy equipment
provides necessary correction of the images for the quantitative
accuracy needed for bone densitometry. Alternatively, a specialized
detector or extremely low form factor image intensifier may be
inserted in the pedestal to be used in lieu of the fluoroscopy
equipment detector. A similar software correction is performed on
an associated computer when a separate detector must be used.
Inventors: |
Mazess, Richard B.;
(Madison, WI) ; Ergun, David L.; (Verona, WI)
; Bisek, Joseph P.; (Madison, WI) |
Correspondence
Address: |
STEVEN J. WIETRZNY
Quarles and Brady LLP
411 East Wisconsin Avenue
Milwaukee
WI
53202
US
|
Family ID: |
27617734 |
Appl. No.: |
09/745704 |
Filed: |
December 21, 2000 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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09745704 |
Dec 21, 2000 |
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09281518 |
Mar 30, 1999 |
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6215846 |
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09745704 |
Dec 21, 2000 |
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09006358 |
Jan 13, 1998 |
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6007243 |
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09006358 |
Jan 13, 1998 |
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PCT/US97/02770 |
Feb 21, 1997 |
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60080164 |
Mar 31, 1998 |
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60011993 |
Feb 21, 1996 |
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Current U.S.
Class: |
378/196 ;
348/E3.045; 348/E5.078; 348/E5.081; 348/E5.088; 378/208 |
Current CPC
Class: |
G01N 2223/419 20130101;
A61B 6/5282 20130101; A61B 6/0421 20130101; H04N 3/2335 20130101;
H04N 5/217 20130101; A61B 6/4441 20130101; G06T 5/006 20130101;
A61B 6/583 20130101; A61B 6/4405 20130101; A61B 6/4225 20130101;
G01N 23/046 20130101; A61B 6/4035 20130101; A61B 6/548 20130101;
H05G 1/44 20130101; G01N 2223/612 20130101; H05G 1/60 20130101;
A61B 6/482 20130101; H04N 5/325 20130101; A61B 6/4488 20130101;
A61B 6/505 20130101 |
Class at
Publication: |
378/196 ;
378/208 |
International
Class: |
H05G 001/02 |
Claims
We claim:
1. A positioner for a fluoroscopy machine of a type having a
radiation source and detector separated along a beam axis and
mounted on an articulated arm, the adapter comprising: a weighted
base supporting the positioner upon a floor; a pedestal extending
upward from the base to provide receiving supports for at least one
of the radiation source and detector so that when the at least one
of the radiation source and detector is received by the receiving
supports, the beam axis is in predetermined orientation with
respect to the top of the pedestal; and a limb positioner attached
to the pedestal between the receiving supports so that a patient's
limb held by the limb positioner is intersected by the beam
axis.
2. The positioner as recited in claim 1 wherein the limb positioner
is removable and is sized to receive the patent's limb, wherein the
limb is an arm, a hand, a leg or a foot.
3. The positioner as recited in claim 1 wherein the limb positioner
is a vertically extending palm support.
4. The positioner as recited in claim 3 wherein the palm support is
attached to a mounting pin sized and positioned to fit within a
guide bore in the pedestal.
5. The positioner as recited in claim 4 wherein the palm support is
removable.
6. The positioner as recited in claim 1 wherein the limb positioner
is a foot cradle having: a calf support plate adjacent to the top
of the pedestal having at least one mounting pin depending downward
sized and positioned to fit within a guide bore in the pedestal; a
sole support plate joined to the calf support plate to define an
obtuse angle; and side gussets spanning the calf support plate and
the sole support plate, the gussets having apertures positioned to
be at the beam axis when the foot cradle is mounted to the
pedestal.
7. The positioner as recited in claim 6 wherein the foot cradle
includes padding material at top surfaces of the calf support
plates.
8. The positioner as recited in claim 6 wherein the foot cradle is
removable.
9. The positioner as recited in claim 1 including a calibration
material attached to the pedestal within the beam axis.
10. The positioner as recited in claim 1 including an anti-scatter
grid attached to the pedestal within the beam axis.
11. The positioner as recited in claim 1 including an occlude
attached to the pedestal within the beam axis.
12. The positioner as recited in claim 1 wherein the receiving
supports are disposed axially along a radius of a cavity, the
receiving supports extending into the cavity to define a radius of
lesser diameter than the cavity according to the size of the
radiation detector.
13. The positioner as recited in claim 1 further comprising an
index guide adjacent to the limb positioner for properly aligning
the radiation detector with the limb positioner on the
pedestal.
14. The positioner as recited in claim 1, further comprising: an
independent detector array located along the beam axis between the
limb positioner and the radiation detector when the radiation
detector is disposed in the receiving supports, the detector array
producing attenuated dual energy signals; and a processor receiving
the attenuated dual energy signals and calculating bone density
measurements.
15. The positioner as recited in claim 14 wherein the detector
array is a set of stimulable phosphor plates.
16. The positioner as recited in claim 14 wherein the detector
array is solid state detector.
17. The positioner as recited in claim 14 wherein the detector
array is a stacked linear array scanning detector.
18. A method of scatter correction for x-ray measurements
comprising the steps of: (a) obtaining an x-ray measurement
providing a plurality of x-ray attenuation values over an area; (b)
preparing a histogram of the x-ray attenuation values, the
histogram indicating the frequency of occurrence of attenuation
values over a range of attenuations; (c) dividing the attenuation
range into at least two regions corresponding to material types;
(d) associating a different deconvolution kernel with each region;
and (e) deconvolving the x-ray measurement with the different
deconvolution kernels depending on the regions of the deconvolved
attenuation values.
19. The method of claim 18 wherein the deconvolution kernels define
a width of deconvolution measured as a number adjacent x-ray
measurements within the deconvolution and wherein the different
deconvolution kernels have different widths.
20. The method of claim 19 wherein the regions corresponding to
greater attenuations, and hence to imaged materials having greater
attenuation, are associated with deconvolution kernels having
greater widths.
21. The method of claim 18 wherein the x-ray measurement is of a
human and wherein the attenuation range is divided into regions
corresponding to materials selected from the group consisting of:
air-only, thin tissue, thick tissue only, thin bone and, and
thick-only.
22. The method of claim 18 wherein the x-ray measurement is of a
human and wherein the deconvolved x-ray measurement is further
analyzed to provide a quantitative measurement of an imaged
object.
23. The method of claim 22 wherein the quantitative measurement is
an indication of bone density.
24. The method of claim 18 wherein during step (e) the
deconvolution kernels are further weighted according to the
intensity of the deconvolved attenuation values.
25. A method of scatter correction for x-ray measurements
comprising the steps of: (a) obtaining an x-ray measurement
providing a plurality of x-ray attenuation measurements over an
area; (b) preparing at least one deconvolution kernel for scatter
reduction in the x-ray measurements; (c) during the application of
the deconvolution kernel to the x-ray measurements, applying a
scale factor to the deconvolution kernel based on the value of the
x-ray measurements being deconvolved.
26. The method of claim 25 wherein the x-ray measurement is of a
human and wherein the deconvolved x-ray measurement is further
analyzed to provide a quantitative measurement of an imaged
object.
27. The method of claim 26 wherein the quantitative measurement is
an indication of bone density.
Description
[0001] This application is based on U.S. provisional application
No. 60/080,164 filed Mar. 31, 1998 and is a continuation in part of
U.S. application Ser. No. 08/814,800 filed Mar. 10, 1997 and is a
continuation in part of U.S. application Ser. No. 09/006,358 filed
Jan. 13, 1998 which is a continuation-in -part of PCT Application
Ser. No. 97/02770 designating the United States filed Feb. 21, 1997
claiming the benefit of provisional application No. 60/011,993
filed Feb. 21, 1996. This provisional application is incorporated
by reference herein.
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
FIELD OF THE INVENTION
[0002] The invention relates generally to x-ray equipment and in
particular to an adapter for bone density measurements as may be
used with compact x-ray fluoroscopy equipment used for orthopedic
and similar procedures.
BACKGROUND OF THE INVENTION
[0003] Portable x-ray fluoroscopy machines provide an x-ray source
held in opposition to an electronic image detector, typically on a
C-arm, so that x-rays from the x-ray source are received by the
image detector. The C-arm may slide through a collar so as to allow
it to be rotated to different angles about the patient. Further,
the collar may be supported by a pivoting arm providing additional
freedom in the positioning of the C-arm.
