Note: Descriptions are shown in the official language in which they were submitted.
BACKGRO~ D OF THE INVENTION
The invention pertains to medical x-ray apparatus,
and more particularly to an x-ray image intensifier tube of
the proximity type for medical x-ray diagnostic use. -~
The common present day x-ray image intensifier
tube is of the electrostatically focused inverter tvpe with
a 100 fold area minified output image size. This conven-
tional inverter type x-ray image intensifier tube typi-
cally has a convexly curved, six to nine inch diameter input
x-ray sensitive screen which converts the x-ray image into
a light image which, in turn, is converted into electrons
which are then accelerated and electrostatically focused onto
an output image screen which is 100 times smaller in area
than the input screen, being typically 0.6 inches to 1.0 inches
in diameter. The displayed image on the output screen can
be optically magnified and coupled to other systems for
radiographic or fluoroscopic purposes~ For example, for
radiographic purposes, the image is optically coupled to
a film camera or a photographic film. For fluroscopic
purposes, the image can be displayed either by using a sys-
tem of mirrors ~nd lenses for direct viewing-or by using a
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¦ c105c~ circuit te evision camer~ and monito~ eor remote viewin3.
~¦ The conversion efficiency oE such a conventional image
31 intensiier system is usually~ around 350,000 to 700,000 erg/cm2-R
4¦ or about 50,000 to lO0,000 cd-sec/m -R, which is about 5,000 to
51 lO,000 times the conversion efficiency of the old-time
61 fluoroscopic screen. Part of this intensification is obtained
71 as true electronic gain, which is about 50 to lO0 times over the
8¦ old-time Eluoroscopic screen. Another factor of lO0 gain is
9¦ obtained through the lO0 fold area minification of the image of
lol the output screen.
11 ¦ The image quality of the conventional inverter type image
12 ¦ intensifier tube is reasona~ly adequate for fluoroscopic use,
13 ¦ but is far short of the requirement for radiographic use. ~e
14 ¦ requirements for radiographic use are established by the
15¦ conventional film-screen system, which demands a 20~ modulation
16¦ transfer function response at between 2 to 3 line pairs per
17¦ millimeter.
18¦ Such conventional Eilm-screen systems are commercially
19¦ available in speeds ranging from 250 R l to 8000 R l. The speed
20¦ is defined as the reciprocal of the x-ray exposure in terms of L
21 t roentgens, R, to the Eilm-screen system to result in a net
22¦ optical density of l.0 on the processed film. The spatial
23¦ resolving ability of the film-screen system is generally
24 inversely proportional to the speed of the system. That is, the
25 higher the spatial resolving ability the lower the speed of the
26 system.
271 While film-screen systems have desirable system speed
2a ¦ qualities, they have the drawback that they require taking full
29 size photos which are difficult to store and which are becoming
30 increasingly more expensive due to the rising cost of the
31
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1 silver halide x-ray film. ~lso, the film cannnt be monitored
2 during exposure to control the dosage or timing.
3 A recent article published by C.B. Johnson in the
4 Proceedings of the Society of Photo Optical Instrumentation
~ngineers, Volume 35, pages 3 - 8 11973), hypothetically
6 suggests that an x-ray sensitive proximity type image
7 intensifier may be designed with an x-ray sensitive conversion
8 screen on one side of a gl3ss support and a photocathode on the
9 other side of the glass support However, the article gives no
specifics concerning the critical parameters or what might be
11 used as the x-ray sensitive conversion screen. ~ow this image
12 intensifier can be designed to result in high conversion
13 efficiency or high resolution was also not discussed.
14 A proximity device using a michrochannel plate ~i~CP) both
1~ as the primary x-ray sensitive conversion screen and as an
16 electron multiplication dPvice was des~ribed by 5. Balter and
17 his associates in Radiology, Volume 110, pages 673 - 676 (1574),
18 and by llanley et al. in U.S. Patent no. 3,394,261* According to
19 an article published by J. Adams in Advances_in Electronics and
Electron Physics, Volume 22A (Academic Press, 1966), pages 139 -
21 153, this type of device has a very low quantum detection
22 efficiency in the practical medical diagnostic x-ray energy
23 range of 30 - 100 Xev. ~he device gain of the Balter article
24 was first reported to be 20 - 3 cd-sec/m2-R which is too low to
be useful as a radiographic or fluoroscopic device. A higher
26 gain device described in the same Balter article exhibite~
27 excessive noise. There is a real question whether a practical
28 self-supporting .~CP plate with uniform gain can be constructed
29 with current technology to sizes beyond five to six inches in
diameter which is not of sufficient size to produce an output
31
32 ~Issued July 23, 1968~
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useful for radiographlc purposes.
Another approach involving proxfmity design was taken by
I.C.P. Millar and his associates and their results were publishe~ in
1) IEl~ Transactions on Electron Devices, Volume ED-18, pages 1101 - 1108
(1971), and 2) A~vances in Electronics and Electron Physics, Volume 33A,
pages 153 - 165 (1972).
Millar's approach again involves the use of a micro-channel
plate (MCP). In this device, however, the MCP is used purely as an
electron multiplication device and not as an x-ray conversion screen.