[0004] When the C-arm is correctly positioned, the x-ray source is
activated and x-rays pass through the patient to be received by the
image detector which provides electronic signals to a video
monitor. For larger mobile C-arm systems, the video monitor is
typically held on a separate cart or may be suspended from the
ceiling on a fixed bracket to be connected to the mobile unit when
the mobile unit is in place.
[0005] With improvements in electronic hardware and in particular
the development of compact image intensifiers and CCD video
cameras, it has become possible to build extremely compact mobile
C-arm systems. Such systems may make use of increasingly powerful
desktop computer technology for image processing and other tasks
and may use compact digital printers for producing images.
BRIEF SUMMARY OF THE INVENTION
[0006] The present invention provides an adapter that may convert a
compact fluoroscopy machine or other mobile x-ray source into a
precision quantitative densitometer suitable for measuring bone
mass or density such as may be helpful in the treatment and
detection of osteoporosis.
[0007] The invention provides a stand to be used with a fluoroscopy
machine, the stand having a cradle for accurately locating the
x-ray source and/or detector with respect to either the patient's
forearm or foot. Special software is loaded to the computer of the
fluoroscopy machine to operate the fluoroscopy machine in a
quantitative dual energy mode and to adapt the fluoroscopy data to
densiometric data. For fluoroscopy equipment not providing for
digital imaging or that may not be easily operated in a dual energy
mode, the invention includes a provision for a separate digital
dual energy detector and if necessary an associated processing
computer.
[0008] Other objects, advantages, and features of the present
invention will become apparent from the following specification
when taken in conjunction with the accompanying drawings.
BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS
[0009] FIG. 1 is a perspective view of the fluoroscopy machine
suitable for use with the present invention showing a C-arm
supporting an image intensifier/video camera and x-ray tube in
opposition for rotation in a vertical plane, the C-arm held along a
mid-line of a cart by an articulated arm attached to the side of
the cart;
[0010] FIG. 2 is a side view in elevation of the cart of FIG. 1
showing a slide attaching the articulated arm to the side of the
cart and showing a four-bar linkage motion of the arm for elevation
of the C-arm;
[0011] FIG. 3 is a top view of the C-arm system of FIG. 1 with the
articulated arm in partial phantom showing the four-bar linkage of
the arm for extending the C-arn toward and away from the cart;
[0012] FIG. 4 is a detail fragmentary view of an outer pivot of the
articulated arm attached to the C-arm such as allows limited
pivoting of a plane of rotation of the C-arm about a vertical
axis;
[0013] FIG. 5 is a detail view of the C-arn of FIG. 1 and the
attached x-ray tube assembly showing the electrical cabling
providing power to an x-ray tube power supply fitting into a groove
in the C-arm and showing an abutment of the anode of the x-ray tube
against the metal casting of the C-arm for heat sinking
purposes;
[0014] FIG. 6 is a schematic block diagram of the fluoroscopy
machine of FIG. 1 showing the path of control of a remote x-ray
tube power supply by a microprocessor and the receipt of data from
the image intensifier/video camera by the microprocessor for image
processing;
[0015] FIGS. 7 and 8 are simplified images such as may be obtained
by the system of FIG. 1 showing portions of the image having moving
elements and portions having stationary elements;
[0016] FIG. 9 is a flow chart of a method of the present invention
providing differently weighted noise reduction to different areas
of the image based on motion in the areas of the image;
[0017] FIG. 10 is a figure similar to that of FIG. 7 showing an
image of a rectilinear grid as affected by pincushion distortion in
the image intensifier and video camera optics such as may provide a
confusing image of a surgical tool being manipulated in
real-time;
[0018] FIG. 11 is a figure similar to FIG. 10 showing equal areas
of the image that encompass different areas of the imaged object,
such variation as may affect quantitative bone density
readings;
[0019] FIG. 12 is a plot of raw image data from the image
intensifier/video camera as is translated into pixel brightness in
the images of FIGS. 7, 8, 10, and 11 by the microprocessor of FIG.
6 according to a non-linear mapping process such as provides noise
equilibrium in the images and maximum dynamic range for clinical
data;
[0020] FIG. 13 is a histogram plotting values of data from the
image intensifier/video camera versus the frequency of occurrence
of data values showing an isolated Gaussian distribution at the
right most side representing unattenuated x-ray values;
[0021] FIG. 14 is a flowchart describing the steps taken by the
programmed microprocessor of FIG. 6 to identify background pixels
and remove them from a calculation of exposure rate used for
controlling the remote x-ray tube power supply of FIG. 6;
[0022] FIG. 15 is a detailed block diagram of the first block of
the flow chart of FIG. 14;
[0023] FIG. 16 is a first embodiment of the second block of the
flow chart of FIG. 14;
[0024] FIG. 17 is a second embodiment of the second block of the
flow chart of FIG. 14;
[0025] FIG. 18 is a detailed flow chart of the third block of the
flow chart of FIG. 14;
[0026] FIG. 19 is a schematic representation of a distorted image
of FIG. 11 and a schematic representation of a corresponding
undistorted image showing the variables used in the mathematical
transformation of the distorted image to correct for rotation and
distortion;
[0027] FIG. 20 is a flow chart of the steps performed by the
computer in correcting and transforming the image of FIGS. 11 and
19;
[0028] FIG. 21 is a perspective view of an occluder placed in an
x-ray beam prior to an imaged object and used for calculating
scatter;
[0029] FIG. 22 is a flow chart of the steps of calculating and
removing scatter using the occluder of FIG. 21;
[0030] FIG. 23 is a cross-sectional view through the occluder of an
imaged object of FIG. 21 along line 23-23, aligned with a graph
depicting attenuation of x-rays as a function distance along the
line of cross-section as well as theoretical attenuation without
scatter and scatter components;
[0031] FIG. 24 is a graphical representation of an adjustment of
calculated scatter from the image of FIG. 23 based on normalizing
points established by the occluder of FIG. 21;
[0032] FIG. 25 is a perspective view similar to that of FIG. 1
showing a C-arm system similar to that of FIG. 1 in position on the
densitometry cradle of the present invention to provide a beam of
x-rays along a horizontal axis across the top of the cradle;
[0033] FIG. 26 is a fragmentary exploded view of the x-ray detector
and x-ray source of the C-arm system of FIG. 25 removed from the
cradle and showing a removable foot positioner also removed from
the cradle;
[0034] FIG. 27 is an enlarged, fragmentary view of FIG. 26 showing
the alternative fitting of a forearm positioner or the foot
positioner within a channel of the cradle along the path of x-rays
between the x-ray source and x-ray detector;
[0035] FIG. 28 is a top plan view of the cradle of FIG. 25 but with
the C-arm removed, showing the forearm positioner in use with a
patient's arm;
[0036] FIG. 29 is a side elevational view of the cradle of FIG.
28;
[0037] FIG. 30 is a top plan view of the cradle of FIG. 28 but with
the foot positioner in use with a patient's leg and showing the use
of an auxiliary pancake image intensifier for use with fluoroscopy
or other x-ray equipment not having suitable dual energy or digital
imaging capabilities;
[0038] FIG. 31 is a cross sectional view of the pancake image
intensifier of FIG. 30 with the protective shrouding removed and
taken along a plane including the x-ray beam axis showing the
compact, high distortion configuration and an attached processing
computer;
[0039] FIG. 32 is a flow chart of software executed by a processing
computer associated with the x-ray detector for providing
quantitative densiometric data;
[0040] FIG. 33 is a block diagram of the steps of scatter
correction of the present invention; and
[0041] FIG. 34 is a block diagram illustrating the removal of line
correlated noise per the present invention.
DETAILED DESCRIPTION OF THE INVENTION
C-arm Support Mechanism
[0042] Referring now to FIG. 1, an x-ray machine 10 per the present
invention includes a generally box-shaped cart 12 having castors 14
extending downward from its four lower corners. The castors 14 have
wheels rotating about a generally horizontal axis, and swiveling
about a generally vertical axis passing along the edges of the cart
12. Castors 14, as are understood in the art, may be locked against
swiveling and/or against rotation.
[0043] With one castor 14 locked and the others free to rotate and
swivel, a pivot point 15 for the cart 12 is established with
respect to the floor such as may be used as a first positioning
axis 11 for the x-ray machine 10.
[0044] Positioned on the top of the cart 12 is a turntable 16
holding a video monitor 18 and attached keyboard 20 for swiveling
about a vertical axis for convenience of the user. The video
monitor 18 and the keyboard 20 may swivel separately so that one
operator may view the video monitor 18 while a second operates the
keyboard 20.
[0045] The video monitor 18 and the keyboard 20 allow for control
of a computer 22 contained in a shelf on the cart 12 open from the
front of the cart 12. The computer 22 may include a general
microprocessor-type processor 23 and a specialized image processor
27 for particular functions as will be described. The computer 22
further includes a number of interface boards allowing it to
provide control signals to various components of the x-ray machine
10 as will be described and to receive x-ray image data. In
addition, the computer 22 receives signals from a foot switch 61
that is used to activate the x-ray system for a brief exposure.
Control of the computer 22 may also be accomplished through a
remote control wand 63 of a type known in the art.
[0046] Referring now also to FIG. 2, attached to the right side of
the cart 12 is a horizontal slide 24 positioned to provide an
attachment point 26 for an articulated arm 19 supporting a
substantially circular C-arm 56, which in turn holds an x-ray tube
68 and an image intensifier 82 and camera 84, in opposition, and
facing each other as will be described below. The function of the
x-ray tube, the image intensifier and the camera are well known in
the prior art in the use of mobile C-arm type x-ray devices used
for image display and are described in U.S. Pat. No. 4,797,907
hereby incorporated by reference as part of the prior art. The
C-arm may be mass balanced, that is to say its weight may be
distributed to reduce its tendency to rotate through collar 54 so
that minimal frictional pressure may be used to prevent it from
moving.
[0047] The articulated arm 19 may be slid horizontally toward the
front of the cart 12 to provide a second positioning axis 25 of the
x-ray machine 10. A first pulley 28 is rotatively fixed in a
vertical plane, attached to the portion of the slide 24 that may
move with respect to the cart 12, and is pivotally attached to a
rigid arm 30 extending toward the front of the cart 12. The other
end of the rigid arm 30 supporting a second pulley 32 is also
mounted to swivel with respect to arm 30. A belt 34 wraps around a
portion of the circumference of each of pulleys 28 and 32 and is
affixed at one point along that circumference to each of the
pulleys 28 and 32 so that pivoting motion of the arm 30 about the
center point 26 of pulley 28 causes rotation of pulley 32 so that
it maintains a fixed rotational orientation with respect to the
cart 12 as pulley 32 and hence C-arm 56 is moved up and down along
a third axis 37. The linkage, so created, is a variation of the
"four bar linkage" well known in the art.
[0048] Helical tension springs (not shown for clarity) balance the
pulley 32 in rotative equilibrium about point 26 against the weight
of the articulated arm 19, C-arm 56, and other devices attached to
the arm 19.
[0049] Attached to pulley 32 is a third pulley 36 extending in a
generally horizontal plane perpendicular to the plane of pulley 32.
The third pulley 32 is attached pivotally to a second rigid arm 40
which at its other end holds another pulley 38 positioned
approximately at the midline 41 of the cart 12. The midline 41
symmetrically divides the left and right sides of the cart 12.
[0050] Portions of the circumference of pulleys 36 and 38 are also
connected together by a belt 44 so as to form a second four bar
linkage allowing pulley 38 to move toward and away from the cart
12, along a fourth positioning axis 45, with pulley 38 and C-arm 56
maintaining their rotational orientation with respect to cart
12.
[0051] Referring now to FIG. 4, pulley 38 includes a center shaft
member 50 having a coaxial outer collar 52 to which belt 44 is
attached. A stop 55 attached to the shaft 50 limits the motion of
the collar 52 in rotation with respect to the shaft 50 to
approximately 26 degrees. Frictional forces between shaft 50 and
collar 52 cause shaft 50 to maintain its rotational orientation
with respect to collar 52 and hence with respect to pulley 36 until
sufficient force is exerted on shaft 50 to displace it with respect
to collar 52. Thus pressure on the C-arm 56 can provide some
pivoting motion of the C-arm about the axis of the pulley along the
fifth positional axis 55.
[0052] Referring now to FIGS. 1, 3 and 4, attached to the shaft 50
is a C-arm collar 52 supporting the arcuate C-arm 56 curving
through an approximately 180 degree arc in a vertical plane
substantially aligned with the midline 41 of the cart 12 as has
been mentioned. The shaft 50 may connect to collar 52 so that the
latter may swivel in about a horizontal axis bisecting the circle
of the C-arm 56. This axis may be aligned with the center of mass
of the C-arm 56 so that there is not a self-righting tendency of
the C-arm or the axis may be placed above the axis of the C-arm so
as to provide for a beneficial self righting action. This motion is
orthogonal to that provided by motion of shaft 50 and may augment
that provided by the castors 14. Techniques of balancing the C-arm
in its various rotational modes, when this is desired, is taught by
U.S. Pat. No. 5,038,371 to Janssen issued Aug. 6th, 1991 and hereby
incorporated by reference as exemplifying the known prior art
understood to all those of ordinary skill in the art.
[0053] As described above, motion of the collar 52 may be had in a
vertical manner by means of the parallelogram linkage formed by
pulleys 28 and 32 of the articulated arm 19 as shown in FIG. 2.
Forward and backward motion away from and toward the cart 12 may be
had by the second four bar linkage formed from pulleys 36 and 38. A
slight pivoting of the C-arm 56 about a vertical axis slightly to
the rear of the collar 52 and concentric with the axis of pulley 38
may be had by means of the rotation between collar 52 and 50 of
FIG. 4. Greater rotation of the C-arm about the vertical axis
passing through pivot point 15 may be had by rotation of the cart
about one of its stationary castors 14. Thus, considerable
flexibility in positioning the C-arm may be had.
[0054] Referring now to FIG. 5, the C-arm 56 is an aluminum casting
having formed along its outer circumference a channel 58 into which
a cable 60 may be run as will be described. C-arm 56 has a
generally rectangular cross-section taken along a line of radius of
the C-arm arc. Each corner of that rectangular cross-section holds
a hardened steel wire 62 to provide a contact point for corner
bearings 64 within the collar 52. The corner bearings 64 support
the C-arm 56 but allow movement of the C-arm 56 along its arc
through the collar 52.
[0055] A cable guide pulley 66 positioned over the channel 58 and
having a concave circumference feeds the cable 60 into the channel
58 as the C-arm moves preventing tangling of the cable 60 or its
exposure at the upper edge of the C-arm 56 when the C-arm 56 is
rotated. The excess length of cable 60 loops out beneath the collar
52.
X-Ray Tube Cooling
[0056] Referring now to FIGS. 5 and 6, the C-arm supports at one
end a generally cylindrical x-ray tube 68 having a cathode 70
emitting a stream of electrons against a fixed anode 72. The
conversion efficiencies of x-ray tubes are such that the anode 72
can become quite hot and typically requires cooling. In the present
invention, the anode 72 is positioned to be bolted against the
aluminum casting of the C-arm 56 thereby dissipating its heat into
a large conductive metal structure of the C-arm 56.
[0057] The x-ray tube 68 is connected to an x-ray tube power supply
74 which separately controls the current and voltage to the x-ray
tube 68 based on signals received from the computer 22 as will be
described. The control signals to the x-ray tube power supply 74
are encoded on a fiber optic within the cable 60 to be noise
immune. Low voltage conductors are also contained within cable 60
to provide power to the x-ray tube power supply 74 from a low
voltage power supply 76 positioned on the cart 12.
[0058] During operation, an x-ray beam 80 emitted from the x-ray
tube 68 passes through a patient (not shown) and is received by an
image intensifier 82 and recorded by a charge couple device ("CCD")
camera 84 such as is well known in the art. The camera provides
digital radiation values to the computer 22 inversely proportional
to the x-ray absorption of the imaged object for processing as will
be described below. Each radiation value is dependent on the
intensity of x-ray radiation received at a specific point on the
imaging surface of the image intensifier 82.
Image Noise Reduction
[0059] Referring now to FIGS. 6 and 7, the data collected by the
CCD camera 84 may be used to provide an image 86 displayed on video
monitor 18. As will be described in more detail below, the CCD
camera receiving a light image from the image intensifier 82 at a
variety of points, provides data to the computer which maps the
data from the CCD camera 84 to a pixel 88 in the image 86. For
convenience, the data from the CCD camera 84 will also be termed
radiation data reflecting the fact that there is not necessarily a
one-to-one correspondence between data detected by the CCD camera
84 and pixels 88 displayed on the video monitor 18.
[0060] The CCD camera 84 provides a complete set of radiation data
for an entire image 86 (a frame) periodically once every "frame
interval" so that real-time image of a patient placed within the
x-ray beam 80 may be obtained. Typical frame rates are in the order
of thirty frames per second or thirty complete readouts of the CCD
detector area to the computer 22 each second.