The conversion factor for Millar's tube is reported to be around 200,000
cd-see/m -R, which is abo~e or higher than needed for fluoroscopic
purposes, but is far too high for radiographic purposes. However, the
output brightness of ~illar's tube also exhibits strong dependence on the
photocathode current density. At around a photocathode current density
of 5 x 10 11 amperes/cm or at the eguivalent x-ray input dose rate of
around 0.6 x 10 3 R/sec, the output brightness of the tube starts to
become sublinear in res~onse with respect to the input ~-ray dose
rate. The sublinear response beco~es worse at higher x-ray dose rate.
This undesirable feature reduces contrast discrimination during fluoro-
seopy and is virtually useless for radiography. Again, it is unknown
whether a large format beyond six inches in diameter, self-supporting
and with uniform gain, MCP can be fabricated.
m e Millar proximity type image intensifier tube has a
glass envelope and an inwardly concave, titanium input window. The
window is deseribed as being 0.3 mm thick. Materials such as titanium,
aluminum and beryllium cause undesirable seattering of the x-rays whieh
reduces the image quality. Furtherm~re, beeause of the relatively
high p~rosity and low tensile strength properties of such materials,
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they cannot be rnade as thick as desirable to .n3xirnize their x-ray trar,~-
rnissive properties. Still another problem with tubes cons,ructed with
such rnaterials for the input window and glass for the tube envelope is
in joining the windch~ to the tube enveloFe. The rnaterials have such
dissimilar ~.errPal expansion properties, among o her differences, as
to preclude their practical ccmroercial use in a large format device.
In all such prior art x-ray image intensification devices
there is the further problern of x-ray back scatter at the output display
screen due to x-rays passing both out of the tube output window and
ccming into the tube through the output window. This can distort the
displayed iroage and pose a danger to the user of the device.
SUMM~RY OF THE INVENTION
I~ne above and other disadvantages of prior art x-ray image
intensifier tubes are overcare by the present invention of an x-ray
sensitive image intensifier tube characterized by an essentially
metallic tube envelope, an inwardly concave, m~etallic input window
in the tube envelope, the input window being made in the preferred
enibodiment of an alloy of iron, chramium and nickel, a flatr directly
viewable output phosphor display screen, a flat scintillator-
photocathode screen which is operated at a negative high Fotential
with respect to the remaining tube components includin~ the tube
envelope and the output display screen. The scintillator-photocathode
screen is suspended parallel to the output screen with insulating
posts in between the input winaow and the output screen. The image
intensifier tube of the invention has a linear response with
respect to input x-ray dose rates in excess of 0.06 R/sec.
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In the preferred embodiments, the brightness gain (conversion
efficiency) is in the range of 500 to 20,000 cd-sec/m-R, the gap spacing
between the scintillator-photocathode screen and the output screen is in
the range of 6 to 25 mm, and the thickness of -the scintillator is in the
ranse of 50 to 600 microns, whereby high x-ray utilization, high gain,
high image quality and low field emissLon are simultaneously obtaLned.
A high Z glass output window reduces x-ray back scatter and
further protects the operator of the tube from the x-rays. A collar of
iron-nickel alloy is fritted to the output window and welded to the tube
envelope fro mounting the output window in the tube enveloFe.
ALthough the image intensifier tube used in the preferred
embodiment of the invention has an essentially flat or planar input x-ray
sensitive screen, it may be slighly curved for the p~ ose of increasing
the mechanical strength of the screen, in othex em~odiments. The tube is
quite thin and ccmpact in size ccmpared to a conventional image intensi-
fier system. The input area can be square, rectangular or circular in
shaFe in the various en}cdmEnts. As discussed above, in a conventional
inverter type image intensifier tube the input screen is limited to a
circuLar disc shaFe and is ccm~only outwardLy curved.
me main advantage of this invention is the absence of three
sources of "unsharpness": the electron optics, the output phosphor
screen, and the external optics. All this is due to the large full-size
output image. Also absent are the shallcwness of the depth of field
of the electron optics and the external optics. Again, this is dlle to
the large full-size output image. The electrical field in the space
be~een the input and output screens of the image intensifier tube of
the present invention is qulte high compared to a conventional tube
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and the cathode region field strength is ab~ut 100 tim~s higher than that
of a conv_ntional tube, thus it is not sensitive to external magnetic
fields and defocusing problems encountered when subjected to bursts of
high intensity, short millisecond duration pulses.
Furthel~ore, since the metallic tube envelope and all of the
basic tube ccmFo~ents except the scintillator-photocathode screen are at
a neutral potential with respect to the output display screen, spurious
electron emission is avoided, resulting in a clearer display.
The absence of some of the sources of "unsharpness" allows
this imrention to improve the performanoe of an image intenslfier tube
in several different ways. For example, much higher gain and patient
dose reduction can be achieved by using a thicker (200 to 600 micron)
input x-ray to light conv,ersion screen and still having acceptable image
resolution. Another example is to provide a radiographic camera by
obtaining a very high image resolution at the output screen through the
use of a 50 to 100 microns thick scintillator screen and a narrower
(6 to 10 mm) photocathode to display screen gap spacing. This output
display can then be photographed.