[0061] Each frame of data is stored in the memory of the computer
22 and held until after complete storage of the next frame of data.
The memory of the computer 22 also holds an average frame of data
which represents an historical averaging of frames of data as will
now be described and which is normally used to generate the image
on the video monitor 18.
[0062] In a typical image 86, there will be some stationary object
90 such as bone and some moving object 92 such as a surgical
instrument such as a catheter. In a second image 86' taken one
frame after the image 86, the bone 90 remains in the same place
relative to the edge of the image 86 and 86', however the surgical
instrument 92 has moved. Accordingly, some pixels 88' show no
appreciable change between images 86 and 86', whereas some other
pixels 88" show a significant change between images 86 and images
86'.
[0063] Referring now to FIG. 9, as data arrives at the computer 22,
the computer 22 executes a stored program to compare current pixels
of the image 86' to the last pixels obtained from image 86 as
indicated by process block 94. This comparison is on a pixel by
pixel basis with only corresponding pixels in the images 86 and 86'
compared. The difference between the values of the pixels 88,
reflecting a difference in the amount of x-ray flux received at the
CCD camera 84, is mapped to a weight between zero and one, with
greater difference between pixels 88 in these two images
corresponding to larger values of this weight w. This mapping to
the weighting is shown at process block 96.
[0064] Thus pixels 88", whose value changes almost by the entire
range of pixel values between images 86 and 86', receive a
weighting of "one" whereas pixels 88' which have no change between
images 86 and 86' receive a value of zero. The majority of pixels
88 being neither unchanged nor radically changed will receive a
value somewhere between zero and one.
[0065] Generally, because the amount of x-ray fluence in the beam
80 is maintained at a low level to reduce the dose to the patient,
the images 86 and 86' will have appreciable noise represented as a
speckling in the images 86 and 86'. This noise, being of random
character, may be reduced by averaging data for each pixel 88 over
a number of frames of acquisition effectively increasing the amount
of x-ray contributing to the image of that pixel.
[0066] Nevertheless, this averaging process tends to obscure motion
such as exhibited by the surgical instrument 92. Accordingly, the
present invention develops an average image combining the values of
the pixels acquired in each frame 86, 86' in which those pixels in
the current image 86' which exhibit very little change between
images 86 and 86' contribute equally to the average image, but
those pixels in the current image 86' that exhibit a great degree
of change between images 86 and 86' are given a substantially
greater weight in the average image. In this process, a compromise
is reached between using historical data to reduce noise and using
current data so that the image accurately reflects changes.
Specifically, the value of each pixel displayed in the image is
computed as follows.
P.sub.i=(1-w)P.sub.i-1+wP.sub.i,t (1)
[0067] where P.sub.i-1 is a pixel in the previous average image, w
is the weighting factor described above and P.sub.i,t is the
current data obtained from the CCD camera 84. This effective merger
of the new data and the old data keyed to the change in the data is
shown at process block 98.
Image Intensifier Distortion
[0068] Referring now to FIG. 10, an image 86" of a rectilinear grid
100 positioned in the x-ray beam 80 will appear to have a barrel or
pincushion shape caused by distortion of the image intensifier 82
and the optics of the CCD camera 84. During a real-time use of the
image 86" by a physician, this distortion may cause confusion by
the physician controlling a tool 102. For example, tool 102 may be
a straight wire shown by the dotted line, but may display an image
86' as a curved wire whose curvature changes depending on the
position of the tool 102 within the image 86. This distortion thus
may provide an obstacle to a physician attempting to accurately
place the tool 102 with respect to an object within the image
86'.
[0069] Referring now to FIG. 1, the distortion of image 86" also
means that two equal area regions of interest 105 (equal in area
with respect to the image) do not encompass equal areas of the
x-ray beam 80 received by the image intensifier 82. Accordingly, if
the data from the CCD camera 84 is used for quantitative purposes,
for example to deduce bone density, this distortion will cause an
erroneous variation in bone density unrelated to the object being
measured.
[0070] Accordingly, the present inventors have adopted a real-time
digital re-mapping of radiation data from the CCD camera 84 to the
image 86 to correct for any pincushion-type distortion. This
remapping requires the imaging of the rectilinear grid 100 and an
interpolation of the position of the radiation data received from
the CCD camera 84 to new locations on the image 86" according to
that test image. By using digital processing techniques in a
dedicated image processor 27, this remapping may be done on a
real-time basis with good accuracy.
[0071] Referring to FIG. 19, there are two types of distortion,
isotropic and anisotropic. Isotropic distortion is rotationally
symmetric (e.g. like barrel and pin cushion distortion).
Anisotropic distortion is not rotationally symmetric. Both types of
distortion and rotation are so-called third order aberrations which
can be written in the form:
Dx=r.sup.2(Du-dv) (2)
Dy=r.sup.2(Dv+du) (3)
[0072] where Dx and Dy are pixel shifts due to distortion; r is the
distance from the correct position to the optical axis and D and d
are distortion coefficients which are constant and u and v are
correct pixel positions.
[0073] Referring also FIG. 2, received image 86 may exhibit pin
cushion distortion evident if an image 86 of the rectilinear grid
100 is made. The distortion is caused by the pixel shifts described
above.
[0074] Equations 1 and 2 may be rewritten as third order
two-dimensional polynomials, the case for equation (1)
following:
x=(a.sub.x+e.sub.xv+i.sub.xv.sup.2+m.sub.xv.sup.3)+(b.sub.x+f.sub.xv+j.sub-
.xv.sup.2+n.sub.xv.sup.3)u+(c.sub.x+g.sub.xv+k.sub.xv.sup.2+o.sub.xv.sup.3-
)u.sup.2+(d.sub.x+h.sub.xv+l.sub.xv.sup.2+p.sub.xv.sup.3)u.sup.3
(4)
[0075] In these polynomials, a.sub.x and a.sub.y govern the x and y
translation of the image, e.sub.x and b.sub.y take care of scaling
the output image, while e.sub.y and b.sub.x enable the output image
to rotate. The remaining higher order terms generate perspective,
sheer and higher order distortion transformations as will be
understood to those of ordinary skill in the art. Thirty-two
parameters are required for the two, third order polynomials. These
parameters may be extracted by a program executed by the computer
in an off-line (non-imaging) mode after imaging the known grid 100
and comparing the distorted image of the grid 100 to the known grid
100 to deduce the degrees of distortion.
[0076] Referring now to FIG. 19 in a first step of the correction
process, the grid 100 is imaged as indicated by process block 160
to determine the exact type of distortion present and to obtain
values for the coefficients a through p of the above referenced
polynomial equations.
[0077] At process block 166, these parameters may be input to the
computer 22 and used at a transformation of received image 86 into
image data 164 as indicated by process block 168. For rotation of
the image 164, new parameters of the polynomials may be entered by
means of hand-held remote control wand 63 shown in FIG. 1.
[0078] The transformation process generally requires a
determination of the pixel shift for each radiation pixel 163 of
the input image 86 which in turn requires an evaluation of the
polynomials whose coefficients have been input. A number of
techniques are known to evaluate such polynomials including a
forward differencing technique or other techniques known in the
art. These transformations provide values of u and v for an image
pixel 170 corresponding to a particular radiation pixel 163.
[0079] After the transformation of process block 168, the u, v
locations of the radiation pixels will not necessarily be centered
at a pixel location defined by the hardware of the video monitor 18
which usually spaces pixels 170 at equal distances along a
Cartesian axis. Accordingly, the transformed pixels must be
interpolated to actual pixel locations as indicated by process
block 172.
[0080] A number of interpolation techniques are well known
including bilateral and closest neighbor interpolation, however in
the preferred embodiment, a high resolution cubic spline function
is used. A given value of an interpolated pixel 170 (P.sub.int) is
deduced from a 4.times.4 block of transform pixels (P.sub.ij) in
which it is centered as follows:
P.sub.int=f(n-2)X.sub.l+f(n-1)X.sub.2+f(n)X.sub.3+f(n+1)X.sub.4
(5)
[0081] where:
X.sub.i=f(m-2)P.sub.i,1+f(m-1)P.sub.i,2+f(m)P.sub.i,3+f(m+1)P.sub.i,4
(6)
[0082] where:
f(x)=(a+2)x.sup.3+-(a+3)x.sup.2+1 for x .epsilon.[0,1];
f(x)=ax.sup.3+-5ax.sup.2+8ax-4a for x .epsilon.[1,2]; (7)
[0083] f(x) is symmetrical about zero. In the preferred embodiment
a=-0.5
[0084] and where m and n are fractions indicating the displacement
of the neighboring pixels P.sub.i,j with respect to P.sub.int in
the x and y directions, respectively.
[0085] At process block 180, the transformed and interpolated image
is displayed.
Noise Equalization
[0086] Referring now to FIG. 12, the radiation data from the CCD
camera 84 are mapped to the brightness of the pixels of the image
86 according to a second transformation. In the preferred
embodiment, this mapping between CCD radiation data and image pixel
brightness follows a nonlinear curve 103 based on the hyperbolic
tangent and being asymptotically increasing to the maximum CCD
pixel value. This curve is selected from a number of possibilities
so that equally wide bands of image pixel brightness 104 and 106
have equal amounts of image noise. The curve 103 is further
positioned to provide the maximum contrast between clinically
significant tissues in the image.
Exposure Control
[0087] The noise in the image 86 is further reduced by controlling
the fluence of the x-ray beam 80 as a function of the density of
tissue of the patient within the beam 80. This density is deduced
from the image 86 itself produced by the CCD camera 84. In response
to the image data, a control signal is sent via the fiber optic
strand within the cable 60 to the x-ray tube power supply 74
positioned adjacent to the x-ray tube 68 (shown in FIG. 5). By
positioning the x-ray tube power supply 74 near the x-ray tube 68,
extremely rapid changes in the power supplied to the x-ray tube 68
may be obtained. Distributed capacitances along high tension cables
connecting the x-ray tube 68 to a stationary x-ray tube power
supply are thus avoided in favor of low voltage cable 60, and the
shielding and inflexibility problems with such high tension cables
are also avoided.
Automatic Technique Control
[0088] Referring now to FIGS. 13 and 14, a determination of the
proper control signal to send to the x-ray tube power supply 74
begins by analyzing the image data 86 as shown in process block
120. The goal is to provide for proper exposure of an arbitrary
object placed within the x-ray beam 80 even if it does not fill the
field of view of the CCD camera 84. For this reason, it is
necessary to eliminate consideration of the data from the CCD
camera 84 that form pixels in the image corresponding to x-rays
that bypass the imaged object and are unattenuated ("background
pixels"). These background pixels may be arbitrarily distributed in
the image 86 and therefore, this identification process identifies
these pixels based on their value. To do this, the computer 22
collects the values of the pixels from the CCD camera 84 in a
histogram 122 where the pixels are binned according to their values
to create a multiple peaked plot. The horizontal axis of the
histogram 122 may for example be from 0 to 255 representing 8 bits
of gray scale radiation data and the vertical axis may be a number
of pixels having a particular value.
[0089] If there is a histogram value at horizontal value 255, and
the maximum gray scale exposure recorded, the entire area of the
histogram 122 is assumed to represent the imaged object only (no
background pixels). Such a situation represents an image of raw
radiation only or a high dose image of a thin object with possible
clipping. In assuming that the whole histogram 122 may be used to
calculate technique without removal of background pixels, a reduced
exposure rate will result as will be understood from the following
description and the peak classification process, to now be
described, is skipped.
[0090] Otherwise, if there are no pixels with the maximum value of
225, the present invention identifies one peak 124 in the histogram
122 as background pixels indicated by process block 120 in FIG. 14.
In identifying this peak 124, the computer 22 examines the
histogram 122 from the brightest pixels (rightmost) to the darkest
pixels (leftmost) assuming that the brightest pixels are more
likely to be the unattenuated background pixels. The process block
120 uses several predetermined user settings as will be described
below to correctly identify the peak 124.
[0091] Once the peak 124 has been identified, the pixels associated
with that peak are removed per process block 126 by thresholding or
subtraction. In the thresholding process, pixels above a threshold
value 138 below the peak 124 are considered to be background pixels
and are omitted from an exposure rate calculation. In the
subtraction method, the peak 124 itself is used as a template to
identify pixels which will be removed.
[0092] At process block 128, an exposure rate is calculated based
on the values of the pixels in the remaining histogram data and at
process block 130, an amperage and voltage value are transmitted
via the cable 60 to the x-ray tube power supply and used to change
the power to the x-ray tube. Generally, if the exposure rate is
above a predetermined value, the amperage and voltage are adjusted
to cut the x-ray emission from the x-ray tube, whereas if the
exposure rate is below the predetermined value, the amperage and
voltage are adjusted to boost the exposure rate to the
predetermined value.
[0093] Referring now to FIGS. 13, 14 and 15, the process of
identifying background pixels will be explained in more detail.
Process block 120 includes as a first step, an identification of a
right most peak 124 in the histogram 122 (shown in FIG. 13) as
indicated by subprocess block 132.
[0094] At succeeding subprocess block 134, this right most peak 124
is compared against three empirically derived parameters indicated
in the following Table 1:
1TABLE 1 Minimum Slope Range (MSR) Minimum necessary pixel range
for which the slope of the peak must be monitonically increasing.
Histogram Noise Level (HNL) Minimum height of the maximum value of
the peak. Maximum Raw Radiation Width Maximum width of the detected
peak (MRRW) with respect to the width of the entire histogram.
[0095] Specifically at subprocess block 134, each identified peak
124 is tested against the three parameters indicated in Table 1. In
the description in Table 1, "width" refers to the horizontal axis
of the histogram 122 and hence a range of pixel values, whereas
"height" refers to a frequency of occurrence for pixels within that
range, i.e., the vertical axis of the histogram 122.
[0096] These first two tests, MSR and HNL, are intended to prevent
noise peaks and peaks caused by bad imaging elements in the CCD
camera 84 or quantization of the video signal in the A to D
conversion from being interpreted as background pixels.
[0097] Peaks 124 with a suitable stretch of monotonically
increasing slope 131 (shown in FIG. 13) according to the MSR value
and that surpass the histogram noise level HNL 133 are evaluated
against the MRRW parameter. This third evaluation compares the
width 135 of the histogram 122 against the width of the entire
histogram 122. The MRRW value is intended to detect situations
where the imaged object completely fills the imaging field and
hence there are no unattenuated x-ray beams or background pixels
being detected. A valid peak 124 will normally have a width 135
more than 33% of the total width of the histogram 122.
[0098] At decision block 136 if the peak 124 passes the above
tests, the program proceeds to process block 126 as indicated in
FIG. 14. Otherwise, the program branches back to process block 132
and the next peak to the left is examined against the tests of
process block 134 until a passing peak is found or no peak is
found. If no peak is found, it is assumed that there are no
background pixels and a raw exposure value is calculated from all
pixels as described above.
[0099] Assuming that a peak 124 passes the tests of Table 1, then
at process block 126 background pixels identified by the peak 124
selected at process block 120 are eliminated.
[0100] In a first method of eliminating background pixels indicated
at FIG. 16, a magnitude threshold 138 within the histogram 122 is
identified. Pixels having values above this threshold will be
ignored for the purpose of selecting an exposure technique. The
threshold 138 is established by identifying the center 140 of the
peak 124 (its maximum value) and subtracting from the value of the
center a value s being the distance between the start of the peak
124 as one moves leftward and the maximum 140. The area under the
histogram 122 for values lower than the threshold 138 is computed
to deduce a raw exposure value which will be used as described
below.
[0101] In a second embodiment, the shape of the histogram peak 124
from the start of the peak as one moves leftward to its maximum 140
is reflected about a vertical line passing through the maximum 140
and subtracted from the histogram peak 124 to the left of the
vertical line. This approach assumes that the peak 124 of the
background pixels is symmetrical and thus this method better
accommodates some overlap between the object pixels and the
background pixels in the histogram 122. Again, the remaining pixels
of the histogram 122 are summed (by integration of the area under
the histogram 122 minus the area of the peak 124 as generated by
the reflection) to provide a raw exposure value.
[0102] Referring now to FIG. 18, the raw exposure value is
transformed by the known transfer characteristics of the CCD camera
(relating actual x-ray dose to pixel value) to produce a calculated
current exposure rate as indicated at process block 144.
[0103] Referring to process block 146, the current exposure rate is
next compared to a reference exposure rate, in the preferred
embodiment being 1.0 mR per frame, however this value may be
refined after further clinical testing. If at process block 148,
the current exposure rate is within a "half fine-tune range" of the
reference exposure rate, then the program proceeds to process block
150, a fine tuning process block, and the amperage provided to the
x-ray tube are adjusted in accordance to the disparity between the
amperage and reference exposure rate. That is, if the current
exposure is greater than the reference exposure rate, the amperage
to the x-ray tube is reduced. The new value of amperage is compared
against a predetermined range of amperage values (maximum beam
current and minimum beam current values) so that the amperage value
may never vary outside of this range.