In the preferred emkodiment of such a radiographic camera
according to the invention, reduction type optics focus the full size
output display onto photographic film which is smaller in diagonal
di~ension than the output display screen. The film sensitivity (G)
is defined as the reciprocal of incident li,aht energy in ergs per square
centimeter (erg/cm ) which is required to produce a net density of 1Ø
More specifically the film sensitivity is chosen to be in the range of
5 to 100 cm /erg. me image intensifier tube is chosen to have a
conversion efficiency (C) in the range of 1,000 to 30,000 er~/cm -R, or,
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if the output phosphor is green e~tting, in the ranc3e of 140 to 4,300
cd-sec/m -R. The fractional light energy (T) emit~ed by the output screen
which is collected by the o~tics and which is transferred to the photo-
~raphic film can be ap~roximat ~ by the relationship:
4f2~1+m)2
where,
t = transmission of the o~ptical system
f = the f number of the optical system, and
m = magnification of the image, or ratio of
image to object size, and
is approxLmately in the range of 1 x 10 to 1 x 10 . In this embo~lment,
the total speed of the camera (S = CTG) in the medical diagnostic region
of the x-ray spectrum, i.e., 30 - 100 Kev, is in -the range of 100 to
10,000 R . In a preferred e~bcdiment, where the conversion efficiency
(C) is in the range of 3,000 to 14,000 erg/cm -R, the system speed (S)
is in the range of 500 to 5,000 R
The x-ray sensitive photographic camera according to one
enkKx~m~nt of the invention is thus designed to have a system speed
which is optimal to take maximum advantage of the amount of information
provided by the incident x-ray quanta such that the recorded image will
have a kalanced image quality for the x-ray information. me image
quality of the photographs produced by the system of the invention
is as good as that of conventional cassette film-screen systems, which
is not achievable with conventional inverter type image intensifier
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systems. Hcwever, with the camera of t}~ inventLon, cmaller than f~
size films can be used with no loss of x-ray info~ tion. Ihis allo~s
for a significant reauction in required storage space for the developed
films. Also the camera can be mDdified by a beam splitting mirror to
simultaneously generate a second photograph of the x-ray information
A second optical system, placed off axis, may also ke used to ~enerate
the second photograph.
One of the significant features of the ca~era system is
that the long focal length, in excess of 100 mm optics in the preferred
em~dinent, the n~n-mLnified output image size and small aper-ture optical
system give the system greater toleran oe for thermal expansions, di-
mensional changes, etc. than a conventional image iiltensifier x-ray camera
system which is extremely sensitive to such changes. Also, the optical
system can be folded so that the camera system can be made more ccmpact,
which is an im~ortant feature in a cramped radiological examination roon,
than can a conventional system of comparable input format size.
Moreover, the image intensifier of the present system allcws
stereo x-ray photographs to ke produced with no image distortion. m is
is primarily due to the fact that the input x-ray conversion screen is
flat (planar) as opposed to the conventional curved input screen of the
other prior art image intensifier tubes.
Still another advantage is that the x-ray sensitive area
input format size of the camera system can be expanded without
sacrificing image quality as would happen with conventional inverter
type image intensifier systems. A still further advantage of
the present system is that it can be easily
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1 ¦photo-timed with a sensing device directly monitorin~ th~
2 ¦output image to obtain consistent e~posure on each ilm.
3 ¦ Unlike the proximity x-ray image intensiiers heretofore
4 ¦discussed, the x-ray image i~tensifier tube of the present
5 ¦invention achieves high conversion efficiency without requiring
6 ¦the use of additional multiplicatio~ means or non-linear
7 responding components, i.e., a micro-channel plate ~etween the
8 ¦output phosphor screen and the photocathode. As a result, the
9 ¦x-ray image intensifier tube of the present invention is
10 ¦mechanically simpler, more reliable and exhibits a linear
11 ¦response with respect to input x-ray dosages in excess of 0.06
12 ¦R/sec, the dosage used for medical diagnostic purposes.
13 1 Among the many advantages of the invention are the light
14 ¦weight, the simplicity of the tube and its compact size. For
15 ¦example, wben the tube is used in a direct view, fluoroscopi^
16 ¦mode, the physician can have easy access to the patient for
17 ¦palpation and can observe the effects of palpation without
18 having to turn away from the patient, as is necessary in the
19 ¦present day systems having an inverter type image intensifier
20 ¦coupled to a television display.
21 ¦ In other embodiments of the invention, such as for use in
22 ¦teaching institutions, for example, it may be desirable to
23 provide remote displays of the output of the x-ray image
24 intensifier tube's large output display screen which is quite
Z5 easily coupled to a silicon intensifier target (SIT) tube type
26 closed circuit television system for remote viewing or for
27 vi~eo recor~ing,
28 Still another advantage is that the x-ray sensitive area
29 ¦input ormat size of the system can be expanded without
30 ¦sacrificing image quality as would happen with conventional
31 ~
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inverter t~e image intensifier systemS-
It is therefore an object of the present in~ention to provide
a pro ~nity type, x-ray sensitive image intensifier tube having a metal
input window which minimizes x-ray self scattering and back scattering
effects.
It is another object of the invention to provide an x-ray
image intensification tube having a flat x-ray conversion input screen
to reduce image distortion.
It is yet another object of the invention to provide a
panel type x-ray irnage intensifier tube of rugged design for medical
diagnostic purposes which minirnizes the danger of injury to the patient
resulting from implosion of the tube.
It is a further object of the invention to provide a panel
type x-ray image intensifier tube having a nearly full size output display
~hich is aligned with that portion of the patient which is being irradiated
by the x-rays.
It is a still further object of the invention to provide
an x-ray image intensifier tube capable of having either a square,
rectangular or circular or other freely shaped input format, and that
; 20 the format size is expandable to 17 x 17 inches.
It lS yet a further object of the invention to provide an
x-ray image intensifier tuke which is not sensitive to the effects of
~ltage drifts, external magnetic fields, and field ~nission.
It is still another object of the present invention to
provide an x-ray radiographic camera having a sy~tem speed and irnage
quality comparable to conventional film screen systems.
; It is also an object of the invention to provide an x-ray
radiographic cam~!ra utilizing a directly viewable reduced size
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film, with a film size smaller than the inp~]t x-ray irnage size;
and
It is yet a further object of the invention to provi~e an
x-ray radiographic camera having lonq focal length optics tc
increase the dimensional stabili~y tolerance of the system.
In accordance with the present invention there is provided
an x-ray sensitive image intensifier tube having a tube envelope,
the tube characterized by a metallic input window in the tube
en~elope, a flat, halide activated, alkaline halide scintillator
screen adjacent the input window for converting the x-ray image
into a light pattern image, a flat photocathode layer parallel
and immediately adjacent to the scintillator screen for emitting
photoelectrons in a pattern corresponding to the light pattern
image, a flat phosphor display screen parallel to and spaced
apart from the photocathode layer with the space between them
being an uninterrupted vacuum, the scintillator screen, the photo-
cathode layer and the display screen all having diagonal dimen-
sions at least equal to the actual size of the x-ray image to be
intensified, and connections for applying an electrostatic potent- :
ial from an external source solelv between the display screen and:
the photocathode layer to accelerate the pattern of photoelectrons
toward the display screen along parallel, straight trajectories ~ :
to impinge upon the output display screen. :
The foregoing and other objectives, features and advantages~ :
: of the invention will be more readily understood upon consideration
of the following detailed description of certain preferred embodi-
ments of the invention, taken in conjunction with the accompanying
drawings.
13 -
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BRIEF D~SCRIPTIiN OF TH~ URAWING8
1 _~ _
3 FI~. l is a diagrammatic illustraiton of a
4 conventional inverter type image intensifier x-ray tube;
FIG. 2 is a diagrammatic illustration of the x-ray image
6 intensifier tube according to the invention;
7 FIG. 3 is a detailed vertical view, in section, of the
8 image intensifier tube of the invention;
9 FIG. 4 is an enlarged, vertical view of the encircled
detail in FIG. 3~ illustrating a cross-section of a portion of
11 the image intensifier tube depicted in FIG. 3;
12 FIG. 5 is a vertical, sectional view, taken generally along
13 the line 5-5 in FIG. 3, of the image intensifier tube according
14 to the invention;
FIG. 6 is a diagrammatic illustration of the x-ray
16 radiographic camera according to the invention; and
17 FIG. 7 is a graph relating the design parameters of the x-
18 ray radiographic camera for a commercially available
19 ¦photographic tilm.
23
26
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1 i DET~ D DE~;CRIE~TIO~i OF Ti~E PRE:FE:KRI~'D l~,~lBODIMI~N'l'
I _
2 ¦ Referring now more particularly to ~IG. 1, a conventional
3 ¦inverter type x-ray image intensifier tube is illustrated. ~n x-
4 ¦ray source 10 generates a beam o~ x-rays 12 which pass thr~ucJh
5 ¦the patient's body 14 and casts a shadow image onto the face of
6 la camera system 16. The camera system includes a conventior,al
7 ¦inverter type image intensifier vacuum tube 18. ~he tube 18 has
81 an outwardly convex input winclow 20 and a correspondingly cGnvex
9 ¦seintillator screen and photocathode assembly 22. The purpcse
10j of this seintillator screen, as is well known to those skil~ed
11 ¦in that art, is to convert the x-ray shadow image into a li~.ht
12 ¦ image, which, in turn, is immediately eonverted by the
13 ¦photocathode layer into a pattern of electrons. This pattern of
14 ¦electrons is electrostatically accelerated by a set of
electrodes 24 and anode ~5 near the display screen 28 and ic
16 ¦focused by this set of electrodes 24 and anode 25 to form an
17 ¦image on the small output screen 28. The electrodes 24 and
18 ¦anode 25 are connected to a bigh voltage source 26 whose other
19 lead is eonneeted to the scintillator and photocathode screen
20 ¦assembly 22. The tube body is made of insulating glass. The
21 ¦image at the output display screen 28 is magnified by a short
22 ¦focal length optical system 30 and is projected onto suitable
231 recording media, such as film 32. The image could also be
24 projected onto the sensitive area of the closed-cireuit
television camera for display on a elosed cireuit monitor in a
26 1uoroscopic mode.
27 The brightness gain of the image by the tube 18 is due
28 ¦partly to the electron acceleration and partly to the result of
29 ¦eleetronie image miniication. This is the result of redueing
the image generated on the scintillator screen 22 down to a
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1 ¦relatively small image at the output display screen 28. The
21 reduced image on the display screen 28 is too small however, to
31 allow direct viewing without optical aids. I~loreovcr, the
41 quality of the image is reduced both by the quality of the
5¦ electron optics and by the quality of the output phosphor screen
¦in the electronic image minification, and by the subse~uent
7¦ enlarging of the output image onto the film or onto the monitor
81 screen by the closed circuit television system.
9¦ Another disadvantage is that because of the curved
10¦ scintillator screen 22, there is a spatial distortion produced
11 ¦ in the image due to x-ray projection on the curved surface and
12 ¦due to the field configuration in the tube. Still another
13 ¦problem is that because of the weak field near the cathode
1~ ¦ region and the multi-electrode arrangement 24, the tube 18 is
15 ¦extremely sensitive to external magnetic fields and voltage
6¦ drifts among the electrodes. Both of these factors can cau~e
.71 distortion and unsharpness in the produced image.