[0104] If at decision block 148, the current exposure rate is
outside of the half fine tune range established at decision block
148, a more substantial adjustment process is undertaken.
Generally, the exposure provided by an x-ray system will follow the
following equation:
X.apprxeq.s.multidot.mA.multidot.kVp.sup.n. (8)
[0105] where:
[0106] s is seconds,
[0107] mA is the amperage provided to the x-ray tube,
[0108] kVp is the voltage provided to the x-ray tube, and
[0109] n is a power factor dependent on the geometry of the machine
and the particular kind of object being imaged.
[0110] Generally, the value of n will not be known in advance.
Accordingly in the more substantial correction process, n is
deduced by obtaining two different exposures for equal
predetermined intervals with different kVp values so that the value
of n may be deduced.
[0111] At decision block 152, it is determined whether a first or
second reference exposure is to be obtained. If the first reference
exposure was just obtained, the program proceeds to process block
154 and a new value of kVp is determined for a second exposure. In
this case, the first exposure used will be that which was employed
to produce the histogram 122 as previously described.
[0112] If the comparison of process block 148 indicated that the
exposure rate was too high, a lower kVp value is selected; and
conversely, if the exposure at process block 148 indicated the
exposure was too low, an increased value of kVp is provided. The
new kVp value for the second exposure must be within a
predetermined range of kVp values established by the user.
Mathematically, the kVp value selected may be described as:
kVp.sub.2=kVp.sub.1+a(dkVp) (9)
[0113] where a is a step factor and
[0114] dkVp is a minimum practical change in tube voltage.
[0115] Two preferred means of selecting may be used: one providing
linear and one providing logarithmic scaling. Such scaling
techniques are well understood to those of ordinary skill in the
art.
[0116] If at decision block 152, a second frame has already been
taken with the new voltage value, then the program proceeds to
process block 156 and the value of n in equation (9) is calculated.
If the value of amperage is held constant between the first and
second frame, the value of n may be determined according to the
following equation: 1 n = log X 2 X 1 / log k Vp 2 kVp 1 ( 10 )
[0117] where X.sub.1 and X.sub.2 are the measured exposure rates at
the first and second frames, respectively and
[0118] kVp.sub.1 and kVp.sub.2 are the two x-ray tube voltages
during the first and second frames.
[0119] At process block 158, this value of `n` is checked against
threshold values intended to detect whether an erroneous value of n
has been produced as a result of `clipping` in the radiation data
used to calculate exposure. As is understood in the art, clipping
occurs when an increased dose of an element of the CCD camera
produces no increase in the camera's output.
[0120] At decision block 158, if the value of n calculated at
process block 156 is greater than or equal to one, it is assumed to
be valid and the program proceeds to process block 160 where kVp
and mA are adjusted by setting mA equal to a maximum reference
value and calculating kVp according to the following equation: 2
kVp new = kVp 2 ( X ref m A 2 X 2 m A ref ) 1 / n ( 11 )
[0121] where kVP.sub.new and mA.sub.new are the settings for the
next frame to be shot.
[0122] If the resulting kVp value conflicts with the minimum
system, kVp, kVp is set to the minimum system value and mA is
calculated according to the following equation using the mA and kVp
value of the second frame. 3 m A new = m A 2 X ref X 2 ( kVp 2 kVp
min ) n ( 12 )
[0123] If the value of n in decision block 158 is less than one,
then at process block 162, n is tested to see if it is less than
zero. This value of n is realized when the exposure rate of the
second frame changes in the opposite direction of the tube voltage.
This suggests a clipped histogram and therefore the program
branches back to process block 154 to obtain a new second frame.
This condition may also arrive from object motion between the first
and second frame.
[0124] On the other hand, if at decision block 162, n is not less
than zero (e.g. n is between zero and 1), the program proceeds to
process block 166. Here it is assumed that because the sensitivity
of the exposure rate on change in kVp is low, there may be some
partial clipping. New values of kVp and mA are then computed and
used with the previous second frame values to calculate a new n as
follows. Generally, if kVp and mA are high, they are both lowered
and if kVp and mA are low, they are both raised.
Scatter Reduction
[0125] Referring now to FIG. 1, the image produced by the present
invention may be used for quantitative analysis including, for
example, that of making a bone density measurement. It is known to
make bone density analyses from x-ray images through the use of
dual energy techniques in which the voltage across the x-ray tube
is changed or a filter is periodically placed within the x-ray beam
to change the spectrum of the x-ray energy between two images. The
two images may be mathematically processed to yield information
about different basis materials within the image object (e.g. bone
and soft tissue). For these quantitative measurements, it is
desirable to eliminate the effect of scatter.
[0126] Referring now to FIG. 23 in imaging a patient's spine 200,
for example, x-rays 202 are directed from an x-ray source 201
through the patient 199 to pass through soft tissue 204 surrounding
a spine 200. Certain of the x-rays 202 are blocked by the spine 200
and others pass through the spine 200 to be recorded at the image
intensifier 206. An attenuation image 208 measured by an image
intensifier measures those x-rays passing through the patient
109.
[0127] A portion 210 of the attenuation image directly beneath the
spine 200 records not only those x-rays 202 passing through the
spine 200 and the soft tissue 204 above and below it, but also
scattered x-rays 212 directed, for example, through soft tissue 204
to the side of the spine 200 but then scattered by the soft tissue
to proceed at an angle to the portion 210 of the attenuation image
208 beneath the spine 200. Because the scattered x-rays 212 do not
carry information about the attenuation of the spine 200, they are
desirably removed from the image 208 prior to its use in
quantitative measurement.
[0128] For this purpose, the present invention uses an occluder 214
being an x-ray transparent plate such as may be constructed of
Plexiglas and incorporating into its body, a plurality of x-ray
blocking lead pins 216. Preferably these pins are placed so as to
project images 218 onto the image 208 received by the image
intensifier 206 in positions outside an image 220 of the spine 200.
Generally therefore, the pins 216 are placed at the periphery of
the occluder 214. The pins 216 are sized so as to substantially
block all direct x-rays from passing through them but so that their
images 218 include a significant portion of scattered x-rays
212.
[0129] Referring now to FIG. 22 at a first step of a scatter
reduction operation with the occluder 214 of FIG. 21, an image is
acquired of the imaged object, for example, the spine 200 and its
surrounding soft tissue 204 (not shown in FIG. 21) including the
images 218 of the pins 216. This acquisition is indicated by
process block 221 of FIG. 22.
[0130] The pins 216 are held in predetermined locations with
respect to the image 208 so that their images 218 may be readily
and automatically identified. Preferably the pins 216 are placed at
the interstices of a Cartesian grid, however, other regular
patterns may be chosen. The image 208 may be corrected for
pincushion type distortion, as described above, so that the
locations of the pins 216 may be readily located in the image based
on their known positions in the occluder 214.
[0131] At each pin image 218, a value 222 indicating the magnitude
of the received x-rays, shown in FIG. 23, may be ascertained. This
value 222 measures the scatter received in the vicinity of image
218 caused generally by the effect of the soft tissue 204 and
possible secondary scatter effects in the image intensifier 206.
Values 222 are recorded, as indicated by process block 224, for
each pin image 218. From these values, a set of normalizing points
are established.
[0132] The image 208 is then used to derive a scatter map.
Referring to FIG. 23, generally the amount of scatter at a given
point will be a function of how many x-ray photons are received at
points adjacent to the given point. For example, comparing the
image 208 to a theoretical scatterless image 228 generally in an
attenuated region 230 of the image 208 (e.g., under the spine 200),
scatter will increase the apparent value in the image 208 as a
result of radiation from nearby low attenuation regions scattering
into the high attenuation region 230. Conversely the apparent value
at a low attenuation region 232 will be decreased because of the
scatter into the high attenuation region.
[0133] A map of the scattered radiation may thus be modeled by
"blurring" the image 208. This blurring can be accomplished by a
low pass filtering of the image 208, i.e., convolving the image 208
with a convolution kernel having rectangular dimensions
corresponding to the desired low pass frequency cut off. The effect
is an averaging of the image 208 producing scatter map 234.