18 ¦ Yet another problem is that because of tne greatly minified
19 ¦output image and the short focal length optics 30, any change in
the positioning of the elements of the optical system with
21 respect to the photosensitive layer of the camera tube or the
22 ¦output screen 28, will render the image out of focus. This can
23 ¦result from vibration or from thermal expansion.
24 ¦ One other major disadvantage of the conventional
25 ~ystem is that because of the curved glass window 20 which
26 is necessary to withstand the pressures due to the vacuum
27 ¦inside the tube 18 and the already very weak ficld strength
28 ~n the cathode region, the system is limited to
29 ~pproximately nine inches in input format for optimum
30 performance. Any greater diameter input will necessitate a muc~l
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1 higher tube voltage and a tl1ickcr input window ~hicl) would
2 cause increased problems due to ion spots inside ~he tube ancl x-
3 ray transmission and scattering in the input window. ~ven in
4 the conventional sized tubes, there is also, of course, the
danger to the patient and the radiologist that the tube might
fracture causing an implosion and resulting ejection of the
7 glass fragments.
8 ReEerring now more particularly to FIG. 2, a panel shaped
9 proximity type x-ray image intensifier tube according to the
invention is illustrated. The image intensifier tube 34
11 comprises a metallic, typically type 304 stainless steel,
12 vacuum tube envelope 36 and a metallic, inwardly concave input
13 window 38. The window 38 is made of a specially chosen metal
14 foil or alloy metal foil in the family oE iron, chromium, and
nickel, and in some embodiments, additionally combinations o~
16 iron or nickel together with cobalt or vanadium. It is
17 improtant to note that these elements are not customarily
18 recognized in the field as a good x-ray window material in the
19 diagnostic region of the x-ray spectrum. ~y making the window
thin, down to O.l mm in thickness, the applicant was able to
21 achieve high x~ray transmission with these materials and at the
22 same time obtain the desired tensile strength. In particular,
23 a foil made of 17-7 PH type of precipitation hardened chromium-
24 nickel stainless steel is utilized in the preferred embodiment.
This alloy is vacuum tight, high in tensile strength and has
26 very attractive x-ray properties: high transmission to primary
27 x-rays, low self-scattering, and reasonably absorbing with
23 respect to patient scattered x-rays. The window 38 is concaved
29 into the tube like a drum head.
The use oE materials which are known for high x-ray
31
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1 transmission such as berylli~m, aluminum an~ titanium for
2 example cause the undesirable scatteting which is present in
3 some prior art proximity type, x-ray image intensifier devices. t
4 One purpose of having a metallic window 3B is that it can
5 be quite large in diameter with respect to the prior art type
6 of convex, glass window 22, as depicted in FIG. 1, without
7 aEfecting the x-ray image quality. In one emoodiment, the
B window measures 0.1 mm thick, 25 cm by 25 cm and withstood Gver
9 100 pounds per square inch of pressure. The input window cOn
10 be square, rectangular, or circular in shape, since it is a
11 high tensile strength material and is under tension rather than
12 compression.
13 The x-ray image passing through the window 38 impinges
14 upon a flat scintillation screen 40 which converts the image
15 into a light image. This light image is contact transformed
16 directly to an immediately adjacent flat photocathode screen 42
17 whicn converts the light image into a pattern of electrons.
18 The scintillator and photocathode screens 40 and 42 comprise a
19 complete assembly 43. The electron pattern on the negatively
20 charged screen 42 is accelerated towards a positively charged
21 1at phosphor output display screen 44 by means of an
22 electrostatic potential supplied by a high voltage source 46
23 connected between the output screen 44 and the photocathode
24 screen 42. Although the display screen 44 is positive with
25 respect to the scintillator-photocathode screen assembly 43, it
26 is at a neutral potentional with respect to tihe remaining
27 elements of the tube, including the metallic envelope 36, to
28 thereby reduce distortion due to field emission. No other
29 elements such as a microchannel plate, for example, are
30 interposed between the output phosphor screen and the
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1 ¦photocathode screcn as is done in some prior embodiments.
2 ¦ The use of such non-line~r devices (with respect to input x-
31 ray dosage~ cause distortion in and o~ themselves but they also
4 ¦increase the deleterious ield emission effects since some o~
51 the elements of the microchannel plate must operate at different
I electrostatic potentials with respect to the output display
7l screen and thereby become sources for spurious electron
8 ¦emission.
9¦ It should be noted that substantially no focusing takes
10 ¦place in the tube 34 as opposed to the prior art type tube l8 in
11 ¦FIG. l. The screen 40, the photocathode layer 42 and the
12 ¦display screen 44 are parallel to each other. Also, the gap
13 ¦spacing between the photocathode 42 and the display screen 44
14 ¦are relatively long, in the range of 6 - 25 millimeters, thereby
15~ reducing the likelihood of field emission and at the same time
16 keeping the electrostatic defocusing to a tolerable level, that
17 ¦is, acound 2.0 to 5.0 line pairs per millimeter.
18 Furthermore, the applied voltage across the gap between
19 Iphotocathode layer 42 and the display screen 44 is in the range
20 ¦lO,000 to 60,000 volts (lO to 60 Kv) which is higher than in
21 ¦~illar's tube, described earlier in this application. In
22 ¦addition, the non-focusing nature of the field avoids the ion
23 spot problem which plagues inverter type tubes. In the
24 preferred embodiments of the invention, the spacing between the
photocathode screen 42 and the output display screen 44 is
25 ¦bet~een 6 mm (at 15 Kv) and 25 mm (at 60 Kv). Thus, the voltage
27 per unit of distance, i.e., the Eield strength, is at least 2
28 Xv/mm. An upper limit to the field stren~th is about 5 Kv/mm.
29 ¦In prior art devices such a high field strength was not
considered feasible for thLs application of an image intensifier
31
32
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1 device becau5e of the field emission problems discussed ~bove
2 and ~hich are obviated in the applicant's device by having all
3 of the tube elements, save for the photocathode-scintillator
4 scleen assembly, be at a neutral potential with respect to the
8 output display screen.
6 The scintillation screen 40 can be calcium tungstate
7 (CaWO4) or sodium activated, cesium iodide (CsI~Na)) or any
8 other type of suitable scintillator material. However, vapor
deposited, mosaic grown scintillator layers are preferred for
the highly desired smoothness and cleanliness. Since such
11 materials and their methods of application are well known to
12 those skilled in the art, see for example, U.S. Patent No.
13 3,82~,763, they will not be described in gr~ater detail.
14 The overall thickness of the scintillator screen 40 is
chosen to be 50 to 600 microns thick to give a higher x-ray
16 photon utilization ability than prior art devices, thereby
17 allowing overall lower patient x-ray dosage levels without a
lB noticeable loss of guality as compared to prior art devices.'
19 This is because the format of the tube and the absence of
several sources of ~unsharpness~ give an extra margin of
21 sharpness to the image which can be traded o~f in favor of lower
22 patient dosage levels with greater x-ray stopping power at tne
23 scintillator screen 40.
24 Similarly, the photocathode layer 44 is also of a material
wPll known to those skilled in the art, being cesium and
26 antim~ny (Cs3Sb) or multi-alkali metal ~combinations o~ cesium,
27 potassium and sodium~ and antimony.
28 The image produced on the phosphor screen 44 is the same
size as the inlput x-ray image. The output phosphor screen 44
3D can be of the well known zinc-cadmium sulfide type ~ZnCdS~Ag))
31 *Issued July 23, 1974 to Ligtenberg et al.
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1 or zinc sulfide type (2nS(Ag)) or a rare earth material like
2 yttrium oxysulfide type (Y2O2S~Tb)) or any other suitable high
3 efficiency blue and/or green emitting phosphor material. The
4 interiorly facing surface of the output screen is covercd with a
metallic aluminum film 48 in the standard manner. The phosphor
6 layer constitu~ing the screen 44 is deposited on a high Z glass
7 output window 50. By high z is meant that the window glass has
8 a high concentration of barium or lead to reduce x-ray back
scatter inside and outside the tube and to shield the
radiologist from both primary and scattered radiation.
11 An important factor in determining the usefulness of any x-
12 ray image intensifier system for medical diagnostic purposes is
13 the conversion efficiency of the tube. The conversion
14 efficiency of the image intensifier tube is measured in terms
of output light energy in ergs per square centimeter per x-ray
16 input dosage of l roentgen (erg/cm -R), which can also be
17 expressed in terms of candlas-second per square meter-roentgen
1B (cd-sec/m2-R) if a green emitting output phosphor like
19 znCdS(Ag) type is used.
Several nine inch diameter working proximity type image
21 intensifier tubes have been constructed according to the
22 invention with a conversion efficiency in the range of 3,',00 to
23 60,000 erg/cm2-R. The output phosphors are of the ZnCdS~Ag)
24 type and thus the conversion efficiency can also be expressed
in photometric terms as S00 to 8000 cd-sec/m -R. This is about
26 equivalent to a brightness gain o~ 50 to 800 times over that of
27 the old-time fluoroscopic screens for example.
28 It is important to compare these results with those
29 reported in the Millar article referred to above~ The overall
conversion efficiency of Millar's tube is 196 to 200 cdm mR
31 ,
32
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:
sec or 196,000 to 200,000 cd-sec/m -R which is obtainecl with the M~P
operating at 10,000 gain. Removing the M~P and its gain would result
in a conversion efficiency around 20 cd-sec/m2-R, which is too low.
Therefore, Millar's article has the effect of leading away from the pre~ent
invention.
Referring now more particularly to FIG. 4, in an enlarged
cross-sectional view, the details of the scintillation and photocathode
screen assembly 43 and the output display screen assembly 44 are illus-
trated. The screen assembly 43 comprises a scintillator layer 40 of
very smooth calcium tungstate or sodium activated cesium iodide which
is vaFor deposited on a smoothly polished nickel plated aluminum substrate
or an anoclized aluminum substrate 52 which faces the input ~lindow 38.
The techniques of such vapor deposition processes are known to those
skilled in the art, see for example, U.S. Patent No. 3,825,763. For
direct viewing purposes, the layer 40 is between 200 to 600 microns
thick. For radiographie purposes, the layer 40 could be thinner
(50 - 200~), i.e., the image eould ke less bright.