[0134] The image used to produce the scatter map 234 is an
attenuation image 208 obtained from the patient 199 without the
occluder 214 in place, or may be an image 208 including the images
218 of the pins 216 but with the latter images 218 removed based on
knowledge of their location. This removal of images 218 may
substitute values of the image 208 at points 239 on either side of
the images 218. The process of driving the scatter map from the
image is indicated by process block 235 of FIG. 24.
[0135] Next as indicated by process block 237, the scatter map 234
is fit to the normalizing points 222 previously determined at
process block 224.
[0136] Referring to FIG. 24, the scatter map 234 is thus normalized
so that the portions 238 of the scatter map 236 located near the
places where the images 218 would fall are given values 222 as
determined at process block 224. This involves a simple shifting up
or down of the scatter map 236 and may employ a "least square" fit
to shift the scatter map 236 to multiple values 222 obtained from
each pin 216. As adjusted, the scatter map 236 is then subtracted
from the image 208 to eliminate or reduce the scatter in that image
as indicated by process block 239.
[0137] The effect of subtracting a low pass filtered or blurred
image properly normalized to actual scatter is to sharpen up the
image 208 but also to preserve its quantitative accuracy. Thus the
present invention differs from prior art scatter reduction
techniques in that it both addresses the variation in scatter
across the image caused by attenuation of x-rays by the imaged
object but also incorporates accurate measurements of scatter in
certain portions of the image.
DENSITOMETER ADAPTER
[0138] Referring now to FIG. 25, a mobile fluoroscopy machine 310
suitable for use with the present invention is similar to that
which has been described above with respect to FIG. 1 with
exceptions that will be apparent from context.
[0139] The mobile fluoroscopy machine 310 includes a mobile cart
312 supporting a computer 314 and monitor and keyboard 317 for
receiving and processing digital x-ray image data. The cart 312
supports on one side an articulating arm assembly 316 terminating
in a rotatable C-arm 318. The C-arm supports, at the ends of the C,
an image intensifier 320 and an x-ray source 322 opposed along an
axis 324 so that the x-ray source 322 projects a cone-beam of x-ray
radiation toward the image intensifier 320 along axis 324.
[0140] The articulating arm assembly 316 is connected to the C-arm
318 through one or more pivotal links 327 so that the axis 324 may
be positioned to be horizontal approximately two feet above the
floor to rest upon or be supported against the upper end of a
supporting pedestal 326 or may be attached to the cart 312.
Referring also to FIG. 26, the pedestal 326 includes a
hemicylindrically concave cradle 328 at its upper surface to
receive a lower portion of the cylindrical image intensifier 320
when the C-arm is so positioned to rest against the pedestal
326.
[0141] The pedestal 326 also provides on its upper surface a
channel 330 extending across the axis 324 between the image
intensifier 320 and the x-ray source 322 when the latter are
positioned on the pedestal 326. The channel 330 may receive a limb
positioner 332 such as may be adapted to support a patient's foot
or arm across the axis 324 for densiometric measurement. The
pedestal 326 may be weighted so as to provide a stable surface for
support of the x-ray source 322 and image intensifier 320 and to
provide adequate support for the patient's limb. The height of the
pedestal 326 is selected to be suitable for either arm or foot
imaging.
[0142] Referring now to FIG. 27 and 30, the channel 330, extending
substantially perpendicularly to axis 324 and has a horizontal
bottom surface 333 pierced by two vertically extending guide holes
334 which may be used to receive and position corresponding pins
337 on one of two limb positioners 332. A foot positioner 336, as
shown in FIG. 26, provides a padded calf support plate 338 fitting
adjacent to the bottom surface 333 and an upwardly extending sole
support 340 forming an obtuse angle with respect to the calf
support plate 338. A cushion 342 on the calf support plate 338 may
be adjusted so as to allow the patient's leg to extend upward
somewhat from vertical for comfort. Gussets 344 span the angle
between the sole plate 340 and calf support plate 338 to fix them
in relative position but include apertures 346 to allow for the
free passage of x-rays through a portion along axis 324 where the
os calcis of the heal will be located.
[0143] When positioned within the channel 330, the foot positioner
336 is also supported by upwardly extending channel sidewalls 348
which serve further to provide an alignment surface for the imaging
face of the image intensifier 320 or other detector array and on
the other side, an alignment surface for an emitting face of the
x-ray source 322. Channel sidewalls 348 are generally radio
translucent so as to permit the passage of x-rays therethrough, but
may include: calibration materials such as are well known in the
art for calibrating dual energy devices, antiscatter grids also
well known in the art, or occluders for evaluating scatter as have
been described above or in the parent applications hereby
incorporated by reference.
[0144] Referring to FIG. 27 and 28, for forearm imaging, the foot
positioner 336 is removed and a palm support 352 is inserted by
means of pin 337 in one of the holes 334 so as to locate a user's
arm resting against the bottom surface 333 with the user's palm
against the palm support 352 such that the bones of the forearm are
placed along the axis 324 for imaging.
[0145] Referring to FIG. 28 and 29, the hemispherical support
cradle 328 may include three radially inwardly extending ribs 354
attached by means of screws or the like to be replaceable. Two of
the ribs 354 are positioned in a horizontal plane to substantially
bisect the image intensifier 320 when it is placed within the
cradle 328. The third rib 354 is positioned at the bottom of the
cradle 328 and is opposed by a rotating locking collar 358 which
may be used to further secure the image intensifier 320 within the
cradle 328. The front edge of the image intensifier is abutted
against the upright face of the dividing barrier so as to precisely
locate it along axis 324. The inner edges of these ribs 354 define
an inner radius 356 of lesser diameter than the cradle 328 that by
proper design of the ribs 354 may be adjusted to conform to the
outer surface of a particular image intensifier 320.
[0146] Referring now again to FIG. 30, some fluoroscopy equipment
will not permit digital imaging or the necessary dual energy
control needed for densitometry. Accordingly, an independent
detector array 360 may be placed within the cradle 328 in lieu of
the image intensifier 320. This detector array 360 may be a pair of
stimulable phosphor plates as are understood in the art with
intermediate filtering so as to provide dual energy readings with a
polychromatic x-ray source. In this way a switching of voltage on
the x-ray source 322, as described above, can be avoided.
Alternatively, the detector array 360 may be a large area solid
state detector or scanning detector assembly such as are understood
in the art including those constructed of amorphous silicon and
thin film transistor technology or those employing active pixel
technology in which C-MOS integrated circuit fabrication techniques
are employed. These detectors may be used with a switched x-ray
source 322 to provide dual energy imaging or may be used in a
stacked configuration with intermediate filtering so as to provide
separate energy measurements, or may be used in a side-by-side
configuration with interleaved detector elements filtered so as to
be selectively sensitive to different energies.
[0147] In the preferred embodiment, and as shown in FIG. 31, the
independent detector array 360 is a "pancake" image intensifier
361, suitably small so as to fit within the space between a
conventional image intensifier 320 and the x-ray source 322.
Referring to both FIG. 30 and 31, the pancake image intensifier 361
includes a vacuum bowl 362 having a planar front surface 364 for
receiving x-rays 366 (normally through the channel sidewalls 348 of
the stand 326).
[0148] According to conventional design, the x-rays 366 pass
through the front surface 364 of the vacuum bowl 362 to strike a
target material 368 to eject electrons 370 into the volume of the
bowl 362. Focusing electrodes 372 direct the electrons to a
phosphor 374 where an image is formed to be received by imaging
array 375 such as a CCD array or camera. The image area of the
phosphor 374 is much smaller than the front surface 364 so as to
reduce the image size to one compatible with the camera. In the
present invention the distance B between the target 368 and the
imaging array 375 (including any optical path through one or more
focusing lenses) is less than or equal to the radial dimension. A
of the front surface 364 gives the pancake image intensifier 361 an
extremely short form factor suitable for practice with the present
invention.
[0149] Hitherto, such form factors were avoided because they are
known to result in severe distortion of the image formed on the
phosphor 374. This distortion is accommodated in the present
invention by means of digital image processing in computer 314
which receives digitized pixel data from scanning electronics 378
connected to the imaging array 375 and corrects it according to the
correction process described above with respect to the pin cushion
correction. Accordingly, the addition of digital signal processing
allows for production of pancake image intensifier 361 in which the
separation of the imaging optics from the front of the image
intensifier is much reduced.
[0150] The above adapter may be modified to use in femur imaging.