As mentioned akove, the purpose of the scintillator screen
40 is to eonvert the x-ray image into a light image. On the surfaee
of the seintillation layer 40 ~hich faces away frcm the substrate 52,
a thin, eonductive, transparent electrode layer 54 sw h as a vap~r de-
posited metallic foil, i.e., titanium or nickel, is deposited and on
- top of this is deposited the photocathode 42. The photocathode layer
42 converts the light image fram the scintillator layer 40 into an
electron pattern image and the free electrons from the photocathode 42
are accelerated by means of the high voltage potential 46 toward the
display sereen 4~, all as mentioned above. me seintillator-photoeathode
-22-
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screen 43 in this invention is suspenc~cl frc~m the tube envelope 36 be~"e-en
the input window 38 and the output screen 44 by se~eral insulatinc3 posts
5~. One or mlore of these posts may be hollow in the center -to allow a high
voltage cable 60 from the source 46 to be inserted to provide the
scintillator-photocathode screen 43 at the layer 54, with a negative
high potential. The remaining parts of the intensificaticn tube in-
cluding the metallic envelope 36, are all operated at ground potential.
This concept of minimizing the surface area which is negative with re-
spect to the output screen results in reduced field emission rate
inside the tube and allows the tube to be operable at higher voltages
- and thus higher brightness gain. It also minimizes the danger of
electrical shock to the patient or workers if one sh~uld sc~ehow come
in contact with the exterior envelope of the tube.
To reduce charges accumulated on the insulating posts 58,
they are coated with a slightly conductive material such as chrGme oxide
w~ich bleeds off the accumulated charge by providing a leakage path of
less than 20 Kv/cm.
m e thick, high atomic number (Z) glass substrate 50 on
which the phosphor dlsplay screen 44 is de~osited forms one exterior
end wall of the vacuum tube envelope 36. This glass substrate 50 is
attached to the tube envelope 36 by means of a collar 54 made of an
iron, nickel, chro~ium alloy, designated to the trade as "Carpenter,
No. 456". Since the thermal coefficient of expansion of this alloy
matches that of the glass and nearly matches that of the tu~e enveloFe
36, the collar 54 can be fritted to the glass substrate 50 and welded
to the tube envelope 36. On the interior surface of the glass
wall 50 is deFosited the phosphor layer 44 which is backed by
-23
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1 ¦ a protective and electron transparent aluminum thin ~ilm 48 to
2¦ prevent light feedback and to provide a uniform potential. It
31 also tends to increase the reElection of the phosphor layer ~4
4I to give a higher li~ht output gain.
5I The essentially all metallic and rugged construction of
61 the tube minimizes the danger of implosion. The small vacuum
7I space enclosed by the tube represents much smaller stored
81 potential energy as compared with a conventional tube which
9¦ further minimizes implosion danger. Furthermore, if punctured,
the metal behaves differently from glass and the air simply
11 ¦leaks in without fracturing or imploding.
12 ¦ The photocurrent drawn by the tube from the power supply
13 ¦ 46 is dependent, of course, on the image surface area of the
14 ¦ scintillator-photocathode screen assembly 43 and the output
15 ¦display screen 44. For a tube used for direct viewing, the
16¦ photocurrent would be 0.4 to 0.8 x 10 amperes/cm at an x-ray
17 ¦dosage level of 1 mR/sec.
18 I The applicant has studied other thin metal alloys in the
19 ¦chromium-nickel stainless steel facility as window materials,
20 ¦and found that these alloys are also better than the well known
21 ¦x-ray window materials like beryllium and aluminum but not as
22 ¦good in overall performance as the 17-7 PH stainless steel.
23 ¦These other materials are: precipitation hardened type 15-7
24 ¦ MO and work hardened type 304.
25 The applicant has also found that thin foils of above- -
26 ¦mentioned alloy windows are very satisfactory for use as x-ray
27 ¦windows in high vacuum devices like x-ray image intensi~ier
2B tube as long as the thickness is under 0.25 mm. ~t 0.125 mm
29 ¦thickness, the x-ray transmission through the 17-7 P~l foil is
94~ for 120 Kvp x-rays filtered with 23 mm aluminum, 88~ for
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1 80 ~vp x-rays filtered with 23 mm aluminum, and 80~ for 60 Kvp
2 x-rays filtered with 23 mm aluminum.
3 Referring now more particularly to FIG. 5, the x-ray
4 camera 100 according to the invention is illustrated. The
camera 100 includes the proximity type image intensifier tube
6 34 described above, a long focal length optical system 138 and
7 a film 140. As mentioned above, in prior art, conventional
8 radiographic, image intensifier systems the optical system
9 magnifies not only the small output image but all the minute
1~ deEects which may be present in the output screen as well,
11 resulting in a need for a more critical manufacturing process.
12 In the present invention the optics 1~8 reduce the size of the
13 image and, correspondingly reduce the apparent size of defects
14 which may be present in the out~ut screen, resulting in a
higher yield, less expensive and less demanding manufacturing
16 process. The originally displayed image at the output screen,
17 however, is much larger than in the conventional tube so that
18 the reduced image at the film 140 is o~ betteL quality than in
19 conventional systems.
The large output image size combined with the long, in
21 excess of 100 mm, focal length of the optical system 13~ in the
22 preferred embodiment makes it less sensitive to thermal
23 expansion than conventional systems. The film 140 is held in a
24 film transport 154 which allows the film to be advanced to take
pictures in a serial manner.
26 ~requently the films are better viewed with the
27 emulsion side facing the radiologist. In order to obtain
28 the proper orientation, a mirror 170 (or 3 mirrors or any
29 odd number of mirrors), shown in hidden line fashion in
~_ ___ 30 FIG. 6, can be inserted into the optlcal path resultlng in
31 ~ ;
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1 a new film holder position. ~irror 170 can also be made of a
2 partially transmissive mirror ~a beam splitter) so that two '.