In this case the pedestal 326 may be eliminated in favor of a
positioner (not shown) attached to the image intensifier 320 or
x-ray source 322 directly. In the former case, the positioner may
provide for a fixed air gap between the patient and the image
intensifier 320 to reduce received scatter. So as to allow free
manipulation of the C-arm 318, the positioner may be a lightweight
plastic radiolucent material and may optionally include a
calibration system such as a flip in phantom for calibration of the
dual energy readings and occluders for scatter correction as has
been described above. Collimation and/or a separate solid state
dual energy image detectors may also be held by the positioner
whose outer surface may guide the positioning of the C arm 318 to
the necessary orientation which need not be horizontal but may be
vertical for fore arm measurements or the like. For femur
measurements, the patient may stand and the C-arm 318 manipulated
appropriately as guided by the positioner.
[0151] Referring now to FIG. 31 and 32, computer 314 includes a
processor 380 and memory 382, the latter of which receives raw
image data in the form of pixels having spatial locations and
brightness values forming images 384. Memory 382 also includes a
processing program 386 providing a general interface and control of
the operation of the fluoroscopy machine 310 and a processing of
images 384 so as to provide a quantitative measure of bone isolated
from soft tissue.
[0152] The processing program 386 can be simply loaded into the
computer 314 for the fluoroscopy machine 310 when the pedestal 326
is to be employed with a fluoroscopy machine 310 providing digital
imaging and x-ray voltage control. If an independent detector array
360 is required, the program 386 may be executed on a computer 314
associated with that independent detector array 360.
[0153] At a first step in the program 386, indicated by process
block 388, the operator of the fluoroscopy machine 310, having
indicated a desire to perform densitometry and having positioned
the C-arm in the pedestal 326, enters patient data that will be
used to identify the image 384 to be collected.
[0154] At succeeding process block 390, data is collected for three
distinct images 384 with: 1) no x-ray exposure, 2) high energy
x-ray exposure, and 3) low energy x-ray exposure. Each of the
exposures is preserved as a separate image file in the memory of
the computer 314. The first exposure is used for correction
routines to be described; the latter two exposures are used to
deduce bone density according to methods well known in the art in
which variations in high energy and low energy absorption are used
to deduce the Compton scattering and atomic number of the material
lying between the x-ray source 322 and the image intensifier 320.
As is understood in the art, these two measurements allow the
amount of bone as opposed to soft tissue located in that image
region to be accurately measured. The data is acquired directly
from the independent detector array 360 or in the event that
stimulable plates are used, a reader may be attached to the
computer 314 so as to acquire the necessary pixel data of an image
384. In the same way a conventional photographic film/filter plate
arrangement may be used.
[0155] At next process block 393, each of these images is corrected
for non-linearity of the detector such as may be determined
empirically at an earlier time by testing the detector according to
methods well known in the art, and which is a function of the
detector and the technology used by the detector. Generally the
testing exposes the detector to different fluences of x-rays and
measures the output of the detector and the correction is intended
to ensure that, for example, a doubling of fluence results in a
doubling of detector output after correction. The correction is
generally simply a scaling of each of the images by a factor that
is a function of the pixel value for each pixel and possibly the
location of the pixel.
[0156] At process block 395, noise related to the particular line
of the detector is removed. Referring to FIG. 34, the imaging array
375 provides a matrix of detector elements 392 arranged in rows and
columns. Normally, either rows or columns are ganged together to be
read out by dedicated read out electronics 391 spanning a
particular row or column. The read out electronics introduces noise
which is imposed upon each detector element 392 of that row and
which is thus line correlated, that is, more highly correlated with
other detector elements 392 of the line than detector elements 392
of different lines. To eliminate line correlated noise, one
detector element 392 in each line is blocked by a lead mask 394 so
as to be shielded from x-rays. A pixel value 396 from this blocked
detector element 392 will provide a value that varies according to
the line noise thus a line correlated noise value 398 may be
deduced and subtracted from the pixel values 400 of the other
detector elements 392 in the line.
[0157] Referring again to FIG. 32, at a succeeding process block
402, veiling glare is removed and the field is flattened. This
former correction attempts to eliminate blurring of the image such
as may be caused by scatter or similar effects within the imaging
array 375. Glare refers generally to a reading that would be
obtained under detector elements 392 that were wholly shadowed by
an occluding absorber on the surface of the array 375. The glare is
a function of the detector technology and is reduced by a
deconvolution process based on an empirically derived deconvolution
kernel according to a number of techniques well known in the
art.
[0158] Also at process block 402 the field is flattened which is to
say the gain variation of the detector elements 392 are normalized
according to an empirically derived normalization map determined at
the factory by exposing the detector to a uniform x-ray elimination
and noting variation and intensities reported in the pixel values
400. At this time, dark currents from the detector elements 392 may
also be eliminated as determined from the no-exposure x-ray image
taken at process block 390.
[0159] Referring now to process block 404, a dynamic scatter
correction may be employed as has been previously described with
respect to FIGS. 21-23. Alternatively referring also to FIG. 33, a
dynamic scatter correction may be employed in which the data of the
image 384 is analyzed so as to create a histogram 406 of pixel
values 400 for the entire image. The histogram may be divided into
regions 408 (five equal regions in the preferred embodiment)
corresponding roughly x-ray paths through: 1) air-only, 2) thin
tissue, 3) thick tissue only, 4) thin bone and, and 5) thick-bone.
Each of these materials will exhibit a different scattering and
hence a different empirically derived scatter kernel 410 may be
assigned to each region 408 with generally the lower density
regions having narrower kernels commensurate with less scatter.
[0160] The selected scatter kernel 410 may be scaled by the pixel
value of the image 384 on a pixel by pixel basis and that kernel,
so scaled, applied to a deconvolver 412 used to deconvolve the
image 384 to produce a deconvolved image 414. A number of
techniques of deconvolution are well known in the art using a fixed
scatter kernel and these same techniques may be used with the
variable scatter kernel 410 described here. During deconvolution
the kernel 410 will be sequentially applied to a set of adjacent
pixel values determined by the width of the kernel. The center
pixel value at any step of the deconvolution will be used to scale
the kernel and to identify the region of the histogram for the
purpose of selecting the kernel 410. In an alternative embodiment,
the kernel may be fixed and simply scaled by the value of the
centermost pixel during de-convolution.
[0161] Referring again to FIG. 32, at process block 416, the images
384 are log corrected reflecting the fact that attenuation is
exponentially related to thickness. The images are now related to
thickness, a dimension which will be important in the ultimate
bone-density determination.
[0162] At following process block 418, speckle may be identified
for certain x-ray detectors 360 that are subject to extremely high
readout values caused by noise which is possibly related to direct
x-ray irradiation of the detector element. Speckle is identified by
a simple thresholding process.
[0163] At next process block 420, path length correction may be
performed based on the geometry of the particular C-arm such as may
vary path length and magnification across the image as is well
understood in the art.
[0164] Similarly at succeeding process block 422, beam hardening,
the well known effect of a spectral shift in a polyenergetic x-ray
beam as it passes through different thicknesses of material, and a
Heel effect correction may be made, the Heel effect correction
referring to a variation in the spectrum of an x-ray beam as a
function of its angle in the cone of x-ray beams. Both of the
corrections are known in the art, but must be employed in the
present invention in order to provide suitable quantitative
accuracy for densitometry.
[0165] At process block 424, the identified speckle of process
block 418 is corrected by eliminating these identified pixels from
subsequent calculation or by replacing them with a local average
value.
[0166] The entire image may then be averaged or low-passed filtered
at process block 426 so as to further reduce noise and to eliminate
unneeded resolution.
[0167] The images are then processed according to well understood
techniques to produce a bone mineral density value at process block
428. This bone mineral density value indicates the amount of bone
material at each pixel of the image largely independent of
surrounding soft tissue. The pixel image may be analyzed in a
number of methods but most simply, as indicated by process block
430, by defining either automatically or manually a desired region
of interest within the image and making a measurement of total bone
density within that region. Automated techniques may look for a
local maximum or minimum of bone density or may use image
recognition type techniques to locate reproducibly a particular
region of the forearm or os calcis. Morphometric analysis may be
applied to the image to detect bone fracture and other techniques
such as texture analysis may be performed according to methods well
known in the art. The results of the analyses and images so
processed may be displayed by the computer 314.
[0168] It is thus envisioned that the present invention is subject
to many modifications which will become apparent to those of
ordinary skill in the art. Accordingly, it is intended that the
present invention not be limited to the particular embodiment
illustrated herein, but embraces all such modified forms thereof as
come within the scope of the following claim.
* * * * *