3 films can be made with a single x-ray exposure.
4 The total system speed of the camera lO0 of the invention
in the medical diagnostic region of the x-ray spectrum, tha~ is
30 to lO0 Kev, is in the range of 500 to 5,000 R . The system
7 speed is defined as the reciprocal of the x-ray radiation
8 dosage incident on the output window of the x-ray image
9 intensifier tube 3~ in terms of roentgens (R) required to
produce a net density of 1.0 on the photographic film 140. The
11 system speed can be expressed by the following slmplified
12 formula S = C T G, where
13 C = conversion efficiency of the image intensifier
14 in terms of output light energy in ergs per
sguare centimeter per x-ray input dosage of
16 l roentgen (erg/cm -R), which can also be
17 expressed in terms of candelas-second per
18 square meter-roentgen (cd-sed/m2-R).
19 T = fractional light emitted by the output screen . .
collected by the optical system transferred
21 to the photographic film which can be
22 approximated by: T = ~ e2 (l+~2 '
23 where t = transmission of the lens,
24 ~ = the f number of lens, m = the
magnification of the image.
26 G = photographic sensitivity of the film in the
27 spectral region of the emission of the output
28 phosphor in terms of the reciprocal of the
29 incident light energy per square centimeter
30 in erg/cm2 which is required to produce a
. .. . _ .. . . . ... . ._ . ..
31 net density of lØ : :
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1 ¦ Therefore, the same system speed can be arrived at through
21 many different combinations of the C, T and G. FIG. 7 is an
31 illustrative example of the interlinking nature of these system
41 parameters. FIG. 7 shows the desired operating region of the
51 invention, the shaded area, Eor a commercial rapid-processable
6I single-emulsion x-ray film marketed by Eastman Kodak Company
7¦ under the brand name of type 2541 RP/FC film. The key
81 parameter of the optical system, f/ ~F , is plotted against
9¦ changes in the conversion efficiency of the image intensifier
tube, C, to achieve the system speed range of 500 to 5,000 p 1
11 ¦ The image magnification of m = 0.6 is selected for the purpGse
12 of illustration. The system speed in this case can be
13 ¦approximated by the formula 2
14 ¦ S = 1.2
15 I
16¦ If a green emitting output phosphor like 2nCdS(Ag) type is
17 ¦used, the conversion efficiency in terms of cd-sec/m2-R may
18 also be used. This scale is also provided in FIG. 7 for
19¦ reference.
20 ¦ It is important to add here that several nine inch
21 ¦diameter working proximity type image intensifier tubes
2Z according to the invention have been constructed with a
23 conversion efficiency in the range of 3,000 to 10,000 ery/cm2-R.
24 The output phosphors are of the ZnCdS(Ag) type and thus the
25 ¦conversion efficiency can also be expressed in photometric
2~ terms as 400 to 1,400 cd-Sec~m -R. A prototype system
27 incorporating these tubes, a f/2 optical system with m = 0.6,
28 and Rodak RP~FC film, achieved image quality of accepted film-
29 screen systems and a system speed in the range of 1,000 to
3,000 R 1.
31
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1 ¦ It is again important to compare these results with those
2 ¦reported in the ~illar article referred to above. The overall
3 ¦conversion eEficiency of ;~illar's tube is more than l00 times
4 ¦the optimum requirement foL radiography. On the other hand,
5 ¦removing the MCP and its gain would result in a conversion
6 ¦efficiency which is too low for radiography purposes.
7 ¦Therefore, Millar's article has the effect of leading away from
8 ¦the radiographic camera system of the present invention.
9 The designed system speed is optimized to take maximum
advantage of the amount of information provided by the incident
11 x-ray quanta, such that the recorded image will have balanced
12 image quality and x-ray information. This avoids the problem of
13 a low system speed, i.e., less than the old photofluorographic
14 camera, where the x-ray information is not fully utilized and
unnecessary patient radiation dosage results. It also avoi~s
16 the problem of an unnecessarily high system speed, as in th~
17¦ case of the conventional inverter type of image intensifier tube
18¦ system, or the ~lillar, ~CP type proximity tube, where the film
~91 is exposed with a very small amount of the x-ray information so
20¦ that the recorded photo contains an insufficient amount of
21 ¦ information with a resulting mottled or grainy picture.
22¦ Referring back to FIG. 6, the beam splitting mirror l70 can
231 be made such that the larger portion of the light beam is
241 directed to one film while the smaller portion of the light beam
is directed to a second film. This arrangement has many
261 advantages in obtaining radiographs in cases where wide latitude
27 of x-ray intensity is encountered. For example, the x-ray
28 intensity after passing through a chest normally would exhibit
291 wide differences between the lung region and the region behind
the heart. In this case, an over-penetrated or over-exposed
32
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09 rjlt It3
1 ¦ record can be made of the region behind the heart on one film
21 and the normal lung field can be recorded on the second film;
3¦ all this done with a single x-ray exposure.
4¦ The terms and expressions which have been employed here are
51 used as terms of description and not of limitations, and there
6¦ is no intention, in the use of such terms and expressions, of
71 excluding equivalents of the features shown and described, or
81 portions thereof, it being recognized that various modifications
¦a~e po:sible wit~in the cope of the inve~tion claimed.
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